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10. 1001/jamacardio. 2016. 2750
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JAMA cardiology
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Potential Strategies to Address the Major Clinical Hurdles Facing Stem Cell Regenerative Therapy for Cardiovascular Disease: A Review
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Importance While progress continues to be made in the field of stem cell regenerative medicine for the treatment of cardiovascular disease, significant barriers to clinical translation still exist that have thwarted the delivery of cell therapy to the bedside. Objective The purpose of this review is to summarize the major current hurdles for the clinical implementation of stem cell therapy and discuss potential strategies to overcome them. Evidence Review Information for this review was obtained through a search of PubMed and the Cochrane database for English language studies published between January 1, 2000 and June 15, 2016. Ten randomized clinical trials and eight systematic reviews were included in this review. Findings One of the major clinical hurdles facing the routine implementation of stem cell therapy is the limited and inconsistent benefit observed thus far. Reasons for this are unclear but may be due to poor cell retention and survival, as suggested by numerous preclinical studies and a handful of human studies incorporating cell fate imaging. Additional cell fate imaging studies in humans are needed to determine how these factors contribute to limited efficacy. Treatment strategies to address poor cell retention and survival are under investigation and include the following: 1) co-administering of immunosuppressive and pro-survival agents, 2) delivering cardioprotective factors packaged in exosomes rather than the cells themselves, and 3) using tissue engineering strategies to provide structural support for cells. If larger grafts are achieved using the aforementioned strategies, it will be imperative to carefully monitor the potential risks of tumorigenicity, immunogenicity, and arrhythmogenicity. Conclusions and Relevance Despite important achievements to date, stem cell therapy is not yet ready for routine clinical implementation. Significant research is still needed to address the clinical hurdles outlined herein before the next wave of large clinical trials is underway.
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No full text available
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10. 1001/jamanetworkopen. 2021. 22607
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JAMA Network Open
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Effect of a Novel Macrophage-Regulating Drug on Wound Healing in Patients With Diabetic Foot Ulcers
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Key Points Question Can the topical application of ON101 cream demonstrate a superior therapeutic benefit in wound healing among patients with diabetic foot ulcers (DFUs) compared with standard care? Findings In this randomized phase 3 clinical trial of 236 patients with DFUs, topical application of ON101 with gauze immediately after debridement demonstrated significant healing efficacy compared with an absorbent dressing in all patients, including those with DFU-related risk factors. Meaning Topical treatment with ON101 resulted in improved healing of DFUs.
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Introduction Approximately 80% of lower limb amputations are preceded by chronic diabetic foot ulcers (DFUs), resulting in a heavy burden of medical care and expenditure. 1, 2 The current treatment for DFUs in clinical practice focuses primarily on local wound care, including debridement, off-loading, infection control, and maintaining a moist environment with dressings, 3, 4 whereas adjunctive therapies such as the use of growth factors, tissue engineering products, hyperbaric oxygen, and negative pressure wound therapies are applied if the DFUs worsen. 5 Although current treatments featuring tissue repair or the use of anti-inflammatory agents might help in closing or controlling the progression of DFUs, most of these treatments are not well supported by clinical evidence or are not recommended for routine care by the International Working Group on the Diabetic Foot. 6 In addition, the annual increase in amputations also suggests that treatment improvement is needed. 7 Diabetic foot ulcers are pathologically complex mostly because the ulceration is undermined by the existence of multiple risk factors, such as poor patient adherence to treatment, severity of the ulcer, ulcer location and duration, vascular condition, control of glycated hemoglobin (HbA 1c ) levels, smoking habits, and kidney dysfunction. 8, 9 These factors impose a significant clinical need for novel and effective interventions to tackle this life-debilitating and life-threatening disease. Accumulating scientific evidence has revealed that targeting macrophage phenotypes might be a potentially effective therapy in DFUs because hyperglycemia increases the ratio of proinflammatory M1 to proregenerative M2 macrophages. 9, 10, 11, 12, 13, 14 ON101 (supplied by Oneness Biotech Co, Ltd; previously given the research code WH-1) exerts its therapeutic effect through regulation of the balance between M1 and M2 macrophages. ON101 is composed of 2 active pharmaceutical ingredients: PA-F4 from an extract of Plectranthus amboinicus and S1 from an extract of Centella asiatica, 2 medicinal plants reported to have significant pharmacological activities in wound healing. 10, 11, 12 With 48 in vitro and in vivo studies performed, these 2 ingredients, which contribute to a synergistic effect on regulation of the M1:M2 macrophage ratio, have been defined and formulated in a cream base using a proprietary formula. One of these ingredients, PA-F4, significantly attenuates M1 macrophages by suppressing the NLRP3-mediated inflammasome pathway and the production of downstream inflammatory cytokines such as interleukin 1β and interleukin 6, 13 which arrest the inflammation phase. On the other hand, the extract of C asiatica has been reported to activate M2 macrophages by increasing collagen synthesis and by stimulating fibroblast proliferation and the migration of keratinocytes. 14, 15 ON101 has been further demonstrated to accelerate wound healing efficiently in a db/db mouse model of diabetes, obesity, and dyslipidemia by decreasing inflammatory M1 macrophage activity and enriching M2 macrophage populations through granulocyte colony-stimulating factor–mediated M2 polarization, which changed the ulcer status from the inflammatory phase to the proliferation and remodeling stages (eFigure 1 in Supplement 2 ). A clinical pharmacokinetic study on 12 patients with DFUs showed that topical administration of ON101 twice daily in single and multiple doses yielded very limited systemic exposure (Kai-Min Chu, MD, PhD, oral communication, September 4, 2017). Thus, the maximum body concentrations from days 1 and 14 were similar, demonstrating that topical ON101 has no obvious accumulation in the body. No treatment-related adverse events were observed. In a clinical research trial conducted in 24 patients with chronic DFUs classified as grade 3 according to the Wagner system, 10 treatment with ON101 for 2 weeks resulted in an approximately 20% reduction in wound size, and no serious adverse events were reported. Of the 21 patients with evaluable data, the mean wound size at baseline was 359 (range, 20-2352) mm 2, decreasing to 293 mm 2 after 2 weeks of ON101 treatment. 10 Another clinical trial was performed with 30 patients with Wagner grade 1 chronic DFUs treated with ON101 for as long as 12 weeks (Yu-Yao Huang, MD, PhD, oral communication, August 22, 2011). The final incidence of healing was 50%. The mean wound area at baseline was 577 (range, 303-1225) mm 2, decreasing to 163 mm 2 after 12 weeks of ON101 treatment. The topical use of ON101 is supported with a safety profile from the manufacturer and has clear therapeutic potential in promoting wound healing based on previous studies. 10 This multicenter, phase 3 randomized clinical trial was designed to evaluate whether ON101 could treat chronic DFUs by comparing it with a standard primary wound care absorbent dressing. Methods We followed adequate and well-controlled studies as categorized by the US Food and Drug Administration 16 to design a randomized, controlled, evaluator-blinded phase 3 trial to evaluate the efficacy of ON101 applied topically twice daily for treating chronic DFUs (the trial protocol is available in Supplement 1 ). This treatment was compared with an absorbent dressing (Hydrofiber; ConvaTec Ltd) as a comparator in the control group for treating chronic DFUs. This multicenter study was performed with institutional review board approval from 21 medical/clinical centers (eTable 1 in Supplement 2 ) with wound care specialty across the US, China, and Taiwan, where these investigational new drug programs were initiated; all patients provided written informed consent at enrollment. The study followed the International Council on Harmonization guideline 17 and the Consolidated Standards of Reporting Trials ( CONSORT ) reporting guideline. From November 23, 2012, to May 11, 2020, we enrolled outpatients with type 1 or 2 diabetes (as defined by World Health Organization criteria) aged 20 to 80 years, with a baseline HbA 1c level of less than 12% measured during screening or within 3 months before randomization (to convert to proportion of total hemoglobin, multiply by 0. 01). The target ulcer classified as grade 1 or 2 based on the Wagner system on the foot (below the ankle) needed to measure from 1 to 25 cm 2 after debridement, without active infection, and present for at least 4 weeks despite receiving standard of care (according to the International Working Group on the Diabetic Foot guidelines 18 ) before randomization. To avoid possible premature discontinuation of the patient treatments during the trial, we excluded patients with an ankle-brachial before randomization; those with necrosis, purulence, or sinus tracts in the target ulcer not removable by debridement during the screening visit; or those with acute Charcot neuroarthropathy as defined by the American Diabetes Association and the American Podiatric Medical Association, which indicates perturbations of bone metabolism. 19 In addition, revascularization procedures aimed at increasing blood flow in the target limb must have been performed at least 4 weeks before randomization. Eligible participants judged by the principal investigators (Y. -Y. H. , N. -C. C. , H. -H. C. , K. -F. H. , K. -Y. T. , H. -L. H. , P. -Y. L. , C. -K. P. , B. S. , C. L. , Y. M. , Y. C. , Y. L. , Y. X. , Q. L. , G. N. , and S. -C. C. ) on completion of the screening period (≤7 days) were assigned to receive ON101 or absorbent dressing for as long as 16 weeks in a 1:1 allocation by a computer-generated block randomization scheme (eMethods 1 in Supplement 2 ). 20 Individual investigators and research staff were blinded to the size of the block and remained blinded to the treatment assignment before randomization, eliminating the possibility of predetermining the prospective participant’s treatment assignment. The investigator was informed of the randomized treatment assignment in a sealed envelope containing the individual treatment code at the baseline visit. The end-of-treatment visit (visit 10) was the visit in the 16th week after randomization or the visit in which complete wound closure was confirmed, whichever happened first. The independent evaluator assessed the degree of wound closure. The independent evaluator and the study statistician were blinded to the participants’ treatment throughout the study until the clinical database had been locked. To ensure masking throughout the trial, a standardized procedure was established including camera settings, photographing and image-encoding, image delivery to the independent evaluator, and outcome assessment based on the digitally encoded images to delink the patients’ identification, treatment groups, visits, or site information. The detailed blinding procedure is described in eMethods 2 in Supplement 2. Interventions Demographic data, medical history, disease status, radiography, and eligibility were evaluated during the screening period (before randomization). Participants were scheduled for return visits every 2 weeks to receive wound cleansing and debridement with an assessment of wound status, wound size measurement, physical examination results, and concomitant medication records throughout the 16 weeks of the study period once the interventions were administered. The principal investigators and nurses were trained to use standardized study materials, ON101 or absorbent dressings, camera setting, and off-loading recommendations. The instruction for use of off-loading devices was given to the patients with plantar ulcers as assessed by the clinical investigators. All adverse events were recorded at every visit once the intervention was applied. Blood samples for laboratory tests (including hematologic and biochemical analysis) were collected at the screening visit, then every 4 weeks during the treatment period and at the last visit of the follow-up period to detect the levels of factors such as alanine aminotransferase and aspartate aminotransferase to measure liver status, creatinine and blood urea nitrogen to measure kidney status, and albumin to measure nutritional status. Levels of HbA 1c and blood glucose were measured to monitor diabetes-related safety concerns. ON101, a topical cream composed of PA-F4 and S1, was supplied by Oneness Biotech Co, Ltd, and manufactured in Taiwan in a facility in compliance with Good Manufacturing Practice certified by the Pharmaceutical Inspection Cooperation Scheme. Participants in the ON101 treatment group were shown how to self-administer the cream twice daily in an amount to cover the target ulcer fully without exceeding 2 mm in thickness at each visit. The absorbent dressing containing sodium carboxymethylcellulose (Aquacel; ConvaTec Ltd) needed to be changed daily or 2 to 3 times weekly subject to exudate level following the product’s instructions or the investigators’ discretion. The only secondary dressing allowed was sterile gauze for both groups. The amount of ON101 used or the frequency of absorbent dressing changes for each patient was recorded at every visit during the treatment period. No systemic prescriptions were contraindicated during the treatment period, whereas topical antimicrobials and antiseptic agents were not allowed. In cases where the target ulcer worsened (defined as Wagner grade 3), the investigators could determine whether to terminate treatment. If the ulcer was judged by the blinded evaluator as having undergone complete epithelialization for 2 consecutive visits during the treatment period (at or before visit 10), the intervention (ON101 or absorbent dressing) was stopped, and a visit 10 was scheduled after this judgment. If the patients were confirmed to have an unhealed target ulcer at visit 10, continual standard of care with the absorbent dressings was provided to them regardless of the allocated group during the 12-week follow-up period. Data Collection and Outcome Measures The primary efficacy outcome was to compare the incidence of complete healing between the 2 groups at the end of the 16-week treatment period. Complete healing, defined as complete epithelialization maintained without drainage or requirement of dressings for at least 2 consecutive visits, was determined by an independent evaluator blinded to the patient’s information and treatment allocation. Secondary ulcer-related outcomes included time to complete ulcer healing (from baseline visit to first 100% re-epithelialization visit), percentage of change in ulcer surface area from baseline (to the latest treatment visit or complete wound closure), percentage of patients with a 50% reduction in ulcer surface area, and incidence of infection of the target ulcer. The exploratory, ulcer-related outcome data included any incidence of ulcer recurrence during the 12-week follow-up period. Target wound size was measured by an investigator using digital planimetry at every visit after any necessary debridement. In addition, efficacy variables were further assessed for subgroups for the incidence of complete healing, characterized according to the prior duration of ulcers recorded at the baseline visit (6 months as a cutoff), 21 ulcer size (5 cm 2 as a cutoff), 22 and HbA 1c level (9% as a cutoff regarded as poor glycemic control according to the definition of the American Diabetes Association). Safety outcomes were used to assess adverse events and clinical laboratory values. Statistical Analysis The sample size was calculated based on the results of ON101 in the previous trial by hypothesizing a 20% superiority in the incidence of wound closure compared with the efficacy of the absorbent dressing (Yu-Yao Huang, MD, PhD, oral communication, August 22, 2011). With a 1:1 randomization ratio in the 2 groups, 236 participants were required to be enrolled to ensure that at least 212 had evaluable data for achieving 80% power with a 2-sided α value of 5% nominal significance. All analyses were performed using SAS software, version 9. 4 (SAS Institute Inc). The intention-to-treat (ITT) principle was applied to the full-analysis set (FAS), which included all randomized patients irrespective of the actual receipt of study intervention and adherence to the protocol or the occurrence of adverse events. The FAS was used to analyze all efficacy and safety data. A modified ITT (mITT) protocol was applied to exclude patients in the FAS with ineligible target ulcers at baseline. The mITT was used for supportive analysis of efficacy data as appropriate. For the primary end point, we used a χ 2 test and a logistic regression model with intervention as a fixed factor, with the baseline ulcer size and Wagner grade adjusted as covariates. The results of the logistic regression model are presented in terms of the odds ratio (OR), with P values and associated 95% CIs. Some outcomes are expressed as the hazard ratio (HR). Exploratory post hoc analyses of pertinent variables, such as ulcer duration, ulcer size, and patients’ HbA 1c levels, were also performed. The time to complete ulcer healing was calculated using the Kaplan-Meier method with a log-rank test. The HRs and 95% CIs were estimated using a Cox proportional hazards regression model. The percentile changes in ulcer surface area and ulcer surface area change from baseline were subjected to regression analysis adjusted by baseline ulcer area and Wagner grades. The incidences of infection of target ulcers and of recurrence were evaluated using the Fisher exact test. The adverse events were regarded as treatment emergent if they occurred after the intervention started. Adverse events, treatment-emergent adverse events, and serious adverse events were summarized by frequency and proportion of total patients by system organ class and by preferred terms. All adverse event–related comparisons between the 2 groups were performed using the Fisher exact test. The clinical laboratory test data were used to tabulate the change in values from baseline and were compared between groups using analysis of covariance. All tests were 2 tailed, and P <. 05 was considered statistically significant. For possible early study termination, an independent data monitoring committee was established to monitor data when the patient numbers reached approximately 50% and 90% of the planned enrollment. The futility or superiority of ON101 cream was assessed by the independent data monitoring committee using the Lan-DeMets alpha-spending approach, in which the boundaries were determined by the type of O’Brien-Fleming spending function. 23 The superiority of ON101 was confirmed by the independent data monitoring committee ( P <. 001, much less than the boundary of 0. 03476) on achieving 90% of the planned enrollment (212 participants with evaluable data) so the interim analysis could proceed. The trial was not terminated despite ON101 achieving superiority in the interim analysis because the 236th patient with evaluable data was already enrolled before this point. Results A total of 236 patients were included in the FAS (175 men [74. 2%]; 61 women [25. 8%]; mean [SD] age, 57. 0 [10. 9] years). The mean (SD) HbA 1c level was 8. 1% (1. 6%) at baseline and did not change significantly at the end of treatment (mean [SD] HbA 1c of 8. 0% [1. 8%] in the ON101 group vs 7. 9% [1. 6%] in the comparator group), and 144 patients (61. 0%) were diagnosed as having had diabetes for more than 10 years. Patients in the FAS were randomly allocated to treatment: 114 (48. 3%) to the comparator group and 122 (51. 7%) to the ON101 group. Sixteen patients (13. 1%) in the ON101 group vs 21 (18. 4%) in the comparator group had an early termination (total of 37) ( Figure 1 ). The instructions for using off-loading devices were given to the patients who were assessed by the clinical investigators. Some patients did not follow the suggestion because of the humidity in Taiwan ( Table 1 ). Among the 236 patients in the FAS, 184 (78. 0%) were classified as having Wagner grade 2 ulcers, 117 (49. 6%) had ulcers in the plantar region, and 64 (27. 1%) had a baseline HbA 1c level of at least 9%. The mean (SD) ulcer size was 4. 8 (4. 4) cm 2, and the mean (SD) prior duration of the target ulcer was 7. 2 (13. 4) months at entry ( Table 1 ). Figure 1. CONSORT Diagram of Study Flow A total of 236 patients were randomized. Absorbent dressing was Hydrofiber (ConvaTec Ltd). To convert glycated hemoglobin (HbA 1c ) to proportion of total hemoglobin, multiply by 0. 01. ABI indicates ankle-brachial index; FAS, full-analysis set; and mITT, modified intention to treat. a Judged by the investigator to be unsuitable for the study for any other reason. Table 1. Baseline Patient Characteristics and Intervention During the Study Characteristic Patient group a ON101 (n = 122) Absorbent dressing (n = 114) All (N = 236) Baseline patient characteristics Age, mean (SD), y 57. 4 (10. 6) 56. 6 (11. 3) 57. 0 (10. 9) Sex Male 93 (76. 2) 82 (71. 9) 175 (74. 2) Female 29 (23. 8) 32 (28. 1) 61 (25. 8) Type 2 diabetes 121 (99. 2) 113 (99. 1) 234 (99. 2) Diabetes duration, y ≤10 55 (45. 1) 37 (32. 5) 92 (39. 0) >10 67 (54. 9) 77 (67. 5) 144 (61. 0) HbA 1c level, % Mean (SD) 8. 1 (1. 5) 8. 1 (1. 8) 8. 1 (1. 6) <9 90 (73. 8) 82 (71. 9) 172 (72. 9) ≥9 32 (26. 2) 32 (28. 1) 64 (27. 1) BMI <25 59 (48. 4) 50 (43. 9) 109 (46. 2) ≥25 63 (51. 6) 64 (56. 1) 127 (53. 8) Hypertension 78 (63. 9) 73 (64. 0) 151 (64. 0) CVD history b 25 (20. 5) 23 (20. 2) 48 (20. 3) Kidney status eGFR, mL/min/1. 73 m 2 ≥60 90 (73. 8) 81 (71. 1) 171 (72. 5) <60 32 (26. 2) 33 (28. 9) 65 (27. 5) ABI, mean (SD) 1. 1 (0. 2) 1. 1 (0. 1) 1. 11 (0. 1) Amputation history c 56 (45. 9) 60 (52. 6) 116 (49. 2) Wound conditions, Wagner grade 1 29 (23. 8) 23 (20. 2) 52 (22. 0) 2 93 (76. 2) 91 (79. 8) 184 (78. 0) Ulcer size, cm 2 Mean (SD) 5. 0 (4. 4) 5. 1 (4. 7) 4. 8 (4. 4) 1-5 88 (72. 1) 77 (67. 5) 165 (69. 9) >5 33 (27. 0) 36 (31. 6) 69 (29. 2) Ulcer duration, mo Mean (SD) 7. 2 (13. 0) 7. 3 (13. 9) 7. 15 (13. 4) <6 86 (70. 5) 79 (69. 3) 165 (69. 9) ≥6 36 (29. 5) 35 (30. 7) 71 (30. 1) Plantar ulcers 64 (52. 5) 53 (46. 5) 117 (49. 6) Intervention during the study Off-loading in plantar ulcer d Use 33 (51. 6) 34 (64. 2) 67 (57. 3) No use 15 (23. 4) 9 (17. 0) 24 (20. 5) Not specified 16 (25. 0) 10 (18. 9) 26 (22. 2) Diabetes medication prescribed Metformin 62 (50. 8) 51 (44. 7) 113 (47. 9) Insulin 67 (54. 9) 67 (58. 8) 134 (56. 8) Any oral hypoglycemic agent 84 (68. 9) 81 (71. 1) 165 (69. 9) Use of antibiotics 30 (24. 6) 26 (22. 8) 56 (23. 7) Abbreviations: ABI, ankle-brachial index; BMI, body mass index (calculated as weight in kilograms divided by height in meters squared); CVD, cardiovascular disease; eGFR, estimated glomerular filtration rate; HbA 1c, glycated hemoglobin. SI conversion factor: To convert HbA 1c to proportion of total hemoglobin, multiply by 0. 01. a Unless otherwise indicated, data are expressed as number (%) of patients. Owing to missing data, numbers may not total column headings or percentages may not total 100. Absorbent dressing was Hydrofiber (ConvaTec Ltd). b Includes ischemic heart disease, coronary artery disease, or cerebral vascular accident with embolic, ischemic, or hemorrhagic stroke. c Due to previous diabetic foot ulcers. d Includes only patients with plantar ulcer. Primary Outcome Seventy-four patients (60. 7%) in the ON101 group vs 40 (35. 1%) in the comparator group achieved ulcer closure within 16 weeks (OR, 2. 84; 95% CI, 1. 66-4. 84; P <. 001) ( Table 2 ). Similar results were also noted in the mITT population, where 73 of 118 patients (61. 9%) in the ON101 group and 38 of 112 (33. 9%) in the comparator group had ulcer closure (OR, 3. 15; 95% CI, 1. 82-5. 43; P <. 001) ( Table 2 and eTable 2 in Supplement 2 ). The independent evaluator assessed the degree of wound closure. Table 2. Primary and Secondary Outcomes a Outcome Patient group OR (95% CI) P value ON101 (n = 122) Absorbent dressing (n = 114) Complete healing, No. (%) FAS 74 (60. 7) 40 (35. 1) 2. 84 (1. 66-4. 84) <. 001 b mITT 73 (61. 9) 38 (33. 9) 3, 15 (1. 82-5. 43) <. 001 b Secondary Change in WSA from baseline to visit 10, mean (SD), % −78. 0 (42. 6) −78. 0 (34. 9) NA. 89 Incidence of patients with 50% reduction in WSA on visit 10, No. (%) 101 (82. 8) 98 (86. 0) 0. 80 (0. 39-1. 62). 53 b Incidence of wound infection 6 (4. 9) 7 (6. 1) NA. 78 Ulcer recurrence, No. (%) c 15 (20. 3) 7 (17. 5) NA. 81 Safety Patients with TEAEs, No. (%) 76 (62. 3) 77 (67. 5) NA. 42 No. of TEAEs 207 235 NA Related TEAEs Patients, No. (%) 7 (5. 7) 5 (4. 4) NA. 77 No. of events 11 5 NA NA Serious TEAEs Patients, No. (%) 14 (11. 5) 9 (7. 9) NA. 39 No. of events 24 14 NA NA Related serious TEAEs in events, No. (%) 0 1 (0. 9) NA <. 48 TEAE leading to death, No. 0 0 NA NA Subgroup analysis Wound closure, No. /total No. (%) HbA 1c level <9% 59/90 (65. 6) 33/82 (40. 2) 2. 81 (1. 50-5. 26) <. 001 b ≥9% 15/32 (46. 9) 7/32 (21. 9) 3. 14 (1. 04-9. 50). 04 b Ulcer size, cm 2 1-5 55/88 (62. 5) 31/77 (40. 3) 2. 46 (1. 31-4. 61). 005 b >5 18/33 (54. 5) 8/36 (22. 2) 4. 09 (1. 42-11. 80). 009 b Ulcer duration, mo <6 mo 62/86 (72. 1) 36/79 (45. 6) 3. 07 (1. 59-5. 95) <. 001 b ≥6 mo 12/36 (33. 3) 4/35 (11. 4) 3. 99 (1. 09-14. 63). 04 b Abbreviations: FAS, full-analysis set; HbA 1c, glycated hemoglobin; mITT, modified intent-to-treat; NA, not applicable; OR, odds ratio; TEAEs, treatment-emergent adverse events; WSA, wound (ulcer) surface area. a The absorbent dressing used was Hydrofiber (ConvaTec Ltd). b Calculated using a logistic regression model. Treatment was the main exposure variable; the baseline wound size in cm 2 and Wagner grade were covariates. c Ulcer recurrence was recorded once the ulcer had healed completely and was observed during the follow-up period. Ulcer duration, ulcer size, and HbA 1c levels are known to be associated with poor prognosis of DFUs. 9, 24, 25 Therefore, a subgroup analysis was conducted on baseline ulcer duration (6 months as a cutoff), baseline ulcer area (5 cm 2 as a cutoff size), and baseline HbA 1c level (9% as a cutoff). The subgroup analysis displayed a significant OR in favor of the ON101 group compared with the comparator group (OR, 3. 14 [95% CI, 1. 04-9. 50; P =. 04] for HbA 1c level ≥9%; OR, 3. 99 [95% CI, 1. 09-14. 63; P =. 04] for ulcer duration ≥6 months; OR, 4. 09 [95% CI, 1. 42-11. 80; P =. 009] for ulcer size >5 cm 2 ) ( Table 2 ). In addition, we subgrouped patients with an ulcer reduction of less than 10% during the screening period and analyzed the primary efficacy variable. The result also favored the ON101 treatment (32 of 64 [50. 0%] vs 18 of 66 [27. 3%]; P =. 02) (eTable 6 in Supplement 2 ). Secondary Outcome Patients in the ON101 group had a better healing rate than those in the comparator group (HR, 1. 80 [95% CI, 1. 23-2. 65; P =. 002]) ( Figure 2 ) in the FAS as well as the mITT population (HR, 1. 91 [95% CI, 1. 29-2. 83; P =. 001]) (eFigure 2 in Supplement 2 ). The cumulative incidence of complete healing at each week also reflected the continual higher probability in the ON101 group for reaching complete wound closure from week 4 onward. The time to reach median population healing was 98 days in the ON101 group, whereas it was undeterminable in the comparator group because ulcers of only 40 patients (35. 1%) healed in this group during the treatment period ( Figure 2 ). The mean reduction in ulcer surface area (from the last treatment visit to baseline) was 78. 0% in both groups (SDs, 42. 6% for the ON101 group and 34. 9% for the comparator group; P =. 89), and the incidence of a 50% reduction in ulcer surface area was not significantly different between both groups (101 of 122 [82. 8%] vs 98 of 114 [86. 0%]) ( Table 2 ). Only a few episodes of target ulcer infection occurred in both groups during the treatment period (6 in the ON101 group and 7 in the comparator group; P =. 78) ( Table 2 ). The incidence of recurrence in completely healed wounds during the follow-up phase was 15 of 74 (20. 3%) in the ON101 group and 7 of 40 (17. 5%) in the comparator group without statistical significance ( P =. 81) ( Table 2 ). Figure 2. Kaplan-Meier Plot of Time to Complete Healing in the Full-Analysis Set Population The survival curve indicates the incidence of ulcers healed at each visit in the full-analysis set population. Complete healing was defined as epithelialization without drainage observed at 2 consecutive visits. A full-analysis set cohort randomly assigned to the absorbent dressing (Hydrofiber; ConvaTec Ltd) group (n = 114) or ON101 group (n = 122) was used for Kaplan-Meier analysis. Adverse Events In terms of safety, there were no clinically significant changes or differences between the 2 treatment groups in hematology, biochemistry (including HbA 1c and fasting glucose levels), or vital signs ( Table 2 and eTable 3 in Supplement 2 ). Treatment-emergent adverse events were reported in 76 patients in the ON101 group and 77 in the comparator group, of whom 7 of 122 (5. 7%) in the ON101 group and 5 of 114 (4. 4%) from the comparator group were considered related to the treatments ( Table 2 and eTable 4 in Supplement 2 ). None of the serious adverse events was related to ON101 treatment, whereas there was 1 case of osteomyelitis reported to be linked to the comparator group in which 1 patient (0. 8%) assigned to ON101 died of septic shock, acute kidney injury, and acute respiratory failure, which were not considered to be related to treatment or to ulcer progression (eTable 5 in Supplement 2 ). Discussion To our knowledge, this study is the first international phase 3 randomized clinical trial of an investigational drug able to regulate M1/M2 macrophage activities in patients with DFUs. ON101 exhibited better efficacy in facilitating the complete healing of DFUs. Hyperglycemia is an underlying cause of chronic DFUs in which the M1-to-M2 macrophage transition is delayed and the inflammatory stage is prolonged. 26, 27 ON101 can restore the balance of M1/M2 macrophages caused by hyperglycemia. The robust efficacy in patients with high-risk factors suggests that ON101 might provide multiple and proactive ways to improve wound healing by promoting the M1-to-M2 transition and thereby accelerating wound healing for ulcers not only in terms of early formation but also with high-risk factors including ulcer duration of at least 6 months, ulcer size greater than 5 cm 2, and an HbA 1c level of at least 9%. The design of this study followed US Food and Drug Administration guidelines. 16 The complete healing rate of the comparator group at 16 weeks (35. 1%) was consistent with the 28. 2% shown by ITT analysis at week 12 disclosed in a previous trial by Jeffcoate et al. 28 This finding verifies the suitability of the design and implementation of this study in conforming to other randomized clinical trials. The application of ON101 after debridement—which can be self-administered at home—indicated the same level of convenience of use as for the absorbent dressing. Despite the statistically significant wound closure and healing rates provided by ON101, the ulcer reduction outcomes, including changes in ulcer area from baseline and rate of 50% reduction in the wound area, were not statistically significant between the 2 groups during the treatment period. This discrepancy possibly arose from the use of 2-dimensional measurements on the wound area without considering the wound depth. In this study, 78. 0% of the ulcers were Wagner grade 2, meaning that they extended into tendon, bone, or capsule. Thus, the measurement of wound area instead of volume might not reflect the actual volumetric change. Similar outcomes were also noted in the pivotal study (study 92-22120-K) of becaplermin (Regranex; Smith & Nephew plc). The use of 3-dimensional measurement tools should be considered in future studies. Limitations This study has some limitations. The first was the open-label design, which did not allow us to mask the interventions to patients or clinical investigators; therefore, blinded evaluation was implemented to minimize any possible bias. Second, the inclusion and exclusion criteria ruled out patients requiring dialysis, which, to a certain extent, reflects some types of patients with DFUs. Using the ankle-brachial index as the sole criterion in judging blood perfusion could not exclude patients with ischemia completely. Last, the lack of a 2-week run-in period was a potential flaw in the design, because possible rapid healers might not have been excluded in the study. To assess whether this factor affected the trial results, a separate analysis of the complete ulcer healing rate was performed by excluding those patients with an ulcer reduction of at least 10% during the screening period, the results of which favored ON101 treatment (32 of 64 [50. 0%] vs 18 of 66 [27. 3%]; P =. 02) (eTable 6 in Supplement 2 ). Conclusions The results of this randomized clinical trial demonstrate a clinically and statistically superior therapeutic efficacy of ON101 in the treatment of DFUs in both FAS and mITT populations in terms of complete healing rate and time to complete healing compared with absorbent dressing. For chronic wounds in patients with high-risk factors, the therapeutic efficacy of ON101 could be sustained in ulcers that last for more than 6 months or measure greater than 5 cm 2 or in patients with high HbA 1c levels. The findings of this study suggest that ON101, a macrophage regulator that behaves differently from moisture-retaining dressings, represents an active-healing alternative for home and primary care of patients with chronic DFUs.
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10. 1001/jamaophthalmol. 2013. 4319
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JAMA ophthalmology
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Accelerated
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Purpose To design patterned, transparent silk films with fast degradation rates for the purpose of tissue engineering corneal stroma, Methods β-sheet (crystalline) content of silk films was decreased significantly by using a short water annealing time. Additionally, a protocol combining short water annealing time with enzymatic pretreatment of silk films with protease XIV was developed. Results Low β-sheet content (17–18%) and enzymatic pre-treatment provided film stability in aqueous environments and accelerated degradation of the silk films in the presence of human corneal fibroblasts in vitro. The results demonstrate a direct relationship between reduced β-sheet content and enzymatic pre-treatment and overall degradation rate of the protein films. Conclusions The novel protocol developed here provides new approaches to modulate the regeneration rate of silk biomaterials for corneal tissue regeneration needs. Translational relevance Patterned silk protein films possess desirable characteristics for corneal tissue engineering, including optical transparency, biocompatibility, cell alignment and tunable mechanical properties, but current fabrication protocols do not provide adequate degradation rates to match the regeneration properties of the human cornea. This novel processing protocol makes silk films more suitable for the construction of human corneal stroma tissue and a promising way to tune silk film degradation properties to match corneal tissue regeneration.
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No full text available
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10. 1001/jamaoto. 2013. 5669
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JAMA otolaryngology-- head & neck surgery
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Defining the critical-sized defect in a rat segmental mandibulectomy model
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Importance Advances in tissue engineering offer potential alternatives to current mandibular reconstructive techniques; however, prior to clinical translation of this technology, a relevant animal model must be used to validate possible interventions. Objective This study aims to establish the critical-sized segmental mandibular defect that does not heal spontaneously in the rat mandible. Design Prospective study using an animal model. Setting Animal laboratory. Participants Sprague-Dawley rats. Interventions Twenty-nine Sprague-Dawley rats underwent creation of one of four segmental mandibular defects: 0-mm, 1-mm, 3-mm and 5-mm. All mandibular wounds were internally fixated with 1-mm microplates and screws and allowed to heal for 12-weeks. Main Outcomes and Measures Mandibles were analyzed with micro-computed tomography (microCT) and bony healing was graded on a semi-quantitative scale. Results Seven animals were utilized in each experimental group. No 5-mm segmental defects successfully developed bony union, whereas all 0-mm and 1-mm defects had continuous bony growth across the original defect on micro-CT. Three of the 3-mm defects had bony continuity, and three had no healing of the bony wound. Bony union scores were significantly lower in the 5-mm defects compared to 0-mm, 1-mm and 3-mm defects (all p < 0. 01). Conclusion and Relevance The rat segmental mandible model cannot heal a 5-mm segmental mandibular defect. Successful healing of 0-, 1- and 3-mm defects confirms adequate stabilization of bony wounds with internal fixation with 1-mm microplates. The rat segmental mandibular critical-sized defect provides a clinically relevant testing ground for translatable mandibular tissue engineering efforts.
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No full text available
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10. 1002/0471143030. cb2309s61
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Current protocols in cell biology / editorial board, Juan S. Bonifacino. . . [et al. ]
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Expanding Mouse Ventricular Cardiomyocytes through GSK-3 Inhibition
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Controlled proliferation of cardiac myocytes remains a major limitation in cell biology and one of the main underlying hurdles for true modern regenerative medicine. Here we provide a technique to robustly expand early fetal-derived mouse ventricular cardiomyocytes on a platform usable for high-throughput molecular screening, tissue engineering or potentially useful for in vivo translational experiments. This method provides a small molecule-based approach to control proliferation or differentiation of early beating cardiac myocytes through modulation of the Wnt/β-catenin signaling pathway. Moreover isolation and expansion of fetal cardiomyocytes takes less than 3 weeks, yields a relatively pure (~70%) functional myogenic population and is highly reproducible.
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No full text available
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10. 1002/1873-3468. 12285
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Febs Letters
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Inherited heart disease – what can we expect from the second decade of human i
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Induced pluripotent stem cells (i PSC s) were first generated 10 years ago. Their ability to differentiate into any somatic cell type of the body including cardiomyocytes has already made them a valuable resource for modelling cardiac disease and drug screening. Initially human i PSC s were used mostly to model known disease phenotypes; more recently, and despite a number of recognised shortcomings, they have proven valuable in providing fundamental insights into the mechanisms of inherited heart disease with unknown genetic cause using surprisingly small cohorts. In this review, we summarise the progress made with human i PSC s as cardiac disease models with special focus on the latest mechanistic insights and related challenges. Furthermore, we suggest emerging solutions that will likely move the field forward.
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Abbreviations ALPK3, alpha kinase 3 ARVC, arrhythmogenic right ventricular cardiomyopathy CMs, cardiomyocytes DCM, dilated cardiomyopathy HCM, familial hypertrophic cardiomyopathy hiPSC, human iPSC hPSC, human pluripotent stem cell iPSCs, induced pluripotent stem cells LQTS, long‐QT syndrome RBM20, RNA‐binding motif protein 20 Since their discovery in 2006 1, induced pluripotent stem cells (iPSCs) have enabled scientists to study the physiological and pathological mechanisms of both development and disease in a new way 2, 3, 4. Inherited cardiovascular disorders and in particular channelopathies have been among the first human diseases studied using iPSCs 5, 6, 7, 8. Indeed, although animal models have been and continue to be essential in advancing the understanding of cardiovascular disease 9, 10, 11, interspecies differences hamper translation of many results directly to humans. Because human iPSCs (hiPSCs) can be derived from virtually any patient of interest and can usually differentiate efficiently into cardiomyocytes (hiPSC‐CMs) and there are well‐established techniques for their functional characterisation in vitro, they have rapidly been exploited for disease modelling and drug screening, and in the future are expected to offer new opportunities for regenerative medicine and personalised medicine. The ambition of the Precision Medicine Initiative is to employ a combination of clinical, genetic or genomic, and molecular data to develop tailored therapies for subgroups of patients 12. The pathogenesis of many inherited cardiac diseases remains insufficiently understood and it has been difficult to account for incomplete penetrance and variable severity. Any methodology that could help deciphering predisposition or causative molecular and cellular mechanisms therefore could be helpful. Because hiPSCs can capture the complex genetic background of a patient, expectations are that they will contribute to this goal. Whether patient‐specific hiPSC‐CMs can provide information that predicts disease penetrance and outcome remains to be determined. However, recent studies using hiPSC‐CMs have lead to optimism with respect to the use of this technology in providing new mechanistic insights into disease pathogenesis 13 and cardiotoxicity 14, 15. In this review we discuss the use of hiPSCs as models of inherited heart diseases, with special focus on underlying disease mechanisms that have not been evident from other approaches, current challenges, and emerging solutions for moving the field forward. Generation of hiPSC‐CMs (for cardiac disease modelling) In using hiPSCs for disease modelling, the first step is reprogramming somatic cells collected from primary tissue samples. Most frequently, hiPSCs are derived from patients with known disease‐causing mutations. An alternative, but increasingly used, approach is to introduce site‐specific genetic changes (including knock‐out and precise nucleotide changes) in wild‐type human pluripotent stem cell (hPSC) lines by gene targeting 4. Even in cases where no causative mutations are yet known, the rationale of using hiPSCs is that they are able to capture any genetic predisposition in patients aside from specific mutations, which are thought to play determinant roles not only in disease manifestation and progression but also in the context of external environmental factors that may precipitate the condition. These could include exercise, cardiac‐ and noncardiac drugs, fever, food supplements and the like 15, 16. Once hiPSCs have been obtained, several methods to induce cardiac differentiation can be used 17, 18. Although these techniques were initially inefficient and not readily transferable across cell lines, there are now a number of more robust protocols available and CMs at > 95% purity can be produced 19, 20. In addition, a number of defined media and commercial kits have become available of late which seem particularly efficient across lines, including several apparently differentiation refractory hiPSC lines. However, it is noteworthy that although the efficiency of differentiation protocols has undergone a multifold increase over recent years as a result of culture condition optimisation, this has not been paralleled by improvements in maturation of the electrophysiological properties of hiPSC‐CMs: resting membrane potential is depolarised, and upstroke velocity and ion channels expression remain low in comparison with adult cardiomyocytes 21, 22, 23. This suggests that optimisation has impacted quantitative rather than qualitative aspects of differentiation. Of note, most of these differentiation protocols result in mixed populations of ventricular‐, atrial‐ and nodal‐like subtypes, with ventricular CMs being the most represented. Some recent studies have succeeded in directing hPSC differentiation towards atrial 24, 25 and pacemaker 26 subtypes, however, their application for studying molecular mechanisms related to disease is still under investigation. Maturation of hiPSC‐CMs Improving maturity in hiPSC‐CMs remains one of the major priorities of the field, since phenotypic immaturity limits their ability to successfully model critical aspects of cardiac disorders including adult‐onset diseases 21. Comparison with human fetal hearts suggests that in vitro ‐derived hPSC‐CMs are similar to first trimester gestational stage CMs with regard to gene expression, structure and function and only in certain culture conditions do they become more similar to second trimester fetal CMs 27, 28. Channelopathies are among the cardiac diseases that suffer least from these limitations, since most (but not all) of the relevant ion channels for the generation of the cardiac action potential are expressed in hiPSC‐CMs. This is the reason why the long‐QT syndrome (LQTS) was one of the first cardiac arrhythmia conditions to be modelled using hiPSC‐CMs 5. Since then, approximately one‐fourth of the publications on cardiac disease modelling have studied LQTS‐causing mutations. Although some key features of other inherited heart diseases, such as catecholaminergic polymorphic ventricular tachycardia (CPVT) 29, arrhythmogenic right ventricular cardiomyopathy (ARVC) 30, familial hypertrophic cardiomyopathy (HCM) 31 and familial dilated cardiomyopathy (DCM) 32 have also been recapitulated, certain molecular mechanisms will only be reproduced when more mature cardiac phenotypes are achieved. Existing hiPSC models of inherited cardiac diseases Human iPSC technology has succeeded in modelling cardiovascular and cardiometabolic diseases with different inheritance patterns: the most common autosomal dominant forms (LQTS 5, 6, 7, 33, CPVT1 29, DCM 32, HCM 31, ARVC 34 ) but also the rarer autosomal recessive forms (CPVT2 35, Jervell and Lange‐Nielsen syndrome (JLNS) 36, Pompe disease 37 ), the X‐linked dominant forms (Danon disease 38, Fabry disease 39 ), the X‐linked recessive forms (Barth syndrome 40, Duchenne muscular dystrophy or DMD 41 ) and also finally the nontypical Mendelian forms (hypoplastic left heart syndrome or HLHS 42 ). All of these examples result from genetic defects with cell‐autonomous mechanisms of action in CMs, which means that the pathological phenotype is evident in the cardiomyoctes expressing the (mutated) gene without the need to interact with other cell types (Fig. 1 ). This may not always be the case and there is an increasing number of examples in which the interaction between two or more cell types is needed to reveal the disease phenotype since the cell expressing the mutation sends defective signals to its neighbours 43. As hiPSC technology advances, the ability to establish heterotypic cultures and complex structures increases, we expect that noncell‐autonomous disease mechanisms will be also recapitulated such as those leading to heart failure due to vascular diseases (thrombosis, atherosclerosis) and myocardial infarction (Fig. 1 ). Figure 1 Cell‐autonomous versus noncell‐aututonomous diseases. hi PSC ‐ CM s have already proven their value in recapitulating cell‐autonomous cardiovascular diseases, such as arrhythmic syndromes ( LQTS, JLNS, CPVT ), cardiomyopathies ( DCM, HCM, ARVC, DMD ), cardiometabolic disorders (Pompe disease, Fabry disease, Danon disease, Barth syndrome). More challenging to be modelled are noncell‐autonomous cardiovascular disorders, such as diabetic cardiomyopathy, heart failure due to vasculature diseases, for example, thrombosis, atherosclerosis or myocardial infarction. Drug screening, toxicology assays and safety pharmacology One of the fundamental applications of hiPSC cardiac disease models is the development of treatments that ideally will eventually be translated into the clinic to cure (reverse) or relieve (delay) disease symptoms, much like that already achieved for some neurodegenerative disorders 44. This approach is highly dependent on understanding the molecular mechanisms underlying the disease, as well as on the sensitivity of the read‐out in the assay that is used for detecting the abnormal phenotype. Testing a limited number of candidate drugs based on underlying disease mechanisms is already proving the fastest way to move forward to clinical application, since it is based on repurposing previously approved compounds for a new disease 45. In the cardiac field this has not yet led to rapid translation from the laboratory bench to patients, partly because cardiovascular diseases are often not as severe and untreatable as many neurodegenerative disorders. As an alternative to repurposing, hiPSC‐CMs can be used as a platform for high throughput drug testing 46, which is most valuable to pharmaceutical companies looking for new drug and disease targets since they often have technologies for automated measurements. In addition to drug screening and drug development, hiPSC‐CMs are now also beginning to demonstrate their value in revealing cardiotoxic effects. In particular, these cells are proving a valuable tool to identify electrophysiological and transcriptional changes related to HDAC inhibitor‐mediated cardiotoxicity 47. Furthermore, Burridge and colleagues have recently shown that patient‐specific hiPSC‐CMs can recapitulate the predisposition of some breast cancer patients to develop late heart failure after exposure to the chemotherapeutic drug doxorubicin 15. Although of significant interest, this study had some limitations: first, relatively few patients were included in each group (four in the in vitro doxorubicin cardiotoxicity assays and only three for the RNA‐seq analysis of the hiPSC‐CMs); second, the retrospective study design and coadministration of additional chemotherapeutic drugs in one patient group might have biased the outcome. Further validation in larger patient cohorts will be needed to determine whether different degrees of severity and early versus late cardiotoxic effects can be detected and whether the same approach proves valid for other patient groups such as those with tumours in other organs or paediatric patients also treated with doxorubicin 48, 49. Nevertheless, the work supports the idea that hiPSCs are able to capture complex genetic backgrounds of patients in a predictive way and therefore might contribute usefully to the realisation of the Precision Medicine Initiative 12. Pathological phenotypes and new mechanistic insights The successful generation of cardiac disease models with hiPSC‐CMs relies on their ability to recapitulate key aspects of CM biology, including their molecular, cellular and physiological properties, and on the scientist’ tools and ability to record and capture these specific features and changes upon pathological or cardiotoxic conditions. For this, appropriate and sensitive read‐out assays have been developed 13 and techniques are being continuously improved 22. Disease‐related phenotypes and read‐out assays During the first years that followed hiPSC discovery, their derivative CMs were used mostly to model known disease phenotypes to explore their potential value in recapitulating maladies and known pharmacological treatments 8. More recently hiPSC‐CMs proved helpful in providing novel mechanistic insights into inherited heart diseases with both known and unknown genetic cause using surprisingly small cohorts. The different assays used to characterise hiPSC‐CMs phenotypes examine parameters such as gene and protein expression, ultrastructural organisation, electrophysiological function, calcium handling, force of contraction and metabolic profile. Here, we discuss some of the latest examples. The various kinds of diseases that have been modelled using hiPSC are illustrated in Fig. 1. Analysis of not only gene expression in hiPSC‐CMs but also the changes that take place during hiPSC cardiac differentiation has offered hints on genes and potential pathways impaired in some inherited heart conditions. In an autosomal dominant form of DCM caused by mutations in the RNA‐binding motif protein 20 gene ( RBM20 ), for example, stage‐specific transcriptome profiling demonstrated early molecular perturbations during cardiogenesis in patient‐specific hiPSCs 50 ; these results suggested that this clinically aggressive form of DCM is a developmental disorder. In addition, using functional assays the authors demonstrated that RBM20‐dependent mis‐splicing of calcium‐handling genes contributed to alterations in the calcium homoeostasis and excitation–contraction coupling. Similarly, whole transcriptome sequencing led to the hypothesis that mitochondria were implicated in DMD cardiac pathogenesis 41 ; subsequent analysis of the metabolic profile demonstrated that indeed apoptosis in DMD hiPSC‐CMs is mainly induced by a mitochondrial network through the proteins DIABLO, XIAP and CASP3 rather than through cytochrome C and CASP9 cascade. The analysis of ultrastructural CM organisation has revealed phenotypes not only in several cardiomyopathies but also in glycogen storage diseases 51. Among these, Pompe disease was one of the first disorders characterised in depth using hiPSC‐CMs, in which both specific features of the cardiomyopathy and the efficacy of recombinant enzyme therapy in patients were faithfully recapitulated 37. However, only more recently did Raval and colleagues discover a specific deficit in the glycan synthesis in the Golgi that was initially revealed by the change in electrophoretic mobility of lysosomal‐associated membrane protein LAMP1 52. Likewise, ultrastructural analysis by electron microscopy revealed the presence of fragmented mitochondria within autophagosomes in hiPSC‐CMs carrying Danon disease, which is caused by lysosomal‐associated membrane protein LAMP2 deficiency 38 ; a consequent increase in oxidative stress and apoptosis was then demonstrated. This study was one of the first attempts to understand the molecular basis of some pathological features that are also characteristic of heart failure. Interestingly, genetic polymorphisms in the cardioprotective enzyme, aldehyde dehydrogenase 2 gene ( ALDH2 ), were studied by examining the metabolic profile of patient‐specific hiPSC‐CMs 53. A new function for this enzyme was demonstrated, namely modulation of cell survival decisions through changes in the oxidative stress in hiPSC‐CMs, although these finding were first observed in patient‐fibroblasts and then later examined in hiPSC‐CMs. Gene expression analysis highlighted overexpression of JUN and consequently the authors were able to restore ROS levels by the Jun N‐terminal kinase (JNK) inhibition. Importantly, no significant differences were identified under normoxic condition, while ischaemia simulation in vitro revealed the phenotype. Calcium influx into the cell triggers further calcium release from the sarcoplasmic reticulum to the cytosol and finally to the sarcomere resulting in cardiomyocyte contraction. The identification of calcium handling abnormalities in hPSC‐CMs harbouring alpha kinase 3 gene ( ALPK3 ) mutations allowed confirmation at the cellular and molecular level of strong genetic evidence that homozygous or bi‐allelic truncating mutations in ALPK3 can cause paediatric cardiomyopathy 54. However, in this casus, the specific role of ALPK3 remained unclear. Channelopathies are most often characterised by examining their electrophysiological and ion channel properties. Molecular profiling coupled with measurements of action potentials and the slow component of the delayed rectifier potassium current ( I Ks ) demonstrated a distinct molecular mechanism of action of two KCNQ1 mutations in JLNS hiPSC‐CMs 36. A recessive phenotype was associated with the amorphic mutation, while a gene dosage‐dependent ion channel protein reduction at the cell membrane explained the presence of a LQTS phenotype in the heterozygously mutated hiPSC‐CMs. Of note, however, the literature reports a wide range of values in the basic electrophysiological properties of hiPSC‐CMs, action potentials differing an order of magnitude and beating rates anywhere between 0. 5 and 1. 5 Hz, but also specific ion currents (e. g. the slow I Ks, and the rapid I Kr components of the delayed rectifier potassium currents, the sodium current I Na, and the L‐type calcium current I CaL ) varying widely even among wild‐type control hiPSC‐CMs 55, 56. Most notably, very different levels of I Ks have been described (ranging from ~ 0. 3 to ~ 2. 5 pA/pF 5, 57 ), variable observation leading to controversial conclusions: on the one hand, I Ks recapitulates physiological behaviour in playing a major role when repolarisation reserve is attenuated 58, 59 ; on the other, it seems to contribute to repolarisation in hiPSC‐CMs even in the absence of sympathetic stimulation 5, 36, 60, 61. The immature phenotype of all stem cell derivatives including hiPSC‐CMs is probably the reason for this variability but, independent of the cause, it is a limitation to extrapolating results obtained using hiPSC‐CMs to native – healthy and diseased – adult human CMs as discussed below. The variability in protocols used for cardiac differentiation and electrophysiology further contribute to making absolute conclusions on human cardiac physiology and disease. Nevertheless, hiPSC‐CMs with ion channel mutations have been able to contribute to understanding these diseases because in many cases they could recapitulate key disease features observed in patients and sometimes indicate underlying pathological molecular mechanism 56, 62. One example in which quantifiable hiPSC‐CM properties were used for drug screening purposes is diabetic cardiomyopathy 63, a complex metabolic condition affecting also the heart. Here the authors built two levels of disease models with hiPSC‐CMs: environmental, by modulating culture conditions to mimic the diabetes chemistry, and genetic, by deriving hiPSC‐CMs from two patients with different disease severities. Interestingly, in the patient‐specific cells, the diabetes phenotype appeared even in the absence of any diabetic trigger, suggesting that hiPSCs indeed capture and recapitulate genetic predisposition. A common limitation of these studies is the small number of patients analysed; to confirm this concept, it will be necessary to validate results independently across larger cohorts. Choosing the right controls The choice of controls is crucial to allow a proper definition and identification of normal versus abnormal phenotypes, including disease‐ and toxic‐specific molecular mechanisms. Each individual harbours many genetic variants in the genome (not only single nucleotide polymorphisms, copy number variations but also heterozygous and homozygous mutations 64, 65 ) that may be functionally interconnected with the genetic defect underlying a disease. Gene targeting enables isogenic hPSC lines to be created that differ only at specific loci, while the rest of the genome remains identical. The advantage of isogenic lines is that any difference in the phenotype is then most likely linked to genetic change since the only difference between the disease and control line is in principle the mutation of interest. With improvement in the methodologies that can be used for precise gene targeting 66, 67, genetically matched (isogenic) hiPSC lines are now becoming the first choice, although in the cardiac field only a few papers have adopted this approach 36, 40, 54, 68, 69, 70 (Fig. 2 ). Hinson and colleagues demonstrated that truncating mutations in the sarcomeric protein, titin, underlie DCM sarcomeric insufficiency 69 ; interestingly, when isogenic hiPSC‐CMs were used, the reduction in force of contraction was still detectable in the mutated CMs but to a lesser extent than when unrelated diseased and control cells were compared. These results confirmed earlier evidence that genetic background can modify disease phenotype. Similarly, we previously generated two pairs of LQTS and control hiPSCs and hESCs harbouring the same KCNH2 mutation 68 ; comparison of genetically matched CMs proved essential for neither under‐ nor overestimating the consequences of the mutation for the cardiac action potentials. Figure 2 summarises the controls used in all studies since first published in 2010. Of note, relatively few have used isogenic controls. Figure 2 Number of publications about hi PSC s and disease modelling using unrelated controls, family matched controls and isogenic controls from 2010 to mid 2016 in the cardiac field. PubMed Advanced Search Builder was used for the literature search using the following builder: [(human pluripotent stem cell) AND (cardiac disease model) NOT review]. Publications on heart regeneration were manually excluded. References from some of the most comprehensive reviews of the field 8, 13, 62, 87 were screened and manually added when not present in the above‐mentioned search. All the References were then screened and classified according to the control used. The complete list and analysis of references is provided in Table S1. Limitation of this representation relates to selection bias. Future challenges It is now clear that hiPSC‐CMs are useful for modelling inherited human cardiac diseases since there are many different examples in which these cells manifest pathogenic features of the disease. However, the predictive and instructive power of hiPSC‐CMs relies on comprehensive and accurate molecular and functional characterisation 13. Challenges that scientists are facing are the ability to model complex‐ and noncell‐autonomous disorders, predict clinical drug response, recapitulate maturation and ageing in vitro along with difficulties in actually recognising mature CMs in culture; for these issues, emerging solutions are discussed. Complex and noncell‐autonomous cardiovascular disorders Many cardiac diseases can be modelled using a single cell type, most often CMs. Ventricular CMs have been the cell type of choice for many diseases, although other cardiac subtypes might be necessary for studying different conditions, for example, nodal and Purkinje cells in conduction disease and atrial CMs in atrial fibrillation. Protocols are becoming available to derive some of these CM subtypes, 25, 26 and we expect that they will soon be used in studying both pathological and cardiotoxic changes. Furthermore, some maladies might benefit from advanced culturing techniques, since certain phenotypes might become evident only under optimised conditions. For example, contractile defects were only uncovered under specific metabolic culture conditions 71 or when engineered tridimensional (3D) microtissues were used 40, 69, 72. Importantly, the human heart is composed not only of CMs but also vascular, smooth muscle and epicardial cells; to better mimic its function, we predict that 3D cardiac tissue structures will be widely implemented, especially where interactions between different cell types might underlie the disease. As an example, ARVC has been modelled in hiPSC‐CMs and these are the major cellular players in the cardiac dysfunction in this disease 30, 34, 73 ; however, the suspected contribution of epicardial cells to fibro‐fatty substitution and the role of inflammation could not so far be studied in two‐dimensional monotypic cultures. The expectation is that complex multicellar structures will be necessary to reflect fully the pathology of the condition. We anticipate that in their second decade, iPSCs will find increasing utility when combined with cardiac tissue engineering. The necessity for better mimics of the multicellular and dynamic conditions of the cardiovascular system that can recapitulate diseases not only with both known or unknown genetic causes but also related to ageing and drug‐induced cardiac damage has already encouraged engineering of three‐dimensional cardiac microtissues. These are beginning to incorporate the different dynamics that reflect blood flow, mechanical stretch and strain and the electrical stimulation. Together with changes in energy substrates, it is expected that these will lead to structurally and functionally mature human myocardium into which biological and biophysical readouts can be built that allow high throughput, real‐time and quantitative measurement of cardiac (patho)physiological status. The hope is that a higher degree of complexity will advance the understanding of how the human heart responds to toxic compounds and disease, improve the integrity of the disease models, and refine the predictability of drug responses. Predicting clinical drug response One of the greatest promises of hiPSC technology, but at the same time its greatest challenge, is in predicting drug responses in a patient‐specific (personsalised) way that disease treatment and prevention can be tailored to the individual. Because hiPSCs capture the genetic background of the person from whom they are derived, they are excellent candidates for recapitulating ‘in a dish’ the variability found among single patients or subgroups of patients. It is one of the few ways forward in coupling genome‐wide association data, which associates disease risk with certain variants in the genome, to proof of causality in humans. This cannot be done in laboratory mice because of the genome differences. The ambition to implement hiPSC‐CMs in precision medicine partially relies on their ability to predict patients’ response to administered drugs. Recent studies provide optimism in this direction 15, 63, 74, although additional consent and focussed investigations will be needed to determine the extent to which individual variability can be distinguished in the hiPSC‐CMs, including mild or severe, acute, early or late responses. Of note, all studies so far have been based on a small number of patients per group and were conducted retrospectively. One goal in the coming years will be to demonstrate that hiPSC‐CMs can be used in prospective study designs, for example, by deriving them from a large cohort of patients (> 200) that are about to undergo a specific drug treatment, applying the same drug to their hiPSC‐CMs, and following‐up over time the patients to find out whether the in vitro responses matched the final clinical outcome. This approach could prove valuable especially in evaluating drug‐induced cardiotoxicity. In addition, a similar approach will ideally be applied in the evaluation of proarrhythmic risk. For example, if hiPSC‐CM‐based platforms for screening arrhythmic events can be combined with genetic‐ and FDA‐collected data for the generation of reliable patient‐specific arrhythmic scores, their real value will become clear in both the choice of individual patient treatment as well as in the drug development process. However, current challenges are not insignificant and suggest that expectations should be tempered in anticipation of more data. The ambition to reduce the incidence of sudden cardiac death as a result of drugs or inherent predisposition may however be a realisable goal in the coming decade. Maturation and ageing A relevant challenge in the field is to find ways to reproduce in vitro the physiological processes of maturation and ageing that the heart naturally experiences from its formation to birth and further during the lifespan of a human being. The heart contracts many millions of times over a lifetime so that defects that are minor in CMs at birth may only be revealed with ageing. The mechanism and process of postnatal CM maturation is incompletely understood and clearly requires environmental factors including hormones, exercise and CM growth by hypertrophy. Approaches used to address this issue include prolonged culture, metabolic manipulation, tissue engineering technologies, electromechanical pacing and other biophysical approaches 21, 22. Promoting adult patterns of metabolic activity already provided a more appropriate basal condition on which to model the response to pathological stimuli, such as in ARVC 34, HCM 71 and diabetic cardiomyopathy 63. Furthermore, anisotropic nanotopography was necessary to distinguish structural differences between control and DMD cardiomyopathy hiPSC‐CMs that were otherwise masked 70. Since both mechanical forces and molecular signalling from nonCM cell types are essential contributors to heart development, formation, ageing and disease progression 75, 76, we anticipate that a combinatorial application of different strategies will likely be most successful in promoting maturation in hiPSC‐CMs. Ideally 3D tissue structures will be developed, where hiPSC‐CMs (subtypes) and other cells are mixed together and organised in microtissues, with or without the addition of extracellular scaffolds, and will be subjected to electrical or mechanical stimulation 77. However, it is still unclear to what extent adult CM properties can be acquired in a culture dish. Nevertheless, some of these improvements in external parameters may contribute to the development of new and reliable methods for screening phenotypic changes also in response to drug treatments. Recognising a mature CM in culture An important question still remains: how do we recognise a mature CM in a culture dish? Our knowledge about human adult CMs relies on that of disease‐free primary tissue, which is scarce and technically challenging to isolate successfully 78, 79, 80. Nevertheless, there is consensus that adult CMs are elongated and rod shaped, the sarcomeres are highly organised, the resting membrane potential is quite negative (−80 to −90 mV), the upstroke velocity rapid (150–350 V·s −1 ), the sarcomeres organised in T‐tubules, the excitation–contraction coupling fast and efficient, the force of contraction relatively strong (10–50 mN·mm −2 ), the mitochondrial content high and the metabolism mainly based on fatty acids (reviewed in 21, 23 ). Ideally a combination of all of these parameters including the structural, molecular and electrophysiological characteristics associated with CM maturation should be assessed to determine whether hiPSC‐CMs resemble myocytes of the human adult heart. However, it is common practice to test only some of these parameters, usually only those that are most important for the disease phenotype to be assessed. We propose that the most informative assays are those based on assessing all aspects of functionality of the cells, including as a minimum the electrophysiology, calcium handling properties, patterns and force of contraction. If, for example, the expression of specific ion channel genes is examined, it is important to bear in mind that this is not always accompanied by a parallel change in the corresponding currents; there are several intermediate steps from transcription to function, including protein synthesis, post‐translational regulation, protein trafficking to the membrane, anchoring of the channel to the membrane, protein turnover and channel regulation by known and unknown accessory proteins and by intracellular signalling 81. Measuring the action potential would then seem more appropriate, since it can give some information on whether the CM population into question displays similar features to adult CMs. Drawbacks of electrophysiological measurements, especially of single‐cell patch clamp, are that they are time consuming and low throughput, and the skills and technology is not readily available in all laboratories. Furthermore, the resulting data refer to the subpopulation of CMs that survived dissociation into single cells; these are usually the most immature in the population. The calcium transients can be measured using calcium‐sensitive dyes and they usually reflect the action potentials, since they are closely related. In this case, complementary information is obtained from the kinetics of the calcium handling, variations of cytosolic calcium concentrations, and the extent of intracellular calcium stores. This type of analysis, much like patch clamp electrophysiology, is also low throughput although optical imaging of voltage and calcium might help increasing the measurement efficiencies. Finally, the force of contraction is another way to determine hPSC‐CM maturation, although it depends on the cell shape and on substrate stiffness 82. Techniques for measuring strain under controlled conditions have been developed 27, 83, 84, 85 and we expect this will increasingly become a parameter that will be evaluated, although specialist technical implementation is required. In summary, we believe that the intended application of hiPSC‐CMs should probably determine the evaluation method to be used for assessing their maturation, but we expect that development of automated methods to analyse voltage and calcium transients and force of contraction simultaneously in both 2D and 3D settings will become a useful tool for disease modelling and drug testing, as well as for testing conditions that may eventually enhance maturation. Additional variables will need to be determined that may play essential roles in modulating CM growth, such as substrate stiffness and specific molecular cues 86 and still it is unclear whether an adult phenotype will ever be completely achieved in vitro. Concluding remarks Patient‐specific models of cardiovascular diseases based on hiPSC‐CMs are proving valuable in advancing our understanding of the complex and sometimes unexpected molecular mechanisms underlying pathological changes. Recent findings provide optimism on the applicability of hiPSC technology to unravel complex disorders, identify cardiotoxic drug effects and ultimately to help defining patient subtypes towards tailored drug treatments. In the future, larger cohorts of patients will be needed from which derive hiPSC‐CMs and their phenotype analysis will tell until which point hiPSC in general, but in particular, their derived CMs can account for variables such as age, gender and medical treatments. Conflict of interest CLM is cofounder of Pluriomics b. v. Supporting information Table S1. Complete list of references used for Fig. 2. Click here for additional data file.
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10. 1002/1873-3468. 12559
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Febs Letters
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Identifying niche‐mediated regulatory factors of stem cell phenotypic state: a systems biology approach
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Understanding how the cellular niche controls the stem cell phenotype is often hampered due to the complexity of variegated niche composition, its dynamics, and nonlinear stem cell–niche interactions. Here, we propose a systems biology view that considers stem cell–niche interactions as a many‐body problem amenable to simplification by the concept of mean field approximation. This enables approximation of the niche effect on stem cells as a constant field that induces sustained activation/inhibition of specific stem cell signaling pathways in all stem cells within heterogeneous populations exhibiting the same phenotype (niche determinants). This view offers a new basis for the development of single cell‐based computational approaches for identifying niche determinants, which has potential applications in regenerative medicine and tissue engineering.
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Abbreviations MS, multiple sclerosis PCST, Prize Collecting Steiner Tree; NSC, neural stem cell Stem cells are indispensable for maintaining tissue homeostasis due to their unique ability to generate more specialized cell types in a well‐coordinated manner depending on the organismal needs. This function depends crucially on the ability of stem cells to make robust cell fate choices such as self‐renewal or differentiation. Multiple cell‐intrinsic and extrinsic factors control this decision‐making process. In this regard, interactions between stem cells and their microenvironment, also known as the niche, determine the stem cell phenotypic states such as quiescent and active stem cells 1. The cellular niche translates information from the neighborhood of the stem cell by transmitting external cues to intracellular signaling events that maintains its cellular state. Schofield in his description of hematopoiesis, proposed the concept of stem cell niche where, a stem cell must be associated ‘with other cells which determine its behavior’ in order to ‘prevent its maturation’; loss of this association was hypothesized to result in differentiation 2. This concept of stem cell niche has evolved over time, and now includes several different supportive stromal cell types, anatomical localization, soluble molecules, as well as physical factors, such as shear stress, oxygen tension, and temperature 3. Involvement of such disparate and stochastically fluctuating components, in addition to feedback regulation of the niche by stem cells, leads to the highly dynamic nature of the niche 1, 4, 5. Stem cells are known to remodel the niche by secreting ECM components and other diffusible factors in response to the signals received from the niche, thus giving rise to feedback regulation of niche–stem cell interactions 4. Such a bidirectional interplay between stem cells and niche is exemplified by the fact that daughter/progenitor cells can serve as niche cells for their parent stem cells in different tissue types 1. These feedback regulatory mechanisms, in addition to the complex bio‐physical characteristics of ECM, contribute to nonlinear stem cell–niche interactions 6. General physiological conditions of tissue and organismal requirements shape the niche effect on stem cell phenotype 7. For instance, healthy tissues under homeostatic conditions are characterized by the tight regulation of stem cell and progenitor cell turnover. However, this tissue‐level homeostasis is often disrupted in case of several diseases such as cancers, neurodegenerative diseases, and cardiac dysfunction. Furthermore, aging is known to contribute toward progressive decline in tissue homeostasis due to degenerative changes in niche‐mediated cues that regulate the stem cell activity 8. In general, complications often arise due to a lack of proper generation of progenitor cells, complete loss of stem cells, and uncontrolled growth of stem/progenitor cells. Deregulated niche components are known to be responsible for several of these defects 9. For such cases, regenerative medicine approaches that rely on transplanting or modulating endogenous stem cells hold immense potential 9. At present, a key challenge in this area includes the limited functional integration (or engraftment) of transplanted stem cells into the target tissue. This has been attributed to the negative regulatory effect of diseased niche on transplanted stem cells 10. In order to overcome this limitation, it is essential to understand those regulatory mechanisms that normally control stem cell functional state in response to the niche. However, the multifactorial complexity of the niche–stem cell interactions is a major roadblock in this direction. Therefore, the role of the niche in maintaining distinct stem cell phenotypic states, and how to influence the niche effect on stem cells to induce transitions among these states constitutes a fundamental problem in stem cell research. Recently, studies have begun to address this issue by explicit characterization of niche components and their interactions with stem cells 11, 12. Despite significant progress in identifying cells that comprise the niche, a comprehensive understanding of all niche components is not yet obtained. This lack of knowledge is predominantly due to the difficulty in obtaining and studying niche cells and factors in vivo. Furthermore, there is a lack of consensus on what actually constitutes the niche and the precise definition of niche components 13, 14, 15. In addition to experimental efforts, a few computational systems biology approaches that model population‐level dynamics of cell–cell interactions have been proposed to study niche regulation of stem cells 16, 17, 18, 19, 20. However, a complete description of stem cell–niche interactions that allows designing strategies for controlling the effect of niche on stem cells is still limited. This is mainly due to incomplete characterization of the niche, fluctuations of the niche components, and a large number of nonlinear interactions between the niche components and stem cells. In this article, we hypothesize that stem cell–niche interactions could be considered as a complex many‐body problem that can be simplified by the concept of mean field approximation. Such a view allows consideration of the net effect of all niche components on stem cells as a constant averaged effect or ‘mean field’. Most existing models consider the niche composition to model stem cell–niche interactions via rate equations. Our approach does not require this knowledge, precisely because it considers that stem cells interact with their niches via a mean field created by all niche components, which ultimately determines the sustained activation/inhibition of specific stem cell‐signaling pathways that maintain their phenotypic states. Application of our view allows the identification of niche‐mediated regulators of stem cell phenotypes by relying on single‐cell profiling data. To support this hypothesis, we use examples of different stem cell systems to illustrate how stem cells maintain their phenotypic state via constant activation or inhibition of certain pathways under homeostatic conditions. Such pathways that determine the stem cell states can be termed as niche determinants, and are expected to be constantly activated/inhibited in all cells within a population sharing the same phenotypic state despite the variability in their molecular profiles. Indeed, knowledge of these niche determinants should enable us to identify target genes whose perturbations can induce transitions between different phenotypic states. Mean field approximation: keeping it simple Mean field theory was initially developed by Pierre Curie and Pierre Weiss in physics for a simplified theory of ferromagnetism 21, 22. They considered a lattice composed of magnetic moments interacting with their nearest neighbors, and proposed to replace the actual interactions experienced by each magnetic moment with the mean interaction (provided by the mean magnetization) by setting the fluctuations around the mean equal to zero. Such an approximation that considers each magnetic moment to be influenced by a mean field created by all their neighboring moments enabled Curie and Weiss to effectively simplify the many‐body interaction problem to a two‐body problem without explicitly accounting for each pairwise interaction. Since its initial proposal, different interpretations of mean field theory have been applied to other disciplines, such as ecology, epidemiology, and protein structure prediction 23, 24, 25. Mean field approximation applied to the stem cell niche Despite the existence of different mean field concepts 24, here, we follow the definition proposed in ferromagnetism. In particular, we hypothesize that stem cells and niche components within a spatial compartment can be viewed as a many‐body interaction system that includes different types of interactions among them (Fig. 1 ). By stem cell niche we invoke the original concept of specialized microenvironment which supports stem cell survival and functions 1, 2. In this regard, even though individual components of the niche can fluctuate, their combinatorial effect on stem cells can be represented by a mean field, which is the average of all the molecular and cellular signals from the niche. A single component of the niche may be perturbed, but it does not form a defective field unless the perturbations spread and completely transforms the entire niche 1. The dynamic equilibrium between the niche and the stem cells is resilient and robust to small perturbations and noise in the individual niche components. Therefore, according to our hypothesis, it is not the interaction between stem cells and individual niche components that determines their state, but rather it is the constant interaction of each stem cell with the mean field that leads to a sustained activation or inhibition of specific stem cell intracellular signaling pathways. This ultimately dictates stem cell function and behavior, governing the choice between quiescence, proliferation, self‐renewal, or differentiation. In this way, not only are discrete fluctuations in niche signals buffered against, but so too are the epigenetic and gene expression heterogeneity that stem cell populations display. According to our view, a given stem cell population (sharing a common phenotype), although exposed to perturbations and noise due to fluctuations in individual niche components in addition to the presence of intrinsic molecular heterogeneity, nonetheless should share commonly activated/inhibited signaling pathways that determine their phenotypic state (Fig. 2 ). Such pathways that determine the stem cell state can be termed as niche determinants (Fig. 2 ). Figure 1 Mean field approximation of stem cell–niche interactions. The mean field approximation considers that each stem cell interacts with its niche via a ‘mean field’ created by all molecular and cellular signals from the niche. The figure depicts the complex nature of stem cell–niche interplay within a spatial compartment. Stem cells (red circles) are entangled in an intricate network of interactions (gray edges) with different niche components ( NC ) (yellow nodes of different shapes). Analyzing the effect of each individual component on stem cell would require consideration of a large number of interactions and fluctuations among them. In the right, the enlarged depiction of a stem cell shows a mean field (yellow cloud) created by the niche components around a stem cell. Figure 2 Niche determinants of stem cell phenotype. Representation of stem cell signaling and gene regulatory network states of a heterogeneous population of stem cells sharing a common phenotypic state. The figure depicts heterogeneity of gene expression at a single‐cell level (red and blue nodes) and the signaling pathways regulating the underlying gene regulatory network. According to the mean field hypothesis, in spite of molecular heterogeneity and fluctuations of niche signals, these cells should share commonly activated/inhibited signaling pathways (niche determinants) that determine their phenotypic state. Such pathways are depicted with red arrows, while the other transient signaling pathway activities not common to all cells in the population are depicted with dashed arrows. The underlying gene regulatory network that maintains the phenotype of these cells is depicted with red and blue nodes representing their expression status. As a consequence of approximating the niche components with an effective mean field, the focus is on identifying sustained signaling (shared within a cellular population) responsible for maintaining the specific stem cell phenotype instead of characterizing the niche explicitly. The proposed approach relies on single‐cell profiling data and works by first identifying the most conserved set of genes (based on the similarity of expression levels at single‐cell resolution) defining that particular phenotype. Subsequently, unique signaling pathways/networks that link the conserved receptors and transcription factors for specific stem cell phenotypes are inferred computationally by relying on network topology and expression levels. Case study: mean field approximation to identify niche determinants of NSCs Based on a mean field approximation hypothesis, we illustrate the applicability of this view of stem cell–niche interactions in order to identify niche determinants of quiescent and active neural stem cell (NSC) phenotypes based on a recently published single‐cell RNA sequencing data 26. The data were obtained from Gene Expression Omnibus ( GSE67833 ). Briefly, these data that we used in our approach are described as follows: mouse subventricular zone NSCs were isolated from their natural environment based on the expression of GLAST and Prom1. The transcriptome of 104 GLAST+/Prom1+ cells were analyzed by single‐cell RNA‐seq using Smartseq2 technology 27. These data were then subjected to principal component analysis followed by unsupervised hierarchical clustering of genes with the highest coordinates in the first four principal components (1844 genes) 26. This analysis partitioned the NSCs into two major clusters. One NSC cluster had Egfr expression (a known marker of active NSCs 28 ) in addition to the expression of cell cycle‐related genes. Based on these attributes, this cluster was defined as active NSCs. On the other hand the cluster that lacked the activation markers were classified as quiescent NSCs. Gene ontology and pathway enrichment analysis revealed that active NSCs were enriched in genes for cell cycle, protein synthesis, and mitosis, whereas glycolytic metabolism was found to be most enriched in quiescent NSCs. Gene ontology and pathway enrichment analysis further divided quiescent and active NSCs into two subpopulations each (quiescent NSC1/2 and active NSC1/2). In our current analysis for the sake of simplicity we considered only quiescent and active NSC populations as a whole without considering the further subpopulations. Our strategy relies on gene expression differences between stem cells displaying different niche‐dependent phenotypes, and aims to infer sustained signaling pathways that are required for stably maintaining their corresponding phenotypes. Moreover, despite the niche‐induced fluctuations in signaling, such pathways must be shared (or conserved) within the cells sharing a common phenotype. However, it must be mentioned that identification of conserved pathways can also result in housekeeping pathways that could be of general importance to a wide variety of cell populations (e. g. , pathways that are important for both quiescent and active NSCs) and therefore could lack cell type specificity. In order to overcome this issue, the approach focuses on uniquely conserved pathways within each population and is different across the populations. Single‐cell gene expression data offer the possibility to identify the set of genes whose expression pattern is conserved within a given phenotype. Such genes are more likely to play a dominant role in phenotype maintenance since their expression pattern is similar at single‐cell level. In the example of NSCs, we first identified the genes exhibiting similar expression pattern within quiescent or active phenotype. For this we employed Shannon entropy 29, which measures the disorder of a system, where lower values indicate similar expression pattern of a given gene. Entropy for each gene, X, is defined by: H X = − ∑ i = 1 n P ( x i ) log 2 p ( x i ) where P(x i ) represents probability of gene expression value x = x i. Entropy calculation was performed using data binning approach and the number of bins ( k ) was determined from the expression data using Sturges' rule 30, given by k = log 2 n + 1, where n is the sample size. After data binning, the computation of entropy was performed using maximum likelihood implementation (entropy. empirical) of the R entropy package. We used an entropy cutoff less than 1 and median expression (FPKM) value greater than 10 to classify the gene as having a conserved expression pattern. Entropy calculation for each gene allowed us to identify quiescent or active phenotype‐specific genes that showed similar expression pattern at a single‐cell level. Next, we sought to identify those signaling pathways that are more likely to be constantly active. For this, we first identified the set of receptors/ligands and transcription factors classified as conserved for quiescent and active NSCs. Entropy calculation based on single‐cell expression levels allowed us to identify the genes that shared a similar expression levels. From that list of genes, transcription factors and transcriptional regulators were identified based on annotation available at Animal TFDB ( http://www. bioguo. org/AnimalTFDB/ ). In the case of receptors, since a complete database of receptor molecules is currently unavailable, we used Gene Ontology classification of receptor activity and plasma membrane (GO:0004872, GO:0005886) to identify genes with possible receptor activity. For the case of secreted ligand molecules we utilized the classification of potential ligands reported in a recent study 31. About 90 and 128 receptors/ligands were identified for quiescent and active NSC phenotypes, respectively. From this, identifying the ones that are most likely to propagate the niche mediated signaling is a challenge. We made use of literature‐curated signaling database Reactome 32 as a background raw signaling network consisting of all reported signaling interactions and employed Prize Collecting Steiner Tree (PCST) formalism to infer the signaling pathways. Interactions reported in the Reactome database were used as the background network from where subsequent Steiner trees were inferred. Reactome consists of curated pathways with molecular interaction data from Reactome Functional Interaction Network and other databases such as IntAct, BioGRID, ChEMBL, iRefIndex, MINT, and STRING. We specifically used Reactome Functional Interaction Network ( http://www. reactome. org/pages/download-data/ ) as they contain information on direction and sign (positive of negative regulatory effect) of the interaction. We consider that the conserved receptors/ligands of a given stem cell phenotype are under the direct influence of the niche. Since the exact mechanisms of the niche effect on the signaling activity are not known, we represent the net effect of the niche by introducing a dummy niche node in the raw signaling network. The external dummy node is incorporated as a way to capture the topologically favorable receptors/ligands (from several expressed ones) that can link it to the TFs specific for quiescent and active NSCs. Furthermore, the dummy node is used as the root node which acts as the starting point for Steiner tree identification, consequently the receptors/ligands will be linked to the dummy node in the inferred Steiner trees. This dummy node is connected to all conserved receptors/ligands for each phenotype under consideration. Therefore, signal transduction from the niche to transcription factor must be propagated through at least one of the conserved receptors. The edges in the signaling interactome were weighted using the gene expression data, where the weights were calculated as, c e = 1 x i x j, where x i and x j are the expression levels of the interacting nodes. We specifically used such a weighting scheme since the objective of the PCST algorithm is to collect as many high prize nodes (genes with high expression) while minimizing the edge weights. Such an edge weighting scheme that inversely correlates with the expression levels will enable collecting those edges where both nodes are highly expressed. In such a weighted raw signaling network, that has a dummy niche node representing the net effect of the niche, we used PCST to infer subnetworks with the dummy niche node as the root or origin node and the conserved transcription factors as the terminal nodes. Steiner Tree formalism has been used earlier to reconstruct active signaling pathways 33, 34. Formally, the PCST problem is defined as, given a graph G = (V, E), representing the raw signaling interactome (where, V denotes the nodes and E denotes the edges), with defined edge costs (weights), c e and node prizes b v find a connected subgraph T = (V′, E′), V′ ⊆ V, E′ ⊆ E, that minimizes the following function: T = min ( E ′, V ′ ) connected ∑ e ∈ E ′ c e − λ ∑ v ∈ V ′ b v The node prizes are computed by b v = |log fold change ( V )| from the gene expression data and c e is the edge weights. The constant λ determines the tradeoff of adding new proteins to the inferred network by balancing the cost of new edges and the prize gained by adding a new protein. We chose λ = 0. 01 for our simulations and employed a heuristic method based on a message‐passing algorithm to infer the PCSTs 33. Basically, minimizing this function implies collecting the largest set of high prize nodes while minimizing the set of high cost edges in a tradeoff tuned by λ that results in a connected subgraph. Since the dummy node is connected only to the conserved receptors of a given cell type, the inferred subnetworks will encompass only those receptors that are both topologically favorable and maximize the expression values of the intermediate nodes. Therefore, from several conserved receptors, one could narrow down to the few linking the transcription factors based on their unique network topological features and expression levels. Employing the above strategy, we identified subnetworks that are likely to maintain the quiescent and active phenotypes of NSCs (Fig. S1). In the case of quiescent NSCs, we identified nine subnetworks with receptors as origins/sources responsible for controlling the expression status of the downstream terminal transcription factors (Fig. S2). Among such identified receptors, the role of Bmpr1b, Notch2, and S1pr1 are known in the case of quiescent NSCs. In fact, BMP signaling is known to maintain the NSC quiescence in an autocrine manner, and further this signaling must be downregulated for the subsequent activation of the quiescent NSCs 26. On the other hand, Notch signaling is known to be involved in a paracrine manner where Notch ligands are expressed by active NSCs and inhibition of Notch signaling increased the active stem cell population 26. Role of S1pr1 in maintaining NSC quiescence has been demonstrated in an independent study where addition of S1pr1 agonist sphingosine‐1‐phosphate significantly affected the activation of quiescent NSCs 28. In the case of active NSCs, we identified Egfr signaling in addition to five other receptor‐mediated signaling pathways(Fig. S3). Moreover, role of Egfr signaling for maintaining active NSCs is well established and in fact Egfr is used as a marker to isolate those cells 28. In principle, such an approach that focuses on sustained signaling pathways conserved within a cellular population could enable identification of niche‐mediated regulators of stem cell phenotypes without the knowledge of niche. Mean field approximation: caveats and comparisons to other models As a result of mean field approximation, transient fluctuations in signaling events that arise due to the dynamic nature of the niche are ignored, as they do not display any functional consequence for the maintenance of stem cell states. In this context, it must be noted that in addition to sustained signals, a cellular niche can also propagate transient, but functionally relevant signals induced by feedback mechanisms to robustly maintain tissue homeostasis 35. Other transient, yet functionally important signals could arise due to perturbations such as cellular injury or genomic mutations. The latter signals generally induce stem cell phenotypic transitions (i. e. , from quiescent to active/proliferative state 5 ), but are less likely to stably maintain the existing stem cell phenotype 36, 37. Therefore, it must be emphasized here that the mean field view of stem cell–niche interactions is valid for identifying the signaling pathways responsible for constant maintenance of cellular phenotypes and not for transient signals that can potentially trigger phenotypic transitions. Furthermore, identification of conserved signaling can provide accurate descriptions of individual cellular behavior only when heterogeneity within a defined population reflects functionally meaningless fluctuations around a single cellular state and not otherwise. Therefore, for the approach to yield accurate results, the characterization of the cellular populations needs to be accurate. Greater emphasis on the identification of sustained signaling pathways that are conserved within a cellular population exhibiting a common phenotype is a major outcome of the mean field approximation of the niche. Even though this outcome appears similar to pathway enrichment analysis that has been routinely utilized over the past decade 38 to identify deregulated (signaling or metabolic) pathways, in actual practice the idea has not been identification of sustained signaling pathways conserved within a cellular population. Moreover, several transient signaling pathways could be identified as deregulated due to indirect effects (of mutations, differences in the niche composition etc. ) and not as a cause for observed phenotypic difference. However, those signaling pathways that are constantly active are more likely to be the cause for stable maintenance of a specific cellular phenotype. Such a view offered by our hypothesis is fundamentally different from the prevailing view, and is often overlooked due to its apparent simplicity. Furthermore, it must be mentioned that computational analysis based on such a view enhances the utility of single‐cell omics data generation and adds value to current development of analytical methods 39 to decipher hidden patterns in such high‐resolution datasets. Given the complexity involved in stem cell–niche interactions, computational systems biology approaches have been useful in modeling their behavior. In fact, computational methods have been proposed to model interactions between stem cells and niche components 16, 17, 18, 19, 20, 40, 41, 42. These methods could be broadly classified into two major categories, (a) methods that aim to capture the population level behavior of stem cell–niche interactions by modeling cell–cell interaction dynamics and (b) construction of intercellular (cell–cell) interaction networks based on gene expression data. The first category of methods model the interaction dynamics of stem and progenitor cells using rate equations that describe the birth and death processes of each cell type and their interdependence on each other 16, 17, 18, 19, 20, 40. Such models are most commonly employed for studying stem cell–niche interaction dynamics and characterizing the system steady‐state properties in order to understand tissue homeostasis, and how perturbations (in the form of diseases) could affect the original steady states. A typical bottleneck in such dynamical models is the lack of knowledge of parameters or probabilities (such as, self‐renewal rate, synthesis rate of differentiated cells, death rates of stem and daughter cells) that govern the system dynamics. In addition to a lack of knowledge on parameters, even the precise composition of the cellular niche is far from being completely known, thereby rendering the development of such dynamical models difficult. Furthermore, these models tend to be powerful for a descriptive analysis of the system dynamics rather than being predictive in nature. In contrast, our proposed approach does not require the explicit knowledge of niche components or the parameters that govern the stem cell–niche interactions to identify niche‐mediated regulators of stem cell phenotype. The second category of models are based on construction of intercellular interaction networks based on gene expression data 41, 42. This approach attempts to build cell–cell interaction networks based on sorting of different cell populations followed by high‐throughput profiling, to define intercellular signaling between phenotypically defined populations of stem, progenitor, and mature cell types. This approach, although not affected by a lack of knowledge on parameters, nevertheless requires sorting and profiling of several cell types to construct the cell–cell interaction network. This is a major limitation since the cell types that truly serve as niche cells in several stem cell systems is not well characterized, and therefore cannot be sorted and profiled easily. However, our proposed strategy requires single‐cell gene expression profiling of only the stem cells with distinct phenotypes (like quiescent and active) and does not require expression profiling of the niche cells. This dramatically simplifies the isolation of the cells, data generation and further downstream analysis since only stem cells are required to be isolated and profiled without the necessity of profiling the niche cells. Although every stem cell system is unique in the way it is regulated by its niche 3, several recent studies in different stem cell systems have observed that stem cell states are determined by constant activation/inhibition of specific pathways by the constitutive influence of its niche 15, 28, 43. The presence of certain constantly activated/inhibited signaling pathways maintained by their niche appears to be the commonality in different stem cell systems. This offers possibilities to address the complexity of stem cell–niche interactions without the explicit niche characterization. Especially, the rapid advancements in single‐cell profiling technologies enable the dissection of cellular populations in greater detail. Moreover, the development of computational systems biology approaches based on the mean field approximation hypothesis finds a natural application of such increasingly available data for identifying signaling pathways that are constantly active in all cells within a population exhibiting the same phenotype. Importantly, identification of such niche determinants has several implications in regenerative medicine. Potential applications for regenerative medicine and tissue engineering The stem cell niche contains a rich and diverse set of cues that impinge constantly on stem cells that can be modulated for therapeutic gain 9, 10. Understanding and characterizing the niche determinants has potential applications in regenerative medicine and stem cell therapies for degenerative diseases of liver, heart, lung, and brain. Limited functional integration of transplanted stem cells into the target tissue possibly due to negative regulatory effect of diseased niche is currently a major challenge 10. In this regard, promoting regeneration by harnessing the latent regenerative potential of endogenous stem/progenitor cells has been used as an alternative regenerative medicine strategy in order to overcome the current translational bottlenecks associated with cell transplantation 44. For example, in the case of multiple sclerosis (MS), a demyelinating disease due to progressive failure of remyelination in the CNS due to aging, endogenous activation of oligodendrocyte precursors by mimicking a youthful microenvironment have been proven useful to promote remyelination in certain MS disease models 45, 46. In order to achieve this, identification of strategies for the activation of endogenous repair mechanisms to promote tissue regeneration in situations in which it does not occur normally is necessary 44. Within this context, the proposed approach for the identification of conserved signaling pathways under diseased and healthy niche conditions (determined by their physiological cues) can enable the development of potential strategies to modulate endogenous stem cell activity by either counteracting the effect of diseased niche or by mimicking the effect of healthy niche in the diseased counterpart. Such intervention strategies would be intended to make endogenous stem cells resistant to the perturbed signals in the diseased state and to sustain long‐term function. Another potential application where the knowledge of niche determinants can provide useful insights is in the area of tissue engineering. In particular, it is relevant in the context of ex vivo tissue engineering, where the main goal is to have the cells surviving and functioning in an optimal environment without necessarily having to replicate the in vivo conditions. In this regard, our proposed approach can enable identification of key factors that are responsible for maintaining a given cellular phenotype in vivo can aid defining better culture conditions for long‐term phenotype maintenance. For instance, long‐term maintenance of primary hepatocytes in a defined culture medium is still a challenge 47. Specifically, identification of a culture system that can facilitate long‐term maintenance of hepatocytes is advantageous for clinical applications such as drug screening and toxicity tests. Conclusions In general, cellular populations with the same functional phenotype exhibit a certain degree of heterogeneity in their molecular profiles due to intrinsic stochasticity in the transcriptional and translational program. Furthermore, the dynamical nature of the niche can perpetuate noisy fluctuations in stem cell signaling pathway activities. Therefore, stem cells face an acute challenge of robustly maintaining their state in the presence of intracellular and extracellular fluctuations, while responding precisely to developmental cues from the niche. The existence of a common stem cell phenotype within a spatial compartment of a tissue, despite the dynamic nature of the niche, seems contradictory. Our mean field view of stem cell–niche interactions provides an explanation for such a seemingly contradictory observation. By focusing on the net effect of the niche created by the mean field after disregarding internal and external fluctuations, it points to the existence of constantly activated/inhibited signaling pathways that maintains the stem cell state in response to the niche. In fact, identification of conserved signaling pathways that are constantly activated/inhibited in all cells in a stem cell population exhibiting the same phenotype will confirm our hypothesis. Furthermore, the development of single‐cell data‐based computational methods relying on a mean field view of the niche can aid in identification of niche determinants by simplifying the complexity of stem cell–niche interactions. Importantly, the knowledge of niche determinants will aid developing regenerative medicine strategies to enhance/modulate stem cell activity for the treatment of injury, disease, or age‐related dysfunctions. In addition, our approach is suitable for identifying factors that can facilitate long‐term maintenance of cells under culture conditions. Thus, combining recent developments in single‐cell technologies and stem cell research with the systems biology approaches discussed here should enable us to more accurately identify niche determinants, which in turn could lead to the implementation of more feasible strategies in regenerative medicine and tissue engineering. Author contributions AdS conceived the idea. SR performed the analysis. Both the authors wrote the manuscript. Supporting information Fig. S1. Inferred Steiner trees for quiescent and active NSCs. It can be seen that the dummy node in the center is the root node that connects with all receptors/ligands. The inverted triangles depict receptor molecules, circles depict signaling intermediates, and squares depict transcription factors. Click here for additional data file. Fig. S2. The figure shows the subnetworks of signaling pathways identified for quiescent NSCs. The inverted triangles depict receptor molecules, circles depict signaling intermediates, and squares depict transcription factors. The experimentally validated signaling pathways are highlighted. Click here for additional data file. Fig. S3. The figure shows the subnetworks of signaling pathways identified for active NSCs. The inverted triangles depict receptor molecules, circles depict signaling intermediates, and squares depict transcription factors. The experimentally validated signaling pathways are highlighted. Click here for additional data file.
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10. 1002/1878-0261. 12323
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Molecular Oncology
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A combined tissue‐engineered/
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Patient‐tailored therapy based on tumor drivers is promising for lung cancer treatment. For this, we combined in vitro tissue models with in silico analyses. Using individual cell lines with specific mutations, we demonstrate a generic and rapid stratification pipeline for targeted tumor therapy. We improve in vitro models of tissue conditions by a biological matrix‐based three‐dimensional (3D) tissue culture that allows in vitro drug testing: It correctly shows a strong drug response upon gefitinib (Gef) treatment in a cell line harboring an EGFR ‐activating mutation ( HCC 827), but no clear drug response upon treatment with the HSP 90 inhibitor 17 AAG in two cell lines with KRAS mutations (H441, A549). In contrast, 2D testing implies wrongly KRAS as a biomarker for HSP 90 inhibitor treatment, although this fails in clinical studies. Signaling analysis by phospho‐arrays showed similar effects of EGFR inhibition by Gef in HCC 827 cells, under both 2D and 3D conditions. Western blot analysis confirmed that for 3D conditions, HSP 90 inhibitor treatment implies different p53 regulation and decreased MET inhibition in HCC 827 and H441 cells. Using in vitro data (western, phospho‐kinase array, proliferation, and apoptosis), we generated cell line‐specific in silico topologies and condition‐specific (2D, 3D) simulations of signaling correctly mirroring in vitro treatment responses. Networks predict drug targets considering key interactions and individual cell line mutations using the Human Protein Reference Database and the COSMIC database. A signature of potential biomarkers and matching drugs improve stratification and treatment in KRAS ‐mutated tumors. In silico screening and dynamic simulation of drug actions resulted in individual therapeutic suggestions, that is, targeting HIF 1A in H441 and LKB 1 in A549 cells. In conclusion, our in vitro tumor tissue model combined with an in silico tool improves drug effect prediction and patient stratification. Our tool is used in our comprehensive cancer center and is made now publicly available for targeted therapy decisions.
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Abbreviations 17AAG 17‐ N ‐allylamino‐17‐demethoxygeldanamycin 2D two‐dimensional 3D three‐dimensional ADME absorption, distribution, metabolism, excretion AICAR 5‐aminoimidazole‐4‐carboxamide ribonucleotide BioVaSc ® Biological Vascularized Scaffold COSMIC Catalogue Of Somatic Mutations In Cancer DRPs differentially regulated proteins DrumPID Drug‐minded Protein Interaction Database Gef gefitinib HE hematoxylin and eosin HPRD Human Protein Reference Database HSP heat shock protein NGS next‐generation sequencing NSCLC non‐small cell lung cancer PK phospho‐kinase RTK receptor tyrosine kinase SMILES Simplified Molecular Input Line Entry Specification SQUAD Stardardized Qualitative Dynamical Modelling Suite TKI tyrosine kinase inhibitor 1 Introduction In the highly mortal lung cancer, next‐generation sequencing (NGS) approaches successfully reveal driver mutations to stratify lung cancer patients for targeted therapies (Buettner et al. , 2013 ). Tyrosine kinase inhibitor (TKI) treatment shows remarkable response rates, exemplified by EGFR inhibitors in patients with activating EGFR mutations (Ciardiello et al. , 2004 ; Paez et al. , 2004 ; Russo et al. , 2015 ). However, often the therapy is only initially successful and then followed by secondary resistance. Unfortunately, tumors with KRAS mutations are primarily resistant to targeted therapies and comprise about 30–40% of all patients (Sequist et al. , 2011 ). Due to poor correlations of preclinical in vitro data to drug efficacy in patients, particularly in the field of cancer (Bhattacharjee, 2012 ), new 3D tumor models arise, such as spheroids, microfluidic devices, organoids, and matrix‐based approaches (Alemany‐Ribes and Semino, 2014 ; Edmondson et al. , 2014 ; Xu et al. , 2014 ). The generally high proliferation rate in 2D cell cultures is one reason for false‐positive predictions of cytostatic compounds (Cree et al. , 2010 ). Decreased proliferation of tumor cells corresponding to clinical specimens was demonstrated on our scaffold (Göttlich et al. , 2016 ; Nietzer et al. , 2016 ; Stratmann et al. , 2014 ) originating from the Biological Vascularized Scaffold (BioVaSc ® ) (Linke et al. , 2007 ; Schanz et al. , 2010 ). It maintains the extracellular matrix, including structures of the basement membrane, enabling physiological anchorage of epithelial cells. Earlier, we combined the tissue‐engineered lung tumor model with its in silico representation to investigate tumor and, thereby, drug‐relevant dependencies – also in the context of resistance (Göttlich et al. , 2016 ; Stratmann et al. , 2014 ). In this study, we introduce a patient stratification tool according to tumor drivers as a promising decision tool for precision medicine in lung cancer. This is exemplified here by studying individual in vitro cell lines and their differing drug responses in 2D and 3D, and by integrating these data in corresponding in silico analyses for target predictions. The tool is generic and provides a rapid stratification pipeline that can support tumor boards to utilize more and more clinically available NGS data from individual patients. We studied how a biological matrix‐based 3D tissue culture allows in vitro drug testing of relevant lung cancer subgroups. To unravel signal cascade outputs in more detail, we investigated apoptosis and proliferation as drug responses. Regarding the EGFR inhibition with the TKI gefitinib (Gef) in a cell line carrying the corresponding biomarker, we observed an enhancement in apoptosis induction compared to 2D. Moreover, we exemplified our stratification tool by looking at responses of two further cell lines (A549, H441) harboring KRAS mutations to the HSP90 inhibitor 17AAG. In contrast to the EGFR inhibition, in this setting only the 3D system could predict no drug efficiency in line with clinical findings. Therefore, we analyzed differences in signaling changes upon treatment between cell lines and between 2D and 3D conditions. Using the experimental data of the 3D tissue model, we created (a) in silico cell line‐specific topologies of the centrally involved proteins including their logical connectivity. Based on these data, (b) dynamic in silico simulations mirrored the differences in cellular responses apparent in the experiments. Considering protein neighbors of central important signaling cascades and cell‐specific mutations from databases resulted (c) in larger in silico networks which were next screened in silico for individual therapeutic options for each cell line. Resulting drug suggestions reflect clinical experiences and include comprehensive FDA‐approved treatment options. In its unique combination, the tool raises hopes of efficiently exploiting upcoming sequence information of patient tumors in the near future for targeted therapy. 2 Results 2. 1 Analysis path To exemplify the process of how a single patient's sequence could be integrated into our new preclinical prediction tool, we chose three cell lines representing different patient subgroups regarding KRAS mutation (HCC827: KRAS wild‐type; A549: KRAS mutant, independent; H441: KRAS mutant, dependent) (Singh et al. , 2009 ). To unravel the complex interdependent signaling network in lung cancer in different mutational backgrounds, we experimentally measured, in a global approach using phospho‐arrays signaling, changes to Gef and 17‐allylamino‐17‐demethoxygeldanamycin (17AAG) treatment in our three different cell lines (Figs 1, 2, 3 ; Table 1 ; Figs S1–S3 ; first simulation is in S2 ; further in silico analyses in S4–S7 ). Firstly, we recognized that with the EGFR inhibitor Gef affected proteins are roughly the same in HCC827 (EGFR mutated) in 2D and 3D models (Table 1 A, Fig. S1 ). By analyzing signaling changes upon the HSP90 inhibitor treatment by phospho‐arrays and western blot in all three cell lines in 2D and in 3D, it became obvious that besides MET, changes between the 2D and the 3D models concern mostly p53 and HSP60 (Table 1 B; Figs 4 A and S3 ). Figure 1 Improved reflection of tumor characteristics by the 3D tissue‐engineered lung tumor model. Tested tumor cell line populations display more homogeneous marker expression in 3D as well as reduced proliferation correlating to tumors. (A) Cells cultured in 2D and 3D conditions shown with immunofluorescence by double stain for E‐cadherin and β‐catenin. Green arrows indicate positive cells and white arrows negative. Scale bars are 50 μm. (B) Paraffin‐embedded adenocarcinoma from patient biopsy was immunofluorescence‐double‐stained against E‐cadherin and β‐catenin. Scale bar is 100 μm. (C) The expression of the proliferation marker Ki67 was detected by immunofluorescence staining of 2D and 3D cultured HCC 827 cells, as well as in vivo tissue from a patient biopsy. Scale bar is 100 μm. Figure 2 Biomarker‐dependent response upon EGFR inhibition is improved in 3D and can also be simulated in silico. (A) Cells cultured in 3D that were either treated with 1 μ m Gef or used as untreated controls were paraffin‐embedded and HE ‐stained. Scale bar is 100 μm. (B) The proliferation rate (proliferative cells per total cell number) was determined by counting Ki67‐positive cells from immunofluorescence staining in 10 images per sample. Total cell number was quantified by DAPI counterstaining. *** P < 0. 001, n ≥ 4. (C) Apoptosis was investigated by M30 CytoDeath™ ELISA. Therefore, supernatants of treated and untreated samples were collected prior to and at 24, 48, and 72 h after treatment. Concentrations of M30 in samples after treatment were normalized to T0 values from samples taken before treatment and related to untreated samples (red line). *** P < 0. 001, n ≥ 4. (D) In silico simulation of the Gef treatment (right, pink curve full on at 1. 0) shows reduced proliferation (right, black curve) only in HCC 827 cells and higher apoptosis (right, gray curve), as compared to untreated cells (left, pink curve switched off at 0. 0). Figure S2 A shows the in silico topology and Fig. S2 B the simulations for A549 and H441. * P < 0. 05, ** P < 0. 01 Figure 3 Effects of the HSP 90 inhibitor 17 AAG diminish in 3D and cannot be aligned to the biomarker KRAS (A549/H441). Strong treatment responses regarding viability, proliferation, and apoptosis can be observed only in 2D conditions. (A) Cells cultured in 2D conditions were treated with different concentrations of the HSP 90 inhibitor 17 AAG. Viability was determined after 3 days of treatment by a CellTiter‐Glo ® Luminescent Cell Viability Assay. n ≥ 4. (B) 3D cultured cells were treated with 0. 25 μ m 17 AAG, paraffin‐embedded, and HE ‐stained. Scale bar is 100 μm. (C) The proliferation rate in 2D and 3D was determined by counting Ki67‐positive cells from immunofluorescence staining in 10 images per sample. Total cell number was quantified by DAPI counterstaining. * P < 0. 05, n ≥ 4. (D) Apoptosis was investigated by M30 CytoDeath™ ELISA. Therefore, supernatants of treated and untreated samples were collected prior to and at 24, 48, and 72 h after treatment. Concentrations of M30 in samples after treatment were normalized to T0 values from samples taken before treatment and related to untreated samples (red line). * P < 0. 05, *** P < 0. 001, n ≥ 4. ** P < 0. 01. Table 1 Comparison of the phosphorylation data showing different regulation between the cell lines in the 2D and 3D system for Gef and 17AAG HCC827 A549 H441 2D 3D 2D 3D 2D 3D (A) Gef treatment a pEGFR ↓ ↓ 0 0 const. const. pErbB2 ↓ 0 0 0 0 0 pMET ↓ ↓ 0 0 const. const. (B) 17AAG treatment b pEGFR ↓ ↓ 0 0 ↓↓↓ ↓↓ pErbB2 ↓ ↓ 0 0 ↓ ↓ pErbB3 ↓ ↓ 0 0 ↓ ↓ pMET ↓↓ ↓ 0 0 ↓↓↓ ↓ pc‐Ret ↓ ↓ 0 0 ↓ ↓ pVEGFR2 0 0 0 0 ↓ ↓ pFGFR3 0 0 0 0 0 ↓ p‐p53 (S46) ↑↑↑ const. 0 0 const. ↑↑↑ HSP60 const. const. ↑ ↑↑ ↑ const. a Based on the RTK array data, this is a qualitative summary of all proteins measured, showing a phosphorylation difference in at least one cell line upon Gef treatment (0 reflects no activation, and const. means no activation change after treatment). Experimental data are shown in Fig. S1. b Based on the western blots (semiquantitative, more than one arrow possible) and RTK array data (qualitative, only one arrow possible), this is a summary of all proteins measured, showing a phosphorylation difference in at least one cell line upon 17AAG treatment (detailed experimental data shown in Figs 4 A and S3A; 0 reflects no activation, const. means no activation change after treatment). John Wiley & Sons, Ltd Figure 4 Signaling changes after HSP 90 inhibition differ between 2D and 3D and between the different cell lines and are integrated into in silico topologies. (A) Cells cultured in 2D and 3D were treated with 0. 25 μ m 17 AAG for 24 h (2D) or 72 h (3D). The signaling changes of different phospho‐proteins were analyzed by western blot. The same lysates were used for the pEGFR and ph‐p53(S46) blots of all three cell lines in 2D and 3D HCC 827 and for ph‐p53(S46) and pMET blots in 3D H441; thus, the same β‐actin loading control is shown below these phospho‐proteins. (B) DRPs from the in vitro 3D system are connected in silico to the central tumor signaling cascade. Here, we show the topology shared between all three cell lines. Colors reflect important input (treatment), signaling proteins, and cellular output (proliferation and apoptosis). Proteins (‘nodes’) from the topology of Stratmann et al. ( 2014 ) are bold rimmed and have an olive background; proteins added specifically to the in silico topology are presented as simple boxes; protein node colors are as in the simulation curves; cell line‐specific proteins (‘nodes’) appear as plus (+). Specific topologies and simulation results for each cell line are given in the Supporting information. Regarding in silico analyses, we first set up cell line‐specific in silico topologies by integrating important signaling nodes that distinguished the cell lines upon Gef and 17AAG treatment into our basic in silico topology (Table 2 ; Stratmann et al. , 2014 ). The nodes of this basic topology are marked in all newly generated topologies with bold printed borders. After the generation of these cell line‐specific in silico topologies, we mirrored the in vitro treatment response of Gef and 17AAG, by applying semiquantitative Boolean simulations using the software squad (Stardardized Qualitative Dynamical Modelling Suite). Based on the logical connectivity of each cellular topology, this software models the dynamic evolution of the included signaling cascades using exponential functions (Di Cara et al. , 2007 ). Furthermore, different activation strengths for each node of the signaling cascade are considered in the simulations that were necessary to adapt the in silico simulation results to the in vitro results for differences of 3D and 2D cultures. Input into the topology of Fig. 4 B is listed in Table 3 A and B for 3D conditions (further network analyses in Tables S1–S3 ) and in Box S2 for 2D conditions. Simulations’ output of 3D conditions is presented in Figs 2 D and S2 for Gef and in Fig. 5 for 17AAG treatment. Simulation results in 2D conditions of KRAS‐mutated cell lines are represented in Fig. S6 for Gef and in Fig. S7 for 17AAG treatment. Table 2 Cell line‐specific proteins introduced in addition to the original topology. a Cell line Gef 17AAG HCC827 (3D) MET cascade + MET cascade + ; Erb2 cascade + ; Erb3 cascade + ; c‐RET cascade + ; HSP90; HSP60; HIF1A; p53 A549 (3D) – HSP90; HSP60 HIF1A; p53 H441 (3D) MET cascade + MET cascade + ; HSP90; HSP60; HIF1A; p53; Erb2 cascade + ; VEGF2 cascade + ; Erb3 cascade + ; c‐RET cascade + ; FGFR3 cascade + a Listed are the proteins extending the network of Stratmann et al. ( 2014 ) that responds upon Gef or 17AAG treatment and which were measured in arrays and western blots. According to interaction analysis p53, HSP60, HIF1A and HSP90 are added as cascades around 17AAG. A plus (+) indicates cell line‐specific protein nodes added according to the experimental data. Further cell line‐specific protein nodes according to COSMIC and relevant to our in silico network as being close to or in our signaling cascades are listed in Table S3 (9 in A549, 18 in H441). Cell‐specific mutations analyzed in detail are shown in Figs 6 and 7. A complete list of all cell line‐specific mutations known is given in Table S2. John Wiley & Sons, Ltd Table 3 Different activation strengths for each node for in silico simulations. Cell line‐specific differences in pathway activities on (A) Gef a and (B) 17AAG a and (C) AMPK activator and HIF1A inhibitor b Cell line Parameter (−) gef (+) gef (A) HCC827 (3D) MET 0. 22 0. 15 EGFR 0. 22 0. 12 EGF‐EGFR 0. 22 0. 12 FLIP 0. 6 0. 5 A549 (3D) KRAS c 0. 413 0. 413 FLIP 0. 6 0. 6 H441 (3D) KRAS c 0. 43 0. 43 EGFR 0. 205 0. 205 MET 0. 205 0. 205 FLIP 0. 4 0. 4 Cell line Parameter (−) 17AAG (+) 17AAG (B) HCC827 (3D) EGFR 0. 25 0. 2 MET 0. 25 0. 2 Stress 0. 6 0. 6 HSP60 0. 4 0. 4 FLIP 0. 7 0. 6 p53 0. 4 0. 4 (Erb2/Erb3/c‐RET) 0. 07 0. 07 Erb2/Erb3/c‐RET 0. 1 0. 07 (EGF‐EGFR) 0. 1935 0. 1935 (MET) 0. 1935 0. 1935 A549 (3D) KRAS c 0. 352 0. 352 Stress 0. 5 0. 5 FLIP 0. 7 0. 33 p53 0. 0 0. 0 HSP60‐act 0. 4 0. 4 H441 (3D) KRAS c 0. 43 0. 43 EGFR 0. 05 0. 01 Erb2/Erb3/c‐RET/FGFR3 0. 04 0. 03 MET 0. 05 0. 02 Stress 0. 7 0. 7 p53‐act 0. 65 0. 65 HIF1‐act 0. 65 0. 65 HSP60‐act 0. 05 0. 05 VEGFR2 0. 35 0. 33 FLIP 0. 75 0. 75 PTEN 0. 41 0. 41 Cell line Parameter (−) AICAR (+) AICAR (−) PX‐478 (+) PX‐478 (C) A549 (3D) KRAS c 0. 345 0. 345 Stress 0. 5 0. 5 FLIP 0. 6 0. 2 p53 0. 0 – p53‐act – 0. 1 low glucose 0. 3 0. 3 HIF1‐act 0. 8 0. 8 mTOR‐act 0. 65 0. 65 H441 (3D) KRAS c 0. 43 0. 43 EGFR 0. 05 0. 05 Erb2/Erb3/c‐RET/FGFR3 0. 04 0. 04 MET 0. 05 0. 05 Stress 0. 7 0. 7 p53‐act 0. 65 0. 65 HIF1‐act 0. 65 0. 65 HSP60‐act 0. 05 0. 05 VEGFR2 0. 35 0. 05 FLIP 0. 75 – casp3‐act – 0. 75 PTEN 0. 41 0. 41 a Cell line‐specific receptor or pathway activity of proteins according to the experimentally determined differences in response behavior (apoptosis, proliferation, RTK, and western blot data); all other proteins were simulated with no specific activation. (−) Treatment activation at stage 0; (+) treatment activation at stage 1. b For the simulation of the AMPK activator AICAR in A549 and the HIF1A inhibitor PX‐478 in H441, we used the cell line‐specific activity from the untreated cells of the 17AAG treatment (Table 3 C); all other proteins were simulated with no specific activation. (−) Treatment activation at stage 0; (+) treatment activation at stage 1. c Constant activation, as there is a KRAS mutation in these cell lines. John Wiley & Sons, Ltd Figure 5 Cell line‐specific in silico simulations for 17 AAG treatment according to data from the 3D system. Simulations of the 17 AAG treatment reflect the in vitro data. Coloring of the curves is according to the network node colors shared for all three cell lines shown in Fig. 4 B. Cell line‐specific pathway differences included are given in Table 2. Top: Simulation of the 17 AAG treatment in HCC 827 cells (right, red curve at full activation) results in slightly induced apoptosis (gray curve at 0. 2) and unchanged proliferation (black curve), as compared to untreated cells (left, red curve at 0. 0, no treatment). Middle: The in silico simulation of the 17 AAG treatment for A549 shows only low apoptosis induction (0. 2); we see no therapeutic effect on proliferation (black curve, dots) compared to untreated cells. However, HSP 60 (black curve, squares) is induced after 17 AAG treatment, similar to the in vitro data. Bottom: In H441 cells, apoptosis is not elevated over time and no effect on proliferation can be obtained. p53 (pink curve) is induced after 17 AAG treatment and correlates with the in vitro data. To reveal – in a systemic approach – further relevant cell line‐specific drug targets in KRAS ‐mutated conditions in the 3D system, we reconstructed two larger in silico networks for A549 and H441 cells. Therefore, we searched in Human Protein Reference Database (HPRD; Table S1) the interacting neighbors of the nine upon 17AAG treatment between A549 and H441 differentially regulated proteins (DRPs) (Table 1 B, Fig. 6 ) and identified individual promising drug targets by mapping these to cell‐specific mutations in COSMIC (Catalogue Of Somatic Mutations In Cancer) generating thereby two cell‐specific networks for A549 and H441 cells. From networks analyses, we expanded the in this study created topology from Fig. 4 B (marked with olive gray background in Fig. 7 A, C) further with gain or loss of function mutations and other important factors by adding activated or inhibited nodes. In subsequent simulations we could predict optimal drug targets in a specific mutational background of KRAS ‐mutated tumors. In Table 3 C, input into topologies and subsequent simulations in Fig. 7 are given. Matching drugs were suggested by screening of our DrumPID (Drug‐minded Protein Interaction Database) (Kunz et al. , 2016 ) screening tool for available target‐specific test substances. Figure 6 KRAS signature development and individual target predictions by generation of HPRD networks. We generated a network around KRAS, according to the experimentally validated DRPs between both KRAS‐mutated cell lines (H441, A549) in the 3D system (Table 1 B; 17 AAG treatment), and included their direct protein interaction partners using the genomewide HPRD. The resulting larger KRAS interaction network includes 556 proteins (= nodes) and 680 protein–protein interactions (= edges), around nine strongly DRPs ( EGFR, ErbB2, ErbB3, MET, FGFR 3, c‐Ret, VEGFR 2, p53, and HSP 60). (A) A Venn diagram compares cell line‐specific mutations. Mapping of cell line‐specific protein mutations (573 for H441 (blue) and 361 for A549 (yellow) from the COSMIC database) against the 556 proteins from the network around KRAS results in 18 H441‐specific mutations and in nine A549‐specific mutations which were included in each cell line‐specific in silico topology to yield the network. Details are given in the Supporting information, and key network differences are shown in B and C. (B) A549‐specific network: represents neighbor proteins that we could target if we consider the experimental data and directly interacting protein neighbors (from HPRD ; functional clusters in Fig. S5A ). As drug targets do not appear for these small modules from key signaling proteins, we considered experimental derived proteins (red) with all first‐degree neighbors, HSP 90 (orange rectangle), and additionally direct neighbors to cell line‐specific mutations (in yellow, suspected ‘driver mutations’). Direct neighbor proteins are labeled in lavender, in cyan are neighbors from neighbors, which are also mutated. The black square ( AMPK, interactor of p53 and LKB 1) indicates a promising drug target (screening procedure given in Box S1 ). (C) H441‐specific network: shows neighbor proteins that we could target, if we consider the experimental data, HSP 90 (orange rectangle) and directly interacting protein neighbors from HPRD (functional clusters in Fig. S5B ). Directly interacting neighbors are shown (lavender, labeling binary interactions). As drug targets do not appear for these small modules from key signaling proteins, we considered all experimental determined nodes (red) with all first‐degree neighbors integrating cell line‐specific H441 mutations (in blue, suspected ‘driver mutations’; EGFR and p53 labeled in red with blue circles as they are array nodes and mutated). Protein interactors according to HPRD are labeled in lavender; in cyan are neighbors from driver mutations, also showing a mutation in H441. The square ( HIF 1A) indicates a promising drug target (screening procedure given in Box S1 ). Figure 7 In silico topologies and simulations of AMPK and HIF 1A treatment. Cell‐specific network extensions according to the experimental data (Table 2 ) are mapped into the shared topology (bold nodes from basic topology from Stratmann et al. ( 2014 ), olive shade for topology nodes from Fig. 4 B). Furthermore, AMPK as a relevant target for A549 (network in (A)) and HIF 1A as a target for H441 (topology in (C)) are included (nodes equivalent to the 17 AAG treatment are deposited in olive, protein node colors are the same as in the simulation curves). Both protein targets were integrated with their direct interacting protein neighbors in the cell‐specific networks to mirror in silico individual therapy. In (B) and (D), the cell‐specific topologies are next simulated dynamically, and selected trajectories of protein node activities were plotted, showing the effects of the potential drug candidate AICAR as an AMPK activator for A549 (B), and the HIF 1A inhibitor PX ‐478 for H441 in (D) to illustrate the in silico screen of different drugs in the two cell line‐specific topologies. (B) Simulation of AMPK activation in A549 cells (right, red curve at stage 1) results in higher apoptosis (pink curve) and reduced proliferation (salmon curve), as compared to untreated cells (left, red curve at 0. 0, no activation). (D) The in silico simulation of the HIF 1A inhibition for H441 (right, olive curve at full activation) shows higher apoptosis (black curve) and reduced proliferation (salmon curve), as compared to untreated cells (left, olive curve at 0. 0, no activation). 2. 2 Tissue‐engineered lung tumor models resemble tumor specimens Firstly, we looked at molecular markers for tumors and tissue differentiation and their variances. We observed homogenous E‐cadherin/β‐catenin localization in fluorescence staining of 3D models as well as tumor specimens, whereas 2D models showed high variation ranging from strongly stained to completely negative tumor cells (Fig. 1 A, B). By comparison to 2D models, we demonstrated with Ki67‐staining the reduction in proliferation rate in 3D models to levels that correlate to lung adenocarcinoma samples (Fig. 1 C). 2. 3 Enhanced biomarker‐dependent drug response to EGFR inhibition in the 3D model As a test for clinically applied biomarker‐guided anti‐EGFR therapy, we compared A549 and H441 (EGFR wild‐type) with HCC827 cells (activating EGFR mutation). Responses upon 3 days of Gef treatment were biomarker‐dependent, as represented by hematoxylin and eosin staining (HE staining) in 3D models (Fig. 2 A), proliferation reduction (Fig. 2 B), and apoptosis induction (Fig. 2 C). Although the proliferation in 3D conditions was reduced to in vivo like rates (Stratmann et al. , 2014 ), treatment with Gef decreased the proliferation further by about 80%, as also observed in 2D. However, biomarker‐related apoptosis induction upon Gef treatment in HCC827 was significantly enhanced in 3D conditions (about 3. 5 to 6‐fold increase), compared to 2D conditions (about 2. 5‐fold increase), which suggests better specificity for the 3D system. Signaling analyses by receptor tyrosine kinase (RTK) and phospho‐kinase (PK) array experiments are provided for the 2D and 3D systems (Fig. S1, Table 1 A). Complementing these biomarker‐dependent drug responses to Gef, we set up an in silico network of key pathways for the proliferative and apoptotic response, to model the observed in vitro responses of each of the three cell lines. An in silico topology was previously developed for HCC827 and A549 (Stratmann et al. , 2014 ). This was extended by those cell line‐specific proteins and pathways (Fig. S2A ; previous network proteins are in bold and with olive background) which showed signaling changes upon the Gef treatment in the experiments (Table 2 ). For the newly investigated H441 cell line, we used the A549 in silico topology as basis. Specifically, we integrated the MET signal transduction cascade for HCC827 and H441. We then applied the squad software to simulate the Gef treatment responses for all cell lines, using initial node stimulations based on the mutational background and the experimental results on protein phosphorylation (prestimulation in Table 3 A; method in the Supporting information). Our simulation of HCC827 with Gef treatment (Fig. 2 D) compared to untreated cells demonstrates, as for the in vitro results, reduced proliferation and higher apoptosis over time. Results of the simulations for A549 and H441 are represented in Fig. S2B. 2. 4 Chemoresistance against HSP90 inhibition in 3D models align to clinical observations 2D models and animal experiments predict HSP90 inhibitor efficiency in KRAS ‐mutated tumors (Acquaviva et al. , 2012 ; Sos et al. , 2009 ). As known from 2D in vitro screens, HCC827, A549, and H441 exhibit different sensitivities to the HSP90 inhibitor 17AAG (Ciocca and Calderwood, 2005 ; Sos et al. , 2009 ). We observed that about 50% of the H441 cells died from 0. 25 μ m 17AAG, which decreased the viability of A549 to 5% and of HCC827 to 35%, as shown by the cell viability assay CellTiter‐Glo ® in 2D (Fig. 3 A). However, the failure of HSP90 inhibitor treatment in a clinical setting of KRAS ‐mutated tumors was reflected in 3D tissue cultures. From HE staining of 3D tumor models, after three days of 0. 25 μ m 17AAG treatment only slight effects were visible in A549 and HCC827, whereas H441 cells were completely unresponsive (Fig. 3 B). Proliferation analysis of the 2D systems predicts KRAS ‐mutated cells to be more responsive to 17AAG than KRAS wild‐type HCC827 cells. This is contrasted by only weak changes between both cell types in 3D tissue culture (Fig. 3 C). A strong apoptotic response upon 17AAG is only observed in KRAS ‐mutated A549 cells in 2D (4 to 6‐fold) but not in 3D models (1 to 2‐fold) (Fig. 3 D). All phosphorylation data from arrays and western blot experiments are summarized and compared in 2D and 3D models in Table 1 for Gef (A) and 17AAG (B) treatment. Protein nodes for in silico topology which were applied later are given in Table 2. 2. 5 Differences in signaling between 2D and 3D conditions upon HSP90 inhibitor treatment as a basis for in silico analyses Signaling responses upon application of the HSP90 inhibitor 17AAG were analyzed by comparing 2D and 3D conditions. Protein activation as observed by RTK arrays was confirmed by western blot (Fig. 4 A) and quantified (Fig. S3B ). Data indicated an inhibition of the EGFR and of MET in HCC827 and in H441 in 2D as well as 3D conditions. In western blot analysis, inhibition of MET was weaker in 3D than in 2D cultures in both cell lines. Interestingly, p53 (S46) was activated in HCC827 with 17AAG treatment in 2D, but stayed constant in 3D conditions. Vice versa, in H441 p53 was activated only in 3D conditions and remained unchanged in the 2D culture. Furthermore, HSP60 was clearly upregulated only in A549 cells under 3D conditions upon 17AAG application. Regulated proteins identified in 3D conditions upon 17AAG treatment include EGFR, ErbB2, ErbB3, MET, c‐Ret, VEGFR2, FGFR3, p53, and HSP60 (Table 1 B, Figs 4 A, and S3A ). Similar to the Gef treatment, we extended our in silico network and topology adding these experimentally measured cell‐specific proteins (Fig. 4 B). Particularly, we included for mirroring 17AAG treatment effects – next to the MET protein – ErbB2, ErbB3, and c‐RET cascade in HCC827 and H441, and also in all three cell line‐specific in silico topologies p53, HSP60, HIF1A, and HSP90, as part of the 17AAG treatment cascade (Table 2 ). For H441, we included further VEGFR2 and FGFR3, as they were downregulated in the arrays in the 3D model upon treatment with 17AAG, in contrast to the other two cell lines (Fig. S3A, Table 1 B). We show only responses for key proteins of all three cells, but we simulated the complete network responses looking at all proteins of the topology. Important aspects of the 3D tissue model upon 17AAG application (red curve at 1) are reflected by in silico simulations (Fig. 5 ): (a) In HCC827 (top), cell proliferation is unchanged and apoptosis is slightly induced compared to the untreated control, (b) in A549 (middle), proliferation is regarded as unchanged and apoptosis is only slightly induced in 3D conditions, and (c) in H441 (bottom), proliferation is unchanged and apoptosis is not induced. Notably, HSP60 is exclusively induced in A549, whereas p53 is upregulated only in H441. Moreover, based on the in silico topology connectivity, in our in silico simulation we found that beside p53 HIF1A is also upregulated in H441. For comparison, the in silico simulations can also be modified to appropriately reflect results of 2D culture. To illustrate this, we focused on the A549 and H441 cell lines and applied the same topology as for 3D, but adjusted activation levels for Gef and 17AAG treatment according to the 2D in vitro conditions (Fig. 4 B and S2A ; Table 1 ; Box S2 ): Essentially, we elevated the value of Raf to simulate higher basic proliferation in 2D, and furthermore, we changed FLIP for the higher apoptotic response in 2D upon 17AAG as this is reported to be important for higher apoptotic resistance when cells grow on collagen (Philippi et al. , 2009 ). Whereas Gef treatment simulation resulted in both KRAS ‐mutated cells in no change of proliferation and apoptosis over time (Fig. S6 ), HSP90 inhibition simulation of 2D conditions revealed in contrast to 3D a lower proliferation and an induced apoptosis over time in A549 (Fig. S7 ). However, the established tool allows us now to test and screen in silico in a systems perspective for tailored therapies according to the cell line‐specific mutational profile and tumor drivers, as detailed in the following section. 2. 6 Generation of in silico protein–protein interaction networks for cell‐specific drug target predictions in KRAS ‐mutated cells Next, as we observed signaling differences between the KRAS ‐mutated A549 and H441 cell lines, we sought to identify a KRAS complementing signature of further potential biomarkers and resultant drug targets for each cell line. For this purpose, we combined experimental data and the cell line‐specific mutational backgrounds with integrated systems biology analysis (Kunz et al. , 2017 ; Naseem et al. , 2014 ), considering direct interacting proteins and available drugs to modulate this extended network. We generated a network around KRAS by considering the DRPs of both KRAS ‐mutated cell lines (A549 and H441) upon 17AAG treatment (Table 1 B) in the 3D system and included their direct interacting proteins according to the genomewide HPRD. The resulting KRAS interaction network includes 556 proteins (= nodes) and 680 protein–protein interactions (= edges) around the nine experimentally DRPs (EGFR, ErbB2, ErbB3, MET, FGFR3, c‐Ret, VEGFR2, p53, and HSP60; Fig. S4A ). Comparing all cell line‐specific mutations known from the genomewide COSMIC database (573 for H441, blue circle, and 361 for A549, yellow circle; Table S2; Fig. 6 A) with this KRAS interaction network, we could match 18 H441‐specific mutations and nine A549‐specific mutations, as parts of our KRAS interaction network (Table S3 ). The two reconstructed cell‐specific KRAS interaction networks for A549 and H441 included these specific mutations, HSP90 as a target of 17AAG, and their direct interaction partners from HPRD (Fig. S4B–D ), which were then analyzed for functional clusters. In the A549‐specific network (322 proteins and 371 protein‐protein interactions, Fig. S4B ; extended network with 795 nodes and 1034 interactions in Fig. S4C ), we found two functional protein clusters with a strong network effect (so‐called hubs) around proteins VEGFR2, MET (experimental measurements) and CBL (mutated), and p53 (experimental) and ARID3A (mutated; Fig. S5A ). Similarly, for the H441‐specific network (903 proteins and 1119 protein–protein interactions; Fig. S4D ), we found two clusters around the proteins PRKACA (mutated) and p53 (experimental and mutated) as well as HSP90AA1, ACTA, and HIF1A (mutated; Fig. S5B ). We compared potential targets in the two cell line‐specific KRAS networks, in terms of their distance and usefulness to modulate cell‐specific signaling cascades. This yielded a highly connected network between interesting tumor drivers (Fig. 6 B for A549, driver mutations in yellow; Fig. 6 C for H441, driver mutations in blue), and cell‐specific biomarker signatures (Table 4, Box S1 ). Other cell‐specific mutations close to the central cascade are indicated by cyan (neighbors of neighbors), and unmutated interactors in lavender circles for A549 and H441, respectively. Regarding ranking of drug targets for potential clinical application, we considered proteins and connections and assigned the priority to direct neighbors, if they could be targeted easily by existing medical drugs, for example, AMPK for A549 and HIF1A for H441. All targets are ranked in Box S1. This drug‐search strategy was made possible by applying our DrumPID (Kunz et al. , 2016 ). Table 4 Overview of potential predictive markers and new therapeutic targets for the KRAS ‐mutated cell lines a John Wiley & Sons, Ltd 2. 7 AMPK and HIF1A targeting in cell‐specific in silico simulations for A549 and H441 cells Subsequently, we investigated in silico the potential therapeutic effect of AMPK as a relevant target for A549 and HIF1A as a target for H441. For this, we integrated the LKB1 cascade for A549 into the cell‐specific in silico topology that simulated the 17AAG therapy (Fig. 7 A), and further considered the connectivity of HIF1A in H441 (Fig. 7 C). Proteins in the network that correspond to the basic topology of Stratmann et al. ( 2014 ) are bold rimmed and proteins that match to the topology from Fig. 4 B have an olive background (Fig. 7 A, C). From our drug screening (Box S1 ), we identified 5‐aminoimidazole‐4‐carboxamide ribonucleotide (AICAR) in A549 as the potential activator of AMPK that is directly modulated by its interactor LKB1, which is specifically mutated in A549 (synonym: STK11; see DrumPID pathway ko04152). Similarly, we found PX‐478 as a selective HIF1A inhibitor for the H441 cell line. Based on this, we simulated their potential therapeutic effect by applying the SQUAD algorithm (Fig. 7 B, D; prestimulations in Table 3 C). For in silico simulations of AMPK activation and inhibition of HIF1A, we found induced apoptosis and reduced proliferation in each cell line over time as a desirable drug effect. 3 Discussion 3. 1 Motivation Personalized treatment strategies have to cope with highly redundant tumor pathways resulting in resistance, whereas combination therapies often show severe toxic side effects (Tannock and Hickman, 2016 ). Therefore, it is necessary to reconsider the design of clinical studies with targeted anticancer approaches. It is critical to understand the underlying dependencies in signaling networks, and to provide tools for exploiting now frequently available sequencing data from patients. Moreover, in the field of oncology, the success rate of preclinical testing is at under 5%, generating enormous financial costs (Bhattacharjee, 2012 ). Even though animal models predict toxicity quite convincingly, they tend to fail in efficacy testing (Greaves et al. , 2004 ; Kubinyi, 2003 ). In particular, for signaling analyses, mice are not adequate models due to inappropriate ligands to some centrally connected human receptors, such as MET (Francone et al. , 2007 ). Next to ethical concerns, these aspects underline the urgent need to develop novel human tumor test systems. Here, we introduce a new concept of in vitro tissue tumor models and in silico analyses to design and test individual biomarker profiles and intervention strategies. This prepares the floor for patient‐tailored clinical studies, required for personalized cancer medicine (Tannock and Hickman, 2016 ). As our main aim is to develop a powerful tool that can be implemented into the clinic by analyzing the patient's sequence data, we investigate as proof of concept in this work three individual lung cancer cell lines with known genome sequence information. However, here we show only exemplary data and not a large validation series. We are aware that exact quantitative estimates require a higher number of experiments and more cell lines with similar driver mutations. Exploring its clinical implementation, currently our in silico tool advises on a case‐by‐case basis the molecular tumor board in our comprehensive cancer center. In particular it supports to identify alternative protein targets when resistance to treatment occurs. Our human 3D tumor model generated by tissue engineering technologies should reduce preclinical failure, as it reflects tumor characteristics better and shows higher predictive accuracy than conventional 2D cultures: it retains the tissue architecture, extracellular matrix components and structures of the basement membrane as unique features for cellular interactions. These are important modifiers of cellular responses (Linke et al. , 2007 ; Philippi et al. , 2009 ; Schanz et al. , 2010 ). In detail, we observed (a) more homogenous staining of E‐cadherin/β‐catenin and lower proliferation rates according to tumor specimens, (b) a biomarker‐dependent apoptosis induction and proliferation reduction by the EGFR inhibitor Gef, and (c) in contrast to other preclinical findings, a reduced response upon HSP90 inhibitor treatment in KRAS ‐mutated tumor cells, which matches observations from clinical studies. Thus, we believe that our in vitro model resolves interpathway dependencies more reliably than 2D or animal models. Individual KRAS in silico networks were established by integrating relevant proteins from in vitro experiments and their interaction partners from HPRD. Our simulations start from a general in silico network for lung cancer, which is refined here to reveal the most relevant protein clusters. By matching cell line‐specific mutations from the COSMIC database, we derived individual drug targets and by screening our custom‐made, protein–drug interaction database DrumPID appropriate drugs (Kunz et al. , 2016 ). 3. 2 HSP90 inhibition in KRAS ‐mutated tumors and correlation of our 3D tissue models and other preclinical models to clinical findings In a previous study of a lung cancer model, we were able to demonstrate a stronger apoptosis induction in the 3D model by Gef, compared to conventional 2D culture (Stratmann et al. , 2014 ). After setting up standard operating procedures (Göttlich et al. , 2016 ), we could predict the clinical failure of HSP90 inhibitor treatment in the context of KRAS mutation, in contrast to other in vitro and in vivo models (Acquaviva et al. , 2012 ; Sos et al. , 2009 ). Heat shock proteins have gained attention in recent years as therapeutic tools, as they are involved in tumor cell proliferation, invasion, and cell death. Their high expression was observed in several cancer entities in clinical settings (Ciocca and Calderwood, 2005 ). Specifically, HSP90 belongs to a family of chaperons important for the function of relevant oncogenic drivers in lung adenocarcinomas. From a genomewide screening of 84 cell lines, KRAS mutation was identified to confer sensitivity to HSP90 inhibition that could also be verified in murine models (Sos et al. , 2009 ). In this screening, the geldanamycin derivatives 17AAG, and 17‐dimethylaminoethylamino‐17‐demethoxygeldanamycin in mice experiments were applied for HSP90 inhibition. However, geldanamycin and its derivates turned out to have safety and pharmacological limitations (Jhaveri and Modi, 2015 ). Another in vitro study showed the effectiveness of HSP90 inhibition in several KRAS ‐mutated non‐small cell lung cancer (NSCLC) lines by ganetespib – a non‐geldanamycin analog with less toxic side effects (Acquaviva et al. , 2012 ). However, single agent HSP90 inhibition by ganetespib failed in NSCLC patients with KRAS ‐mutated tumors. Combination therapy trials with docetaxel (GALAXY 1 and 2) led to better outcomes in patients with adenocarcinomas, than docetaxel single agent therapy, but not in the subgroup of KRAS ‐mutated tumors (Bhattacharya et al. , 2015 ). Recently, ALK, ROS1, and RET kinase gene rearrangements have been suggested to predict efficacy by targeting HSP90 (Rothschild, 2015 ; Sang et al. , 2013 ; Socinski et al. , 2013 ). 3. 3 In silico simulations of cell responses and development of a predictive KRAS signature In this study, in silico analyses for KRAS signature development are executed in three steps: Set up of cell‐specific in silico topology with logical Boolean connectivity (software tool celldesigner ; http://www. celldesigner. org ; Funahashi et al. , 2008 ) Cell‐specific dynamic in silico simulations of tumor cell responses (software tool squad ) For systematic drug‐target identification we generate larger cell‐specific protein–protein interaction networks considering neighbors of the central cascades (using data from HPRD) and cell line‐specific mutations (using data from COSMIC). For drug suggestions we apply the database tool DrumPID. In detail, we explain here the three above mentioned distinct types of in silico analyses: With the term ‘ in silico topology’, we considered a previously published knowledge‐based network which focuses mainly on kinase cascades (Stratmann et al. , 2014 ) and integrated here cell‐specific differences as additional nodes (proteins) derived from experimental data to specifically mirror the effects of Gef treatment and HSP90 inhibition. For this, we measured by phospho‐arrays and western blot signaling changes as drug responses as well as differences in proliferation and apoptosis in the different cell lines in 2D and 3D conditions. Missing parts of the cascade or modulatory crosstalk are filled in according to expert knowledge and public databases. These represent only the key parts of the signaling cascades. We used the tool ‘CellDesigner’ to set up the in silico topologies and to bring them into a machine‐readable format as done before for other cell types (Schlatter et al. , 2012 ). ‘ in silico simulations’ with the SQUAD tool predict the systemic response of a tumor cell upon a specific treatment which depends on the tumor cell topology and the activation/inactivation of its integrated nodes. As input the activation level of certain nodes can be set between zero as an inhibitory effect (inactivation) and one as an activating effect (activation). Furthermore, mutations can be integrated that stay independent from upstream signaling events at a certain value in case of gain or loss of function mutations. Also differences in 2D and 3D conditions can be simulated by adjusting the nodes’ values to experimentally measured levels, that is, phosphorylation determined by western blot. Some of the nodes summarize also global cellular responses, for example, ‘stress’. Importantly, also the drug responses proliferation and apoptosis are integrated in the topology as nodes. Values of the other nodes must be adjusted until the level of proliferation and apoptosis comply with the in vitro observations. Traditionally, differential equations for detailed kinetic modeling look at biological responses (Di Cara et al. , 2007 ; Dwivedi et al. , 2015 ; Robubi et al. , 2005 ). However, this requires then detailed kinetic information on individual kinases. This is not necessary in our approach, as the squad modeling software interpolates automatically exponential functions between our protein network nodes fitting signal transmission and logical connectivity (Di Cara et al. , 2007 ). We previously applied this combination to study cancer (Göttlich et al. , 2016 ; Stratmann et al. , 2014 ), infection biology (Audretsch et al. , 2013 ; Naseem et al. , 2012 ), and different tissues (Brietz et al. , 2016 ; Czakai et al. , 2017 ; Philippi et al. , 2009 ). For drug targeting, we looked systematically at larger protein–protein interaction networks; in particular, we collected all neighbors of upon 17AAG treatment between DRPs of both KRAS ‐mutated cell lines. To this network, we matched cell‐specific mutations from the COSMIC database. These larger networks we term here ‘ in silico networks’. The cell‐specific networks were then scrutinized to identify most promising treatment targets considering their relation to highly connected proteins in the network that are called ‘hubs’. A robust drug prediction algorithm collates information from several large‐scale databanks including chemical information according to Simplified Molecular Input Line Entry Specification (SMILES) notation and basic drug pharmacokinetics ADME (absorption, distribution, metabolism, excretion) rules (DrumPID, Kunz et al. , 2016 ). We used our reconstructed cell‐specific networks (Fig. 6 B, C) and screened which drugs according to DrumPID (Kunz et al. , 2016 ) influence apoptosis and proliferation in a cell‐specific manner. Targets are ranked by the effect strength, closeness to central cascades and druggability (Box S1 ). Subsequently, we simulated the potential therapeutic effect on apoptosis and proliferation focusing on AICAR (AMPK activator) and PX‐478 (HIF1A inhibitor) as top candidates and integrated their specific connectivity to the central cascades in the in silico topology. However, other drug target candidates can also be simulated, but for each simulation the individual targets and side targets of the drug has to be considered. Further testing of predictions is required to confirm suggested targets regarding clinical relevance. So far, neither HCC827, nor A549, nor H441 lung cancer cell lines have been analyzed by such a comprehensive in silico approach. 3. 4 Experimentally measured differences between the 2D and 3D system and in silico analyses Besides a higher chemoresistance in the case of HSP90 inhibitor treatment, we observed in 3D lower reduction of MET upon 17AAG treatment and an inverse regulation of p53, when compared to 2D conditions. Next to semiquantitatively evaluated western blot experiments, we present data from two phospho‐array screens (RTK, PK) as a starting point for further analyses (Figs S1 and S3A ). Importantly, in our 3D experiments we observed in contrast to 2D conditions upon 17AAG treatment an upregulation of HSP60 exclusively in A549, and an activation of p53 only in H441, which we were able to achieve also in our in silico simulations. However, the literature reports HSP60 inhibition by HSP90 and p53 inactivation by HSP60 (Ghosh et al. , 2008 ), which would explain our experimental observations in A549 and H441. The reason why HSP60 upregulation and a lack of p53 expression in A549 has a small effect on apoptosis in this setting, could be due to reduced HIF1A activation upon 17AAG treatment, as predicted by our in silico simulation. This reduced activation could stem from the inhibition of HSP90. HIF1A is not completely silenced in the simulation, due to its connection to LKB1 via mTOR in the in silico topology, as according to the COSMIC database, in A549 LKB1 carries a loss of function mutation. Furthermore, our in silico topology illustrates that this LKB1 mutation should lead to reduced AMPK activation and, thereby, also reflects the nonproliferative effect of 17AAG treatment via the mTOR signaling pathway. On the other hand, apoptosis induction is also blocked in H441. Induction of p53 upon inhibition of HSP90 should have no apoptotic effect due to the loss of function mutation of p53 identified in the COSMIC database. Furthermore, in our simulation we can see that HIF1A is still activated in H441 following HSP90 inhibition. This could be due to a mutation inside this gene that leads to an inhibitory effect on apoptosis and favors proliferation. Box S2 shows that we can use the same topology to simulate 2D results in silico, but differences in protein activation have to be taken into account to appropriately simulate the stronger apoptotic as well as proliferative responses upon 17AAG treatment observed in 2D cultures. As we could correctly simulate the observed responses upon Gef and 17AAG treatment for 2D and 3D conditions we could support that nodes in our topology are so far connected correctly. The in silico screen can reveal new dependencies, as high quality databases consider cancer‐subtype‐specific mutations and their interacting proteins along with all available drugs to directly attack the mutated protein or one of its neighbor. AICAR and PX‐478 are given as attractive examples (top ranked; see Supporting information) and their therapeutic effect on apoptosis and proliferation is simulated. However, other drugs can also be used by integrating other drug target candidates by considering its individual targets and side targets. Subsequently, the targets and drugs can be integrated in the in silico topology by considering its specific connectivity to the central cascades and further in silico simulated with SQUAD. 3. 5 Exemplified target and drug candidate prediction for A549 cells From the newly established KRAS network, based on proteins that exhibit changes in signaling between A549 and H441 in 3D conditions and cell‐specific mutations, the for A549 unique mutation LKB1 stands out. Our drug–protein interaction database DrumPID (Kunz et al. , 2016 ) identifies drugs that modulate the query protein directly, or one of its directly interacting neighbor proteins. This database tool identifies AMPK in our analyses as a potential drug target in A549 cells which can be modulated by the drug AICAR. AMPK protein is a direct interaction partner of LKB1 (Fig. 6 B) (Fay et al. , 2009 ; Rattan et al. , 2005 ; Tang et al. , 2011 ). AMPK activation using the drug AICAR, an analog of AMP, leads to tumor growth arrest in our in silico simulation for A549 cells (Fig. 7 A, B; nodes from first topology are olive‐shaded in 7 A). Moreover, AICAR shows promising results in the clinical phase 1/2 for chronic lymphatic leukemia (Van Den Neste et al. , 2013 ). In addition, the approved anticancer agent pemetrexed is known to indirectly activate AMPK by the accumulation of ZMP in LKB1‐null lung cancer (Rothbart et al. , 2010 ). 3. 6 Exemplified target and drug candidate prediction for H441 cells Regarding H441 cells, we also screened the H441 protein interaction network around the KRAS signature for potential drugs targeting either the protein or its direct neighbor. HIF1A was the highest‐ranked target (Box S1 ), as it is altered in H441 cells according to COSMIC data and is involved in a signaling loop (Greijer and van der Wall, 2004 ) (Fig. 6 C). Inhibition of HIF1A using PX‐478 shows an antitumor effect in our in silico simulations (Fig. 7 C, D; nodes from first topology are olive‐shaded in 7C). Studies demonstrated that inhibition of HIF1A shows promising therapeutic effects in human xenograft models (Welsh et al. , 2004 ). 3. 7 Application of the combined in vitro / in silico tool For clinical application, patient tumors have to be sequenced first, or at least tested by PCR or microarrays, to confirm that the driver mutation profile matches those in our cell lines. Notably, primary tumor cell culture is still challenging and has to be optimized for its utilization in routine personalized approaches. 4 Conclusion Predictive gene signatures were identified in a combined, tissue‐engineered, 3D lung tumor model with improved clinical correlation and a Boolean in silico approach that integrated measured cell‐specific differences in drug responses. We established cell line‐specific networks that depend on individual mutation patterns. This enabled better understanding of the interdependencies between single signaling cascades to prevent treatment resistance. Exemplified by the KRAS ‐mutated cell lines A549 and H441, we demonstrated how our analysis tool could lead to individual signature development, based on in vitro / in silico investigations on signaling, interaction partners from the HPRD, and sequence data from COSMIC. The limited number of direct interference points with the proliferative and/or apoptosis signaling cascade suggests and ranks best cell line‐specific targets, implying future therapies according to NGS data, tailored to the individual cancer mutation profile. Translated into clinical application, our lung cancer cell line‐specific examples suggest for patient stratification to determine not only the KRAS mutation status, but also to test for LKB1, p53, and HIF1A. Such cancer‐specific prescreening could distinguish among individual mutational subgroups to improve patient stratification and the design of clinical studies. 5 Materials and methods 5. 1 Cell culture HCC827 and A549 cell lines were purchased from DSMZ (Braunschweig, Germany), H441 from ATCC (LGC Standards GmbH – Germany Office, Wesel, Germany). A549 and H441 cells were cultured in RPMI + 10% FBS, HCC827 cells in RPMI + 20% FBS. Cells were monitored for pathogen infections at regular intervals. For a 2D culture, cells were either grown on glass coverslips in well plates until they had reached a confluency of about 70% or were cultured for 5 days in 12‐well plates or 6‐cm petri dishes. For a 3D culture, 1 × 10 5 tumor cells were grown for 14 days on the SISmuc (see Section 5. 3 ) that was fixed between two metal rings, as described in the literature (Göttlich et al. , 2016 ; Moll et al. , 2013 ; Stratmann et al. , 2014 ). Both 2D and 3D cultures were performed under standard conditions (37 °C, 5% CO 2 ). 5. 2 Treatment with Gef and 17AAG After 1 day in a 2D and 11 days in a 3D culture, cells were treated with either 1 μ m Gef (Iressa™, AstraZeneca, Wedel, Germany; Selleckchem) or 0. 01, 0. 05, 0. 1, 0. 25, 0. 5 or 1 μ m 17AAG (17‐ N ‐allylamino‐17‐demethoxygeldanamycin, Tanespimycin; Selleckchem) for 72 h, with a medium change after the first 48 h of treatment. 5. 3 Porcine material The SISmuc consisting of porcine small intestine submucosa (SIS) and mucosa (muc) was used as a scaffold for all 3D culture experiments. It was prepared from the BioVaSc ® as described in the literature (Linke et al. , 2007 ; Schanz et al. , 2010 ). All explantations were in compliance with the German Animal Protection Laws (§4(3), supervised by the institute's animal protection officer, all animals received proper care according to the National Institute of Health standards (NIH publication no. 85e23, revised 1996)), and as approved by the institutional animal protection board. 5. 4 Human material Human lung tumor tissue was provided by the Department of Thoracic Surgery of the University Hospital of Wuerzburg (local ethics approval: 182/10, 25. 11. 2015). 5. 5 Histology and immunofluorescence Cells cultured on glass slides in 2D were fixed in 4% paraformaldehyde for 10 min, cells in a 3D culture for 2 h, and the human lung tumor tissue overnight at 4 °C. The SISmuc samples, as well as the tumor tissue, were embedded in paraffin and sectioned at 3 μm thickness for hematoxylin–eosin (HE) and immunofluorescence staining. The primary antibodies E‐cadherin (#610181; BD Transduction Laboratories, Heidelberg, Germany), β‐catenin (#ab32572; Abcam, Cambridge, UK), and Ki67 (#ab16667; Abcam) were diluted 1 : 100 and incubated overnight at 4 °C. Secondary antibodies conjugated with fluorescent dyes Alexa 555 or 647 were diluted 1 : 400 and incubated for 1 h at room temperature. Nuclei were counterstained by DAPI dissolved in a Mowiol embedding solution. Pictures were taken with a digital microscope (BZ‐9000; Keyence Deutschland GmbH, Neu‐Isenburg, Germany). 5. 6 Cell proliferation To determine the proliferation rate, cells cultured in 2D and 3D were stained against Ki67. Ten nonoverlapping images of 3D sections and five nonoverlapping images of 2D cultures were taken. Quantification of the proliferation rate was performed as described in the literature (Göttlich et al. , 2016 ). 5. 7 M30‐elisa Apoptosis was determined from supernatants taken from untreated and treated tumor models during the last 4 days of the culture. M30 CytoDeath™ ELISA (Peviva) was performed according to the manufacturer's instructions. All samples were measured in duplicates. 5. 8 Western blot and phospho‐RTK and PK arrays Cells were lysed in modified RIPA buffer (137 m m NaCl, 50 m m NaF, 20 m m Tris/HCl pH 8. 0, 2 m m EDTA, 10% (v/v) glycerol, 1% (v/v) NP‐40, 0. 5% (w/v) DCA, 0. 1% (w/v) SDS, 1 m m Na 3 VO 4, and 1× protease inhibitor cocktail (Sigma‐Aldrich, Darmstadt, Germany)), or in the provided lysis buffers of the respective array kit. For western blot analysis, protein samples (27 μg per lane) were separated electrophoretically in a 10% SDS/gel and blotted on a 0. 2‐μm nitrocellulose membrane (Whatman, Fisher Scientific GmbH, Schwerte, Germany). The primary antibodies pEGFR (#ab32430; Abcam), pMet (#3077; Cell Signaling Technology, Frankfurt a. Main, Germany), phospho‐p53 (S46) (#2521; Cell Signaling Technology), HSP60 (#ab46798; Abcam), and β‐actin (#3700; Cell Signaling Technology) were incubated in NFDM or a 1% BSA overnight at 4 °C. Secondary anti‐mouse or anti‐rabbit IgG antibodies conjugated to horseradish peroxidase (#JAC‐111035144 or #JAC‐115035146; Jackson ImmunoResearch, Cambridgeshire, UK) were incubated for 1 h at room temperature. Bands were visualized using the Pierce ECL Western Blotting kit (Thermo Scientific, Breda, Netherlands). Phospho‐RTK and PK arrays were performed according to the manufacturer's instructions. Western blot and array membranes were imaged at the imaging station FluorChem Q (Biozym Scientific, Hessisch Oldendorf, Germany). Gray values were determined with the related image acquisition and analysis software alphaview (version 3. 2. 2. 0; Proteinsimple, San Jose, CA, USA). 5. 9 Statistical analysis of the experimental data The nonparametric Kruskal–Wallis test and post hoc Wilcoxon rank‐sum test were used for statistical analysis of proliferation and apoptosis results. P < 0. 05 was considered as significant. Statistical analysis was carried out with the open‐source software r (The Comprehensive R Archive Network). 5. 10 Bioinformatics analysis 5. 10. 1 Network analysis Bioinformatics analyses combined cell culture array and western blot data for the Gef and 17AAG treatment in the 3D system with information from databases to build up an individual network for each cell line. We extended the original in silico topology (Stratmann et al. , 2014 ) by integrating proteins listed in Table 2. 5. 10. 2 Dynamic simulation For the 3D in vitro system, we simulated the Gef and 17AAG treatment using the squad software (Di Cara et al. , 2007 ), by taking the pathway activity differences into account (Table 2 ) while running the simulation (prestimulations in Table 3 ). For the 2D system we focused on the A549 and H441 cell lines (prestimulations in Box S2 ). SQUAD represents the network topology (activation, inhibition) using logical Boolean operator (AND, OR, NOT) and interpolates them by applying mathematical e‐functions. The resulting network effects are visualized in a graph as changes of state over an arbitrary time, allowing in silico simulations of different network scenarios. Simulation protocols were written using the SQUAD function ‘perturbator’ (prestimulation option in the simulation menu of the software), in which the value for the drug Gef and 17AAG were set to an initial state of 0 and 1 (reflecting either no treatment or standard treatment, respectively), and experimental nodes and mutations were adjusted (Table 3 and Box S2 ). All parameters for the proteins (‘nodes’) in the network without experimental or mutational regulation were set as an active node pulse (state = 0 and time = 0) that changes, depending on interconnectivity in the cell‐specific network. 5. 10. 3 Software for visualization To set up the silico topology we used the celldesigner software tool. For visualizing the network, we used cytoscape version 2. 8. 3 (Shannon et al. , 2003 ). The cytoscape software is an open‐source platform for visualization and analysis of biological networks using several plug‐ins (Saito et al. , 2012 ; Shannon et al. , 2003 ). We analyzed the reconstructed cell line‐specific networks for functional modules (‘clusters’) using the Cytoscape plug‐in MCODE (Bader and Hogue, 2008 ). Potentially available drugs were selected using our previously developed DrumPID (Kunz et al. , 2016 ). The following methods were applied, as detailed below: An in silico signaling network is invariably a simplified view of the biological complexity. We focus here on the major cascades relevant for the output. The bioinformatics analysis combined the cell culture array and western blot data with the systems biology network analysis approaches and information from databases. The robustness of the simplified networks of central tumor cascades in cell‐specific in silico networks was also verified by considering removing and adding protein nodes at the rim of the network. This did not affect signaling responses, whereas changing central hub protein nodes strongly affected the results. Cell compartmentalization (e. g. , divergent cytosolic and mitochondrial processes), multiphosphorylation processes and complex formations were not included. This limits such approaches to semiquantitative descriptions of the sequential order, strength, and respective duration of events (Brietz et al. , 2016 ; Göttlich et al. , 2016 ). However, the simulations allow an in silico overview of the lung tumor topology and important drug responses, such as changes in individual cell line‐specific responses. 5. 10. 4 Cell line‐specific network reconstruction For establishing cell line‐specific signaling networks, we always used the same central cascades based on our previously published in silico topology (Göttlich et al. , 2016 ; Stratmann et al. , 2014 ) for HCC827 and A549 cell lines. In contrast, we extended the in silico network for the additionally introduced KRAS ‐mutated H441 cell line. 5. 10. 5 Simulation protocol For dynamic simulation, we first looked at the effects of Gef and next 17AAG treatment in the tissue in vitro system, revealed by phospho‐RTK arrays and western blots. As the model is semiquantitative, weaker or stronger biological activation has to be taken into account with values between 0 (reflects no activation) and 1 (reflects full activation) to model key input. We fitted parameters to the results obtained from experiments (data‐driven modeling) and optimized the fit in iterative cycles of new simulations and new experiments. This included prestimulations according to mutations known from their pharmacological behavior. We hence simulated the proteins as network nodes with the parameters given in Table 3 and Box S2. Additional information on the details of the bioinformatics analysis is given in the Doc. S1 and the Supporting information figures and boxes. Author contributions CG performed all the experiments in the study, supported in some by SLN. MK performed all the bioinformatics work including simulations, network analysis, and large‐scale data analysis. In several places and in particular for the initial model set up, CZ worked on the bioinformatics analysis. GD and SLN supervised the experimental work. HW gave expert analysis and advice on all experimental work. HW, GD, and TD supervised CG, and TD and MK supervised CZ. TD and GD supervised the bioinformatics analysis, analyzed all data, and made suggestions for new experiments, simulations, data comparisons, and other analyses. TD and GD drafted the manuscript together, and all authors contributed to the iterations and agreed to the final version of the manuscript. TD and GD led and guided the study. Data availability All data and simulation protocols for the study are made available with the publication (paper plus all Supporting information). Supporting information Fig. S1. Signaling is unchanged in gefitinib responsive HCC827 cells in 2D and 3D. Fig. S2. In silico model and simulation for the gefitinib treatment in A549 and H441. Fig. S3. Signaling changes in 2D and 3D after treatment of different cell lines with the HSP90 inhibitor 17AAG. Fig. S4. Biological network analyses on the KRAS‐mutated cell lines for 17AAG in the 3D system. Fig. S5. Functional cluster analyses of the cell line‐specific networks. Fig. S6. Cell line‐specific in silico simulations for gefitinib treatment in A549 and H441 according to data from the 2D system. Fig. S7. In silico simulations for 17AAG treatment in A549 and H441 according to data from the 2D system. Box S1. Ranking and comparison of all cell‐specific mutations for KRAS signature development and individual target predictions. Box S2. Cell line‐specific differences modeled in 2D. Click here for additional data file. Doc. S1. Additional information for the bioinformatics analyses Click here for additional data file. Table S1. For the generation of networks we downloaded the HPRD which contains 9620 protein nodes and 39185 protein–protein interaction edges (release 9 from April 13, 2010). Table S2. For the identification of a KRAS signature of potential markers we downloaded cell line‐specific mutations from the COSMIC database (A549: Sample Name: A549, Sample ID: 905949; H441: Sample Name: NCI‐H441, Sample ID: 908460). Table S3. Mapping of the COSMIC mutations to the KRAS‐mutated network results in 18 H441‐ and 9 A549‐specific overlapping proteins (nodes). Click here for additional data file.
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10. 1002/1878-0261. 13037
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Molecular Oncology
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Lineage‐specific mechanisms and drivers of breast cancer chemoresistance revealed by 3D biomimetic culture
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To improve the success rate of current preclinical drug trials, there is a growing need for more complex and relevant models that can help predict clinical resistance to anticancer agents. Here, we present a three‐dimensional (3D) technology, based on biomimetic collagen scaffolds, that enables the modeling of the tumor hypoxic state and the prediction of in vivo chemotherapy responses in terms of efficacy, molecular alterations, and emergence of resistance mechanisms. The human breast cancer cell lines MDA‐MB‐231 (triple negative) and MCF‐7 (luminal A) were treated with scaling doses of doxorubicin in monolayer cultures, 3D collagen scaffolds, or orthotopically transplanted murine models. Lineage‐specific resistance mechanisms were revealed by the 3D tumor model. Reduced drug uptake, increased drug efflux, and drug lysosomal confinement were observed in triple‐negative MDA‐MB‐231 cells. In luminal A MCF‐7 cells, the selection of a drug‐resistant subline from parental cells with deregulation of p53 pathways occurred. These cells were demonstrated to be insensitive to DNA damage. Transcriptome analysis was carried out to identify differentially expressed genes (DEGs) in treated cells. DEG evaluation in breast cancer patients demonstrated their potential role as predictive biomarkers. High expression of the transporter associated with antigen processing 1 ( TAP1 ) and the tumor protein p53‐inducible protein 3 ( TP53I3 ) was associated with shorter relapse in patients affected by ER + breast tumor. Likewise, the same clinical outcome was associated with high expression of the lysosomal‐associated membrane protein 1 LAMP1 in triple‐negative breast cancer. Hypoxia inhibition by resveratrol treatment was found to partially re‐sensitize cells to doxorubicin treatment. Our model might improve preclinical in vitro analysis for the translation of anticancer compounds as it provides: (a) more accurate data on drug efficacy and (b) enhanced understanding of resistance mechanisms and molecular drivers.
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Abbreviations ABC ATP‐binding cassette Ct threshold cycle DEGs differentially expressed genes ECMs extracellular matrices mTORC1 mammalian target of rapamycin complex 1 SEM scanning electron microscopy 1 Introduction Drug resistance is the major cause of treatment failure in cancer patients, ultimately leading to the death of patients with advanced‐stage cancers [ 1 ]. For decades, candidate drugs for clinical development have been tested on monolayer cultures using inert plastic supports. While this approach has benefit from a hit identification perspective, it presents some limitations: the fast‐replicating phenotype of cancer cells in this model enhances their sensitivity to antiproliferative drugs producing overestimated responses, while complex stimuli from the tumor microenvironment that are involved in multidrug resistance are lacking [ 2 ]. The association of a defective DNA repair machinery and exposure to oxidative stress affects cell genetic and epigenetic characteristics, altering key biological functions like cell metabolism and resistance profile [ 3 ]. This at least partially explains why the 90% of promising drug candidates that proved effective and safe in preclinical trials ultimately fail in the clinical setting [ 4 ]. To improve the success rate of current trials, efforts have been directed on two main areas: advances in precision medicine to introduce predictive biomarkers for the selection of patients that would benefit most from specific targeted therapies, and the development of engineered models to enhance the predictivity of preclinical tests through the recapitulation of key tumor features involved in therapy response [ 5 ]. In this regard, a plethora of three‐dimensional (3D) culture systems mimicking physical, chemical, and biological elements of the tumor microenvironment have been developed up to date [ 6, 7, 8, 9, 10 ]. These models allow to recreate and dissect: (a) the penetration and distribution of drugs in 3D structures [ 11, 12, 13, 14 ]; (b) how mechanical stimuli as the flow of extracellular fluids that generate flux of compounds and removal of metabolites impact drug efficacy [ 15 ]; (c) drug delivery strategies mediated by the targeting of extracellular matrices (ECMs) [ 16 ]; (d) the induction of resistance mechanisms by the tumor microenvironment and tumor stroma [ 17, 18 ] including immune cells [ 19, 20 ]; and (e) the preservation in native environments of primary cell phenotypes for precision medicine [ 21, 22 ]. In particular, it has been extensively demonstrated that culturing cancer cells in these engineered systems generates an enhanced understanding of resistance mechanisms that are generally difficult to study in monolayer conditions [ 23, 24, 25, 26, 27 ]. Drug resistance might be derived from intrinsic cancer cell characteristics or emerge during the tumor evolution driven by a number of different mechanisms, such as drug efflux or inactivation, target alteration, inhibition of cell death, or epithelial to mesenchymal transitions [ 28 ]. The subpopulation of drug‐resistant cells is often responsible for cancer relapse that follows the remission period after treatment. These cells persist in patients and might migrate to distant sites initiating metastasis [ 29 ]. Therefore, it is fundamental that drug development and screening processes have evolved from monolayer culture systems toward more complex and relevant models that can help predict clinical drug resistance [ 30 ]. The combination of these innovative experimental approaches with new generation genomic and proteomic technologies will help identify novel resistance mechanisms and study therapies that can overcome this process and target cells that are not susceptible to current treatments [ 31 ]. Breast cancer is one of the most frequent cancers and common causes of cancer‐related deaths among women [ 32 ]. Despite the advances in breast cancer treatment in early and metastatic phase, medical therapies still fail in patients due to pharmacological resistance, resulting in disease progression, recurrence, and reduced overall survival [ 33 ]. We previously established a 3D technology based on biomimetic scaffolds that mimic the hierarchically organized structure of extracellular collagen, a matrix protein that is present in almost every tissue of the body [ 34 ]. Biocompatible materials with high degree of ECM biochemical mimicking have been successfully used in different tissue engineering applications [ 35 ], as native ECMs provide fundamental stimuli affecting cell function during pathophysiological events, including cancer development and evolution [ 36, 37, 38, 39 ]. Our scaffolds enabled the modeling of the tumor hypoxic niche and its contribution to disease progression. We implemented this platform for the identification of lineage‐specific drug resistance mechanisms. Here, we have applied this approach in two established breast cancer models and identified mechanisms not yet fully described in literature. Clinically relevant biomarkers were investigated and generated to predict doxorubicin efficacy in patients. 2 Materials and methods 2. 1 Collagen scaffold synthesis All chemicals were purchased from Sigma‐Aldrich (St. Louis, MO, USA). The collagen scaffolds were synthesized and characterized as previously described [ 34 ]. Type I collagen was suspended in acetic acid, precipitated to pH 5. 5, and cross‐linked with 1, 4‐butanediol diglycidyl ether. An established freezing and heating ramp (from 25 °C to −25 °C and from −25 °C to 25 °C in 50 min under vacuum conditions, P = 0. 20 mbar) produced the scaffold's porosity ensuring proper pore size, interconnectivity, and orientation. Scaffolds were sterilized in 70% ethanol for 1 h and then washed three times in sterile Dulbecco's Phosphate‐Buffered Saline (Life Technologies, Carlsbad, CA, USA). Porosity and pore size of the scaffold were determined as previously described [ 34 ]. 2. 2 Cell seeding and culture The human breast cancer cell lines MDA‐MB‐231 and MCF‐7 were obtained from the American Type Culture Collection (Rockville, MD, USA). Cells were maintained in DMEM medium with 10% fetal bovine serum, 1% penicillin‐streptomycin, and 1% glutamine (PAA, Piscataway, NJ, USA) at 37 °C in a 5% CO 2 atmosphere. For monolayer cultures, 6 × 10 5 cells were seeded in 25‐cm 2 flasks. For 3D cultures, 5 × 10 6 cells suspended in 50 µL of culture medium were dropped onto the upper surface of each scaffold (1 × 9 mm) in 24 multiwell plates. Scaffolds were plunged in PBS to maintain them hydrated. Before cell seeding, scaffolds were dried out through the elimination of liquid using sterile tips. Seeding was reached by soaking of the cell suspension in dry scaffolds. After cells were allowed to adhere for 1 h at 37 °C, the culture medium was gently added in each well. After 24 h, scaffolds were gently placed in a six multiwell plate. The medium was replaced daily. For the in vivo study, luciferase‐transfected MDA‐MB‐231 and MCF‐7 cells were maintained in DMEM high glucose without sodium pyruvate with 10% FCII (Fetal Clone II; Hyclone, Logan, UT, USA), 1% penicillin/streptomycin, 1% glutamine, and 800 µg·mL −1 of Geneticin (G418 Invitrogen, Waltham, MA, USA) for selection of luciferase and cultured as previously described [ 40 ]. 2. 3 Scanning electron microscopy and confocal microscopy Cells in collagen scaffolds were imaged by scanning electron microscopy (SEM) and laser confocal microscopy as previously described [ 34 ]. For SEM, the samples were washed in 0. 1 m sodium cacodylate buffer pH 7. 4 and fixed in 2. 5% glutaraldehyde 0. 1 m sodium cacodylate buffer for 2 h at 4 °C. Before imaging, samples were dehydrated in a series of ethanol, dried in a dessicator overnight, and sputter‐coated with platinum. The Nova NanoSEM 230 (FEI, Hillsboro, OR, USA) was used to acquire all the images. For confocal microscopy cells were fixed in 4% paraformaldehyde for 20 min and stained with Alexa Fluor™ 546 Phalloidin and DAPI (Thermo Fisher Scientific, Waltham, MA USA) for 1 h at 4 °C. For lysosomes detection, cells were collected by trypsinization for monolayer cultures or by digestion in Collagenase type I (Merck Millipore, Burlington, MA, USA) for 3D culture. Cells were then stained with 75 n m LysoTracker™ Green DND‐26 (Thermo Fisher Scientific) for 30 min at 37 °C and cytospinned onto glass slides. For yH2AX immunofluorescence staining, Phospho‐Histone H2A. X (Ser139) (mAb #9718 Cell Signaling, Beverly, MA, USA) was used (1 : 400) and detected with secondary antibody Alexa Fluor™ 488. Images were acquired with an N‐SIM E laser confocal microscope (Nikon Corporation, Tokyo, Japan) and performed at 20× magnification. 2. 4 Transcriptome analysis The expression profile of 3D‐cultured MDA‐MB‐231 and MCF‐7 cells treated for 72 h with 4 µg·mL −1 doxorubicin was compared with that of untreated cells. Gene expression analysis was carried out using the Illumina HumanHT‐12 v4 Expression BeadChip (Illumina, San Diego, CA, USA). For RNA extraction, the scaffolds were fragmented into small pieces, while monolayer cultured cells were collected by tripsinization. RNA quality control was performed through an electrophoretic run on Agilent Bioanalyzer using the Agilent RNA 6000 Nano Kit (Agilent Technologies, Santa Clara, CA, USA). Total RNA samples were processed using the Ambion Illumina TotalPrep RNA Amplification kit. Beadchips were hybridized and processed following the Illumina Whole Genome Gene Expression Direct Hybridization Assay protocol. Fluorescence data generated by the iScan were analyzed with the Ilumina genomestudio software package. Data normalization was performed using the Robust Spline Normalization (RSN) algorithm. The identification of differentially expressed genes (DEGs) was addressed using Linear Models for Microarray Data (LIMMA) and empirical Bayes methods together with false discovery rate correction of the P ‐value (Benjamini–Hochberg). Statistically significant DEGs ( P. adj. value < 0. 01) have been selected according to a |LogFC| > 1. We used KEGG (Kyoto Encyclopedia of Genes and Genomes), REACTOME, and Gene Ontology (GO) tools to test for the enrichment of any pathway/terms that may be related to the drug resistance phenotypes. For each tool, we have taken into consideration the first twenty terms sorted by adjusted P ‐value. 2. 5 Quantitative real‐time reverse transcriptional‐PCR (qRT‐PCR) Total mRNA was isolated using TRIzol Reagent (Invitrogen) following the manufacturer's instructions and reverse‐transcribed using the iScript cDNA Synthesis Kit (Bio‐Rad, Hercules, CA, USA). The final mixture was incubated at 25 °C for 5 min, at 42 °C for 20 min, at 47 °C for 20 min, at 50 °C for 15 min, and at 85 °C for 5 min. Real‐Time PCR was performed on the 7500 Real‐Time PCR System using the SYBR Select Master Mix (Applied Biosystems, Foster City, CA, USA). Primers sequences are reported in Table S1. Amplification was performed in a final volume of 20 µL containing 2× Gene expression master Mix (Applied Biosystem), 2 µL of cDNA in a total volume of 20 µL. The reaction mixtures were all subjected to 2 min at 50 °C, 10 min at 95 °C followed by 40 PCR cycles at 95 °C for 15 s and 60 °C for 1 min for overall markers. The amount of transcripts was normalized to the endogenous reference genes β‐actin and HPRT and expressed as n‐fold mRNA levels relative to a calibrator using a comparative threshold cycle (Ct) value method (∆∆Ct). The RNA extracted from untreated cells was used as the calibrator. 2. 6 Flow cytometry Cells were collected by trypsinization for monolayer culture or by digestion in Collagenase type I (Merck Millipore) for 3D culture. To determine cell viability, cells were stained with 50 µ m calcein‐AM and 2 m m ethidium homodimer‐1 (Invitrogen). For lysosomes quantification, cells were stained with 50 n m LysoTracker™ Green DND‐26 (Thermo Fisher Scientific). The TUNEL assay was performed with the In Situ Cell Death Detection Kit (Roche, Basel, Switzerland) according to the manufacturer's protocol. The cell suspensions were analyzed on the BD FACS CantoI (Beckmann Coulter, Brea, CA, USA). 2. 7 Immunohistochemical analysis Scaffolds were fixed in neutral buffered formalin, dehydrated by incubation in scaling ethanol solutions (30–100%), and embedded in paraffin as previously described [ 34 ]. Paraffin blocks were sliced with a rotating microtome (Leica Biosystems, Wetzlar, Germany) at 5 µm thickness, and sections were mounted onto Superfrost Plus microslides (Thermo Fisher Scientific, Waltman, MA, USA). Hematoxylin and eosin staining was performed to assess the scaffold architecture, cell morphology, and distribution. Immunostaining for anti‐HIF‐1α (1 : 500, Abcam) was performed using the Ventana Benchmark XT staining system (Ventana Medical Systems, Tucson, AZ, USA) with the Optiview DAB Detection Kit (Ventana Medical Systems). 2. 8 Western blot Proteins were isolated with a lysis buffer composed of 50 m m Tris/HCl (pH 8), 150 m m NaCl, 1% Triton X‐100, and 0. 1% SDS, supplemented with the Halt Protease and Phosphatase Inhibitor Cocktail (Thermo Fisher Scientific). The protein content was quantified using the BCA protein assay kit (Thermo Fisher Scientific). For each sample, an equal amount of protein was loaded on Bolt™ 10% Bis‐Tris Plus Gels (Life Technologies) and transferred to polyvinylidene fluoride membranes through Trans‐Blot ® Turbo™ blotting system (Bio‐Rad). The membranes were blocked in 5% nonfat dry milk PBS with 0. 1% Tween 20 (Sigma‐Aldrich) for 2 h at room temperature. Then, the membranes were incubated overnight with primary antibodies at 4 °C. The following antibodies were used: anti‐CASP3 (1 : 1000; Cell Signaling Technology) and anti‐vinculin (1 : 1000; Thermo Fisher Scientific). After two washes, the membranes were incubated for 1 h at room temperature with horseradish peroxidase‐conjugated secondary antibody followed by the visualization of the proteins with a chemidoc XRS system (Bio‐Rad). 2. 9 Doxorubicin testing Doxorubicin treatment was performed in monolayer cultures or in the 3D scaffolds at the following concentrations: 0. 8, 1. 6, and the human plasma peak concentration 4 μg·mL −1 [ 41, 42 ]. Doxorubicin hydrochloride solution was diluted in culture media. Cells were cultured for 24 h before exposure to the drug. Cell viability was assessed after 72 h of treatment (according to the terminal half‐life of doxorubicin) [ 43 ] by MTT assay directly in the scaffolds or in the culture wells. Briefly, controls and drug‐treated samples were incubated with 0. 5 mg·mL −1 of MTT solution (Sigma‐Aldrich) in DMEM for 2 h at 37 °C. Cell viability was determined by reading the absorbance at 550 nm. Survival percentages were calculated as the average absorbance of cells at each doxo doses over the absorbance of untreated cells. In vivo experiments were performed through orthotopically injections into the right mammary fat pad of 6‐week‐old female immunodeficient NU/NU nude mice (Crl:NU‐Foxn1nu) purchased from Charles River Laboratories. 2 × 10 6 MDA‐MB‐231 and MCF‐7 cells were marked with Luciferase and suspended in 100 µL Matrigel (BD) before the injection. Mice were maintained under pathogen‐free conditions and on low‐fluorescence diet according to the guidelines set forth by the National Institutes of Health. Tumor growth was followed by in vivo bioluminescence imaging using the Xenogen IVIS 200 In Vivo Bioluminescence Imaging System (PerkinElmer, Waltham, MA, USA) every 2–3 days after cells injection. Tumor volume was assessed at each time point by caliper measurement. When the tumors reached an average volume of 70 mm 3, animals were randomly assigned to either control or doxo group (5 mice per experimental group). Doxorubicin hydrochloride was dissolved in saline and administered daily by intraperitoneal injection at the doses of 0. 2, 0. 08, and 0. 04 mg·kg −1 (dosages were selected according to the human plasma peak of doxorubicin from pharmacokinetic clinical data and converted to mice equivalent surface area) [ 44 ], while control animals were injected with the same volume of saline. After 3 days, the treatment was stopped and after 1 week animals were sacrificed (Fig. S1 ). The percentages of cell survival were calculated normalizing the average volume of treated tumors versus the average volume of untreated controls. Tumors were collected, fragmented into small pieces, and stored in TRIzol at −80 °C for RNA extraction or in lysis buffer at −80 °C for protein extraction. For both in vivo and in vitro data, the IC 50 values were calculated from the nonlinear regression of the dose–log response curves. 2. 10 Statistics For each experiment, at least three biologically independent replicates were performed. Data were presented as mean ± standard deviation (S. D. ), or mean ± standard error of the mean (S. E. M. ), as specified. N indicates the number of replicates. The differences between groups were assessed by two‐tailed Student's t ‐test or Mann–Whitney test, as stated, and accepted as significant when P < 0. 05. 2. 11 Study approval All animal procedures were reviewed and approved by the Institutional Animal Care and Use Committee (IACUC) of the Houston Methodist Research Institute (HMRI) protocol number AUP 0614‐0033. 3 Results 3. 1 Efficacy of doxorubicin in 3D culture is analogous to in vivo response Breast cancer cells cultured in collagen scaffolds recreated a tissue‐like organization with distinct cell phenotypes, as previously observed [ 34 ]. MCF‐7 grew in discrete round clusters with a tightly cohesive structure and displayed an epithelial morphology (Fig. 1A–C ). MDA‐MB‐231 grew homogeneously dispersed within the scaffold's pores showing a spindle mesenchymal morphology (Fig. 1D–F ). We compared the efficacy of doxorubicin (doxo), one of the most used chemotherapy agent for the treatment of breast cancer patients [ 45 ], in 3D culture, standard monolayer, and in vivo. Both cell lines showed within the scaffold a decreased sensitivity to the drug compared with cells in monolayer, as shown by the higher rates of survival at all tested concentrations (Fig. 1G ) and by the IC 50 values (Fig. 1H ). Efficacy in the scaffold was analogous to that demonstrated by cells growing in vivo in terms of survival rates, dose–response curves, and IC 50 values. Conversely, cells cultured in monolayer were much more sensitive to the drug. Notably, MCF‐7 demonstrated to be relatively insensitive to doxorubicin treatment in 3D and in vivo conditions being completely resistant at the highest dose: Cells treated with 4 µg·mL −1 doxorubicin showed a survival percentage near to 100% and the IC 50 concentration was not reached (Fig. 1H ). Fig. 1 Characterization and drug sensitivity of breast cancer cells in 3D collagen scaffold. (A, D) SEM micrograph of collagen scaffolds showing the porous surface of the material cellularized with MCF‐7 (A) and MDA‐MB‐231 (B) ( n = 3). (B, E) Hematoxylin‐and‐eosin–stained histological sections of MCF‐7 (B) and MDA‐MB‐231 (E) within the scaffolds at day 7 ( n = 3). (C, F) Confocal microscopy images of MCF‐7 (C) and MDA‐MB‐231 (F) within the scaffold at day 7 ( n = 3). Cells are stained with DAPI (blue) and phalloidin (red). Scale bars for all pictures: 100 µm. (G) Percentages of survival of MCF‐7 and MDA‐MB‐231 after 72 h of treatment with different concentrations of doxorubicin in monolayer culture (2D), within the scaffold (3D), or orthotopically implanted into a murine model ( in vivo ). Data represent mean ± S. D. ( n = 3 for in vitro data, n = 6 for in vivo data). (H) Nonlinear fit of log–dose responses curves and IC 50 calculation. Decreased sensitivity was not caused by impaired drug penetration in inner scaffold areas: No significant differences were observed in the mean fluorescence intensity of doxorubicin between cells in core or edge regions of the scaffold, with the exception of the 0. 8 µg·mL −1 dose in MDA‐MB‐231 ( P = 0. 032) (Fig. S1 b). 3. 2 Lineage‐specific signaling pathways are activated in doxorubicin‐treated cells within the scaffold Diverse signaling pathways were found to be modulated in the two cell lines in response to doxorubicin administration. Transcriptome analysis demonstrated that MCF‐7 cultured in the scaffolds and treated with 4 µg·mL −1 doxo, at which cells were totally resistant, showed upregulation of the systemic lupus erythematosus and p53 signaling pathways (Fig. 2A, B ). Between the most DEGs TAP1, TP53I3, GADD45G, GADD45B, and S100P were found. The expression levels of selected candidate DEGs were quantified by qPCR analysis (Table S1 ). All genes resulted significantly upregulated in treated cells compared with controls ( P = 0. 0461 for TAP1, P = 0. 0075 for TP53I3 and P = 0. 0085 for S100P ) (Fig. 2C ). In MDA‐MB‐231, we observed the upregulation of the glycosylphosphatidylinositol (GPI)‐anchor biosynthesis and lysosome pathways and downregulation of pathways involving cell cycle, endocytosis, spliceosome, RNA degradation, and Pathogenic Escherichia coli infection (Fig. 2D, E ). Between the most deregulated genes, LAPTM4A, LAPTM4B, PRKCZ, LAMP2, RAB40C, RAB22A, and MMP3 were found. The qPCR analysis on selected DEGs (Table S2 ) confirmed that LAPTM4A, LAPTM4B, LAMP2, RAB40C, RAB22A, and MMP3 were significantly upregulated in treated samples compared with controls, while PRKCZ was downregulated ( P = 0. 01119 for LAPTM4A, P = 0. 01833 for LAPTM4B, P = 0. 02232 for LAMP2, P = 0. 01293 for RAB40C, P = 0. 00779 for RAB22A, P = 0. 03562 for MMP3 and P = 0. 00169 for PRCKZ ) (Fig. 2F ). The identified markers were not affected when cells were treated with doxorubicin in monolayer culture with the exception of PRKCZ that resulted downregulated in treated samples ( P = 0. 03844), as observed within the scaffolds (Fig. S1 c). Fig. 2 Transcriptomic data analysis of breast cancer cells treated with doxorubicin. (A, D) Gene count of significantly altered pathways identified in MCF‐7 (A) or MDA‐MB‐231 (D) treated with doxorubicin within 3D collagen scaffolds. (B, E) Log fold change of DEGs in each identified pathway for MCF‐7 (B) or MDA‐MB‐231 (E). (C, F) Relative expression levels from qPCR data of candidate DEGs belonging to the identified pathway for MCF‐7 (C) and MDA‐MB‐231 (F) treated with doxo. The values are relative to untreated control samples. Data represent mean ± S. D. ( n = 3). * P < 0. 05, two‐tailed Student's t ‐test. 3. 3 Reactivation of caspase 3 and p53 signaling induction in drug‐resistant MCF‐7 In MCF‐7, the selection of a doxorubicin‐resistant subpopulation from parental cells was observed, as demonstrated by the detection of caspase 3 protein in cells cultured in the scaffold (Fig. 3A ). MCF‐7 are normally known to express a truncated isoform of caspase 3, while the drug‐resistant sublines express the full‐length transcript [ 46 ]. Cell selection process was partially independent from drug exposure as cells expressing full‐length caspase 3 were present also in control samples, although to a lower extent. Despite the expression of caspase 3, cells within the scaffold showed lower levels of apoptotic cell death after drug exposure, compared with cells in monolayer ( P = 0. 01434 for the 0. 8 µg·mL −1 dose) (Fig. 3B ). We, thus, investigated the activation of DNA damage response. Doxorubicin treatment induced an increase in the accumulation of γH2AX foci, which is indicative of DNA double‐strand breaks, in both monolayer and 3D‐cultured cells. However, DNA damage response activation and apoptosis detection were significantly lower in cells treated within the scaffold compared with monolayer, with faster resolution of γH2AX foci (Fig. 3C ). The number of cells positive for γH2AX and the average number of foci were significantly lower within the scaffold, in particular at later time points ( P = 0. 0032 at 2 h, P = 0. 0048 at 24 h, and P = 0. 0098 at 48 h after treatment for the number of positive cells; P = 0. 0032 at 2 h, and P = 0. 0151 at 48 h after treatment for the average number of foci) (Fig. 3C ). Conversely, in monolayer resolution of γH2AX foci of DNA, breaks occurred at a later time point (48 h), and apoptotic cells were detected after 24 h (arrowed) (Fig. 3C ). Caspase 3 expression and induction of p53 signaling (Fig. 3D ) in cells treated within the scaffold did not result in significant DNA damage response activation and apoptosis. Fig. 3 Mechanism of drug resistance in MCF‐7 cultured within the scaffold. (A) Western blot for caspase 3 in MCF‐7 untreated or treated with different concentrations of doxorubicin for 72 h in monolayer culture (2D) or within the scaffold (3D). (B) Percentages of apoptotic MCF‐7 after 72 h of treatment with different concentrations of doxorubicin in 2D or 3D cultures. Data represent mean ± S. D. ( n = 3). * P < 0. 05, two‐tailed Student's t ‐test. (C) Immunofluorescence staining of γH2AX in MCF‐7 cells untreated (ctr) or treated with doxorubicin for 2, 6, 24, and 48 h in 2D or 3D cultures; quantification of the percentages of γH2AX‐positive cells and of the average number of γH2AX foci per cell. Data represent mean ± S. E. M ( n = 5). * P < 0. 05, two‐tailed Student's t ‐test. Scale bars for all pictures: 20 µm. Arrows indicate apoptotic cells. (D) Schematic representation of doxorubicin effects in MCF‐7 cell line cultured within the scaffold. The most significantly altered pathway implicated in drug resistance with the list of relative DEGs are reported in the box. Green is indicative of upregulation. 3. 4 MDA‐MB‐231 show reduced doxorubicin uptake and lysosomal confinement of the drug MDA‐MB‐231 treated with doxorubicin in the scaffold showed a reduced intracellular accumulation of the drug compared with cells in monolayer, as demonstrated by lower doxorubicin fluorescence signal detected by flow cytometry at different time points after administration (Fig. 4A ). These data were confirmed by fluorescence microscopy analysis of treated cells recovered from monolayer culture or from the scaffold (Fig. S2 A). In monolayer, doxorubicin signal decreased over time, while it was relatively constant for cells in 3D. Interestingly, we observed the presence of a cell side population characterized by a low‐fluorescence intensity both in the doxorubicin and in the calcein‐AM fluorescence channels (Fig. S2 B). Calcein is known to be extruded by the multidrug transporter MDR‐1 before the intracellular conversion to its fluorescent‐free isoform and provides an efficient experimental method to determine the activity of MDR‐1 in cancer cells [ 47 ]. The percentage of cells showing this phenotype was significantly lower in monolayer culture ( P = 0. 045 and P = 0. 039 at 0 and 4 µg·mL −1 dose, respectively). Moreover, we found that the lysosomal content of cells cultured within the scaffold was significantly higher compared with cells in monolayer (Fig. 4B and Fig. S2 C). Confocal microscopy analysis demonstrated the colocalization of the fluorescence signal of doxorubicin with labeled lysosomes (Fig. 4C ). These mechanisms were partially independent from drug exposure, as all observed phenotypes were detected also in untreated samples, although to a lower extent. Taken together, these data are consistent with the reduced endocytosis and the induction of lysosomal pathway observed by transcriptome analysis in cells treated within the scaffold (Fig. 4D ). Fig. 4 Mechanism of drug resistance in MDA‐MB‐231 cultured within the scaffold. (A) Doxorubicin median fluorescence intensity detected by flow cytometry in MDA‐MB‐231 after 6, 24, 48, and 72 h of treatment with different doxo concentrations in monolayer culture (2D) or within the scaffold (3D). Data represent mean ± S. D. ( n = 3). (B) Histogram plot of MDA‐MB‐231 stained with lysotracker (LT) green in 2D or 3D cultures and median fluorescence intensity of LT green in control cells or cells treated with 4 µg·mL −1 doxorubicin after 24, 48, and 72 h. Data represent mean ± S. D. ( n = 3). (C) Confocal microscopy images of MDA‐MB‐231 treated with doxorubicin within the scaffold. Red is doxorubicin autofluorescence and green is LT green signal. Scale bar is 10 µm (D) Schematic representation of doxorubicin effects in MDA‐MB‐231 cell line cultured within the scaffold. The most significantly altered pathways implicated in drug resistance with the list of relative DEGs are reported in the box. Green is indicative of upregulation. Red is indicative of downregulation. 3. 5 Hypoxia is involved in doxorubicin resistance We previously observed that our 3D model allows for the creation of a hypoxic core environment that guides multiple phenotypic changes in breast cancer cells [ 34 ]. To address the correlation between hypoxia and the emergence of chemotherapy resistance in 3D‐cultured cells, we performed treatment in the presence of Resveratrol, a hypoxia inhibitor that reduces hypoxia‐mediated HIF‐1α accumulation [ 48 ]. A preincubation of 24 h with 50 µ m Resveratrol was performed prior to drug administration. While Resveratrol did not affect cancer cell proliferation either in the scaffold or in monolayer cultures (Fig. 5A ), it markedly decreased HIF‐1α expression in both cell lines (Fig. 5B ). Hypoxia inhibition was found to re‐sensitize 3D‐cultured cells to doxorubicin, as proved by the marked decrease in survival percentages demonstrated by MCF‐7 at the highest drug dose and by MDA‐MB‐231 at the doses of 0. 8 and 1. 6 µg·mL −1 (Fig. 5C ). Sensitivity was not completely restored as that observed for monolayer cultured cells, suggesting that other 3D‐related phenotypes, such as the slower rate of proliferation, might contribute to the induction of resistance. Fig. 5 Role of hypoxia in 3D‐induced drug resistance. (A) Survival percentages (day 7) of MCF‐7 and MDA‐MB‐231 cultured in monolayer (2D) of within the scaffold (3D) in the absence (CTR) or presence of a hypoxia inhibitor (HI). Data represent mean ± S. D. ( n = 5). (B) HIF‐1α expression in histological sections of MCF‐7 and MDA‐MB‐231 cultured within the scaffold in control conditions (CTR) or in the presence of an hypoxia inhibitor (HI). Scale bars: 50 µm. (C) Percentage of survival of MCF‐7 and MDA‐MB‐231 after 72 h of treatment with different concentrations of doxorubicin in monolayer culture (2D), within the scaffold (3D) and in the presence (3D HI) of a hypoxia inhibitor. Data represent mean ± S. D. ( n = 3). (D) Relative expression levels from qPCR data of candidate DEGs in MCF‐7 untreated or treated with 4 µg·mL −1 doxorubicin under control conditions (ctr) or in the presence of hypoxia inhibition (HI). Data represent mean ± S. D. ( n = 3). * P < 0. 05, two‐tailed Student's t ‐test. (E) Relative expression levels from qPCR data of candidate DEGs in MDA‐MB‐231 untreated or treated with 4 µg·mL −1 doxorubicin under control conditions (ctr) or in the presence of hypoxia inhibition (HI). Data represent mean ± S. D. ( n = 3). * P < 0. 05, two‐tailed Student's t ‐test. Moreover, the hypoxia inhibition reduced the expression level of all the biomarkers involved in resistance mechanisms. After doxorubicin treatment under hypoxia, TAP1 and S100P were not significantly overexpressed in MCF‐7 cell line, in contrast to control conditions (Fig. 5D ). Conversely, expression of TP53I3 was significantly enhanced even when hypoxia was inhibited. In MDA‐MB‐231, the differences in expression of LAPTM4A, LAPTM4B, LAMP2, RAB40C, and RAB22A between control and treated samples resulted not significant under hypoxia inhibition, in contrast to what observed for samples treated under control conditions (Fig. 5E ). On the contrary, a significant downregulation of PRKCZ and overexpression of MMP3 was observed even when hypoxia was inhibited. Interestingly, when cancer cells are pretreated with hypoxia inhibitors, there is a significant downregulation of most of the markers in response to doxorubicin treatment in cells. This result suggests a direct correlation between hypoxia and drug exposure in the acquired resistance by MDA‐MB‐231. 3. 6 The identified biomarkers are involved in doxorubicin response in vivo and in breast cancer patients In order to understand their translational significance, we analyzed the expression of all biomarkers involved in resistance in orthotopic tumors generated in murine models and in a cohort of breast cancer patients from a public dataset. In in vivo samples, ER+ tumors treated with doxorubicin displayed significant upregulation of TP53I3 and S100P, compared with untreated samples. Also TAP1 was upregulated by treatment although the data were not statistically significant (Fig. 6A ). In MDA‐MB‐231, we found a significant upregulation of LAPTM4B, RAB40C, MMP3, and downregulation of PRKCZ after tumor treatment with doxorubicin. Conversely, expression of LAPTM4A, LAMP2 and RAB22A resulted not significantly affected (Fig. 6B ). Fig. 6 Expression levels of the identified biomarkers in vivo and in breast cancer patients. (A) Relative expression levels from qPCR data of candidate DEGs in MCF‐7 untreated or treated with 4 µg·mL −1 doxorubicin in an orthotopic murine model. Data represent mean ± S. D. ( n = 3). * P < 0. 05, two‐tailed Student's t ‐test. (B) Relative expression levels from qPCR data of candidate DEGs in MDA‐MB‐231 untreated or treated with 4 µg·mL −1 doxorubicin in an orthotopic murine model. Data represent mean ± S. D. ( n = 3). * P < 0. 05, two‐tailed Student's t ‐test. (C) Expression levels of TAP1 and TP53I3 in ER‐positive breast cancer patients and LAMP1 and LAMP2 in triple‐negative breast cancer patients in relation to response to anthracycline treatment. Patients were classified as responder or nonresponder according to the 5‐year relapse‐free survival. ROC curves of TAP1 and TP53I3 as predictor of response to anthracycline treatment in ER‐positive breast cancer patients, and LAMP1 and LAMP2 as predictor of response to anthracycline treatment in triple‐negative breast cancer patients. * P < 0. 05, two‐tailed Student's t ‐test, Mann–Whitney test. For the expression analysis in breast cancer patients, we used the online transcriptome‐level validation tool for predictive biomarkers ROC Plotter that integrates 3104 breast cancer patients with treatment and response data [ 49 ]. We investigated the correlation between expression of the biomarkers identified in our screening and response to anthracycline regimens by means of relapse‐free survival at 5 years. Each biomarker was investigated in patients with the matching molecular subgroups. In ER + breast tumor, we found that patients who did not respond to therapy showed a higher expression of TAP1 and TP53I3 compared with responders. High expression of TAP1 and TP53I3 was associated with shorter relapse‐free survival after treatment (Fig. 6C ). In triple‐negative breast cancer, we found that patients who did not respond to therapy showed a higher expression of LAMP1, but lower expression of LAMP2 compared with responders. High expression of LAMP1 was associated with shorter relapse‐free survival after treatment (Fig. 6C ). All other identified markers did not show a significant deregulation in this dataset of patients (data not shown). 4 Discussion Engineered 3D models are generating increasing knowledge on drug sensitivity and on mechanisms of resistance acquisition in cancer cells, while offering high‐throughput analyses and cost‐efficient screenings [ 50, 51, 52, 53, 54, 55 ]. These innovative experimental models have represented a groundbreaking innovation for the clinical translation of anticancer agents. Here, we used a 3D technology based on biomimetic collagen scaffolds, enabling the modeling of the tumor hypoxic niche, to identify and describe mechanisms and drivers of chemotherapy resistance in breast cancer. Firstly, we demonstrated that in vitro results from our 3D model were comparable to those obtained using murine tumor xenografts. The activity of doxorubicin, one of most used chemotherapy agent for the treatment of breast cancer patients, tested in 3D was predictable of in vivo response. Conversely, efficacy was significantly overestimated when tested in monolayer culture. In vivo models remain the gold standard for preclinical drug development, despite showing the important drawbacks of time‐consuming, high cost, and availability depending on the tumor type [ 56 ]. The development of more reliable in vitro systems is reducing the amount of animals required for pharmacological trials, allowing to generate data with comparable translational value. In particular, an interesting observation was that the ER + luminal A cell line demonstrated poor responsiveness to doxorubicin within our cancer model and in tumor xenografts. This is consistent with emerging clinical evidence that indicates the potential lack of benefit from anthracycline chemotherapy in patients with ER + luminal A breast tumors [ 57, 58, 59, 60 ]. Through our 3D model, we next described the mechanisms of resistance which were specifically activated in the two molecular subgroups of breast cancer. In ER + positive cells, the selection of a drug‐resistant subpopulation was observed. The presence of inherently resistant subclones in parental MCF‐7 cells, characterized by the expression of full‐length CASP3, has already been demonstrated [ 46 ]. Here, we showed that culturing in our biomimetic model results in the selection and propagation of this resistant subclone. This subpopulation shows overexpression of TP53I3 and TAP1 correlated to multidrug resistance in human cancers [ 61, 62 ] and with the presence of hypoxic conditions [ 63 ]. TP53I3 has been found to be involved in mitotic progression regulation in non–small‐cell lung cancer [ 62 ], while TAP1 is a member of the superfamily of ATP‐binding cassette (ABC) transporters [ 61 ]. In particular, TAP complex possesses characteristics of a xenobiotic transporter and the TAP dimer contributes to the atypical MDR phenotype of human cancer cells, mediating the translocation of hydrophobic antitumor agents into the endoplasmic reticulum lumen [ 61 ]. These cells displayed also reduced DNA damage response, despite expression of caspase 3, indicating a potential increased ability of DNA repair. This was further suggested by the enhanced expression of the GADD45 family, members of the p53 signaling pathway, and mediators of demethylation and DNA excision repair [ 64 ]. Triple negative cells were able to reduce the intracellular drug accumulation through different processes: the downregulation of endocytic pathway components and the selection of a side subpopulation displaying the ability of extruding calcein and doxorubicin. Side population cells have been identified in several human cancers and are defined as cells capable of extruding dyes, such as Hoechst 33342, through the ABC transporters [ 47, 65, 66 ]. These cells were found, not only to possess increased drug resistance but also to display stem‐like properties [ 67 ]. Culturing within our 3D environment results in a significant selection of side population cells offering the possibility to further understand their functional and molecular characteristics. In addition to reduced intracellular drug accumulation, we found these cells to activate the lysosomal pathway and to accumulate doxorubicin inside lysosomes. This mechanism of resistance, identified in cisplatin‐treated cancer cells [ 68 ], has not yet been described for anthracyclines and hold an interesting potential. It has been demonstrated that mammalian target of rapamycin complex 1 (mTORC1), a downstream effector of oncogenic pathways, directly regulates the lysosomal biogenesis [ 69 ]. Several compounds able to suppress mTORC1 functions, as everolimus and temsirolimus, have been developed and are currently in clinical practice [ 69 ]. Combinatorial regimens, by counteracting the development of resistance, can be more effective than single therapy and should be considered as the best treatment option for many cancer patients [ 29 ] in order to prevent the increasing prevalence of drug resistance [ 70 ]. Here, we provide preliminary data to support the clinical rationale to explore the combination of doxorubicin and mTOR inhibitors for the treatment of triple‐negative breast cancer patients. However, further analyses are needed to support this approach. In both cell lines, the resistant subpopulations emerged independently from doxorubicin exposure denoting intrinsic mechanisms, while treatment enhanced the observed phenotypes. Indeed, we demonstrated that the pretreatment of cancer cells with a hypoxia inhibitor hamper the upregulation of the identified markers related to drug resistance in response to doxorubicin exposure, suggesting a direct correlation between hypoxia and drug treatment. Hypoxia showed a central role in promoting resistance acquisition, as the blocking of HIF‐1α partially restored drug sensitivity in both breast cancer subtypes and decreased the molecular alterations induced by treatment. The role of hypoxia in cancer drug resistance is well documented. It has been demonstrated that hypoxia can confer resistance by regulating a number of signaling pathways as apoptosis, autophagy, DNA damage, mitochondrial activity, p53, and drug efflux [ 71 ]. In breast cancer, it has been recently demonstrated that resistance is connected to an increased plasticity of cells mediated by hypoxia [ 72 ]. Therefore, the possibility to model this process when screening anticancer agents demonstrates a crucial value and will help gaining new insights into mechanisms and molecular drivers of drug resistance [ 73, 74, 75 ]. Finally, profiling of treated cancer cells within the scaffold led to the identification of candidate predictive biomarkers. Several evidences indicates that engineered 3D models can be useful approaches to study and identify drug resistance mechanisms to anticancer agents [ 76, 77, 78, 79, 80 ]. Here, we demonstrate that the molecular changes identified through our biomimetic model are (a) predictive of in vivo molecular alterations on tumor xenografts and (b) demonstrate clinical predictive potential. Indeed, some of the biomarkers identified in our screening showed a significant value in predicting the 5‐year relapse rate of patients with breast cancer treated with anthracyclines regimens. Although our analysis shows some limitations, as the lack of standardization of patient characteristics in public datasets, it provides a proof‐of‐concept of the clinical value of these biomarkers. Therefore, further validation in independent cohorts of patients, ideally considering a neoadjuvant setting, is warranted. 5 Conclusion These findings suggest that our model might support in vitro trials for the translation of targeted therapies and anticancer compounds as it provides (a) more relevant data on efficacy and (b) enhanced understanding of resistance acquisition, one the major causes of chemotherapy failure in cancer patients [ 1 ]. Our cancer model recreates the emergence of resistance fostered by a hypoxic niche and allows for the investigation of potentially unexplored mechanisms involved in therapy response. This approach may offer therapeutic targets for the design of combinatorial therapies and introduce new predictive biomarkers for precision medicine. Conflict of interest The authors declare no conflict of interest. Author contributions CL, TI, and LM designed the study. CL, ADV, CS, GM, CC, AB, ADL, FLM, FF, MT, and ET acquired and analyzed the data. CL, DA, LM, and TI conceived all the experiments and interpreted the results. CL, LM, and TI drafted the manuscript. All authors read and approved the final manuscript. Peer Review The peer review history for this article is available at https://publons. com/publon/10. 1002/1878‐0261. 13037. Supporting information Fig. S1. Schedule of doxorubicin administration in orthotopic murine models, doxorubicin localization in the 3D model, and DEGs expression in monolayer cells. (a) Schematic representation of the schedule of doxorubicin administration in orthotopic murine breast cancer models, generated by the xenotransplantation of MCF‐7 and MDA‐MB‐231. (b) Median fluorescence intensity of doxorubicin detected by immunofluorescence in MCF‐7 and MDA‐MB‐231 within core or edge regions of the scaffold after 72‐h treatment. Data represent mean ± S. D. ( n = 20) * P < 0. 05, two‐tailed Student's t‐test. (c) Relative expression levels from qPCR data of candidate DEGs belonging to the identified pathway for MCF‐7 and MDA‐MB‐231 treated with doxo in monolayer cultures. The values are relative to untreated control samples. Data represent mean ± S. D. ( n = 3). * P < 0. 05, two‐tailed Student's t‐test. Click here for additional data file. Fig. S2. Lysosomal‐mediated doxorubicin resistance in MDA‐MB‐231. (a) Representative images and median fluorescence intensity of doxorubicin detected by immunofluorescence in MDA‐MB‐231 cultured in monolayer (2D) or within the scaffolds (3D) after 72 h treatment with different doses. Scale bar is 20 µm. (b) Flow cytometry scatter plot of 2D‐ and 3D‐cultured MDA‐MB‐231 untreated or treated with 4 µg/ml doxo: Samples were double stained with Calcein AM and Ethidium Bromide. SP indicate a side population negative for both signals. On the right, percentages of doxorubicin‐ and calcein‐negative cells (side population) in 2D or 3D‐cultured MDA‐MB‐231 untreated or treated with doxo. Data represent mean ± S. D. ( n = 3). * P < 0. 05, two‐tailed Student's t‐test. (c) Inverted microscopy images of MDA‐MB‐231 treated with doxorubicin in monolayer culture (2D) or within the scaffold (3D). Red is doxorubicin autofluorescence and green is LT green signal. Scale bar is 20 µm. Click here for additional data file. Table S1. List of DEGs found in MCF‐7. Table S2. List of DEGs found in MDA‐MB‐231. Click here for additional data file.
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10. 1002/2211-5463. 12250
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Role of the TGF‐β pathway in dedifferentiation of human mature adipocytes
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Dedifferentiation of adipocytes contributes to the generation of a proliferative cell population that could be useful in cellular therapy or tissue engineering. Adipocytes can dedifferentiate into precursor cells to acquire a fibroblast‐like phenotype using ceiling culture, in which the buoyancy of fat cells is exploited to allow them to adhere to the inner surface of a container. Ceiling culture is usually performed in flasks, which limits the ability to test various culture conditions. Using a new six‐well plate ceiling culture approach, we examined the relevance of TGF‐β signaling during dedifferentiation. Adipose tissue samples from patients undergoing bariatric surgery were digested with collagenase, and cell suspensions were used for ceiling cultures. Using the six‐well plate approach, cells were treated with SB431542 (an inhibitor of TGF‐β receptor ALK5) or human TGF‐β1 during dedifferentiation. Gene expression was measured in these cultures and in whole adipose tissue, the stromal–vascular fraction (SVF), mature adipocytes, and dedifferentiated fat (DFAT) cells. TGF‐β1 and collagen type I alpha 1 (COL1A1) gene expression was significantly higher in DFAT cells compared to whole adipose tissue samples and SVF cells. TGF‐β1, COL1A1, and COL6A3 gene expression was significantly higher at day 12 of dedifferentiation compared to day 0. In the six‐well plate model, treatment with TGF‐β1 or SB431542, respectively, stimulated and inhibited the TGF‐β pathway as shown by increased TGF‐β1, TGF‐β2, COL1A1, and COL6A3 gene expression and decreased expression of TGF‐β1, COL1A1, COL1A2, and COL6A3, respectively. Treatment of DFAT cells with TGF‐β1 increased the phosphorylation level of SMAD 2 and SMAD 3. Thus, a new six‐well plate model for ceiling culture allowed us to demonstrate a role for TGF‐β in modulating collagen gene expression during dedifferentiation of mature adipocytes.
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Abbreviations COL1A1 collagen type I alpha 1 COL1A2 collagen type I alpha 2 COL6A3 collagen type 6 alpha 3 DFAT dedifferentiated ECM extracellular matrix KRH Krebs–Ringer‐Henseleit OM omental SC subcutaneous SVF stromal–vascular fraction TGF‐β transforming growth factor beta Adipose tissue contains connective tissue matrix, preadipocytes, immune cells, and mature adipocytes 1. Mature adipocytes are specialized in lipid storage and are generated from the differentiation of mesenchymal stem cells committed to preadipocytes 2. Matsumoto et al. 3 demonstrated that adipocytes can dedifferentiate into precursor cells to acquire a fibroblast‐like phenotype using ceiling culture. This method is based on the buoyancy of adipocytes, which allows them to adhere to the top inner surface of a reversed culture flask that is completely filled with medium 4. In 2000, Zhang et al. 5 proposed another culture technique in which adipocytes adhere to the underside of a floating piece of glass. More recently, Jumabay et al. obtained dedifferentiated (DFAT) cells from adipocytes using a method that did not require attachment of the cells to a plastic surface as opposed to ceiling culture. In this experimental model, the isolated adipocytes are incubated in the culture medium for 24 h and then transferred to another dish containing a filter. After 5 days, the filter is removed and the DFAT cells sink through the filter to the bottom of the dish 6. We have recently described a modified version of the ceiling culture approach in six‐well plates, allowing us to decrease the number of cells used and test a larger number of culture conditions 7. TGF‐β1 has often been described as an important regulator of adipocyte physiology because of its role in inhibiting adipogenesis 8, 9, 10. This growth factor also plays a major role in adipose tissue remodeling through the induction of ECM protein‐coding genes such as collagens 11, 12, 13. We have recently demonstrated that genes coding for proteins of the extracellular matrix including COL1A1, COL1A2, COL6A3 were significantly upregulated during the dedifferentiation process. In that analysis, we did not observe significant changes in gene expression of other types of collagens including COL4A3, COL5A1, COL5A2, COL6A1, COL6A2, COL8A2, COL20A1 14. Moreover, we have previously shown that gene expression of matrix metalloproteinase 1 (MMP1), fibroblast‐activated protein (FAP), dipeptidyl peptidase IV (DPP4), and transforming growth factor β1 (TGF‐β1) was strongly induced during dedifferentiation. Finally, recent data support a role for the TGF‐β pathway in the dedifferentiation of human pancreatic islet β cells in vivo 15. Although many groups have used the ceiling culture approach, many aspects of this unique cellular process remain to be characterized. Our previous reports 7, 16 suggest that the TGF‐β pathway may be involved in the dedifferentiation process and contribute to the generation of a proliferative cell population that could be useful in cellular therapy or tissue engineering. From the technical standpoint, previous culture systems limited our ability to test various incubation conditions during the process. To the best of our knowledge, no study has ever attempted to modulate the molecular, metabolic, or secretory attributes of the cells during dedifferentiation. Our objective was to implement a new six‐well plate culture system and modulate the dedifferentiation process by targeting the TGF‐β pathway and its effects on the expression of collagens. We hypothesized that the TGF‐β pathway is a significant modulator of COL1A1, COL1A2, and COL6A3 gene expression during the dedifferentiation of mature adipocytes. Materials and methods Tissue sampling We have complied with all mandatory laboratory health and safety procedures in the course of the experimental work presented in this paper. The project was approved by the Research Ethics Committee of the Institut Universitaire de Cardiologie et de Pneumologie de Québec (IUCPQ). Written informed consent was obtained from tissue donors prior to sampling through the management framework of the IUCPQ Obesity Tissue Bank. Adipose tissue samples were obtained from men and women undergoing bariatric surgery. Portions of adipose tissues were digested with collagenase type I in Krebs–Ringer‐Henseleit (KRH) buffer for up to 45 min at 37 °C according to a modified version of the Rodbell method 17. Adipocyte suspensions were filtered through nylon mesh and washed three times with KRH buffer to obtain a purified population of mature adipocytes. The residual KRH buffer of adipocyte isolation, which contained the stromal–vascular fraction, was centrifuged and the pellet was washed in DMEM‐F12 culture medium supplemented with 10% fetal bovine serum, 2. 5 μg·mL −1 amphotericin B, and 50 μg·mL −1 gentamicin. The stromal–vascular cell fraction was then filtered through 140‐μm nylon mesh to remove endothelial/mesothelial cells, placed in culture plates, and cultured at 37 °C under a 5% CO 2 atmosphere. Medium was changed every two to three days. Isolated mature adipocytes were used for ceiling culture, whereas the stromal–vascular fraction was seeded in standard culture flasks containing DMEM‐F12 culture medium supplemented with 2. 5 % fetal bovine serum. In previous studies, we have demonstrated that adipocytes can successfully dedifferentiate independent of their fat depot origin (SC or OM). We also demonstrated that the dedifferentiation process is relatively independent of obesity level, sex, or age of the cell donor 14, 16. Consistent with these results, data from SC and OM samples were combined in our analyses. Ceiling culture When no treatment was used during the dedifferentiation process, isolated mature adipocytes were counted and 500 000 cells were added to a T‐25 flask completely filled with DMEM‐F12 culture medium supplemented with 20% fetal bovine serum, 2. 5 μg·mL −1 amphotericin B, and 50 μg·mL −1 gentamicin. Flasks were incubated upside down at 37 °C, in 5% CO 2 for 7 days. Cells floated up and adhered to the top inner ceiling surface of the flask. After seven days, the medium was removed and the flasks were inverted so that the cells were on the bottom until day 12 in the same medium 7. The medium was changed every three days. For gene expression analysis, cultures from each patient were harvested at days 4 and 7 of the dedifferentiation process as conducted previously 14. One flask per depot per patient was reversed at day 7 and maintained in culture for an additional 5 days (corresponding to day 12). Time points were chosen based on our observations that harvesting cells at day 4 provides a population of round cells that has completely adhered to the flask, while many cells at day 7 begin to be elongated. All ceiling cultures performed in the laboratory were reversed at day 7. Day 12 represents a time point at which the majority of cells have a fibroblast‐like morphology. When DFAT cells were needed for western blot experiments, cultures were maintained in standard condition for more than 12 days and subcultured when cells reached confluence. For other experiments, a ceiling culture model in six‐well plates was used 7. To do so, 8 mL of DMEM/F12 20% fetal bovine serum was added to each well containing a 1/2‐inch plastic bushing. A glass slide was put on top of the bushing and mature adipocytes were seeded under each coverslip (Fig. 1 ). These mature adipocytes floated and then adhered to the top slide. They were examined at specific time points during the dedifferentiation process. This model required a smaller number of cells and allowed us to use various media during the dedifferentiation process 7. The effectors were added to the media, and when appropriate, the glass slides with the adherent cells were flipped in a new well and then harvested for RNA extraction. Figure 1 (A) Picture of the six‐well plate ceiling culture model. Panels 1, 2, and 3: 8 mL of DMEM/F12 20% fetal bovine serum was added to each well containing a 1/2‐inch plastic bushing; panels 4 and 5: A glass slide was put on top of the bushing; panel 6: Mature adipocytes were seeded under each coverslip. (B) Cells floated and adhered to the slides. They can be studied at specific time points. RNA extraction and real‐time quantitative RT‐PCR Total RNA was isolated from SVF cells, whole adipose tissue, and DFAT cells from five donors using the RNeasy lipid tissue extraction kit (Cat No. /ID: 74804) and digested with RNase‐free DNase (Qiagen, Mississauga, ON, Canada) to remove all traces of DNA. RNA extraction was also performed from dedifferentiation time course experiments at day 0 (freshly isolated adipocytes), day 4, and day 7 of ceiling culture and at day 12 from three donors using the QIAGEN RNeasy extraction kit (Cat No. 330503) and in treated DFAT cells. To assess RNA quantity and quality, an Agilent Technologies 2100 Bioanalyzer and RNA 6000 Nano LabChip kit (Agilent, Mountain View, CA, USA) were used. For each sample, cDNA was obtained using the QuantiTect reverse transcriptase kit (Cat No 205311). The following sequences were used for quantitative PCR (forward/reverse): ATP synthase, H+ transporting, mitochondrial F1 complex, O subunit (ATP5O): 5′‐AACGACTCCTTGGGTATTGCTTAA‐3′/5′‐ATTGAAGGTCGCTATGCCACAG‐3′, glucose‐6‐phosphate dehydrogenase (G6PD): 5′ GCAGGGCATTGAGGTTGGGAG‐3′/5′‐GATGTCCCCTGTCCCACCAACTCTG‐3′, transforming growth factor β1 (TGF‐β1): 5′‐AAG TTG GCA TGG TAGCCC TT‐3′/5′‐CCC TGG ACA CCA ACT ATT GC‐3′, transforming growth factor β2 (TGF‐β2): 5′‐CTC CAT TGC TGA GAC GTC AA‐3′/5′‐ATA GAC ATG CCG CCC TTC TT‐3′, transforming growth factor β3 (TGF‐β3): 5′‐CAC ATT GAA GCG GAA AAC CT‐3′/5′‐AAA TTC GAC ATG ATC CAG GG‐3′, collagen type I alpha 1 (COL1A1): 5′‐CAC ACG TCT CGG TCA TGG TA3′/5′‐AAG AGG AAG GCC AAG TCG AG‐3′, collagen type I alpha 2 (COL1A2): 5′‐AGC AGG TCC TTG GAA ACC TT3′/5′‐GAA AAG GAG TTG GAC TTG GC‐3′, collagen type 6 alpha 3 (COL6A3): AAG TGC CGA TGT TTC CTC AT3′/5′‐TAA TTG AAT CGA GGA GCC CA‐3′. Housekeeping gene expression (ATP5O and G6PD) was measured in each sample. Results are expressed as ΔCt relative to housekeeping gene expression. Graph bars represent mean values of ΔCt values, and error bars are the standard error means (SEM). Only G6PD‐normalized results are shown but both housekeeping genes yielded similar results. TGF‐β1 recombinant treatment and TGF‐β receptor 1 inhibitor Mature adipocytes were counted so that 50 000 cells were seeded in six‐well plates for ceiling culture in 20% serum. At day 4, slides with adherent adipocytes were reversed into a new plate containing 2 mL of 5% serum in each well. At day 5, cells ( n = 7) were treated with 5 ng·mL −1 recombinant human TGF‐β1 or vehicle (0. 1% bovine serum albumin) for 24 h (Ref Cat 100‐21) or with 1 μM SB431542 ( n = 8) (Cat. No 1614), an inhibitor of the TGF‐β receptor ALK5 or with vehicle (dimethylsulfoxide, DMSO 20 mg·mL −1 ) for 24 h. The cells were then harvested in phenol buffer (Cat No. /ID: 79306) for RNA extraction. Three replicates were cultivated for each condition and pooled together at day 6 into phenol buffer for RNA extraction and RT‐PCR quantification. Western blotting and antibodies Proteins were extracted from the organic phase of the RNA phenol/chloroform extraction. First, 100% ethanol was added to the organic phase and incubated for 5 min. After centrifugation (4500 rpm, 2 min, 4 °C), the supernatant was incubated for 10 min with 1. 5 mL isopropanol. The pellet was washed three times with 1. 5 mL ethanol/0. 3 m guanidine with 20‐min incubations and one additional wash without guanidine. The pellet was then incubated at 65 °C in Tris pH 7. 4–6%/SDS until complete dissolution. Sonication was performed as a final disruption step. Protein samples (30 μg) were run on a 10% SDS/PAGE and transferred to nitrocellulose membrane. We used a human TGF‐β1 antibody (RD System, Cat No. AB‐246‐NA) and the Smad2/3 Antibody Sampler Kit (Cell Signaling Technology, Beverly, MA, USA, Cat No. 12747). β‐Tubulin was used as a loading control (Cell Signaling Technology, Cat No. 2146). Densitometric analyses were performed using image j software (NIH, Bethesda, MD, USA). Statistical analyses Statistical analyses were performed using jmp software (SAS Institute Inc, Cary, NC, USA). Expression levels of transcripts were expressed as ΔCT relative to G6PD expression (mean value ± SEM). Comparison of gene expression between the SVF fraction, whole adipose tissue, adipocytes, and DFAT cells was made using ANOVA followed by a Tukey post hoc test. Differences in mRNA expression and protein expression between control and treated cells were tested using matched paired t‐test analyses. Results Gene expression We first measured gene expression of TGF‐β1, TGF‐β2, TGF‐β3, COL1A1, COL1A2, and COL6A3 in whole adipose tissue, SVF cells, and DFAT cells by real‐time quantitative RT‐PCR. As shown in Fig. 2, expression of TGF‐β1, TGF‐β3, and COL1A1 was significantly higher in DFAT cells compared to the SVF ( P = 0. 02, P = 0. 01, and P = 0. 02, respectively), whereas trends were observed for a similar pattern with TGF‐β2, COL1A2, and COL6A3 ( P = 0. 10, P = 0. 08, P = 0. 06, respectively). TGF‐β1, COL1A1, and COL6A3 gene expression was significantly higher in DFAT cells compared to whole adipose tissue ( P = 0. 05, P = 0. 01, P = 0. 02, and P = 0. 03, respectively), and there was a trend for higher expression of COL1A2 in DFAT cells ( P = 0. 08). Figure 2 Expression levels of (A) TGF‐β1, (B) TGF‐β2, (C) TGF‐β3, (D) COL1A1, (E) COL1A2, and (F) COL6A3 in whole adipose tissue, SVF, and DFAT cells (day 12) ( n = 5 donors). Values are mean ± SEM. We then examined the expression of these transcripts during the dedifferentiation process at day 0, corresponding to mature adipocytes, and at days 4, 7, and 12. As shown in Fig. 3, TGF‐β1, COL1A1, and COL6A3 gene expression increased significantly from day 0 to day 12 of the process ( P = 0. 01, P = 0. 02, P = 0. 02). The increase in COL1A2, TGF‐β2 gene expression from day 0 to day 7 was significant ( P = 0. 05, P = 0. 02, and P = 0. 02). A trend was observed for higher expression at day 12 compared to day 0 for COL1A2 ( P = 0. 06). Protein expression of TGF‐β1 was confirmed by western blotting at days 0, 4, 7, and 12 in the SC and OM depots (data not shown). Figure 3 Expression levels of (A) TGF‐β1, (B) TGF‐β2, (C) TGF‐β3, (D) COL1A1, (E) COL1A2, and (F) COL6A3 at various time points during the dedifferentiation process ( n = 3 donors). Values are mean ± SEM. TGF‐β1 treatment Using the six‐well plate culture system, we tested the effect of serum starvation on the cells. However, when adipocytes were plated in the well without serum, they did not adhere to the upper slides (data not shown). We then used DMEM/F12 supplemented with 5% serum to test the effect of 5 ng·mL −1 recombinant TGF‐β1 or vehicle for 24 h on collagen gene expression in DFAT cells from seven patients. As shown in Fig. 4, supplementation with TGF‐β significantly increased TGF‐β1, TGF‐β2, COL1A1, and COL6A3 gene expression compared to 5% serum alone ( P < 0. 05 for all). The treatment had no significant effect on TGF‐β3 and COL1A2 gene expression. Figure 4 Expression levels of (A) TGF‐β1, (B) TGF‐β2, (C) TGF‐β3, (D) COL1A1, (E) COL1A2, and (F) COL6A3 in dedifferentiating adipocytes incubated in media containing 5% serum or 5% serum supplemented with TGF‐β1 (5 ng·mL −1 ) ( n = 7 samples). Values are mean ± SEM. TGF‐β receptor 1 inhibitor We used our six‐well plate model to investigate whether endogenous inhibition of TGF‐β signaling by SB431542, a TGF‐β receptor ALK5 inhibitor, would downregulate collagen transcripts during the dedifferentiation process. At day 5 of the ceiling culture, cells from n = 8 donors were treated with 1 μm SB431542 or with vehicle for 24 h and gene expression was measured by real‐time quantitative RT‐PCR. Fig. 5 shows that treating cells with the inhibitor significantly decreased expression of TGF‐β1, COL1A1, COL1A2, and COL6A3 ( P = 0. 04, P = 0. 04, P = 0. 04, and P = 0. 03, respectively). Trends were observed for a decrease in TGF‐β2 and TGF‐β3 gene expression. Figure 5 Expression of (A) TGF‐β1, (B) TGF‐β2, (C) TGF‐β3, (D) COL1A1, (E) COL1A2, (F) COL6A3 after treatment with SB431542 or vehicle (CTRL) at day 5 of the dedifferentiation process ( n = 8 samples). Values are mean ± SEM. The activation of the serine/threonine kinase pathway by TGF‐β ligands leads to phosphorylation of some members of the intracellular signaling transduction cascade 18. Protein phosphorylation of SMAD 2 and SMAD 3 as well as SMAD 2, SMAD 3, and SMAD 4 levels was measured using western blot analysis in the DFAT cells incubated with or without (vehicle control) 5 ng·mL −1 TGF‐β1. As shown in Fig. 6, treatment with TGF‐β1 significantly increased phosphorylation level of SMAD 2 and SMAD 3 ( P = 0. 001 and P = 0. 03, respectively). Protein levels of SMAD 2, SMAD 3, and SMAD 4 were not affected by TGF‐β treatment. Figure 6 Protein level of (A) SMAD 2, (B) phospho‐SMAD 2, (C) SMAD 3, (D) phospho‐SMAD 3 and (E) SMAD 4 measured in DFAT cells incubated with or without 5 ng·mL −1 recombinant TGF‐β1 (P‐SMAD 2 and SMAD 2, n = 9; P‐SMAD 3, SMAD 3, and SMAD 2, n = 7 samples). Values are mean ± SEM. Protein expression relative to β‐tubulin. Bands from a representative blot are shown. Discussion The aim of this study was to test a new ceiling culture system and modulate the process of dedifferentiation using various incubation conditions targeting the TGF‐β pathway. Adipose tissue expresses various types of collagen other than COL1A1, COL1A2, and COL6A3 19, and the presence of the three subunits of type VI collagen is required for the stable synthesis of collagen VI 20. However, following the results of our previous study demonstrating upregulation of COL1A1, COL1A2, and COL6A3 gene expression during dedifferentiation 21, we chose to examine the effects of TGF‐β on these transcripts specifically. We first demonstrated that TGF‐β1, COL1A1, and COL6A3 gene expression was significantly higher in DFAT cells compared to whole adipose tissue. We also observed a general increase in TGF‐β1, COL1A1, COL1A2, and COL6A3 gene expression over time during dedifferentiation. Using the six‐well plate model, we found that incubation of cells with TGF‐β (5 ng·mL −1 ) during dedifferentiation significantly increased TGF‐β1, TGF‐β2, COL1A1, and COL6A3 gene expression compared to 5% serum alone ( P < 0. 05 for all). Furthermore, when treated with the TGF‐β receptor ALK5 inhibitor, we observed a significant decrease in TGF‐β1, COL1A1, COL1A2, and COL6A3 gene expression during the process, showing that our treatment was effective. Finally, recombinant TGF‐β significantly increased the phosphorylation levels of SMAD 2 and SMAD 3 in DFAT cells. The six‐well plate model of ceiling culture allowed us to treat cells during the dedifferentiation process by targeting and modulating the TGF‐β pathway. The ceiling culture method was first described by Sugihara et al. 4 to study the biology of adipocytes. Ceiling culture has since become the standard strategy to dedifferentiate mature adipocytes. This culture system allows the cells to be maintained in culture for long periods of time and to efficiently proliferate. It has proven useful, but the large number of cells required in flask cultures makes it impossible to study the metabolic, molecular, and secretory attributes of the cells under various incubation conditions. Here, we show that our six‐well plate model could be helpful to understand the physiological process of dedifferentiation and to identify the triggering factors. Some authors have put forward the hypothesis that dedifferentiation is caused by limited gas exchange, high serum concentrations, and cell–plastic contact. However, the six‐well plate model that we have developed allows for gas exchanges, which rules out a predominant effect of hypoxia. On the other hand, the high serum conditions may be an important aspect of the culture because it may contain high concentrations of growth factors. Here, we tested lower serum concentrations (5%) and still observed cell adherence and dedifferentiation. However, when the cells were cultured in serum‐free medium, they could not adhere to the surface. The six‐well plate model represents a relevant approach to examine the impact of various culture environments on the cells. The response of the cells to the treatments with agonists and antagonists of the TGF‐β pathway using the six‐well plate model supports a role for TGF‐β in modulating expression of extracellular matrix components during dedifferentiation. TGF‐β is a well‐known potent inducer of ECM protein‐coding genes such as fibronectin and collagens 22. This multifunctional cytokine has been described as a key factor in matrix remodeling in various physiological and pathological processes 23, 24, 25, 26, 27. However, to the best of our knowledge, we are the first to describe a role for this pathway in human adipocyte dedifferentiation. Due to the limitations in the amount of material, we could not prove a causal impact of TGF‐β/SMAD signaling during dedifferentiation. However, concomitant with the high expression levels of TGF‐β and collagens in DFAT cells, we show that TGF‐β signaling is effective at the end of the process. The increase in SMAD 2/3 phosphorylation following the treatment with recombinant TGF‐β indeed proves that the pathway remains active once the cells are dedifferentiated. We used high concentrations of a soluble and active TGF‐β. Thus, we omitted the essential activation steps of this pathway. Furthermore, active TGF‐β is cleared from the extracellular space if it does not associate rapidly with surface signaling receptors. However, our data demonstrate that the six‐well plate model allows for detailed characterization of the cells regarding a well‐known pathway, and suggest that it could be used to examine other cell programs during dedifferentiation. Dedifferentiation contributes to the generation of a proliferative cell population that could be useful in cellular therapy or tissue engineering. Modulating the secretory, molecular, or metabolic characteristics of the cells could potentially increase the efficiency of the process. The extent of dedifferentiation is difficult to measure and we do not have a recognized marker for this process yet. We have previously used cell size of the remaining mature cells, but we had reported that qPCR was, to date, a more sensitive method to detect short‐term changes in the process 16. Cell size has little sensitivity for short‐term incubations and effects that may not dramatically alter the process. Accordingly, we observed changes in collagen gene expression with SB431542, but we were unable to detect qualitative or quantitative changes in the visual aspects of the cultures. More studies are needed to address whether the six‐well plate model can be used to modulate the extent of dedifferentiation. Some limitations need to be acknowledged. As mentioned, we were unable to address whether TGF‐β can modulate the extent of dedifferentiation. Because TGF‐β1 gene expression was very low at the beginning of the process, it would be surprising that it would induce dedifferentiation, even if a previous study has shown that TGF‐β1 is crucial for the dedifferentiation of cancer cells to cancer stem cells in the context of osteosarcoma 5. It was also demonstrated that TGF‐β1 contributes to the loss of the myofibroblast phenotype 28. As mentioned, TGF‐β1 inhibits adipogenesis 8, 10 and we cannot rule out that the treatment with TGF‐β1 increased the proliferative rate of DFAT cells. The observed increase in TGF‐β signaling and collagen gene expression could reflect cell composition of the culture. As the dedifferentiation process takes place, an increasing number of fibroblasts could contribute to the increase in TGF‐β secretion and signaling in DFAT cells. The changes observed in collagen gene expression following the treatment with TGF‐β1 or with TGF‐β signaling inhibitor may indirectly suggest a modulation of fibroblast proliferation. In conclusion, the six‐well plate culture system will help understand and modulate the dedifferentiation process longitudinally instead of focusing exclusively on the resulting DFAT cells. To the best of our knowledge, this study is also the first to show a role for TGF‐β and the collagens during human mature adipocyte dedifferentiation in vitro. Our findings could potentially contribute to a more extensive characterization of the dedifferentiation process with significant interest for tissue engineering and cell‐based therapy. Author contributions JAC was involved in data acquisition, analysis and interpretation of data, manuscript writing, revision of the manuscript, and final approval. JL and MP were involved in data acquisition, analysis and interpretation of data, revision of the manuscript, and final approval. SM and OL were involved in clinical aspects, sample acquisition, review of the manuscript, and approval. JF was involved in interpretation of data, revision of the manuscript, final approval, and study supervision. AT obtained study funding, was involved in design and conduction of the study, data collection and analysis, interpretation of data, manuscript writing, revision of the manuscript, final approval, and study supervision. Conflict of interest AT receives research funding from Johnson & Johnson Medical Companies for studies unrelated to this manuscript.
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10. 1002/2211-5463. 12317
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FEBS Open Bio
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Coculture of endothelial progenitor cells and mesenchymal stem cells enhanced their proliferation and angiogenesis through
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The beneficial effects of combined use of mesenchymal stem cells ( MSC s) and endothelial progenitor cells ( EPC s) on tissue repair and regeneration after injury have been demonstrated, but the underlying mechanism remains incompletely understood. This study aimed to investigate the effects of direct contact coculture of human bone marrow‐derived EPCs (hEPCs)/human bone marrow‐derived MSCs ( hMSC s) on their proliferation and angiogenic capacities and the underlying mechanism. hEPC s and hMSC s were cocultured in a 2D mixed monolayer or a 3D transwell membrane cell‐to‐cell coculture system. Cell proliferation was determined by Cell Counting Kit‐8. Angiogenic capacity was evaluated by in vitro angiogenesis assay. Platelet‐derived growth factor‐ BB ( PDGF ‐ BB ), PDGF receptor neutralizing antibody ( AB ‐ PDGFR ), and DAPT (a γ‐secretase inhibitor) were used to investigate PDGF and Notch signaling. Cell proliferation was significantly enhanced by hEPC s/ hMSC s 3D‐coculture and PDGF ‐ BB treatment, but inhibited by AB ‐ PDGFR. Expression of cyclin D1, PDGFR, Notch1, and Hes1 was markedly enhanced by PDGF ‐ BB but inhibited by DAPT. In vitro angiogenesis assay showed that hEPC s/ hMSC s coculture and PDGF ‐ BB significantly enhanced angiogenic capacity, whereas AB ‐ PDGFR significantly reduced the angiogenic capacity. PDGF ‐ BB increased the expression of kinase insert domain receptor ( KDR, an endothelial marker) and activated Notch1 signaling in cocultured cells, while DAPT attenuated the promoting effect of PDGF ‐ BB on KDR expression of hEPC s/ hMSC s coculture. hEPC s/ hMSC s coculture enhanced their proliferation and angiogenic capacities. PDGF and Notch signaling pathways participated in the promoting effects of hEPC s/ hMSC s coculture, and there was crosstalk between these two signaling pathways. Our findings should aid understanding of the mechanism of beneficial effects of hEPC s/ hMSC s coculture.
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Abbreviations 2D two‐dimensional 3D three‐dimensional AM adipogenic medium ECs endothelial cells EPCs endothelial progenitor cells FACS fluorescence‐activated cell sorting FITC fluorescein isothiocyanate HBMECs human brain microvascular endothelial cells hEPCs human bone marrow‐derived EPCs HLA‐DR human leukocyte antigen hMSCs human bone marrow‐derived MSCs HUVECs human umbilical vein endothelial cells MSCs mesenchymal stem cells N1ICD Notch1 intracellular domain During tissue repair, adequate angiogenesis is a crucial precondition for success in repair and functional recovery of injured tissue due to the requirement for nutrients and oxygen from blood supply 1, 2. Currently, cell‐based therapeutic strategies have been widely applied in tissue repair and regeneration after injury 3. Endothelial progenitor cells (EPCs) are a group of circulating cells derived from bone marrow, adult peripheral blood, and umbilical cord blood, which can differentiate into mature endothelial cells (ECs), and play important roles in angiogenesis, neovascularization, and vascular endothelial repair 4. Mesenchymal stem cells (MSCs) are multipotent stem cells capable of self‐renew, differentiating, and participating in angiogenesis. Both MSCs and EPCs take part in vascularization and tissue repair and have been widely employed for cell‐based therapy in both preclinical studies 5, 6 and clinical trials 7, 8. Even though many previous studies utilized a single type of cells for transplantation, however, some studies have suggested that coculture or transplantation of more than one type of cells may improve the biological properties of stem cells 9, 10. It has been demonstrated that coculture of human umbilical vein endothelial cells and MSCs can form a vascular tissue‐like network in vitro through the induction of VEGF production 11. When MSCs were cocultured with ECs, the secreted TGF‐β stimulated their differentiation into pericytes and smooth muscle cells, both of which were involved in blood vessel formation 12. Furthermore, cell‐to‐cell interaction promotes rat MSC differentiation into EC via activation of TACE/TNF‐alpha signaling 13. These findings indicated that cell–cell coculture may activate multiple signaling pathways to enhance the biological effects of MSCs. To improve the therapeutic efficacy, therefore, several studies involving a combination of EPCs and MSCs for coculture or cotransplantation have been reported. The purpose of the combined use strategy is to achieve synergistic effects on angiogenesis and tissue regeneration. It has been revealed that direct cell‐to‐cell contact in MSC‐based coculture significantly enhanced the biological properties of MSCs 14. Fu et al. 15 have demonstrated that coculture of peripheral blood‐derived MSCs and EPCs on strontium‐doped calcium polyphosphate scaffolds can enhance osteogenic and angiogenic markers and generate a vascularized engineered bone. Likewise, Li and Wang 16 have reported that canine bone marrow‐derived MSCs and EPCs cocultured in a direct contact coculture system promote osteogenesis and angiogenesis. Very recently, Sun et al. 17 conducted a meta‐analysis including five controlled preclinical studies on cotransplantation in animal models of disease. Their results showed that compared with MSC‐alone group, cotransplantation of EPCs and MSCs significantly enhanced angiogenesis, bone regeneration, vessel revascularization, and tissue repair in cerebrovascular disease model 17. However, although the beneficial effects of combined use of MSCs and EPCs have been demonstrated, the underlying molecular mechanism is still not fully understood. PDGF and Notch signaling pathways have been shown to be involved in the pre‐ and postnatal vasculogenesis and/or angiogenesis 18, 19. PDGF signaling is critical for vascular development and blood vessel homeostasis 20. It has been shown that PDGFs are potent mitogens for mesenchymal cells and are involved in angiogenic induction 21. PDGF signaling is important for differentiation and growth of MSCs 22. PDGF isoforms exert their biological effects through the activation of two tyrosine kinase receptors, PDGFR‐α and PDGFR‐β, which are expressed on MSCs and EPCs 23. Meanwhile, in addition to improving angiogenic functions, exogenous PDGF‐BB also strongly induced the proliferation and migration of MSCs 24. On the other hand, Notch signaling is a highly conserved, cell–cell signaling pathway involved in cell differentiation 25. Binding with Notch ligands such as Delta and Jagged‐1 induces a proteolytic cleavage of the Notch receptors by the γ‐secretase, ultimately leading to transactivation of the promoters of target genes, such a HES and HEY families 26, 27. Notch signaling also mediates intercellular signals that affect proliferation, survival, and differentiation of ECs 28. In angiogenesis, Notch signaling is essential for vascular development and is involved in the determination of arteriovenous vessel fate, whereas alterations in Notch signaling lead to abnormal vascular development at multiple stages 29. Taken together, all these findings suggested that both PDGF and Notch signaling pathways play important roles in the proliferation and angiogenesis of stem cells. Nevertheless, it remains to be investigated whether PDGF and Notch signaling pathways participate in the biological functions of cocultured MSCs and EPCs. To elucidate the molecular mechanism underlying the promoting effect of combined transplantation of EPCs and MSCs on angiogenic capacity, this study aimed to investigate the effects of human bone marrow‐derived EPCs (hEPCs)/human bone marrow‐derived MSCs (hMSCs) coculture on their proliferation and angiogenic capacities in vitro and whether PDGF and Notch signaling pathways play a role in the biological effects of EPCs and MSCs coculture. Utilizing hEPCs and hMSCs in a two‐dimensional (2D) monolayer mixed and 3D transwell membrane cell‐to‐cell coculture systems, the above issue was investigated. Results Characterization of isolated hEPCs and hMSCs hEPCs and hMSCs were isolated from human bone marrow samples. The isolated hEPCs exhibited vascular‐like cells on Day 7, changing toward a spindle‐shaped endothelium‐like morphology and assembling in clusters with cobblestone‐like arrangement on Day 14 (Fig. 1 A). hEPCs were positive for Dil‐acLDL and FITC‐UEA‐I staining with a double positive rate of 92. 7 ± 6. 0% (Fig. 1 B). FACS analysis showed that hEPCs were positive for mesenchymal markers CD133 and CD34 as well as endothelial markers KDR, VE‐cadherin, E‐selectin, and vWF (Fig. 1 C). In addition, the positive cells for KDR, E‐selectin, and vWF significantly increased on Day 14 as compared to Day 7, suggesting a more mature EC phenotype (Fig. 1 D, all P < 0. 05). hEPCs possessed active angiogenic potential in tubule synthesis on Day 14 (Fig. 1 E). Figure 1 Characterization of isolated hEPC s and hMSC s. (A) Microscopic images of hEPC s on Day 7 and Day 14; (B) hEPC s were double‐stained with Dil‐ac LDL and FITC ‐ UEA ‐I and observed by a fluorescence microscopy. Images showed Dil‐ac LDL and FITC ‐ UEA ‐I‐labeled hEPC s; (C) FACS analysis for markers of CD 133, CD 34, KDR, VE ‐cadherin, E‐selectin, and vWF expressions on hEPC s on Day 7 and Day 14; (D) quantification of FACS data of hEPC s on Day 7 and Day 14, n = 6 for each group, # P < 0. 05; (E) HEPC s exhibited angiogenic tubule‐like formation on Day 14; (F) microscopic images of passage 3 hMSC s; (G) FACS analysis for markers of CD 105, CD 73, CD 90, CD 45, CD 14, CD 19, CD 34, and HLA ‐ DR on hMSC s; (H) osteogenic, adipogenic, and chondrogenic differentiation assays for hMSC s on Day 21. hMSC s were stained with ARS, Oil red O, and toluidine blue for calcium deposits, lipid droplets, and proteoglycan, respectively. Passage 3 hMSCs exhibited typical fibroblastic morphology of hMSCs (Fig. 1 F). As shown in Fig. 1 G, hMSCs were positive for mesenchymal markers of CD105, CD73, and CD90, but negative for hematopoietic markers CD45, CD14, CD19, and CD34, and human leukocyte antigen (HLA‐DR). Furthermore, the osteogenic, adipogenic, and chondrogenic differentiation assay on Day 21 showed that the cultured hMSCs were of multilineage differentiation potential (Fig. 1 G). These results suggested that the cultured hMSC possessed characteristics of mesenchymal stem cells. hEPCs/hMSCs coculture and PDGF‐BB enhance proliferation To investigate the effects of hEPCs/hMSCs coculture on the proliferation, a 2D (directly mixed monolayer) coculture system was employed. As shown in Fig. 2 A, compared to hEPCs or hMSCs alone, there was no significant effect on cell proliferation of cocultured cells on Days 3, 6, and 9 (all P > 0. 05). PDGF‐BB treatment significantly enhanced proliferation in single‐cell culture and coculture group, whereas neutralizing antibody against PDGF receptor β (AB‐PDGFR) significantly inhibited proliferation of three groups of cells (all P < 0. 05). Western blot showed that coculture decreased the protein expressions of cyclin D1 as compared with two single‐cell culture groups (both P < 0. 05). Meanwhile, the protein levels of cyclin D1 and PDGFR in three groups were consistently upregulated and downregulated by PDGF‐BB and AB‐PDGFR, respectively (all P < 0. 05, Fig. 2 B). These observations indicated that PDGF signaling was implicated in the proliferation of hEPCs and hMSCs. Figure 2 hEPC s coculture and PDGF ‐ BB enhanced proliferation of hMSC s. (A) Cell viability was determined in a 2D mixed monolayer coculture on Days 3, 6, and 9 as described in Materials and methods. n = 6 for each group, * P < 0. 05, compared to hEPC s alone; # P < 0. 05, compared to hMSC s alone; ▵ P < 0. 05, compared to hEPC s+ hMSC s; (B) protein expression levels of cyclin D1 and PDGFR in all groups on Day 6 were determined by western blot. * P < 0. 05, compared to coculture group, # P < 0. 05, compared to corresponding untreated control; (C) the proliferation of hMSC s on Day 6 in a 3D cell‐to‐cell coculture system, where the hEPC s were cultured on the opposite side of transwell membrane; (D) cell numbers were counted in six random fields (magnification 200×); * P < 0. 05, compared to hEPC s alone; # P < 0. 05, compared to hEPC s+ hMSC s; ▵ P < 0. 05, compared to hEPC s+ PDGF ‐ BB ; ○ P < 0. 05, compared to hEPC s+ AB ‐ PDGFR ; (E) protein expression level of cyclin D1 and PDGFR in all groups on Day 6 was determined by western blot. * P < 0. 05, compared to hMSC s alone, # P < 0. 05, compared to corresponding untreated control. To eliminate the monolayer coculture‐induced contact inhibition in 2D coculture system, a 3D cell‐to‐cell coculture system was used. As shown in Fig. 2 C, D, hEPCs coculture on the opposite side of transwell membrane significantly promoted hMSC proliferation on Day 6, compared to hMSCs cultured alone ( P < 0. 05). The proliferation trends of PDGF‐BB‐ and AB‐PDGFR‐treated cells were consistent with those in 2D coculture. As shown in Fig. 2 E, western blot showed that 3D coculture significantly increased cyclin D1 level ( P < 0. 05). PDGF‐BB significantly promoted cyclin D1 and PDGFR expression in both hMSC‐alone and coculture groups (except for PDGFR in hMSC‐only group), whereas AB‐PDGFR significantly inhibited the expressions of these two proteins (compared to untreated counterparts, all P < 0. 05, Fig. 2 E). PDGF and Notch signaling pathways were involved in the effect of hEPCs/hMSCs coculture on cell proliferation Next, we determined whether Notch signaling pathway plays a role in the molecular mechanism underlying the proliferation‐promoting effect of hEPCs/hMSCs coculture. As shown in Figs 3 A and 2 D coculture significantly increased the expression of Notch1 as compared with the single type of cell culture groups (both P < 0. 05), indicating that Notch signaling pathway was implicated in direct contact culture. To further confirm the involvement of Notch signal in the proliferation of hEPCs/hMSCs coculture, a γ‐secretase inhibitor, DAPT, was used to block Notch signaling. DAPT significantly decreased the expressions of cyclin D1, Notch1, and Hes1 (a transcriptional target of Notch signaling) in hMSC‐only group and the coculture groups (Fig. 3 A, all P < 0. 05). Figure 3 Notch and PDGF signaling participated in the proliferative effects of hEPC s/ hMSC s coculture. (A) In the 2D coculture system, the protein expressions Notch1, Hes1, and cyclin D1 of cells with or without DAPT treatment were determined by western blot. * P < 0. 05, compared to coculture group, # P < 0. 05, compared to corresponding untreated control. (B) Protein levels of PDGFR, Notch1, and Hes1 in hEPC s/ hMSC s coculture on Day 6 in the presence of PDGF ‐ BB or DAPT. * P < 0. 05, compared to untreated control. (C) In the 3D coculture, the protein levels of PDGFR, Notch1, and Hes1 in hMSC ‐alone or coculture groups on Day 6 in the presence of PDGF ‐ BB or AB ‐ PDGFR were determined by western blot. * P < 0. 05, compared to hMSC s group, # P < 0. 05, compared to corresponding untreated control. As our results also revealed that PDGF signaling was implicated in the proliferation of hEPCs and hMSCs, we further investigated whether there was a relationship between PDGF and Notch signaling pathways in cocultured hEPCs and hMSCs. The results showed that in addition to protein level of PDGFR, PDGF‐BB significantly enhanced the expressions of Notch1 and Hes1 (all P < 0. 05, Fig. 3 B). Likewise, when Notch signaling was inhibited with DAPT, expression levels of PDGFR, Notch1, and Hes1 declined coincidently (all P < 0. 05, Fig. 3 B). These findings suggested that there was a relationship between PDGF and Notch signaling pathways in cocultured cells. In the 3D coculture system, hEPCs/hMSCs coculture significantly increased the expressions of PDGFR and Hes1 on Day 6 (both P < 0. 05, Fig. 3 C). In addition, PDGF‐BB further elevated the levels of Notch1, Hes1, and PDGFR in cocultured cells (all P < 0. 05, Fig. 3 C). When PDGFR was blocked, the expression of PDGFR showed no significant differences between hMSC‐alone and cocultured groups ( P > 0. 05, Fig. 3 C). These observations suggest that the activation of PDGF and Notch signaling pathways was associated in 3D coculture cells. hEPCs/hMSCs coculture and PDGF‐BB enhanced angiogenic capacity To determine the effect of hEPCs/hMSCs coculture on their angiogenic capacity, in vitro angiogenesis assay was used. The data of angiogenesis assay at 12, 24, and 48 h are shown in Fig. 4 A, 4 B, and 4 C, respectively, and the tubules quantification data are shown in Fig. 4 D. Compared with hEPCs or hMSCs alone, hEPCs/hMSCs coculture significantly improved capillary‐like formation at all the three time points (all P < 0. 05, Fig. 4 D). Images demonstrated a significant increase in tubule cross‐sectional diameter and junction area in coculture group as compared with either hEPCs or hMSC‐alone group at three time points (Fig. 4 A–C). In addition, PDGF‐BB treatment significantly improved the amount and diameter of tubules in all groups, whereas PDGFR‐β antibody significantly reduced the angiogenic capacity in all groups as compared with their untreated counterparts (all P < 0. 05, Fig. 4 D). These results suggested that hEPCs/hMSCs coculture and PDGF‐BB enhanced angiogenic capacity. Figure 4 hEPC s coculture and PDGF ‐ BB enhanced angiogenic capacity of hMSC s. In vitro angiogenesis was performed in hEPC s, hMSC s, and cocultured cells with or without PDGF ‐ BB or Ab‐ PDGFR at 12 (A), 24 (B), and 48 (C) hours. hMSC s treated with endothelial cell growth medium ( EGM ‐2) were used as a positive control ( EGM ‐2‐ PC ). (D) The average number of tubules was determined in five independent fields at 100× magnification for each well. * P < 0. 05, compared to hEPC ‐alone group; # P < 0. 05, compared to hMSC ‐alone group; ▵ P < 0. 05, compared to coculture group. PDGF and Notch signaling pathways were involved in angiogenesis of cocultured cells As PDGF‐BB enhanced angiogenic capacity, we further investigated whether Notch signaling pathways were also involved in the angiogenic capacity of cocultured cells. As shown in Fig. 5 A, hEPCs/hMSCs coculture and PDGF‐BB markedly increased the levels of KDR (an endothelial marker) and PDGFR (except for KDR in coculture group, all P < 0. 05). Compared to hMSCs or hEPCs alone, hEPCs/hMSCs coculture markedly increased expression of Notch1 protein ( P < 0. 05, Fig. 5 A), suggesting that PDGF and Notch signaling pathways were involved in angiogenesis of hEPCs/hMSCs coculture. We further investigated the relationship between PDGF and Notch signaling pathways. When PDGF signaling was activated by PDGF‐BB, the KDR level elevated in cocultured cells, along with increased activation of Notch1/Hes1 signaling (all P < 0. 05 compared with untreated control, Fig. 5 B). Meanwhile, when Notch signaling was inhibited with DAPT, Notch1 and KDR expression declined significantly as compared with the untreated cells (both P < 0. 05, Fig. 5 B). Furthermore, DAPT attenuated the promoting effect of PDGF‐BB on endothelial differentiation of hEPCs/hMSCs cocultured cells (all P < 0. 05). These findings indicated that both Notch and PDGF signaling pathways participated and crosstalked with each other in the angiogenesis‐promoting effect of hEPCs/hMSCs coculture. Figure 5 PDGF and Notch signaling pathways were involved in angiogenesis of cocultured cells. (A) Protein levels of KDR, Notch1, and PDGFR in endothelial differentiation of hEPC s, hMSC s, and cocultured cells with or without PDGF ‐ BB or Ab‐ PDGFR on Day 3. * P < 0. 05, compared to coculture group. (B) Protein levels of KDR, Notch1, and Hes1 in hEPC s/ hMSC s coculture with or without PDGF ‐ BB and DAPT inhibitor on Day 3. * P < 0. 05, compared to normal‐ DAPT (‐) group. # P < 0. 05, compared to PDGF ‐ BB ‐ DAPT (‐) group. Coculture of hEPCs enhanced the proliferation of hMSCs To accurately evaluate the effect of hEPCs coculture on the proliferation of hMCSs, DiD‐stained hMCSs and DiO‐stained hEPCs were used in the proliferation assay. The results showed that coculture of hEPCs significantly enhanced the proliferation of hMSCs at Day 2, Day 4, and Day 6 as compared with hMCS‐only control (all P < 0. 05, Fig. 6 A). Figure 6 The effect of hEPC s coculture on the proliferation of hMCS s. (A) A coculture of DiD‐stained hMCS s (6 × 10 3 cells) and DiO‐stained hEPC s (6 × 10 3 cells) (coculture group) and a hMCS ‐only group (6 × 10 3 of DiD‐stained hMCS s) were seeded onto a six‐well plate. At Day 2, Day 4, and Day 6, the cells were observed under a fluorescence microscope (three wells for each group). For each well, five fields (at 100×) were randomly chosen and photographed, followed by quantification of cell fluorescence using image j software ( NIH, USA ). (B) Comparisons of in vitro angiogenic capacity between hEPC s/ hMSC s coculture (10 5 or 10 4 of each cell type) and hMSC ‐only (10 5 cells) groups. * P < 0. 05; ** P < 0. 01, compared to MSCs group. As for angiogenic capacity, hEPCs/hMSCs coculture (10 5 of each cell type) group can form better capillary networks as compared with the hMSC‐only (10 5 cells) group at 24 h (Fig. 6 B). Moreover, although the cell numbers were halved, hEPCs/hMSCs coculture (10 4 of each cell type) group remained to form better capillary networks than the hMSC‐only (10 5 cells) group, suggesting that coculture with hEPCs could improve the formation of capillary networks in the hEPCs/hMSCs coculture. Discussion In the current study, we investigated the effects of hEPCs/hMSCs direct contact coculture on their proliferation and angiogenic capacities and their underlying mechanism. The results showed that cell proliferation was significantly improved by hEPCs/hMSCs 3D‐coculture and PDGF‐BB treatment, but significantly inhibited by AB‐PDGFR. The expressions of cyclin D1, PDGFR, Notch1, and Hes1 were significantly enhanced by PDGF‐BB but inhibited by DAPT. In vitro angiogenesis assay showed that hEPCs/hMSCs coculture and PDGF‐BB significantly enhanced angiogenic capacity, whereas AB‐PDGFR significantly reduced the angiogenic capacity in all groups. PDGF‐BB increased endothelial marker KDR expression and activated Notch1 signaling in cocultured cells, while DAPT attenuated the promoting effect of PDGF‐BB on endothelial differentiation of hEPCs/hMSCs coculture. Taken together, these findings suggested that hEPCs/hMSCs coculture enhanced their proliferation and angiogenic capacities and both PDGF and Notch signaling pathways participated and crosstalked with each other in these promoting effects. To our best knowledge, this is the first study reporting the roles of PDGF and Notch signaling pathways in promoting effects of hEPCs/hMSCs coculture on their biological functions. Our proliferation data showed that hEPCs/hMSCs coculture significantly improved proliferation in 3D transwell chamber but not in 2D mixed monolayer coculture system, which is line with previous studies 30, 31. One possible explanation may be that in the 2D‐coculture, the direct contact between hMSCs and hEPCs induced a contact inhibition. However in the 3D‐coculture, the contact inhibition was prevented due to the complete separation between hEPCs and hMSCs by the membrane in transwell chamber 30. In addition, the membrane pores allowed a variety of hEPCs/hMSCs secreting growth‐promoting factors to pass through the membrane, leading to enhanced proliferation. MSCs represent a promising cell source for angiogenic therapies due to their capacity to differentiate into ECs and to form capillary networks 32, 33. Consistent with previous reports 15, 16, our in vitro angiogenesis assay showed that coculture of hEPCs and hMSCs markedly increased the thickness of capillary‐like structure and junction area at all three time points as compared with hMSC‐alone group. Particularly, at 48 h, when cell death was observed in both hEPCs and hMSC‐alone groups, cocultured cells remained to maintain good tubule morphology. These observations suggest that hEPCs/hMSCs coculture not only promoted the tubule formation but also prolonged cell survival, which can further improve the angiogenic potential. Our data showed that PDGF‐BB significantly enhanced cell proliferation through upregulation of PDGFR‐β and cyclin D1. In addition, PDGF‐BB played a synergistic role in hEPCs/hMSCs coculture to further enhance proliferation, which is in agreement with other reports 34, 35. Meanwhile, when PDGFR‐β was blocked, cell proliferation, as well as the levels of PDGFR‐β and cyclin D1, declined coincidently, demonstrating that PDGF signaling was involved in the proliferation. The previous study has reported the enhancing effect of PDGF‐BB on vascularization 36. Our in vitro angiogenesis data showed that PDGF‐BB and hEPCs/hMSCs coculture also exhibited a synergistic effect on the angiogenic capacity. The promoting effects of hEPCs/hMSCs coculture on proliferation and angiogenesis may be attributed to the fact that both EPCs and MSCs secrete substantial amounts of PDGF‐BB 37, 38. Notch signaling has been reported to participate in the proliferation and angiogenesis 28. Our data showed that direct contact coculture markedly increased Notch1 expression, and DAPT inhibited expression of cyclin D1. Meanwhile, the expression of Notch1 significantly increased in endothelial differentiation of cocultured cells. DAPT decreased the level of KDR in endothelial differentiation of cocultured cells, indicating suppression of cell endothelial differentiation. These data supported that Notch signaling pathway was implicated in the promoting effects of hEPCs/hMSCs coculture on the proliferation and angiogenesis. The association between PDGF and Notch signaling pathways has been reported previously. PDGFR‐β has been shown to be a target of Notch signaling gene in vascular smooth muscle cells 20. It has been reported that Notch1 signaling is an upstream of PDGF‐B transcription in human brain microvascular endothelial cells 39. On the other hand, PDGF‐BB can affect Notch1 activation and Notch1–Furin interaction 40. These findings suggest that crosstalk between PDGF and Notch signaling pathways is common in cell biological regulations. The current study showed that when PDGF signaling was activated in hEPCs/hMSCs coculture, Notch1/Hes1 levels also elevated simultaneously, resulting in an improvement of cell proliferation and endothelial differentiation. Coincidentally, when PDGF signaling was blocked, Notch1/Hes1 levels were also reduced, leading to inhibited cell growth and angiogenesis. On the other hand, DAPT attenuated the promoting effect of PDGF‐BB on endothelial differentiation of hEPCs/hMSCs cocultured cells. Taken together, these observations indicated that there was a crosstalk between PDGF and Notch signaling pathways in the cell proliferation and endothelial differentiation of cocultured hEPCs/hMSCs. There are still some limitations in this study. First, in the coculture system, we cannot distinguish the individual contribution of hMSCs/hEPCs both in the western blot and in the angiogenic assays because we cannot conduct cell separation before assays. However, we can partially evaluate the individual contribution of hMSCs/hEPCs by comparing the western blot/angiogenic assay data between the coculture group and the hMSCs or hEPCs single culture group. In addition, we did not use Notch signaling activators to comprehensively evaluate the role of Notch signaling in hEPCs/hMSCs coculture. The angiogenic potential was assessed only in the 2D‐coculture, but not in the 3D‐coculture system. Furthermore, the in vitro findings of this study remain to be further validated using an in vivo model. All these limitations should be addressed in the following study. In summary, our results showed that hEPCs/hMSCs coculture demonstrated significant enhancement effects on cell proliferation and angiogenic capacity in direct contact coculture. These promoting effects were involved in the crosstalk between PDGF and Notch signaling pathways. Our findings may be useful for the development of future applications in tissue engineering and therapeutic strategy. Materials and methods Isolation and culture of hEPCs and hMSCs Bone marrow was collected from the drill holes of pedicle during the operations of patients with spine internal fixation (age range 22–56 years; mean age 43 years) with lumbar degenerative diseases (degenerative lumbar spondylolisthesis and lumbar spinal stenosis with instability). Informed consent was obtained from the patients for bone marrow collection, and all the procedures were performed in accordance with the guidance and approval of a research ethics committee in the First Affiliated Hospital of Sun Yat‐sen University. Mononuclear cells were collected by Ficoll density gradient centrifugation (1. 077; GE Health, Fairfield City, CT, USA) from the bone marrow at 300 g for 25 min. The nucleated cells were collected from the defined layer at the interface, diluted with two volumes of PBS, centrifuged twice at 100 g for 5 min, and finally resuspended in basal medium. For the isolation of hEPCs, the collected cells were cultured in dishes coated with fibronectin and induced by EGM‐2 MV Single‐Quots (Cambrex, East Rutherford, NJ, USA) at 37 °C with 5% CO 2 in humidified air at a density of 5 × 10 5 cm −2. After three days, nonadherent cells were washed out with PBS and cultured to Day 14. At Day 7 and Day 14, immunofluorescence staining and flow cytometry were applied to identify hEPCs. Quantitative fluorescence‐activated cell sorting (FACS) was performed on a FACS Vantage SE flow cytometer (Becton Dickinson, Lake Franklin, NJ, USA). The ability of tube formation by hEPCs was determined by in vitro angiogenesis assay. For the isolation of hMSCs, the collected mononuclear cells were resuspended in basal medium at a density of 2 × 10 5 cm −2 and maintained in a humidified atmosphere of 95% air and 5% CO 2 at 37 °C. Basal medium consisted of Dulbecco's modified Eagle's medium with Glutamix‐1, sodium pyruvate, 4500 mg·L −1 glucose, and pyridoxine (DMEM; Gibco, BRL, Gaithersburg, MD, USA) supplemented with 10% inactivated fetal bovine serum (FBS; Gibco, BRL) and 1% penicillin/streptomycin (Sigma, St. Louis, MO, USA). After three days, the medium was changed to remove all nonadherent cells. Thereafter, the medium was changed twice a week until subconfluence. When 80% confluent, the cells were detached using 0. 125% trypsin/5 m m ethylenediaminetetraacetic acid (Sigma), then placed in the basal medium and expanded with 1 : 3 ratio. The third passage of hMSCs was used in all the experiments. Flow cytometry was used to identify hMSC phenotypes at all samples. The capacity of multilineage differentiation of hMSCs, including osteogenesis, chondrogenesis, and adipogenesis, was detected for further identification at Day 21. The first‐passage (P1) hEPCs and the third‐passage (P3) hMSCs were used in all the following experiments. Double staining for hEPCs Passage 1 hEPCs were incubated with 2. 5 μg·mL −1 of 1, 1′‐dioctadecyl 3, 3, 3′, 3′‐tetra‐methylindo‐carbocyanine‐labeled acetylated low‐density lipoprotein (DiI‐acLDL; Invitrogen, Carlsbad, CA, USA) and 10 μg·mL −1 of fluorescein isothiocyanate (FITC)‐conjugated factor VIII, ulex europaeus agglutinin‐1 (Sigma) for 3 h at 37 °C, and the cells were examined under a fluorescence microscope (Nikon, Tokyo, Japan). Flow cytometric analysis hEPCs were characterized for immunophenotype using monoclonal antibodies (MoAbs) specific for CD133, CD34, KDR, vWF, E‐selectin, and VE‐cadherin at Days 7 and 14. hMSCs were incubated specifically with CD105, CD73, CD29, CD44, CD45, CD14, HLA‐DR, and CD90 antibodies at the third passage. All antibodies were purchased from Pharmingen/Becton Dickinson (PharMingen, San Diego, CA, USA). Cells were detached using trypsin/EDTA for 5 min, immediately washed with PBS to remove trypsin, and resuspended at 10 6 mL −1. Cell suspension (100 mL) was incubated at 4 °C for 10 min with 15% FBS, followed by incubation with the specific antibody at 4 °C for 30 min. Cells were washed with PBS. At least 10 000 events were analyzed by flow cytometry (FACScali‐bur; Becton Dickinson, Milan, Italy) using cell quest software. Osteogenic, adipogenic, and chondrogenic differentiations of hMSCs For osteogenic differentiation, hMSCs at the third passage were seeded at a concentration of 0. 8 × 10 4 cm −2 in a six‐well plate. When confluent, the cells were detached using 0. 125% trypsin/5 m m EDTA and then placed in the basal medium. After 1‐day culture, the medium was replaced with osteogenic medium consisting of basal medium supplemented with 10 n m dexamethasone, 10 m m β‐glycerol‐phosphate, and 0. 2 m m ascorbic acid (all from Sigma). At Day 21, cells were fixed with ice‐cold 70% ethanol for 30 min, washed with PBS for three times, and stained with Alizarin Red S (40 m m, PH 4. 2; Sigma) for 30 min, and rinsed with PBS for three times. The dish area was observed with a light microscope. For adipogenic differentiation, cells were seeded at a concentration of 2. 5 × 10 4 cm −2 in a six‐well plate. Adipogenic differentiation was induced with adipogenic medium, containing DMEM, 10% FBS, 10 −6 m dexamethasone, 0. 2 m m indomethacin, 10 μg·mL −1 insulin, and 100 ng·mL −1 3‐isobutyl‐L‐methylxanthine (all from Sigma Immunochemicals, Sigma‐Aldrich Co. , St. Louis, MO, USA). At Day 21, cells were examined for the presence of lipid vacuoles using Oil Red O staining. Briefly, cells were fixed in 10% formaldehyde in phosphate buffer for 1 h, washed with 60% propylene glycol for 3 min, stained with 0. 18% Oil Red O (Sigma) for 10 min, rinsed with water, and counterstained with hematoxylin for 10 min. For chondrogenic differentiation, hMSCs (2. 5 × 10 5 ) were centrifuged in a 15‐mL polypropylene Falcon tube to form a pellet. Chondrogenic medium consists of DMEM, 10% FBS, 37. 5 mg·mL −1 ascorbic acid, 10 n m dexamethasone, 1 : 100 ITS premix (BD Biosciences, San Jose, CA, USA), and 10 ng·mL −1 human recombinant TGF‐β3 (R&D Systems, Minneapolis, MN, USA). The medium was changed twice a week. Presence of proteoglycan (PG) was analyzed by means of 0. 1% toluidine blue staining (Sigma). Determining the proliferation in hEPC/hMSC coculture Cell proliferation was determined by 2‐(2‐methoxy‐4‐nitrophenyl)‐3‐(4‐nitrophenyl)‐5‐(2, 4‐disulfo‐phenyl)‐2H‐tetrazolium, monosodium salt (WST‐8) assay kit (CCK‐8; Dojindo, Mashikimachi, Japan). Briefly, WST‐8 was added to each well for 4 h before the measurement. The absorbance at 450 nm was measured using a microplate reader (Thermo Fisher Scientific, Waltham, MA, USA). For PDGF‐BB (PeproTech, Rocky Hill, NJ, USA) treatment, cells were cultured in 2% FBS DMEM medium containing 2 ng·mL −1 PDGF‐BB for 6 days. To block the PDGF receptors, 20 μg·mL −1 of neutralization antibody against PDGFR‐β (R&D Systems) was used in corresponding groups. To inhibit Notch signaling, 50 μmol·L −1 DAPT (γ‐secretase inhibitor, dissolved in dimethyl sulfoxide, purchased from Sigma) was used. Two kinds of coculture methods were utilized in this assay (Fig. 7 ). The first coculture system was a traditional mixed monolayer (2D) system. Briefly, hEPCs and hMSCs were mixed (1 : 1 ratio) in DMEM supplemented with 10% FBS and 1% penicillin/streptomycin (Invitrogen) and seeded in 96 wells. At Days 3, 6, and 9, CCK‐8 was used for cell proliferation detection according to the manufacturer's instructions. The second coculture system was a three‐dimensional (3D) cell‐to‐cell method using a six‐well culture plate and inserts (Corning, Sullivan Park, NY, USA), containing a polyethylene terephthalate track‐etched membrane with 0. 4‐μm pores at the bottom of inserts. hMSCs were seeded onto the membrane of each culture insert that has 4. 2 cm 2 of available culture area at 1 × 10 4 cells. hEPCs were cocultured with hMSCs through direct cell‐to‐cell contact in monolayer on the opposite side of the membrane, the bottom of the culture insert. After culture for 6 days, the average number of cells in ten random microscopic 200× fields was determined manually. Figure 7 A schematic flow diagram of this study. hEPC s and hMSC s were isolated from human bone marrow. The isolated hEPC s were characterized with in vitro angiogenic potential and the specific markers. The isolated hMSC s were characterized with the multilineage differential potential and the specific surface markers. Then hEPC s and hMSC s were cocultured in a two‐dimensional (2D) monolayer mixed or 3D transwell membrane cell‐to‐cell coculture systems. The proliferation and angiogenic capacities of cells were assessed, and the underlying mechanism was investigated. To accurately evaluate the effect of hEPCs coculture on the proliferation of hMCSs, hMCSs and hEPCs were stained with DiD and DiO fluorescent dyes (Thermo Fisher Scientific) for membrane labeling, according to the manufacturer's protocol, and were used in the proliferation assay. A coculture of DiD‐stained hMCSs (6 × 10 3 cells) and DiO‐stained hEPCs (6 × 10 3 cells) (coculture group) was seeded onto a six‐well plate. Meanwhile, a hMCS‐only group (6 × 10 3 of DiD‐stained hMCSs) was also seeded onto a six‐well plate and incubated in 37 °C incubator. At Day 2, Day 4, and Day 6, the cells were observed under a fluorescence microscope (three wells for each group). For each well, five fields (at 100×) were randomly chosen and photographed, followed by quantification of cell fluorescence using image j software (NIH, USA). Western blot analysis In order to further investigate cell proliferation, cocultures of hMSCs and hEPCs were harvested and total proteins were extracted for western blot analysis at Day 6. hEPCs, hMSCs, and coculture cells were collected and lysed with ice‐cold lysis buffer (Merck, Novagen, Temecula, CA, USA). Protein concentrations were determined with the Bio‐Rad assay system (Bio‐Rad, Hercules, CA, USA). Thirty micrograms of total proteins was separated on a 10% SDS/polyacrylamide gel. Bands on the gels were transferred onto polyvinylidene fluoride membranes (Millipore Corp, Billerica, MA, USA). The membranes were blocked for 1 h with 5% nonfat milk or BSA in PBS with 0. 1% Tween 20. Blots were incubated with primary antibody overnight at 4 °C followed by secondary antibodies for 1 h each at room temperature. Immunoreactive bands were visualized with Chemiluminescence Reagent Plus (Thermo Fisher Scientific) and exposed to X‐ray film (Fujifilm, JP). The membranes were then incubated with stripping buffer (Pierce, Holmdel, NJ, USA) for 30 min at 37 °C, reblocked, and reprobed with β‐actin as a loading control. The proteins were detected with specific cyclin D1, Notch1 intracellular domain (N1ICD, cat. no. ab83232), Hes1, KDR (1 : 1000, all above from Abcam, Cambridge, UK) and PDGF receptor β antibodies (Ab‐PDGFR, 1 : 1000, from Cell signaling Technology, Danvers, MA, USA). Western blot bands were quantitated using quantity one software (Bio‐Rad, USA). In vitro angiogenesis assay for hEPC/hMSC coculture Angiogenesis assay for capillary‐like tube formation was performed with an In Vitro Angiogenesis Assay Kit (Chemicon, Temecula, CA, USA) according to the manufacturer's instructions. Briefly, ECMatrix solution was thawed on ice overnight, mixed with 10× ECMatrix™ diluents, and placed in a 96‐well tissue culture plate at 37 °C for 1 h to allow the matrix solution to solidify. After starved for 12 h, 1 × 10 4 cells containing hEPCs and hMSCs mixed with a 1 : 1 ratio were seeded on the top of the solidified matrix solution in each 24‐well with 2% FBS DMEM. At 12, 24, and 48 h, tubule formation was inspected under an inverted light microscope at 100× magnification. Tubule formation was defined as a structure exhibiting a length four times its width. Five independent fields were assessed for each well, and the average number of tubules/100× fields was determined. For endothelial differentiation in coculture, hEPCs and hMSCs were mixed (1 : 1 ratio) in EC growth medium (EGM‐2) supplemented with 2% FBS and seeded in 1 μg·mL −1 fibronectin‐coated wells. For control conditions, non‐cocultured cells in the same medium were used. After 3‐day endothelial induction, hMSCs and hEPCs coculture were harvested and total proteins were extracted for western blot analysis. Statistical analysis All values are presented as mean ± standard deviation of the mean. Data were analyzed using the Student's t ‐test or a general linear model two‐way ANOVA with a post hoc Tukey test to compare between groups as appropriate. The value of P < 0. 05 was considered statistically significant. All analyses were performed using ibm spss version 17 (IBM Corporation, Somers, NY, USA). Author contributions We declare that all the listed authors have participated actively in the study and all meet the requirements of the authorship. DS and TL designed the study and wrote the protocol. TL, LZ, and WG performed research/study. LZ contributed important reagents. MG and JR managed the literature searches and analyses. HY and KW undertook the statistical analysis. DS and TL wrote the first draft of the manuscript.
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10. 1002/2211-5463. 12634
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FEBS Open Bio
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Blueberry juice protects osteocytes and bone precursor cells against oxidative stress partly through
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Oxidative stress and abnormal osteocyte apoptosis are often related to dysregulation of bone turnover and chronic bone loss, and so fruit and vegetables with high antioxidant potential may play an important role in the prevention and/or management of osteoporosis. Osteocytes are the main regulators of bone remodelling. For the first time, we demonstrate here that blueberry juice ( BJ ), obtained from Vaccinium myrtillus, rich in polyphenols, shows antioxidant and antiosteoclastogenic properties in MLO ‐Y4 osteocytes. We report that BJ prevents oxidative stress‐induced apoptosis and reverses the increase in receptor activator of nuclear factor κB ligand and sclerostin expression, crucial factors for osteoclast activation and bone resorption. BJ is also able to prevent oxidative stress‐induced cell cytotoxicity in bone marrow mesenchymal stromal cells ( MSC s), which are considered to be an important tool for cell therapy in bone disorders. No significant difference in preventing these events was observed between BJ and blueberry dry extract containing equal amounts of total soluble polyphenols. We have also shown that blueberry acts as both an antioxidant and an activator of sirtuin type 1, a class III histone deacetylase involved in cell death regulation and considered a molecular target for blocking bone resorption without affecting osteoclast survival. Overall, these novel data obtained in osteocytes and MSC s may help us clarify the mechanisms by which blueberry counteracts oxidative stress‐induced damage in bone remodelling and osteogenesis at the cellular and molecular level. Our findings are consistent with the reported beneficial effects of blueberry on bone tissue reported in animal studies, which suggest that blueberry may be a useful supplement for the prevention and/or management of osteoporosis and osteogenic process.
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Abbreviations Ac‐p53 acetylated‐p53 BB blueberry BE blueberry dry extract BJ blueberry juice ERK1/2 extracellular signal‐regulated kinase 1/2 FBS fetal bovine serum H 2 DCFDA 2′, 7′‐dichlorodihydrofluorescein diacetate JNK c‐Jun N‐terminal kinase MSC marrow mesenchymal stromal cell MTT 3‐(4, 5‐dimethylthiazol‐2‐yl)‐2, 5‐diphenyl‐tetrazolium bromide NBT Nitroblue tetrazolium PARP‐1 poly‐(ADP‐ribose) polymerase 1 RANKL receptor activator of nuclear factor κB ligand ROS reactive oxygen species SIRT1 sirtuin type 1 TSP total soluble polyphenols Fruits and vegetables rich in phytochemicals, such as phenolic acids, are essential for bone formation and health 1, 2, 3, and their daily consumption may be important in increasing the bone mass peak 4, 5. These antioxidant compounds scavenge reactive oxygen species (ROS) and reduce oxidative stress in many diseases 3, 5, 6 including bone diseases and, in particular, osteoporosis 7, 8, 9. This latter disease is mainly due to abnormal activation of osteoclasts, inhibition of osteoblast activity and an increase of osteocyte apoptosis due to oxidative stress, which may be the result of oestrogen deficiency and/or other bone diseases 10, 11, 12. Indeed, redox balance regulation in bone cells affects bone turnover and remodelling 13, 14, 15, and microdamage, oxidative stress, and abnormal osteocyte apoptosis are related to an imbalance of the remodelling process with the consequent altered bone formation and a decrease in mineral density 16, 17, 18. The osteocytes are in close contact with the blood flow and contribute to the transfer of endogenous and exogenous compounds to osteoblasts and osteoclasts, regulating their formation and activity 16, 19. Moreover, increased production of ROS damages stem cell self‐renewal and differentiation toward tissue‐specific lineages including bone tissue 20, 21, 22, 23. Given their antioxidant activity, natural polyphenols act as powerful modulators of mesenchymal stromal stem cells, which are considered a useful tool for studying bone regeneration processes and bone tissue engineering treatments 20, 21, 24. Effectively, nutritional approaches to antioxidant strategies can be useful for the treatment and the prevention of bone loss 1, 2, 3, 4, considering the possible adverse side effects associated with anti‐resorptive drugs, such as bisphosphonate or oestrogen therapy. In fact, these therapies reduce significantly bone loss and the osteoclastogenesis but do not restore a normal bone remodelling process 14, 25. In view of this, there is a rising demand for use of natural products as adjuvant therapy or supplementation in combination with medical therapy to promote the restoration of normal bone metabolism. However, there are few data that correlate the antioxidant properties of plants containing bioactive phytochemicals to molecular mechanisms of biological processes involved in the modulation and regulation of bone formation and regeneration in the condition of oxidative stress. In particular, no data are reported on the protective effects of blueberry (BB) phytochemicals against oxidative stress‐induced damage on osteocyte activity and on bone precursor cell viability. Indeed, we previously demonstrated in a murine osteocyte‐like cell line, MLO‐Y4, with similar phenotype and many characteristics of mature osteocytes 26, 27, that thiol antioxidants such as glutathione, N ‐acetylcysteine and lipoic acid prevent and down‐regulate both apoptosis and osteoclastogenic factors induced by intracellular oxidative stress 18. BBs and, in particular, Vaccinium mirtillus, contain high concentrations of well‐characterized anthocyanins, phenolic acids, coumarins, flavonols, flavanols, and phenolic compounds with antioxidant activity 28, 29. Recently, a role for the bioactive compounds contained in Vaccinium mirtillus as ‘functional food' for dietary supplementation has been suggested 28, 29. Studies on animal and cell models reveal a positive association of high intake of BB polyphenols with antioxidant properties and high bone mass 30. Indeed, diet‐derived phenolic acids promote bone growth and regulate osteoblast and adipocyte lineage commitment and differentiation in young mice 2, 23, and diets containing BBs prevent osteoporosis in ovariectomized rats 31, 32. However, the molecular mechanisms through which these act are still little known. The aim of this study was to evaluate the ability of natural compounds contained in BB juice (BJ) and BB dry extract (BE) to preserve osteocyte activity and bone precursor cell regeneration in the presence of oxidative stress, and to identify possible biological mechanisms and targets on which BB phytochemicals can act to stimulate bone formation and to maintain normal bone remodelling in bone diseases related to oxidative stress. For this study, MLO‐Y4 osteocyte‐like cells and bone mesenchymal stromal stem cells (MSCs) were used. MLO‐Y4 constitutes a model to study in vitro osteocyte viability and apoptosis in response to microdamage and bone diseases 26, 27, 33, whereas MSCs are considered an important tool for cell therapy in bone disorders due to their ability to differentiate into various tissues including bone tissue 20, 21. The results demonstrate both in osteocytes and in MSCs, cultured in serum deprivation, that BJ and BE are able to reduce ROS levels and to prevent apoptosis and cytotoxicity due to oxidative stress. Moreover, in starved osteocytes they prevent the up‐regulation of receptor activator of nuclear factor κB ligand (RANKL) and sclerostin, osteoclastogenic factors related to apoptosis and bone resorption. The effects of BJ and BE are in part mediated by activity of SIRT1, which has been proposed as a potential target to restore a normal bone remodelling process and for anabolic therapies against excessive bone resorption in osteoporosis. Results Effect of BJ and BE on ROS production in starved MLO‐Y4 cells and in cell‐free model In MLO‐Y4 cells, oxidative stress was induced by serum deprivation (starved cells), and two different BB preparations, BJ and BE, were used given that BBs are commercialised in different ways, mainly as fresh or frozen products but also as juice or dry extract. Previously, we demonstrated a remarkable increase of ROS after 4 and 24 h in starved MLO‐Y4 cells 18, as reported in the present study in Fig. 1 A. In these experimental conditions, the antioxidant effect of BJ containing various concentrations (from 7. 5 to 60 μg·mL −1 ) of total soluble polyphenols (TSP) was measured. Figure 1 A shows that the lowest concentrations (7. 5–15 μg·mL −1 ) reduced significantly ROS levels after 4 h by about 25% and the highest concentrations (30–60 μg·mL −1 ) by about 50%, as compared to starved cells. ROS reduction elicited by BJ treatments after 24 h significantly and gradually increased from 25% to 50%, reaching the maximum effect at 30 μg·mL −1 TSP (Fig. 1 A). Next, we compared the BJ antioxidant effect to that of BE at this concentration of TSP. As shown in Fig. 1 B, no difference was observed between BJ and BE after both 4 and 24 h of treatment. Effectively, BJ and BE also showed a similar antioxidant capacity when superoxide anion radical scavenging activity was measured in a cell‐free model using the same concentration of TSP (30 μg·mL −1 ) (Fig. 1 C). Figure 1 Antioxidant effect of BJ and BE on intracellular ROS in MLO ‐Y4 cells and in a cell‐free model. (A, B) Intracellular ROS, detected by measuring the fluorescence intensity of the probe 2′, 7′‐dichlorodihydrofluorescein diacetate (H 2 DCFDA ), were measured in MLO ‐4Y cells cultured for 4 and 24 h in complete medium (C, control) or in serum‐free medium (S, starved cells). Starved cells were treated or not with BJ at various concentrations (μg·mL −1 ) of total soluble polyphenols ( TSP ) (A), or with 30 μg·mL −1 TSP of BJ or BE (B), as reported in Materials and methods. (C) The xanthine/xanthine oxidase system was used for O 2 − · production and nitroblue tetrazolium ( NBT ) was used as target for the detection of scavenging activity of O 2 − · by BJ and BE in a cell‐free model, as reported in Materials and methods. In (A, B), ROS data, normalized on total protein content, are expressed as fold‐increase over the respective control values and are the mean ± SEM of four experiments performed in duplicate. In (C), O 2 − · scavenging activity is expressed as absorbance arbitrary units (A. U. ) and the data are the mean ± SEM of three experiments performed in duplicate. Data were evaluated by using one‐way ANOVA followed by Bonferroni's post hoc test. □ P ≤ 0. 001 compared to 4 h untreated starved cells; * P ≤ 0. 05; ** P ≤ 0. 001 compared to the respective untreated starved cells; ○ P ≤ 0. 05 compared to 30 and 60 μg·mL −1 TSP treated cells; ∆ P ≤ 0. 05 compared to 24 h 7. 5 μg·mL −1 TSP treated cells; ● P ≤ 0. 01 compared to control cells. Effect of BJ and BE on oxidative stress‐induced RANKL and sclerostin expression and apoptosis in starved MLO‐Y4 cells The effect of BJ and BE on oxidative stress‐induced overexpression of osteoclastogenic factors, such as RANKL and sclerostin, was investigated. Figure 2 A shows that RANKL and sclerostin up‐regulation due to oxidative stress in starved MLO‐Y4 cells 18 was prevented by the treatment with 30 μg·mL −1 TSP of BJ or BE, the concentration at which the maximum antioxidant effect was obtained. Since RANKL and sclerostin overexpression have been correlated to c‐Jun N‐terminal kinases (JNK) and extracellular signal‐regulated kinases 1/2 (ERK1/2) activation 18, we evaluated the effect of BJ and BE on the oxidative stress‐induced activation of these kinases. Phosphorylation of both JNK and ERK1/2 was significantly reduced confirming the ability of BJ and BE to counteract these effects (Fig. 2 B). Cell apoptosis, induced by oxidative stress after 24 h of starvation as previously reported 18, was assayed by detecting oligonucleosomes and phosphatidylserine on the external surface of the plasma membrane. Both methods of analysis showed that the treatment with 30 μg·mL −1 TSP of BJ or BE significantly prevented apoptosis, which decreased by about 70–80% as compared to starved cells (Fig. 3 A, B). The protective effect was confirmed by the decrease of 17 kDa caspase‐3 active form that derives from proteolytic cleavage of inactive 32 kDa procaspase‐3 induced by oxidative stress (Fig. 4 A). Similarly, the levels of cleaved poly‐(ADP‐ribose) polymerase 1 (PARP‐1), a natural substrate of caspase‐3, decreased in cells treated with BJ or BE (Fig. 4 B). Altogether, these results show that BB treatments are effectively able to prevent osteocyte apoptosis. Figure 2 Effect of BJ and BE on RANKL and sclerostin expression, and JNK and ERK phosphorylation in MLO ‐Y4 cells. RANKL and sclerostin expression, and ERK and JNK phosphorylation were measured in MLO ‐Y4 cells cultured for 24 h in complete medium (C, control) or in serum‐free medium (S, starved cells). Starved cells were treated or not with 30 μg·mL −1 of total soluble polyphenols ( TSP ) of BJ and BE, as reported in Materials and methods. RANKL and sclerostin (A), p‐ ERK and p‐ JNK (B) were measured by western blot analysis and values are normalized with β‐actin, ERK and JNK bands obtained by densitometric analysis, respectively. Blots are representative of four experiments and data, expressed as fold‐increase over the respective control, are reported as mean ± SEM at the bottom. Data were evaluated by using one‐way ANOVA followed by Bonferroni's post hoc test. * P ≤ 0. 01; ** P ≤ 0. 001 compared to the respective control, BJ ‐ and BE ‐treated cells. Figure 3 Effect of BJ and BE on apoptosis in MLO ‐Y4 cells. Apoptosis was measured in MLO ‐4Y cells cultured for 24 h in complete medium (C, control) or in serum‐free medium (S, starved cells). Starved cells were treated or not with 30 μg· mL −1 of total soluble polyphenols ( TSP ) of BJ and BE, as reported in Materials and methods. Apoptosis data, relative to mono‐ and oligonucleosomes released into the cytoplasmic fraction from 10 4 cells (A), or relative to phosphatidylserine on the external plasma membrane (B), are expressed as fold‐increase over control values and are the mean ± SEM of four experiments performed in duplicate. Data were evaluated by using one‐way ANOVA followed by Bonferroni's post hoc test. * P ≤ 0. 05 compared to the respective control cells; ** P ≤ 0. 001 compared to the respective control, BJ ‐ and BE ‐treated cells. Figure 4 Effect of BJ and BE on active 17 kDa caspase‐3 and PARP ‐1 cleavage in MLO ‐Y4 cells. Active 17 kDa caspase‐3 and PARP ‐1 cleavage were measured in MLO ‐4Y cells cultured for 24 h in complete medium (C, control) or in serum‐free medium (S, starved cells) by western blot analysis. Starved cells were treated or not with 30 μg· mL −1 of total soluble polyphenols ( TSP ) of BJ and BE, as reported in Materials and methods. Densitometric analysis of active 17 kDa caspase‐3 (A) and cleaved PARP ‐1 (B) values were normalized with β‐actin bands. Blots are representative of three experiments and data, expressed as fold‐increase over control, are reported as mean ± SEM at the bottom. Data were evaluated by using one‐way ANOVA followed by Bonferroni's post hoc test. * P ≤ 0. 001; ** P ≤ 0. 05 compared to the respective control, BJ ‐ and BE ‐treated cells. Involvement of SIRT1 in antiosteoclastogenic and antiapoptotic effect of BJ and BE in starved MLO‐Y4 cells This study investigated also the role of sirtuin type 1 (SIRT1), a class III histone deacetylase involved in the regulation of apoptosis 34 and activated in mammals by dietary blueberry 35, 36. Figure 5 A shows that in our experimental conditions the starvation did not affect SIRT1 expression, which, on the contrary, was remarkably increased by BJ or BE treatment as compared to control and starved cells. Subsequently, SIRT1 activity was investigated by measuring levels of acetylated‐p53 (Ac‐p53), a substrate of SIRT1 also involved in the apoptotic event 34. As reported in Fig. 5 B, BJ or BE was able to decrease Ac‐p53 levels as compared to starved cells. This effect was removed completely by EX527, a specific inhibitor of SIRT1, at the concentration able to inhibit SIRT1 activity 37. The involvement of SIRT1 in an antiosteoclastogenic and antiapoptotic effect elicited by BJ and BE was determined in MLO‐Y4 cells transfected with specific SIRT1 siRNA or in the presence of EX527. Figure 5 C shows SIRT1 decrease in control cells after 24 h of transfection. SIRT1 down‐regulation induced a significant increase in RANKL, sclerostin and apoptosis levels in 24‐h‐starved cells treated with BJ or BE (Fig. 5 D, E). Similarly, SIRT1 inhibition due to EX527 increased apoptosis in starved MLO‐Y4 cells (Fig. 5 E). Figure 5 Effect of SIRT 1 on RANKL and sclerostin expression and apoptosis in MLO ‐Y4 cells treated with BJ and BE. SIRT 1 expression (A, C) and Ac‐p53 levels (B) were determined in MLO ‐4Y cells in the absence or in the presence of 10 μ m EX 527. RANKL and sclerostin expression and apoptosis were measured in cells transfected with SIRT 1 si RNA or Scr si RNA (D, E) or in the presence of 10 μ m EX 527 (E). Cells were cultured for 24 h in complete medium (C, control) or in serum‐free medium (S, starved cells). Starved cells were treated or not with 30 μg· mL −1 of total soluble polyphenols ( TSP ) of BJ and BE, as reported in Materials and methods. Ac‐p53, RANKL and sclerostin levels were measured by western blot analysis and values are normalized with β‐actin bands obtained by densitometric analysis, and blots are representative of three experiments. Apoptosis data were relative to mono‐ and oligonucleosomes released into the cytoplasmic fraction from 10 4 cells. The data are expressed as fold‐increase over control or starved cell values and are the mean ± SEM of three experiments performed in duplicate. Data were evaluated by using one‐way ANOVA followed by Bonferroni's post hoc test. * P ≤ 0. 001 compared to BJ ‐ and BE ‐treated cells without EX 527 or with Scr si RNA ; ○ P ≤ 0. 01 compared to the respective untreated starved cells with EX 527 or with SIRT 1 si RNA ; ∆ P ≤ 0. 001 compared to untreated starved cells. Effect of BJ on ROS production and viability in serum‐deprived MSCs Since bone formation requires the recruitment, proliferation and osteogenic differentiation of mesenchymal progenitors, we extended our investigation to MSCs. In particular, we investigated the antioxidant ability of BJ and its protective effect against oxidative stress‐reduced viability of MSCs. The antioxidant effect of BJ, containing various concentrations of TSP (7. 5–30 μg·mL −1 ), was tested in MSCs cultured for 24 h in the presence of reduced concentrations of serum (0. 1%, 0. 5%) in order to promote oxidative stress. Indeed, as reported in Fig. 6, serum deprivation determined an increase of ROS production as compared to cells incubated in medium with 10% fetal bovine serum (FBS). A significant ROS increase was also observed in cells cultured with 0. 1% FBS as compared to values from cells cultured in 0. 5% FBS (Fig. 6 A). In the presence of 0. 5% FBS, BJ treatment (7. 5–30 μg·mL −1 TSP) reduced significantly ROS levels at all concentrations used, and 30 μg·mL −1 TSP reduced ROS levels to about 70% as compared to untreated cells (Fig. 5 A). A similar effect was also obtained in cells cultured in 0. 1% FBS (Fig. 6 A). The decrease of serum content also induced changes in cell morphology and density (Fig. 6 B), and the decrease of cell viability was significantly related to serum deprivation (Fig. 6 C). Figure 6 Effect of serum deprivation and BJ on intracellular ROS production, morphology and cell viability in MSC s. MSC s were cultured for 24 h in growth medium with 10%, 0. 5%, 0. 1% or no FBS, and treated or not with various concentrations of total soluble polyphenols ( TSP ) of BJ, as reported in Materials and methods. (A) Intracellular ROS production was detected as described in Figure 1 and the data are expressed as fold‐increase over values obtained in cells cultured in 10% FBS. (B) Morphological images of MSC s representative of at least three independent experiments; scale bar: 25 μm. (C) Cell viability was detected by 3‐(4, 5‐dimethylthiazol‐2‐yl)‐2, 5‐diphenyl‐tetrazolium bromide ( MTT ) and the absorbance values are reported as arbitrary units (A. U. ) from 10 4 cells. The data are the mean ± SEM of three experiments performed in triplicate. Data were evaluated by using one‐way ANOVA followed by Bonferroni's post hoc test. * P ≤ 0. 05 compared to cells grown in 10% FBS ; ○ P ≤ 0. 05 compared to cells grown in 0. 5% or 0. 1% FBS without BJ TSP. Since the incubation of MSCs in a medium containing 0. 1% FBS promoted an extensive cell death, leading to a 4‐fold increase of the sub‐G 0 /G 1 fraction as compared to cells with 10% FBS (4. 5 ± 0. 1% versus 1. 02 ± 0. 08%, n = 4, P < 0. 05), the ability of BJ to counteract cell damage was determined in MSCs cultured in 0. 5% FBS in which the reduction of the sub‐G 0 /G 1 fraction was only 2‐fold (2. 5 ± 0. 2% versus 1. 02 ± 0. 08%, n = 4, P < 0. 05). As shown in Fig. 6 C, BJ (15–30 μg·mL −1 TSP) was able to preserve cell viability by approximately 75% as compared with untreated cells, whereas surprisingly, a higher concentration of TSP (60 μg·mL −1 ) did not have a similar protective effect. Effect of BJ on cytotoxicity and involvement of SIRT1 in serum‐deprived MSCs Next, we evaluated the role of SIRT1 in BJ's ability to prevent MSC cytotoxicity. MSCs were incubated with 30 μg·mL −1 TSP of BJ and in serum‐deprived medium (0. 5% FBS) in the absence or in the presence of EX527 for 24 h. As shown in Fig. 7 A, a significant reduction of cytotoxicity (of about 40%) was observed in the presence of BJ TSP compared to untreated cells. Notably, similarly to what observed in osteocytes, the pharmacological inhibition of SIRT1 prevented the protective effect elicited by BJ (Fig. 7 A), suggesting the involvement of this deacetylase activity also in these progenitor cells. On the other hand, while the expression of SIRT1 was induced by BJ in MLO‐Y4 cells (Fig. 5 A), the protein expression was not changed by BBs in these cells (Fig. 7 B). Figure 7 Effect of SIRT 1 on toxicity in MSC s treated with BJ. (A) Cell cytotoxicity was determined in MSC s cultured for 24 h in 0. 5% FBS in the absence or in the presence of 30 μg· mL −1 of total soluble polyphenols ( TSP ) of BJ and with or without 10 μ m EX 527, as reported in Materials and methods. The absorbance values are expressed as arbitrary units (A. U. ) from 10 4 cells and are the mean ± SEM of four experiments performed in quadruplicate. (B) SIRT 1 expression was determined by western blot analysis and values are normalized with β‐actin bands obtained by densitometric analysis. Blots are representative of three experiments. Data were evaluated by using one‐way ANOVA followed by Bonferroni's post hoc test. * P ≤ 0. 05 compared to cells grown in 0. 5% FBS ; ° P ≤ 0. 05 compared to cells grown in 0. 5% FBS with TSP and without EX 527. Discussion The consumption of BBs is considered an important contribution to the diet; this is due to the abundance of various classes of polyphenolic compounds that make this fruit rich in anti‐inflammatory, anti‐hypertensive, antimicrobial, and anticancer properties 29. Studies on bioavailability of BB polyphenols show that some of these are absorbed and present in human plasma for a variable period after intake. In fact, plasma concentrations of many of these compounds increase significantly following the administration of diets containing blueberries, although many of the phytochemicals are modified after absorption 38, 39, 40. Plasma kinetic profiles of BB polyphenols for 2 h after a BJ or BB smoothie intake have been reported 41. Some orally administrated polyphenols, such as anthocyanins or cyanidins, can be absorbed as glycosides and/or aglycones in intact form and as such they are found in the blood 42, 43, 44. Nevertheless, many studies evidence that intake amount, chemical structure, enterohepatic circulation and other individual factors, such as age, gender, gut microbiota, and genetic polymorphisms, play a role in BB polyphenol bioavailability evidencing a complex metabolic fate of these compounds not necessary involving the loss of their function 45, 46. Among the bioactive substances present in BBs, flavonoids, specifically anthocyanins, have been proven to have a strong antioxidant capacity 28, 29, 30, 40. In this study, the effect of BBs was evaluated in starved MLO‐Y4 cells, a condition that mimics in vitro a metabolic situation of oxidative stress that may be similar to what occurs in vivo in the bone environment after microdamage and oestrogen deficiency 12, 14, 18, 26, 47. Previously, it has been demonstrated that oxidative stress‐induced apoptosis by starvation in MLO‐Y4 cells is involved in the up‐regulation of osteoclastogenic factors 18. In fact, thiol antioxidants inhibit ROS production due to starvation and prevent both apoptosis and the increase of osteoclastogenic factors. A relationship among these events has been found 18. The present study demonstrates for the first time in MLO‐Y4 cells the antioxidant, antiapoptotic and anti‐osteoclastogenic properties of both BJ and BE containing equal amounts of TSP. In particular, BJ promotes a concentration‐dependent antioxidant effect after 24 h of starvation. Moreover, a similar antioxidant effect between BJ and BE is found in both a cellular and a cell‐free model. This indicates that BB effects are independent of the type of preparation (juice or extract) containing the same TSP amount. Like in starved MLO‐4Y cells, BJ shows the same antioxidant activity also in MSCs cultured at low serum concentrations, a condition that induces intracellular oxidative stress in these cells. We speculate that the beneficial antioxidant effect of BBs is mainly due to the polyphenolic fraction; in fact, dietary antioxidant polyphenols activate osteoblast function promoting bone growth and inhibiting osteoclast differentiation 1, 2, 3, 4, 5, 31, 32. Literature data show that ROS might be involved in the pathogenesis of bone loss and that nutritional approaches to antioxidant strategies may prevent this 2, 6, 12, 14. Indeed, dietary BBs protect against ovariectomized‐induced osteoblast death 31 and regulate osteoblast differentiation 2. This agrees with the preventive effect of BJ and BE on oxidative stress‐induced up‐regulation of RANKL and sclerostin and osteocyte apoptosis, and this preventive effect occurs by reducing significantly ROS‐induced JNK and ERK1/2 activation present in starved MLO‐4Y cells, as previously reported 18, 41. It has been reported that the up‐regulation of RANKL and sclerostin occurs in osteocytes under various conditions, including bone pathological alterations 12, 16, 19. In fact, RANKL increases osteoclast differentiation and bone resorption, whereas sclerostin, specifically produced in mature osteocytes, is a negative regulator of Wnt/β‐catenin signalling that inhibits osteoblast activity and osteogenesis 16, 48. Our data regarding the down‐regulation of RANKL expression by BJ and BE agree with the inhibitory effect of polyphenol extracts on increased RANKL expression in response to tumour necrosis factor‐α‐induced oxidative stress 4. However, no data are reported on the regulation of sclerostin expression in osteocytes by polyphenols or plant extracts. It is interesting that the effects observed on the osteoclastogenic factors in osteocytes are obtained using concentrations of TSP similar to the polyphenol concentrations of dried plums, which suppress RANKL expression and enhance osteoblast activity 49. The ability of BJ and BE to protect osteocytes from apoptosis is very important considering that abnormal osteocyte apoptosis is closely related to the expression of osteoclastogenic factors and high bone turnover, both events involved in bone loss and osteoporosis 14, 16, 17. Indeed, in young rats fed a diet containing different amounts of BB (from 1% to 10%) there is a significant increase in bone mass and inhibition of osteoclast differentiation associated to RANKL and related to amounts of administered BBs 50. Moreover, polyphenol‐derived phenolic acid present in serum from BB diet‐fed rats is bioactive, stimulating osteoblast differentiation and bone mineralization through Wnt signalling 2, 23. Previous data demonstrated that in MLO‐Y4 cells, starvation‐induced apoptosis is closely related to increased mitochondrial ROS production, which, through JNK activity, induces caspase‐3 activation 18, 47. Furthermore, this study demonstrates that the antiapoptotic ability (apoptosis reduction of about 70–80%) of BJ and BE is higher than their antioxidant capacity (ROS reduction of about 50%). This may agree with the involvement of SIRT1 in the antiapoptotic effect of BJ and BE, which is able to up‐regulate the expression of SIRT1 and increase its activity. These events are also related to BB prevention of RANKL and sclerostin up‐regulation due to oxidative stress. Thus, the protective effect of BJ and BE is due both to their direct antioxidant action and to the up‐regulation of SIRT1, which does not appear to be a redox‐regulated mechanism. In fact, the expression of SIRT1 does not change in untreated starved cells in the presence of oxidative stress as compared with the control. Indeed, recently it has been demonstrated that SIRT1 overexpression may prevent H 2 O 2 ‐induced apoptosis in osteoblast cells by decreasing the activity of caspase‐3 via down‐regulation of the forkhead box O/β‐catenin signalling pathway 51. Many data indicate that the effects of phytocompounds may be due not only to their ROS scavenger activity but also to specific interactions with proteins involved in intracellular signalling pathways related to the regulation of osteoblast and osteoclast activity 30. Indeed, dietary blueberry increased SIRT1 levels in mammals 36, and SIRT1 overexpression has also been related to down‐regulation of sclerostin in ovariectomized female mice 52 and to the inhibition of RANKL‐induced osteoclast differentiation without affecting osteoclast survival 53. Therefore, SIRT1 has been proposed as a potential target for anabolic therapies aiming to block bone resorption and to restore a normal remodelling process. Altogether, these data suggest that the antiapoptotic and antiosteoclastogenic role of BBs in starved MLO‐Y4 cells is due to their antioxidant properties and to SIRT1 activity. Finally, it has been reported that during tissue regeneration, stem cell proliferation is sensitive to changes of environmental conditions and, in particular, is reduced in the presence of oxidative stress 21, 23. Thus, our findings on BJ's ability to protect MSCs from damage induced by ROS increase are interesting and worthy of further investigation. Moreover, since metabolites of phenolic acids, derived from vegetable polyphenols, are able to stimulate MSCs versus osteoblast differentiation 23, BJ's effect in MSCs could be related to its polyphenolic components. However, it is noteworthy that in MSCs the highest concentration of TSP does not have a beneficial effect on cell viability, different from what was observed in MLO‐Y4 cells, indicating a possible harmful action in MSCs due to the high concentration of TSP. This may be due to the different sensitivity to the action of BJ of the two types of cells; in fact, MLO‐Y4 cells are highly differentiated and mature osteoblasts, while on the other hand, MSCs are undifferentiated cells and, perhaps, more sensitive either to toxic effects due to oxidative stress or to possible harmful effects of high concentrations of BB phytochemicals. These findings indicate that in the MSCs the beneficial effect of BJ occurs within a certain concentration range of TSP. Moreover, BJ's action in these cells seems to depend in part on SIRT1 activity and this, together with the ability of SIRT1 to promote MSCs' osteogenic differentiation 54, may be very important considering that MSCs are currently considered among the best candidates for cell‐based therapy in regenerative medicine 20, 24. Given that BJ in MSCs does not activate SIRT1 expression, we suggest that its involvement in BJ's protective effect can be due to a post‐translational mechanism(s) in the regulation of SIRT1 activity induced by BJ. It is necessary to further investigate the mechanisms by which SIRT1 in osteocytes as well as in MSCs can contribute to the protective effect of BJ against oxidative stress damage. Conclusions The results of this study demonstrate, for the first time, in osteocytes, cells in close contact with blood capillaries and considered the major regulators of bone remodelling, a significant relationship between the antioxidant activity of BBs and molecular events related to apoptosis and expression of osteoclastogenic factors induced by oxidative stress. Other novel data show that BBs protect MSCs, important cells for bone regeneration, against reduction of viability and cytotoxicity due to oxidative stress. Moreover, the protective effects of BBs both in osteocytes and in MSCs are in part mediated by SIRT1. Indeed, this enzyme is considered a possible target for anti‐resorptive drugs 55 and for anabolic therapies for osteoporosis 52, 53, 56 ; reduced SIRT1 expression has been associated with osteoporotic hip fracture 57. The present study also reports that two different preparations of BB (juice or extract) containing the same TSP amount show similar effects due to the complex mixture of polyphenols and/or to other bioactive phytochemicals for some of which the bioavailability has been demonstrated 38, 39, 40, 42, 43, 44. This may be interesting, considering that most studies report the effects on bone metabolism of isolated natural compounds, although individuals do not consume isolated molecules but fruits and vegetables rich in many different types of polyphenols and phytochemicals. Overall, these novel data in osteocytes and MSCs may contribute to explain at the cellular and molecular level the protective effects of BB phytochemicals against the damage caused by oxidative stress in bone remodelling and regeneration. Indeed, beneficial anabolic effects of BBs in bone tissue have been reported in animal studies, which suggest blueberries to be a useful supplement for the prevention and/or management of osteoporosis and the osteogenic process 2, 3, 31, 32, 50. Materials and methods Preparation of blueberry juice and solubilized extracts BBs, harvested in August 2017/2018 in Tuscany Apennines, were frozen freshly picked in aliquots of 100 g each and used when necessary to prepare BJ by homogenization in a refrigerated Waring blender. The fruit mixture was then filtered under vacuum and centrifuged at 20 000 g for 10 min to remove insoluble particles. Aliquots of BJ were stored at −20 °C until use. BE was prepared by solubilizing the dry extract, obtained by Aboca SpA (Sansepolcro, Italy), in phosphate‐buffered saline (PBS). Determination of total soluble polyphenols The TSP fraction in BJ and BE was quantified with the Folin–Ciocalteu reagent according to a slightly modified method of Singleton and Rossi using gallic acid as the standard 58. Briefly, 100 μL commercial Folin–Ciocalteu reagent (Merck KGaA, Darmstadt, Germany) diluted 1 : 10 in distilled water was added to 20 μL of sample or standard placed in 96‐well plates. After 5 min, the reaction was stopped by adding 80 μL of saturated Na 2 CO 3 solution. Samples and standard were kept in the dark at 25 °C for 2 h. Subsequently, the absorbance was measured in a microplate reader at 765 nm. The TSP concentration in BJ, obtained from 100 gr of BB fresh weight, was 1. 6 ± 0. 1 mg·mL −1 (about 50 mg/100 g of fresh weight). TSP concentration in the solution obtained from BB dry extract was 2. 3 ± 0. 2 mg·mL −1 (about 460 mg/100 g of dry extract). In literature, the range of TSP is from 48 to 304 mg/100 g of fresh BB weight 40, and this range strictly depends on the cultivar, growing conditions and maturity, and the estimation may vary depending on the analytical procedure. Determination of antioxidant capacity by measuring superoxide anion radical scavenging activity The antioxidant capacity of BJ and BE, both containing 50 μg·mL −1 of TSP, was assayed evaluating the Nitroblue tetrazolium (NBT) reduction mediated by superoxide anion produced by the xanthine/xanthine oxidase system, as previously described 59. Briefly, the superoxide anion, generated by the reaction catalysed by xanthine oxidase in the presence of xanthine, induced the reduction of NBT that represents a target for detection of O 2 − · scavenging capacity. The coloured reaction, due to the reduction of NBT with the O 2 − ·, was detected at 560 nm [Perkin Elmer (Waltham, MA, USA) spectrophotometer]. Antioxidant activity of BJ and BE samples inhibits the colour change. For this assay, reagents used were purchased from Sigma‐Aldrich (St Louis, MO, USA). Cell culture and treatment MLO‐Y4 osteocyte‐like cells (a gift from Dr L. Bonewald, University of Missouri‐Kansas City) were cultured at 37 °C in a 5% CO 2 humidified atmosphere in α‐MEM supplemented with 5% calf serum (HyClone, GE Healthcare, Chicago, IL, USA), 5% FBS (HyClone, GE Healthcare), 2 m m l ‐glutamine, 72 mg·L −1 penicillin and 100 mg·mL −1 streptomycin (complete medium). MLO‐Y4 cells were grown in complete medium to 70–80% confluence, and then incubated for 1 h in the presence or not of BJ containing TSP at different final concentrations (from 7. 5 μg·mL −1 to 60 μg·mL −1 ), or in the presence of BE containing TSP at the final concentration of 30 μg·mL −1. Subsequently, complete medium was removed, and for another 4 or 24 h the cells were cultured in serum‐free medium (starved cells) in the presence or not of BJ or BE and in fresh complete medium (control). Experiments with EX527 (Sigma‐Aldrich), a specific inhibitor of SIRT1 37, were performed in cells cultured in complete medium in which the inhibitor was added for 30 min at the final concentration of 10 μ m. After removing the complete medium, the cells were cultured for another 24 h in serum‐free medium (starved cells) and in fresh complete medium (control) in the presence or not of EX527. Starved cells with or without the inhibitor were treated or not with BJ or BE as reported above. The MSCs were purchased from ATCC (Manassas, VA, USA), expanded and cultured as previously reported 51. The cells were plated at low density (3–5000 cells·cm −2 ) and incubated for 24 h in 10%, 0. 5% or 0. 1% or no FBS in the presence or not of BJ at concentrations reported for osteocytes. Experiments with EX527 were performed as previously reported for osteocytes. Some treatments were performed in cells transiently transfected with 75 n m mouse SIRT1 siRNA corresponding to two DNA target sequences of mouse SIRT1 (5′‐GUUACUGCAGGAGUGUAAA[dT][dT]‐3′; 5′‐UUUACACUCCUGCAGUAAC[dT][dT]‐3′) (Sigma‐Aldrich) or scrambled siRNA (Universal Negative Control #1, Sigma‐Aldrich), using Lipofectamine RNAiMAX™ (Invitrogen, Carlsbad, CA, USA) according to the manufacturer's instructions. The ability of siRNA to silence SIRT1 expression levels was checked in control cells 24 h after transfection. In experiments with EX527, 0. 008% final concentration of DMSO was present in control and in all treated and untreated cells. Determination of intracellular ROS The intracellular levels of ROS were measured in MLO‐Y4 cells and in MSCs seeded in 12‐well plates and treated with BJ and BE as reported above. One hour before the end of the various treatments, the cell‐permeant probe 2′, 7′‐dichlorodihydrofluorescein diacetate (H 2 DCFDA; Invitrogen) was added in culture medium. Once deacetylated by esterases, the probe is rapidly oxidized to a highly fluorescent compound in the presence of ROS. After PBS washing, adherent cells were lysed in RIPA buffer (50 m m Tris/HCL pH 7. 5, 1% Triton X‐100, 150 m m NaCl, 100 m m NaF, 2 m m EGTA, phosphatase and protease inhibitor cocktail), centrifuged at 20 000 g (ALC PK121R, Thermo Fisher Scientific, Waltham, MA, USA) for 10 min, and analysed immediately by fluorescence spectrophotometric analysis at 510 nm. Data were normalized to total protein and expressed as fold‐increase over the control values. Cell apoptosis assay MLO‐Y4 cells, seeded in six‐well plates and treated with BJ and BE, as reported above, were used to assess apoptosis by using Cell Death Detection ELISAplus Kit (Roche Laboratories, Nutley, NJ, USA), and annexin V–FITC Kit (Miltenyl Biotec GmbH, Bergisch Gladbach, Germany), according to the manufacturer's instructions. The specific increase of mono‐ and oligonucleosomes released by 10 4 cells in the cytoplasmic fractions was detected, and data are expressed as fold‐increase over the control values using the following ratio: absorbance of the sample/absorbance of the control values. Cells treated with annexin V–FITC were analysed using a flow cytometer (FACSCalibur; BD Biosciences, San Jose, CA, USA). cellquest ™ software (version 3. 3; BD Biosciences) was used for analysis relative to the phosphatidylserine present outside the plasma membranes and data are expressed as fold‐increase over the control values. Cell viability and cytotoxicity assays The MSCs were incubated for 24 h in DMEM containing 10% FBS in the absence or in the presence of compounds as reported above. Oxidative stress was induced by reduction of serum in the culture medium. Cell viability was evaluated by non‐radioactive cell assay (MTT) (CellTiter 96 ® Assay; Promega Corp. , Madison, WI, USA) according to the manufacturer's instructions and as previously reported 60, 61. Cell cytotoxicity was determined by using CytoTox 96 ® Non‐Radioactive Cytotoxicity Assay (Promega), a colorimetric assay that measures lactate dehydrogenase, a stable cytosolic enzyme that is released upon cell lysis. Apoptotic cell fraction was determined after fixation and propidium iodide staining by TALI ® cytometry (Life Technologies, Carlsbad, CA, USA) 61. Western blot analysis The phosphorylation of ERK1/2 and JNK, and the expression of activated caspase‐3, PARP‐1, SIRT1, RANKL, sclerostin and Ac‐p53 were performed by western blot in MLO‐Y4 cells seeded in 60 mm tissue culture dish treated as reported above. Cells were lysed for 30 min at 4 °C in ice‐cold RIPA buffer and centrifuged at 20 000 g for 10 min. Equal amounts of total proteins (40–60 μg) were loaded in each line and were subjected to SDS/PAGE on 10% gel and electrotransferred to poly(vinylidene difluoride) membrane (GE Healthcare). Membranes were probed with specific primary antibody anti‐caspase‐3 or anti‐phospho‐ERK1/2 or anti‐phospho‐JNK or anti‐PARP‐1 (Cell Signalling Technology, Danvers, MA, USA) or anti‐SIRT1 or anti‐RANKL or anti‐sclerostin (Santa Cruz Biotechnology, Inc. , Dallas, TX, USA), or anti‐Ac‐p53 (Thermo Fisher Scientific). Subsequently, after stripping the membranes were reprobed with anti‐ERK1/2 or anti‐JNK or anti‐β‐actin for normalization of densitometric values. Secondary antibodies conjugated to horseradish peroxidase were used to detect antigen–antibody complexes using a chemiluminescence reagent kit (Clarity Western ECL Substrate, Bio‐Rad, Hercules, CA, USA). imagej software (National Institutes of Health, Bethesda, MD, USA) was used to perform quantitative analyses, and the values of the bands were expressed as fold‐increase relative to control values. Protein assay Protein concentrations were determined by the bicinchoninic acid solution protein reagent assay (Thermo Fisher Scientific) using bovine serum albumin as the standard (Sigma‐Aldrich). Statistical analysis Each experiment was performed a minimum of three times. Data are expressed as means ± SEM and statistical significance was determined by one‐way ANOVA with Bonferroni's multiple comparison test, using prism software (GraphPad Software Inc. , La Jolla, CA, USA). P ≤ 0. 05 was considered statistically significant. Conflict of interest The authors declare no conflict of interest. Author contributions CG, EM, TI, MLB, and MTV designed the experiments and analysed the data. VD, GM, FP, and LCM performed the experiments. GB performed the apoptosis experiments with the annexin V–FITC kit. VD, EM, TI, and MTV wrote the manuscript.
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10. 1002/2211-5463. 12792
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FEBS Open Bio
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All‐trans retinoic acid and human salivary histatin‐1 promote the spreading and osteogenic activities of pre‐osteoblasts
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Cell‐based bone tissue engineering techniques utilize both osteogenic cells and biomedical materials, and have emerged as a promising approach for large‐volume bone repair. The success of such techniques is highly dependent on cell adhesion, spreading, and osteogenic activities. In this study, we investigated the effect of co‐administration of all‐trans retinoic acid (ATRA) and human salivary peptide histatin‐1 (Hst1) on the spreading and osteogenic activities of pre‐osteoblasts on bio‐inert glass surfaces. Pre‐osteoblasts (MC3T3‐E1 cell line) were seeded onto bio‐inert glass slides in the presence and absence of ATRA and Hst1. Cell spreading was scored by measuring surface areas of cellular filopodia and lamellipodia using a point‐counting method. The distribution of fluorogenic Hst1 within osteogenic cells was also analyzed. Furthermore, specific inhibitors of retinoic acid receptors α, β, and γ, such as ER‐50891, LE‐135, and MM‐11253, were added to identify the involvement of these receptors. Cell metabolic activity, DNA content, and alkaline phosphatase (ALP) activity were assessed to monitor their effects on osteogenic activities. Short‐term (2 h) co‐administration of 10 μ m ATRA and Hst1 to pre‐osteoblasts resulted in significantly higher spreading of pre‐osteoblasts compared to ATRA or Hst1 alone. ER‐50891 and LE‐135 both nullified these effects of ATRA. Co‐administration of ATRA and Hst1 was associated with significantly higher metabolic activity, DNA content, and ALP activity than either ATRA or Hst1 alone. In conclusion, co‐administration of Hst1 with ATRA additively stimulated the spreading and osteogenicity of pre‐osteoblasts on bio‐inert glass surfaces in vitro.
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Abbreviations ALP alkaline phosphatase ATRA all‐trans retinoic acid FAK focal adhesion kinase Fmoc fluorenylmethoxycarbonyl Hst1 histatin‐1 LVBD large‐volume bone defects pNPP p ‐nitrophenyl phosphate RXR retinoid X receptor Large‐volume bone defects (LVBD) may severely influence aesthetics and musculoskeletal functions. Due to the limited healing capacity of bone tissues, the osseous repair of LVBD can be problematic 1. For treatment purposes, autologous bone grafts are still considered as the gold standard. However, their application is confined by limited graft supply, donor site pain and morbidity, infections, and poor cosmetic outcomes 2. As alternative options to autologous bone grafts, allografts, xenografts, and synthetic materials have been developed and adopted as bone‐defect‐filling materials 3. However, most of these materials need to be premixed with autologous bone grafts to obtain osteogenic cells. In such cases, the disadvantages of autologous bone grafts remain. To approach these challenges, cell‐based tissue engineering techniques that integrate osteogenic cells and biomedical materials have emerged as a promising approach for bone repair 4. However, the chance of success is, however, highly dependent on the interactions of the osteogenic cells with the cell‐scaffold surfaces 5. At first, cell–substrate interactions are critical for the determination of cell fates, such as proliferation, quiescence, or apoptosis 6. Furthermore, surface adhesion and osteogenic cell proliferation are indispensable for initiation of bone regeneration 7. Consequently, tremendous efforts have been made to develop a large variety of techniques (e. g. , immobilized RGD peptide on titanium surface 8 and femtosecond laser‐induced micropattern and Ca/P deposition 9 ) to modify surface chemistry and/or topography of various biomedical materials in order to improve their cell–substrate interactions 5, 10. These material‐specific approaches, however, render their broad applicability limited. In comparison, the other strategies – cell‐targeting techniques that directly promote cellular response to materials – have become highly attractive as they do not require surface modifications of materials, thus bearing a broader applicability. A promising candidate cell‐targeting agent to promote cell–substrate interactions is histatin‐1 (Hst1), a member of a large histidine‐rich salivary peptide family. Our previous findings show that Hst1 can significantly promote the attachment, spreading, and migration of various cell types including epithelial, endothelial, and osteogenic cells 11, 12, 13, 14, 15. Our recent data confirm that Hst1 can promote the spreading of osteogenic cells on both bio‐inert glass and titanium surface 14, 15, 16, 17, which suggests a promising application potential of Hst1 in the cell‐based bone tissue engineering. In a previous study, we found that a 3‐day treatment of all‐trans retinoic acid (ATRA), an active metabolite of vitamin A, can cause the uniform alignment and stretch of cell skeleton (Fig. S1 ). This finding inspired us to apply ATRA to promote cell spreading. ATRA, the active metabolite of vitamin A, is known to act as regulator of many physiologic processes 18. It plays a role in a wide range of biological processes mediated through binding and activation of the nuclear receptors, such as the RA receptor (RAR) and retinoid X receptor (RXR). There are three subtypes of RAR (α, β, and γ) and three subtypes of RXR (α, β, and γ). RARs are bound and activated by ATRA, while RXRs are bound and activated by the 9‐cis‐RA only 19. Heterodimers of activated RAR and RXR act as ligand‐dependent transcription factors. On the other hand, it was found that a 3‐day treatment of ATRA also results in significantly reduced osteogenic differentiation of pre‐osteoblast cells and bone marrow stromal cells 20, 21. Consequently, in the present study, we analyzed in vitro the effect of a short (2 h) co‐application of ATRA and Hst1 in order to amplify the stimulating effect of Hst1 on the spreading of osteogenic cells on the one hand and to avoid the decrease in osteogenic potential on the other hand. Materials and methods Study design The effect of a short (2 h) co‐administration of ATRA and Hst1 on cell spreading was evaluated. Thereafter, we used specific inhibitors of retinoic acid receptor alpha (RARα), RARβ, and RARγ, that is, ER‐50891, LE‐135, and MM‐11253, respectively, to identify the involvement of RARs. Furthermore, we examined the effects of a short co‐administration of ATRA and Hst1 on the osteogenic potentials of pre‐osteoblast cells, such as metabolic activity, DNA content (indicator for proliferation), and alkaline phosphatase (ALP) activity (early marker of osteogenic differentiation). Preparation of histatin‐1 Histatin‐1 was manufactured by solid‐phase peptide synthesis using 9‐fluorenylmethoxycarbonyl (Fmoc) chemistry as described previously 15, 22. Hst1 was purified to at least 95% by high‐performance liquid chromatography (RF‐HPLC, Dionex Ultimate 3000; Thermo Scientific, Breda, the Netherlands). The authenticity was confirmed by mass spectrometry with a Microflex LRF MALDI‐TOF (Bruker Daltonik GmbH, Bremen, Germany) as previously described 15, 22. Fluorescently labeled Hst1 was prepared using the fluorogenic dye ATTO‐647N (ATTO‐TEC GmbH, Siegen, Germany). The ε‐amino group of the side chain of lysine residue number 17 (lys17, K of Hst1 after removal of the specific protective lysine derivative, Fmoc‐Lys(ivDde)‐OH, by hydrazine (2% hydrazine hydrate)) was coupled to equimolar amount of the dye. Cell culture and chemicals MC3T3‐E1, a mouse pre‐osteoblast cell line, subclone 4 (CRL‐2593, American Type Culture Collection, ATCC, Manassas, VA, USA), was cultured in alpha‐minimum essential medium (α‐MEM; Gibco, Thermo Fisher Scientific, Paisley, UK) supplemented with 10% FBS (Gibco, Thermo Fisher Scientific) and 1% penicillin/streptomycin (Sigma, St. Louis, MO, USA). Cells were cultured in humidified oxygen‐controlled 37 °C incubator with 5% CO 2. Passages between 4 and 7 were used for experiments. Measurement of cell spreading on glass surface Cells were treated with serum‐free medium for 24 h before being detached by 0. 05% trypsin (Gibco, Thermo Fisher Scientific). Growth medium contained 2% FBS was used to inactivate the effect of trypsin and to resuspend the cells. MC3T3‐E1 was seeded on coverslips (20 mm in diameter; Thermo Scientific, Braunschweig, Germany) in 12‐well plates at a density of 6 × 10 4 cells/well. Cells were treated either with 0, 1, 10, or 20 µ m ATRA (Sigma‐Aldrich) or Hst1 or co‐administered 10 µ m ATRA and Hst1. To investigate the role of potential signaling pathways, 10 µ m RARα antagonist (ER‐50891; R&D, Bio‐Techne, Minneapolis, MN, USA), 10 µ m RARβ antagonist (LE‐135; R&D, Bio‐Techne), and 10 µ m RARγ antagonists (MM‐11253; R&D, Bio‐Techne) were supplemented in cell spreading assays. Cells were photographed every 20 min for 3 h using a microscope (EVOS FL; Thermo Fisher Scientific) equipped with a LPlanFL PH2 20× using the phase‐contrast setting or the Cy5 light cube (628/40 and 692/40 nm, excitation and emission filters, respectively). Relative cell spreading surface area was quantified by measuring the surface area of cells' filopodia and lamellipodia using a manual point‐counting method 23 (Fig. S2 ). Each assay was performed in triplicate and repeated twice. Fluorescent staining of spreading cells Cell spreading on glass surface was performed as described in the section of cell spreading assay. 1. 5 h after seeding, cells were fixed, dehydrated, and stained with FITC‐Phalloidin. Fluorescent micrographs were randomly taken using a fluorescent microscope (Leica Microsystems GmbH, Wetzlar, Germany) with excitation/emission wavelengths (nm) of 496/516. On the micrographs, spreading surface of each cell was estimated using the above‐mentioned point‐counting method. More than 20 cells per group were calculated. Cell metabolic activity Subconfluent growing cells were plated on glass coverslips (diameter, 10 mm; Thermo Scientific, Germany) in 48‐well plate in a density of 1. 5 × 10 4 cells/well. Cells were treated with either 10 µ m Hst1 or ATRA, or cells were treated with premixed ATRA and Hst1 for 2 h at 37 °C. After washing with 1× PBS, cells were treated with α‐MEM with 10% FBS which was refreshed on a daily basis. PrestoBlue™ Cell Viability Assay was adopted to evaluate cell viability using the reducing ability of cells (Invitrogen Corporation, Carlsbad, CA, USA). In short, 1/10th volume of PrestoBlue™ reagent was added to cells in culture medium and incubated for 30 min at 37 °C. Results were measured by reading fluorescence intensity with the Multiskan FC (Thermo Scientific) using a fluorescence excitation wavelength of 560 nm and an emission wavelength of 590 nm. Each assay was performed in triplicate and repeated twice. DNA quantification The CyQUANT Proliferation Assay Kit (Molecular Probes, Waltham, MA, USA) was employed to monitor the proliferation of pre‐osteoblasts. Subconfluent growing cells were plated on glass coverslips (diameter, 10 mm; Thermo Scientific, Germany) in a 48‐well plate at a density of 1. 5 × 10 4 cells/well. Cells were treated with either 10 µ m Hst1 or ATRA, or co‐administered ATRA and Hst1 for 2 h. After washing with 1× PBS, cells were treated with α‐MEM with 10% FBS which was changed every day. The cells were retrieved right after or 5 days after the short treatment. Subsequently, the freshly prepared 100 μL CyQUANT solution was added to the well to measure the optical density with excitation at 480 nm and emission at 520 nm using a plate reader (Synergy, BioTek™, Winooski, VT, USA). Each assay was performed in triplicate and repeated twice. Alkaline phosphatase assays Quantitative determination of ALP activity was done using the p ‐nitrophenyl phosphate (pNPP) liquid substrate method. Cells were suspended in serum‐free media in the presence or absence of 10 µ m Hst1 or 10 µ m ATRA or both and then seeded on glass coverslips (10 mm in diameter; Thermo Scientific, Germany) in 48‐well plates at a density of 5 × 10 4 cells/well. Two hours after seeding, the media were changed to 10% FBS‐containing α‐MEM, and subsequently, cells were cultured for 1 more day. Thereafter, cells were treated with α‐MEM containing 2% FBS. After 3 days, cells were lysed in distilled water using a freeze–thaw method and harvested with a cell scraper. Cell lysates were centrifuged at 250 g for 5 min at room temperature, and supernatants were incubated with 1. 86 mg·mL −1 pNPP for 1 h at 37 °C in the dark. After 1 h, 100 μL 300 m m NaOH solution was added; then, absorbance at 405 nm was measured using Multiskan FC (Thermo Scientific, Rockford, IL, USA), and ALP activity was calculated according to the standard curve. The total protein was assessed using Pierce BCA Protein Assay Kit (Thermo Fisher Scientific, Rockford, IL, USA) for normalizing the ALP activity 20. Each assay was performed in quadruplicate and repeated twice. Statistical analysis Data were plotted using graphpad prism (GraphPad Software version 6. 0, La Jolla, CA, USA) and analyzed by one‐way ANOVA with Bonferroni's post hoc test for multiple comparisons. For the data from different groups at different time points in Fig. 2 C, we used two‐way ANOVA to analyze the data with Tukey test for multiple comparisons. Results were reported as mean ± standard deviation (SD). A value of P < 0. 05 was considered as statistical significance. Results and Discussion ATRA and Hst1 promote the spreading of osteogenic cells on bio‐inert glass surface subsection At a concentration of 10 µ m, ATRA significantly promoted the spreading of pre‐osteoblasts on bio‐inert glass surfaces in comparison with the control (no ATRA) (Fig. 1 A, B). Hst1 at 10 and 20 µ m significantly promoted the spreading of osteogenic cells in comparison with the control (no Hst1) (Fig. 1 C, D). Thereafter, we performed a pilot experiment to check the effect of co‐administered 10 µ m Hst1 and ATRA of different concentrations (e. g. , 0. 1, 1, and 10 µ m ). Only 10 µ m ATRA and 10 µ m Hst1 resulted in a significant cell spreading area than 10 µ m Hst1 effect (data not shown). Therefore, we adopted the combination of 10 µ m ATRA and 10 µ m Hst1. Our data showed that there was no significant difference between the promoting effects of 10 µ m Hst1 and 20 µ m Hst1. In this light, it was chosen to further use 10 µ m ATRA and 10 µ m Hst1 in the following experiments. Figure 1 (A) Light micrographs depicting the spreading of pre‐osteoblasts in the presence or absence of 10 µ m ATRA. Bar = 50μm. (B) Folds of cell spreading surface area in the presence or absence of 1, 10, and 20 µ m ATRA. Data are shown as mean ± SD ( n = 6). (C) Light micrographs of the cell spreading in the presence or absence of 10 µ m Hst1. Bar = 50μm. (D) Folds of cell spreading surface area in the presence or absence of 1, 10, and 20 µ m Hst1. Data were shown as mean ± SD ( n = 6). Data were plotted using graphpad prism (GraphPad Software version 6. 0) and analyzed by one‐way ANOVA with Bonferroni's post hoc test for multiple comparisons. ** P < 0. 01 comparing with control group; n. s. , no statistically significant difference. Effects of ATRA and Hst1 co‐administration on spreading of pre‐osteoblasts Sixty minutes post‐treatment, the co‐administration of ATRA and Hst1 resulted in significantly larger spreading of pre‐osteoblasts (2. 6‐fold) compared to the individual counterparts, viz. Hst1 (1. 9‐fold) and ATRA (1. 7‐fold) alone (Fig. 2 A, B). In the subsequent time‐course assay, the promoting effect of 10 µ m ATRA and 10 µ m Hst1 became significant from 40 to 100 min and more pronounced at 160 min (Fig. 2 C). Moreover, the surface area of cells treated with ATRA and Hst1 (861 ± 206 μm 2 ) was also significantly higher than those treated with ATRA (723 ± 182 μm 2, P < 0. 05) or Hst1 alone (665 ± 185 μm 2, P < 0. 001) (Fig. 3 A, B). Figure 2 (A) Light micrographs depicting the spreading of pre‐osteoblasts in the presence or absence of 10 µ m Hst1 or 10 µ m ATRA. Bar = 50μm. (B) Folds of spreading surface area of pre‐osteoblasts in the presence or absence of 10 µ m Hst1 or 10 µ m ATRA. Data are shown as mean ± SD ( n = 6). Data were plotted using graphpad prism (GraphPad Software version 6. 0) and analyzed by one‐way ANOVA with Bonferroni's post hoc test for multiple comparisons. * P < 0. 05; *** P < 0. 001. (C) Time‐dependent cell spreading surface area [expressed in folds with the value of the control group (no Hst1 and no ATRA) at first time point as 1] in the presence or absence of 10 µ m Hst1 or 10 µ m ATRA. Data were shown as mean ± SD ( n = 6). Data were analyzed by two‐way ANOVA with Tukey test for multiple comparisons. § P < 0. 05 indicating a significant difference compared with the values in the groups of Hst1 or ATRA; & P < 0. 05 indicating a significant difference compared with the value in the control group at the same time point; # P < 0. 05 indicating a significant difference compared with the value in the same treatment group at the earlier time point. Figure 3 (A) Fluorescent micrographs depicting the spreading of pre‐osteoblasts (stained with FITC‐Phalloidin) in the presence or absence of 10 µ m Hst1 and 10 µ m ATRA on bio‐inert glass surface. Bar = 50μm. (B) The surface area per pre‐osteoblast in the presence or absence of 10 µ m Hst1 and 10 µ m ATRA on bio‐inert glass surface. Data were plotted using graphpad prism (GraphPad Software version 6. 0) and analyzed by one‐way ANOVA with Bonferroni's post hoc test for multiple comparisons. Data were shown as mean ± SD ( n > 20). * P < 0. 05; *** P < 0. 001; **** P < 0. 0001. The antagonists of RARα and RARβ suppressed the promoting effect of ATRA and Hst1 on cell spreading The antagonists of RARα (ER‐50891) and RARβ (LE‐135) significantly suppressed the promoting effect of the co‐administered ATRA and Hst1 on the spreading of pre‐osteoblasts (Fig. 4 A). Consistent with above‐mentioned results, 10 µ m ATRA significantly elevated the promoting effects of 10 µ m Hst1 ( P < 0. 05), which could be nullified by the pretreatment of 10 µ m ER‐50891 (Fig. 4 B) or 10 µ m LE‐135 (Fig. 4 C). Figure 4 Folds of spreading surface area of pre‐osteoblasts that were treated with (A) co‐administered 10 µ m ATRA and Hst1 with or without the pretreatment with MM [10 µ m MM‐11253, the antagonist of retinoic acid receptor (RAR)γ], ER (10 µ m ER‐50891, the antagonist of RARα), and LE (10 µ m LE‐135, the antagonist of RARβ); (B) 10 µ m Hst1 in the presence or absence of 10 µ m MM or 10 µ m ATRA; (C) co‐administered 10 µ m Hst1 in the presence or absence of 10 µ m LE135 or 10 µ m ATRA. Data were plotted using graphpad prism (GraphPad Software version 6. 0) and analyzed by one‐way ANOVA with Bonferroni's post hoc test for multiple comparisons. Data were shown as mean ± SD ( n = 6). * P < 0. 05; *** P < 0. 001. Co‐administration of ATRA and Hst1 upregulated the osteogenic activities of pre‐osteoblasts Two‐hour treatment of 10 μ m Hst1 significantly enhanced the metabolic activity of pre‐osteoblasts already after 1 day, in contrast to ATRA (Fig. 5 A). Furthermore, the co‐administration of ATRA and Hst1 resulted in significantly higher metabolic activity in comparison with Hst1 alone ( P < 0. 05) (Fig. 5 A). Directly after seeding, the DNA content in cells treated with Hst1 and ATRA was significantly higher than those stimulated by Hst1 or ATRA alone. Five days after seeding, the DNA content in the group of ATRA alone and Hst1 alone was significantly higher compared to the control group. The co‐administration of ATRA and Hst1 resulted in a significantly higher DNA content than those in the groups of ATRA or Hst1 alone (Fig. 5 B). Three days postseeding, the ALP activity of the cells treated with Hst1 and ATRA was about 2. 6‐fold higher ( P < 0. 001) than those in the groups of Hst1 alone, ATRA alone, or control (Fig. 5 C). Figure 5 (A) Folds of the metabolic activities of pre‐osteoblasts within 5 days after a short (2 h) treatment with either no Hst1, no ATRA (control), or 10 µ m Hst1, or 10 µ m ATRA, or co‐administered 10 µ m ATRA and 10 µ m Hst1 with α‐MEM containing 10% FBS during seeding ( n = 6). (B) Folds of DNA content at 0 day and 5 days after response graph of DNA content after a short (2 h) treatment with either no Hst1, no ATRA (control), or 10 µ m Hst1, or 10 µ m ATRA, or co‐administered 10 µ m ATRA and 10 µ m Hst1 with α‐MEM containing 10% FBS during seeding ( n = 6). (C) Folds of ALP activity at 3 days after a short (2 h) treatment either without Hst1 or ATRA (control), or with 10 µ m Hst1, or 10 µ m ATRA, or co‐administered 10 µ m ATRA and 10 µ m Hst1 with α‐MEM containing 2% FBS during seeding ( n = 8). Data were plotted using graphpad prism (GraphPad Software version 6. 0) analyzed by one‐way ANOVA with Bonferroni's post hoc test for multiple comparisons. Data were shown as mean ± SD. * P < 0. 05; ** P < 0. 01; *** P < 0. 001. Surface adhesion, spreading, proliferation, and differentiation of osteogenic cells are critical steps for their respective success within cell‐based bone tissue engineering techniques 24. Previously, we found that Hst1 promoted the spreading of osteogenic cells on both bio‐inert substrates and titanium SLA surfaces in vitro. In this study, we found that the co‐administration of ATRA and Hst1 significantly increased cell spreading efficiency compared to the presence of Hst1 only. In line with previous work, cell spreading was used as a key parameter to evaluate efficacy of surface compatibility for osteogenic cells by different bioactive agents in vitro. In our previous studies, we have used the percentage of spreading cells or cell index as parameters to evaluate cell spreading 14, 15. The former parameter indicates the percentage of cells that initiate protrusion, and the latter parameter qualifies the impedance of cells that proportionally correlate to, but not directly show, cell spreading extent. In contrast, in the current study, we adopted a point‐counting method 25 to directly measure the surface area of cell spreading parts, which could directly reflect the newly formed cell–substrate contact area. We further subtracted the area of nuclei to purely evaluate area of spreading part, which helped us to more precisely evaluate the spreading extent. Most of the current methods to promote cell–substrate interaction are focused to modify the surface chemistry and/or topography 26. However, due to the large variety of biomaterials, there is still an apparent great need for broadly applicable approaches to promote cell spreading on biomaterials. Previously, we and others showed that Hst1 promoted adhesion, spreading, and migration of various epithelial cells from different origins, such as mucosa 12, 16, gingiva 14, 15, cornea 13, and skin 27, endothelial cells 14, 17, and osteogenic cells 11, 14. All together, these findings underline a non‐cell type‐specific character of Hst1 rendering a promising application potential for tissue engineering purposes. Next to Hst1, ATRA was used as agent to promote cell spreading for cell‐based tissue engineering techniques. Numerous studies have demonstrated that the RA signaling pathway, which is mediated via RAR and/or RXR, can modulate the expression of genes involved in cell growth, 28 energy metabolism, 24 and immune responses 29, 30. In a previous study, we found that a 3‐day treatment of ATRA caused uniformly‐directionally alignment of actin in vitro (Fig. S2 ). Notably, it was reported that ATRA increased the adhesion and spreading of pancreatic stellate cells via RARβ‐dependent signaling, thereby inhibiting cancer cell invasion 31. In this process, ATRA‐treated pancreatic stellate cells formed larger focal adhesion complexes, spread faster, attained a larger spreading area, attached stronger to the ECM (extracellular matrix), and displayed significantly larger and brighter focal adhesion complexes (both for talin and paxillin) in comparison with untreated control cells 31. Here, we showed that RARs were potentially involved for the promoting effect of ATRA on the spreading of pre‐osteoblasts. For this purpose, we adopted specific antagonists of RARα, RARβ, and RARγ and found that the antagonists of RARα (ER‐50891) and RARβ (LE135), but not RARγ (MM‐11253), significantly suppressed cell spreading induced by co‐administered ATRA and Hst1. Furthermore, we found that the antagonists of RARα (ER‐50891) and RARβ (LE135) abolished the amplification by ATRA of Hst1's effects on cell spreading. These findings suggested that ATRA affects cell spreading by RARα‐ and RARβ‐dependent signaling. This may be consistent with the reports that the agonists of RARα and RARβ, but not RARγ, activated focal adhesion kinase (FAK) and paxillin in breast cancer cells 32. Concerns may be raised for using ATRA since it was previously shown to have negative effects on the adhesion and migration of epithelial cells 33. Treatment of 0. 1–1 μ m ATRA for 1 h significantly inhibited the adhesion of retinal pigment epithelial cells. Furthermore, it was found that ATRA significantly inhibited the spreading of retinal pigment epithelial cells with suppressed FAK, suggesting that ATRA's effect is highly cell type‐dependent. Consequently, caution must be taken for extrapolating these data to osteogenic cells. Another concern may be that ATRA may inhibit the osteogenic activities of pre‐osteoblasts 34 and bone marrow stromal cells 35. In our previous studies, we showed that a long‐term (3–21 days) treatment with ATRA significantly reduced cell proliferation, metabolic activity, protein expression, osteocalcin expression, and extracellular matrix mineralization of osteogenic cells 20, 34, 35. In our current study, we showed that the short‐term (2 h) treatment of either Hst1 alone or ATRA alone did not result in significantly higher DNA content compared to the control. Surprisingly, the DNA content in the group of the co‐administered ATRA and Hst1 was significantly higher than those in the groups of ATRA alone, Hst1 alone, or control, which suggested the co‐administration of ATRA and Hst1 synergistically promoted cell attachment. Our data further showed that the 2‐h co‐administration of ATRA and Hst1 resulted in significantly enhanced metabolic activity of pre‐osteoblasts within the monitoring time span (5 days) than either ATRA or Hst1 alone. Moreover, neither ATRA alone nor Hst1 alone had any effect on ALP activity, suggesting that neither of them significantly influence osteogenic differentiation of pre‐osteoblasts. In contrast, the combination of ATRA and Hst1 significantly enhanced ALP activity. These data indicate that such a treatment with ATRA and Hst1 is potentially suitable to promote both the cell–substrate interactions of pre‐osteoblasts and enhance their osteogenic differentiation. The underlying molecular mechanisms of ATRA and Hst1 co‐administration on ALP activity remain to be elucidated. Possibly, the activation of p38 MAPK signaling pathway may be involved. Recently, we found that specific p38 MAPK inhibitors abolish the promoting effect of Hst1 (data not shown) in this type of pre‐osteoblasts, suggesting that Hst1 could activate p38 signaling. It is well‐established that p38 MAPK is a key mediator for many drugs to upregulate ALP activity in pre‐osteoblasts 36, 37, 38. On the other hand, ATRA is also found to transiently activate p38 signaling 39, 40. Although a short‐term treatment of either Hst1 or ATRA seemed not sufficient to induce ALP activity, in combination Hst1 and ATRA significantly upregulated ALP (Fig. 5 ), suggesting an additive stimulating effect on p38 MAPK signaling. Further studies are needed to confirm this hypothesis. With the inspiration of the ALP result, we, thereafter, performed an experiment of extracellular matrix mineralization with osteogenic medium to check the effect of the 2‐h co‐administration of ATRA and hst1. We found that the 2‐h co‐administration of ATRA and hst1 was associated with a higher (without statistical difference) mineralization at 21 days post‐treatment than the control group (data not shown). In fact, the result is not so surprising since the effect of a 2‐h treatment can quickly taper during the 21‐day culture period with osteogenic medium (10% FBS‐containing α‐MEM with beta‐glycerophosphate and l ‐ascorbic acid‐2‐phosphate as supplements). Consequently, the short‐term co‐administration of ATRA and hst1 can show a significant effect only in the initial cellular events of osteogenic activities, such as cell adhesion, spreading, proliferation, and early differentiation. For late (e. g. , osteocalcin expression) and final differentiation (extracellular matrix mineralization), osteoinductive growth factors, such as bone morphogenetic proteins, are highly needed. Finally, in this study we used a mouse MC3T3‐E1 cell line. Although MC3T3‐E1 cells are widely used as a cell model for pre‐osteoblasts, these effects must be replicated in systems that bear more relevance for the human physiological situation, such as primary mesenchymal stem cells and osteoblasts. Furthermore, caution should be taken to extrapolate the current in vitro results to in vivo situation. Animal studies are highly needed to confirm the promoting effect of co‐administered ATRA and Hst1. In summary, our current study showed that Hst1 and ATRA co‐administration positively influenced the spreading, cellular metabolic activity, proliferation, and osteogenic differentiation of pre‐osteoblasts. Based on these observations, we postulated that such combined treatment may be supportive for cell‐based bone tissue engineering techniques. Conflict of interest The authors declare no conflict of interest. Author contributions GW and ECIV contributed to conceptualization; WS, GW, HL, AS, and DM participated in investigation; GW and JGMB collected resources; WS, AS, and GW performed formal analysis; KN and FJB curated the data; WS, HL, FJB, and GW wrote the original draft; AS, WS, DM, FJB, JGMB, HL, KN, and ECIV edited the manuscript; and WS, HL, GW, and ECIV acquired funding. Supporting information Fig. S1. Fluorescent micrographs depicting the spreading of pre‐osteoblasts (stained with FITC‐Phalloidin) with or without a treatment with 1 µM ATRA for 3 days. Bar = 50μm. Fig. S2. Graph depicting a point‐counting method to measure the surface area of cell spreading. The grid was randomly put on the light micrographs of cells during spreading for the point‐counting method. The filopodia and lamellipodia (red arrow) was included for calculating the cell spreading area with the exclusion of the relatively constant peri‐nuclear area (within red dot circle). Bar = 50μm. 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10. 1002/2211-5463. 12937
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FEBS Open Bio
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Hyperoside ameliorates periodontitis in rats by promoting osteogenic differentiation of BMSCs via activation of the NF‐κB pathway
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Hyperoside has been reported to possess anti‐inflammatory properties. Here, we confirmed that hyperoside exhibits potential therapeutic properties against periodontitis via promotion of proliferation and osteogenic differentiation of rat bone mesenchymal stem cells through activation of the NF‐κB signaling pathway. Therefore, our study provides evidence that hyperoside may have potential as a novel therapeutic option for periodontitis treatment.
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Abbreviations ALP alkaline phosphatase BMSC bone mesenchymal stem cell CCK‐8 Cell Counting Kit‐8 HGF human gingival fibroblast IL‐17 interleukin‐17 LPS lipopolysaccharide NF‐κB nuclear factor‐κB rBMSC rat bone mesenchymal stem cell SD Sprague Dawley TNF‐α tumor necrosis factor‐α Periodontitis is one of the most common chronic inflammatory diseases and is caused by pathogenic bacteria on the tooth surface characterized with a series of symptoms, including gingival bleeding, periodontal bag, alveolar bone absorption and even tooth loss [ 1, 2 ]. As is well known, the local alveolar bone damage caused by the excessive activation of osteoclasts around the roots is closely related to the occurrence of periodontitis, which is one of the obvious histopathological signs of periodontitis [ 3, 4, 5 ]. Therefore, more and more researchers have paid attention to the damage of bone destruction caused by periodontitis. In recent years, with the development of regenerative medicine and tissue engineering technology, the focus of periodontitis treatment has shifted from the control of soft tissue inflammation to the control of pathological bone loss, so as to achieve periodontal tissue regeneration [ 6, 7, 8 ]. At present, a variety of regeneration methods have been proposed to treat periodontitis, including guided tissue regeneration, application of platelet‐rich plasma, enamel matrix derivative and so on [ 9, 10, 11 ]. However, most of these treatment methods have shown variable and unpredictable clinical results. In addition, factors that affect the inflammatory process, such as poor oral hygiene, extraoral infection and systemic diseases, especially osteoporosis, diabetes and cardiovascular diseases, may be related to this high degree of variability [ 12, 13, 14 ]. Bone mesenchymal stem cells (BMSCs) are a kind of stem cell that can differentiate in many directions and can self‐renew and differentiate into various cell types necessary for tissue regeneration [ 15, 16 ]. Recent studies have also reported the effects of BMSCs on reducing inflammatory response and producing cytokines related to tissue regeneration. When exposed to an inflammatory microenvironment, exogenous BMSCs can promote tissue formation by self‐differentiation or activation of endogenous progenitor cells [ 17 ]. Nevertheless, BMSCs also can be affected by the inflammatory process. Inflammatory or immune‐related cytokines, such as tumor necrosis factor‐α (TNF‐α), interferon‐γ and interleukin‐17 (IL‐17), have been reported to affect the differentiation and induce apoptosis of BMSCs [ 18 ]. Therefore, it is of great clinical significance to clarify the role and mechanisms of BMSCs in periodontal tissue regeneration. Recent studies have demonstrated that the nuclear factor‐κB (NF‐κB) signaling pathway, as one of the main signal transduction pathways, exhibited an essential role in the development and progress of periodontitis [ 19 ]. Without external stimulation, NF‐κB binds to inhibitory factor IκB and exists in the cytoplasm in an inactive state. When host cells are stimulated, the upstream kinase IKK of IκB is activated, further phosphorylating, ubiquitinating and finally degrading IκB. NF‐κB, without IκB inhibition, is activated and rapidly nuclear translocated to participate in the progression of periodontitis [ 20 ]. It has been reported that TLR4/MyD88 can activate the NF‐κB signaling pathway and induce the secretion of IL‐6, IL‐8, inducible nitric oxide synthase (iNos) and cyclooxygenase‐2, leading to periodontitis [ 21 ]. In addition, periodontal pathogens, such as fusobacterium and actinomycetes, can stimulate the NF‐κB signaling pathway of macrophages to induce the expressions of TNF‐α and IL‐6, whereas the NF‐κB‐specific blocker can obviously inhibit this reaction process [ 22 ]. Therefore, the NF‐κB signaling pathway plays an important role in the pathogenesis of periodontitis. Nowadays, activation of the NF‐κB signaling pathway by gene intervention or small‐molecule drugs to activate downstream multiple effectors has become a new target of periodontitis therapy [ 23 ]. In recent years, the application of Chinese herbal medicine in the treatment of periodontitis has attracted the attention of numerous investigators because of its wide source, few side effects and low price [ 24 ]. A great quantity of studies has confirmed that many Chinese herbal medicines have biological effects, such as inhibiting the release of various inflammatory factors, such as IL‐13 and TNF‐α, regulating the differentiation of osteoclasts, adjusting the immune activity of the host and so on [ 25 ]. The effective ingredients with positive activities extracted from Chinese herbal medicine may become a candidate drug for the treatment of periodontitis [ 26 ]. Hyperoside, as an active compound, widely exists in many Chinese herbal medicines and belongs to flavonol glycosides, with physiological activities such as anti‐inflammation, softening blood vessels, diuresis and so on [ 27, 28, 29 ]. However, the detailed role and possible mechanisms of hyperoside in the treatment of periodontitis have not been clarified yet. Therefore, this study was designed to investigate the role and possible mechanisms of hyperoside in periodontitis, and we found that hyperoside could ameliorate periodontitis, which may be achieved by promoting osteogenic differentiation of BMSCs through activating the NF‐κB signaling pathway. Our results provide new insights into the biological functions and underlying mechanisms of hyperoside in periodontitis and will give a novel therapeutic target for treating periodontitis. Materials and methods Experimental animals A total of 24 male Sprague Dawley (SD) rats weighing 200 ± 20 g were purchased from the Animal Center of Nanjing Medical University (Nanjing, Jiangsu province, China), and all animals were housed under standard environmental conditions at controlled temperature (22 ± 2 °C), humidity (50 ± 10%) and light (12 h light–dark cycle) with free access to standard diet and water. All procedures for animal care and use were in accordance with the National Institutes of Health guidelines and approved by the Institutional Animal Care and Use Committee of Nanjing Medical University (1801012). Establishment of the periodontitis model in rats The periodontitis model was established following the protocols described previously [ 30 ]. Twenty‐four SD rats were randomly divided into two groups of 12 rats each (periodontitis group and periodontitis + hyperoside group). First, 24 male SD rats were anesthetized by intraperitoneal injection of 7% chloral hydrate at a dose of 3 mg·kg −1 according to body weight. Then SD rats laid on their back, and their limbs were fixed. Iodophor was used to disinfect the area of maxillary anterior teeth, and orthodontic ligation wire (0. 2 mm) was used to wrap the upper anterior teeth. Furthermore, 30 μL lipopolysaccharide (LPS) solution was injected into the ligated gingiva, and then the injection site was pressed with sterile cotton swabs for 30 s; this was done once every 48 h, three times in total. The rats in the periodontitis + hyperoside group were given hyperoside (20 mL, 200 mg·mL −1 ) at the site of LPS injection, whereas the rats in the model group were given corn oil as control. Drug injection was repeated every other day on three separate days. Micro‐CT analysis The mandible specimens were collected and scanned with micro‐CT (Siemens, Munich, Germany). The parameters we set are X‐ray source (80 kV), the 360 rotational steps (500 ms per time) and a node current (500 μA). Images were analyzed with 3D scanning software (DataViewer/CT Analyzer/CT volume). Hematoxylin and eosin and Masson staining assays Tissues were fixed in 4% paraformaldehyde, embedded in paraffin and sectioned into 4 μm. Then the samples were stained with hematoxylin for 10 min and treated with eosin or Masson for 5 min at room temperature. Histological images were acquired under a light microscope (Olympus, Tokyo, Japan) at the magnification of 200×. ELISA assay and alkaline phosphatase activity detection Periodontal tissues were added with PBS (pH 7. 4), followed by homogenization thoroughly with a homogenizer and centrifugation for about 20 min (4 °C, 400 g ). Supernatant was collected to analyze the release of cytokines. The levels of TNF‐α (JEB‐13718; Jin Yibai, Nanjing, China), IL‐1β (JEB‐13503; Jin Yibai), IL‐6 (JEB‐14081; Jin Yibai) and alkaline phosphatase (ALP; Jiancheng, Nanjing, China) were measured according to the manufacturer’s directions. Isolation and culture of BMSCs SD rats were decapitated and killed, and then soaked in 75% alcohol for 15 min. The long bone of SD rats was taken under sterile condition, and the femoral side was removed in D‐Hank’s balanced salt solution containing 100 mg·mL −1 streptomycin and 100 U·mL −1 penicillin (Gibco, Grand Island, NE, USA). Bone marrow was washed with α‐MEM medium and collected in a 25‐cm culture bottle. Then bone marrow was blown repeatedly and slightly to disperse the cells as much as possible and put into a jacket cell incubator at 37 °C with 5% CO 2 for a week [ 31 ]. Identification of BMSCs First, BMSCs were cultured and passaged normally, and the morphological characteristics of BMSCs were observed under the microscope at P0 (Original cells), P1 (First‐passage cells), P2 (Second‐passage cells) and P3 (Third‐passage cells), respectively. Moreover, isolated BMSCs were collected and counted with a density of 1 × 10 6 per milliliter. Following that, BMSCs were incubated with corresponding antibodies, including CD34, CD45, CD73, CD90 and CD105, at 4 °C for 30 min with low‐speed shaking. Subsequently, the cells were washed with PBS three times to remove the antibodies on the cell surface. Then the cells were resuspended by adding 500 mL PBS, and the expressions of cell surface markers were detected by flow cytometer (BD Biosciences, San Jose, CA, USA) [ 32 ]. Cell Counting Kit‐8 assay For cell viability, BMSCs were seeded in 96‐well plates at a density of 1 × 10 4 cells per well for 24 h and then treated with hyperoside with indicated days. Then cell viability was examined by Cell Counting Kit‐8 (CCK‐8) kit (Beyotime Biotechnology, Shanghai, China) based on the specification. The absorbance was detected at 490 nm by microplate reader (Tecan Infinite M200 Micro Plate Reader; LabX, Midland, ON, Canada). EdU assay The proliferation of BMSCs treated with hyperoside was assessed by EdU assay. BMSCs (1 × 10 4 per well) were cultured in 96‐well plates and treated with hyperoside. Then treated BMSCs were fixed with 4% paraformaldehyde, Triton X‐100 was used to permeabilize the nuclear membrane and treated BMSCs were blocked with goat serum for 1 h. Further, treated BMSCs were stained according to the manufacturer’s suggestions. Cell‐cycle distribution and cell apoptosis BMSCs with the density of 1 × 10 6 per milliliter were fixed with precooled 70% EtOH at 4 °C overnight. PBS was used to wash cells three times, and 100 μL RNase A was added to mix the solution fully for 30 min at 37 °C. Forward, 400 μL propidium iodide (PI) dye solution was added to mix well, and incubated for 30 min at 4 °C. Then, the cell‐cycle distribution was detected by flow cytometer (BD Biosciences). The role of hyperoside in the apoptosis of BMSCs was measured by the Annexin V–FITC kit (Beyotime Biotechnology). In brief, BMSCs were treated with hyperoside, harvested with trypsin, washed with PBS and resuspended in 500 μL binding buffer. Next, BMSCs were treated with 5 μL Annexin V–FITC and 10 μL PI for 15 min in the dark. Finally, apoptotic cells were determined by flow cytometer (BD Biosciences). Alizarin red and ALP staining assays BMSCs were collected from each group, washed with PBS and fixed with 10% neutral formalin for 10–30 min. Then the fixed solution was removed, and BMSCs were washed twice by PBS. Further, alizarin red staining solution or ALP staining solution was added to cover the sample for 1–5 min. After that, the cells were washed twice with PBS and observed under inverted microscope. For the quantification of alizarin red, 10% cetylpyridinium chloride was used to dissolve relevant calcium composition. Quantitative real‐time PCR Total RNA was separated from BMSCs and tissues by TRIzol reagent (Invitrogen, Carlsbad, CA, USA), and cDNA was synthesized by TaqMan Reverse Transcription Kit (Applied Biosystems, Foster City, CA, USA). Routine quantitative real‐time PCR was implemented using ABI 7300‐fast RT PCR system (Applied Biosystems) with SYBR Green PCR Kit (Qiagen, Hilden, Germany) based on the specifications. Relative expressions of mRNAs were evaluated by 2 ‐ Δ Δ C t method with Gapdh as internal reference. The primer sequences used were listed as follows: Runx2 forward, 5′‐CATGGCCGGGAATGATGAG‐3′, and reverse, 5′‐TGTGAAGACCGTTATGGTCAAAGTG‐3′; Osx forward, 5′‐TGCCAATGACTACCCACCC‐3′, and reverse, 5′‐TGCCCACCACCTAACCAA‐3′; Alp forward, 5′‐CAGTGGTATTGTAGGTGCTGTG‐3′, and reverse, 5′‐TTTCTGCTTGAGGTTGAG GTTAC‐3′; Ocn forward, 5′‐CATGAGAGCCCTCACA‐3′, and reverse, 5′‐AGAGCGACCCTAGAC‐3′; Gapdh forward, 5′‐TGGAGTCTACTGGCGTCTT‐3′ and reverse, 5′‐TGTCATATTTCTCGTGGT TCA‐3′. Western blot assay Proteins were extracted from BMSCs and tissues by radioimmunoprecipitation assay lysis buffer and quantified by BCA kit (Beyotime Biotechnology). Proteins were separated by 10% SDS/PAGE and transferred into poly(vinylidene difluoride) membranes (Millipore, Bedford, MA, USA). Then the membranes were blocked with 5% nonfat milk in TBST at room temperature for 1 h and treated with the primary antibody at 4 °C overnight. Membranes were washed with TBST three times and probed with horseradish peroxidase‐conjugated secondary antibody (1: 2000, ab6728; Abcam, Cambridge, MA, USA) for 1 h at room temperature. Finally, protein blots were visualized by enhanced chemiluminescence kit (Millipore) and quantified using imagej software (National Institutes of Health, Bethesda, MD, USA, version 4. 3). The primary antibodies were as follows: RUNX2 (ab192256, 1 : 1000; Abcam, Cambridge, UK), OSX (ab22552, 1 : 1000; Abcam), ALP (ab83259, 1 : 1000; Abcam), Dentin sialophosphoprotein (DSPP) (ab216892, 1 : 1000; Abcam), p‐P65 (ab86299, 1 : 1000; Abcam), P65(ab16502, 1 : 1000; Abcam), p‐IκBα (ab133462, 1 : 1000; Abcam), IκBα (ab32518, 1 : 1000; Abcam) and (glyceraldehyde‐3 phosphate dehydrogenase) (ab181602, 1 : 1000; Abcam). Statistical analysis Data were analyzed by graphpad prism 5. 0 (GraphPad Software Inc. , San Diego, CA, USA) and presented as mean ± SD. One‐way ANOVA followed by Tukey’s post hoc test was used to compare differences among multiple groups. P < 0. 05 was indicated as statistically significant. Results Effects of hyperoside on periodontitis in rats To explore the possible role of hyperoside in the development and progression of periodontitis, we constructed the periodontitis rat model and treated with hyperoside. First, we measured the rat body weight before and after treatment with hyperoside, and the data of Fig. 1A showed that the body weight of rats from the model group and hyperoside group was significantly increased compared with those before treatment, but there was no significant difference between the two groups. Then micro‐CT was performed to investigate the effects of hyperoside on the absorption of alveolar bone of rats with periodontal disease, and the data of Fig. 1B indicated that in the model group, alveolar bone resorption reached one‐third of the middle root zone and the root bifurcations were completely exposed, whereas hyperoside treatment could be the absorption of alveolar bone. In addition, hematoxylin and eosin staining was used to investigate the effect of hyperoside on periodontal tissue pathology of rats with periodontal disease, and the data of Fig. 1C showed that the gingival morphology and structure of rats in the model group had disappeared, the epithelium in the sulcus was eroded and a large number of inflammatory cells were infiltrated. In contrast, after hyperoside intervention, the periodontal disease of rats was significantly improved, the gingival morphology and structure of rats were partially disappeared and the infiltration of inflammatory cells was significantly relieved. Besides, Masson staining was adapted to investigate the effects of hyperin on collagen fibers in periodontal disease rats, and the results of Fig. 1D showed that the collagen fibers in periodontal tissues of the model group were disordered and even broken, and after hyperoside intervention, the density of collagen fibers in periodontal tissue was uniform and arranged orderly. Finally, ELISA was carried out to evaluate the effects of hyperoside on the production of proinflammation factors in periodontitis rats, and the data of Fig. 1E demonstrated that the production of TNF‐α, IL‐1β and IL‐6 was overexpressed in periodontitis rats, whereas hyperoside treatment could notably inhibit the production of TNF‐α, IL‐1β and IL‐6. These data suggested that hyperoside ameliorated periodontitis in rats. Fig. 1 Effects of hyperoside on periodontitis in rats. (A) Body weight of rats from the model group and hyperoside group was measured before and after experiment. (B) The absorption of alveolar bone of periodontitis rats treated with hyperoside was detected with micro‐CT. (C) Periodontal tissue pathology of periodontitis rats treated with hyperoside was evaluated by hematoxylin and eosin staining assay. (D) The expression of collagen fibers in periodontitis rats treated with hyperoside was evaluated by Masson staining assay. (E) The levels of TNF‐α, IL‐1 and IL‐6 in periodontitis rats treated with hyperoside were determined by ELISA assay ( n = 3, means ± SD). Data are representative of three independent experiments and were analyzed using ANOVA multiparametric test. *** P < 0. 001. Scale bars: 50 μm. B, bone tissues; cf, collagen fibers. Isolation and identification of rat BMSCs To explore the underlying mechanisms of hyperoside on the development and progression of periodontitis, first, we isolated rat bone mesenchymal stem cells (rBMSCs) from the SD rats, which were further identified with morphological observation and flow cytometry analysis. As shown in Fig. 2A, at the beginning, the cells were suspended in the medium, round and different in size. Then the cells began to adhere to the wall, and the cells were long fusiform and spindle shaped after culture for 6 h. After 48 h of culture, adherent cells were increased and a small number of cell colonies formed. After 7 days, a large number of cells were spindle or spindle shaped, and the number of colonies was increased significantly. After 14 days, the cells were laid on the whole cell culture plate, regular cobblestone‐like. After that time, the cell proliferation was active and the morphology could not be changed significantly. In addition, the data of Fig. 2B indicated that CD73, CD90 and CD105 were positively expressed, whereas CD34 and CD45 were negatively expressed in isolated cells. These results confirmed that the cells isolated from SD rats were rBMSCs. Fig. 2 Isolation and identification of rBMSCs. (A) Morphological observation of rBMSCs of P0‐P3. (B) Flow cytometry analysis showed that rBMSCs were positive for CD73, CD90, CD105, CD34 and negative for CD45. Scale bars: 100 μm. rBMSCs, rat bone mesenchymal stem cells; CD, cluster differentiation Effects of hyperoside on viability, proliferation, apoptosis and cell‐cycle distribution in rBMSCs To investigate the effects of hyperoside on the functions of rBMSCs, first, ALP activity and western blot were performed to evaluate the effects of hyperoside on the expression of ALP, and the data of Fig. 3A, B showed that ALP was highest expressed when the concentration of hyperoside was 40 μg·mL −1. Therefore, hyperoside with the concentration of 40 μg·mL −1 was chosen for further study. Thus, to evaluate whether hyperoside could affect the proliferation of rBMSCs, first, we performed CCK‐8 assay to evaluate the effects of hyperoside on viability of rBMSCs, and the data of Fig. 3C indicated that after treatment with hyperoside, the viability of rBMSCs was obviously enhanced. Besides, EdU assay was adapted to evaluate the suppressive effects of hyperoside on proliferation in rBMSCs. As shown in Fig. 3D, hyperoside could significantly up‐regulate the EdU‐positive cells in rBMSCs. In addition, the effects of hyperoside on cell‐cycle distribution of rBMSCs were evaluated by flow cytometry assay, and the results of Fig. 3E suggested after the intervention of hyperoside, the distribution of rBMSCs in G1 phase was decreased significantly. Further, western blot was used to detect the protein expression levels related to proliferation, including Ki67 and PCNA, and the data of Fig. 3F showed that compared with the blank control group, the expression level of Ki67 and PCNA in the hyperoside intervention group was significantly higher. Moreover, the effects of hyperoside on apoptosis of rBMSCs were evaluated by flow cytometry assay, and the results of Fig. 3G suggested hyperoside had no significant impact on the apoptosis of rBMSCs. These data suggested that hyperoside promoted proliferation in rBMSCs. Fig. 3 Effects of hyperoside on viability, proliferation, apoptosis and cycle distribution in rBMSCs. (A) ALP activity was performed to evaluate the effects of hyperoside with different concentrations on the activity of ALP in rBMSCs. (B) Western blot assay was performed to evaluate the effects of hyperoside with different concentrations on the expression of ALP protein in rBMSCs. (C) CCK‐8 assay performed that the viability of rBMSCs was obviously enhanced in the hyperoside group compared with the control group. (D) EdU assay was performed to evaluate the effects of hyperoside on proliferation in rBMSCs. (E) Flow cytometry assay was performed to evaluate the effects of hyperoside on cell‐cycle distribution in rBMSCs. (F) Western blot assay was performed to evaluate the effects of hyperoside on the protein expressions of Ki67 and PCNA in rBMSCs. (G) Flow cytometry assay was performed to evaluate the effects of hyperoside on apoptosis in rBMSCs ( n = 3, means ± SD). Data are representative of three independent experiments and were analyzed using ANOVA multiparametric test. *P < 0. 05, ** P < 0. 01, *** P < 0. 001, compared with that of control group. Scale bars: 200 μm. Effects of hyperoside on osteogenic differentiation both in vivo and in vitro In the development of periodontitis, the balance between osteoblasts and osteoclasts plays an important role. First, we collected periodontal tissues, and quantitative RT‐PCR and western blot were performed to evaluate the mRNA and protein expressions of RUNX2, OSX, ALP and DSPP, and the data of Fig. 4A, B showed that, compared with the model group, hyperoside could significantly promote the mRNA and protein expressions of RUNX2, OSX, ALP and DSPP. Further, quantitative RT‐PCR and western blot were carried out to detect the effects of hyperoside on the expressions of osteogenic differentiation‐related indexes, including RUNX2, OSX, ALP and OCN in rBMSCs, and the data of Fig. 4C, D suggested that hyperoside significantly promoted the mRNA and protein expressions of RUNX2, OSX, ALP and OCN. Further, alizarin red staining was used to investigate the effects of hyperoside on the mineralization of rBMSCs, and the results of Fig. 4E showed that calcium deposition was produced in rBMSCs after hyperoside intervention. These data suggested that hyperoside promoted osteogenic differentiation both in vivo and in vitro. Fig. 4 Effects of hyperoside on osteogenic differentiation both in vivo and in vitro. (A, B) Quantitative RT‐PCR (A) and western blot assays (B) were performed to evaluate the mRNA and protein expressions of RUNX2, OSX, ALP and DSPP in periodontitis rats treated with hyperoside. (C, D) Quantitative RT‐PCR (C) and western blot (D) were carried out to detect the effects of hyperoside on mRNA and protein expressions of RUNX2, OSX, ALP and OCN in rBMSCs. (E) Alizarin red staining was used to investigate the effects of hyperoside on the mineralization in rBMSCs ( n = 3, means ± SD). Data are representative of three independent experiments and were analyzed using ANOVA multiparametric test. **P < 0. 01, *** P < 0. 001 compared with that of the model group or control group. Scale bars: 200 μm. Effects of hyperoside on the expression of the NF‐κB signaling pathway both in vivo and in vitro The NF‐κB signaling pathway, as one of the main signal transduction pathways, exhibited an essential role in the development and progression of periodontitis. First, western blot was performed to evaluate the effects of hyperoside on the expressions of the NF‐κB signaling pathway in vivo, and the data of Fig. 5A, C showed that, compared with the model group, hyperoside could significantly promote the expressions of the NF‐κB signaling pathway. Similarly, western blot was carried out to detect the effects of hyperoside on the expressions of the NF‐κB signaling pathway in vitro, and the data of Fig. 5B, D showed that hyperoside could significantly promote the expressions of the NF‐κB signaling pathway. These data suggested that hyperoside activated the NF‐κB signaling pathway both in vivo and in vitro. Fig. 5 Effects of hyperoside on the expression of the NF‐κB signaling pathway both in vivo and in vitro. Western blot was performed to evaluate the effects of hyperoside on the expressions of the NF‐κB signaling pathway both (A) in vivo and (B) in vitro. (C, D) Histograms showed the quantification of band intensities in (A) and (B), respectively ( n = 3, means ± SD). Data are representative of three independent experiments and were analyzed using ANOVA multiparametric test. ** P < 0. 001. Effects of hyperoside on biological functions in rBMSCs mediated by the NF‐κB signaling pathway Further, to explore whether the NF‐κB signaling pathway was involved in the effects of hyperoside on the biological functions in rBMSCs, we chose BMS345541, a NF‐κB signaling pathway inhibitor, to inhibit the NF‐κB signaling pathway in rBMSCs treated with hyperoside. As shown in Fig. 6A, BMS345541 could inhibit the expressions of the NF‐κB signaling pathway activated by hyperoside. Then the results of CCK‐8 showed that BMS345541 reversed the effects of hyperoside on the viability of rBMSCs shown in Fig. 6B. The results of Fig. 6C, D revealed that BMS345541 restored the effects of hyperoside on the osteogenic differentiation‐related indexes detected by quantitative RT‐PCR and western blot analysis. The data of Fig. 6E, F indicated that BMS345541 restored the effects of hyperoside on the osteogenic differentiation and cell mineralization detected with ALP and alizarin red staining assays. These data suggested that hyperoside promoted proliferation and osteogenic differentiation in rBMSCs mediated by the NF‐κB signaling pathway. Fig. 6 Effects of hyperoside on biological functions in rBMSCs mediated by the NF‐κB signaling pathway. (A) Western blot was performed to evaluate the effects of BMS345541 on the expressions of the NF‐κB signaling pathway in rBMSCs. (B) CCK‐8 assay showed that BMS345541 reversed the effects of hyperoside on the viability of rBMSCs. (C, D) Quantitative RT‐PCR (C) and western blot (D) were carried out to detect the effects of BMS345541 on mRNA and protein expressions of RUNX2, OSX, ALP and OCN in rBMSCs. (E) ALP staining was performed to evaluate the effects of BMS345541 on osteogenic differentiation in rBMSCs. (F) Alizarin red staining indicated that BMS345541 restored the effects of hyperoside on the cell mineralization of rBMSCs ( n = 3, means ± SD). Data are representative of three independent experiments and were analyzed using ANOVA multiparametric test. ** P < 0. 01, *** P < 0. 001 compared with that of the control group; ## P < 0. 01, ### P < 0. 001 compared with that of the hyperoside group. Scale bars: 200 μm. Discussion At present, some drugs, which have been widely used for the treatment of periodontitis clinically, have a certain toxic side effect, such as changes of oral microflora, further leading to tooth discoloration. In addition, they can cause gastrointestinal disorders, usually accompanied by diarrhea, nausea and vomiting, or rarely phototoxicity and accumulation in bones and teeth [ 33, 34 ]. Based on these, a large number of studies have shown that the application of Chinese herbal medicine is a new trend in the prevention and treatment of periodontitis, with fewer adverse effects on patients. For example, paeonol exhibited protective effects against periodontitis through regulation of the Nrf2/NF‐κB/NFATc1 signaling pathway [ 35 ]. Standardized Boesenbergia pandurata extract and its active compound panduratin A could obviously improve LPS‐induced periodontal inflammation and alveolar bone loss in rats [ 36 ]. Moreover, mangiferin significantly ameliorated Porphyromonas gingivalis ‐induced experimental periodontitis by regulating NF‐κB and JAK1–STAT1/3 signaling pathway [ 37 ]. Therefore, more and more studies on natural products or phytochemicals are expected to become acceptable substitutes or supplements for the treatment of periodontitis. Hyperoside is a kind of traditional Chinese medicine with anti‐inflammatory, immunoregulatory and antioxidant activities, and hyperoside can play a role by regulating the transmission of inflammatory factors, immune molecules and related signal pathways, which has been widely used in clinic [ 27, 28, 29 ]. Therefore, this study was designed to investigate the possible role and underlying mechanisms of hyperoside in the development and progression of periodontitis, and the in vivo study results showed that hyperoside obviously ameliorated periodontitis achieved by improving absorption of alveolar bone, inflammatory infiltration and collagen fibers. These data were similar to previous studies, which further confirmed the positive effects of hyperoside on periodontitis. Periodontal therapy aims to achieve complete regeneration of these structures, especially alveolar bone. So far, several methods have been used to achieve this goal, such as the use of various bone grafts, growth factors and barrier membranes [ 38 ]. The regeneration of alveolar bone depends on the differentiation of BMSCs or periodontal ligament stem cells into osteoblasts, and osteoblasts are the key to the formation of new bones. Inducing and activating more osteoblasts can promote the regeneration of alveolar bone [ 39 ]. BMSCs have a strong ability of self‐renewal, so that they can produce the same cells as themselves in a long period of time. Studies have shown that some aging BMSCs still have strong proliferation ability, even at 80 ℃ [ 40 ]. Cell proliferation is related to many factors, not only the influence of cytokines but also the culture matrix and cell density. BMSCs can obtain higher amplification at low cell density (1. 5–3 per cm 2 ), whereas at high density (12 per cm 2 ), the results were just the opposite. These experimental results showed that the self‐renewal ability of BMSCs in vivo and in vitro may not be the same [ 41 ]. Although BMSCs have a wide range of proliferation and differentiation ability in vitro, BMSCs are generally in a static state in vivo. Only specific environmental changes or external stimuli, such as damage and tissue degeneration, can activate BMSC proliferation and directional differentiation signal transduction [ 42, 43 ]. As expected, in our study, we found that hyperoside can promote the proliferation of BMSCs. Multidifferentiation potential is considered to be an important biological feature of BMSCs. In vitro experiments showed that BMSCs can be differentiated into many known cell lineages under appropriate induction conditions, such as osteoblasts, adipocytes, chondrocytes, cardiomyocytes and so on [ 44, 45 ]. The process of osteogenesis in vitro has four stages: transformation, cell proliferation, cell aggregation and secretion, and extracellular matrix calcification. Maniatopoulos et al. [ 46 ] reported for the first time that BMSCs obtained from bone marrow of adult rats can form calcified bonelike tissue in vitro, and X‐ray diffraction analysis confirmed that the calcified tissues have a hydroxyapatite‐like structure, which proved that BMSCs cultured in vitro can differentiate into osteoblasts. Ouyang et al. [ 47 ] reported that BMSCs grew in sheet shape and arranged closely after adding ascorbic acid into the culture medium. In addition, BMSCs were implanted into the damaged site by combining with the grafted bone, and after 3 weeks, the structure of the implant was similar to that of the normal periosteum, and osteogenic differentiation was observed by morphological, histological and immunohistochemical methods [ 47 ]. Interestingly, hyperoside could promote osteogenic differentiation of BMSCs. Similarly, in vivo studies showed that hyperoside could promote osteogenic differentiation of periodontal tissues. NF‐κB, as an excellent aggregation point of various signal transduction pathways, plays an important role in immune response, inflammatory response, cell proliferation and apoptosis. Sustained or overexpressed NF‐κB is closely related to the occurrence and development of inflammatory diseases. The expressions of NF‐κB transcription factor (p50/p65) and IκB in the nucleus and cytoplasm of healthy patients and periodontal tissues were detected by Huang et al. [ 48 ], and it has been found that the activation rate of NF‐κB in periodontal tissues (75. 90%) was significantly higher than that in normal tissues (5. 30%), whereas the expression of IκB periodontal tissues (5%) was significantly lower than that in normal tissues (50%), showing that NF‐κB played an important role in periodontitis. It has been demonstrated that dietary flavonoid hyperoside may have potential as a therapeutic agent for NF‐κB activation [ 49 ]. In the study of IL‐1β‐induced inflammatory mediators and NF‐κB‐induced effects on human gingival fibroblasts (HGFs), Vardar‐Sengul et al. [ 50 ] found that the gene expression changes induced by IL‐1β were consistent with the pathological changes of periodontitis, including the increase of inflammatory factors, driving factors, transcription factors, matrix metalloproteinases and adhesion molecules, especially the increase of NF‐κB‐dependent antiapoptotic factors. Therefore, activation of NF‐κB in periodontitis can prevent apoptosis and decrease HGF activity, indicating that NF‐κB played a role in HGF [ 50 ]. In this study, we found that hyperoside promoted the expressions of the NF‐κB signaling pathway both in vivo and in vitro. To further verify role of the NF‐κB signaling pathway in occurrence and development of periodontitis, we chose a NF‐κB signaling pathway inhibitor, BMS345541, to inhibit the NF‐κB signaling pathway, and we found that BMS345541 reversed the effects of hyperoside on the biological functions of BMSCs. In conclusion, we confirmed that hyperoside exhibited potential therapeutic properties against periodontitis by promoting proliferation and osteogenic differentiation of BMSCs by activation of the NF‐κB signaling pathway. Therefore, our study provides evidence that hyperoside may emerge as a therapeutic option for periodontitis treatment. Conflict of interest The authors declare no conflict of interest. Author contributions TX and XW conceived and designed the study, collected and assembled data, and wrote the manuscript; these two authors contributed equally to this work. YY and ZZ performed data analysis and interpretation. CY and NZ reviewed the data. JY conceived and designed the study, provided financial support and study material, performed data analysis and interpretation, and approved the final version of the manuscript. All authors read and approved the manuscript.
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10. 1002/2211-5463. 13002
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FEBS Open Bio
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Platelet lysate induces chondrogenic differentiation of umbilical cord‐derived mesenchymal stem cells by regulating the lncRNA H19/miR‐29b‐3p/SOX9 axis
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H19 expression is induced in platelet lysate‐stimulated human umbilical cord‐derived mesenchymal stem cells (hUCMSCs), and H19 acts as a sponge of miR‐29b‐3p to up‐regulate the expression of the master chondrocyte transcription factor SRY‐related high‐mobility‐group box 9, thereby promoting platelet lysate‐induced chondrogenic differentiation of hUCMSCs.
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Abbreviations hUCMSC human umbilical cord‐derived mesenchymal stem cell lncRNA long noncoding RNA MSC mesenchymal stem cell Mut mutated NC negative control OA osteoarthritis PL platelet lysate PRP platelet‐rich plasma qRT‐PCR quantitative real‐time PCR ROCR regulator of chondrogenesis RNA SOX9 SRY‐related high‐mobility‐group box 9 TGF‐β transforming growth factor β WT wild‐type Osteoarthritis (OA) is one of the major joint diseases worldwide mainly characterized by articular cartilage degeneration [ 1 ]. Articular cartilage is an important component of the synovial joints and lacks the capacity for self‐regeneration [ 2 ]. Mesenchymal stem cells (MSCs) have emerged as a relevant cell source for cartilage repair in recent years [ 3, 4 ]. Zheng et al. [ 5 ] observed that human chondrocytes could be successfully induced by coculture of human umbilical cord‐derived mesenchymal stem cells (hUCMSCs) with rabbit chondrocytes. Platelet‐rich plasma (PRP), a blood extraction product widely used in clinical practice, can not only enhance tissue healing ability and promote tissue cell proliferation and differentiation but also greatly accelerates the healing process of tissue injuries [ 6, 7, 8 ]. Platelet lysate (PL), a purified product of PRP after lysis [ 9 ], contains high levels of growth factors important for MSC proliferation and differentiation, such as transforming growth factor β (TGF‐β), fibroblast growth factor and platelet‐derived growth factor [ 10 ]. Among these, TGF‐β is critically important to induce chondrogenic differentiation of MSCs [ 11 ]. A previous study demonstrated that PL induced chondrogenic differentiation of hUCMSCs, concomitant with elevated expression of SRY‐related high‐mobility‐group box 9 (SOX9), a cartilage‐specific transcription factor that plays crucial roles in chondrocyte differentiation [ 12 ]. However, the underlying mechanism is still not clear. Long noncoding RNAs (lncRNAs) are non‐protein‐coding transcripts longer than 200 nucleotides and play important regulatory roles in various physiological and disease processes [ 13 ]. Recent studies have suggested that lncRNA plays a critical role in chondrogenic differentiation of MSCs. For example, Barter et al. [ 14 ] reported that lncRNA regulator of chondrogenesis RNA (ROCR) induced chondrogenic differentiation of MSCs by up‐regulating SOX9 expression. H19, a kind of lncRNA, has been shown to promote the osteogenic differentiation of MSCs [ 15, 16 ] but inhibit the differentiation MSCs into adipocytes [ 17 ]. Existing evidence has demonstrated that H19 expression can be induced by TGF‐β that presents in PL [ 10, 18 ], suggesting that H19 might be involved in the PL‐induced chondrogenic differentiation of hUCMSCs. In recent years, many studies have indicated that H19 is implicated in tenogenic differentiation [ 19 ], tumor cell epithelial–mesenchymal transition [ 20, 21 ] and myocardial ischemia–reperfusion injury [ 22 ] by functioning as a miR‐29b‐3p sponge via competing endogenous RNA activity. It should be noted that miR‐29b‐3p is predicted to bind to the 3′‐UTR of the master chondrocyte transcription factor SOX9 by bioinformatics software ( http://starbase. sysu. edu. cn/ ). Thus, we hypothesized that H19 might be involved in PL‐induced chondrogenic differentiation of hUCMSCs by regulating the miR‐29b‐3p/SOX9 axis. Materials and methods Preparation of human PLs Human PLs were prepared from PRP obtained from five donors as previously described [ 12 ]. In brief, 20 mL of peripheral blood was collected from each healthy volunteer and pooled. PRP was isolated from peripheral blood and then subjected to three cycles of freeze–thaw (frozen at −80 °C and thawed at 37 °C) to obtain PL. PL was centrifuged at 15 000 g for 20 min at 4 °C, filtered through 0. 22‐μm filters and stored at −20 °C until use. Written informed consent was obtained from all volunteers. This experiment was approved by the Ethics Committee of First Affiliated Hospital of Harbin Medical University. The study methodologies conformed to the standards set by the Declaration of Helsinki. Isolation of hUCMSCs hUCMSCs were isolated from umbilical cords of five cesarean‐delivered full‐term neonates using the explant method [ 12 ]. Informed consent was obtained from the parent/legal guardian. In brief, umbilical cords were cut into smaller pieces of ~5 × 5 cm, transferred to a 25‐cm 2 flask and maintained in Dulbecco’s modified Eagle’s medium (Gibco, Thermo Fisher Scientific, Inc. , Waltham, MA, USA) supplemented with 5% PL, 100 μg·mL −1 streptomycin and 100 U·mL −1 penicillin at 37 °C in a humidified atmosphere containing 5% CO 2. The medium was replaced for the first time after 7 days of culture and thereafter every 3–4 days. The umbilical cord segments were removed on the 14th day, and adherent cells reaching 75–85% confluence were detached by 0. 05% trypsin/EDTA and further cultivated until the third passage. hUCMSCs at passage 2 were used in the following experiments. Immunophenotypic characterization of hUCMSCs hUCMSCs at passage 2 were characterized by flow cytometry. In brief, cells were washed twice with PBS, digested by 0. 05% trypsin and suspended at a concentration of 5 × 10 5 cells per tube. Cells were then stained at 4 °C with the following antibodies: CD34‐phycoerythrin (Invitrogen, Carlsbad, CA, USA), CD45‐FITC (Abbexa, Cambridge, UK), CD44‐FITC (Miltenyi Biotec, San Diego, CA, USA) and CD105‐phycoerythrin (Invitrogen). After 30 min of incubation in the dark, cells were washed with PBS and analyzed using flow cytometry (Beckman Coulter, Brea, CA, USA). Chondrogenic differentiation hUCMSCs at passage 2 (1 × 10 6 cells) were cultured in a chondrogenic differentiation medium supplemented with Dulbecco’s modified Eagle’s medium, 1 × insulin‐transferrin‐selenium, 50 mg·L −1 ascorbic acid, 100 n m dexamethasone and 5% PL for 21 days. The culture medium was replaced twice a week. Cell infection and transfection The miR‐29b‐3p precursor‐expressing lentiviral particles (Lenti‐miR‐29b‐3p) and control lentivirus particles (Lenti‐miR‐NC), H19‐expressing lentivirus (Lenti‐H19) and control lentivirus particles (Lenti‐NC), and sh‐H19‐expressing lentivirus (Lenti‐sh‐H19) and sh‐NC‐expressing lentivirus (Lenti‐shRNA) were purchased from Hanbio (Shanghai, China). The miR‐29b‐3p mimic, mimic NC, miR‐29b‐3p inhibitor and inhibitor NC were purchased from GenePharma (Shanghai, China). The hUCMSCs at 70–80% confluence were infected with these lentiviruses or transfected with the mimics and/or inhibitors using Lipofectamine 3000 (Invitrogen) according to the manufacturer’s instructions. Cells were harvested 48 h later and subjected to further experiments. Alcian blue staining After infection with the indicated lentiviruses, hUCMSCs were cultured in a chondrogenic differentiation medium for 21 days as described earlier. Then chondrogenic differentiation was assessed by Alcian blue staining. In brief, the pellets were fixed, dehydrated, embedded in paraffin and sectioned at 5 μm thickness. For detection of proteoglycan and mucopolysaccharides, sections were deparaffinized, rehydrated and then stained with 1% Alcian blue (Sigma‐Aldrich, St. Louis, MO, USA) for 30 min. Luciferase reporter assay The fragments of H19 containing the predicted wild‐type (WT) or mutated (Mut) miR‐29b‐3p binding sites were amplified by PCR and inserted into a pGL3‐luciferase vector (Promega, Madison, WI, USA), generating pGL3‐H19 WT/pGL3‐H19 Mut luciferase constructs. Similarly, the fragments of SOX9 3′‐UTR containing the predicted WT or Mut miR‐29b‐3p binding sites were amplified by PCR and inserted into a pGL3‐luciferase vector, generating pGL3‐SOX9 WT or pGL3‐SOX9 Mut constructs. The primers were as follows: H19 WT construct primer forward, 5′‐ ACTGGAATTCTTACTTCCTCCACGGAGTCG‐3′; H19 WT construct primer reverse, 5′‐ACTGCTCGAGTGTTCCGATGGTGTCTTTGA‐3′; SOX9 WT 3′‐UTR construct primer forward, 5′‐ACTGGAATTCTCAGGCTTTGCGATTTAAGG‐3′; SOX9 WT 3′‐UTR construct primer reverse, 5′‐ ACTGCTCGAGAGGCAGGAGGAAATGCACTA‐3′. HEK 293T cells at 75–85% confluence were cotransfected with these constructs, miR‐29b‐3p mimic/mimic NC and pRL‐TK (a plasmid expressing Renilla luciferase) using Lipofectamine 2000 (Invitrogen). The luciferase activities 24 h after transfection were analyzed using a luciferase reporter assay system (Promega). Quantitative real‐time PCR Total RNA was extracted from cultured hUCMSCs using TRIzol Reagent (Invitrogen) and subjected to reverse transcription reactions using a PrimeScript RT Reagent Kit (Takara, Dalian, China). The expression levels of miR‐29b‐3p were determined using the miRNA quantitative real‐time PCR (qRT‐PCR) kit (GeneCopoeia, Rockville, MD, USA), whereas the expression levels of H19 and SOX9 were detected using the SYBR premix (Takara) on Applied Biosystems 7500 PCR system. Relative expression levels of candidate genes were calculated by the 2 ‐ Δ Δ C t method and normalized to U6 (internal control for miR‐29b‐3p) or GAPDH (internal control for H19 and SOX9). The primers were as follows: H19 forward, 5′‐TCTGAGAGATTCAAAGCCTCCAC‐3′; H19 reverse, 5′‐GTCTCCACAACTCCAACCAGTG‐3′; SOX9 forward, 5′‐ AGGAAGCTCGCGGACCAGTAC‐3′; SOX9 forward, 5′‐ GGTGGTCCTTCTTGTGCTGCAC‐3′; miR‐29b‐3p forward, 5′‐ACACTCCAGCTGGGTAGCACCATTTGAAATC‐3′; miR‐29b‐3p reverse, 5′‐ TGGTGTCGTGGAGTCG‐3′; GAPDH forward, 5′‐ GTCTCCTCTGACTTCAACAGCG‐3′; GAPDH reverse, 5′‐ ACCACCCTGTTGCTGTAGCCAA‐3′; U6 forward, 5′‐ TGCGGGTGCTCGCTTCGGCAGC‐3′; U6 reverse, 5′‐GTGCAGGGTCCGAGGT‐3′. Western blot The whole‐cell lysates were extracted from hUCMSCs using the radioimmunoprecipitation assay lysis buffer (Beyotime, Haimen, China). Then the protein samples were loaded and separated with 10% SDS/PAGE gels and transferred to poly(vinylidene difluoride) membrane. After being blocked with 5% nonfat milk, the membranes were incubated with primary antibodies (1 : 1000) at 4 °C overnight, followed by the horseradish peroxidase‐conjugated secondary antibodies (1 : 2000; Abcam, Cambridge, MA, USA). Blots were examined by an ECL Detection kit (Pierce Biotechnology, Rockford, IL, USA). The primary antibodies used were as follows: anti‐collagen II (#ab239007), anti‐aggrecan (#ab36861), anti‐SOX9 (#ab185230) and anti‐α‐tubulin (#ab7291) (1 : 1000 dilution; all from Abcam). Statistical analysis All statistical analyses were performed using spss 22. 0 (SPSS, Inc. , Chicago, IL, USA). Statistical significance was analyzed using the Student's t ‐test between two groups and one‐way ANOVA among three or more groups. A P value <0. 05 was considered statistically significant. Results H19 expression was increased during PL‐induced chondrogenic differentiation of hUCMSCs hUCMSCs were isolated from umbilical cords of cesarean‐delivered full‐term neonates and identified by flow cytometry. Data revealed that the obtained hUCMSCs were positive for CD105 and CD44, but negative for CD45 and CD34 (Fig. 1A ), confirming that the cells were MSCs. hUCMSCs were then cultured in 5% PL‐supplemented chondrogenic differentiation medium for 0, 7, 14 and 21 days. The results from qRT‐PCR analysis demonstrated significant up‐regulation of H19 expression during PL‐induced chondrogenic differentiation of hUCMSCs (Fig. 1B ). Fig. 1 H19 expression was increased during PL‐induced chondrogenic differentiation of hUCMSCs. (A) hUCMSCs at passage 2 were characterized by flow cytometry. Forward scatter (FSC) and side scatter (SSC) distribution of gated cells (left). The cells were positive for CD105 and CD44, but negative for CD45 and CD34 (right). (B) qRT‐PCR analysis of H19 expression in hUCMSCs cultured in 5% PL‐supplemented chondrogenic differentiation medium for 0, 7, 14 and 21 days. * P < 0. 05, ** P < 0. 01, vs. 0 day (ANOVA). The quantitative statistics were presented as the mean ± standard deviation ( n = 3). H19 promoted PL‐induced chondrogenic differentiation of hUCMSCs To determine the role of H19 in PL‐induced chondrogenic differentiation of hUCMSCs, we overexpressed and silenced H19 in hUCMSCs, and then cultured them in 5% PL‐supplemented chondrogenic differentiation medium for 21 days. The overexpression and silencing efficiencies of H19 were confirmed by qRT‐PCR analysis (Figs 2A and S1 A). Importantly, Alcian blue staining demonstrated that H19 overexpression led to increased production of proteoglycan in the differentiated cells, whereas H19 silencing yielded the opposite results (Fig. 2B ). Furthermore, the protein levels of chondrogenic markers (collagen II and aggrecan) in the differentiated cells were significantly up‐regulated by H19 overexpression but down‐regulated by H19 silencing (Fig. 2C ). These findings indicated that H19 promoted PL‐induced chondrogenic differentiation of hUCMSCs. Fig. 2 Effects of H19 overexpression and silencing on PL‐induced chondrogenic differentiation of hUCMSCs. (A) qRT‐PCR analysis of H19 expression in hUCMSCs at 48 h after infection with Lenti‐NC, Lenti‐H19, Lenti‐shRNA and Lenti‐sh‐H19. (B, C) The infected hUCMSCs as described in (A) were cultured in 5% PL‐supplemented chondrogenic differentiation medium for 21 days. (B) Chondrogenic differentiation was assessed by Alcian blue staining (scale bar: 100 μm). (C) The protein levels of collagen II and aggrecan in hUCMSCs were examined by western blot. * P < 0. 05, ** P < 0. 01, vs. Lenti‐NC; # P < 0. 05, ## P < 0. 01, vs. Lenti‐shRNA (Student's t ‐test). The quantitative statistics were presented as the mean ± standard deviation ( n = 3). H19 acted as a sponge of miR‐29b‐3p to up‐regulate SOX9 expression Next, we explored the molecular mechanism by which H19 promotes PL‐induced chondrogenic differentiation of hUCMSCs. The results from qRT‐PCR analysis revealed that miR‐29b‐3p expression was notably down‐regulated, whereas SOX9 mRNA and protein levels were significantly up‐regulated during PL‐induced chondrogenic differentiation of hUCMSCs (Fig. 3A–C ). Furthermore, H19 overexpression led to a marked increase in SOX9 mRNA and protein levels. In contrast, H19 silencing significantly down‐regulated SOX9 expression (Fig. 3D–E ). In addition, the luciferase activity was significantly decreased in cells cotransfected with H19 WT reporter and miR‐29b‐3p mimic (Fig. 3F ), indicating that H19 directly interacted with miR‐29b‐3p. We also observed decreased luciferase activity in cells cotransfected with SOX9 WT reporter and miR‐29b‐3p mimic (Fig. 3G ), suggesting that 3′‐UTR of SOX9 was directly targeted by miR‐29b‐3p. Moreover, transfection with miR‐29b‐3p mimic resulted in a significant decrease in mRNA and protein levels of SOX9. On the contrary, inhibition of miR‐29b‐3p noticeably increased SOX9 expression (Fig. 4A, B ). Together, these results manifested that H19 might act as a sponge of miR‐29b‐3p to up‐regulate SOX9 expression. Fig. 3 Interaction between miR‐29b‐3p and H19 or SOX9. qRT‐PCR analysis of (A) miR‐29b‐3p expression and (B) SOX9 mRNA level, and (C) western blot analysis of SOX9 protein level in hUCMSCs cultured in 5% PL‐supplemented chondrogenic differentiation medium for 0, 7, 14 and 21 days. * P < 0. 05, ** P < 0. 01, vs. 0 day (Student's t ‐test). (D, E) qRT‐PCR analysis of SOX9 mRNA level (D) and western blot analysis of SOX9 protein level (E) in hUCMSCs infected with Lenti‐NC, Lenti‐H19, Lenti‐shRNA and Lenti‐sh‐H19. ** P < 0. 01, vs. Lenti‐NC; ## P < 0. 01, vs. Lenti‐shRNA (Student's t ‐test). (F) The WT (H19 WT) and mutation (H19 Mut) of binding sites between H19 and miR‐29b‐3p. Results of luciferase activity assay verified the direct binding between H19 and miR‐29b‐3p. (G) The WT (SOX9 WT) and mutation (SOX9 Mut) of binding sites between SOX9 and miR‐29b‐3p. Results of luciferase activity assay verified the direct binding between SOX9 3′‐UTR and miR‐29b‐3p. ** P < 0. 01, vs. mimic NC (Student's t ‐test). The quantitative statistics were presented as the mean ± standard deviation ( n = 3). Fig. 4 H19 promoted PL‐induced chondrogenic differentiation of hUCMSCs by regulating the miR‐29b‐3p/SOX9 axis. (A, B) qRT‐PCR analysis of SOX9 mRNA level (A) and western blot analysis of SOX9 protein level (B) in hUCMSCs transfected with miR‐29b‐3p mimic, miR‐29b‐3p inhibitor or corresponding controls. ** P < 0. 01, vs. mimic NC; ## P < 0. 01, vs. inhibitor NC (Student's t ‐test). (C, D) hUCMSCs were coinfected with Lenti‐NC/Lenti‐H19 and Lenti‐miR‐NC/Lenti‐miR‐29b‐3p, and then cultured in 5% PL‐supplemented chondrogenic differentiation medium for 21 days. (C) Chondrogenic differentiation was assessed by Alcian blue staining (scale bars: 100 μm). (D) The protein levels of SOX9, collagen II and aggrecan in hUCMSCs were examined by western blot. * P < 0. 05, vs. Lenti‐miR‐NC + Lenti‐NC; # P < 0. 05, vs. Lenti‐miR‐NC + Lenti‐H19 (ANOVA). The quantitative statistics were presented as the mean ± standard deviation ( n = 3). H19 promoted PL‐induced chondrogenic differentiation of hUCMSCs by regulating the miR‐29b‐3p/SOX9 axis Finally, we determined whether H19 promotes the PL‐induced chondrogenic differentiation of hUCMSCs by regulating the miR‐29b‐3p/SOX9 axis. To this purpose, we coinfected hUCMSCs with Lenti‐NC/Lenti‐H19 and Lenti‐miR‐NC/Lenti‐miR‐29b‐3p, and then cultured them in 5% PL‐supplemented chondrogenic differentiation medium for 21 days. The results of the qRT‐PCR analysis showed that H19 expression was increased postinfection with Lenti‐H19 but decreased postinfection with Lenti‐miR‐29b‐3p (Fig. S1 B). Alcian blue staining revealed that Lenti‐miR‐29b‐3p infection reduced production of proteoglycan in the differentiated cells and attenuated the H19 overexpression‐mediated promotion of proteoglycan content (Fig. 4C ). Furthermore, the up‐regulation of protein levels of SOX9, collagen II and aggrecan in the H19‐overexpressing cells was rescued following Lenti‐miR‐29b‐3p infection (Fig. 4D ). Collectively, these observations suggested that H19 promoted PL‐induced chondrogenic differentiation of hUCMSCs by regulating the miR‐29b‐3p/SOX9 axis. Discussion MSCs are multipotent stem cells with the potential to differentiate into chondrocytes, and MSC‐derived cartilage tissue engineering is a clinical method used for OA treatment [ 23 ]. PRP is a blood extraction product widely used in clinical practice to accelerate the wound healing process of some tissues, such as muscle [ 6 ], cartilage [ 7 ] and tendon [ 8 ]. Prepared from PRP, PL has been shown to reduce immunogenicity [ 10 ] and promote MSC chondrogenic differentiation, providing broad application prospects for allogeneic or xenogeneic transplantation in cartilage‐related diseases [ 24, 25 ]. A previous study clearly showed that PL induced the differentiation of hUCMSCs into chondrocytes [ 12 ]. In this study, we demonstrated for the first time the molecular mechanism by which PL induces MSC chondrogenic differentiation; that is, PL induced chondrogenic differentiation of hUCMSCs by regulating the H19/miR‐29b‐3p/SOX9 axis. A number of recent studies have suggested that lncRNA plays a critical role in chondrogenic differentiation of MSCs. An increasing number of lncRNAs have been identified with a potential role in chondrogenesis, such as ROCR [ 14 ], differentiation antagonizing non‐protein‐coding RNA [ 26 ], maternally expressed 3 [ 23 ] and ADAMTS9 antisense RNA 2 [ 27 ]. For instance, lncRNA urothelial cancer‐associated 1 (UCA1) promotes chondrogenic differentiation of human bone marrow‐derived MSCs via regulating the miR‐145‐5p/SMAD5 and miR‐124‐3p/SMAD4 axis [ 28 ]. lncRNA H19 has been shown to promote the osteogenic differentiation of MSCs [ 15, 16 ] and inhibit the differentiation MSCs into adipocytes [ 17 ]. Yang et al. [ 18 ] demonstrated that in the bovine mammary alveolar cell–T cell line, TGF‐β induced up‐regulation of H19, which promoted epithelial‐to‐mesenchymal transition and contributed to mammary gland fibrosis. PL contains a relatively large number of growth factors, including TGF‐β that is critical for chondrogenic differentiation [ 10 ]. Thus, we speculated that H19 might be involved in the PL‐induced chondrogenic differentiation of hUCMSCs. To our knowledge, our findings in this study provided the first evidence that H19 overexpression promoted, whereas H19 silencing attenuated the chondrogenic differentiation of PL‐treated MSCs. Evidence indicates that TGF‐β1 up‐regulated H19 expression through activating the phosphoinositide‐3‐kinase/AKT pathway in epithelial cells [ 29 ]. Whether the phosphoinositide‐3‐kinase/Akt pathway or other transcriptional regulation mechanisms are involved in the PL‐induced H19 up‐regulation remains to be further investigated. Several studies have established that H19 regulates the expression and biological functions of certain miRNAs by acting as a competing endogenous RNA or a miRNA sponge [ 30, 31 ]. For example, H19 contributes to oxidative damage repair in the early age‐related cataract by sponging miR‐29a [ 32 ]. miR‐29b‐3p not only plays an important role in tumor initiation and progression [ 33 ] but also participates in DNA damage response [ 34 ], tissue fibrosis [ 35 ] and chondrogenic differentiation [ 36 ]. In this study, we performed bioinformatics analysis and luciferase reporter assay to analyze the interaction between miR‐29b‐3p and H19 or SOX9. Results showed that H19 can function as a sponge RNA for miR‐29b‐3p that negatively regulates the expression of target gene SOX9. SOX9 gene is a cartilage‐specific transcription factor that controls the expression of numerous chondrocyte genes (e. g. collagen II and aggrecan) and plays essential roles in chondrocyte differentiation and cartilage formation [ 37, 38 ]. Our results demonstrated that H19 positively regulated SOX9 expression and promoted MSC chondrogenic differentiation, and these effects were rescued by miR‐29b‐3p mimic. Conclusion The findings in this study demonstrate that H19 expression is induced in PL‐stimulated hUCMSCs and that H19 acts as a sponge of miR‐29b‐3p to up‐regulate the expression of the master chondrocyte transcription factor SOX9, thereby promoting PL‐induced chondrogenic differentiation of hUCMSCs. These data provide new experimental evidence for elucidating the mechanism of MSC in cartilage repair and a novel therapeutic option for OA. Conflict of interest The authors declare no conflict of interest. Author contributions BC designed the study. BC and XD participated in the experiments. BC and XD contributed to the data analysis. BC drafted the paper. All authors approved the paper. Supporting information Fig. S1. Expression of H19 in hUCMSCs postinfection with the indicated lentiviruses. (A) qRT‐PCR analysis of H19 expression in hUCMSCs infected with Lenti‐NC, Lenti‐H19, Lenti‐shRNA and Lenti‐sh‐H19 after culture in 5% PL‐supplemented chondrogenic differentiation medium for 21 days. ** P < 0. 01, vs. Lenti‐NC; ## P < 0. 01, vs. Lenti‐shRNA (Student’s t ‐test). (B) qRT‐PCR analysis of H19 expression in hUCMSCs coinfected with Lenti‐NC/Lenti‐H19 and Lenti‐miR‐NC/Lenti‐miR‐29b‐3p after culture in 5% PL‐supplemented chondrogenic differentiation medium for 21 days. ** P < 0. 01, vs. Lenti‐miR‐NC + Lenti‐NC; # P < 0. 05, vs. Lenti‐miR‐NC + Lenti‐H19 (ANOVA). The quantitative statistics were presented as the mean ± standard deviation ( n = 3). Click here for additional data file.
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10. 1002/2211-5463. 13043
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FEBS Open Bio
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miR‐196b‐5p inhibits proliferation of Wharton's jelly umbilical cord stem cells
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Human umbilical cord mesenchymal stem cells can be obtained from different parts of the umbilical cord, including Wharton's jelly. Transplantation of Wharton's jelly umbilical cord stem cells (WJCMSCs) is a promising strategy for the treatment of various diseases. However, the molecular mechanisms underlying the proliferation of WJCMSCs are incompletely understood. Here, we report that overexpression of miR‐196b‐5p in WJCMSCs suppresses proliferation and arrests the cell cycle in G0/G1 phase, whereas knockdown of miR‐196b‐5p promotes WJCMSC proliferation and cell‐cycle progression. Moreover, miR‐196b‐5p overexpression resulted in decreased levels of Cyclin A, Cyclin D, Cyclin E and cyclin‐dependent kinases 2 and increased levels of p15 INK4b, whereas miR‐196b‐5p knockdown had the opposite effects. In conclusion, our data suggests that miR‐196b‐5p inhibits WJCMSC proliferation by enhancing G0/G1‐phase arrest.
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Abbreviations BMMSC bone marrow mesenchymal stem cell CDK cyclin‐dependent kinase CFSE (CFDA‐SE) 5, 6‐carboxyfluorescein diacetate, succinimidyl ester GAPDH glyceraldehyde‐3 phosphate dehydrogenase iPSC induced pluripotent stem cell MSC mesenchymal stem cell qRT‐PCR quantitative RT‐PCR SA‐β‐gal senescence‐associated β‐galactosidase SWATH‐MS sequential window acquisition of all theoretical mass spectra WJCMSC Wharton's jelly umbilical cord stem cell Mesenchymal stem cells (MSCs) have a wide range of applications in regenerative medicine and immunotherapy because of their good ability of proliferation, multidirectional differentiation and immunoregulation [ 1, 2 ]. With the isolation and identification of bone marrow MSCs (BMMSCs), embryonic stem cells and induced pluripotent stem cells (iPSCs), these stem cells have become an important part of regeneration therapy [ 3, 4, 5 ]. However, the low proportion of BMMSCs in bone marrow, the ethical issues of embryonic stem cells extraction from embryos and the tumorigenicity of iPSCs limit their clinical application [ 6, 7, 8 ]. Therefore, it is imperative to find a useful source of stem cells. Human umbilical cord MSCs can be obtained from different parts of the umbilical cord, including cord lining, perivascular region and Wharton’s jelly [ 9 ]. Wharton’s jelly umbilical cord stem cells (WJCMSCs) were used by most studies for their more prominent biological characteristics [ 10, 11 ]. First, the collection of WJCMSCs process is noninvasive and can prevent the risk for infection. In terms of the biological properties, WJCMSCs have higher proliferation potential than other MSCs. Furthermore, WJCMSCs have higher differentiation ability and faster self‐renewal ability than BMMSCs because of their unique gene expression profile, producing almost no teratomas and maintaining at an earlier embryonic stage [ 12, 13 ]. Moreover, some studies revealed that WJCMSCs are a good candidate for allogeneic transplantation because of their lower immunogenicity than BMMSCs [ 14 ]. The immunological characteristics of WJCMSCs also suggest that they may be a choice to treat autoimmune diseases such as type 1 diabetes [ 15 ]. Due to these advantages, WJCMSCs are the ideal source of stem cell for regenerative medicine and immunotherapy and have broad application prospects. The growth of WJCMSCs was influenced by the culture in vitro. When expanded to passage 10, the growth rate of WJCMSCs was slower than those at passage 5 [ 16 ]. Cell proliferation is important for cell therapy and tissue engineering because sufficient cell numbers are the prerequisite to ensure their regenerative and therapeutic effect. However, the molecular mechanisms of WJCMSC proliferation are vague. Therefore, it is of great significance to explore the molecular mechanisms of WJCMSC proliferation for better application of WJCMSCs in the future. miRNAs are kinds of noncoding single‐stranded RNA molecules, which play a significant role in cell proliferation, cell apoptosis, cell cycle and differentiation by inducing target mRNA cleavage or translational inhibition [ 17, 18, 19, 20 ]. Previously, our miRNA array data showed that some miRNAs, including miR‐196b‐5p, were differentially expressed in WJCMSCs, adipose‐derived stem cells (ADSCs), BMMSCs, stem cells from apical papilla (SCAPs), dental pulp stem cells (DPSCs) and periodontal ligament stem cells (PDLSCs), which was confirmed by real‐time quantitative RT‐PCR (qRT‐PCR) (Fig. S1 ). These differentially expressed miRNAs may be key regulators controlling the proliferation and differentiation of these stem cells and warrant further study. miR‐196b gene is located in highly evolutionarily conserved regions of human chromosome 7 between HOXA9 and HOXA10 genes [ 21 ]. miR‐196b can regulate cell proliferation, invasion and migration in many tumors [ 22, 23 ]. miR‐196 might be involved in the process of chondrogenic differentiation of human iPSCs and was significantly up‐regulated in chondrocytes derived from human iPSCs [ 24 ]. Recently, it was revealed that miR‐196b inhibited human leukemia stem cell growth by targeting Cdkn1b [ 25 ]. However, the function of miR‐196b‐5p in WJCMSC proliferation is unclear. In this study, the physiological function of miR‐196b‐5p on WJCMSC proliferation was investigated. We confirmed that miR‐196b‐5p suppressed cell proliferation and blocked G0/G1 phase of WJCMSCs, indicating that miR‐196b‐5p can be served as a potential target and help for clinic application of WJCMSCs in the future. Materials and methods Cell culture The research involving human stem cells in this study complied with the International Society for Stem Cell Research ‘Guidelines for the Conduct of Human Embryonic Stem Cell Research’. Human WJCMSCs were purchased from ScienCell Research Laboratories (Carlsbad, CA, USA). Cells were cultured as shown previously [ 26 ]. Cells at generations 3–5 were used in the subsequent experiments. Induction of senescence and senescence‐associated β‐galactosidase staining To induce premature senescence, we treated WJCMSCs with H 2 O 2 (100 m m ) for 4 h, washed with PBS and continued to incubate for 24 h. A senescence‐associated β‐galactosidase (SA‐β‐gal) staining kit (Cell Senescence Testing Kit; GenMed Scientifics Inc. , Shanghai, China) was used following the manufacturer’s protocol. In brief, cells were washed and fixed with 1× Fixative Solution for 10 min at room temperature. Then the cells were incubated at 37 °C with β‐galactosidase staining solution (pH 6. 0) for 24 h. The number of SA‐β‐gal‐positive cells was selected in 10 randomly chosen fields, and the percentage of positive cells was calculated from three independent experiments. Synthesis of miRNA and construction The lentivirus miR‐196b‐5p mimic, miR‐196b‐5p inhibitor and negative control (Consh) were obtained from GenePharma (Suzhou, China). Virus transfection was performed as described previously [ 27 ]. The sequences are listed in Table 1. Table 1 Sequences used in the study. F, forward; R, reverse. Gene symbol Sequence (5′–3′) miR‐196b‐5p mimic TAGGTAGTTTCCTGTTGTTGGG miR‐196b‐5p NC mimic GTTVTCCGAACGTGTCACGT miR‐196b‐5p inhibitor CCCAACAACAGGAAACTACCTA miR‐196b‐5p NC inhibitor CAGUACUUUUGUGUAGUACAA miR‐196b‐5p F: GCGTAGGTAGTTTCCTG miR‐196b‐5p R: GAGCAGGCTGGAGAA U6 F: GCTTCGGCAGCACATATACT U6 R: GAGCAGGCTGGAGAA PTGS2 F: ATGCTGACTATGGCTACAAAAGC PTGS2 R: TCGGGCAATCATCAGGCAC PON2 F: GTTGGACCGGCATTTCTAT PON2 R: CATTTGCCCAGTGTAAGTTCAAG METTL3 F: AGATGGGTAGAAAGCCTCCT METTL3 R: TGGTCAGCATAGGTTACAAGAGT GAPDH F: ACAACTTTGGTATCGTGGAAGG GAPDH R: GCCATCACGCCACAGTTTC John Wiley & Sons, Ltd RNA isolation and real‐time qRT‐PCR Total RNA was extracted from transfected WJCMSCs by TRIzol (Invitrogen, Carlsbad, CA, USA). The levels of miRNA or mRNA in WJCMSCs were detected by using a Hairpin‐it™ microRNA and U6 snRNA Normalization RT‐PCR Quantitation Kit (GenePharma) or QuantiTect SYBR Green PCR kit (Qiagen, Hilden, Germany). U6 and GAPDH ( glyceraldehyde‐3 phosphate dehydrogenase ) were used to normalize miRNA or mRNA levels. The primer sequences are listed in Table 1. 5, 6‐Carboxyfluorescein diacetate, succinimidyl ester assay Following the CellTrace™ 5, 6‐carboxyfluorescein diacetate, succinimidyl ester (CFSE) Cell Proliferation Kit protocol (Invitrogen), CFSE assay was determined as shown previously [ 27 ]. In brief, the suspension of cells was labeled with CFSE and inoculated into a six‐well plate at 1. 0 × 10 5 cells per plate for 72 h. The proliferation cells were fixed with formaldehyde and analyzed by flow cytometry (FACSCalibur; BD Biosciences, New Jersey, NJ, USA). Calculation of proliferation index was carried on by ModFit LT (Verity Software House, Topsham, ME, USA). Cell‐cycle assay The transfected WJCMSCs were collected and fixed with cold 70% alcohol at 4 °C overnight, washed and resuspended with PBS. According to the operating protocol, the cells were added with 100 µg·mL −1 RNase A at 37 °C for 30 min and finally 100 µg·mL −1 Propidium (PI) (Sigma‐Aldrich, St. Louis, MO, USA) stained for 20 min away from light at 4 °C. DNA content was evaluated by ModFit LT. The proliferation index was calculated as PI = (S + G2/M)/(G0/G1 + S + G2/M). Western blot Total proteins were resolved from transfected WJCMSCs, and SDS‐polyacrylamide gel tests were performed as shown previously [ 28 ]. The primary antibodies in this study were Cyclin A (Cat. No. SAB4503499; Sigma‐Aldrich), Cyclin D (Cat. No. 05‐152; Merck Millipore, Darmstadt, Germany), Cyclin E (Cat. No. 05‐363; Merck Millipore), cyclin‐dependent kinase 2 (CDK2; Cat. No. 05‐163; Merck Millipore), CDK4 (Cat. No. MAB8879; Merck Millipore), p15 INK4B (Cat. No. 4822; Cell Signaling Technology, Boston, MA, USA) and GAPDH (Cat. No. G8795; Sigma‐Aldrich). imagej 1. 52 V (National Institutes of Health, Bethesda, MD, USA) was used to quantify the bands related to GAPDH expression and to normalize the total protein loaded in each lane. Statistical analysis Statistical calculations were implemented using spss 10. 0 statistical software (SPSS Inc, Chicago, IL, USA). Statistical significance was analyzed by Student’s t ‐test or one‐way ANOVA; P ≤ 0. 05 was regarded as statistically significant. Results The expression of miR‐196b‐5p is increased in senescent cells Senescence was induced with H 2 O 2 in WJCMSCs as previously described to measure the effect of aging on the expression of miR‐196b‐5p [ 29 ]. SA‐β‐gal staining and quantitative analysis results showed that the percentage of SA‐β‐gal + cells in the WJCMSCs + H 2 O 2 group increased significantly compared with the control group (Fig. 1A, B ). qRT‐PCR was used to measure the expression of miR‐196b‐5p in senescent cells, and the results indicated that the level of miR‐196b‐5p in the WJCMSCs + H 2 O 2 group was significantly increased compared with the WJCMSCs group (Fig. 1C ). Fig. 1 The expression of miR‐196b‐5p in senescent cells. (A, B) Senescence‐associated β‐galactosidase (SA‐β‐gal) staining (blue) and quantitative analysis of senescent WJCMSCs induced with H 2 O 2. Scale bar: 100 μm. (C) Real‐time RT‐PCR showed the expression of miR‐196b‐5p increased in senescent WJCMSCs. U6 was used as an internal control for miR‐196b‐5p. Student’s t ‐test was used to analyze the statistical significance. All error bars signify standard deviations ( n = 3). * P ≤ 0. 05, ** P ≤ 0. 01. Overexpression of miR‐196b‐5p suppresses cell proliferation and arrests cell cycle in G0/G1 phase in WJCMSCs To explore the role of miR‐196b‐5p on WJCMSC proliferation, we infected the cells with miR‐196‐5p mimics. After 3 days of 1 μg·mL −1 puromycin treatment, overexpression of miR‐196‐5p in WJCMSCs was confirmed by qRT‐PCR (Fig. 2A ). Then, we conducted CFSE assays, and the results showed that the number of cells proliferating to the sixth generation (Fig. 2B ) and the proliferation index (Fig. 2C ) in the miR‐196b‐5p mimics group were significantly reduced compared with the control group. The results indicated that miR‐196‐5p mimics inhibited cell growth of WJCMSCs. Furthermore, cell‐cycle assays were performed to confirm whether the effect of miR‐196‐5p on WJCMSC growth was associated with cell‐cycle distribution. Compared with the control group, the number of cells in the miR‐196b‐5p‐overexpressed group increased significantly in G0/G1 stage but decreased significantly in S and G2/M stages (Fig. 3A, B ). Moreover, cell‐cycle proliferation index detection results also showed that miR‐196b‐5p mimics suppressed cell proliferation of WJCMSCs (Fig. 3C ). Fig. 2 Overexpression of miR‐196b‐5p suppressed cell proliferation in WJCMSCs. (A) Real‐time RT‐PCR showed the efficiency of miR‐196b‐5p overexpression in WJCMSCs. (B, C) CFSE assay results showed that the number of cells proliferating to the sixth generation and the proliferation index in the miR‐196b‐5p mimics group were lower than those in the control group. U6 was used as an internal control for miR‐196b‐5p. Student’s t ‐test was used to analyze the statistical significance. All error bars signify standard deviations ( n = 3). * P ≤ 0. 05, ** P ≤ 0. 01. Fig. 3 Overexpression of miR‐196b‐5p induced G0/G1 phase arrest in WJCMSCs. (A, B) The flow cytometer analysis results showed an increased cell percentage in G0/G1 phase and a decreased cell percentage in S and G2/M phase in miR‐196b‐5p‐overexpressed WJCMSCs. (C) The cell‐cycle proliferation index was calculated based on flow cytometer results. Student’s t ‐test was used to analyze the statistical significance. All error bars signify standard deviations ( n = 3). * P ≤ 0. 05, ** P ≤ 0. 01. Knockdown of miR‐196b‐5p promotes cell proliferation and accelerates cell‐cycle progression in WJCMSCs To further explore the role of miR‐196b‐5p on WJCMSC proliferation, we transfected miR‐196b‐5p inhibitors and Consh into WJCMSCs. After 3 days of 1 μg·mL −1 puromycin treatment, the knockdown efficiency of miR‐196b‐5p in WJCMSCs was detected by qRT‐PCR (Fig. 4A ). CFSE assays showed that the number of cells with higher proliferating algebra (Fig. 4B ) and the proliferation index (Fig. 4C ) in the miR‐196b‐5p inhibitors group were significantly increased compared with the control group. The results indicated that miR‐196b‐5p inhibitors increased cell growth. Cell‐cycle assays revealed that compared with the control group, miR‐196b‐5p inhibitors significantly attenuated the G0/G1 phase ratio of WJCMSCs but significantly augmented the S and G2/M phase ratio (Fig. 5A, B ). Then the cell‐cycle proliferation index further confirmed that miR‐196b‐5p inhibitors promoted WJCMSC proliferation (Fig. 5C ). Fig. 4 Knockdown of miR‐196b‐5p promoted cell proliferation in WJCMSCs. (A) Real‐time RT‐PCR showed the knockdown efficiency of miR‐196b‐5p in WJCMSCs. (B, C) CFSE assay results showed that the number of cells with higher proliferating algebra and the proliferation index in the miR‐196b‐5p inhibitors group were higher than those in the control group. U6 was used as an internal control for miR‐196b‐5p. Student’s t ‐test was used to analyze the statistical significance. All error bars signify standard deviations ( n = 3). * P ≤ 0. 05, ** P ≤ 0. 01. Fig. 5 Knockdown of miR‐196b‐5p accelerated cell‐cycle progress in WJCMSCs. (A, B) The flow cytometer analysis results showed a decreased cell percentage in G0/G1 phase and an increased cell percentage in S and G2/M phases in miR‐196b‐5p knockdown WJCMSCs. (C) The cell‐cycle proliferation index was calculated based on flow cytometer results. Student’s t ‐test was used to analyze the statistical significance. All error bars signify standard deviations ( n = 3). * P ≤ 0. 05, ** P ≤ 0. 01. miR‐196b‐5p regulates expression of cell‐cycle‐related proteins in WJCMSCs Next, to determine the mechanism of miR‐196b‐5p on cell cycle, we measured the protein expression level of cell‐cycle‐related regulators by western blot and quantitative analysis. Under miR‐196b‐5p overexpression, the protein levels of Cyclin A, Cyclin D, Cyclin E and CDK2 decreased significantly, whereas p15 INK4b level increased significantly in comparison with the control group (Fig. 6A, B ). Furthermore, miR‐196b‐5p knockdown up‐regulated significantly the expression of Cyclin A, Cyclin D, Cyclin E and CDK2 and down‐regulated significantly the expression of p15 INK4b compared with the control group (Fig. 6C, D ). However, there was no significant difference in CDK4 level between miR‐196b‐5p mimics, miR‐196b‐5p inhibitors and the control groups (Fig. 6A–D ). Fig. 6 miR‐196b‐5p regulated the expressions of Cyclin A, Cyclin D, Cyclin E, CDK2 and p15 INK4b in WJCMSCs. (A, B) Western blot and quantitative analysis results showed the expressions of Cyclin A, Cyclin D, Cyclin E, CDK2, CDK4 and p15 INK4b in miR‐196b‐5p‐overexpressed WJCMSCs. (C, D) Western blot and quantitative analysis results showed the expressions of Cyclin A, Cyclin D, Cyclin E, CDK2, CDK4 and p15 INK4b in miR‐196b‐5p knockdown WJCMSCs. GAPDH was used as an internal control for western blot and quantitative analysis. Student's t ‐test was used to analyze the statistical significance. All error bars signify standard deviations ( n = 3). * P ≤ 0. 05, ** P ≤ 0. 01. Sequential window acquisition of all theoretical mass spectra results analysis To further elucidate the mechanism of miR‐196‐5p in WJCMSCs, we used sequential window acquisition of all theoretical mass spectra (SWATH‐MS) to identify differentially expressed proteins induced by miR‐196‐5p. A total of 163 proteins were found to be significantly differentially expressed between miR‐196‐5p inhibitors and the control groups, including 56 up‐regulated and 107 down‐regulated proteins (Table S1 ). PTGS2 and METTL3 were found to be up‐regulated, whereas PON2 was down‐regulated. We speculate that these proteins may be the direct or indirect targets of miR‐196b‐5p, regulating the proliferation of WJCMSCs. Then we detected the mRNA level of PTGS2, METTL3 and PON2 in miR‐196b‐5p inhibitors and the control group. In the miR‐196b‐5p inhibitors group, the mRNA levels of PTGS2 and METTL3 were significantly increased (Fig. 7A, B ), whereas the mRNA level of PON2 was significantly decreased (Fig. 7C ). These results were consistent with the results of SWATH‐MS and confirmed the reliability of the SWATH‐MS data. Fig. 7 The mRNA levels of differentially expressed proteins in WJCMSCs induced by miR‐196‐5p. (A–C) Real‐time RT‐PCR analyzed the mRNA level of PTGS2, METTL3 and PON2 in miR‐196b‐5p inhibitors and the control group. GAPDH was used as an internal control for real‐time RT‐PCR. Student's t ‐test was used to analyze the statistical significance. All error bars signify standard deviations ( n = 3). * P ≤ 0. 05. Discussion WJCMSCs are under investigation in various clinical treatment trials for regenerative medicine due to their characteristics, such as fast proliferation, rapid self‐renewal, good DT stability, low immunogenicity and feasible harvest process. Preclinical studies have shown that the therapeutic mechanisms of WJCMSCs mainly include paracrine, cell substitution and cell contact, among which paracrine is recognized as the primary mechanism [ 30, 31 ]. The basic requirement of this mechanism for clinical application is to ensure sufficient quantity and quality of cells. It has been widely recognized that the therapeutic effect of WJCMSCs will be enhanced when improving the subculture efficiency. However, studies have shown that the phenotype of WJCMSCs may be affected by culture conditions and continuous passage when the cells are expanded in vitro before transplantation [ 32 ]. Therefore, it is necessary to develop methods to enhance the proliferative ability of WJCMSCs to promote their clinic application. In this study, we found that the expression of miR‐196b‐5p was significantly increased in senescent WJCMSCs, suggesting that miR‐196b‐5p may be associated with the decline of proliferation ability in aged cells. Indeed, we found that overexpressed miR‐196b‐5p inhibited WJCMSC proliferation and reduced cells of S and G2/M phases through blocking cells in G0/G1 phase, whereas miR‐196b‐5p knockdown promoted WJCMSC growth by accelerating cell‐cycle progress into S and G2/M phases. Cell proliferation is strictly controlled by cell cycle, which involves a series of complex cascade events [ 33 ]. During G1/S transition, cells are blocked in the G0/G1 phase, which means a prolonged initiation time for DNA synthesis. However, the increase of G2/M phase indicates accelerated cell mitosis and cell proliferation [ 34 ]. Our study suggested that miR‐196b‐5p might play an important role in regulating cell cycle and cell proliferation of WJCMSCs, and it may be a potential therapeutic target for improving subculture efficiency of WJCMSCs. However, the underlying mechanism is still unclear and needs further study. The important mechanism of cell growth is mainly regulated by cell‐cycle regulatory proteins, including cyclins, CDKs and CDK inhibitors. CDK4/6 and CDK2 are activated by Cyclin D binding to CDK4/6 or Cyclin E to CDK2, but are inactive without their homologous cyclin partners [ 35 ]. In cell‐cycle regulation, the key in G1 phase is the binding of Cyclin D and CDK4/6, which drives the start of the cell cycle [ 36 ]. Cyclin E is essential for the control of the cell cycle at the G1/S transition and combines to CDK2 to make the cell cycle enter into S phase from the late G1 phase [ 37 ]. Cyclin A is induced at the G1/S boundary and binds to CDK2 in S phase and participates in the progress of S phase [ 38 ]. p15 INK4B is a member of the CDK inhibitor family, which can delay the progress of the G0/G1 phase through inhibiting the binding of Cyclin D and CDK4/CDK6 [ 39 ]. Our study showed that overexpression of miR‐196b‐5p down‐regulated Cyclin A, Cyclin D, Cyclin E and CDK2 and up‐regulated p15 INK4b, whereas knockdown of miR‐196b‐5p up‐regulated Cyclin A, Cyclin D, Cyclin E and CDK2 and down‐regulated p15 INK4b, which is consistent with the results of Li et al. [ 40 ]. We speculate that miR‐196b‐5p may attenuate the G1/S transition through regulating the combination of Cyclin A and Cyclin E to CDK2, and Cyclin D and p15 INK4b may be involved. In brief, our results showed that miR‐196b‐5p may block G0/G1 phase by down‐regulating Cyclin A, Cyclin D, Cyclin E and CDK2 and up‐regulating p15 INK4b, which needs to be further verified. Many studies show that miRNAs act directly or indirectly on transcripts encoding proteins related to cell proliferation and cell cycle [ 18 ]. To find the target of miR‐196b‐5p and further explain its molecular mechanism on WJCMSC proliferation, we conducted SWATH‐MS analysis in miR‐196‐5p inhibitors and the control groups. The result suggests that several differential proteins are related to cell proliferation or cell cycle, among which the expressions of PTGS2 and METTL3 were up‐regulated, whereas the expression of PON2 was down‐regulated. It was revealed that knockdown of PTGS2 (also known as cyclooxygenase‐2) could increase cells in G0/G1 phase and decrease cells in S phase, thus inhibiting the proliferation and growth of Capan‐2 cells [ 41 ]. METTL3 promoted tumor growth in bladder cancer via modulating pri‐miR221/222 maturation by an m6A‐dependent manner [ 42 ]. PON2 was involved in the regulation of C12‐HSL on the mitochondrial energy production and function of LS174T cells and the inhibition of cell proliferation [ 43 ]. Then we used qRT‐PCR to detect the mRNA levels of PTGS2, METTL3 and PON2 in miR‐196b‐5p inhibitors and the control group. The results showed that the expression levels of the selected mRNAs were consistent with the SWATH‐MS results. We speculate that PTGS2, METTL3 and PON2 might be the direct or indirect targets of miR‐196b‐5p in regulating the proliferation of WJCMSCs, which needs further exploration. Conclusions This study indicated that miR‐196b‐5p suppressed cell growth by blocking the cell cycle of WJCMSCs in G0/G1 phase through down‐regulating Cyclin A, Cyclin D, Cyclin E and CDK2 and up‐regulating p15 INK4b. This study contributed to reveal a novel molecular mechanism on WJCMSC proliferation and laid a foundation for better use of WJCMSCs in future clinical applications. Conflict of interest The authors declare no conflict of interest. Author contributions XH was responsible for collection and assembly of data, data analysis and interpretation, manuscript writing and final approval of the manuscript. HY, HL, YC and CZ contributed to data collection. RS and ZF were responsible for conception and design, manuscript writing and revising, financial support and final approval of the manuscript. All authors have read and approved the final version of the manuscript. Supporting information Fig. S1. The expression of miR‐196b‐5p in WJCMSCs, ADSCs, BMMSCs, SCAPs, DPSCs, and PDLSCs. QRT‐PCR showed that the expression of miR‐196b‐5p in WJCMSCs, ADSCs, and BMMSCs increased significantly compared with SCAPs, DPSCs, and PDLSCs. U6 was used as an internal control for miR‐196b‐5p. One‐way ANOVA was used to analyze statistical significance. All error bars signify standard deviations ( n = 3). ** P ≤ 0. 01. Click here for additional data file. Table S1. SWATH‐MS in miR‐196b‐5p inhibitors WJCMSCs. Click here for additional data file.
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10. 1002/2211-5463. 13119
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FEBS Open Bio
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The role of biotechnology in the transition from plastics to bioplastics: an opportunity to reconnect global growth with sustainability
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Building new value chains, through the valorization of biomass components for the development of innovative bio‐based products (BBPs) aimed at specific market sectors, will accelerate the transition from traditional production technologies to the concept of biorefineries. Recent studies aimed at mapping the most relevant innovations undergoing in the field of BBPs (Fabbri et al. 2019, Final Report of the Task 3 BIOSPRI Tender Study on Support to R&I Policy in the Area of Bio‐based Products and Services, delivered to the European Commission (DG RTD)), clearly showed the dominant position played by the plastics sector, in which new materials and innovative technical solutions based on renewable resources, concretely contribute to the achievement of relevant global sustainability goals. New sustainable solutions for the plastic sector, either bio‐based or bio‐based and biodegradable, have been intensely investigated in recent years. The global bioplastics and biopolymers market size is expected to grow from USD 10. 5 billion in 2020 to USD 27. 9 billion by 2025 (Markets and Markets, 2020, Bioplastics & Biopolymers Market by Type (Non‐Biodegradable/Bio‐Based, Biodegradable), End‐Use Industry (Packaging, Consumer Goods, Automotive & Transportation, Textiles, Agriculture & Horticulture), Region ‐ Global Forecast to 2025), and this high growth is driven primarily by the growth of the global packaging end‐use industry. Such relevant opportunities are the outcomes of intensive scientific and technological research devoted to the development of new materials with selected technical features, which can represent feasible substitutes for the fossil‐based plastic materials currently used in the packaging sectors and other main fields. This article offers a map of the latest developments connected to the plastic sector, achieved through the application of biotechnological routes for the preparation of completely new polymeric structures, or drop‐in substitutes derived from renewable resources, and it describes the specific role played by biotechnology in promoting and making this transition faster.
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Abbreviations BBPs bio‐based products CE circular economy LVHV low‐volume high‐value Introduction Over the last 150 years, the industrial system has been uniquely driven by a linear model, which is based on the production of goods starting from fossil raw materials, their commercialization, their use, and final disposal as waste to be discharged or incinerated [ 1, 2 ]. The traditional plastic sector shows an iconic representation of this linear model mainly in reference to the packaging field, with its large volume production almost exclusively intended for products for single or very short‐term use, followed by a fast transition to the waste state. For many reasons (including technical restrictions for commingled plastics recycling [ 3 ] or lack of infrastructure dedicated to recycling distributed over the territory), for a long time plastic waste was not considered a resource to be valued, but mainly a global problem giving rise to huge negative externalities. In 2014, Valuing Plastic, a report by the UN Environment Programme and the Plastics Disclosure Project, estimated the total natural capital cost of plastics in the consumer goods industry at USD 75 billion, of which USD 40 billion was related to plastic packaging, exceeding the profit pool of the plastic packaging industry [ 4 ]. In the face of such alarming indicators, the need for a fast transition for the plastic sector, shifting toward a sustainable model, is clearly understandable. The main routes toward a new circular economy (CE) for plastics have been already identified, and the Ellen MacArthur Foundation played an eminent role by giving their clear definition, reported in the document The New Plastic Economy published in 2016 [ 5 ]. Three main areas of intervention were identified: (a) creating an effective after‐use plastic economy, by promoting efficient recycling and a new design strategy inspired and driven by reuse and recyclability; (b) reducing the uncontrolled release of plastics into natural systems, by promoting more efficient waste collection and the development of infrastructure dedicated to waste treatment and valorization; (c) decoupling plastics from fossil resources, by promoting plastics derived from renewable resources. Renewably sourced plastics decouple the production of plastics from fossil resources by sourcing the virgin feedstock either from captured greenhouse gases (GHG‐based) or biomass (bio‐based). Bio‐based plastics can be produced from different generations of feedstock [ 6, 7 ]: 1st generation: Biomass from plants that are rich in carbohydrates and that can be used as food or animal feed (e. g. , sugar cane, corn, and wheat). 2nd generation: Biomass from plants that are not suitable for food or animal feed production. They can be either nonfood crops (e. g. , cellulose) or waste materials from first‐generation feedstock (e. g. , waste vegetable oil, bagasse, or corn stover). 3rd generation: Biomass derived from algae, which has a higher growth yield than either first‐ or second‐generation feedstock, and therefore has been allocated its own category. Based on their physical and chemical properties, renewably sourced plastics can be divided into two categories: drop‐ins and completely new materials. Drop‐ins are identical replicates of the currently used traditional plastics, except that they are derived from renewable resources instead of being produced from fossil oil. The most relevant example is that of bio‐polyethylene (bio‐PE), which offers the same properties and features of standard PE, but is polymerized from ethylene monomer obtained by the dehydration of bioethanol, fermented from sugarcanes [ 8 ]. Drop‐ins can be directly introduced into the existing value chains, as they can deliver exactly the same level of technical performance as standard plastics. On the other hand, completely new bio‐based plastics can be derived from renewable resources without having a fossil‐based counterpart. This includes the examples of polylactic acid (PLA) or poly(hydroxyalkanoate)s (PHA), which offer a completely new set of properties but need a tailored approach for their introduction into production chains. Bio‐based plastics indeed offer multiple advantages, ranging from decoupling plastics production from fossil resources, to decreasing carbon dioxide emissions. Bio‐based plastics could also act as a carbon sink throughout their life cycle. For plastics synthesized using carbon from captured GHG, this looks rather obvious [ 9 ]. For bio‐based plastics coming from vegetable source, this happens indirectly: Plants capture carbon dioxide from the atmosphere as they grow and this carbon is then harnessed in the polymer [ 10 ]. Bio‐PE GHG savings is up to 0. 60 kg CO 2 e per kg for corn‐derived PE, and 3. 4 kg CO 2 e per kg for switchgrass‐derived polyethylene compared to GHG emissions of standard fossil‐based PE [ 11 ]. In recent years, two main studies have been funded by the European Commission for the identification of innovative pathways from renewable feedstocks to innovative bio‐based products (BBPs), including biofuels and biochemicals, with biopolymers being a subcategory of the latter. The first study focused on the sugar platform and was completed in 2015 [ 12 ]. Its main objective was to provide a evidence base regarding the production of biofuels and biochemicals from the sugar platform, offering the following outcomes: An assessment of the status of the different pathways, mapping their suitable feedstocks and potential products, and identifying technology opportunities, enablers and barriers to commercialization; An assessment of European developments, and how competitive European industry is likely to be versus other world regions; For a defined set of ten case studies, an analysis of production costs, and comparison of business cases against current technologies in the market; A sustainability assessment using key criteria such as GHG emissions, land use, safety issues, and other environmental and socio‐economic factors; The identification of current research gaps and R&D needs—with a focus on recommending measures that will accelerate the introduction of large‐scale demonstration facilities. This study on the sugar platform selected the 25 most relevant BBPs, out of which the vast majority (16 products) were clearly developed through a biotechnological (biological or intracellular) route. Almost all of these biotechnological BBPs were somehow related to the plastics sector, some being biopolymers (PLA, PHA, bio‐PE) and some others being key monomers for innovative bioplastics (e. g. , iso‐butene (which is the building block for elastomers), and 1, 4‐butanediol (BDO), which is a building block for biopolyesters, bioelastomers, and biopolyurethanes). The latest study, named BIOSPRI, focused on a variegated set of biomass components, either available in large volumes (natural rubber, lignins, plant fibers, vegetable oils, and animal fats) or in low volumes but offering high added‐value (terpenes, polyelectrolytes), in addition to urban waste (organic fraction of municipal solid waste and urban waste water) [ 1 ]. Similarly to the previous study dedicated to the sugar platform, the BIOSPRI study also clearly evidenced the key role played by biotechnology in the development of innovative BBPs starting from the above‐mentioned selection of biomass platforms. The study offered a mapping of more than one hundred emerging innovations derived thereof, and claimed as a main outcome of the study that bio‐based innovative solutions somehow related to the plastics sector certainly occupy the most relevant positions in the ranking of the top‐emerging ones. In fact, the BIOSPRI study down‐selected the 20 most innovative BBPs out of the overall mapping for the development of detailed case studies. Down‐selection of the top‐emerging BBPs was achieved by applying five assessment criteria to the more than 100 products mapped. Assessment was, respectively, based on (a) level of technological readiness (TRL) at least reaching the pilot scale (i. e. , TRL 5); (b) presence of an active marketplace; (c) EU‐based development; (d) degree of innovativeness; and (e) market potential. The final list of the selected top‐emerging 20 BBPs again included many products strictly related to the plastics sector and developed through a biotechnological approach, such as polyamide‐12 (PA12), third‐generation chitosan (also called biotechnological or fungal chitosan), PHA from urban waste, and more. In the following section, a description is offered of some of the most relevant BBPs described by the BIOSPRI study, which represent significant innovation in the plastics sector and are exclusively derived from biotechnological methods. Bio‐based polyamide 12 (PA12), fungal chitosan and PHA obtained from vegetable oils and animal fats, are three different kinds of biopolymers deserving deployment in the next 5–10 years, and are currently under development at the pre‐industrial scale level thanks to the driving role of biotechnology in the transition from plastics to bioplastics. Results from the BIOSPRI tender study on support to R&I policy on bio‐based products Polyamide‐12 (PA12) Introduction Aliphatic polyamides (PA) are industrially synthesized by ring‐opening polymerization of cyclic monomers (lactams) or by step‐growth polycondensation of diacids/diesters with diamines, or ω‐amino acids/esters. PA are engineered polymers, used as high‐performance materials [ 13 ], and they include the PA 6 to 12, and PA4, 6 and PA6, 6, PA6, 10, and PA6, 12 [ 14 ]. As an example, PA6 is synthesized by ring‐opening polymerization of caprolactam and PA6, 6 is synthesized by polycondensation of adipic acid with hexamethylenediamine. Some bio‐based PA have been already developed such as PA11, PA10, 10, and partially bio‐based variations using the same monomers, commercialized under the trade name Rilsan T® by the company Arkema, or similarly by EVONIK and other suppliers. Aliphatic polyamides are semi‐crystalline polymers. The regular spatial alignment of amide groups allows a high number of hydrogen bonding to develop when chains are aligned. The crystalline regions contribute to the hardness, yield strength, chemical resistance, creep resistance, and temperature stability, while the amorphous areas contribute to the impact resistance and ductility. Aliphatic polyamides are versatile plastics for engineering and excellent fiber materials. As per their application, aliphatic polyamides are categorized into two divisions: polyamide fibers and polyamide thermoplastics. Polyamides fibers are mainly used in carpets, apparel, tire reinforcement, and in other industrial applications. They are used in racing car tires and airplane tires owing to their excellent strength, adhesion to rubber, and fatigue resistance in these demanding applications. Molecular weight, in the case of PA6, 6 fibers, is in the range of 12 000–15 000 for apparel fibers and 20 000 for tire yarn is preferred. Polyamides thermoplastics are important engineering plastics because of their toughness over a wide range of temperatures. In addition, they have good resistance to impact and abrasion, organic solvents, and petroleum products. They are used in many automotive applications such as gears and bearings. Reinforced PA are used for exterior body compartments such as fender extensions, decorative louvers, filler plates, head lamp housings, cross‐over panels, and many other applications. In the electrical and electronic field, polyamides are used in making plugs, sockets, switches, and connectors. PA12 has a low concentration of amide moieties compared to other commercially available polyamides. PA12 absorbs very little moisture, has excellent resistance to chemicals (e. g. , hydraulic fluids, oil, fuels), dampens noise and vibration, and is highly processable. Product description PA12 is a long‐chain linear polyamide belonging to the aliphatic polyamide groups [ 15 ] and is characterized by 12 carbon atoms between two nitrogen atoms of the two amide groups (Fig. 1 ). It presents a melting point of 180 °C and medium to high viscosity grade [ 16 ]. Fig. 1 Polyamide 12. The PA12 currently commercialized is exclusively based on fossil sources. It is prepared on an industrial scale via ring‐opening polymerization of lauryl lactam converted after multiple step synthesis starting from butadiene [ 17 ]. It can also be prepared from ω‐amino‐dodecanoic acid [ 18 ]. UBE in Japan is using 12‐amino‐dodecanoic acid in its manufacturing process for PA12. For bio‐based PA12, the biomass‐derived monomers ω‐amino lauric acid (ALA) or 12‐amino‐dodecanoic acid (saturated amino acid) are used and they represent the fully renewable alternative to the petroleum‐based lauryl lactam (LL) derived from butadiene. Conventional fossil‐based PA12 and fully bio‐based PA12 offer identical properties apart from the input use and production process; therefore, bio PA12 is to be considered as a drop‐in solution. At present, fatty acids derived from renewable resources have been proven to be suitable monomers and building blocks for the production of bioplastics, including PA12 [ 19, 20 ]. As an example, ω‐amino undecanoic acid, which is synthesized from 10‐undecenoic acid derived from castor oil, is currently used for the industrial production of bio‐based PA11, commercially available from Arkema [ 17 ]. The German chemical company Evonik aims to produce bio‐based ALA for the synthesis of PA12; a pilot‐scale plant entered into operation in 2013, in Slovakia. The biotechnological manufacturing process starts from palm kernel oil (PKO) that the company was already using as a platform for the production of other chemicals. PKO contains a large amount of C12 lauric acid and over 50% of it is saturated. The fruit palm tree containing PKO is mainly produced in Asia. Arkema is investigating an alternative process, using the C11 undecenoic acid methyl ester produced from castor oil, and cross‐metathesis with acrylonitrile. Production processes for bio‐based PA12 Three different routes have been proposed for the preparation of fully bio‐based PA12: fermentation; cross metathesis from C10 or C11 unsaturated esters; from bio‐based butadiene, through the synthesis of the monomer dodecanelactame. The fermentation method is the one completely related to the biotechnological approach, while the second and third methods take advantage of the application of traditional chemical treatments for the synthesis of bio‐based PA12. As an example, in the bio‐based butadiene route, the major issue is that bio‐butadiene becomes available, but after that there is no need to change the standard production process which is currently applied to petrol‐based butadiene for the synthesis of PA12. Here, butadiene is treated with a Ziegler‐type catalyst system to yield the cyclic trimer, cyclododeca‐1, 5, 9‐triene. This may then be hydrogenated to give cyclododecane, which is then subjected to direct air oxidation to give a mixture of cyclododecanol and cyclododecanone. Treatment of the mixture with hydroxylamine yields the corresponding oxime, which on treatment with sulfuric acid rearranges to form the monomer dodecanelactame [ 21 ]. In the cross‐metathesis approach, the C12 amino‐ester can be produced through olefin metathesis. The successful application of cross metathesis to synthesize PA12 precursors from oleic acid has been reported by the literature. For example, the method reported by Rupilius & Ahmad [ 22 ] provides a short and simple route for production of PA12 from PKO. In addition, Abel and coauthors [ 23 ] demonstrated that cross metathesis could be applied directly to crude fatty acid methyl ester extracts from algal biomass. However, while metathesis can be regarded as bio‐based process since the raw material used is renewable, the underlying chemical reactions could create more undesired by‐products and hazardous waste than alternative organic reactions. A completely renovated approach is offered by biotechnology through the fermentation route. The process starts with the vegetable PKO, which generates a new alternative PA12 precursor which is ALA‐equivalent, namely ω‐amino‐dodecanoic acid methyl ester (ADAME). ADAME can be considered as the renewable alternative to petroleum‐based LL. ADAME can be subsequently polymerized to an identical PA12 biopolymer with respect to the one obtained from ALA. This new C12 monomer was obtained from lauric acid methyl ester via Escherichia coli bacterial fermentation [ 15 ] via a three‐step cascade, in which bio‐based dodecanoic acid methyl ester (DAME) hydroxylation and alcohol oxidation both catalyzed by the alkane monooxygenase AlkBGT from Pseudomonas putida GPo1 were followed by terminal amination by means of Chromobacterium violaceum ω‐transaminase CV2025. Thereby, dodecanedioic acid monomethyl ester (DDAME) was formed as a major by‐product. Biomass yield and feedstock availability Palm oil is a plant that produces two different types of oils, the palm oil from mesocarp and PKO from the seed of kernel with different chemical and physical properties. The fatty acid (lauric saturated C12) content in PKO is about 48. 2% [ 24 ]. One of the advantages compared to other crop oil is that palm oil yields about from 4. 09 to 0. 5 tonne per ha of PKO [ 25 ], which sound competitive compared to the 0. 37 tonne per ha of soybeans oil, 0. 5 tonne per ha of sunflower oil and 0. 75 tonne per ha of rapeseed oil. In terms of availability, nowadays Indonesia and Malaysia account for 85% of global palm oil production. Indonesia is expected to continue to be the leading producers with expansion in land for palm oil; on the other hand, in Malaysia expansion is expected to slow in view of limited land availability. Nevertheless, more and smaller producer countries have emerged within the palm oil market, including Cameroon, Colombia, Costa Rica, Cote d‘Ivoire, Ghana, Guatemala, Honduras, Nigeria, Papua New Guinea, and Thailand. Commercial relevance and future perspectives of bio‐based PA12 Important drivers leading the current and future trends in PA markets are increasing oil prices, that will make conventional polymers more and more expensive, and the expected increase in automotive sales in emerging economies [ 26 ]. The PA market at the global level is projected to reach USD 30. 76 Billion by 2021, at a CAGR of 4. 1% from 2016 to 2021 [ 27 ] and USD 32. 7 billion by 2025 [ 28 ]. In general, the demand for the fastest growing type of sector of the polyamide market by 2021 is projected to be the bio‐based and specialty polyamide sector [ 27 ]. While the global PA market was valued at USD 25. 14 Billion in 2016 [ 26 ], the bio‐based PA market was valued at USD 29 454 million in 2012 and USD 110. 5 million in 2016 [ 27 ]. The largest market is represented by the Asia‐Pacific area and is followed by Europe and then North America where we find the fastest growing sectors of automobile, packaging, electronics, and consumer goods, including retail. The Chinese, Japanese, Indian, Brazil, and Russian markets are also driving the growth of this market. The growth of the transportation industry due to rising disposable income levels, mainly in China and India, is likely to drive regional markets. This picture is expected to be preserved in the future and will still be dominated by the Asia‐Pacific region, which shows the highest growth rate, although new market alliances will possibly stabilize the overall business. Korea, Taiwan, and Japan are expected to create important opportunities for specialty PA manufacturers, due to increasing demand in PA 12‐based applications. On the other hand, Central and South America will show only moderate growth in the future [ 26 ]. Value proposition and sustainability It is ambiguous whether PA based on renewable raw materials are in fact more sustainable than other types when considering factors such as landscape consumption [ 29 ], use of agrochemicals and fertilizers, water, transport, working conditions, and the impact on food security [ 30 ]. In terms of sustainability, consumers are demanding verification of the way palm oil has been produced. The Round Table of Palm Oil Supply Chain Certification systems addresses this issue by reducing the risk of nonsustainable palm oil use by consumers while further driving the mainstream trade of sustainable palm oil, according to the company [ 31 ] (Green Chemical Blog, 2015). In order to allow future sustainable development of the palm oil industry, it is crucial to consider the increment of palm tree yield per hectare and the value addition of the oil. At the same time, it is important to reduce production costs. In this respect, special attention should be given to the improvement of oil palm varieties with higher yields and good oil quality, and which are compact in architecture, better adapted to climate change and exhibit higher tolerance to diseases [ 32, 33 ]. Alternative vegetable oils to support the process have been tested, such as coconut oil [ 34 ], but their current technology readiness level is still lower with respect to the PKO route. Switching from fossil‐based to bio‐based PA12, sustainably produced, would support the economies of countries growing palm oil, in addition to reducing the carbon footprint of the wide variety of final products that are produced using PA12 as input. Third‐generation Chitosan Introduction Chitin is a very abundant biopolymer in nature [ 35 ], being the second most abundant polysaccharide after cellulose. It can be obtained from many sources such as exoskeletons of crustaceans, clams (endo‐skeleton of cephalopods), cell walls of fungi and microalgae, insect exoskeletons, yeast, and the spines of diatoms [ 36, 37, 38 ]. At present, crustacean shells represent the most important source of chitin for industrial production, due to the availability of marine waste material from the seafood processing industry [ 39 ], and well‐established chemical processes of deacetylation. In fact, chitin use on a large scale is limited due to its water and (most) solvent insolubility. Deacetylation to chitosan represents the most relevant modification of chitin to a water‐soluble derivative, with much wider potential applications. Different degrees of deacetylation are possible, and therefore several grades of chitosan can be derived from chitin, offering a set of different chemical, physical, and mechanical properties [ 40, 41 ]. The interest in chitosans as advanced functional biopolymers comes from their special features which make them suitable for specialty applications and development of high added‐value polymer solutions. Biomedical applications are of primary importance because of their biocompatibility, biodegradability, and nontoxicity [ 42 ], added to their intrinsic antimicrobial activity [ 43 ] and low immunogenicity, which clearly point to an immense potential for future development. These biopolymers can be easily processed into gels, sponges, membranes, beads, and scaffold forms [ 44 ]. Chitosan‐based nanomaterials, including nanofibers, nanoparticles, and nanocomposite scaffolds for tissue engineering [ 45 ], wound dressing, drug delivery [ 46 ] and cancer diagnosis represent huge potential for future development in advanced materials for medical use. Chitosans also find relevant uses in other sectors, such as bioremediation [ 47, 48 ] and food preservation [ 49 ]. Product description Chitosan, classified by the European Commission as a ‘basic substance’ in 2014, is a renewable polysaccharide. Specifically, it is an aminoglucopyran. It is a linear cationic heteropolymer (Fig. 2 ). In its dissolved form, only achieved at ph < 5, 7, thanks to the positive charge distributed along the biopolymer chain, chitosan gives rise to versatile uses based on its chelating, antimicrobial, gelling and film‐forming properties [ 50 ]. Fig. 2 Chitin and chitosan. With reference to the source biomass and the chemical characteristics, three kinds of chitosan can be defined: Chitosans derived from animal chitin of marine origin, deacetylated through chemical (soda) treatment is referred to as first‐generation chitosan. These chitosans benefit from a low‐cost, easy‐processing route, but suffer from scarce control over the degree of deacetylation and the broad distribution of the biopolymer molecular weights [ 51 ]. If chitosans come from the enzymatic processing of animal chitin, then they are called second‐generation chitosans [ 52 ]. These biopolymers are better defined than 1st generation chitosan in terms of their degree of polymerization and acetylation, and are therefore more suitable for the development of reliable products with good batch‐to‐batch consistency for industrial applications. A mix of molecules showing different chain length, varied degree of acetylation, and different acetylation patterns characterize both first‐ and second‐generation chitosans. Finally, third‐generation chitosan, also called biotechnological or fungal chitosan, is a high‐purity non‐animal‐sourced biopolymer, derived exclusively from renewable, non‐GMO sources such as algae and fungi, without any synthetic manipulations [ 53 ]. Source biomass is mainly represented by waste from agro‐alimentary industries. Regarding the chemical characteristics, the molecular weight and degree of deacetylation of fungal chitosan can be controlled by varying the fermentation conditions, while it is rather randomly obtained with chemical treatment of animal‐sourced chitin. Specifically, compared with chitosan from crustaceans, characterized by a high molecular weight (1. 5 × 10 6 Da), fungal chitosan produces a medium‐low molecular weight (1–12 × 10 4 Da) with molecular weight homogeneity [ 54 ], and nonrandom patterns of acetylation, and therefore a well‐defined chemical structure, known cellular modes of action, and better defined biological activities [ 55 ]. It is expected that these chitosans will create important opportunities for the market in the future for advanced applications, mainly in the medical field, due to its high purity and a well‐defined chemical composition. Furthermore, as regards raw material supply, due to dis‐continuous supply and seasonal variations of marine sources, the use of fungi could represent an advantageous alternative. The overall differences between first‐ and second‐generation chitosans compared to fungal chitosan are reported in Table 1. Table 1 Generations and properties of chitosans. Chitosan generation Characteristics Source biomass Applications 1st/2nd generation Animal derived (crustaceans) High molecular weight (1. 5 × 10 6 Da) Random patterns of acetylation Low purity Seasonal variation Marine wastes Nonadvanced applications (i. e. , agriculture, wastewater treatment, and in general all the applications that not require high volume and/or high purity) 3rd generation Nonanimal derived (derived from fungi and algae) Medium‐low molecular weight (1‐12 × 10 4 Da) Low molecular weight homogeneity Nonrandom patterns of acetylation Shrimp‐protein and heavy metal free High purity Nonseasonal variation Wastes from agro‐alimentary industries and from Biotech industries Advanced applications (i. e. , medicine, cosmetics, health care) and when a high and constant quality is strictly required John Wiley & Sons, Ltd The properties of fungal chitosans compared with those of a commercial chitosan derived from crab shells are shown in Table 2. The degree of deacetylation is an important parameter affecting the physicochemical properties of chitosan; higher degrees of deacetylation induce higher concentration of positive charges along the polymer chain. This characteristic makes the biopolymer more suitable for food applications as a coagulating or chelating agent, as clarifying agent or as antimicrobial agent. Chitosan with lower molecular weight was reported to reduce the tensile strength and elongation of the chitosan membrane but to increase its permeability. Chitosan from fungi and algae shows more homogeneity in terms of chemical composition but at present is less available and more expensive compared to marine chitosan. To date, small quantities of these molecules are made available on the market; few attempts have been made to obtain chitosan with reliable and predictable biological functionalities in amounts that are appropriate to the market request. It is currently not feasible for high‐volume and low‐cost applications. Table 2 Properties of fungal chitosans [ 57 ]. Chitosan from Degree of deacetylation (%) Molecular weight (Da) Viscosity (cP) Crab shell 97. 9 ± 0. 9 9. 4 × 10 5 372. 7 Aspergillus niger 90. 0 ± 2. 1 1. 4 × 10 5 6. 2 Rhizopus oryzae 87. 9 ± 2. 1 6. 9 × 10 4 3. 5 Lentinus edodes 86. 5 ± 2. 2 1. 9 × 10 5 5. 8 Pleurotus sajo‐caju 83. 8 ± 0. 1 1. 1 × 10 5 5. 6 Zygosaccharomyces rouxii 85. 1 ± 1. 1 2. 7 × 10 4 3. 3 Candida albicans 83. 8 ± 0. 8 1. 1 × 10 5 3. 1 John Wiley & Sons, Ltd Production process Chitosans can be produced by several species of fungi. The most common method for the production of chitosan from mycelia of fungi involves fungi growth by solid‐state fermentation on carbon‐rich substrates, mainly agro‐waste such as sweet potatoes or others. The chitosan is then extracted by enzymatic treatment [ 56 ]. Species of filamentous fungi, such as Aspergillus niger, Rhizopus oryzae, Lentinus edodes, and Pleurotus sajo‐caju, but also yeast strains, such as Zygosaccharomyces rouxii TISTR5058 and Candida albicans TISTR5239, were investigated for their ability to produce chitosan in complex media [ 57 ]. A remarkable characteristic of the composition of zygomycetes is the high concentration of chitosan in their cell walls. In R. oryzae, it amounts to 42% of the total cell mass [ 58 ], and deriving chitosan from these fungi might thus offer an alternative to chitosan production involving deacetylation of chitin from marine crustacean shells. Deacetylation of crustacean shells is normally conducted through harsh alkaline hydrolysis at high concentration and temperature, which entails a long processing time, environmental pollution, and inconsistent physicochemical properties of the produced chitosan. Production of chitosan from zygomycetes under milder controlled conditions could yield a readily available and much more consistent product. Fungal chitosan was produced at 10–140 mg·g −1 cell dry weight and had a degree of deacetylation of 84–90% and a molecular weight of 2. 7 × 10 4 –1. 9 × 10 5 Da with a viscosity of 3. 1–6. 2 centipoises (cP). R. oryzae TISTR3189 was found to be the producer of the highest amounts of chitosan, suitable for commercial deployment of this production strategy for chitosans. Biomass yield and feedstock availability As compared to marine sources, chitosan production using fungi is minor, and the fungi showing the highest portion of chitosan appears to be the mucoraceous fungi [ 54, 57 ]. Chitosan content strongly varies and depends on the specific species of fungi, as reported in Table 3. According to Ghormade et al. [ 54 ], there exist three major sources of fungi that can be exploited at a commercial level: (a) waste fungal biomass from biotech industries where thousands of tons of waste fungal are produced per year, (b) fungi containing high amounts of chitosan grown by fungal fermentation, and (c) value addition to existing mycotech products. Table 3 Chitosan content in different fungi species [ 57 ]. Fungal species Chitosan content (%) Aspergillus niger 11 Rhizopus oryzae 38 Lentinus edodes 3. 3 Zygosaccharomyces rouxii 3. 6 Candida albicans 4. 4 John Wiley & Sons, Ltd In the first group of sources, it can be mentioned that more than 80 000 tons of waste A. niger biomass are generated per year from the production of citric acid. Additionally, it can be estimated that in the brewing and baking industries where S. cerevisiae and S. carlbergenis genuses are traditionally used, biomass availability is more than 130 tonne yeast lees/year. For penicillin production, 1 tonne of penicillin implies around 8–10 tonne/year of waste that can be used for chitosan production. In the second group of sources involving fermentation of fungus, major contents can be found in Gongronella butleri, Mucor rouxii, and Absidia coerulea, amongst zygomycetous fungi. More specifically, Benjaminiella poitrasii, Cunnighamella blackesleeanus, Gongrenella butleri, Mortierella isabelina, Rhizopus delemar, and Rhizopus stolonifera, amongst others, show a chitosan content ranging between 6. 7% and 10. 4% production. Commercial relevance and future perspectives The current amount of third‐generation chitosan produced is currently very low so that market trends are extremely hard to forecast. However, both the chitin and chitosan markets, at the global level, have been growing massively in recent years, due to the expansion of their application domain, which is even larger if co‐products are considered. Applications include biopharmaceutical, cosmetics, biotechnological and biomedical, agriculture, and both food and nonfood industries. Some features justify positive expectations for the future, such as low toxicity, excellent biocompatibility, versatile biological activity, and complete biodegradability. In addition, the third‐generation chitosan derived from fungi and algae presents many advantages compared with 1st and 2nd generation, primarily less allergenic tendency. Other factors supporting the market development of chitosan are as follows: The enhancement of the water treatment sector resulting from high demand for removal of metals and chemicals from wastewater, including pesticides, surfactants, phenol, and polychlorinated biphenyls. The demand for the compound in food and beverage application in Europe, which is expected to grow at significant rate due to rising demand for the wine‐processing industry. The development expected in the cosmetics industry due to the rising demand for bio‐derived personal care products. The cosmetic industry was the largest application sector of chitosan in North America in 2015 [ 59 ]. On the other hand, the growth of value‐added chitosan‐based products is limited by the availability of a sustainable supply chain: the process chemistry for bulk chitosan manufacturing is currently not very environmentally friendly. Green technologies for chitosan modification are facing the challenge of economic viability which hinders its current development. According to existing analysis, we can consider a CAGR of 14% after 2020 to be reasonable, declining over the following years due to a contained effect of market saturation, assuming a $2. 0 billion volume in 2016 and 155 000 metric tons of volume reached in 2022 [ 60 ]. Price estimates for chitosans and chitin are not clear and notably vary amongst the studies analyzed. However, the source of chitosan derivation strongly affects final price; also, chitosans derived from fungi and algae are notably more expensive than animal‐sourced ones. Chitosan from fungal sources is less available compared to marine chitosan sources, which implies higher production costs and therefore very high sale price. A current share of 0. 2% of nonanimal chitin is envisaged for the total chitin market, growing slightly over time as soon as production costs are able to decrease. Overall price dynamics could mainly depend on two opposite forces. On the one hand, the increasing demand for high‐quality chitosan products will put pressure on their prices. On the other hand, expected improvements in technologies and chemical processes will reduce marginal production costs and limit the price increase. A modest annual price increase is therefore expected in time (1%). Value proposition and sustainability The production of chitosan from fungi can be made more economically advantageous by the exploitation of a varied set of possible co‐products and by‐products. As reported by Kuk [ 61 ], both chitin and chitosan can be used as starting materials for the production of high added‐value chemicals, such as mono‐, di‐, and oligosaccharides that are derived with additional hydrolysis processes [ 62 ]. Examples are chitobiose, the N‐acetylglucosamine monomer, and glucosamine salts via enzymatically catalyzed hydrolysis reactions of the obtained chitin/chitosan. An important implementation of chitosanase is the preparation of chitosan oligosaccharide from chitosan [ 63 ], which is able to prevent the accumulation of fat in internal organs, plays a remarkable role in liver function and in stimulating the immunological system, and may find pharmaceutical application in formulations for the control of cholesterol accumulation. For mushroom industry waste, where stalks, which represent ~ 25–33% of the weight of fresh mushrooms, are normally used as low economic value animal feed, a trade‐off exists. Indeed, this biowaste material could be utilized to produce vitamin D and chitosan as co‐products of the industry of high‐quality mushrooms [ 64 ]. Therefore, it is clear that one of the most relevant aspects related to the production of chitosan from fungi is that large amounts of waste can be valued, and seasonal availability of marine sources can be overcome through the rapid, easy, and cheap cultivation of fungal species. Chitosans, recently used in wastewater treatment to remove heavy metals from polluted sediment [ 65 ], also respond to the increasing need of addressing environmental and pollution problems relative to rivers and lakes showing sediment contamination. On the same path, chitosans with a high deacetyletic degree can support lower food waste as they show potential for extending the shelf life of refrigerated fish fillets, due to their inhibitory properties, thus contributing to mitigating the global problem of food loss [ 49 ]. PHA from renewable oils and fats Introduction Polyhydroxyalkanoates (PHA) are intracellular biopolyesters accumulated as a carbon supply or for energy storage by various microorganisms. The homopolymer poly(3‐hydroxybutyrate) (PHB) and its copolymers containing valerate units (PHBV) or hexanoate units (PHBH) represent the most diffused types of PHA. PHA macromolecular chains can be synthesized from numerous carbon‐rich substrates by the biosynthetic action of selected prokaryotic microorganisms. Sugar‐rich substrates, such as sugar beets molasses or sugar cane bagasse, have been widely studied for the industrial production of PHA [ 12 ]. All industrially relevant production worldwide is based on these kinds of food sources for microorganisms ( http://www. bio‐on. it. http://www. biomer. de. http://www. kaneka. co. jp ). The function of PHA granules inside bacterial cells is that of carbon and energy storage, which can be degraded when necessary by several microorganisms producing depolymerizing enzymes. This makes PHA degradable to water and carbon dioxide (or methane, under anaerobic conditions) in all biologically active environments such as soil, open waters (i. e. , rivers and lakes, seas and oceans), compost, and sewage [ 66, 67 ]. A variety of waste streams different from sugars have been tested in earlier decades to produce PHA in economically sound ways. Examples of used carbon sources are different kinds of agro‐industrial food waste [ 68, 69, 70 ], organic municipal waste [ 71 ], and activated sludge coming from waste water treatment [ 72 ]. It is noteworthy that the molecular structure of PHA chains accumulated inside bacterial cells, and their consequent physical and mechanical properties as plastic materials, does not directly vary based upon the kind of biomass feedstock used as a carbon source for bacterial fermentation. It is only dependent on the bacterial stream used for the fermentation process, its operative conditions, and the number of carbon atoms available in the chemical moieties of the feedstock, which are used as building blocks by bacteria. In the search for alternative carbon sources for PHA bacterial fermentation, which may be more affordable with respect to sugars, attention has focused on vegetable oils [ 73, 74 ] and animal fat waste, such as residues from slaughterhouses [ 75 ], and glycerol as the main residue from biodiesel production [ 76, 77 ], but also carbon‐rich wastewater from different industrial activities [ 78 ]. Product description The chemical and physical properties of PHA biopolymers are very similar to those characterizing petroleum‐derived commodity polymers, such as polypropylene and polyethylene. This makes PHA a very good substitute for conventional plastics [ 79, 80 ], and their uses and applications can vary accordingly to their specific molecular structure and formulation with additives. Depending on the chain length in the PHA subunit (monomer), the hydrophobicity and a number of other properties including the glass transition temperature, the melting point, and level of crystalline color, can vary. Normally, short‐length PHA are hard crystalline materials; medium‐chain length PHA are ductile plastics and have a much lower melting point and glass transition temperature; and PHA with longer pendant groups are rather elastomeric. Notwithstanding the bacterial strain and carbon‐rich substrates selected for their synthesis, PHA always show spontaneous and complete biodegradability in several environmental conditions, ranging from composting industrial and home facilities to open waters, that is, rivers, seas, oceans, and wet lands. This makes PHA particularly attractive because this feature of complete biodegradability is coupled with their 100% renewable and biotechnological origin [ 81 ]. For the homopolymer 3‐hydroxybutyrate) (PHB), which represents the most‐studied example of biodegradable polyesters belonging to the family of PHA, its use is mainly as a substitute for rigid commodity plastics, such as HDPE, PP, and ABS. It is highly crystalline, and therefore, it is optically opaque, highly stiff, and resistant to tensile stress. PHB suffers from being quite brittle, and its processing on standard industrial equipment can be achieved upon optimization of processing parameters, mainly to accommodate its rheological and thermal properties. Several PHB‐based grades, fully or partially biodegradable, can be developed by polymer modification or melt blending with several other commercial plastics and bioplastics. The resulting level of biodegradability depends on the composition of the blend. At present, PHB still has higher production costs with respect to its fossil‐based commodity plastics (HDPE and PP mainly), but it holds advantages mainly related to the full biodegradability, renewable origin, and possibility to be obtained by the valorization of waste. The most important features of PHA are its biodegradable capacity under aerobic and anaerobic conditions. Degradation often occurs upon exposure to soil, compost, or marine sediment. However, the biodegradation rate depends on factors such as exposed surface area, moisture, temperature, pH, and molecular weight. Biocompatibility is perhaps the most important property for applications in the medical field [ 82, 83 ]. In addition, biopolyesters are inert, water‐insoluble, not affected by moisture and indefinitely stable in air. In pharmaceutical applications, PHA synthase is the key enzyme for PHA biosynthesis and the same can be exploited for the development of many chiral derivatives [ 84 ]. Moreover, the food and beverage industries take advantage of PHA satisfactory barrier properties. Water‐resistant surfaces, moisture vapor barriers, and UV‐resistant layers based on PHA can be used in the packaging industry [ 85 ]. Last but not least, PHA may also find relevant applications in engineering sectors, such as sensors [ 86, 87 ] or construction [ 88 ]. Production process Microbial PHAs are bacterial polyesters produced by enzymatic reactions inside microbial cells from acetyl‐co‐enzymeA (acetyl‐CoAs) by PHA synthase, which is a substrate‐specific enzyme present in the cell cytosol. PHA derived from animal fat and vegetable oils can be synthesized by different microorganisms [ 89 ]. Microorganisms able to transform triglycerides into PHA polymer chains, accumulated in the form of granules inside the bacterial cell cytosol, should be characterized by two main features: being able to produce a lipase enzyme to hydrolyze the triglycerides to liberate the long‐chain fatty acids, and to then be transformed into PHA chains by β‐oxidation. In fact, if the carbon substrate used are oils or lipids and are utilized though the fatty acid pathway, then the oxidation of enoyl‐CoA to (R)‐3‐hydroxyacyl‐CoA takes place due to catalyses initiated by the (R)‐specific enoyl‐CoA hydratase (PhaJ). (R)‐3‐hydroxyacyl‐CoA acts as a substrate for the PHA synthase (PhaC) enzyme and is the immediate precursor of PHA biosynthesis [ 89, 90 ]. The use of vegetable oils in the production of PHA is rather straightforward, and almost no kind of pretreatments are required [ 89, 91 ]. Any remaining food traces will eventually contribute to the overall carbon supply for the microorganisms. This represents a remarkable difference with respect to biodiesel production, in which waste oils need to be pretreated using esterification and filtration before being used. The most efficient bacterial strains identified for the intracellular accumulation of PHA using oil and fat carbon substrates are Cupriavidus necator H16, C. necator ATCC 17699, R. eutropha Re2133, and C. nectator Re2058/pCB113 [ 89 ], all of which show an accumulated PHA content higher than ~ 70% with respect to the dry cell weight. Maximum values of rather 90% (w/w) were observed for C. necator, which is the top performing strain when vegetable oils are used as carbon substrates to feed the fermentation process [ 89 ]. Solid animal fats, characterized by triacylglycerol‐containing long‐chain fatty acyl groups, require triacylglycerol‐utilizing bacteria, which can secrete lipases. Lipases will release fatty acids in the fermentation media, which are then transformed into PHA polymer chains through β‐oxidation [ 89, 92 ]. In general, microorganisms find it more difficult to utilize animal fats than liquid vegetable oils, mainly due to the physical solid state of animal fats at room temperature and under fermentation conditions. Thus, pretreatment of animal fats is usually necessary to bring them into a more accessible physical state, that is, liquid; this can be made, for instance, by pre‐emulsification and heating. In 1996, Cromwick [ 93 ] found that P. resinovorans could produce PHA polymers on unhydrolyzed tallow, with a PHA content of 15% of the cell dry weight. However, the ester of the fatty acids from tallow with methanol can be fermented using P. citronellolis, with a productivity for medium‐chain length mcl‐ PHA of 0. 036–0. 050 g/(L*h) and PHA contents of 20. 1–26. 6% (wt) [ 94 ]. This demonstrates that PHA can be also produced from saturated biodiesel fractions (SFAE) stemming from waste animal fats. This waste biomass source is available in notable amounts in Europe and in other parts of the world; in Europe alone, animal lipids from slaughtering and the animal‐processing industry amount to more than 500 000 t per year. Separating the SFAE fractions from biodiesel enhances its performance as a fuel, while the recovered waste fractions can be advantageously valued as feedstock for the biotechnological production of PHA [ 94 ]. Biomass yield and feedstock availability There are many potential advantages in using plant oils instead of sugars as carbon sources for PHA production. The superior source biomass yield is of course of primary importance for a possible industrial up‐scaling of the process. While the maximum yield reported for PHA production starting from glucose is ~ 0. 3–0. 4 g of PHA per gram of sugar, that value increases at 0. 6–0. 8 g of PHA when plant oils are used for feeding the fermentation process. This is due to the higher content of carbon present in oils per weight compared to sugars [ 95 ]. When comparing the efficacy of WCO with fresh vegetable oil as carbon sources, similar values are found for the growth rate of microorganisms and PHA accumulation. Yields can vary from 0. 94 g PHA /g oil when WCO is used, to the slightly lower value of 0. 87 g PHA /g oil for fresh vegetable oils. In any case, yields are very high and PHA accumulation stays in the range 85–75% w/w of the dry biomass content, respectively. Therefore, the WCO is slightly more efficient in terms of yield when compared to pure vegetable oils [ 89, 96 ]. At the global level, the feedstock availability for production of PHA from oily biomass is extremely relevant. It was estimated that a 10 million gallon per year biodiesel plant would have the potential of producing 20. 9 ton PHB [ 77 ]. Considerable amounts of oily waste rich in lipid composition are generated from various developed industrial processes, including animal fats, oil mills, dairy foods, and processing of food material. The Energy Information Administration (EIA) estimates that ~ 11 billion liters of waste vegetable oils are generated annually in the USA. In the European Union, the aggregate amount of waste vegetable oils produced per year is accounted as ~ 1 billion liters. In 2010, the United States produced around 2. 7 billion tons of tallow and grease. Thus, large amounts of animal fats with a high fatty acid content are being generated and can be used as a potential feedstock for producing PHA [ 89 ]. Commercial relevance and future perspectives Amongst the innovative plastic materials currently under development, PHA are considered as the most promising bioplastics, deserving deployment, and increased utilization over the next few years [ 97 ]. The PHA market is expected to reach an estimated volume of 45. 49 tons by 2027, based on a growth rate of 7. 60% for the period between 2020 and 2027 [ 98 ]. Increased use of PHA in various application sectors, including high added‐value sectors such as the biomedical and the cosmetics industries, is expected to drive the PHA market in the next years. Commercial PHA are sold as commodity bioplastics for the large volume production of disposable goods (packaging, and food service items such as plastic straws, cutlery, trays, and bottles) and for agriculture (mulching films, nutrients carriers), and for high added‐value markets such as the biomedical and the cosmetics industry, 3D printing, and chemical additives. The vast availability of renewable and cost‐effective raw materials such as bagasse, zein, casein, plant starch, and many more is encouraging the growth of the PHA market. Government regulations and policies against single use plastic is a major factor for the growth of the PHA market. Moreover, with the emergence of new raw materials and growth in the Asia‐Pacific region will create further opportunities for the PHA market. The spontaneous biodegradability of PHA under environmental conditions also represents a main driver for the market growth of PHA. Value proposition and sustainability PHA produced by bacterial fermentation using vegetable oils or animal fat waste biomass perfectly adhere to the principles of CE and circular bioeconomy. In addition to contributing to the reduced depletion of fossil resources, PHA are fully derived from renewable resources, can be produced from different types of biowaste through low impact biotechnological routes, and offer spontaneous biodegradation in an open environment. On the social side, all of these features together induce positive thinking about plastics in modern society, supporting acceptance and understanding of the opportunities by the end‐users. With regard to evaluation of environmental impact, life cycle assessment methodologies have been widely applied to study PHA. The emission of GHGs during production of PHA showed a reduction of up to 200% and use of fossil energy was reduced by 95%, contributing to a reduction in the amount of waste to be managed [ 89, 99 ]. Conclusions BIOSPRI, a recent tender study on support to R&I policy in the field of BBPs, commissioned by the European Commission DG RTD, selected the 20 most promising BBPs currently under development, and in the framework of this selection, many products are directly related to the plastics sector. The findings are evidence for the strategic role of biotechnology in driving and boosting the transition from fossil‐based plastics to bioplastics obtained from renewable resources. Three relevant examples were discussed in this review: (a) PA12, which represents an innovative bioplastic mainly for engineering applications; (b) fungal chitosan, which offers relevant features for advanced applications in the biomedical field and for decontamination of waters and soils; and (c) PHA derived from vegetable oils and animal fats, which represents an example of bioplastic fulfilling the complete set of requirements of the circular bioeconomy. Biotechnology is the key enabling technology for the development of these innovative biomaterials, and clearly supports their transition to full technological maturity and commercial accessibility. Valorization of waste and forward‐looking management of critical raw materials are main drivers for the further development of the bioplastics here discussed, and biotechnological methods offer increasing opportunities for the whole plastic sector. Conflict of interest The authors declare no conflict of interest. Author contributions PF, FF, LB, FC and DV contributed to conception, design and development of the study. PF, FF, LB, FC, DV, MDE, and DM equally contributed to the organization and writing of the manuscript, revised, read, and approved the submitted version. Data accessibility Data and information used to elaborate this review paper are accessible through the Publication Office of the European Union, at the following link: https://op. europa. eu/en/publication‐detail/‐/publication/15135e98‐81c2‐11e9‐9f05‐01aa75ed71a1/language‐en/format‐PDF/source‐search.
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10. 1002/2211-5463. 13336
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FEBS Open Bio
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Dental pulp tissue engineering is a promising alternative treatment for pulpitis and periapical periodontitis, and dental pulp stem cells (DPSCs) are considered to be the gold standard for dental seed cell research. Periapical lesions harbor mesenchymal stem cells with the capacity for self‐renewal and multilineage differentiation. However, it remains unknown whether these periapical lesion‐derived stem cells (PLDSCs) are suitable for dental pulp tissue engineering. To investigate this possibility, PLDSCs and DPSCs were isolated using the tissue outgrowth method and cultured under identical conditions. We then performed in vitro experiments to investigate their biological characteristics. Our results indicate that PLDSCs proliferate actively in vitro and exhibit similar morphology, immunophenotype and multilineage differentiation ability as DPSCs. Simultaneously, PLDSCs exhibit stronger migrative ability and express more vascular endothelial growth factor and glial cell line‐derived neurotrophic factor than DPSCs, and PLDSC‐derived conditioned medium was more effective in tube formation assay. The mRNA expression levels of immunomodulatory genes HLA‐G, IDO and ICAM‐1 were also higher in PLDSCs. However, regarding osteo/odontogenic differentiation, PLDSCs showed weaker alkaline phosphatase staining and lower calcified nodule formation compared to DPSCs, as well as lower expression of ALP, RUNX2 and DSPP, as confirmed by a quantitative RT‐PCR. The osteo/odontogenic protein expression levels of DSPP, RUNX2, DMP1 and SP7 were also higher in DPSCs. The present study demonstrates that PLDSCs demonstrate potential use as seed cells for dental pulp regeneration, especially for achieving enhanced neurovascularization.
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Abbreviations ALP alkaline phosphatase CCK‐8 Cell Counting Kit‐8 CFU‐F colony‐forming units‐fibroblasts DMEM Dulbecco’s modified Eagle’s medium DMP1 dentin matrix protein 1 DPSCs dental pulp stem cells DSPP dentin sialophosphoprotein GDNF glial cell line‐derived neurotrophic factor HGF hepatocyte growth factor HLA‐G human leukocyte antigen G HUVECs human umbilical vein endothelial cells ICAM‐1 intercellular adhesive molecule‐1 IDO indoleamine 2, 3‐dioxygenase MSCs mesenchymal stem cells PLDSCs periapical lesion‐derived stem cells qRT‐PCR quantitative RT‐PCR RUNX2 RUNX family transcription factor 2 SP7 Sp7 transcription factor VEGF vascular endothelial growth factor Recent advances in the field of dental pulp regenerative engineering have resulted in bench‐to‐bedside treatments [ 1, 2 ]. Mesenchymal stem cells (MSCs), especially oral‐derived MSCs, are promising tools in dental pulp tissue engineering as a result of their readily accessibility. Among them, dental pulp stem cells (DPSCs) were first characterized by Gronthos et al. [ 3 ] and are still the most frequently studied type of cells, being considered as the gold standard for dental seed cells research. Subsequently, a variety of dental MSCs, including stem cells from human exfoliated deciduous teeth [ 4 ], stem cells from apical papilla [ 5 ], dental follicle stem cells [ 6 ] and periodontal ligament stem cells [ 7 ], have been cultivated from the oral cavity. Nevertheless, it is still imperative to find alternative sources of seed cells as a result of the vulnerability of MSCs in vitro, difficulty of acquiring a sufficient number of well‐functioning cells and ethical problems pertaining to the cultivation of these cells. Following the discovery of the immunomodulatory ability of MSCs, it has been reported that inflammation not only has little impact on the survival of MSCs, but also can active de novo MSCs as well as attract MSCs from other sites to contribute to the healing process [ 8 ]. Furthermore, it was also reported that MSCs from inflammatory conditions varied in characters such as proliferation, migration, differentiation potential, growth factors secretion and immunomodulatory ability [ 9, 10, 11 ]. Periapical periodontitis, a consequence of the inflammatory response to the pulp necrosis and bacterial invasion, was recently found to be a potential source of cells exhibiting MSC properties [ 12, 13 ]. Therefore, the pathological tissues that were once thrown away have now been recognized as a rich source of MSCs. However, these periapical lesion‐derived stem cells (PLDSCs) have not been investigated in the context of dental pulp regeneration. For use in dental pulp regeneration, the pro‐angiogenesis ability plays a pivotal role because the dentin–pulp complex is nourished only by few blood vessels entering through the narrow root canal apex [ 14 ]. The pro‐angiogenesis ability of dental MSCs has already been confirmed in many studies [ 14, 15 ]. Therefore, as oral‐derived stem cells, PLDSCs might act as promising pro‐angiogenesis seed cells for dental pulp engineering. Furthermore, osteo/odontogenic differentiation ability and neurogenesis induced ability are also required for seed cells to achieved complete dentin–pulp complex regeneration. However, whether the characteristics of PLDSCs are suitable for dental pulp engineering remain unknown. To clarify this issue, we cultured PLDSCs and DPSCs under identical condition and compared multiple aspects of them. Materials and methods Cell isolation and culture Patients ( n = 3) aged 18–25 years and with chronic periapical periodontitis were informed and had signed written informed consent in the study. All subsequent protocols were approved by the Ethical Committee of Shanghai Ninth People’s Hospital. The study methodologies conformed to the standards set by the Declaration of Helsinki. Before microscopic endodontic surgery, complete blood counts, coagulation function and the presence of infectious diseases were examined. In addition, all patients underwent periodontal treatment 1 week prior to surgery. The surgeon and surgical protocols remained the same for all enrolled subjects. During the surgery, the removed inflamed tissue was placed in sterile phosphate‐buffered saline (PBS) and then transferred to the laboratory within 10 min. The whole tissue was then washed several times with sterile PBS containing 100 U·mL −1 penicillin and 100 mg·mL −1 streptomycin (Invitrogen, Waltham, MA, USA) until the mixture became transparent. Then, the tissue outgrowth method was applied. Briefly, the tissue was placed in a 10‐mm dish for mechanical disruption. The samples were then minced into small pieces (approximately 0. 1 mm in diameter) using sterile ophthalmic scissors. The minced tissues were transferred to 10‐mm culture dishes, followed by the application of sterile cover glasses dipped in sterile petroleum jelly to the four corners for tissue stability. High‐glucose Dulbecco’s modified Eagle’s medium (DMEM (Gibco, Walthem, MA, USA) supplemented with 10% FBS (Gibco), 100 U·mL −1 penicillin and 100 mg·mL −1 streptomycin (Invitrogen) was used to culture the cells. The culture dishes were incubated at 37 °C and 5% CO 2, and the medium was changed every 3 days. The cells were passaged when they reached about 100% confluence. To isolate DPSCs, premolars or wisdom teeth without cavities for orthodontic reasons were extracted from healthy patients aged 18–25 years ( n = 3), followed by gentle separation of the dental pulp. The culturing steps were similar to those for PLDSCs. To acquire an objective baseline, we used passage 3 of both cell types to perform the following experiments. Flow cytometry assay DPSCs and PLDSCs were identified via the cell surface antigens CD45, CD31, CD34, CD44, CD29, CD73, CD90 and CD105 using a flow cytometry assay. Cell were fixed in 4% phosphate‐buffered paraformaldehyde and incubated with CD45‐FITC (Invitrogen), CD31‐PE (Becton Dickinson, Franklin Lakes, NJ, USA), CD34‐PE (BD Bioscience, San Jose, CA, USA), CD44‐FITC (Invitrogen), CD29‐PE (Invitrogen), CD73‐PE (eBioscience, San Diego, CA, USA), CD90‐FITC (eBioscience) and CD105‐PE (eBioscience) antibodies for 45 min. Then, the cells were washed twice with PBS and analyzed via a FACSCalibur flow cytometer (Becton Dickinson). Immunocytochemistry To detect the expression of vascular endothelial growth factor (VEGF) and glial cell line‐derived neurotrophic factor (GDNF), a sterile round cover glass for cell growth was placed into a 24‐well plate. DPSCs and PLDSCs were seeded at a density of 2. 5 × 10 4 cells·mL −1 onto the cover glass. After culturing the cells for 3 days with growth medium (DMEM supplemented with 10% FBS, 100 U·mL −1 penicillin and 100 mg mL −1 streptomycin), they were fixed with 4% paraformaldehyde for 20 min. Cells were permeabilized using 1% Triton X‐100 and 3% BSA was used to block non‐specific binding. Then, the samples were incubated with primary antibodies against VEGF (dilution 1 : 100; MA5‐13182; Thermo Fisher, Waltham, MA, USA) and GDNF (dilution 1 : 200; #711074; Thermo Fisher), Alexa Fluor 555‐labeled donkey anti‐mouse IgG secondary antibodies (dilution 1 : 500; A0460; Beyotime, Shanghai, China) and Alexa Fluor 555‐labeled donkey anti‐rabbit IgG secondary antibodies (dilution 1 : 500; A0453; Beyotime) was then applied respectively. The cytoskeleton was stained with fluorescein isothiocyanate–phalloidin (dilution 1 : 200; #40735ES75; Yeasen, Shanghai, China) for 1 h at room temperature, followed by the addition of 4′, 6‐diamidino‐2‐phenylindole (C1005; Beyotime) to stain the nuclei. Images were acquired using a confocal laser scanning microscope (Leica, Wetzlar, Germany). In vitro osteo/odontogenic differentiation Both cell types were detached with 0. 25% trypsin–EDTA, resuspended in growth medium and seeded at a density of 10 × 10 4 cells·well −1 into a 24‐well plate. After reaching 90% confluence, the medium was replaced with osteo/odontogenic medium containing 10% FBS, 0. 2 m m l ‐ascorbic acid‐2‐phosphate, 100 n m dexamethasone, 10 m m β‐glycerophosphate, 100 U·mL −1 penicillin and 100 mg·mL −1 streptomycin. The medium was changed every 3 days. After culturing the cells for 7, 14, 21 and 28 days, they were washed three times with PBS and fixed with 4% paraformaldehyde for 20 min. A 1% Alizarin Red S (A5533; Sigma‐Aldrich, St Lois, MO, USA) solution, dissolved in isopropanol and filtered through a 0. 22‐μm filter, was added at a volume of 1 mL·well −1 for 15 min to detect the presence of calcified nodules. Then, the samples were washed with PBS until the water became transparent, followed by observation under an inverted phase‐contrast microscope and simultaneous imaging. In vitro adipogenic differentiation To assess the adipogenic differentiation capacity of PLDSCs and DPSCs, cells were incubated in adipogenic differentiation medium kit (HUXDP‐90031; Cyagen, Santa Clara, CA, USA) supplemented with 10% FBS, 100 U·mL −1 penicillin, 10 μg·mL −1 streptomycin, 12 m m l ‐glutamine, 10 μ m insulin, 200 μ m indomethacin, 1 μ m dexamethasone and 0. 5 m m 3‐isobutyl‐1‐methylxanthine. After culturing the cells for 3 weeks, they were fixed and stained with oil red O stain for 30 min. Images were captured under an inverted phase‐contrast microscope. In vitro chondrogenic differentiation To investigate the chondrogenic differentiation capacity of PLDSCs and DPSCs, a chondrogenic differentiation medium kit (HUXMA‐90041; Cyagen) was applied in accordance with the manufacturer’s instructions. In brief, 4 × 10 4 cells were transferred to a centrifuge tube and resuspended with chondrogenic differentiation medium and cultured for 24 h. When the cells aggregated, the medium was changed gently with fresh chondrogenic differentiation medium. After 28 days of induction, the cell aggregates were fixed and embedded in paraffin, and cross‐sections were stained with Alcian blue. Images were then captured using an optical microscope. Cell proliferation assay A Cell Counting Kit‐8 (CCK‐8) (Dojindo, Shanghai, China) assay was used to assess proliferation. DPSCs and PLDSCs were digested with 0. 25% trypsin–EDTA, resuspended in culture medium and then seeded into a 96‐well plate at 5000 cells·well −1 with five replicate wells each. All cells were cultured in DMEM for 1, 3, 5, 7, 9 and 11 days. At each time point, CCK‐8 solution (dilution 1 : 100 with serum‐free DMEM) was added into the wells and incubated for 1 h at 37 °C. After incubation, attenuance was determined at 450 nm. To compare the proliferation ability of the cells, relative attenuance values were used for normalization and proliferation curves were plotted. Colony‐forming units‐fibroblasts (CFU‐F) assay The cells were digested and seeded in a six‐well plate at a density of 1000 cells·well −1 in triplicate. After 14 days of culture, the cells were fixed with 4% paraformaldehyde and then stained with 1% crystal violet (C0121; Beyotime). Aggregates of more than 50 cells were scored as one colony. The number of colonies of each cell type was then counted using an inverted phase‐contrast microscope. The colony‐forming rate was determined as: colonies/1000 × 100%. Transwell assay Both cell types were resuspended with serum‐free DMEM and seeded at a density of 2. 5 × 10 4 cells·mL −1 in the upper chamber of a 0. 22‐μm transwell plate ( n = 3). The bottom chamber was filled with DMEM supplemented with 10% FBS. After incubating the plate for 24 h, a cotton swab was used to wipe off the cells remaining on the upper chamber. Then, the cells were fixed and stained with 1% crystal violet for 15 min. The stained cells were counted under an inverted phase‐contrast microscope, with simultaneous imaging. Alkaline phosphatase (ALP) staining ALP is an important enzyme involved in the early stages of osteogenesis and odontogenesis. The BCIP/NBT Alkaline Phosphatase Color Development Kit (C3206; Beyotime) was used to detect ALP expression in accordance with the manufacturer’s instructions. DPSCs and PLDSCs were seeded into 24‐well plates at a density of 2. 5 × 10 4 cells·mL −1 in triplicate. After reaching more than 90% confluence, the culture medium of the experimental group was exchanged with osteo/odontogenic medium containing 10% FBS, 0. 2 m m l ‐ascorbic acid‐2‐phosphate, 100 n m dexamethasone, 10 m m β‐glycerophosphate, 100 U·mL −1 penicillin and 100 mg·mL −1 streptomycin. After 3 and 7 days of culture, the cells were fixed and stained with ALP working solution at 37 °C for 20 min in dark condition. The reaction was stopped using PBS, and images were captured using an inverted phase‐contrast microscope. Quantitative RT‐PCR (qRT‐PCR) DPSCs and PLDSCs were seeded in a six‐well plate at a density of 10 × 10 4 cells·mL −1. After reaching more than 90% confluence, the culture medium in experimental groups was replaced with osteo/odontogenic medium (containing 10% FBS, 0. 2 m m l ‐ascorbic acid‐2‐phosphate, 100 n m dexamethasone, 10 m m β‐glycerophosphate, 100 U·mL −1 penicillin and 100 mg·mL −1 streptomycin) and cultured for 3, 7 days, while the control groups replaced with growth medium. At each time point, the total RNA of DPSCs and PLDSCs was extracted using the Trizol reagent after reaching more than 90% confluence. All RNA was then reverse‐transcribed to cDNA using a reverse transcription kit (RR036A; Takara, Beijing, China). RT‐PCR was performed using a quantitative real‐time PCR system (Roche, Basel, Switzerland) with the settings: 95 °C for 30 s for one cycle, followed by 40 cycles of 95 °C for 10 s and 60 °C for 30 s. The comparative ΔCt method was used to calculate the relative expression levels of ALP, RUNX2, DSPP, VEGF, GDNF, IDO‐1, HLA‐G, HGF and ICAM‐1 genes. The ACTB gene was used for normalization. All primers were commercially synthesized (Sangon, Shanghai, China) and are listed in Table 1. Table 1 Primers used in the qRT‐PCR. Gene name Accession number Product length (bp) Primer (5′‐ to 3′) Primer sequence VEGF NM_001025366. 3 186 Forward TGACAGGGAAGAGGAGGAGA Reverse CGTCTGACCTGGGGTAGAGA RUNX2 NM_001015051. 4 127 Forward CACTGGCGCTGCAACAAGA Reverse CATTCCGGAGCTCAGCAGAATAA GDNF NM_000514. 4 182 Forward CGAACTCTTGCCCCTGACCT Reverse ACAGCCACGACATCCCATAAC IDO NM_002164. 6 124 Forward CTGTTCCTTACTGCCAACT Reverse TCCATGTTCTCATAAGTCAGG HLA‐G NM_001363567. 2 144 Forward CTGAGATGGAAGCAGTCTT Reverse GCTCCCTCCTTTTCAATCT HGF NM_001010931. 3 184 Forward AGACCAATGTGCTAATAGATGTA Reverse GCAGTTTCTAATGTAGTCTTTGT ICAM‐1 NM_000201. 3 69 Forward AGCTTCGTGTCCTGTATGGC Reverse TTTCTGGCCACGTCCAGTTT ACTB NM_001101. 5 186 Forward TGGCACCCAGCACAATGAA Reverse CTAAGTCATAGTCCGCCTAGAAGCA DSPP NM_014208. 3 157 Forward AGTGACAGCCAGAGCAAG Reverse CCTATCCCATTACCAAACT ALP NM_001127501. 4 137 Forward CCTTGTAGCCAGGCCCATTG Reverse GGACCATTCCCACGTCTTCAC John Wiley & Sons, Ltd Western blot analysis DPSCs and PLDSCs were culturing in growth medium or osteo/odontogenic medium (containing 10% FBS, 0. 2 m m l ‐ascorbic acid‐2‐phosphate, 100 n m dexamethasone, 10 m m β‐glycerophosphate, 100 U·mL −1 penicillin and 100 mg·mL −1 streptomycin) for 7, 14 and 21 days. At the time point, total protein was harvested using RIPA lysis buffer (WB, 0102; Biotechwell, Shanghai, China) supplied with PMSF; the cell lysates were then centrifuged at 15 984 g for 15 min at 4 °C, and the supernatant was collected for protein analysis. The protein concentration was determined using a BCA Protein Assay (P0012S; Beyotime). Equal amounts of cell lysates were subjected to 10% SDS/PAGE and transferred to poly(vinylidene difluoride) (FFP24; Beyotime) membranes. The membranes were blocked with 5% non‐fat milk and incubated with primary antibodies against RUNX family transcription factor 2 (RUNX2) (dilution 1 : 500; #12556S; Cell Signaling Technology, Danvers, MA, USA), dentin sialophosphoprotein (DSPP) (dilution 1 : 500; sc‐73632, Santa Cruz Biotechnology, Santa Cruz, CA, USA), dentin matrix protein 1 (DMP1) (dilution 1 : 500; ab103203; Abcam, Cambridge, UK), Sp7 transcription factor (SP7) (dilution 1 : 1000; PA5‐115697; Thermo Fisher), VEGF (dilution 1 : 100; MA5‐13182; Thermo Fisher), GDNF (dilution 1 : 200; 711074; Thermo Fisher) and GAPDH (WB0197; Biotechwell) and were then incubated with HRP‐conjugated secondary antibodies (dilution 1 : 5000; Biotechwell). Finally, the protein bands were visualized using ECL Plus reagents (Biotechwell) and analyzed using imagej ( https://imagej. nih. gov/ij ). Preparation of conditioned medium DPSCs and PLDSCs were seeded at a density of 5000–6000 cells·cm −2 into culture dishes. After reaching 70%–80% confluence, the cells were washed three times with PBS and then incubated with serum‐free DMEM containing 1% penicillin–streptomycin for 48 h. The supernatant was collected and centrifuged at 4 °C at 3000 g for 3 min, followed by centrifugation at 1500 g for 5 min, filtration using 0. 22‐μm filters and storage at −80 °C until further use [ 16 ]. Tube formation assay To assess the pro‐angiogenesis ability of the cells, an in vitro tube formation assay was conducted. Human umbilical vein endothelial cells (HUVECs) obtained from the cell bank at the Chinese Academy of Science were seeded at a density of 20 × 10 4 cells·mL −1 into 96‐well plates pre‐coated with 60 μL of growth factor reduced Matrigel (356230; BD Bioscience) in triplicate. Then, 100 μL of serum‐free DMEM (control group) and DPSC‐ or PLDSC‐conditioned medium was added. After incubation for 6 h, images of five random fields were acquired under an inverted phase‐contrast microscope and analyzed using image j. Statistical analysis All experiments were performed at least three times on three different patients. When normal data distribution was confirmed, data were analyzed using Student’s t ‐test or one‐way analysis of variance with prism, version 7. 0 (GraphPad Software Inc. , San Diego, CA, USA). P < 0. 05 was considered statistically significant. Results Identification of mesenchymal stem cell properties of DPSCs and PLDSCs Both DPSCs and PLDSCs adhered to plastic and exhibited a homogeneous spindle‐like shape from passage 1 (Fig. 1A ). No obvious differences in morphology were observed. Oil red O and Alcian blue staining confirmed the adipogenic and chondrogenic differentiation of DPSCs and PLDSCs (Fig. 1B ). Furthermore, the flow cytometry results showed that both DPSCs and PLDSCs expressed high levels of CD29 (> 95%), CD44 (> 95%), CD73 (> 95%), CD90 (> 95%) and CD105 (83. 2% and 79%, respectively), and expressed low levels of hematopoietic marker CD31 (< 2%), CD34 (< 2%) and CD45 (< 2%) (Fig. 1C ). Fig. 1 Identification of mesenchymal stem cell properties of DPSCs and PLDSCs. (A) Images obtained from phase‐contrast microscopy showing PLDSCs with a spindle‐like shape and adhesive properties similar to those of DPSCs. Scale bar = 500 μm. (B) Oil red O and Alcian blue staining showing that both DPSCs and PLDSCs can achieve adipogenic and chondrogenic differentiation. Scale bar = 100 μm (upper); 50 μm (below). (C) Immunophenotype of different cell surface markers. Comparison of proliferative and migrative ability of PLDSCs and DPSCs After 3 days of culture, the relative attenuance values at 450 nm rapidly increased for DPSCs and PLDSCs, and both growth curves had a sigmoid shape, indicating that PLDSCs could be effectively expanded in vitro. In line with previous study [ 17 ], DPSCs had stronger proliferation ability, and both cell types entered the platform stage after 9 days of culture (Fig. 2A ). The CFU‐F assay revealed that DPSCs and PLDSCs formed similar colonies after 14 days of culture and there were no significant differences between the two cell types (Fig. 2B ). Considering that the stem cells suitable for tissue engineering should have migration ability, we next performed the transwell assay. After 24 h of culture, we found an increased number of migrated cells in the PLDSC group compared to the DPSC group (Fig. 2C ). Fig. 2 Comparison of the proliferative and migrative abilities of cells. (A) Growth curve showing the relative attenuance value from the CCK‐8 assays at different time points ( n = 3, Student’s t ‐test). (B) Results of the colony‐forming unit assay; there was no statistical difference between DPSCs and PLDSCs on day 14 ( n = 3, Student’s t ‐test). (C) Results of the transwell assay; after 24 h of culture, PLDSCs exhibited stronger migration ability ( n = 3, Student’s t ‐test). Scale bar = 100 μm. ** P < 0. 01. Data are shown as the mean ± SD. Comparison of osteo/odontogenic abilities of PLDSCs and DPSCs ALP is an enzyme that plays an important role in the early stages of osteogenesis and odontogenesis. After culturing in osteo/odontogenic medium for 3 and 7 days, ALP expression was enhanced in both DPSCs and PLDSCs. However, ALP expression dramatically increased in a time‐dependent manner in DPSCs, whereas only a slight increase was observed in PLDSCs (Fig. 3A ). Alizarin red staining revealed that calcified nodules appeared in DPSCs after 7 days of induction, whereas there was no calcified deposition until day 14 in PLDSCs. Both DPSCs and PLDSCs could form calcified nodules by 4 weeks after induction, although there were fewer calcified nodules in PLDSCs (Fig. 3B ). The qRT‐PCR results showed that mRNA expression levels of ALP, RUNX2 and DSPP were significantly higher in DPSCs at day 7 (Fig. 3C ). Western blot analysis revealed that the protein levels of DSPP, RUNX2, DMP1 and SP7 were also higher in DPSCs after 7, 14 and 21 days of osteo/odontogenic induction (Fig. 3D ). Fig. 3 Comparison of the osteo/odontogenic capacities of the cells. (A) ALP staining after osteo/odontogenic induction for 3 and 7 days. Scale bar = 500 μm. (B) Alizarin red staining showing the presence of calcified nodules after 7, 14, 21 and 28 days of osteo/odontogenic induction. Scale bar = 500 μm. (C) Gene expression patterns of ALP, RUNX2 and DSPP after osteogenic induction for 3 and 7 days; the expression level was normalized to that of ACTB ( n = 3, Student’s t ‐test). * P < 0. 05, ** P < 0. 01. (D) Western blotting detection (upper) and imagej analysis (below) of DSPP, DMP1, RUNX2 and SP7 after osteogenic induction for 7, 14 and 21 days. Data are shown as the mean ± SD. Pro‐angiogenesis ability of PLDSCs and DPSCs We next examined the pro‐angiogenic ability of the cells via qRT‐PCR analysis and found that the mRNA expression levels of VEGF were higher in PLDSCs than in DPSCs (Fig. 4A ). Western blot results also showed that protein levels of VEGF were higher in PLDSCs (Fig. 4B ). The results of immunofluorescence staining for VEGF were consistent with those of qRT‐PCR and western blotting; the fluorescence intensity for VEGF was higher in PLDSCs than in DPSCs (Fig. 4C ). A tube formation assay was then performed to evaluate the pro‐angiogenic ability of different conditioned media on HUVECs. After 6 h, there was little tube formation in the control group. However, tubular structures were formed when HUVECs were incubated with DPSC‐ and PLDSC‐conditioned medium (Fig. 4D ). Furthermore, analysis of various tube formation parameters, including the number of junctions, nodes, segments and meshes; total segment length; and total meshes area, as analyzed using imagej, indicated that PLDSCs‐conditioned medium was more effective for initiating angiogenesis (Fig. 4E ). Fig. 4 Comparison of pro‐angiogenesis properties of DPSCs and PLDSCs. (A) qRT‐PCR was used to compare the expression levels of VEGF ; the expression levels were normalized to those of ACTB ( n = 3, Student’s t ‐test). ** P < 0. 01. (B) Western blotting detection of VEGF expression. (C) Immunofluorescence staining showing the protein expression of VEGF in the cytoplasm of DPSCs and PLDSCs. Scale bar = 50 μm. (D) Representative images for each group via inverted phase‐contrast microscopy. Scale bar = 100 μm. (E) Analysis of tube formation parameters (number of junctions, nodes, segments and meshes; total segment length; total mesh area) using imagej ( n = 5, one‐way analysis of variance). * P < 0. 05 with serum‐free, & P < 0. 05 with DPSCs‐CM. DPSCs‐CM, dental pulp stem cells‐conditioned medium; PLDSCs‐CM, periapical lesion‐derived stem cells‐conditioned medium. Data are shown as the mean ± SD. Neurotrophic ability of PLDSCs and DPSCs To evaluate neurotrophic ability, we examined the expression of GDNF, which is expressed in dental MSCs. The qRT‐PCR revealed the higher expression of GDNF in PLDSCs than in DPSCs (Fig. 5A ). Western blot analysis showed that protein levels of GDNF were higher in PLDSCs (Fig. 5B ). In addition, immunofluorescence staining showed a stronger fluorescence intensity for GDNF in PLDSCs than in DPSCs (Fig. 5C ). Fig. 5 Comparison of neurotrophic ability of DPSCs and PLDSCs. (A) qRT‐PCR was used to compare the expression level of GDNF between PLDSCs and DPSCs; the expression level was normalized to that of ACTB ( n = 3, Student’s t ‐test). (B) Western blotting results of GDNF expression. (C) Immunofluorescence staining for GDNF expression in DPSCs and PLDSCs. Scale bar = 50 μm. ** P < 0. 01. Data are shown as the mean ± SD. Comparison of the immunomodulatory ability of the cells We assessed the expression levels of immunomodulatory genes IDO, HLA‐G, HGF and ICAM‐1 to compare the immunomodulatory abilities of the two cell types. PLDSCs expressed higher levels of HLA‐G, IDO and ICAM‐1 than DPSCs. However, no significant difference was observed for expression of HGF (Fig. 6 ). Fig. 6 Comparison of the immunomodulatory abilities of the cells. Expression of immunomodulatory genes HLA‐G, IDO, HGF and ICAM‐1 were evaluated via qRT‐PCR; ACTB was used for normalization ( n = 3, Student’s t ‐test). ** P < 0. 01. Data are shown as the mean ± SD. Discussion Derived from the mesoderm, dental MSCs are investigated intensely for their strong proliferation and adhesion abilities, as well as multilineage differentiation capacity [ 18 ]. However, well functioned dental MSCs are still difficult to mass produce to meet clinical demands and the properties vary under different circumstances. As a result, it is imperative to find more alternative seed cells for tissue engineering. Recently, periapical lesions were found to harbor cells exhibiting MSC‐like properties, suggesting that the use of these pathological tissues should be re‐evaluated [ 12, 17, 19, 20, 21 ]. However, the MSCs isolated from periapical lesions had not been investigated from the viewpoint of dental pulp regeneration. Therefore, we aimed to more closely examine the properties of PLDSCs by comparing them with DPSCs, the most widely used seed cells for dental pulp tissue regeneration to date [ 22 ]. Using a tissue outgrowth method, the present study obtained spindle‐like and plastic‐adhesive cells from periapical lesions. Under different conditioned media, these types of cell could achieve osteogenic, adipogenic and chondrogenic differentiation. In addition, flow cytometry results showed that PLDSCs were positive for CD29, CD105, CD44, CD73 and CD90 and negative for hematopoietic markers CD31, CD34 and CD45, which was similar to that for DPSCs. However, it was also reported that the different culture methods and passage numbers will alter the properties and marker profile of MSCs [ 23, 24 ]. Therefore, to acquire an objective baseline, we used passage 3 for both cell types when conducting our experiments. For the purposes of stem cell therapy and tissue regeneration, which require a tremendous number of cells, in vitro cell proliferation ability is of great importance [ 25 ]. It is well known that dental tissue‐derived MSCs, particularly DPSCs, have strong proliferation ability [ 25 ]. In the present study, we found similar results indicating that DPSCs had stronger proliferative ability [ 17 ], with the proliferation rate of DPSCs being 2. 07 ± 0. 04‐fold compared to that of PLDSCs on day 11. Given that the CCK‐8 assay is an indirect method for evaluating proliferation, we then performed a CFU‐F assay. The results of the CFU‐F assay showed that PLDSCs had a self‐renewal ability similar to that of DPSCs. Taken together, PLDSCs could be cultured and actively proliferate in vitro. Furthermore, the migration ability of cells should be taken into consideration because injured tissues secrete chemokines to attract stem cells from other sites and initiate tissue repair. In the present study, we showed that PLDSCs were more active and had higher migration ability than DPSCs. This finding may be explained by the distinct environment of PLDSCs, where many pro‐inflammatory cytokines are present. For example, the pro‐inflammatory chemokine RANTES/CCL5 has been reported to increase the migration ability of periodontal ligament stem cells isolated from inflamed periodontal ligaments [ 26 ]. For the regeneration of the dentin–pulp complex, seed cells must be capable of osteo/odontogenic differentiation. It was reported that different isolation methods, culture environment, commercial osteo/odontogenic induced medium and inflammatory stages will also lead to different osteo/odontogenesis results [ 23, 27, 28 ]. In the present study, we isolated MSCs by the outgrowth method and found that PLDSCs showed weaker osteo/odontogenic differentiation capacities than DPSCs, which is consistent with the results for alveolar bone resorption symptoms in patients with periapical periodontitis. Similarly, inflammation was reported to compromise the osteo/odontogenic capacity of MSCs. Lee et al. [ 29 ] reported that DPSCs from inflamed dental pulp may have decreased osteo/odontogenic capacity compatred to DPSCs from normal dental pulp. In vitro studies using Pg lipopolysaccharide or tumor necrosis factor‐α to simulate the inflammatory environment also showed the decreased osteo/odontogenic ability of dental MSCs [ 30, 31 ]. By contrast, a comparative study using an enzyme digestion method to isolate MSCs showed that PLDSCs had stronger osteogenic differentiation ability than DPSCs and dental follicle stem cells [ 17 ]. Other studies applying Escherichia coli lipopolysaccharide or a mixture of pro‐inflammatory cytokines to simulate inflammatory conditions in vitro have demonstrated increased ALP activity and calcification ability in bone marrow stem cells, adipose stem cells and DPSCs [ 11, 32 ]. The controversy regarding inflammation and the osteo/odontogenic capacity of MSCs may be explained by the complexity of the inflammatory niche, which contains various cytokines and immune cells. Therefore, to achieve better osteo/odontogenesis of PLDSCs, further studies are needed to investigate the potential regulatory cytokines or immune cells in the different stages of periapical lesions. It is widely accepted that oxygen can only diffuse approximately 200 μm through tissues [ 33 ]; therefore, sufficient vessel networks are vital for the survival of engineered regenerative tissues; this is particularly true for dental pulp, which only has a small opening (< 1 mm) for blood vessels to enter [ 14 ]. The incorporation of growth factors or the co‐culture of endothelial cells with pro‐angiogenic cells are the two major approaches used to solve this problem [ 22 ]. Among the various grow factors, VEGF can be secreted by dental MSCs in a paracrine way [ 15 ] and has a strong impact on blood vessel initiation, being effective for enhancing vessel density in the dental pulp [ 34 ]. Yet, no research has focused on the pro‐angiogenesis potential of PLDSCs. In the present study, the secretion of VEGF by DPSCs and PLDSCs was confirmed via qRT‐PCR, western blotting and immunofluorescence; we observed that VEGF expression was higher in PLDSCs than in DPSCs. It has been reported that different stress conditions will give rise to different secretions and concentrations of pro‐angiogenesis growth factors [ 35 ]. Of these conditions, hypoxia is considered to be a common driving force for the dramatic increase in VEGF expression observed in injured and inflamed dental pulp [ 36 ]. In response to hypoxia, dental pulp cells increase the expression of hypoxia‐inducible factor 1 and then mediate the increased transcription of pro‐angiogenesis factors, including VEGF, platelet‐derived growth factor AB and angiopoietin [ 14 ]. A recent study sequencing RNA from 10 periapical periodontitis tissues also revealed that the hypoxia‐inducible factor 1 pathway was activated [ 37 ]. Therefore, it is reasonable to assume that the PLDSCs in the present study expressed higher levels of VEGF and that PLDSC‐conditioned medium induced increased tubular structure formation in HUVECs because the PLDSCs resided in a hypoxic inflammatory condition before surgery. This enhanced pro‐angiogenic ability is promising with respect to addressing insufficient angiogenesis in dental pulp engineering. Neurogenesis is essential for functional pulp regeneration. Nerves in the pulp chamber help adjust the masticatory force when chewing and nourish blood vessels inside the pulp by secreting multiple growth factors. DPSCs are able to differentiate into neural lineage and secret neurotrophic factors, being promising in neural repair and regenerative in nerve diseases [ 38 ]. PLDSCs were also capable of differentiating into neuronal cells [ 13 ]. In the present study, we chose GDNF to evaluate its neurotrophic function. GDNF is an important neurotrophin involved in the survival of neurons, differentiation of neuroblasts and neuritogenesis [ 39 ]. Furthermore, it was reported that GDNF was able to exert anti‐pro‐inflammation function in renal interstitial fibrosis and inflammatory bowel disease [ 40, 41 ]. GDNF upregulation was found in lipopolysaccharide‐induced nigral inflammation [ 42 ]. In the present study, we showed similar results indicating that PLDSCs expressed higher GDNF levels than DPSCs, suggesting the potential use of PLDSCs in regenerating dental pulp nerves, or even in other nerve diseases [ 43, 44 ]. Although autologous stem cell‐based tissue engineering is the most ideal approach for tissue regeneration, it is not suitable for all patients because the self‐renewal and differentiation abilities of stem cells depend on the health status of the host. Therefore, an allogenic stem cell‐based approach is still an alternative, and an immunomodulatory effect should also be considered when applying seed cells. MSCs are able to secret immunomodulatory factors and suppress immunocytes directly via a cell–cell interaction [ 45 ]. Indoleamine 2, 3‐dioxygenase (IDO), human leukocyte antigen G (HLA‐G), hepatocyte growth factor (HGF) and intercellular adhesive molecule‐1 (ICAM‐1) have been implicated in MSC‐mediated immunomodulation. IDO is an enzyme that converts tryptophan into kynurenine and leads to anergy in dendritic cells and T cells [ 46 ]. DPSCs from inflamed pulp have been shown to exhibit increased IDO expression compared to DPSCs from healthy pulp [ 47 ]. Another crucial cytokine, HLA‐G, is a nonclassical MHC class I involved in immunomodulation by inhibiting the activation of natural killer cells and effector T cells [ 48, 49 ]. A previous study reported that simulating the inflammatory environment of dental MSCs by incubating the cells with the pro‐inflammatory cytokine interferon‐γ results in increased HLA‐G expression in an interferon‐γ‐dependent manner [ 50 ]. Co‐culture experiments with peripheral blood mononuclear cells also drive the increased expression of IDO, soluble HLA‐G and HGF in MSCs [ 51, 52 ]. ICAM‐1 also functions in MSC‐mediated immunomodulation by inhibiting DC maturation and the T‐cell response [ 53, 54 ]. Studies have shown the improved immunosuppressive ability of MSCs with increased ICAM‐1 expression; however, this effect was impaired when ICAM‐1 was inhibited or knocked down in MSCs [ 55, 56 ]. PLDSCs was also reported to inhibit proliferation, function and pro‐inflammation cytokines secretion of immune cells [ 17, 57 ]. Yet, whether this inflammatory niche of periapical lesions alters the immunomodulatory ability of PLDSCs remains unknown. In the present study, the expression levels of IDO, HLA‐G and ICAM‐1 were higher in PLDSCs than in DPSCs, suggesting that the inflammatory environment of periapical lesions is beneficial for the immunomodulatory ability of PLDSCs and that this ability could be retained after culturing in vitro. This property of PLDSCs may be utilized in allogenic stem cell therapy or allogenic stem cell‐based dental pulp tissue engineering to achieve a weak graft versus host response. In summary, the present study demonstrates that PLDSCs might be optional seed cells for dental pulp regeneration, especially for dental neurovascularization with less ethical problems. As shown in Table 2, PLDSCs can actively proliferate in vitro and have weaker osteo/odontogenic ability, but have stronger migration, pro‐angiogenesis, neurotrophic and immunomodulatory abilities, compared to DPSCs. Therefore, they have promising application in autologous or allogenic stem cell‐based dental pulp regeneration. Beyond their application in dental disorders, they are also promising with respect to neurological disease [ 43 ]. However, for partial pulpectomy investigations with the aim of regenerating only dentin, DPSCs are recommended. Accordingly, the osteo/odontogenic differentiation ability of PLDSCs and correlations with different types of periapical lesions require more investigation. Furthermore, additional studies are warranted aiming to elucidate the mechanisms regulating the modified neurovascular induction and immunomodulatory abilities of these cells, their potential risk of tumorigenesis, and their in vivo effects. Table 2 Summary of the properties of DPSCs and PLDSCs DPSCs PLDSCs Morphology Spindle‐like shape Immunophenotype CD29(+)CD105(+)CD44(+)CD73(+)CD90(+)/CD31(–)CD34(–)CD45(–) Adipogenic and chondrogenic differentiation No obvious differences Proliferation Stronger Migration Stronger Osteo/odontogenesis Stronger Pro‐angiogenesis Stronger Neurotrophic ability Stronger Immunomodulatory ability Stronger John Wiley & Sons, Ltd Conflict of interests The authors declare that they have no conflicts of interest. Author contributions WPL and MYM conducted most of the experiments and drafted the manuscript. NH contributed to the collection of clinical samples and data analysis, and also revised the manuscript. JW and JH revised the manuscript. SSG contributed to study conception and design and critically revised the manuscript. All authors read and approved the final manuscript submitted for publication.
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10. 1002/2211-5463. 13352
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Bone marrow‐mesenchymal stem cell‐derived extracellular vesicles affect proliferation and apoptosis of leukemia cells
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Mesenchymal stem cells (MSCs) have been proposed to have potential for tissue engineering and cell therapy due to their multilineage differentiation potential and ability to secrete numerous paracrine factors, including extracellular vesicles (EVs). Increasing evidence has demonstrated that MSC‐derived EVs (MSC‐EVs) are able to induce the repair of tissue damage and regulate the immune system. However, their role in cancer development is still unclear. Reports have suggested that whether MSC‐EVs have an inhibitory or promoting effect on cancer is dependent on the type of cancer. In this study, the role of MSC‐EVs in the regulation of leukemic cell growth in vitro was investigated. The EVs were collected from conditioned media of MSCs by ultrafiltration using a 10 kDa molecular weight cutoff (MWCO) filter. The isolated MSC‐EVs were comprised of microvesicles and exosomes, as examined by the size of vesicles and exosomal proteins, CD81 and flotillin‐1. Cell proliferation, cell cycle status, apoptosis, and gene expression were examined in the leukemic cell lines NB4 and K562 after treatment with MSC‐EVs. Suppression of cell proliferation and induction of apoptosis was observed. Gene expression analysis revealed differential expression of apoptotic‐related genes in NB4 and K562. MSC‐EVs increased the expression of BID and BAX and decreased expression of BCL2, indicating the induction of intrinsic apoptosis in NB4. In contrast, MSC‐EVs increased the expression of the death receptor gene TRAILR2 and cell cycle regulator genes P21 and CCNE2 in K562. In conclusion, MSC‐EVs partially induce leukemic cell apoptosis, and thus may have potential for the development of supportive therapies for leukemia.
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Abbreviations BAX BCL2‐associated X apoptosis regulator BCL2 apoptosis regulator BID BH3 interacting domain death agonist CCNE2 cyclin E2 FAS Fas cell surface death receptor GAPDH glyceraldehyde‐3‐phosphate dehydrogenase P21 cyclin‐dependent kinase inhibitor 1A (CDKN1A) P53 tumor protein p53 PUMA p53 upregulated modulator of apoptosis TRAILR2 TNF‐related apoptosis‐inducing ligand receptor 2 Mesenchymal stem cells (MSCs) are multipotent stem cells that have the ability to self‐renew and differentiate into multiple cell types in the mesodermal lineage. A variety of tissue types is found to be the source of MSCs [ 1 ]. MSCs have been extensively studied during the past decade, in particular, in the field of regenerative medicine. In addition to MSC themselves, accumulating evidence reveals that MSCs exert effects on surrounding cells through the paracrine activity by secretion of various soluble factors including extracellular vesicles (EVs) [ 2, 3 ]. Generally, there are three types of EVs released from the cells―exosomes, microvesicles, and apoptotic bodies that are classified by the mechanism of biogenesis. However, the EVs can be classified based on their characteristics that are examined in vitro according to the Minimal Information for Studies of Extracellular Vesicles (MISEV) criteria [ 4 ]. Among the subtypes of EVs, exosomes and microvesicles have been investigated and are proposed to use as a cell‐free therapeutic approach that could overcome the limitations of cell therapy. Exosomes, the smallest type of EVs (typically 40–100 nm), are generated from the internal vesicles of multivesicular bodies (MVBs), which are subsequently released into the extracellular space, while microvesicles (50–1000 nm) are shed by outward blebbing of the plasma membrane [ 5 ]. EVs serve as a vehicle carrying a variety of molecules, including proteins, lipids, metabolites, and nucleic acid (microRNA, mRNA) to function in intercellular communication in normal physiological and also pathological conditions. In recent years, MSC‐derived EVs (MSC‐EVs) have demonstrated favorable results for the treatment of various diseases, including cancers. However, the role of MSC‐EVs in cancer is still controversial. They exhibited antitumor activity in some cancer types such as pancreatic cancer [ 6 ], prostate cancer [ 7 ], and glioma [ 8 ] while they showed promoting effects on the others [ 9, 10, 11 ]. This variation was suggested to be in accord with the type of cancer as well as the source of MSCs and genetic modification of MSCs, which need to be clarified before clinical use. Leukemia is a result of uncontrolled proliferation of abnormal immature hematopoietic cells, which further accumulate in the bone marrow and interrupt normal hematopoiesis. There are several types of leukemia, classified by characteristics of abnormal blood cell types that respond to therapy differently. Hematopoietic stem cell transplantation (HSCT) is considered the curative treatment in postremission therapy of leukemia. The roles of MSCs in HSCT have been investigated in various conditions. Evidence revealed a favorable effect of MSCs on the prevention of graft‐versus‐host disease (GVHD) in patients with hematological malignancies undergoing allogeneic HSCT [ 12, 13 ]. Interestingly, it has been demonstrated that MSC‐EVs recapitulated the immunomodulatory roles of MSCs in the GVHD condition indicating alternative MSC products for clinical use [ 14, 15 ]. However, the role of MSC‐EVs on leukemic cells has not been clearly elucidated. Importantly, there are a limited number of previous studies that demonstrated the activity of unmodified MSC‐EVs on leukemic cells. In the present study the effect of MSC‐EVs on leukemic cells was investigated. The MSCs were isolated from bone marrow and MSC‐EVs were collected from conditioned medium of MSC culture. The leukemic cell lines, NB4 and K562, which are derived from acute promyelocytic leukemia and chronic myeloid leukemia, respectively, were used in this study. The results revealed the inhibitory effect of MSC‐EVs on both leukemic cells, thus supporting the antitumor activity of MSC‐EVs. Materials and methods Cell culture Bone marrow‐derived MSCs (BMMSCs) were isolated from subjects during the operation. The procedure was approved by the Ethical Committee of the Faculty of Medicine, Ramathibodi Hospital, Mahidol University (MURA2017/603). The study methodologies conformed to the standards set by the Declaration of Helsinki. Written informed consent was obtained from all subjects before sample collection. The mononuclear cells were isolated from bone marrow aspiration by density gradient centrifugation using Histopaque‐1077 (Sigma‐Aldrich, St. Louis, MO, USA). The isolated mononuclear cells were washed with phosphate‐buffered saline (PBS) and seeded into a tissue culture flask in a growth medium containing Dulbecco's modified Eagle's medium (DMEM, Gibco, Grand Island, NY, USA), 10% fetal bovine serum (FBS, Merck, Darmstadt, Germany), and 1% penicillin/streptomycin (Gibco). The cells were cultured for 3 days before removing the floating cells. The adherent cells were cultured until 80% confluence and were ready to be passaged using 0. 5% trypsin‐EDTA (Gibco). Bone marrow‐derived MSCs were characterized for the typical MSC markers including surface marker expression and multilineage differentiation potency as previously described [ 16 ]. Briefly, the expression of CD73, CD90, CD105, CD34, and CD45 was assessed by flow cytometry. Osteogenic and adipogenic differentiation was determined by culturing the cells in osteogenic and adipogenic differentiation medium for 14–21 days before staining with Alizarin Red S and Oil Red O, respectively. The leukemic cell lines, NB4 and K562, were purchased from ATCC (Rockville, MD, USA). Leukemic cells were cultured in Roswell Park Memorial Institute 1640 Medium (RPMI, Gibco) supplemented with 10% FBS and 1% penicillin/streptomycin. Cells were passaged at a ratio of 1:4 every 2 days. The cell viability was determined by trypan blue staining (Gibco) during culture. Isolation of EVs from conditioned medium of MSCs Bone marrow‐derived MSCs at passage 4–5 were seeded into a tissue culture flask at a density of 5 × 10 5 cells/T75 and cultured in growth medium for 2 weeks. The medium was removed and the cells were washed several times with PBS. After discarding the PBS, 10 mL of serum‐free medium was added to the flask and the cells were cultured for 48 h before collecting the conditioned medium (CM). The CM was centrifuged at 300 × g for 5 min to remove cell debris followed by filtration with a 0. 22‐µm filter. To remove apoptotic bodies, the CM was centrifuged at 10, 000 × g for 1 h and the pellet was discarded. The supernatant was collected and ultrafiltration performed using 10 kDa MWCO filter (Amicon, Beverly, MA, USA) to collect the EVs. At the last step of ultrafiltration, the column was filled with PBS and recentrifuged to suspend the EVs in PBS (Fig. 1E ). The EV protein was quantified using the BCA assay (Pierce, Rockford, IL). The size of the EVs was examined by flow cytometry comparing it with the standard fluorescent polystyrene particles with sizes ranging from 0. 22 µm to 1. 35 µm (SheroTech, Lake Forest, IL, USA). Fig. 1 Characterization of bone marrow‐derived MSCs and MSC‐EVs. (A) The MSCs display fibroblast‐like morphology. (B) Alizarin Red S staining shows the matrix mineralization, the characteristic of osteogenic differentiation. (C) Oil Red O staining shows the differentiated adipocytes. (D) Flow cytometric analysis of cellular markers including CD73, CD90, CD105, CD34, and CD45 of the MSCs is shown. The signal from the unstained control is presented in the white histogram. The percentage of each CD marker’s positive cells are presented as mean ± SD ( n = 4). (E) Schematic diagram shows the process of EV isolation from 48 h CM of MSCs. (F) Flow cytometric analysis shows the pattern of the standard fluorescent polystyrene particles that was gated by FITC intensity followed by the forward scatter (FSC) and side scatter (SSC) plot. The size of the MSC‐EVs was determined using FSC and SSC plots according to the standard particles. (G) Exosomal proteins, CD81 and flotillin‐1, of MSC‐EVs were examined by western blot ( n = 4). Scale bar, 200 µm. Western blot Exosomal protein makers were determined using western blot. Fifteen micrograms of EVs protein were resolved with sodium dodecyl sulfate‐polyacrylamide gel electrophoresis (SDS‐PAGE, Sigma‐Aldrich) and blotted to a polyvinylidene difluoride (PVDF) membrane (Merck Millipore, Bedford, MA, USA). The blotted membrane was blocked with 5% skim milk (Sigma‐Aldrich) for 1 h at room temperature followed by staining with anti‐CD81 (1:250, cat. 10630D, Invitrogen, La Jolla, CA, USA) and antiflotillin‐1 (1:2000, cat. F1180, Sigma‐Aldrich) at 4 °C overnight. The membrane was washed with buffer several times to remove primary antibodies before staining with horseradish peroxidase (HRP)‐conjugated secondary antibody for 1 h at room temperature. The chemiluminescent signal was developed using ECL Prime western blotting detection reagent (GE Healthcare, Chicago, IL, USA) and visualized by chemiluminescent detection instrument (Bio‐Rad, Hercules, CA, USA). Proliferation assay Proliferation of leukemic cell lines after treatment with MSC‐EVs was examined by MTT assay. The leukemic cell lines were seeded in triplicate at a density of 1. 5 × 10 4 cells/well of the 96‐well plate and treated with 0, 50, and 100 µg·mL −1 MSC‐EVs in a reduced‐serum medium (RPMI, 2% FBS, 1% penicillin/streptomycin) for 48 h. Fifty microliters of 1 mg·mL −1 MTT reagent (Invitrogen) was added to the well and the plate was incubated for 4 h in a humidified 37 °C incubator. To dissolve formazan crystals formed in the cells, 100 µL of 10% SDS in 0. 01 M hydrochloric acid was added to the well and incubated overnight for complete dissolution. The optical density at 540 nm was measured using a spectrophotometer microplate reader. The data were collected as the absorbance value and presented as percent of control ( n = 3). Apoptosis analysis Cell apoptosis was examined using FITC Annexin V apoptosis detection kit (BD Biosciences, San Jose, CA, USA) and analyzed by flow cytometry. The leukemic cell lines were treated with 0 and 100 µg·mL −1 of MSC‐EVs for 48 h. The cells were collected, washed twice with cold PBS, and suspended in 1x binding buffer at a concentration of 1 × 10 6 cell·mL −1. One hundred microliters of cell suspension were mixed with 5 µL of FITC‐conjugated Annexin V and 5 µL of Propidium iodide (PI). After incubation for 15 min at room temperature, 400 µL of 1x binding buffer was added to the suspension and performed flow cytometric analysis within 1 h using FACS Canto II and FACSDiva software (BD Biosciences, San Jose, CA, USA). Cell cycle analysis The cell cycle was examined using PI staining and analyzed by flow cytometry. The leukemic cell lines were treated with 0 and 100 µg·mL −1 of MSC‐EVs for 48 h. The cells were collected, washed with cold PBS, and suspended in 1 mL of cold PBS. To fix the cells, cell suspension was added dropwise to 4. 5 mL of cold 70% ethanol, mixed thoroughly, and incubated for 12–24 h at 4 °C. Cell suspension was centrifuged at 500 × g, 5 min to discard the ethanol. After washing with PBS, the cell pellet was suspended in 1 mL of PI staining solution containing 10 µg·mL −1 PI, 100 µg·mL −1 RNase A, and 0. 1% Triton X‐100 in PBS (all from Sigma‐Aldrich). The cells were incubated for 15 min at room temperature before flow cytometric analysis using FACS Canto II and facsdiva software (BD Biosciences). Gene expression The leukemic cell lines were treated with 0, 50, and 100 µg·mL −1 of MSC‐EVs for 48 h. Total RNA was harvested using TRIzol reagent according to the manufacturer’s instruction (Invitrogen). The quantity and purity of RNA was examined by OD at 260 nm and the ratio of OD 260/280, respectively, using NanoDrop (ThermoFisher, Waltham, MA, USA). One microgram of RNA was reverse transcribed to cDNA using RevertAid first strand cDNA synthesis kit (Thermo Scientific, Waltham, MA, USA). Real‐time polymerase chain reaction (PCR) assay was carried out by CFX Bio‐Rad using SYBR green master mix (Kappa) and the primers specific to the interested genes (Table 1 ). The primer sequences were designed by a primer designing tool (NCBI, Bethesda, MD, USA), Primer 3, and BLAST. The nucleotide sequences of the genes were obtained from the NCBI database. The gene expression was normalized to glyceraldehyde‐3‐phosphate dehydrogenase ( GAPDH, housekeeping gene) and presented as a relative level to control. Table 1 Primer sequences. Gene Primer sequences (5’–3’) Product size (bp) BID Forward: AGACTGATGGCAACCGCAG 133 Reverse: GGGATGCTACGGTCCATGCT BAX Forward: AGGATGCGTCCACCAAGAAG 137 Reverse: AGCTGCCACTCGGAAAAAGA BCL2 Forward: TCCTGCATCTCATGCCAAGG 191 Reverse: TCCCAGAGGAAAAGCAACGG PUMA Forward: GGATGAAATTTGGCATGGGGT 168 Reverse: TAAGGGCAGGAGTCCCATGA FAS Forward: AATAAACTGCACCCGGACCC 192 Reverse: AGAAGACAAAGCCACCCCAA TRAILR2 Forward: TAAGTCCCTGCACCACGAC 190 Reverse: CCACTGTGCTTTGTACCTGATTC P53 Forward: CCTCTCCCCAGCCAAAGAAG 100 Reverse: GCCTCATTCAGCTCTCGGAA P21 Forward: GATGAGTTGGGAGGAGGCAG 156 Reverse: CTGAGAGTCTCCAGGTCCAC CCNE2 Forward: GCTGGTCTGGCGAGGTTTT 248 Reverse: AATGCAAGGACTGATCCCCC GAPDH Forward: CAACTACATGGTTTACATGTTCCAA 206 Reverse: CAGCCTTCTCCATGGTGGT John Wiley & Sons, Ltd Statistical analysis Data are presented as mean ± standard deviation (SD). Statistical analysis was assessed using graphpad Prism v. 5. 00 for Windows ( graphpad software, San Diego, CA, USA). Student’s t ‐test was used to evaluate the significant difference between two groups while the ANOVA test with Tukey's multiple comparison test was used to evaluate the difference of multiple groups. The difference was considered statistically significant at P < 0. 05. Results Characterization of bone marrow‐derived MSCs and MSC‐EVs After culture in vitro, bone marrow‐derived MSCs showed a plastic adherent property with fibroblast‐like morphology (Fig. 1A ). Typical characteristics of MSCs were examined according to the minimal criteria defined by the International Society for Cellular Therapy [ 17 ]. The MSCs demonstrated trilineage differentiation potency when cultured in differentiation induction medium as shown by positive staining for Alizarin Red S, Oil Red O staining, and expression of the chondrogenic gene (Fig. 1B, C, Fig. S1 ). Immunophenotypic analysis revealed that more than 95% of MSCs were positive for CD73, CD90, CD105 and less than 2% were positive for CD34 and CD45 (Fig. 1D ). The characterized MSCs from four donors were continuously grown for EV isolation separately. After differential centrifugation of the CM to remove apoptotic bodies that are large vesicles (>1 µm), the EVs were collected by ultrafiltration with 10 kDa MWCO filter (Fig. 1E ). By this isolation technique, 1–2 mg of MSC‐EV protein was collected from 100 mL CM. The size of the isolated MSC‐EVs examined by comparison with the standard polystyrene particles was ~200 nm to 1 µm (Fig. 1F ). Although the conventional flow cytometer could not distinguish the exosomes, which are smaller than 200 nm from the noise, the isolated MSC‐EVs expressed the specific exosomal proteins, CD81, and flotillin‐1 as examined by western blot (Fig. 1G ). These results indicated the mixture of exosomes and microvesicles in the isolated MSC‐EVs. Effect of MSC‐EVs on the proliferation of leukemic cells The leukemic cell lines, NB4 and K562, were treated with 50 and 100 µg·mL −1 MSC‐EVs for 48 h before assessing cell proliferation by the MTT assay. The proliferation of NB4 was not different from control after treatment with 50 and 100 µg·mL −1 MSC‐EVs. In K562, treatment with 50 and 100 µg·mL −1 MSC‐EVs exhibited a significant decrease in the growth potential compared with control, suggesting the inhibitory effect of MSC‐EVs on the proliferation of K562 (Fig. 2 ). Fig. 2 Effect of MSC‐EVs on the proliferation of leukemic cells. NB4 and K562 were treated with MSC‐EVs at concentrations of 50 and 100 µg·mL −1 for 48 h. The experiment was performed three times using EVs collected from three lines of MSCs ( n = 3). Cell proliferation was examined by MTT assay. The MTT results (absorbance values) are presented as percentage to the control group of each experiment before statistical analysis (mean ± SD). * P < 0. 05 versus control. Statistical significance was determined using one‐way ANOVA followed by Tukey's multiple comparison test. Effect of MSC‐EVs on cell cycle status and apoptosis of leukemic cells To investigate whether the inhibitory effect of MSC‐EVs on leukemic cell proliferation involved the regulation of the cell cycle, the cell cycle phase of NB4 and K562 was examined by PI staining. Leukemic cells were treated with 100 µg·mL −1 MSC‐EVs for 48 h before PI staining. The results revealed that MSC‐EVs influenced the cell cycle phase of leukemic cells. Sub‐G1 population that was relevant to apoptotic cell death was found to be significantly increased, while the S phase population were decreased in NB4 after MSC‐EV treatment (Fig. 3A ). In K562, MSC‐EVs significantly decreased the S phase population, while they slightly increased the sub‐G1 population; however, it was not significantly different from the control ( P = 0. 0751) (Fig. 3B ). The ability of MSC‐EVs to induce leukemic cell apoptosis was examined by Annexin‐V/PI staining. In NB4, the early and late apoptosis were not different between MSC‐EVs and the control group (Fig. 3C ). In K562, the early apoptosis was significantly increased after MSC‐EV treatment; however, the total apoptosis was not different from the control group ( P = 0. 0774) (Fig. 3D ). Fig. 3 Effect of MSC‐EVs on cell cycle and apoptosis of leukemic cells. Histogram shows the cell cycle analysis according to the intensity of PI staining in (A) NB4 and (B) K562 after treatment with 100 µg·mL −1 MSC‐EVs for 48 h. Graphs present the percentage of the cells in each phase ( n = 3). * P < 0. 05 versus control. Flow cytometric analysis of Annexin‐V/PI staining of (C) NB4 and (D) K562 after treatment with 100 µg·mL −1 MSC‐EVs for 48 h. Graph presents the percentage of early (Annexin‐V+, PI‐), late (Annexin‐V+, PI+), and total (Annexin‐V+) apoptotic cells of leukemic cells ( n = 3). * P < 0. 05 versus control. Statistical significance was determined using Student’s t ‐test. MSC‐EVs increased the expression of apoptotic‐related genes The expression of genes involved with apoptosis and cell cycle arrest was examined in NB4 and K562 after MSC‐EV treatment for 48 h. In NB4, MSC‐EVs significantly increased the expression of the proapoptotic genes, BID and BAX, while it decreased the expression of the antiapoptotic gene, BCL2, compared with the control. In contrast, the expression of BID, BAX, and BCL2 in K562 after MSC‐EV treatment was not different from the control. Interestingly, MSC‐EVs increased the expression of the death receptor, TRAIL2 in K562. The P53 and PUMA were likely induced; however, there were no significant differences between MSC‐EVs and the control group both in NB4 and K562. The expression of the cell cycle regulator gene, P21, and its downstream target gene, CCNE2, were significantly increased in K562 after MSC‐EV treatment, while they were not affected in NB4 (Fig 4 ). Fig. 4 Relative expression of genes involved with apoptosis and cell cycle arrest in leukemic cell after MSC‐EV treatment. The level of mRNA expression was normalized to GAPDH, a housekeeping gene. The expression of each gene is presented as relative expression compared to the control group (mean ± SD). * P < 0. 05 versus control ( n = 3). Statistical significance was determined using one‐way ANOVA followed by Tukey's multiple comparison test. Discussion During the last decade, MSC‐EVs have received growing interest in regenerative medicine due to their potential to regulate immune systems and stimulate tissue regeneration [ 18 ]. In cancer, the results obtained from numerous studies demonstrated a dual role of MSC‐EVs on cancer development and progression that might depend on the type and stage of the cancer as well as the source of MSCs and the methodology of EV isolation [ 19 ]. Those evidences have led to attentive consideration of using MSC‐EVs in cancer therapy. In the present study, EVs were collected from bone marrow‐derived MSCs by ultrafiltration with a 10 kDa MWCO filter that has been used as an alternative method to ultracentrifugation to reduce vesicle loss and damage [ 20 ]. The MSC‐EVs that comprised exosomes and microvesicles showed the suppressive effect on leukemic cells growth in vitro. Although the effect was not clearly demonstrated, the MSC‐EVs partially induced apoptosis and cell cycle arrest in NB4 and K562. These results were similar to the previous studies that revealed the antiproliferative and proapoptotic effect of MSC‐EVs on leukemic cells, although the EVs were derived from different sources and techniques [ 21, 22, 23 ]. To reveal the potential use of EVs in a cancer treatment application, the combination of MSC‐EVs and chemotherapeutic agents has been studied. The MSC‐EVs could increase the sensitivity of chemotherapeutic drugs to induce leukemic cell apoptosis suggesting the advantage of MSC‐EVs in supportive treatment, particularly in the case with chemo‐resistance [ 24 ]. Even though the cellular characteristics of NB4 and K562 after exposure to MSC‐EVs were similar, the apoptosis‐associated genes expression analysis was different between NB4 and K562. In NB4, the MSC‐EVs increased the expression of the proapoptotic genes, BID and BAX, while it decreased the expression of the antiapoptotic gene, BCL2, indicating the induction of the intrinsic apoptosis pathway. The MSC‐EVs rather induced extrinsic apoptosis in K562 through increasing the expression of TRAILR2, the death receptor. Furthermore, MSC‐EVs increased the expression of P21 and its downstream target gene, CCNE2, in K562, indicating the induction of cell cycle arrest in K562. P53 has been identified to play a central role in regulating apoptosis and cell cycle arrest. P53 has diverse roles in modulating apoptotic pathways by direct induction of several proapoptotic genes in the BCL2 family ( PUMA, NOXA, BID, BAD, BAX, BAK ) and death receptor genes ( FAS, TRAILR ) that lead to intrinsic and extrinsic apoptosis, respectively. In addition, p53 negatively controls cell cycle by inducing cyclin‐dependent kinase inhibitor (p21) resulting in G1/S cell cycle arrest [ 25 ]. In the present study, several target genes of p53 were found to be activated in NB4 and K562 after MSC‐EV treatment, suggesting that the apoptosis might be mediated, at least in part, by the p53‐mediated pathway. However, the expression of the P53 gene was not increased by the MSC‐EVs, suggesting that there might be other pathways involved, which needs further study. Various signaling molecules and microRNAs (miRNAs) have been investigated to be involved in the antitumor activity of MSC‐EVs. Most of the studies focused on the function of miRNAs carried by the MSC‐EVs. The study of miRNA profiling revealed a diversity of miRNAs in MSC‐EVs, some of which have been reported to play roles in cancer‐regulatory activity by targeting several cancer‐survival pathways [ 26 ]. Additionally, genetically modified MSC‐EVs with overexpression of certain miRNAs demonstrated a beneficial effect in antitumor activity, while unmodified MSC‐EVs showed the variation of results [ 27 ]. This evidence implies that to use MSC‐EVs as an effective cancer therapeutic tool, genetic modification of MSC‐EVs is required. Owing to the controversial effects of MSC‐EVs on cancer cells, the use of MSC‐EVs in cancer therapy has to be considered. As reported here, although MSC‐EVs alone were insufficient to kill leukemic cells, they did not show the promoting effect on leukemic cell growth in vitro. Moreover, the MSC‐EVs partially induce leukemic cell apoptosis, which would be beneficial for suppression of minimal residual disease, in particular during treatment with HSCT. In our finding, the study with NB4 and K562 enlightens the influence of MSC‐EV on some types of leukemic cell lines. The molecular mechanism causing cell apoptosis seemed to be different between both cell lines. Therefore, further studies involving more advanced techniques, testing with other leukemic cell lines, as well as the primary leukemic cells are highly needed. In addition, the combination of MSC‐EVs with known leukemia chemotherapeutic drugs is interesting in order to investigate the beneficial roles of MSC‐EVs as synergistic effects with the former treatments. Moreover, as in a review by Bailey et al [ 27 ], the result from an in vivo model might be different from the in vitro study, especially when using unmodified MSC‐EVs. Thus, the functional test the in vivo model is required before clinical application. Conflict of interest The authors declare no conflicts of interest. Author contributions AS and JP designed the study. JP performed the experiments and analyzed the data. TT collected the BM tissue. PO, DT, and SB participated in the interpretation of the data. AS, JP, and WK wrote the article. Supporting information Fig. S1. Chondrogenic differentiation potential of BMMSCs. Click here for additional data file.
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10. 1002/2211-5463. 13615
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Bone marrow mesenchymal stem cell‐derived exosomes reduce insulin resistance and obesity in mice via the
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Obesity is a common chronic metabolic disease that induces chronic systemic inflammation in the body, eventually leading to related complications such as insulin resistance (IR), type 2 diabetes mellitus, and metabolic syndromes such as cardiovascular disease. Exosomes transfer bioactive substances to neighboring or distal cells through autosomal, paracrine, or distant secretion, regulating the gene and protein expression levels of receptor cells. In this study, we investigated the effect of mouse bone marrow mesenchymal stem cell‐derived exosomes (BMSC‐Exos) on high‐fat diet obese mice and mature 3T3‐L1 adipocyte models of IR. BMSC‐Exo treatment of obese mice promoted their metabolic homeostasis, including reduction of obesity, inhibition of M1‐type proinflammatory factor expression, and improvement of insulin sensitivity. In vitro analysis revealed that BMSC‐Exos improved IR and lipid droplet accumulation in mature 3T3‐L1 adipocytes treated with palmitate (PA). Mechanistically, BMSC‐Exos cause increased glucose uptake and improved IR in high‐fat chow‐fed mice and PA‐acting 3T3‐L1 adipocytes by activating the phosphoinositide 3‐kinases/protein kinase B (PI3K/AKT) signaling pathway and upregulating glucose transporter protein 4 (GLUT4) expression. This study offers a new perspective for the development of treatments for IR in obese and diabetic patients.
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Abbreviations BMSC bone marrow mesenchymal stem cell BMSC‐Exos bone marrow mesenchymal stem cell‐derived exosomes FABP4 fatty acid‐binding protein 4 FFA free fatty acid GLUT4 glucose transporter protein 4 GTT glucose tolerance tests H&E hematoxylin and eosin HFD high‐fat diet IL‐6 interleukin‐6 IR insulin resistance ITT insulin tolerance tests iWAT inguinal white adipose tissue MDI adipocyte differentiation agent NS normal saline NTA nanoparticle tracking analysis PA palmitate PI3K/AKT phosphoinositide 3‐kinases/protein kinase B scWAT subcutaneous white adipose tissue TEM transmission electron microscope TNF‐α tumor necrosis factor‐α Obesity has become a global epidemic. According to the World Health Organization, obesity has almost tripled since 1975, with approximately 650 million people diagnosed with obesity in 2016. The obese population is expected to increase to 12 billion by 2030 [ 1 ]. Recent data show that obesity is also strongly associated with novel coronavirus pneumonia. Obese patients are more susceptible to the virus than the general population, and obesity can exacerbate symptoms, cause poor prognosis and significantly increase mortality in patients [ 2 ]. In addition, obesity can also induce chronic systemic inflammation in the body, eventually leading to a series of related complications, such as insulin resistance (IR), type 2 diabetes and cardiovascular disease, and other metabolic syndromes [ 3, 4, 5 ]. Therefore, it is urgent to prevent the treatment of obesity and control the further development of its related complications. Obesity refers to a state in which the body's energy intake is more significant than its energy consumption, disrupting the balance of energy metabolism in the body and causing an excess accumulation of white adipose tissue (WAT) under the skin and in the internal organs. More than 200 genes have been linked to the development of obesity [ 6, 7 ]. For example, leptin, a product encoded by the obesity gene, is produced primarily by adipocytes [ 8 ]. Leptin is overexpressed at the gene level in the adipose tissue of obese individuals [ 9 ]. Leptin, a cytokine, is elevated in circulating levels in obese patients and can lead to hypo‐inflammation [ 10 ]. Fatty acid binding protein 4 (FABP4, also known as aP2), a cytoplasmic fatty acid chaperone, is expressed mainly in adipocytes and bone marrow cells [ 11 ]. The high expression of FABP4 in the obese state of the body exacerbates many immunometabolic diseases, including diabetes and IR [ 12, 13 ]. In mouse models and humans, circulating FABP4 levels correlate with the incidence of metabolic disease, and lowering FABP4 levels or activity is associated with improved metabolic health [ 14 ]. White adipose tissue is traditionally considered the body's primary energy storage site. However, many studies have found that WAT is also a dynamic endocrine organ that secretes various cytokines, regulates communication within WAT and between WAT and other organs and cells, and participates in the body's metabolic homeostasis [ 15, 16, 17 ]. In an obese state, adipose tissue releases tumor necrosis factor‐α (TNF‐α) and interleukin‐6 (IL‐6), causing chronic inflammation in the adipose tissue, where large numbers of M1 proinflammatory macrophages are recruited to produce inflammatory factors that lead to IR in the adipose tissue [ 18 ]. In conclusion, abnormalities in the function of the adipose organs are a critical factor in the body's obesity and IR. Insulin resistance is a condition in which the biological effect of insulin on target tissues is impaired, the efficiency of promoting glucose uptake and utilization is reduced, and the body compensates by producing excess insulin to produce hyperinsulinemia and maintain blood glucose stability [ 19 ]. Many factors can lead to IR, and obesity is one of the most important. The metabolic inflammation caused by obesity starts after WAT. Many inflammatory factors will circulate to the liver, muscle, and other insulin‐sensitive organs, interfering with the PI3K/AKT insulin signaling pathway conduction and leading to systemic IR [ 20, 21, 22 ]. With the continuous development of cell therapy technology, the quintessential role of stem cells in weight loss has been corroborated by extensive studies, and MSCs have especially attracted much attention. Studies have found that MSCs are a kind of pluripotent stem cells with proliferation, renewal, and multi‐directional differentiation, which have the functions of anti‐inflammation, immune regulation, inhibiting fibrosis, and promoting angiogenesis [ 23, 24 ]. Furthermore, it has a wide range of sources, which can be isolated from bone marrow, umbilical cord, fat, amniotic fluid, placenta, synovium, and synovial fluid [ 25 ]. Bone marrow mesenchymal stem cells (BMSCs) are mainly distributed in the femur, tibia, and iliac crest, which is easy to isolate and culture from tissues. BMSCs are considered to be one of the ideal seed cells in the field of tissue engineering since that has a strong ability for self‐renewal and genetic modification [ 26, 27 ]. It has been pointed out that MSCs repair tissue damage not only by differentiating to achieve the regeneration of damaged cells but more importantly, by releasing signaling molecules to damaged tissues through the mechanism of paracrine vesicles to enable tissue function to be restored [ 28, 29 ]. Exosomes are essential components of these paracrine vesicles and are important carriers of signal communication between stem cells and target cells [ 30 ]. Exosomes are extracellular vesicles 40–150 nm in diameter, and their contents include components such as nucleic acids, proteins, and enzymes. Exosomes are secreted by a variety of cells throughout the body, and the expression of internal components of exosomes varies depending on the source cell type and the environment in which they are located [ 31, 32, 33, 34 ]. At present, it is believed that exosomes secreted into extracellular space are recognized by target cells through ligand‐receptor binding, subsequently enter the cells through endocytosis, releasing endogenous signaling molecules to complete information transmission, and ultimately regulate target cell functions, such as promoting tissue repair and immune regulation [ 35, 36, 37 ]. The therapeutic effects of exosomes have been demonstrated in most tissues and organs. For instance, exosomes can render it possible to improve myocardial ischemia–reperfusion injury [ 38 ], promote angiogenesis to prevent diabetic nephropathy [ 39 ], promote nerve cell repair [ 40 ], and so on. The project will focus on the ameliorative effects of BMSC‐Exos on obesity‐induced inflammation and IR, providing new ideas and targets for preventing and treating obesity, IR, and related metabolic syndromes. Materials and methods Experimental animals and sample collection C57BL/6 mice (7 weeks old, male) were purchased from the Experimental Animal Center of Shanxi Provincial People's Hospital and were maintained under constant conditions (temperature, 22 ± 3 °C; humidity, 40–50%). After 1 week of acclimatization, mice were divided into a group given a normal diet (NCD), a group of mice fed a high‐fat diet (HFD) of 60% of total calories (HDF); and a group of mice fed a HFD of 60% of total calories administered BMSC‐Exos treatment (HDF + Exosome). Mouse chow was purchased from Jiangsu Xietong Pharmaceutical Bio‐engineering (Jiangsu, China). The HDF and HDF + exosome groups were fed a HFD for 12 weeks to induce obesity. During the last 4 weeks of HFD feeding, the HDF + exosome group was treated with BMSC‐Exos, administered with an intraperitoneal injection of 50 μg of BMSC‐Exos every 3 days per animal for a total of 4 weeks. HFD or NCD mice fed with normal saline (NS) were used as controls, and their body weight and dietary intake were recorded weekly. After the intervention, mice were executed under anesthesia. Inguinal white adipose tissue (iWAT) and subcutaneous white adipose tissue (scWAT) were collected and weighed, with some tissue fixed in 4% paraformaldehyde and the rest stored at −80 °C until analysis. All animal care and experimental protocols complied with the Animal Management Rule of the Ministry of Health, People's Republic of China (Documentation No. 55, 2001) and the Guide for the Care and Use of Laboratory Animals published by the United States National Institutes of Health (Publication No. 85‐23, Revised 1996), and the Global Research Animal Guide. All animal operations were carried out in accordance with the ‘Guidelines for the Care and Use of Laboratory Animals of Shanxi Agricultural University’ and were approved by the Animal Medicine Committee of Shanxi Agricultural University [SXAU‐EAW‐2020M0725]. Glucose tolerance and insulin tolerance tests (GTT and ITT) For GTT, mice fasted without water for 12 h, glucose (2 g/kg body weight) was injected intraperitoneally, tail blood was taken at 0, 30, 60, 90, and 120 min, respectively, and blood glucose values were measured at different times to calculate glucose tolerance. For ITT, mice fasted without water for 4 h and were injected insulin intraperitoneally (1 U·kg −1 body weight). Tail blood was taken at 0, 30, 60, 90, and 120 min, and blood glucose values were measured at different times to calculate the insulin tolerance of the mice. Hematoxylin and eosin The adipose tissue was fixed with 4% paraformaldehyde for 24 h. After dehydrating and being transparent with gradient ethanol and xylene, it was embedded with wax for 4 h. Tissue sections with a thickness of 0. 8 μm were prepared by an automatic rotary slicer (RM2265; Leica, Wetzlar, Germany, Japan) after embedding tissue into wax blocks. The adipose tissue sections of NCD‐NS, HDF‐NS, and HDF‐exosome groups were collected, dewaxed, and rehydrated with xylene and gradient ethanol. Tissue sections were stained with hematoxylin and eosin (H&E) and sealed with neutral glue finally. Cell culture Mouse bone marrow mesenchymal stem cells and 3T3‐L1 cells were purchased from American Type Tissue Culture (ATCC) (Maryland, USA). The cells were cultured in a high glucose medium (01‐052‐1ACS; BI, Kibbutz Beit‐Haemek, Israel) containing 10% fetal bovine serum (0510; Sciencell, San Diego, California, USA) and in a constant temperature incubator (5% CO 2, 37 °C). Cell treatments After 3T3‐L1 cells were fully fused (day 1), they were induced to differentiate towards adipocytes with induction culture medium containing 0. 5 m m isobutylmethylxanthine (IBMX), 0. 25 μ m dexamethasone (Dex), and 10 μ m insulin. After 2 days of induction (day 3), cells were shifted to the insulin‐containing differentiation culture medium, and every 2 days with a change of differentiation culture, 3T3‐L1 cells were for a total of 4 days of stimulation. Mature 3T3‐L1 were adipocytes obtained for use in subsequent experiments. The inducing differentiation agents involved are referred to as MDI in the following. Fully differentiated 3T3‐L1 adipocytes were pretreated with BMSC‐Exos (10 and 20 μg·mL −1 ) for 24 h. Subsequently, fatty acid‐free 10% bovine serum albumin medium containing 1 m m palmitate (PA) was incubated for 24 h. BMSC‐Exos continued to be administered during this procedure. PA was added to simulate the pathological condition of lipotoxicity. In addition, to investigate the insulin signaling pathway, 3T3‐L1 adipocytes were stimulated with 100 n m of insulin during the last 15 min of PA action to demonstrate the effect of PA and BMSC‐Exos on the insulin‐activated signaling pathway. Purification of exosomes The BMSCs were cultured in fresh DMEM without FBS (basal medium) for 24–32 h until reaching about 85% confluency. When the number of dead cells increased under the inverted microscope, The above culture was stopped as the phenomenon appeared that the number of dead cells increased under the inverted microscope, collecting the cell's supernatant of exosome‐rich ones. The collected cell supernatant was centrifuged at low speed at 300 g, 10 min, 4 °C. Subsequently, the supernatant was centrifuged again at 2000 g, 10 min, 4 °C to collect the supernatant, at which point the precipitate was dead cells and apoptotic debris. Based on the above operation, we collected the supernatant at 10 000 g, 30 min, 4 °C, while the precipitate was discarded, at which point the precipitate was more giant vesicles. After centrifugation at 100 000 g, 90 min, 4 °C, the supernatant was carefully aspirated to leave the precipitate washed with PBS buffer (30 mL) and resuspended before centrifugation at 100 000 g, 90 min, 4 °C. The precipitate obtained after centrifugation is resuspended in 100 μL sterile PBS buffer and is ready for immediate use or storage at −80 °C. Transmission electron microscope We dropped 10 μL exosome solution on copper mesh, incubated at room temperature for 10 min and rinsed with sterile distilled water, and absorbed excess liquid with absorbent paper. After absorbing 10 μL drops of 2% uranyl acetate on the copper mesh for 1 min of negative staining, the floating solution was blotted off with filter paper aiming at better results of incandescent drying for 2 min. Finally, the copper mesh was observed under a transmission electron microscope (TEM), generally with 80 kV imaging. Nanoparticle tracking analysis The scattered light of nanoparticles in nanoparticle suspensions was detected after laser irradiation. The concentration of nanoparticles and their size and mass were calculated by counting the number of scattered particles as well as analyzing the particle trajectory of the exosomes. PKH67 With PKH67 dye (Sigma; PKH67GL, GER, Saint Louis, Germany), BMSC‐Exos were labeled, which were subsequently added to 3T3‐L1 cells for 24 h, then cells were fixed with 4% paraformaldehyde for 30 min, sealed with anti‐fluorescence attenuated blocking slices containing DAPI (S2110; Solarbio, Beijing, China), and observed with a confocal microscope (FV1000; Olympus, Tokyo, Japan). PKH67 staining was performed by utilizing a standard protocol to see the standard procedure for details. Cell CCK‐8 The cells containing 100 μL of the total system were seeded into the 96‐well plate (701001; NEST, Wuxi, JIangsu, China). After 12 h of preculture, differentiated 3T3‐L1 cells were treated with different concentrations of BMSC‐Exos for 24 h, and 10 μL CCK‐8 solution (40203ES80; YEASE, Shanghai, China) was added to per well at that time. The 450 nm absorbance was read using a microplate reader (filter maxF5; Molecular Devices, Sunnyvale, Silicon Valley Center, USA) after another 2 h of incubation. Oil Red O staining According to the manufacturer's instructions, the Adipogenesis Assay Kit Cell‐Based (ab133102; Abcam, Cambridge, UK) was used to analyze the contents of each group's lipid droplets. Briefly, the cells were washed twice with washing solution before adding lipid droplet analysis Oil Red O solution to the cells, and the staining was observed microscopically after incubating the cells for 20 min at room temperature, after which the stained lipid droplets were detected by reading the absorbance at 490 nm with an enzymatic standard. Western blotting Extracted protein samples were measured for concentration, and loadings were calculated using BCA (P0011; Beyotime, Shanghai, China). SDS/PAGE electrophoresis was performed, and the target proteins' gels were transferred to PVDF membranes (ISEQ00010; Millipore, Boston, American Massachusetts, USA). The gels were closed with 5% skimmed milk powder (abs9175; Absin, Shanghai, China) blocking solution for 2 h and incubated overnight at 4 °C. Primary antibodies include: anti‐Akt (1 : 1000, #40569, Rabbit; SAB, College Park, Maryland, USA), Phospho‐AKT (Ser473) (1 : 5000; 66444‐1‐Ig, Mouse; Proteintech, Wuhan, China), anti‐PI3K (1 : 1000, T40064, Rabbit; Abmart, Shanghai, China), Phospho‐PI3K (1 : 1000, T40065, Rabbit; Abmart), anti‐Leptin (1 : 1000, PA6011, Rabbit; Abmart), anti‐FABP4 (1 : 5000, 12802‐1‐AP, Rabbit; Proteintech), anti‐GLUT4 (1 : 5000, 66848‐1‐lg, Mouse; Proteintech), anti‐α‐tubulin (1 : 2000, 11224‐1‐AP, Rabbit; Proteintech), anti‐CD9 (1 : 1000, 20597‐1‐AP, Rabbit; Proteintech), and anti‐TSG101 (1 : 2000, 28283‐1‐AP, Rabbit; Proteintech). The target bands were incubated with HRP‐conjugated anti‐rabbit IgG (1 : 20 000, CW0156S; CWBIO, Beijing, China) for 1 h and then washed six times with TBST for 5 min each. The ECL luminescent solution detected target bands (CW0049S; CWBIO). The density of the target protein bands was normalized according to the density of α‐tubulin protein in the same sample. Real‐time fluorescence quantitative PCR Primer sequences were designed using Primer Bank, and primer synthesis was performed by Shanghai General Biological Company (Shanghai, China). RNA was extracted using RNAiso Plus (9109; Takara, Japan), and the quality was tested by 1% agarose gel electrophoresis. cDNA was synthesized according to the instructions of the reverse transcription kit (R223‐01; Vazyme, Shanghai, China). SYBR (Q711‐02; Vazyme), DEPC water, cDNA, and upstream and downstream primers were mixed proportionally into a 10 μL system for a polymerase chain reaction. All samples were processed on the real‐time step one software system in triplicate (ABI QuantStudio5, Thermo Fisher Scientific, Massachusetts, USA). Results were calculated from ΔΔ C T values. The primer sequences for qRT‐PCR used were: 5′‐TTGCTGACAGGATGCAGAAG‐3′ and 5′‐ACATCTGCTGGAAGGTGGAC‐3′ for β‐actin, 5′‐TCAAGCAGTGCCTATCCAGAAAGTC‐3′ and 5′‐GGGTGAAGCCCAGGAATGAAGTC‐3′ for Leptin, 5′‐AAGGTGAAGAGCATCATAACCCT‐3′ and 5′‐TCACGCCTTTCATAACACATTCC‐3′ for FABP4, 5′‐CACTTCACAAGTCGGAGGCT‐3′ and 5′‐CTGCAAGTGCATCATCGTTGT‐3′ for IL‐6 5′‐CCTGTAGCCCACGTCGTAG‐3′ and 5′‐GGGAGTAGACAAGGTACAACCC‐3′ for TNF‐α. Statistical analysis Experiment results were presented as mean ± standard error of the mean and analyzed with one‐way analysis of variance via graphpad prism 8. 3 software (GraphPad Software, San Diego, USA). Differences between the treatment group and the normal group were conducted using Student's t ‐test. image leb (Bio‐Rad, Hercules, California, USA) was used to analyze the results of western blot analysis. The relative expression level of the target protein was calculated from the ratio of the target to the internal reference. qRT‐PCR results were calculated based on the ΔΔ C T value. Result Identification of BMSC‐Exos Following the isolation of exosomes from BMSC supernatant cultures by ultracentrifugation, we did observe a large number of vesicles with intact membrane structure by transmission electron microscopy, which met the internationally certified criteria for the characteristics of exosomes, showing a round or oval shape with approximately 40–150 nm in diameter, a lightly stained center and clear edges of the vesicles, and low electron‐density material was seen in the lumen of the vesicles (Fig. 1A ). Nanoparticle tracking analysis (NTA) showed an average particle size of 109. 4 nm and a concentration of 7. 1 × 10 9 Particles·mL −1 (Fig. 1B ) and clearly presented the Brownian motion of BMSC‐Exos in solution (Fig. 1C ). By performing western blotting, we observed positive expressions of the exosomes surface markers CD9 and TSG101 (Fig. 1D ). The above results show that exosomes were successfully extracted from the supernatant of BMSC cells. To delve into the properties of exosomes, we labeled BMSC‐Exos with a PKH67 kit and added them to 3T3‐L1 cells at a concentration of 10 μg·mL −1 for 12 h. Subsequently, the nuclei were fixed and stained, and exosomes with green fluorescence were clearly presented to be taken up by 3T3‐L1 under confocal microscopy (Fig. 1E ). Fig. 1 Characteristics of BMSC‐Exos. (A) The morphology of BMSC‐Exos showed a bilayer spherical vesicle structure mounted by TEM (bars = 200 nm). (B) NTA results showed that the average particle size of BMSC‐Exos was 109. 4 nm. (C) Screenshot of Brownian motion video of BMSC‐Exos. (D) Western blotting detected the protein expressions of TSG101 and CD9 in BMSC‐Exos. (E) PKH67‐labeled BMSC‐Exos (green) was taken up by 3T3‐L1 cells (bars = 40 μm). BMSC‐Exos alleviates obesity, metabolic disorders, and inflammation in HDF‐fed mice Throughout the experiment, we monitored the changes in the body weight of the mice in each group. The results showed that the body weight of the mice in the HFD group was significantly higher than those in the NCD group, and BMSC‐Exos mitigated the persistent weight gain in the HDF‐fed mice (Fig. 2A, B ). We counted the food intake of each group of mice during the BMSC‐Exos intervention and did not find any statistical significance (Fig. 2C ). One of the key factors contributing to IR is obesity. We found that HDF‐fed mice exhibited severe glucose intolerance and IR. Administration of BMSC‐Exos significantly improved glucose tolerance and insulin sensitivity in HFD‐fed mice (Fig. 2E–H ). Obesity is characterized by hypertrophy and hyperplasia of adipose tissue. We analyzed iWAT and scWAT in each group of mice. We found that HDF feeding resulted in a significant increase in the weight of both types of fat as a percentage of body weight. In contrast, exosome‐treated groups decreased iWAT weight as a percentage of body weight (Fig. 2D ). H&E staining showed that adipocytes in iWAT and scWAT were significantly hypertrophied in the HDF group mice compared to the NCD group mice, while continuous administration of BMSC‐Exos significantly improved the hypertrophy of adipocytes in the HDF group mice (Fig. 2I–K ). Leptin and FABP4 were highly expressed in WAT as obesity genes. Leptin, FABP4 protein, and mRNA levels were significantly higher in the iWAT of obese mice in the HFD group compared to mice in the NCD group, and BMSC‐Exos suppressed the levels of Leptin, FABP4 protein, and mRNA to some extent in obese mice (Fig. 2L–N ). Fig. 2 BMSC‐Exos alleviates obesity, metabolic disorders, and inflammation in HDF‐fed mice. (A–O) C57BL/6 mice were given either NCD or HDF for 12 weeks. Mice with HDF were given 50 μg of BMSC‐Exos by intraperitoneal injection every 3 days in the last 4 weeks. Equal amounts of NS were administered intraperitoneally to NCD and HDF mice as control. (A) Body weight change ( n = 6 per group). (B) Last count of body weight of each group of mice before sampling ( n = 6 per group). (C) Mean daily food intake per mouse in the HDF‐NS and HDF‐exosome groups during the injection of BMSC‐Exos exosomes. (D) Weight of iWAT and scWAT as a percentage of body weight for each group of mice ( n = 6 per group). (E, F) Glucose tolerance test (GTT) and statistics of the relative area under the curve ( n = 6 per group). (G, H) Insulin tolerance test (ITT) and statistics of the relative area under the curve ( n = 6 per group). (I–K) iWAT and scWAT H&E staining and corresponding area statistics for each group (bars = 100 μm). (L–N) Protein and mRNA levels of iWAT obesity‐related genes (Leptin and FABP4) in various groups of mice ( n = 3 per group). (O) mRNA levels of iWAT inflammation‐related genes (IL‐6 and TNF‐α) in all groups of mice ( n = 3 per group). The error bars indicate the SEM, whereas comparisons between two groups were performed by an unpaired Student's test, * P < 0. 05; ** P < 0. 01; *** P < 0. 001. NCD‐NS, normal diet mice injected with normal saline; HDF‐NS, high‐fat‐fed mice injected with normal saline; HDF + Exosome, high‐fat‐fed mice injected with exosome. Obesity is a chronic state of low‐grade inflammation, usually accompanied by the accumulation of macrophages in WAT, which secrete large amounts of inflammatory factors. In the present study, the expression of IL‐6 and TNF‐α proinflammatory factors was significantly increased in the iWAT of HFD‐fed mice, while exosome treatment reduced the mRNA levels of IL‐6 and TNF‐α in obese mice to some extent (Fig. 2O ). These results suggest that BMSC‐Exos is essential in alleviating obesity, metabolic disorders, and inflammation. BMSC‐Exos improves PA‐induced lipid droplet accumulation and obesity in mature 3T3‐L1 adipocytes In order to more fully characterize the inhibitory effect of BMSC‐Exos on obesity, BMSC‐Exos was used to mature 3T3‐L1 adipocytes for relevant experiments. The different concentrations of BMSC‐Exos did not produce toxic effects on the cells (Fig. 3A ). We performed Oil Red O staining to assess the effect of BMSC‐Exos on PA‐induced lipid accumulation in 3T3‐L1 adipocytes. The results showed that PA led to adipocyte hypertrophy and massive lipid accumulation, whereas BMSC‐Exos showed a dose‐dependent alleviation of adipocyte hypertrophy and lipid accumulation caused by PA (Fig. 3B, C ). In addition, PA‐induced high expression of the adipocyte obesity genes Leptin and FABP4, while BMSC‐Exos also significantly and dose‐dependently reduced Leptin and FABP4 protein levels (Fig. 3D, E ). These results suggest that BMSC‐Exos dose‐dependently attenuated PA‐induced. Fig. 3 BMSC‐Exos improves PA‐induced lipid droplet accumulation and obesity in mature 3 T3‐L1 adipocytes. (A–E) 3T3‐L1 preadipocytes were added to MDI to differentiate them into mature adipocytes. Based on this, no or added BMSC‐Exos (10 and 20 μg·mL −1 ) were pretreated for 24 h, followed by adding PA to induce adipocytes in the obesity model. (A) CCK‐8 cytotoxicity assay. Different concentrations of BMSC‐Exos did not cause toxic effects on cells ( n = 6 per group). (B, C) Analysis of lipid droplet accumulation in each group by cellular oil red O staining ( n = 4 per group; bars = 100 μm). (D, E) Protein levels of obesity‐related genes (Leptin and FABP4) in various groups of cells ( n = 3 per group). The error bars indicate the SEM, whereas comparisons between two groups were performed by an unpaired Student's test, * P < 0. 05; ** P < 0. 01; *** P < 0. 001. Ctrl, control group; MDI, adipocyte‐induced differentiation group; MDI + PA, adipocyte‐induced differentiation group with the addition of palmitic acid; MDI + PA + 10 μg·mL −1 exosome, adipocyte‐induced differentiation group treated with palmitic acid and 10 μg·mL −1 of exosome; MDI + PA + 20 μg·mL −1 exosome, adipocyte‐induced differentiation group treated with palmitic acid and 20 μg·mL −1 of exosome. BMSC‐Exos regulates insulin sensitivity through the activation of the PI3K/AKT signaling pathway The PI3K/AKT signaling pathway plays a crucial role in IR. In order to further investigate the mechanism of BMSC‐Exos' role in obesity alleviation and IR, we have studied it accordingly. P‐PI3K, P‐AKT, and GLUT4 protein levels were significantly downregulated in the iWAT of HFD‐fed obese mice compared to NCD group mice, resulting in a blocked insulin signaling pathway and reduced glucose uptake and utilization in iWAT, which may lead to the pathological state of IR. Continuous administration of BMSC‐Exos to obese mice resulted in some upregulation of P‐PI3K, P‐AKT, and GLUT4 protein levels, resulting in improved insulin sensitivity (Fig. 4A–D ). Fig. 4 BMSC‐Exos regulates insulin sensitivity by activating the PI3K/AKT signaling pathway. (A–D) Mice were treated as described in Fig. 2, and iWAT proteins were collected for assay. Expression of each group of insulin signaling pathway‐related proteins (P‐PI3K, P‐AKT, and GLUT4) was analyzed by western blot ( n = 3 per group). (E–H) Mature 3T3‐L1 adipocytes were pretreated with BMSC‐Exos (10 and 20 μg·mL −1 ) for 24 h. Subsequently, they were treated with 1 m m PA for 24 h and incubated with 100 n m insulin for 15 min before sample collection. Expression of each group of insulin signaling pathway‐related proteins (P‐PI3K, P‐AKT, and GLUT4) was analyzed by western blot ( n = 3 per group). The error bars indicate the SEM, whereas comparisons between two groups were performed by an unpaired Student's test, * P < 0. 05; ** P < 0. 01. To investigate the effect of BMSC‐Exos on PA interference with the insulin signaling pathway in 3T3‐L1 adipocytes, we examined P‐PI3K, P‐AKT, and GLUT4 protein levels using western blot. The results showed that the expression of P‐PI3K, P‐AKT, and GLUT4 protein levels in 3T3‐L1 adipocytes was significantly enhanced by insulin alone. In contrast, PA exposure disrupted the insulin pathway, and this effect was reversed dose‐dependently by BMSC‐Exos (Fig. 4E–H ). Discussion Obesity and IR are closely related and mutually reinforcing. When the body is obese, the ability of insulin to inhibit lipolysis and reduce plasma free fatty acid (FFA) concentrations is significantly impaired, leading to an increase in the rate of lipolysis and a chronic increase in plasma FFA concentrations [ 41 ]. Lipotoxicity occurs when triglycerides and their hydrolysis products, FFAs, in the blood exceed adipose tissues' metabolic and storage capacity. Large amounts of triglycerides and FFAs are transferred to non‐adipose tissues, where ectopic deposition occurs and causes tissue damage. Lipotoxicity can lead to the dysfunction of various metabolic pathways in adipose tissue and surrounding organs (liver, muscle, heart, etc. ), resulting in the pancreatic cell the dysfunction and IR [ 42 ]. Increased blood lipid levels, changes in fatty acid metabolism, and alterations in intracellular signaling all contribute to IR in adipose tissue, muscle, and the liver. In this study, C57BL/6 mice were given a HFD for 12 weeks, and a mouse model of high‐fat obesity was successfully constructed, and the constructed obese mice showed signs of IR by GTT and ITT. Exosomes are rich in quite various bioactive substances, such as nucleic acids, proteins, lipids, amino acids, and metabolites [ 43 ]. Exosomes as essential members of intercellular communication networks, embodied in their phospholipid bimolecular structure protecting their internal bioactive components from degradation or dilution to a certain extent, and also their biological functions such as transferring their bioactive substances to neighboring or distant cells through autocrine, paracrine, or telecine secretion to regulate the gene and protein expression levels of recipient cells [ 44 ]. Xu et al. [ 45 ], found that the pancreatic β‐cell‐derived exosome miR‐26a improved insulin sensitivity and protected β‐cell function. Wu et al. [ 46 ], showed that the liver‐derived exosome miR‐130a‐3p inhibits adipogenesis and thus lipid and glucose metabolism, mainly by downregulating the expression of fatty acid synthase (FASN) and PPARγ. In the present study, mouse bone marrow mesenchymal stem cell‐derived exosomes (BMSC‐Exos) were delivered to HDF‐fed mice. The results showed that BMSC‐Exos reduced body weight and iWAT accumulation and expression of obesity genes (Lpetin and FABP4) in obese mice. BMSC‐Exos also effectively alleviated systemic IR in obese mice. In addition, BMSC‐Exos was used in a PA‐induced obese 3T3‐L1 adipocyte model and effectively inhibited the accumulation of cellular lipid droplets and the expression of obesity genes. The PI3K/AKT signaling pathway is an IR‐related signaling pathway involved in various activities, including proliferation, differentiation, regulation, and glucose transport. It is also closely associated with IR‐related type 2 diabetes [ 47, 48, 49, 50 ]. The study of the PI3K/AKT signaling pathway has helped to provide insight into the mechanisms involved in IR. Glucose metabolism depends on the cellular uptake of glucose, and GLUT is a class of carrier proteins embedded in cell membranes to transport glucose and is widely distributed in various tissues. When the PI3K/AKT signaling pathway is activated, GLUT4 is transferred from the cell to the cell membrane, increasing glucose uptake and helping to alleviate the symptoms of IR [ 51 ]. There are relevant studies demonstrating the role of exosomes in obesity‐associated IR. Yu et al. [ 52 ] found that adipocyte‐derived exosome miR‐27a reduced the expression of IRS‐1 and glucose transporter protein GLUT4 in skeletal muscle cells by targeting PPARγ, suggesting that adipose tissue‐derived miR‐27a may play a vital role in the development of obesity‐induced IR in skeletal muscle. Our data suggest that iWAT in HDF‐fed obese mice shows signs of IR, as observed by a significant blockage of the PI3K/AKT signaling pathway and a substantial decrease in GLUT4 protein expression. We also observed a similar PI3K/AKT signaling pathway blockage in PA‐induced mature 3 T3‐L1 adipocytes. However, when BMSC‐Exos was applied to IR models in obese mice and mature 3 T3‐L1 adipocytes, it was found that the exosomes relieved the blocked PI3K/AKT signaling pathway and inhibited GLUT4 protein expression to a certain extent, resulting in increased glucose uptake by the cells and some relief of IR symptoms. Exosomes act as mediators to deliver content from the mother to the recipient cells, affecting human pathophysiology. Exosomes offer enormous advantages in the treatment of obesity and IR. Firstly, exosomes have low cellular immunogenicity and can avoid causing immune rejection of the organism. Secondly, exosomes contain similar content to parental cells and can act in place of parental cells. Finally, exosomes can act as a vehicle for IR by loading drugs. This project's shortcoming is that it does not provide insight into the role of specific components of BMSC‐Exos (microRNAs, proteins) in obesity‐related metabolic diseases. Nevertheless, our data suggest that exosomes derived from mouse bone marrow mesenchymal stem cells effectively alleviate obesity‐associated IR symptoms. In future, we can consider the use of exosomes to diagnose and treat obesity‐associated metabolic diseases. Conflict of interest The authors declare no conflict of interest. Author contributions Most of the experiments were performed by HS; HS and XH wrote the manuscript; XH revised the manuscript; YS, HZ, and YZ performed part of the in vitro experiments; BW and JL performed part of the in vivo experiments; YY, XY, WH, and LX provided technical assistance; HW and XL designed the experiments and provided funds. All authors read and approved the final manuscript.
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10. 1002/2211-5463. 13650
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FEBS Open Bio
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Endothelial cell‐derived extracellular vesicles induce pro‐angiogenic responses in mesenchymal stem cells
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Angiogenesis is a central component of vital biological processes such as wound healing, tissue nourishment, and development. Therefore, angiogenic activities are precisely maintained with secreted factors such as angiopoietin‐1 (Ang1), fibroblast growth factor (FGF), and vascular endothelial growth factor (VEGF). As an element of intracellular communication, extracellular vesicles (EVs)—particularly EVs of vascular origin—could have key functions in maintaining angiogenesis. However, the functions of EVs in the control of angiogenesis have not been fully studied. In this study, human umbilical vein endothelial cell line (HUVEC)‐derived small EVs (<200 nm; HU‐sEVs) were investigated as a potential pro‐angiogenic agent. Treating mesenchymal stem cells (MSCs) and mature HUVEC cells with HU‐sEVs induced their tube formation under in vitro conditions and significantly increased the expression of angiogenesis‐related genes, such as Ang1, VEGF, Flk‐1 (VEGF receptor 2), Flt‐1 (VEGF receptor 1), and vWF (von Willebrand Factor), in a dose‐dependent manner. These results indicate that HU‐sEVs take part in angiogenesis activities in physiological systems, and suggest endothelial EVs as a potential therapeutic candidate for the treatment of angiogenesis‐related diseases.
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Abbreviations AFM atomic force microscopy Ang1 Angiopoietin‐1 ATPS aqueous two‐phase system CANX calnexin DMEM Dulbecco's Modified Eagle's Medium FAME fatty acid methyl ester FBS fetal bovine serum FGF fibroblast growth factor Flk‐1 VEGF receptor 2 Flt‐1 VEGF receptor 1 GC–MS gas chromatography–mass spectrometry GO Gene Ontology HU‐Cell HUVEC HU‐sEVs human umbilical vein endothelial cell line‐derived small EVs HUVEC human umbilical vein endothelial cell MIS Microbial Identification System MSCs mesenchymal stem cells MTS 3‐(4, 5‐di‐methyl‐thiazol‐2‐yl)‐5‐(3‐carboxy‐methoxy‐phenyl)‐2‐(4‐sulfo‐phenyl)‐2H‐tetrazolium NC negative control NTA nanoparticle tracking analysis PANTHER Protein Analysis Through Evolutionary Relationships PBS phosphate‐buffered saline PC positive control PEG polyethylene glycol PSA penicillin/streptomycin/amphotericin qPCR quantitative polymerase chain reaction SEM scanning electron microscopy sEV small extracellular vesicles TSG101 tumor susceptibility gene 101 VEGF vascular endothelial growth factor vWf von Willebrand factor Angiogenesis is the process of vascular growth to provide nourishment to tissues and organs. Angiogenic activities happen throughout the lifespan of an organism, starting with the rapid expansion of the vascular system during embryonic development. After embryonic development, the frequency of angiogenic activities falls significantly and only occurs under tight restrictions in a localized manner [ 1 ], such as during the regeneration of wounds and the ovarian cycle [ 2 ], and during bone development and fracture healing [ 3 ]. An imbalance of angiogenesis is the cause for many pathological conditions. Chronic wounds encountered frequently in diabetes patients are primarily caused by the inadequate formation of neovascularization and dysregulation in angiogenesis activities due to lasting inflammation in endothelial cells and high glucose levels in blood [ 4, 5 ]. This disruption prevents the adequate transportation of micronutrients and oxygen to the wound site, which prevents diabetic wounds from healing [ 6 ]. Angiogenesis is tightly controlled by several molecular signals. VEGF (vascular growth factor also known as VEGFA) and corresponding receptor proteins such as VEGFR1 (VEGF receptor 1, Flt‐1) and VEGFR2 (VEGF receptor 2, flk‐1) carry out vascular endothelial cell proliferation and angiogenesis [ 7, 8 ]. Angiopoietin‐1 (Ang‐1) also displays a crucial function in angiogenesis. It has been demonstrated that a lack of Ang‐1 causes vascular deficiency in mice [ 9 ]. In addition, von Willebrand factor (vWF) is a multifunctional glycoprotein, which has a crucial role in homeostasis and regulates angiogenesis [ 10, 11 ]. Cells producing these signals may secrete them packed in extracellular vesicles (EVs). Extracellular vesicles are lipid‐bound vesicles secreted by all cell types. Their primary function is to carry cellular signals and cargo such as proteins, nucleic acids, and other molecules between cells [ 12 ]. In addition to transporting molecules between cells, the EV itself may initiate responses at cells that interact with it by binding to cell membrane proteins with its own membrane proteins [ 13 ]. Uniquely, the EVs may transfer membrane proteins between cells by fusing their membrane to the recipient cell, effectively joining the vesicle membrane with the cellular membrane of its target [ 14 ]. Since their initial discovery, studies have discovered that EVs take an active part in almost all cellular functions. EVs partake in biological functions such as differentiation [ 15 ], immune response [ 16 ], angiogenesis [ 17 ], and pathological functions such as tumorigenesis [ 18 ], and immune avoidance [ 19 ]. Due to the critical functions of angiogenesis, the discovery of effective pro‐angiogenic agents is crucial and necessary. In this study, human umbilical vein endothelial cell (HUVEC)‐derived small EV (<200 nm) (human umbilical vein endothelial cell line‐derived small EVs, HU‐sEVs) were used as a potential pro‐angiogenic agent. HU‐sEVs were isolated, and their physical, proteomic, and lipidomic profiles were characterized. The effects of HU‐sEVs on the cellular viability, tube formation capacity, and the expression of angiogenesis‐related genes were evaluated with an in vitro study and tube formation assays. In addition, sEVs were chosen in this study because they are smaller than 200 nm and are easier to use in tissue engineering. Material and methods Cell culture conditions HUVEC (ATCC CRL‐1730) and mesenchymal stem cell (MSC; #PT‐5025, Poietics™ Human Dental Pulp Stem Cells) cells were used for this study. HUVEC cells were cultured in Dulbecco's Modified Eagle's Medium (DMEM, #41966‐029, Invitrogen, Gibco, UK) supplemented with 10% fetal bovine serum (FBS, #10500‐064, Invitrogen, Gibco, UK) and 1% penicillin/streptomycin/amphotericin (PSA, Invitrogen, Gibco, UK). MSC cells were cultured according to the manufacturer's instruction a complete media consisting of Basal Medium (#PT‐4927) with supplements (SingleQuot Kit, #PT‐4514). Incubation conditions of the cells were at 37 °C in a humidified atmosphere with 5% CO 2. Media collection HUVEC cells were cultured at T‐150 cell culture flasks with Dulbecco's Modified Eagle's Medium (DMEM) containing 10% FBS and 1% penicillin–streptomycin–amphotericin until they reached 80% confluence. Media was discarded, and the cells were washed with phosphate‐buffered saline (PBS) to remove EVs of FBS origin. Cells were then cultured in 30 mL of serum‐free DMEM for 18 h. Media was collected, and cells were given complete media to prevent serum deprivation. Cells were allowed to recover for at least 2 days before another collection. Conditioned media were stored at 20 °C before further use. Isolation of HUVEC‐EVs Human umbilical vein endothelial cell line‐derived small EVs were isolated via aqueous two‐phase system (ATPS) [ 19 ] isolation and density cushion ultracentrifugation [ 20 ]. ATPS were used in in vitro studies, while analytical studies that could have been affected by ATPS polymers were conducted with density cushion ultracentrifugation. For both methods, collected media were first cleared from contaminants including dead cells and cellular debris using differential centrifugation, 300 g and 2000 g for 10 min, and 20 000 g for 30 min, respectively, and were then filtered through a 0. 22 μm filter for further removal. For ATPS isolation, resulting supernatants were then mixed at a 1 : 1 v/v ratio with ATPS isolation solution, consisting of 3. 35 w/v polyethylene glycol (PEG) (25–45 kDa, Sigma, #81310, St. Louis, MO, USA) and 1. 65 w/v dextran (450–650 kDa, Sigma, #81392) in PBS. Two washing solutions were also prepared for each sample, which are 1 : 1 dilution of the ATPS isolation solution with PBS. Samples and their washing solutions were centrifuged at 1000 g for 10 min for phase separation. Upper phase of the samples, which equates to 80% of their total volume, was then removed and replaced with the upper phase of a washing solution, and was centrifuged at 1000 g for 10 min for phase separation. The washing procedure was repeated again for each sample with the second set of washing solutions. After the final centrifugation, EV containing bottom phases of the samples (which equate to the 10% of the starting volumes) were extracted. For density cushion ultracentrifugation, 10 mL of contaminate‐free plasma samples was layered above 1. 5 mL of 1 m sucrose solution in a 12. 5 mL ultracentrifugation tube. Samples were then centrifuged at 100 000 g for 80 min using a SW 40i ultracentrifugation rotor (Beckman Coulter, Brea, CA, USA). After the centrifugation, top layer was removed, and 1 mL of the sucrose layer was collected from the bottom to ensure the EV‐containing phase remained unmixed with the contaminants of the upper phase. Scanning electron microscopy Human umbilical vein endothelial cell line‐derived small EVs samples were imagined with a scanning electron microscope to determine their size distribution and morphology. Samples were diluted 1 : 100 and dried on glass microscope slides. Samples were then washed away from excess dextran by dropping 100 μL of chilled methanol to the center of the dried sample, and letting it dry. Samples were coated with gold with a sputter coater (BAL‐TEC SCD 005, Switzerland) and imagine with SEM Zeiss EVO 40 (Jena, Germany). Nanoparticle tracking analysis Size distribution of HU‐sEV and their quantification were done using nanoparticle tracking analysis (Nanosight NS300, Malvern Instruments, Malvern, UK). Samples were diluted to fit the concentrations suggested by the manufacturer. Video capture was done with the low‐volume sample chamber for 60 s at camera level 16. Chamber was flushed with distilled water between each capture. A total of five captures were taken for each sample. Data analysis was done using nanoparticle tracking analysis ( nta ) software version 3. 4. Characterization of HU‐sEV surface markers using flow cytometry Surface markers of HU‐sEV were analyzed using bead‐assisted flow cytometry. HU‐sEV were adhered onto aldehyde/sulfate latex beads (4% w/v, 4 μ m, Thermo Fisher, A37304, Waltham, MA, USA). HU‐sEV were incubated with the beads for 15 min at RT on a shaker for proper binding. Bound HU‐sEV were diluted with 200 μL of PBS with 2% BSA to block non‐specific antibody binding. Glycine (Merck, Darmstadt, Germany) was added to the solution to reach 100 m m concentration, incubating for 30 min at RT shaking. Samples were then diluted with 800 μL of cold PBS and centrifuged for 2700 g for 3 min to pellet the HU‐sEV. Resulting pellet was resuspended with 500 μL of PBS and aliquoted into 100 μL in test tubes and incubated with corresponding antibodies. Markers for CD9 (Biolegend, 124808, San Diego, CA, USA), CD63 (Biolegend 143904), CD81 (Biolegend, 349506), and HSP70 (Biolegend 648004) were at a 1 : 1000 dilution and incubated overnight. Primary antibodies of Alix (Abcam, ab186429, Cambridge, UK), tumor susceptibility gene 101 (TSG101) (Abcam, ab209927), and calnexin (CANX) (Abcam, ab203439) were incubated overnight, centrifuged at 2700 g for 3 min to wash the samples of excess antibodies and then resuspended in 100 μL of 1 : 100 dilution Alexa Fluor 488 (Abcam, ab150077). Analysis of EVs was done with Becton Dickinson (BD) FACS Calibur Flow Cytometry System (Becton Dickinson, San Jose, CA, USA). Capillary western blot Protein expression profile of HU‐sEV was shown by capillary western blot (Wes, Protein Simple; San Jose, CA, USA). Experimental procedure was carried out according to the manufacturer's instructions. Cell lysate was used as a positive control (PC) to test the functionality of the antibodies used. One to 2 μg total protein was added from cell lysate or Hu‐sEV to the capillary cartridges (12–230 kDa Wes Separation Module 8 × 13 capillary cartridges, Cat#SM‐W002 and 2–40 kDa Wes Separation Module 8 × 13 capillary cartridges, Cat#SM‐W009) to each capillary. Using the Wes system, proteins that correspond with the primary antibodies of Flotillin 1 (D2V7) (1 : 50, Cat#18634), Alix (3A9) (1 : 50, Cat#2171), GM130 (D6B1) (1 : 50, Cat#12480) (Cell Signalling Technology; Denver, MA, USA), CD9 (1 : 50, Cat#10626D), CD81 (1 : 50, Cat#10630D) (Thermo Fisher Scientific; USA) and HRP conjugated Serum Albumin (1 : 50, Cat#ab18193) (Abcam) were detected automatically. Secondary antibodies used were Anti‐Rabbit IgG, HRP‐linked (1 : 1000, Cat#7074) and Anti‐Mouse IgG, HRP‐linked (1 : 1000, Cat#7076) (Cell Signalling Technology; Denver, MA, USA). Calculation and analysis of protein expression were based on the gel‐like images produced by the Compass for sw software (Version 4. 0, Protein Simple). Atomic force microscopy analysis Visual imaging of the purified EV samples was performed with atomic power microscopy. While preparing EV samples, they were diluted 1 : 100 with 18 megohm deionized water and dried overnight on an ultra‐flat silicon surface. Subsequently, the prepared samples were scanned using an aluminum‐coated cantilever (NSC36‐B) in contact mode in the Park System XE‐100 atomic force microscopy (AFM). Topographic scanning in EV samples was taken at a scan rate of 2 Hz and a resolution of 256 × 256 pixels. The acquired images were analyzed via xei software (Park Systems, Santa Clara, CA, USA). Fatty acid methyl ester assay Fatty acid profiling of EVs purified from HUVEC cells was performed by fatty acid methyl ester (FAME) analysis. In the analysis, primarily, the total fatty acids of the EVs were isolated and then they were converted to FAME derivatives by transesterification reaction. Samples were prepared according to the manufacturer's instructions. Analysis of prepared samples and identification of fatty acids were performed with Agilent Tech GC‐Midi 6890 N device. In the last stage of the sample preparation process, the upper phase was transferred to GC vial and inserted for gas chromatography–mass spectrometry (GC–MS) analysis. Fatty acid peaks in chromatogram were identified in the Sherlock Database using the Sherlock Microbial Identification System (MIS). Proteomic profiling Proteomic profiles of HU‐sEVs were characterized using mass spectroscopy. Briefly, HU‐sEVs were resuspended in SDS‐PAGE sample buffer, and protein concentrations were measured by performing a Bradford assay (Bio‐Rad, Foster City, CA, USA), measured by a Nanodrop 1000 Spectrophotometer (Thermo Fisher). 12% SDS‐PAGE gel electrophoresis was used for protein separation. Separated proteins were precipitated and concentrated using ReadyPrep 2‐DE Cleanup Kit (Bio‐Rad) per the manufacturer's instructions. Proteins were fixed overnight in a fixation solution (40% methanol, 10% acetic acid, colloidal Coomassie Brilliant Blue G‐250). Proteins were then recovered from the gels using an in‐gel tryptic digestion kit (Thermo Fisher). Analysis of the digested peptides was performed with an nLC‐MS/MS using an Ultimate 3000 RSLC nanosystem (Dionex, Thermo Fisher) coupled with a Q Exactive mass spectrophotometer (Thermo Fisher). MS spectra were obtained with the following settings: resolution of 70. 000, scan range of 40–2000 m/z, spray voltage of 2. 3 kV, target automatic gain control of ‘AGC’ 3 × 10 6, and maximum injection time of 60 ms. The top 10 precursor ions were selected by data‐dependent acquisition for MS/MS analysis. Protein candidates were identified using proteome discoverer 2. 2 (Thermo Fisher). Resulting proteins were queried against the Uniprot/Swissprot database for identification. Cell viability assay MTS (3‐(4, 5‐di‐methyl‐thiazol‐2‐yl)‐5‐(3‐carboxy‐methoxy‐phenyl)‐2‐(4‐sulfo‐phenyl)‐2H‐tetrazolium) (#G3582, CellTiter96 AqueousOne Solution; Promega, Southampton, UK) assay was done to measure the effects of HU‐sEV on cell viability of MSC at Days 1, 5, and 10. MSCs were seeded onto 96‐well plates at a concentration of 5 × 10 3 cells/well. The next day, HU‐sEV were applied onto cells with various particle numbers including 2. 50E+09, 5. 00E+09, 10. 00E+09, 15. 00E+09, 20. 00E+09, and 25. 00E+09 np·mL −1. After each incubation time point (Days 1, 5, and 10), 10 μL MTS solution, and 100 μL PBS/glucose solution onto each well. The plate was incubated for 1 h at 37 °C. Absorbance (nm) was measured at 495 with an ELISA plate reader (Biotek, Winooski, VT). Experiments were carried out both technically and biologically with at least three replicates. MSC differentiation using HU‐sEV Mesenchymal stem cell cells were seeded on 6‐well plates at a number of 5 × 10 4 cells/well as three replicates for each group. After 24 h, morphology of the cells was checked, and the differentiation process was started. Three different concentrations of HU‐sEV (2. 50E+09, 5. 00E+09, and 10. 00E+09 np·mL −1 ) were prepared in MSC culture media to differentiate MSC cells. Apart from the treatment groups, MSC incubated in culture media was used as a negative control (NC) of differentiation, and endothelial HUVEC cells were used as a positive control of differentiation. Three different concentrations of HU‐sEV containing differentiation media were replenished every other day to MSC cells for 10 days. At the end of the 10th day, cells were collected to confirm and measure differentiation levels of HU‐sEV‐induced MSC. Quantitative polymerase chain reaction HU‐sEV‐treated MSCs were collected to total RNA isolation. Then, cDNAs were synthesized from the RNA for use as a template in PCR technique. To evaluate the levels of angiogenesis‐related markers in the HU‐sEV treated cells, quantitative polymerase chain reaction (qPCR) was done. As markers, Ang1, VEGF, Flt1, Flk1, and vWF were selected (Table 1 ). HUVEC was used as a PC. SYBR Green (Applied Biosystem, Waltham, MA, USA) was used for carrying out the reaction. iCycler qPCR system (Bio‐Rad, CFX Real Time System, Singapore) was run for analysis. Experiments were carried out both technically and biologically with at least three replicates. Table 1 qPCR primer sequences. Flt 1, VEGFR1; Flk 1, VEGFR2. Primer Sequences 5′–3′ Ang 1 CATTCTTCGCTGCCATTCTG GCACATTGCCCATGTTGAATC VEGF TTGCCTTGCTGCTCTACCTCCA GATGGCAGTAGCTGCGCTGATA Flt 1 GAGGAGGATGAGGGTGTCTATAGGT GTGATCAGCTCCAGGTTTGACTT Flk 1 GCCCTGCTGTGGTCTCACTAC CAAAGCATTGCCCATTCGAT vWF CCTTGAATCCCAGTGACCCTGA GGTTCCGAGATGTCCTCCACAT 18S RNA CGGCTACCACATCCAAGGAA GCTGGAATTACCGCGGCT Tube formation assay Tube formation assay was used as a model experiment for evaluating angiogenesis capacity of the cells under tested agent. HUVEC cells were cultured onto the 150 μL Matrigel‐coated 24‐well plate at a concentration of 1 × 10 5 cells/well. Three different particle numbers of the HU‐sEV were tested on HUVEC cells. After 7 h, tube‐like structures were qualified and quantified in the aspect of different points including tube length, total loops, and branching points with Wimasis, 2016. wimtube : Tube Formation Assay Image Analysis Solution. Release 4. 0. Similarly, to HUVECs, MSCs were seeded onto 150 μL Matrigel‐coated 24‐well plates at a concentration of 1 × 10 5 cells/well and then treated with HU‐sEV. Bioinformatics analysis Gene Ontology (GO) enrichment analysis of the HU‐sEV proteome in Fig. 4 was performed using Protein Analysis Through Evolutionary Relationships [ 21 ] (PANTHER), using the Fisher's exact test with Bonferroni correction. Fold enrichment of proteins was calculated using the Homo sapiens reference database. Percentage of identified proteins was calculated by dividing the number of proteins categorized under a particular GO term with the number of uniquely mapped IDs. Protein–protein associations of the HU‐sEV proteome were visualized using string [ 22 ] and cytoscape [ 23 ] software. Parts of the Graphical Abstract and Fig. 1 were drawn by using pictures from Servier Medical Art. Servier Medical Art by Servier is licensed under a Creative Commons Attribution 3. 0 Unported License. Fig. 1 Graphical demonstration of HU‐sEV characterization according to MISEV criteria. Statistical analysis All data for Figs 4, 5, 6, 7, 8 were statistically analyzed by one‐way ANOVA and Tukey's post hoc test using graphpad prism version 8. 0. 0 for Windows, GraphPad Software, San Diego, California USA, www. graphpad. com. The values of * P < 0. 05, ** P < 0. 01, and *** P < 0. 001 were accepted as significant. Results Physical and biochemical characterization of HU‐sEVs Extracellular vesicles isolated from HUVEC cells were characterized according to the standards presented in MISEV2018 [ 24 ] (Fig. 1 ). NTA measurements of HU‐sEVs showed a polydisperse population that was <200 nm in diameter. Particle diameters measured from the E‐SEM micrographs correlated to the NTA measurements (Fig. 2A, B ). Besides, the morphological characteristics of HU‐sEVs were analyzed by AFM (Fig. 2C ). According to the microscopic images, the sphere‐like shape of HU‐sEVs was observed favorably to inherent features of EVs. Flow cytometry of HU‐sEVs show that they positively express transmembrane proteins CD9, CD63, and CD81; and membrane binding proteins HSP70, TSG101, and Alix (Fig. 2D ). Western blots of HU‐sEVs confirmed the presence of CD9, CD81, and Alix. Western blots further show that proteins GM130 and serum albumin, which are used as a control against protein contaminants for EV isolations, were not present in HU‐sEV isolations (Fig. 2E ). Fig. 2 Characterization of HU‐sEV. (A) SEM micrographs of HU‐sEVs. (B) Size distribution of HU‐sEVs with NTA and Brownian motion image. (C) AFM image. Analysis of EV biochemical characterization of HU‐sEVs with (D) flow cytometry, and (E) Capillary Western Blotting (Wes). We performed network interaction analysis for HU‐sEV proteins (Table S1 ) using string (Fig. 3 ) and kegg (Table S2 ) softwares. According to proteomic analysis, the association diagram of these proteins with each other and with the relevant pathways was conducted (Fig. 3A ). Additionally, diagrams of enriched proteins serving in ‘Extracellular’ (Fig. 3B ) and ‘Cytoskeletal’ (Fig. 3C ) compartments were shown. The relevant information about the interactions of enriched protein groups are given in Fig. 3D, E. Besides, network statistics of these interactions also were shown in Fig. 3F. Fig. 3 Proteomic analysis of HU‐sEVs. STRING relation scheme of HU‐sEV (A) total proteins, (B) extracellular, and (C) cytoskeletal compartment‐related proteins. (D) Protein–protein associations were represented in the Edges. (E) Significant proteins, splice isoforms, or post‐translational modifications were represented as Network nodes and (F) the network stats related to these interactions were represented. Gene Ontology enrichment profiles of EV proteomes can be useful in understanding the nature and biological functions of the EV (Fig. 4 ). HU‐sEVs were compared with the human reference background, which represent the average expression levels of different protein groups (Fig. 4A ). Analyses revealed that HU‐sEVs were enriched in proteins under categories such as ‘mesenchyme migration’, ‘mesenchyme morphogenesis’, and ‘tissue migration’. Of the identified proteins, 37. 8% were related to cytoskeleton organization (Fig. 4B ). Furthermore, around 83–78% of the proteins were annotated under ‘vesicle’, ‘extracellular exosome’, ‘extracellular vesicle’, and similar categories. Molecular functions of the HU‐sEV proteins were primarily structural (Fig. 4C ). Fig. 4 Gene Ontology of HU‐sEV according to functional enrichment networks, (A) cellular component (green), (B) molecular function (red), and (C) biological process (blue). Bars represent the percentage of identified protein for each GO term, while the line graph indicates fold enrichment of said terms compared with the reference human proteome. Pie charts show the percentage of identified proteins for the terms with the highest hierarchical orders. To identify fatty acid content expressed from HUVEC cell pellet and HU‐sEV, three replicate samples were analyzed by gc‐fame software (Fig. 5 ). Fatty acid ratios obtained from HU‐Cell (HUVEC) GC‐FAME analysis were detected and the ratios are caprylic acid 16. 59%, myristic acid 7. 72%, palmitic acid 28. 78%, (7Z)‐14‐methyl‐7‐hexadecenoic acid 9. 44%, oleic acid 8. 08%, cis‐vaccenic acid 4. 58%, and stearic acid 24. 8% (Fig. 5A ). HU‐sEV fatty acid ratios are caprylic acid 16. 1%, myristic acid 12. 9%, palmitic acid 34. 2%, γ‐linolenic acid 14. 05%, and stearic acid 22. 75% (Fig. 5B ). In addition, fatty acid distribution was determined according to the saturation and other modification properties of fatty acids (Fig. 5C ). Fig. 5 Comparative FAME results. Quantitative results of fatty acid content in (A) HU‐Cell and (B) HU‐sEV samples. (C) Fatty acid ratio of HU‐Cell and HU‐sEV samples. BFA, branched fatty acid; HFA, hydroxy fatty acid; MUFA, mono‐unsaturated fatty acid; PUFA, polyunsaturated fatty acid; SFA, saturated fatty acid; UNSFA, unsaturated fatty acid. Effects of HU‐sEV on morphology and cell viability During angiogenesis, mesenchymal stem cells (MSCs) form capillary tube‐like structures. To study their pro‐angiogenic potentials, various concentrations of HU‐sEVs were added to MSC cultures, and the cultures were observed for any morphological changes over a 10‐day period. Over this duration, cells in groups treated with HU‐sEVs, and not the control groups, exhibited morphological changes that lead to the formation of circular intercellular spaces reminiscent of tubes. The frequency and degree of the morphological changes were dose dependent of HU‐sEVs, and increased with increasing concentrations (Fig. 6A ). A cell viability experiment was performed to determine the nontoxic doses of HU‐sEVs to be used in the differentiation assay. HU‐sEVs did not display any cytotoxic or proliferative effects on MSCs for the tested concentrations (Fig. 6B ). Fig. 6 Effect of HU‐sEV on MSC cell morphology and cell viability. (A) HU‐sEV‐treated cells formed a round structure as the dose increased and gained tubular shape, qualitative analysis evaluated by bright field 5× magnification light microscope images. (B) Graphical demonstration of cell viability of HU‐sEV‐treated MSCs in a dose‐dependent manner at 1st, 5th, and 10th days. Values are reported as the means ± SD. HU‐sEV can induce tube formation capacities of MSCs and HUVEC Tube formation assay was used to determine the effects of different concentrations of the HU‐sEVs on angiogenesis capacities of the HUVECs and MSCs (Fig. 7 and Fig. S1 ). All used concentrations of the HU‐sEVs significantly increased the tube lengths (Fig. 7A, B ) and total loop number (Fig. 7A, C ) compared with the control group. These effects were observed in a concentration‐dependent manner, where the highest concentration of the HU‐sEV gave the best result. When the HU‐sEV were used on the MSCs in tube formation assay, only the highest concentration of the HU‐sEV displayed significantly positive effect in terms of tube lengths (Fig. 7A, B ) and total loops (Fig. 7A, C ) compared with the control group. Fig. 7 Tube formation of endothelial cells induced by three different concentrations of HU‐sEV for 7 h. (A) qualitative analysis evaluated by bright field 10× magnification light microscope images and quantitative analysis conducted by using Wimasis wimtube software. (B) Tube length and (C) loop values are also obtained from wimtube Software. The data were statistically analyzed using one‐way ANOVA with Tukey post hoc test. The data were presented as mean values ± SD. n = 3. ** P < 0. 01 and *** P < 0. 001. HU‐sEVs induce endothelial differentiation of MSCs At the end of the differentiation protocol, 2. 50E+09, 5. 00E+09, and 10. 00E+09 np·μL −1 of HU‐sEV treated groups, non‐treated MSC and HUVEC cells were collected from culture. The fatty acid profile of HU‐sEV treated MSCs and the expression levels of angiogenesis‐related genes such as Ang1, VEGF, Flt1, Flk1, and vWF were evaluated (Fig. 8 ). Non‐treated MSCs were used as a NC group, and HUVEC were used as a positive control group. In the results, it was observed that the fatty acid composition of MSCs altered noticeably in the result of HU‐sEVs treatment (Fig. 8A ). Accordingly, newly emerged fatty acids like oleic acid, cis‐vaccenic acid, and arachidonic acid were observed in the fatty acid composition of EV‐treated MSCs. Besides, these increasing fatty acids were not observed in the fatty acid profile of HUVECs. Furthermore, caprylic acid, one of the fatty acids in HUVECs, could not be observed in the EV‐treated MSCs. On the contrary, as a result of HU‐sEV treatment, the percentages of myristic acid and γ‐linolenic acid decreased in the MSCs and so began resembling the fatty acid composition of HUVECs. In the gene expression analyses, all data were calculated and expressed as the fold change according to the NC group. While the highest expression level of the Ang1 gene was observed in the PC group (37. 85 ± 3. 66), its expression level was significantly increased in the 2. 50E+09, 5. 00E+09, and 10. 00E+09 np·μL −1 groups compared with the NC group to 6. 83 ± 4. 65, 8. 32 ± 4. 4, and 22. 3 ± 2. 4, respectively. Besides, administration of the 2. 50E+09 (2. 84 ± 0. 23), 5. 00E+09 (3. 29 ± 0. 25), and 10. 00E+09 (3. 65 ± 0. 25) np·μL −1 HU‐sEV upregulated VEGF levels compared with the PC (1. 71 ± 0. 15) and NC groups. Flk‐1 expression levels were significantly increased in the PC group (3. 67 ± 0. 17) compared with all other groups and just the highest particle number of HU‐sEV (1. 75 ± 0. 15) caused significant upregulation of the Flk‐1 expression compared with NC group. Flt‐1 expression levels were markedly upregulated, when the cells were exposed to the 2. 50E+09 (70. 89 ± 9. 23), 5. 00E+09 (97. 85 ± 12. 26), and 10. 00E+09 (103. 6 ± 25. 7) np·μL −1 HU‐sEV compared with the PC (42. 85 ± 11. 57) and NC group. 5. 00E+09 and 10. 00E+09 np·μL −1 HU‐sEV caused overexpression of the vWF to 4. 74 ± 0. 54 and 9. 90 ± 3. 02, respectively, compared with the PC (2. 54 ± 1. 72) and NC groups (Fig. 8B ). Fig. 8 Effects of different dose of HU‐sEV treatment on (A) fatty acid profile and (B) gene expression levels of angiogenesis markers in MSCs as non‐differentiated negative control group and HUVEC as angiogenic positive control group * P < 0. 05. (HUVEC, human embryonic vascular endothelial cells as positive vascular control; MSC, mesenchymal stem cells as untreated control). The data were statistically analyzed using one‐way ANOVA with Tukey post hoc test. The data were presented as mean values ± SEM. n = 3. Discussion Continuation of normal angiogenic activities is necessary for the health of individuals. In healthy individuals, a balance of inducing and suppressing factors keeps angiogenic activities at an equilibrium [ 25 ]. Failure in maintaining this balance can lead to fatal complications such as cerebral ischemia—the absence of nutrients and oxygen through the blood to the brain tissue [ 26 ]. Induction of angiogenesis in the infarct region acts as a natural defense mechanism as a result of the transport of nutrients and oxygen to the ischemic tissue [ 27 ]. It was determined that an increased density of capillaries in the brain was a factor that prolonged the life span of ischemia patients correlated with the enhancement in angiogenesis activities [ 28 ]. In a 2016 study, the preventive application of a herbal mixture called MLC901 before ischemic damage to mice improved their survival rates by inducing angiogenesis, leading to reductions in infarct areas [ 29 ], with other in vivo studies reporting similar results [ 29, 30 ]. Thus, it is clear that the increment of angiogenesis activities after the stroke leads to positive results in postischemia recovery. Maintaining angiogenic activities without interruptions or alterations is necessary for the healthy being of individuals. Therefore, the determination of whether the angiogenesis activities occurring are normal or pathological is of the utmost importance. A balance of inducing and suppressing factors keeps angiogenic activities at an equilibrium [ 31 ]. Failures in maintaining this balance can lead to fatal complications such as cerebral ischemia, characterized by the absence of nutrients and oxygen through the blood to the brain tissue [ 32 ]. Therefore, it is very important to regulate angiogenesis positively at the beginning of the healing process of patients who recover from an ischemic crisis. Induction of angiogenesis in the infarct region acts as a natural defense mechanism as a result of the transport of nutrients and oxygen to the ischemic tissue [ 27 ]. Besides, it was determined that an increased density of capillaries in the brain was a factor that prolonged the life span of ischemia patients correlated with the enhancement in angiogenesis activities [ 28 ]. In a 2016 study, the herbal mixture called MLC901 was applied as preventive purposes for 5 weeks before ischemic damage to mice [ 29 ]. In the recovery process after the ischemic damage, it was observed that survival rates of MLC901‐treated mice increased and the cerebral infarct regions decreased. Also, it has been determined that herbal treatment increases the proliferation of brain endothelial cells in the infarct area and effectuates the neovascular structures. Also, it was determined that similar results were obtained in other in vivo studies [ 29, 30 ]. Thus, it is clear that the increment of angiogenesis activities after the stroke leads to positive results in postischemia recovery. However, when this equilibrium state slips on one side, angiogenesis‐related diseases begin to emerge. The most known examples of these diseases are ischemia, chronic inflammation, and cancer. These sorts of diseases are usually treated by inducing neovascular restoration due to the decrease in angiogenesis activities that result from the downregulation of pro‐angiogenic factors [ 31, 32 ]. Besides, a new vessel formation is an important point that has to be solved for tissue engineering applications. Even if all procedures necessary for regeneration are successfully carried out, an incomplete angiogenesis—and thus inability to provide adequate oxygen and nutrients—may cause a failure in tissue regeneration. Small EVs (sEV) are nano‐sized vesicles secreted by all cell types that carry cellular cargo between cells. Composition of an EVs cargo depends on their cell‐of‐origin and the current physiological state of the cell, allowing them to take part in a variety of physiological functions [ 33 ]. Their unique nature has led to a field of rapidly expanding research, employing EVs as potential therapeutics [ 34 ], diagnosis of diseases [ 35 ], and delivery of drugs and nucleic acids [ 36 ]. HU‐sEV were positively characterized according to their surface markers. Positive presence of transmembrane proteins such as CD9, CD63, and CD81 supports the presence of a lipid membrane, while Alix, TSG101, and HSP70 supports the presence cytoplasmic cargo, which in tandem supports the presence of intact EVs in our samples. Western blots of CD9, CD81, and flotillin confirmed the presence of these markers in HU‐sEVs, while the negative GM130 and serum albumin results established that HU‐sEVs samples were not contaminated by co‐isolating proteins [ 24 ]. While researchers' study with stem cells to investigate the pro‐differentiation, pro‐angiogenic or pro‐migratory effects of any agent, the first thing to be done is the determination of sub‐lethal doses of the agent on the cells [ 37 ]. We observed that all applied groups of the HU‐sEV did not cause any significant effect on cell viability. Thus, the three lowest doses of the HU‐sEV were chosen to evaluate the pro‐angiogenic effects of them in a dose‐dependent manner. Besides, during endothelial differentiation of the MSCs, the cells gain capillary tube‐like structure and intercellular spaces occur. Administrations of sEV also caused formation of this structure in a dose‐dependent manner. Gene Ontology enrichment of HU‐sEVs revealed a high degree of enrichment in proteins related to tube formation. The majority of HU‐sEVs proteins had functions in cytoskeleton organization, and the presence of proteins related to biological processes related to tube formation, such as migration or morphogenesis, were enriched >100 and 43. 6‐fold, respectively, compared with the reference background. HU‐sEVs may induce angiogenesis and tube formation in vivo through these proteins. This is further supported by the fact that the majority of identified proteins were annotated as showing structural molecular activity. As expected, a large percentage of the identified proteins was annotated as of EV origin. Proteins that would be present if the EV population was contaminated by cell death, such as of nuclear or mitochondrial origin, were absent in the EV proteome. When we compare the fatty acid contents of the cell and the exosome, the ratios of saturated fatty acid and unsaturated fatty acid do not differ significantly, but the ratios in the content of FAs change with an individual examination. While FAs with polyunsaturated fatty acid properties, which increase fluidity in particular, disappear in the exosome, the proportion of FAs that provide fluidity does not change with the emergence of branched fatty acids in the exosome when they are not present in the cell. Although there is no exosomal FAME analysis data that we can compare in the literature, the data obtained from our fatty acid analyzes have shown that the lipid profiles of exosomes are more saturated and more stable. Although this feature gives a more robust structure to exosomes, it has reduced its use as a drug‐loading system. According to the general profile of HU‐sEV, fluidity did not decrease and the saturated FAs did not increase, providing it with a more flexible structure. Although the decrease in fluidity reduces the half‐life of the EV, it has increased the possibility of loading drugs, nucleic acids, and similar substances into the EV. And this feature has given the EV the ability to be used as a carrier system. Tube formation assay is a model experiment for evaluating angiogenesis capacity of the cells [ 38 ]. HUVECs, which are model cells for angiogenesis studies, were used. In addition, MSCs have angiogenesis capacity [ 39 ]. Matrigel has a unique chemical composition to induce angiogenesis of the cells and measurements of the total length and loop indicate angiogenesis capacity of the cells under certain conditions like sEV administration [ 40 ] in Matrigel assay. We observed that while angiogenesis capacity of the HUVEC is more than MSCs according to integrity of the tube structure, sEV treatment displayed different effects onto the cells. All particle numbers of the sEV increased total tube length and loop in HUVEC; however, only the highest particle number of the sEV caused an increase in MSCs. When evaluating the differentiation capacity of cells, changes in the cell membrane of MSC cells can be a criterion for angiogenicity. The change in the FA ratios formed on the cell membrane by different doses of HU‐sEVs given on MSCs gives information about the differentiation of cells. In particular, the presence of oleic and cis‐vaccenic unsaturated fatty acids in MSCs treated with Hu‐sEV increases the fluidity of the cell membrane, and this feature facilitates the cells to gain vascularization properties and to change the cell structure more easily. To trigger the vascular maturation in vitro, VEGF and Ang‐1 exhibit positive roles [ 41 ]. Besides, VEGF has interactions such as Flt1 and Flk1 [ 42, 43, 44 ]. The von Willebrand factor's main role is hemostasis, which plays a critical role in blood vessel formation [ 10, 11 ]. The highest particle numbers of the HU‐sEV increased all these critical gene expressions, while other applications caused partial changes in the related gene expressions. HUVECs displayed a pro‐angiogenic effect in a dose‐dependent manner in the aspect of the gene expression levels. In conclusion, small EVs isolated from HUVEC cells induce pro‐angiogenic responses in MSCs. Treatment with HU‐sEVs increased the viability and tube‐forming potential of MSCs. These could be the result of the HU‐sEVs inducing the increased expression of pro‐angiogenic genes in treated MSCs. The pro‐angiogenic activity of the sEV was also confirmed on the HUVECs in a tube formation assay. According to these results, HU‐sEVs are promising candidates for the treatment of vascular diseases and wound healing studies. However, to reinforce this claim, the efficacy of the HU‐sEVs should also be tested in vivo. As a future aspect, increasing the neovascularization potential of these EVs and triggering the cells in the environment for vascularization will lead to the use of these EVs in many fields, especially tissue engineering. In tissue engineering‐oriented studies, when it is desired to repair the damaged tissue or to create a new 3D tissue, even if the target tissue can be created, the inability to induce vascularization causes the tissue to deteriorate because the tissue cannot be fed properly. In addition, HU‐EVs have high biocompatibility and biodegradable properties that can be used both in local applications and in systems that can be mixed with scaffold and can be used as cell‐free therapy. Conflict of interest The authors declare no conflict of interest. Author contributions PNT, HA, and FŞ performed the experimental design of this study and the construction and analysis of the experiments. HA, PNT, OKK, and BTB performed cell culture, media collection, exosome isolation, and characterization. HA, OKK, BTB, TBH, and EAA performed molecular experiments and tube formation assay. PNT, EAA, and TBH structured the preparation of the figures. BTB, EAA, FŞ, HA, and OKK performed grammar correction and final writing. All authors read and approved the final manuscript. Supporting information Figure S1. Tube formation of MSC induced by three different concentrations of HU‐sEV for 7 h. (a) qualitative analysis evaluated by bright field 10X magnification light microscope images and quantitative analysis conducted by using Wimasis WimTube software. (b) Tube length and (c) loop values are also obtained from WimTube Software. The data were statistically analyzed using one‐way ANOVA with Tukey post hoc test. The data were presented as mean values ± SD. n = 3. *p < 0. 05 and ** p < 0. 01. Table S1. HU‐sEV proteins Table S2. KEGG software analysis.
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10. 1002/acm2. 12489
| 2,018
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Journal of Applied Clinical Medical Physics
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A dosimetric study on the use of 3D‐printed customized boluses in photon therapy: A hydrogel and silica gel study
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Abstract Purpose The aim of the study was to compare the dose differences between two kinds of materials (silica gel and hydrogel) used to prepare boluses based on three‐dimensional (3D) printing technologies and commercial bolus in head phantoms simulating nose, ear, and parotid gland radiotherapy. Methods and materials We used 3D printing technology to make silica gel and hydrogel boluses. To evaluate the clinical feasibility, intensity modulated radiation therapy (IMRT) plans were created for head phantoms that were bolus‐free or had a commercial bolus, a silica gel bolus, or a hydrogel bolus. Dosimetry differences were compared in simulating nose, ear, and parotid gland radiotherapy separately. Results The air gaps were smaller in the silica gel and hydrogel bolus than the commercial one. In nose plans, it was shown that the V 95% (relative volume that is covered by at least 95% of the prescription dose) of the silica gel (99. 86%) and hydrogel (99. 95%) bolus were better than the commercial one (98. 39%) and bolus‐free (87. 52%). Similarly, the homogeneity index (HI) and conformity index (CI) of the silica gel (0. 06; 0. 79) and hydrogel (0. 058; 0. 80) bolus were better than the commercial one (0. 094; 0. 72) and bolus‐free (0. 59; 0. 53). The parameters of results (HI, CI, V 95% ) were also better in 3D printing boluses than in the commercial bolus or without bolus in ear and parotid plans. Conclusions Silica gel and hydrogel boluses were not only good for fit and a high level of comfort and repeatability, but also had better parameters in IMRT plans. They could replace the commercial bolus for clinical use.
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1 INTRODUCTION In the radiotherapy of skin carcinomas, parotid gland carcinomas, or chest wall after mastectomy in breast carcinomas, there is a deficiency in the dose of the skin or superficial tissue due to the dose build‐up effect of the high energy X (γ) ray. The bolus can effectively improve the skin dose from 10% to 40% to nearly 100% in 6 MV X‐ray radiotherapy. 1 The commercially available flat‐form boluses (hereafter, “commercial boluses”) are commonly used in the clinic, but they are difficult to make full contact with irregularly shaped patient skin. This lack of contact can result in the presence of air cavities between the bolus and the patient skin, which can compromise the accuracy of the surface dose. 1, 2, 3 Conventional boluses include wet gauze, Vaseline, Truwax, Polyflex colloid, Jeltrate Plus, thermoplastic film, among others. 4 They can be made into a suitable shape as needed to partially solve the air gap problem, but the existence of defects such as rough process, poor tissue uniformity, poor repeatability, and being easily moved, leads to a great deal of uncertainty in the actual radiation dose of superficial lesions, thus affecting the treatment efficacy. However, the use of three‐dimensional (3D) printing technology could help to create a patient specific bolus which facilitates correspondence with the patient skin, yielding agreement between the planned and delivered doses, but few studies have evaluated the silica gel and hydrogel bolus based on 3D‐printed technology in photon radiation therapy to our knowledge. In this study, the bolus shells were fabricated by 3D printing technology and silica gel and hydrogel were selected as filling materials. The detailed physical characteristics and clinical feasibility of silica gel and hydrogel bolus were compared in simulating nose, ear, and parotid gland radiotherapy separately. 2 MATERIALS AND METHODS 2. A Bolus materials The commercial bolus was Bolx‐I (CIVCO Medical Solution, Orange City, FL, USA) with a density of 1. 03 g/cm 3 and a main component of polymer gel. The boluses based on 3D printing technology were composed of silica gel or hydrogel. Silica gel was based on α, ω‐dihydroxypolysiloxane, ethyl polysilicate as the cross‐linking agent, and chloroplatinic acid as catalyst and synthesized at room temperature. The monomer of hydrogel was methacrylic acid and was initiated by ultrasonic polymerization. 2. B Physical property verification Two kinds of boluses were placed on the standard solid water phantom computed tomography (CT) scanning was performed and then exported to a treatment planning system (Pinnacle version 9. 8). Subsequently, the mean density of the bolus was measured randomly by 10 points. To evaluate the physical characteristics of the boluses, percentage depth dose (PDD) curves of two kinds of different materials were measured with a standard solid water phantom. The photon beam energy was set to 6 MV (Elekta Synergy) and the prescribed dose to the reference point was set to 100 MU with a 10 × 10 cm field, source to surface distance of 100 cm. The dose distributions and PDD curves were calculated using the treatment planning system (TPS) for each plan obtained from the physical evaluation with the water‐equivalent phantom. The calculated doses, D 0. 5cm, D 1cm, D 1. 5cm, D 1. 8cm, D 2. 0cm, D 2. 5cm, D 3. 5cm, and D 5. 5cm were compared for each case. 2. C Bolus fabrication The head phantom fitted with thermoplastic underwent CT scans (GE Healthcare, Waukesha, WI, USA) with a slice thickness of 1. 25 mm. The images were transferred to the pinnacle treatment planning system (Philips, Amsterdam, the Netherlands). We outlined the CTV nose, CTV ear and CTV parotid, separately. Then expand 0. 5 cm margin to PTV which avoids exceeding the patient surface. In the plan optimization, PTV was used as objective structure and some ring around the PTV were made to constrain the dose outside PTV. Target volume of 95% is normalized to the prescription dose after optimization. These CT sets without a bolus were used to create the virtual bolus structure according to the PTV in the TPS [Fig. 1 (a)], and the 3D Slicer version 4 software was used to extract bolus 3D point cloud data. Creo software (PTC, Boston, MA, USA) was used to reconstruct the 3D information to generate bolus images, which was converted and saved as a stereo lithography (STL) file, a commonly used file format in 3D printing. The data was further optimized and designed using Magics software (Materialise, Leuven, Belgium) [Fig. 1 (b)], and the bolus shell and marker points were designed using Creosoftware and 3‐matic software (Materialise) [Fig. 1 (c)]. Then, the final STL file of the bolus shell was printed with two copies in polylactic acid (PLA) by the MakerBot Replicator II printer (MakerBot Industries LLC, Brooklyn, NY, USA) [Fig. 1 (d)], which took about 3–4 hr. The two shells were filled with hydrogel and silica gel separately. The hydrogel took about 10 min, but the silica gel needed about 10 hr at room temperature of 25°. The shells were removed by softening with a hot air gun after the solidification of hydrogel [Fig. 1 (e)] and silica gel [Fig. 1 (f)]. Figure 1 The procedure of making hydrogel and silica gel boluses based on 3D printing technology. (a) Virtual bolus structure created in the TPS, (b) the reconstructed and optimized bolus, (c) the designed bolus shell, (d) the bolus shell fabricated by 3D printing, (e) hydrogel bolus, (f) silica gel bolus. 2. D Plan evaluation The head phantom was kept in the same position and was fixed with thermoplastics in which the commercially bolus was placed outside the thermoplastic and the silica gel or hydrogel was placed inside the thermoplastic. Four sets of CT images were collected (thickness 2. 5 mm): (a) without bolus, (b) commercial bolus, (c) silica gel bolus, and (d) hydrogel bolus. For the plan comparison, the PTV was outlined in the image of bolus‐free images and fused to the other three sets of CT images to ensure the consistency of the PTV. The IMRT plans which were given the same prescription dose ( V 95% = 6000 cGy) were designed in CT images without a bolus and were copied to the other three plans. All the plans were re‐optimized under the condition that all parameters are consistent. In the plan optimization, collapsed cone convolution (CCC) algorithm was used for dose calculation and the objective value of maximum dose, uniform dose, minimum dose and minimum DVH were gave to the PTV. 0. 3 × 0. 3 cm dose calculation grid was choose in order to have sufficient sampling to determine depth of D max. We then performed an analysis and comparison of the four plans: the maximum size of the gap between skin and bolus, PTV D max, D mean, D 2%, D 50%, D 98%, homogeneity index (HI, (( D 2% − D 98% )/ D 50% )), conformity index (CI, (TV 95% × TV 95% )/( T × V 95% )), 5 and comparison of the four Dose‐volume Histogram (DVH) graphs. 3 RESULTS 3. A Physical evaluation We compared the physical properties of two types of 3D printed bolus materials: silica gel and hydrogel. The mean density of silica gel and hydrogel were 1. 15 and 1. 04 g/cm 3, respectively. The depth of the maximum dose ( d max ) for the silica gel and hydrogel boluses were 1. 650 and 1. 645 cm, respectively. There were good agreements between the PDD curves for the different fillings of 3D bolus materials (Fig. 2 ) with 10 cm thickness. The differences were less than 1. 0% compared to the PDD data at all measurement points on the PDD curves, except D 0. 5cm (the dose at a depth of 0. 5 cm). Figure 2 PDD curves at central axis comparing TPS‐calculated data for silica gel and hydrogel plans to the physical evaluation using the water‐equivalent phantom. The thicknesses of two boluses were 10 cm. 3. B Fabrication of boluses We used a 3D reconstruction program to design the bolus shells, and the 3D printing technology was used to print the shells. The shells were filled with silica gel or hydrogel and then separated after curing to obtain the boluses. For the phantom treatment plans, the silicone gel, and hydrogel boluses were close to the mold body, and the outer U cover was fixed; thus the gap was small, and the maximum gap was 2 mm [Figs. 3 (c), 3 (d), 4 (c), 4 (d)]. The commercial bolus was a square‐shaped, large area, placed outside the positioning thermoplastic mask, which was not as well fixed in the location, less repeatable, and with a larger gap; the maximum gap was almost 1 cm [Figs. 3 (b), 4 (b)]. Figure 3 Comparison of plans for simulating nasal radiotherapy with different boluses. (a) Without bolus, (b) commercial bolus, (c) silica gel bolus, (d) hydrogel bolus Figure 4 Comparison of plans for simulating ear radiotherapy with different materials. (a) Without bolus, (b) commercial bolus, (c) silica gel bolus, (d) hydrogel bolus. 3. C Results of plan comparison A head phantom was used to simulate nose, ear, and parotid tumor radiotherapy. In the absence of bolus, the dose near to the patient surface is insufficient, and the maximum dose will be increased in order to achieve a prescription dose of 95% of the target volume. In contrast, using of bolus pushes the build‐up region away from the patient, thus reducing the minimal dose to target and the 95% target volume is relatively easy to reach the prescribed dose so that the maximum dose is not required to increase (Table 1, Table 2 ). In nose treatment plans (Table 1, Fig. 3 ), the D max (71. 66 Gy) and HI (0. 59) of PTV without a bolus were much higher than other plans, while the CI (0. 63) was less than other plans, indicating that there were hot spots in the target and poor uniformity of target. When the boluses were used, the conditions were much better. We observed that the V 95% of silica gel bolus (99. 86%) and hydrogel bolus (99. 95%) were better than the commercial bolus (98. 39%) and without a bolus (87. 52%). Similarly, the HI and CI of the silica gel bolus (0. 06; 0. 79) and hydrogel bolus (0. 058; 0. 80) were better than the commercial bolus (0. 094; 0. 72) and without a bolus (0. 59; 0. 53). The bolus made using 3D printing skills had better results than the commercial boluses. The parameters of HI, CI, and V 95% of the hydrogel bolus plan were slightly better than that of the silica gel bolus. Table 1 Parameters of nose radiotherapy PTVnose D max D mean D 2% D 50% D 98% HI CI V 95% Without bolus 7166 6002 6680 6116 3060 0. 59 0. 53 87. 52 Commercial bolus 6419 6121 6328 6140 5750 0. 094 0. 72 98. 39 Silica gel bolus 6456 6121 6318 6114 5950 0. 060 0. 79 99. 86 Hydrogel bolus 6393 6106 6292 6100 5936 0. 058 0. 80 99. 95 John Wiley & Sons, Ltd Table 2 Parameters of Ear radiotherapy PTVear D max D mean D 2% D 50% D 98% HI CI V 95% Without bolus 7312 6096 6950 6254 3650 0. 53 0. 51 82. 51 Commercial bolus 6958 6135 6686 6176 4600 0. 34 0. 53 92. 20 Silica gel bolus 6335 6086 6258 6082 5876 0. 06 0. 67 99. 84 Hydrogel bolus 6103 6103 6306 6106 5800 0. 08 0. 56 99. 10 John Wiley & Sons, Ltd We could also conclude that the parameters of results (especially HI, CI, V 95% ) were better in 3D printing boluses than in commercially available boluses or without a bolus in ear and parotid plans. Two kinds of materials of 3D printing boluses were also compared in the simulated plans. In ear plans (Table 2, Fig. 4 ), the HI, CI, and V 95% of the hydrogel bolus plan were 0. 08, 0. 56, and 99. 1%, respectively. Which were much better results than those of the commercial bolus (HI = 0. 34, CI = 0. 53, V 95% = 92. 20%), but only slightly superior to the silica gel bolus (HI = 0. 06, CI = 0. 67, V 95% = 99. 84%). In the parotid plan sets (Table 3, Fig. 5 ), the hydrogel bolus (HI = 0. 07, CI = 0. 69, V 95% = 99. 56%) and silica gel bolus (HI = 0. 06, CI = 0. 67, V 95% = 99. 65%) were slightly better than the commercial bolus (HI = 0. 15, CI = 0. 69, V 95% = 96. 12%). There was not much difference between the hydrogel bolus and silica gel bolus with respect to HI, CI, and V 95% values. Table 3 Parameters of parotid radiotherapy PTVparotid D max D mean D 2% D 50% D 98% HI CI V 95% Without bolus 7022 6098 6670 6146 4700 0. 32 0. 55 88. 79 Commercial bolus 6492 6109 6370 6162 5450 0. 15 0. 69 96. 12 Silica gel bolus 6333 6092 6268 6080 5870 0. 06 0. 67 99. 65 Hydrogel bolus 6331 6075 6268 6070 5848 0. 07 0. 69 99. 56 D max and D mean represent the maximum and average values of target dose respectively; D 2%, D 50% and D 98% represent the corresponding dose of target volume of 2%, 50%, and 98% respectively; HI = ( D 2% − D 98% )/ D 50%. John Wiley & Sons, Ltd Figure 5 Comparison of plans for simulating parotid radiotherapy with different materials. (a): Without bolus, (b) commercial bolus, (c) silica gel bolus, (d) hydrogel bolus. 3. D DVH curves The DVH curves for all the plans (ear, nose, parotid) were shown in Fig. 6. In nose [Fig. 6 (b)] and parotid plans [Fig. 6 (c)], the DVH curves for the hydrogel bolus were slightly better than silica gel and both of them were significantly better than the commercial bolus. However, in the ear treatment plan, the DVH curves of silica gel were better than hydrogel, which was consistent with the parameters above. Figure 6 DVH curves for all the plans. (a) Ear plans, (b) nose plans, (c) parotid plans. 4 DISCUSSION In radiotherapy, the commercial boluses, which have different degrees of gaps because of their poor shape on the irregular surface. The effect of the air gap on surface dose reduction is related to factors such as field size, incident angle, ray energy, and patient characteristics. 4 In recent years, with the continuous progress of 3D printing technology, 3D printing skills have becoming more and more widely used in the medical field, especially in plastic surgery, oral and maxillofacial surgery, and orthopedics. 6 Application in the radiation therapy field has also gradually increased, especially in the production of boluses, most of which have been used in electron radiation therapy. 7, 8, 9, 10 Indeed, many electron treatments do not involve a planning CT and instead rely entirely on a visual “clinical setup” of the patient anatomy within the room. This simplifies the procedure of making boluses, but we need to further facilitate accurate and reproducible alignment of patient anatomy. 11 In this study, we used 3D printing skills to create individually customized boluses, which were designed to compensate for the irregular surface in photon IMRT radiotherapy. In general, two ways have been reported of making a bolus in past studies. One method was to print a bolus directly with 3D printing materials after the design stage. Polylactic acid (PLA) whose physical density was 1. 19 kg/m 3, was a commonly used printing material, which had been demonstrated to be a bolus material in a previous study. 12 Studies reported that the doses of 3D printed PLA bolus in phantom simulating radiotherapy of breast cancer after radical resection were more uniform than with the commercial bolus. 13, 14 Acrylonitrile butadiene styrene (ABS) copolymer is another printing material commonly used except PLA, but both the two materials are too hard and with poor comfort. More importantly, the different infill percentage of these two materials corresponds to different densities, which may lead to discrepancies between the calculated and measured dose distribution. 15 Another method is to print the shell of the bolus and then fill it with other soft materials. Richard et al. 16 printed the shell in PLA using the 3D printer and filled it with silicone rubber for non‐melanoma skin cancer electron beam radiotherapy. Silicone rubber has the advantage when making a bolus due to its excellent biocompatibility, chemical stability, and good mechanical properties, but its density is 1. 1–1. 2 g/cm 3 which differs from that of human tissue. In this study, we tried different filling materials, which were silica gel and hydrogel. Silica gel has the same physiochemical characteristics as silicone rubber. Hydrogels are widely used in biomedical fields, for example, scaffolds for tissue engineering, vehicles for drug delivery, actuators for optics and fluidics, and model extracellular matrices for biological studies. 17 The density of hydrogel is similar to that of human tissue, but it is friable and of poor mechanical strength compared with silica gel or silicone rubber. Therefore, this study compares the dosimetric merits and demerits of two different materials. In nose radiotherapy (Table 1 ), silica gel bolus and hydrogel bolus plans were much better than commercial bolus or without a bolus. The commercial bolus was a square with a thickness of 5 mm, which had a gap more than 6 mm when placed on an irregular surface [Figs. 3 (b), 4 (b)], but the air gaps of the 3D printed bolus were smaller, which was consistent with other studies. 3 Although the boluses had the same conformation of nose, the dose distribution of the hydrogel bolus was more uniform and was slightly superior in D max, D mean, HI, CI, and V 95% than the silica gel bolus. We know the CI value is 0–1, and the greater the value, the better conformity; the HI value reflects the uniformity of dose in the target area, the lower the HI value, the better the homogeneity. We speculated that the advantages in the nose treatment plan in hydrogel maybe due to its similar density with human tissue. However, in the ear treatment, we observed that silica gel was better than hydrogel with respect to HI, CI, and V 95%, and silica gel had smaller low‐dose areas (like 30 and 20 Gy dose areas) than hydrogel. It was possible that when the head model adopts a supine position, the hydrogel was more easily deformed and not closely jointed with the surface of the body because of gravity; behind the ear, the air gap was particularly large (Fig. 4 ). However, the hydrogel was much better than the commercially available bolus, because the commercial bolus and ear space exceed 1 cm with a poor contrast and a gravity effect (Fig. 4 ). In the parotid treatment, the plans for the two materials were quite similar and a bit better than the commercial one. The air gap between the commercial one and the surface was as big as in the ear treatment (Fig. 5 ). To our knowledge, hydrogel is flexible, odorless, biologically nontoxic, and highly transparent, but it has not been used as a bolus in radiotherapy because of its physical characteristics. Hydrogels tend to lose water and undergo deformation, which is not suitable for long‐time use. The traditional polymer hydrogel is usually formed by chemical cross‐linking. The uneven dispersion of the chemical cross‐linking agent leads to an uneven gel network, and the gel is very fragile, which greatly limits its application. 18, 19 We must solve these two problems for the clinical application of hydrogel. We used polyol polyurethane membrane to cover the hydrogel surface to prevent contact with air, thereby preventing dehydration. Because of its poor strength, many studies have reported methods to increase its strength, such as nanocomposite hydrogel 20, 21 and double‐network hydrogel. 22, 23, 24 Clinical application of strong, tough, and responsive hydrogels is the directions of future development, but more improvement is needed in biosafety and biocompatibility. Most of the novel hydrogels have strong hydrophilicity, which is not conducive to affinity with cells or biological tissues. Therefore, how to improve the biological function of hydrogel is also a problem to be overcome, or polymeric gel could be used which has been reported to have been used for bolus. 11 In addition, the 3D printing materials which currently can be directly printed are relatively hard, and if we can directly print out 3D material which meets the special requirements of a bolus, we will greatly simplify the production process and promote commercialization. The next step is to try to use new materials to print tissue bolus, such as polycaprolactone (PCL), 25 which has already been used in medical applications. 5 CONCLUSIONS In this study, we used 3D printing skills to create individually customized boluses, which were designed to compensate for the irregular surface in photon IMRT radiotherapy. The dosimetric differences of hydrogel, silica gel, and commercial boluses were compared in head phantoms simulating nose, ear, and parotid gland radiotherapy. Silica gel and hydrogel boluses were not only good for fit and a high level of comfort and repeatability, but also had better dose parameters in IMRT plans. They may replace the commercial bolus for clinical use. CONFLICT OF INTEREST The authors have no relevant conflict of interest to report.
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10. 1002/adbi. 201700056
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Advanced biosystems
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Proangiogenic Activity of Endometrial Epithelial and Stromal Cells in Response to Estradiol in Gelatin Hydrogels
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Biomaterial vascularization remains a major focus in the field of tissue engineering. Biomaterial culture of endometrial cells is described as a platform to inform the design of proangiogenic biomaterials. The endometrium undergoes rapid growth and shedding of dense vascular networks during each menstrual cycle mediated via estradiol and progesterone in vivo. Cocultures of endometrial epithelial and stromal cells encapsulated within a methacrylamide-functionalized gelatin hydrogel are employed. It is reported that proangiogenic gene expression profiles and vascular endothelial growth factor production are hormone dependent in endometrial epithelial cells, but that hormone signals have no effect on human telomerase reverse transcriptase (hTERT)-immortalized endometrial stromal cells. This study subsequently examines whether the magnitude of epithelial cell response is sufficient to induce changes in human umbilical vein endothelial cell network formation. Incorporation of endometrial stromal cells improves vessel formation, but co-culture with endometrial epithelial cells leads to a decrease in vascular formation, suggesting the need for stratified cocultures of endometrial epithelial and stromal cells with endothelial cells. Given the transience of hormonal signals within 3D biomaterials, the inclusion of sex hormone binding globulin (SHBG) to alter the bioavailability of estradiol within the hydrogel is reported, demonstrating a strategy to reduce diffusive losses via SHBG-mediated estradiol sequestration.
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No full text available
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10. 1002/adbi. 201700083
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Advanced biosystems
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Embedded Spheroids as Models of the Cancer Microenvironment
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To more accurately study the complex mechanisms behind cancer invasion, progression, and response to treatment, researchers require models that replicate both the multicellular nature and 3D stromal environment present in an in vivo tumor. Multicellular aggregates (i. e. , spheroids) embedded in an extracellular matrix mimic are a prevalent model. Recently, quantitative metrics that fully utilize the capability of spheroids are described along with conventional experiments, such as invasion into a matrix, to provide additional details and insights into the underlying cancer biology. The review begins with a discussion of the salient features of the tumor microenvironment, introduces the early work on non-embedded spheroids as tumor models, and then concentrates on the successes achieved with the study of embedded spheroids. Examples of studies include cell movement, drug response, tumor cellular heterogeneity, stromal effects, and cancer progression. Additionally, new methodologies and those borrowed from other research fields (e. g. , vascularization and tissue engineering) are highlighted that expand the capability of spheroids to aid future users in designing their cancer-related experiments. The convergence of spheroid research among the various fields catalyzes new applications and leads to a natural synergy. Finally, the review concludes with a reflection and future perspectives for cancer spheroid research.
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No full text available
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10. 1002/adbi. 201800101
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Advanced biosystems
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Stereolithographic 4D Bioprinting of Multiresponsive Architectures for Neural Engineering
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4D printing represents one of the most advanced fabrication techniques for prospective applications in tissue engineering, biomedical devices, and soft robotics, among others. In this study, a novel multiresponsive architecture is developed through stereolithography-based 4D printing, where a universal concept of stress-induced shape transformation is applied to achieve the 4D reprogramming. The light-induced graded internal stress followed by a subsequent solvent-induced relaxation, driving an autonomous and reversible change of the programmed configuration after printing, is employed and investigated in depth and details. Moreover, the fabricated construct possesses shape memory property, offering a characteristic of multiple shape change. Using this novel multiple responsive 4D technique, a proof-of-concept smart nerve guidance conduit is demonstrated on a graphene hybrid 4D construct providing outstanding multifunctional characteristics for nerve regeneration including physical guidance, chemical cues, dynamic self-entubulation, and seamless integration. By employing this fabrication technique, creating multiresponsive smart architectures, as well as demonstrating application potential, this work paves the way for truly initiation of 4D printing in various high-value research fields.
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No full text available
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10. 1002/adbi. 201900089
| 2,019
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Advanced biosystems
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Senescent Cells with Augmented Cytokine Production for Microvascular Bioengineering and Tissue Repairs
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Controlled delivery of cytokines and growth factors has been an area of intense research interest for molecular and cellular bioengineering, immunotherapy, and regenerative medicine. In this study, we show that primary human lung fibroblasts chemically induced to senescence (cell cycle arrest) can act as a living source to transiently produce factors essential for promoting vasculogenesis or angiogenesis, such as VEGF, HGF, and IL-8. Co-culture of senescent fibroblasts with HUVECs in a fibrin gel demonstrated accelerated formation and maturation of microvessel networks in as early as three days. Unlike the usage of non-senescent fibroblasts as the angiogenesis-promoting cells, this approach eliminates drawbacks related to the overproliferation of fibroblasts and the subsequent disruption of tissue architecture, integrity, or function. Co-culture of pancreatic islets with senescent fibroblasts and endothelial cells in a gel matrix maintains the viability and function of islets ex vivo for up to five days. Applying senescent fibroblasts to wound repair in vivo led to increased blood flow in a diabetic mouse model. Together, this work points to a new direction for engineering the delivery of cytokines and growth factors that promote microvascular tissue engineering and tissue repairs.
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No full text available
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10. 1002/adbi. 201900137
| 2,020
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Advanced biosystems
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Decorating 3D Printed Scaffolds with Electrospun Nanofiber Segments for Tissue Engineering
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Repairing large tissue defects often represents a great challenge in clinics due to issues regarding lack of donors, mismatched sizes, irregular shapes, and immune rejection. Three-dimensional (3D) printed scaffolds are attractive for growing cells and producing tissue constructs because of the intricate control over pore size, porosity, and geometric shape, but the lack of biomimetic surface nanotopography and limited biomolecule presenting capacity render them less efficacious in regulating cell responses. Herein, we report, for the first time, a facile method for coating 3D printed scaffolds with electrospun nanofiber segments. The surface morphology of modified 3D scaffolds changes dramatically, displaying a biomimetic nanofibrous structure, while the bulk mechanical property, pore size and porosity are not significantly compromised. The short nanofibers-decorated 3D printed scaffolds significantly promote adhesion and proliferation of pre-osteoblasts and bone marrow mesenchymal stem cells (BMSCs). Further immobilization of bone morphogenetic protein-2 (BMP-2) mimicking peptides to nanofiber segments-decorated 3D printed scaffolds show enhanced mRNA expressions of osteogenic markers Runx2, Alp, OCN, and BSP in BMSCs, indicating the enhancement of BMSCs osteogenic differentiation. Together, the combination of 3D printing and electrospinning is a promising approach to greatly expand the functions of 3D printed scaffolds and enhance the efficacy of 3D printed scaffolds for tissue engineering.
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No full text available
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10. 1002/adbi. 202000046
| 2,020
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Advanced biosystems
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Elastic Biomaterial Scaffold with Spatially Varying Adhesive Design
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In order to secure biomaterials to tissue surfaces, sutures or glues are commonly used. Of interest is the development of a biomaterial patch for applications in tissue engineering and regeneration that incorporates an adhesive component to simplify patch application and ensure sufficient adhesion. A separate region dedicated to fulfilling the specific requirements of an application such as mechanical support or tissue delivery is also desirable. Here, we present the design and fabrication of a unique patch with distinct regions for adhesion and function, resulting in a biomaterial patch resembling the Band-Aid. The adhesive region contains a novel polymer which we synthesized to incorporate a molecule capable of adhesion to tissue, dopamine. The desired polymer composition for patch development was selected based on chemical assessment and evaluation of key physical properties such as swelling and elastic modulus, which were tailored for use in soft tissue applications. The selected polymer formulation, referred to as the adhesive patch (AP) polymer, demonstrated negligible cytotoxicity and improved adhesive capability to rat cardiac tissue compared to currently used patch materials. Finally, the AP polymer was used in our patch, designed to possess distinct adhesive and non-adhesive domains, presenting a novel design for the next generation of biomaterials.
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No full text available
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10. 1002/adbi. 202000133
| 2,021
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Advanced biosystems
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Reconfigurable Microphysiological Systems for Modeling Innervation and Multitissue Interactions
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Tissue-engineered models continue to experience challenges in delivering structural specificity, nutrient delivery, and heterogenous cellular components, especially for organ-systems that require functional inputs/outputs and have high metabolic requirements, such as the heart. While soft lithography has provided a means to recapitulate complex architectures in the dish, it is plagued with a number of prohibitive shortcomings. Here, concepts from microfluidics, tissue engineering, and layer-by-layer fabrication are applied to develop reconfigurable, inexpensive microphysiological systems that facilitate discrete, 3D cell compartmentalization, and improved nutrient transport. This fabrication technique includes the use of the meniscus pinning effect, photocrosslinkable hydrogels, and a commercially available laser engraver to cut flow paths. The approach is low cost and robust in capabilities to design complex, multilayered systems with the inclusion of instrumentation for real-time manipulation or measures of cell function. In a demonstration of the technology, the hierarchal 3D microenvironment of the cardiac sympathetic nervous system is replicated. Beat rate and neurite ingrowth are assessed on-chip and quantification demonstrates that sympathetic-cardiac coculture increases spontaneous beat rate, while drug-induced increases in beating lead to greater sympathetic innervation. Importantly, these methods may be applied to other organ-systems and have promise for future applications in drug screening, discovery, and personal medicine.
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No full text available
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10. 1002/adbi. 202000168
| 2,021
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Advanced biology
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Engineering Advanced In Vitro Models of Systemic Sclerosis for Drug Discovery and Development
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Systemic sclerosis (SSc) is a complex multisystem disease with the highest case-specific mortality among all autoimmune rheumatic diseases, yet without any available curative therapy. Therefore, the development of novel therapeutic antifibrotic strategies that effectively decrease skin and organ fibrosis is needed. Existing animal models are cost-intensive, laborious and do not recapitulate the full spectrum of the disease and thus commonly fail to predict human efficacy. Advanced in vitro models, which closely mimic critical aspects of the pathology, have emerged as valuable platforms to investigate novel pharmaceutical therapies for the treatment of SSc. This review focuses on recent advancements in the development of SSc in vitro models, sheds light onto biological (e. g. , growth factors, cytokines, coculture systems), biochemical (e. g. , hypoxia, reactive oxygen species) and biophysical (e. g. , stiffness, topography, dimensionality) cues that have been utilized for the in vitro recapitulation of the SSc microenvironment, and highlights future perspectives for effective drug discovery and validation.
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1. Introduction Systemic sclerosis (SSc; also termed scleroderma) is an autoimmune disease that is characterized by the distinctive pathogenetic triad of microvascular dysfunction, dysregulation of innate and adaptive immunity, and generalized fibrosis in multiple organs. [ 1 ] SSc has the highest case-specific mortality of any of the autoimmune rheumatic diseases, with more than half of patients dying as a direct consequence of the pathology. [ 2 ] For this reason, SSc is a disease with a high unmet clinical need. Although intensive research in the last years has improved the understanding of the disease, only one drug, nintedanib, has thus far been approved by the Food and Drug Administration (FDA) for the treatment of SSc-associated interstitial lung disease. [ 3 ] Furthermore, there are no generally accepted therapies for skin and organ fibrosis, which are known to be key manifestations of SSc, leaving the need for novel antifibrotic therapeutic strategies in SSc. [ 4 ] Although significant strides have been achieved using various animal models, these systems are expensive for the purposes of routine drug development/screening studies and have limited replicability due to different physiology and genetics in comparison to the human disease. [ 5 ] These shortcomings of animal models impose the need for standardize protocols to increase reproducibility [ 6 ] and development of more reliable, clinically relevant in silico and in vitro models. [ 7 ] In vitro cell-based studies have been proven to be valuable tools in drug discovery programs, especially due to their low cost and high speed of testing compounds. Cell cultures represents an immense value in the investigation of cellular and functional aspects of disease processes for improved therapeutic interventions. [ 8 ] The most commonly utilized model systems are based on conventional 2D monolayer cultures exposed to exogenous profibrotic stimuli, commonly transforming growth factor-β (TGF-β). [ 9 ] However, customarily used cell-based models frequently fail to give predictable and reliable data for in vivo responses. A critical component of this failure results from the lack of recapitulating the native in vivo microenvironment. In the SSc scenario, the histopathological and physicochemical cues of the disease microenvironment are critical for the stimulation of biological functions mediated by cell signaling. [ 10 ] Multiple cytokines, chemokines, and signal transduction pathways are implicated in the progression of SSc as well as structural features of the extracellular matrix (ECM) such as stiffness, viscoelasticity, and topography. Moreover, although 2D culture systems are simple and economical, they do not consider the spatial organization of cells within the 3D architecture of organs and do not replicate native cell–cell and cell–ECM interactions and signaling pathways. [ 11 ] Recent advancements in 3D cell culture technologies and tissue engineering strategies have made it possible to engineer advanced physiologically relevant 3D in vitro models not only to study disease mechanism and progression, but also to use as a platform to design new therapeutic compounds and to screen for drug efficacy and safety. [ 12 ] This review summarizes the utility and limitations of various animal models of SSc and focuses on the most recent advances in in vitro SSc models, highlighting the crucial role of biological, biochemical, and biophysical cues in mimicking SSc microenvironment. The potential of bioengineered tissues as in vitro models to investigate molecular and cellular mechanisms involved in the onset and progression of systemic sclerosis and/or to serve as screening platforms to test novel pharmaceutical therapies for the treatment of the disease, will be discussed. In addition, we will shed light to the next challenges and future directions that must be addressed toward an effective 3D in vitro model for SSc. 2. Mechanisms and Pathophysiology of Systemic Sclerosis (SSc) Systemic sclerosis is a complex chronic and often progressive disease characterized by vascular alterations, inflammation and autoimmunity, and multisystemic excessive fibrosis ( Figure 1 ). Although skin fibrosis is the distinguishing hallmark, the pathological involvement of the viscera including the lungs, gastrointestinal tract, kidneys and heart determines the clinical outcome. [ 13 ] Patients are characterized by subtypes based on the extent of skin involvement, with limited cutaneous systemic sclerosis (lcSSc) and diffuse cutaneous systemic sclerosis (dcSSc) subsets delineated on the basis of distal or proximal skin involvement. [ 14 ] The pathogenesis of systemic sclerosis is complex and remains elusive. An interplay between genetic factors and environmental events, such as job-related exposures to silica dust, vinyl chloride and organic solvents or viruses and other infectious agents, is likely to play a part in the origin of the disease. [ 15 ] The onset of vascular injury in SSc includes endothelial activation and vascular damage, thickening of the vessel wall due to intimal and smooth muscle cell proliferation, and finally vessel narrowing and obliteration, which lead to tissue ischemia, oxidative stress and ultimately organ dysfunction. [ 16 ] Infiltration of inflammatory cell within the lesions is common in patients with early-stage disease, and inflammatory and immune cells are an important source of profibrotic mediators such as TGF-β, platelet-derived growth factor (PDGF), interleukin 1 (IL-1) and interleukin 6 (IL-6). Figure 2 depicts these multiple pathologic processes. Dysregulation of both innate and adaptive immunity is also a prominent factor that contributes to systemic sclerosis pathogenesis. Antinuclear antibodies are present in up to 95% of SSc patients and specific autoantibodies, such as antitopoisomerase 1, anticentromere, and anti-RNA polymerase III antibodies, directed against intracellular nuclear components, are present in over 75% of patients. [ 17 ] Besides the presence of autoantibodies, evidence of dysregulated immune responses are represented by inflammatory cells and inflammatory molecules in target tissues such as the skin and lungs and a prominent type I interferon (IFN) signature in circulating and tissue-infiltrating immune cells. [ 18 ] The pathologic hallmark of SSc is extensive fibrosis involving multiple organs, which can lead to significant organ failure. [ 19 ] Fibrosis is characterized by replacement of normal tissue architecture with rigid and mechanically stressed connective tissue rich in collagen and other ECM macromolecules, such as elastin, glycosaminoglycan, and fibronectin. [ 20 ] The abnormal accumulation of ECM results from increased synthesis by activated fibroblasts, enhanced assembly and deposition catalyzed by prolyl and lysyl-oxidase and transglutaminase 2 and aberrant ECM degradation. [ 21 ] The differentiation of fibroblasts into myofibroblasts is a critical step in the onset of fibrosis. Myofibroblasts are specialized fibroblasts that acquire characteristics of smooth muscle cells, including the expression of α-SMA. In contrast to physiological wound healing, where myofibroblasts are present only transiently within granulation tissue before undergoing apoptosis, myofibroblasts in SSc are persistent. These contractile cells secrete not only matrix proteins, but also TGF-β and other profibrotic components, and thus further promote ECM deposition and remodeling. [ 22 ] In addition to activation and proliferation of resident fibroblasts, other sources of activated fibroblasts include recruitment of circulating fibrocytes and the differentiation from epithelial cells. [ 23 ] Epithelial cells have been demonstrated to trans differentiate into fibroblasts and myofibroblasts, undergoing an epithelial to mesenchymal transition (EMT) in response to TGF-β and other growth factors and/or cytokines during the development of fibrosis. [ 24 ] In addition to epithelial cells, endothelial cells have also been shown to transdifferentiate into fibroblasts through endothelial to mesenchymal transition (EndMT). [ 25 ] Other sources of fibroblasts include trans differentiation of pericytes and adipocytes. [ 26 ] Scientific advances have considerably augmented the understanding of the pathophysiological mechanism of SSc. The antifibrotic drugs nintedanib and pirfenidone have been approved for the treatment of patients with idiopathic pulmonary fibrosis, and nintedanib recently received approval for SSc-ILD, but there are still a dearth of effective anti-fibrotic agents for the full array of SSc manifestations. [ 27 ] There is therefore an urgent unmet need to develop new anti-fibrotic therapies for use in SSc. 3. Animal Models of SSc Animals models have been extensively used to study the complex mechanisms involved in the pathogenesis of SSc and ultimately to bring new insight for the development of therapeutic strategies. Recent years have seen a plethora of genetic, transgenics and induced animal models that have contributed to our knowledge of the initiating events of systemic sclerosis ( Table 1 ). [ 28 ] Genetic animal models spontaneously develop mutations to the genome with manifestations similar to those of SSc. One of the best-characterized genetic animal models of SSc are tight skin-1 (Tsk-1) mice, in which a tandem duplication in the gene for fibrillin 1 (Fbn1), a mediator of elastic fibers assembly, is responsible for the pathogenic phenotype. In heterozygous mice, this mutation leads to thickening of the subcutaneous tissue (hypodermis) and endothelial cell apoptosis. Fibrosis in these mice develops from excessive production of ECM by activated fibroblasts upon activation of the TGF-β pathway. [ 29 ] A related genetic model of SSc is tight skin 2 (Tsk-2) mouse, which presents mutations in the gene for type III collagen alpha. [ 30 ] Tsk-2 mice demonstrate increased type I and III collagen, which lead to abnormal ECM deposition, and an inflammatory dermal mononuclear cell infiltrate. [ 31 ] Transgenic mouse models with the pathological cascade of SSc have been established to further understand the process of the disease. The transgenic mouse model overexpressing the Fos-related antigen-2 (FRA-2) gene showed many of the important factors resulting in the vascular damage and progressive skin and lung fibrosis of systemic sclerosis as well as pulmonary hypertension. [ 32 ] A TNF-transgenic model was recently shown to develop spontaneous and severe pulmonary hypertension and have genomic overlap with SSc-PAH but to lack systemic fibrosis. [ 33 ] Another model is represented by urokinase-type plasminogen activator receptor (uPAR)-deficient mice. Urokinase-type plasminogen activator receptor is a glycosylphosphatidylinositol-anchored cell surface receptor which concentrates its ligand, urokinase-type plasminogen activator (uPA), at the cell–matrix interface. The uPA/uPAR complex promotes the fibrinolysis and the degradation of other ECM, serving as a key regulator of ECM homeostasis and angiogenesis. uPAR deficient mice reproduce the fibrotic and vascular features of SSc, such as increase collagen content and perivascular inflammatory cells infiltration in skin and lungs. [ 34 ] Inducible animal models are quicker and easier to evaluate than genetic models and offer valuable clues to study the role of selected target molecules in the developmental process of SSc. The bleomycin model of fibrosis is probably the most utilized model of SSc. Bleomycin was originally isolated from the fungus streptomyces verticillus, and is often used as an anti-tumor medication for the treatment of various kinds of malignancy. Bleomycin hydrolase inactivates bleomycin by hydrolyzing the amide bond in the β-aminoalanineamide moiety. Bleomycin-induced toxicity occurs predominantly in the lungs and the skin, due to the deficiency of the enzyme in these organs. [ 35 ] For this reason, the bleomycin mouse model of fibrosis has been frequently used to replicate common features of SSc such as dermal or pulmonary fibrosis. Local dermal injections of bleomycin in mice induced collagen synthesis at the injection site over 4 weeks. The overall effects were found to be systemic because the lung similarly showed increased collagen synthesis. [ 36 ] Bleomycin can also be delivered via the intratracheal route, resulting in severe pulmonary fibrosis. [ 37 ] Although the bleomycin model replicates critical aspects observed in SSc, this model lacks the typical autoantibody patterns present in the pathology and bleomycin induced fibrosis was found to be strain specific. [ 38 ] Another inducible animal models of SSc is represented by the hypochlorous mouse model (HOCl). Repeated intradermal injections of hypochlorous acid generates hydroxyl radicals, which lead to enhanced synthesis of collagen in the lung and skin tissues. In addition, this model mimics the pathological damages observed in the systemic sclerosis kidneys and induces antitopoisomerase antibodies. [ 39 ] The mechanism of action of hypochlorous acid-induced fibrosis is not fully understood, thus restricting its commonality of use. Another model is represented by the angiotensin II-inducible model of fibrosis. Angiotensin II (Ang II) is a vasoactive peptide that induces vascular constriction, water and salt retention, and high blood pressure. [ 40 ] Subcutaneous injection of Ang II induced both inflammation and fibrosis in the skin by accumulating activated fibroblasts and promoting EndMT of circulating blood cells. However, it is not known whether these animal models developed autoantibodies specific for systemic sclerosis. [ 41 ] These animal models offer essential clues for the improved knowledge of the molecular mechanisms of SSc pathology and the identification of potential therapeutic targets for the treatment of this disease. As explained above, none of the currently available models encompasses all aspects of SSc in humans. Therefore, multiple models should be utilized when studying drug efficacy to account for deficiencies and limitations of single models, resulting however in high costs while not guaranteeing clinical translatability. In addition, although animal models predict biological relevant pharmacokinetic responses to drug administration, their different physiology and genetics from humans hamper the exact recapitulation of the human diseases. [ 28a ] There have been numerous drugs which have been successful in animal models which have not performed well in clinical trials given the complexity of SSc and imperfection of each of the models. [ 42 ] The high number of animals required during preclinical studies remain an ethical issue, besides being cost-intensive and laborious. [ 43 ] Thus, it is becoming imperative the need to develop more predictable in vitro models that can mimic aspects of human in vivo cellular behavior. 4. SSc: the Need for Advanced In Vitro Models In the last few years, progress in the understanding of the pathogenesis of SSc energized the design of numerous promising clinical studies. Several recent reviews have summarized therapies for SSc that are currently in clinical trials and shed light on novel potential therapeutic targets for the management of the disease. [ 44a, 44b, 4, 44c, 44d ] For example, tocilizumab, a humanized monoclonal antibody against the human IL-6 receptor-α, has shown encouraging results by improving both skin and lung fibrosis [ 45 ] and has reached phase III clinical trials. Another novel promising therapy is seen in trials of the endocannabinoid receptor type 2 agonist lenabasum. This synthetic molecule has emerged as a potent modulator both of skin and lung inflammation have antifibrotic potential as well. [ 46 ] Nevertheless, despite the positive signs of clinical response in subsets of patients, these two clinical trials have been unsuccessful in meeting their endpoints and failed to gain regulatory approval. [ 47, 46 ] Furthermore, considering the emerging of new potential therapeutics, along with the repurposing of existing drugs, clinical trials in SSc are more active than ever. However, the limited numbers of patients available for trials poses the need to refine pre-clinical research in order to select the optimal drug candidates with the best chance of clinical success. In light of the limitations associated with animal models, cell systems and in vitro tissue equivalents represents precious tools to investigate the disease’s molecular pathways and to generate a platform for drug screening for early-phase studies. In vitro cell-based models are an important element of the drug discovery process. In contrast to cost-intensive animal models, assays using cultured cells are simple, fast, and cost-effective as well as versatile and easily reproducible. [ 48 ] The efficacy of an in vitro model is determined by its capacity to closely replicate relevant characteristics of the in vivo microenvironment. [ 49 ] Different approaches can be utilized to develop in vitro models that recapitulate the SSc phenotype. One approach comprises the use of cells isolated from healthy donors, which are converted into a disease-specific phenotype by the addition of profibrotic modulators during the culture time to induce the expression and secretion of fibrotic markers and increase the deposition of ECM. Nevertheless, the use of exogenous stimuli does not result in a disease-activating mechanism. Due to short culture periods that do not model disease progression, cells do not acquire the full disease-specific patterns of gene expression and are fundamentally limited in representing the complexity of the disease. [ 50 ] For this reason, cells derived from SSc patients have become one of the most important materials in the study of the pathology. For example, fibroblasts derived from SSc skin lesions have been demonstrated to secrete an abnormal amount of ECM proteins (collagens, fibronectin) and fibrogenic modulators (TGF-β, CTGF) and fibrotic markers (α-SMA) in vitro. [ 51 ] Despite the tremendous utility of patient-derived cells, in vitro studies are limited by challenges including availability of patient donor cells, particularly in a rare disease such as SSc. In vitro expansion of scleroderma fibroblasts has been associated with loss of the SSc phenotype over time in culture, showing a marked decrease in collagen production in fibroblasts cultured for up to ten passages, [ 52 ] and a reduction of the disease transcriptional signature after four passages. [ 53 ] Moreover, patient-derived cells showed high heterogeneity with regard to inflammatory as well as fibrotic signatures, which can be lost during cell proliferation into any daughter lineage, leading to cell pools that do not recapitulate the variety of cells in vivo. [ 54 ] Therefore, especially given the complexity of SSc, it is imperative to achieve a system that allows for spatiotemporal control over the biological, biochemical, and biophysical cues of the in vitro extracellular microenvironment to properly mimic the pathological condition of SSc ( Figure 3 ). 5. Biological Cues 5. 1. Growth Factors Supplementation A multitude of soluble growth factors are implicated in systemic sclerosis ( Table 2 ) and TGF-β is commonly recognized as the master regulator of fibrosis. [ 55 ] TGF-β belongs to a superfamily of proteins that includes bone morphogenetic proteins (BMPs), growth differentiation factors (GDFs) activins, inhibins, myostatin, nodal and anti-Mullerian hormone (AMH) proteins. [ 56 ] There are three isoforms of TGF-β (TGF-β1, TGF-β2, and TGF-β3), which contain highly conserved regions but diverge in several amino acid regions. The three TGF-β isoforms function through the same receptor heterodimers, TGF-β receptor type 1(TGFR-1) and TGF-β receptor type 2 (TGFR2) and activate the same canonical mothers against decapentaplegic homologue (SMAD)-2–SMAD3 signaling pathway. [ 57 ] In this review we refer solely to the TGF-β family, and in particular to the TGF-β1 isoform, unless otherwise stated. The bioavailability of TGF-β is regulated by its secretion from macrophages and other cells as an inactive precursor, which is then converted to its biologically active matrix-bound latent form via integrin-mediated processes. [ 58 ] TGF-β is a master regulator of fibroblast phenotype and function. Upon TGF-β stimulation, fibroblasts are become activated and undergo phenotypic transition into myofibroblasts, which leads to excessive matrix deposition and unbalance between matrix synthesis/degradation signals. [ 59 ] Furthermore TGF-β plays an important role in the EMT of epithelial cells to myofibroblasts. [ 24, 60 ] In addition, TGF-β can play a role in the vasculopathy observed in SSc. TGF-β stimulates the expression of vascular endothelial growth factor (VEGF) and endothelin-1 (ET-1) in endothelial cells, [ 61 ] thereby mediating the vasoconstriction seen in patients with SSc. [ 62 ] Thus, the complex effects of TGF-β on proangiogenic and antiangiogenic factors partially explain the complex vascular phenotype seen in patients with SSc. The delivery of TGF-β to in vitro fibrosis platforms is a critical element in studying fibrotic mechanisms. The delivery in vitro is generally performed by simple addition of soluble TGF-β to the culture medium. For instance, studies proved that the supplementation of TGF-β to human skin fibroblasts increased the deposition of collagen type I. [ 63 ] In response to the need for an effective therapeutic for dermal fibrosis, a plethora of TGF-β stimulated in vitro models have been used as screening platform for antifibrotic molecules. As one example of many, human dermal fibroblasts stimulated with TGF-β1 to induce differentiation into profibrotic myofibroblast cells were used to assess the antifibrotic potential of Pirfenidone (PFD), a synthetic molecule already FDA-approved for the treatment of idiopathic pulmonary fibrosis. It was shown that PFD inhibited fibrogenic signals of TGF-β by abrogating p38-mediated MAPK activation, downregulating the transcription of profibrotic genes, such as type I and type III collagen, and blocking myofibroblast differentiation. [ 64 ] One of the limitations of this approach consists in the fact that the bioavailability of TGF-β in vivo is controlled by ECM-mediated integrins and mechanosensing mechanisms that inhibit or activate its binding to the corresponding receptors. Therefore, the development of culture conditions with tailored (patho-) physiological substrate stiffness could better mimic TGF-β in vivo bioavailability and may advance the relevance of its use in vitro culture models. [ 65 ] Another critical growth factor involved in the pathogenesis of SSc is connective tissue growth factor (CTGF). CTGF is a cysteine-rich matricellular protein that functions in combination with TGF-β to enhance fibrotic responses. CTGF is not normally expressed in dermal fibroblasts, but is constitutively overexpressed by fibroblasts present in skin and pulmonary fibrotic lesions of scleroderma patients. The overexpression of CTGF promotes fibroblast proliferation, myofibroblast differentiation, and matrix deposition. [ 66 ] Moreover, CTFG plays a role in leading endothelial cells to transdifferentiate toward myofibroblasts. [ 67 ] Supplementation of CTGF in the culture media was showed to stimulate fibroblastic cell proliferation and ECM synthesis. [ 68 ] A recent study demonstrated that knockdown of CTGF using a mesoporous silica nanoparticle-based small interfering RNA (siRNA) delivery system prevented collagen deposition, activation and differentiation of fibroblast. [ 69 ] Platelet-derived growth factors (PDGFs) have been demonstrated to have a critical role in fibrosis. PDGF is secreted by platelets, endothelial cells, macrophages, and fibroblasts that function as potent mitogens and chemoattractants for mesenchymal progenitor cells. [ 66a ] Elevated expression of PDGF and its receptors has been found in scleroderma skin and lung tissues, and contributes to persistent fibrosis by the generation of reactive oxygen species and consequent fibroblast activation. [ 70 ] The supplementation of PDGF to normal human dermal fibroblasts in vitro has been shown to increase the mRNA and protein levels of matrix metalloproteinase 1 (MMP-1) and tissue inhibitor of metalloproteinase (TIMP)-1, but not type I collagen, fibronectin, or TIMP-2. Additionally, PDGF induced the mitogenic and migratory activity of human dermal fibroblasts in a dose-dependent manner. [ 71 ] Another important mediator involved in the pathogenesis of SSc is endothelin 1 (ET-1). Endothelins are potent vasomodulatory peptides produced by endothelial cells, macrophages, fibroblasts and other cell types and can function as downstream mediators of TGF-β responses. [ 72 ] ET-1 signaling via endothelin receptors A (ETRA) and B (ETRB) on fibroblasts induces fibroblast migration and myofibroblast differentiation. Primary cultured dermal fibroblasts from SSc patients and healthy controls treated with ET-1 upregulated collagen type I, CTGF, type I plasminogen activator inhibitor (PAI-1) and pAkt in a time-dependent manner within 72 h. [ 73 ] In addition, the synergistic treatment of endothelial cells isolated from patients with SSc with ET-1 and TGF-β induced activation of the endothelial to mesenchymal transition (EndMT) process. Treatment with macitentan (MAC), an ET-1 receptor antagonist which is clinically used in pulmonary hypertension, prevented EndMT and fibroblast accumulation. [ 74 ] It is worth to note that, due to the multifactorial nature of SSc, a multitude of interconnected growth factor-activated molecular pathways occur across multiple tissues, leading to aberrant signaling crosstalk (which includes also cytokines, chemokines, adipokines, neurotrophins, and metabolites) and ultimately to organ alterations. For this reason, results obtained in vitro from individual cell populations exposed to single growth factors need to be treated with caution, representing only a minimal fraction of SSc complexity. 5. 2. Cytokines Supplementation A wide range of cytokines have been found to be potent regulators of tissue fibrosis and endothelial damage. [ 75 ] The interleukin (IL)-1 family is a group of 11 proinflammatory and anti-inflammatory cytokines that have been reported to be involved in the pathogenesis in SSc. For example endogenous IL-1α expression by SSc fibroblasts has been demonstrated to increase the expression of IL-6 and PDGF, determining the abnormal function of SSc fibroblasts. [ 76 ] IL-6 is a pleiotropic and pro-inflammatory cytokine that is produced by activated immune cells and stromal cells, including T cells, macrophages, endothelial cells and fibroblasts, and is associated with a wide range of biological functions. [ 77 ] In particular, IL-6 is a potent inducer of matrix production in fibroblasts by increasing TGF-β expression, TIMP-1 synthesis and myofibroblast differentiation, resulting in collagen accumulation. [ 78 ] Treatment with anti-IL-6 therapy (tocilizumab) modified the biological characteristics of dermal fibroblasts derived from SSc patients, restoring functional properties, and reversing TGF-β-activated molecular pathways which were present prior to treatment. [ 79 ] IL-4 is a type 2 cytokine activated by CD4 + and CD8 + T cells and mast cells. [ 80 ] IL-4 has been demonstrated to be a profibrotic cytokine participating in cutaneous, cardiac fibrosis, pulmonary fibrosis, and hepatic fibrosis. IL-4 supplementation was shown to induce fibroblasts proliferation, myofibroblasts differentiation and collagen production in vitro. [ 81 ] IL-13 is another type 2 cytokine that is increased in the serum and lesional tissue of patient affect by SSc. [ 82 ] Supplementation of primary dermal fibroblasts with IL-13 stimulated cell proliferation and ECM synthesis. [ 83 ] IL-17 has been reported to be increased in the peripheral blood and target organs, including skin. It amplifies inflammatory responses by inducing the production of IL-6, CCL2 and CXCL8 (IL-8), MMPs and the expression of adhesion molecules in stromal cells including fibroblasts and endothelial cells. [ 84 ] Supplementation of IL-17 enhanced the proliferation of dermal fibroblasts and induced the expression of adhesion molecules and IL-1 production in endothelial cells in vitro. [ 85 ] TNF-α is a proinflammatory cytokine which has been reported to be elevated in patients with SSc and favors the development of pulmonary fibrosis and pulmonary arterial hypertension. [ 86 ] TNF-α supplementation induced high levels of IL-6 in SSc-derived fibroblasts, participating in the self-perpetuation of inflammation during SSc. [ 87 ] In summary, a myriad of soluble mediators is involved in the pathogenesis and progression of SSc, which are secreted by several cell populations according to the stage of the disease. In order to properly supplement these molecules in vitro, continued efforts to understand native pathophysiological signaling pathways will be necessary. This requirement will help to recapitulate the concentrations and spatial and temporal distributions of bioactive factors during these processes. To date, biomaterial-based GF delivery systems have been optimized to provide differential immobilization efficiency and release kinetics. [ 88 ] While these systems have been extensively used in regenerative medicine, we also foresee their utility for the design of in vitro SSc models. 5. 3. Serum Supplementation The beneficial effect of animal sera has long been recognized as means to promote in vitro cell attachment, expansion, maintenance, and proliferation by providing essentials nutrients and growth factors. [ 89 ] However, animal derived sera, such as fetal bovine serum (FBS) or fetal calf serum (FCS), have several technical disadvantages associated with to batch-to-batch variation, xenoimmunization, and possible contamination with mycoplasma, viruses, endotoxins, and prions. [ 90 ] These limitations can affect the phenotype and the behavior of cells expanded in culture, preventing a quality-by-design approach. [ 91 ] Human-derived sera can replace FBS and FCS supplemented media and can create an in vitro microenvironment that more accurately resembles the human environment. [ 92 ] It has been demonstrated that human dermal fibroblasts viability cultured in human serum (HS) supplementation was much higher compared to FBS supplementation. Furthermore, gene expression analysis showed that fibroblasts cultured with HS supplementation maintained expression of collagen type I, increased expression of collagen type III and fibronectin, and reduced expression of a-smooth muscle actin (a-SMA) compared to FBS. [ 93 ] Serum derived from patients affected by SSc has been demonstrated to contain characteristic serum autoantibodies, profibrotic chemokines and growth factors such as IL-4, IL-17, and CTGF, and sonic hedgehog (SHH), which stimulate fibroblast-to-myofibroblast transition and promote dermal fibrosis. [ 94 ] In addition, a recent study identified a discriminant metabolic profile between the serum derived from patients affected by SSc and healthy patients, suggesting the importance of SSc serum not only as a diagnostic tool for the diagnosis and classification of the disease. [ 95 ] Very few studies have thus far investigated the effect of SSc sera in vitro. In particular, immune complexes containing scleroderma-specific autoantibodies derived from patients’ serum, have been showed to elicit proinflammatory and profibrotic effects in skin fibroblasts. [ 96 ] Another study demonstrated that autoantibodies purified from SSc-patient sera directed to platelet-derived growth factor receptor (PDGFR) were able to induce growth and a pro-fibrotic state in vascular smooth muscle cells through the epidermal growth factor receptor (EGFR). [ 97 ] Another study showed that H 2 O 2 production by endothelial cells and fibroblasts was higher after incubation with SSc sera than with healthy sera. Moreover, this model allowed to test the efficacy of bosentan and N-acetylcystein potentiated 5-fluorouracil (5FU) on the inhibition of oxidative stress. [ 98 ] Treatment with SSc sera has been reported to induce EndoMT of dermal microvascular endothelial cells by reducing the expression of endothelial markers such as CD31 and VE-cadherin, and upregulating of mesenchymal markers, including α-SMA and collagen type I. [ 99 ] Despite these interesting results on the use of patient-derived serum in cell culture, a limited number of investigations have assessed the potential of SSc serum as a tool to recreate the pathological microenvironment in vitro. 5. 4. Cocultures In the native tissue milieu, various cell populations interact between each other, stimulating different signaling pathways, and thus influencing numerous aspects of cell function. In vitro coculture models have been developed to recapitulate the in vivo physical contact and paracrine signaling between cell types. Coculture systems can be carried out either by directly seeding different cell types together in the same culture dish, or indirectly, using transwell inserts, whereby cells are located in the same media, without being in contact. [ 100 ] Multiple cell types, including epithelial cells, fibroblasts, endothelial cells, pericytes and leukocytes respond to noxious stimuli in the pathogenesis of SSc. Several in vitro models have been established to investigate the effect of complex intracellular interactions during pathogenic events. For example, in order to elucidate the influence of SSc epidermal keratinocytes on dermal fibroblasts, dispase-dissociated epidermal layers were directly cocultured within a collagen-embedded fibroblast gel. Results showed that SSc epidermal sheets strongly stimulated fibroblast activation, causing gel contraction and induction of CTGF via IL-1, ET-1, and TGF-β. The addition of exogenous IL-1 receptor antagonist (IL-1ra) blocked gel contraction by SSc epidermis, suggesting a potential therapeutic implication. [ 101 ] Another study investigated the effect of dermal fibroblasts in the impairment of angiogenesis in SSc. By oversecreting pigment epithelium-derived factor (PEDF), a major antiangiogenic factor, SSc-derived fibroblast suppressed tube formation when cocultured with human dermal microvascular endothelial cells (MVECs) in an angiogenesis in vitro assay. PEDF knockdown in SSc fibroblasts reversed this process and rescued the number of tubule formed by MVECs. This pathway may present a promising target for new therapeutic interventions in SSc. [ 102 ] 6. Biochemical Cues 6. 1. Hypoxia Tissue hypoxia is a characteristic feature of SSc and contributes directly to the progression of the disease. [ 103 ] It was demonstrated that fibrotic lesions in the skin of SSc patients exhibit significantly decreased oxygen levels in comparison to SSc nonfibrotic skin and the skin of healthy individuals. [ 104 ] Molecular responses to hypoxia are regulated by the transcription factor hypoxia-inducible factor 1 (HIF-1). While HIF-1 is rapidly degraded after translation under normoxic conditions, its activity increases exponentially after exposure to low oxygen levels. The activation of HIF-1 thus plays a critical role in the transcriptional activation of downstream signaling involved in cell proliferation, angiogenesis and fibrogenesis. [ 105 ] Tissue oxygenation in SSc is impaired by microvascular alterations and reduced capillary density, which result in a decrease of the blood flow and poor oxygen supply. [ 104, 106 ] Oxygen supply is further reduced by accumulation of ECM, which impairs diffusion from blood vessels to cells. [ 107 ] Chronic tissue hypoxia causes a vicious cycle by overexpression of VEGF, which in turn leads to aberrant vessel formation and TGF-β activation, thereby increasing tissue fibrosis. [ 108 ] Hypoxia induces multiple ECM proteins in dermal fibroblasts in vitro, such as thrombospondin 1, collagens, fibronectins, insulin-like growth factor binding proteins, and transforming growth factor β-induced protein, in a time-dependent manner. [ 109 ] Numerous studies have contributed to the understanding of the role of hypoxia on the molecular mechanism of SSc. For instance, it has been showed that dermal fibroblasts stimulated with hypoxia (1% oxygen tension) showed increased CTGF expression via activation of HIF-1α, contributing to the progression of fibrosis. [ 110 ] Hypoxia can also drive epithelial-mesenchymal transition (EMT). It was demonstrated that severe hypoxia (1. 5% oxygen tension) as well as moderate hypoxia (3% oxygen tension) induced the expression of α-smooth muscle actin (α-SMA) and vimentin and decreased the expression of E-cadherin in alveolar epithelial cells (AEC). In addition, hypoxia increased the levels of TGF-β, and preincubation of cells with an inhibitor of the TGF-type I receptor kinase prevented the hypoxia-induced EMT, suggesting that the process was TGF-β dependent. [ 111 ] Therapeutic strategies targeting hypoxia have been tested in vitro. 2-Methoxyestradiol (2-ME), a potent inhibitor of HIF-1α, inhibited the fibrotic effect of hypoxia on SSc fibroblasts by down-regulating CTGF and collagen I through the PI3K/Akt/mTOR/HIF-1α signaling pathway. In addition, 2-ME induced apoptosis and inhibited proliferation of SSc fibroblasts. [ 112 ] Despite the importance of hypoxia in SSc, conventional in vitro conditions expose the cells to a non-physiological hyperoxic environment (20% oxygen tension) that is far from recapitulating the pathological microenvironment. 6. 2. Reactive Oxygen Species Oxidative stress, as defined by an imbalance between oxidants (reactive oxygen and nitrogen species (ROS/RNS) and antioxidants, is consistently observed in patients with SSc. [ 113 ] Unpaired electrons make free radicals highly reactive and among them, the superoxide radical (·O 2 − ), hydrogen peroxide (H 2 O 2 ), hydroxyl radical (·OH − ), hypochlorous acid (HOCl) and peroxynitrite (ONOO − ) are key oxidative molecules within the ROS family. [ 114 ] Several oxidative stress biomarkers, such as nitric oxide, malondialdehyde (MDA-a marker of lipid peroxydation), asymmetric dimethylarginine (ADMA) and hydroperoxides are elevated in the blood of SSc patients compared to healthy controls. [ 115 ] Oxidative stress has been demonstrated to cause the activation and damage of ECs, leading to vascular hyperreactivity, apoptosis and impaired angiogenesis. [ 116 ] Increased ROS generation has been reported to mediate TGF-β-induced EndMT. [ 117 ] ROS have been shown to support chronic inflammation and autoimmunity through the genesis of neoepitopes and the activation of T and B lymphocytes and macrophages. [ 118 ] Permanent overproduction of ROS stimulates the proliferation and activation of fibroblasts and their synthesis of ECM. [ 119 ] Also, fibroblasts from SSc have been shown to maintain an overproduction of ROS in SSc through the upregulation of nicotinamide adenine dinucleotide phosphate (NADPH) oxidase (NOX)-2 and NOX-4 proteins. [ 120 ] NOX is a family of enzymes involved in the generation of ROS, acting via the transfer of a single electron from NADPH to oxygen. [ 121 ] The activation of NOX has been demonstrated to trigger fibroblast proliferation and expression of type I collagen genes in SSc cells, thereby maintaining an autocrine feedback mechanism of ROS generation. [ 119 ] As oxidative stress impacts many aspects of the pathophysiology of SSc, several in vitro studies assessed the potential of natural and synthetic antioxidants in the supportive treatment of SSc. The antioxidant epigallocatechin-3-gallate (EGCG), a polyphenol present in green tea extracts, has been shown to reduce oxidative stress and the fibrotic effects on activated dermal fibroblasts from SSc patients. [ 122 ] The antioxidant effect of kaempferol, a natural flavonoid, was investigated on H 2 O 2 -induced intracellular ROS accumulation in SSc fibroblasts and suppressed the intracellular accumulation of ROS and reduced H 2 O 2 -induced apoptosis. [ 123 ] N-acetylcysteine (NAC), a scavenger of free radicals and a precursor of the major antioxidant glutathione, inhibits fibroblast proliferation and collagen synthesis [ 119 ] and reduces peroxynitrite (ONOO − ) synthesis by activated lung macrophages from SSc patients in vitro. [ 124 ] Although ROS seem to have an important role in fibrosis, a therapeutic strategy utilizing antioxidants is not yet clear and further investigations are needed to further elucidate the mechanisms linking ROS dynamics and SSc pathogenesis. 7. Biophysical Cues 7. 1. Substrate Stiffness In addition to biological and biochemical signals, the dysregulation of biophysical properties of the tissue microenvironment in skin, lung, and other organs have been associated with the progression of fibrosis in SSc. [ 125 ] Excessive deposition of ECM increases tissue stiffness and reduces the elasticity of affected organs, leading to mechanical stress ( Table 3 ). [ 126 ] Tissue stiffness can be measured as the elastic modulus, defined as the resistance to deformation, and expressed as the magnitude of a stress (compression, elongation, or shear force, normalized to area) divided by the strain (deformation) induced by the stress. [ 127 ] Increased matrix stiffness anticipates the development of fibrosis, which suggests that tissue stiffening may induce the early activation of myofibroblasts. [ 4 ] Matrix stiffness orchestrates fibrosis by controlling fundamental profibrotic mechanisms including mechano-activation of myofibroblasts via mechano-transduction pathways. Mechano-transduction involves cell surface integrins and changes in cytoskeletal tension mediated by focal adhesion kinase (FAK) and RHO-associated kinase (ROCK). These signals activate the downstream effectors YAP (Yes-associated protein), TAZ (transcriptional coactivator with PDZ-binding motif) and myocardin-related transcription factor (MRTF) which increase fibroblast activation and further perpetuate the fibrotic process ( Figure 4 ). [ 128 ] Lung fibroblasts cultured on stiff substrates showed an increase in proliferation and differentiation into myofibroblasts in comparison to soft substrates. [ 129 ] Moreover, high substrate stiffness increased the synthesis of ECM and the expression of α-SMA and decreased the expression of matrix proteolytic genes and prostaglandin E2 (PGE-2). [ 129 ] The increase in myofibroblasts’ contractility leads to further matrix remodeling and stiffening and activation of TGF-β, amplifying the signal and leading to a positive feedback loop. [ 130 ] TGF-β1 is secreted and stored in the ECM in a latent multiprotein complex with latency associated peptide (LAP) and latent TGF-β1 binding proteins (LTBPs). [ 131 ] The activation of TGF-β1 first requires the binding of αv integrin to arginine-glycine-aspartic acid (RGD) sequences in LAP. [ 132 ] Mechanical forces exerted by cells on LAP via integrin-based adhesions then lead to changes in the conformational state of this complex, which ultimately release and activate TGF-β1 for receptor binding. [ 133 ] Culture systems with tailored mechanical properties have been engineered to mimic aspects of diseased tissues and investigate the mechanism of myofibroblast activation. For instance, stiff collagen hydrogels promoted stress fibers formation, smooth muscle actin (SMA) expression and TGF-β1-induced response in human fibroblasts. [ 134 ] Another study identified that a stiff polyacrylamide substrate induced lung myofibroblast differentiation through actin cytoskeletal remodeling-mediated activation of megakaryoblastic leukemia factor-1 (MKL1), which transduces mechanical stimuli from the ECM, leading to the induction of a fibrogenic program. [ 135 ] A recent study showed that the activity of the transient receptor potential vanilloid 4 (TRPV4) channel was increased when cells were plated on stiff collagen-coated polyacrylamide gel matrices within the pathophysiological range seen in diseased/fibrotic dermal tissue. Genetic ablation or pharmacological antagonism of the TRPV4 channel abrogated Ca 2+ influx and matrix stiffness-induced myofibroblast differentiation, evidencing that therapeutic inhibition of TRPV4 activity may provide a targeted approach to the treatment of scleroderma. [ 136 ] Given the increasing recognition that ECM stiffness is a major factor contributing to the progression of SSc, it has become evident that the identification of optimal substrate stiffness which replicates pathological fibrotic conditions will enable the development of more precise mechano-therapeutic interventions for tissue stiffening. Therapeutic strategies targeting mechanotransduction signaling mediated by integrins, [ 132b ] FAK, [ 137 ] ROCK, [ 138 ] or YAP/TAZ [ 139 ] have been promisingly tested in preclinical and clinical studies. 7. 2. Substrate Topography The topographical organization of the ECM significantly influences cell morphology and behavior, including growth, adhesion, and migration. [ 140 ] Cells sense their underlying topography via focal adhesion interactions, pushing and pulling against the matrix and activating a cascade of cellular and molecular events, which ultimately influence gene and protein expression. [ 141 ] Morphological changes to the dermal collagen organization and focal adhesion complexes have been reported in skin biopsies from SSc patients. These changes are characterized by the presence of highly aligned collagen bundles, and a loss of normal “basket-weave” collagen organization that is characteristic of the healthy dermis. [ 142 ] The alteration of the stiffness of the ECM significantly contributes to the alignment of the collagen fibrils, and further amplifies the fibrotic process. [ 143 ] The ability of cells to mechanically sense these changes may be due to the deposition and realignment of new collagen fibrils in which cells generate myosin-generated tensile forces applied through cell-matrix adhesions. [ 144 ] In vitro studies have demonstrated that increasing stress fiber formation and ECM alignment increase the elastic modulus of the fibroblast-populated collagen gels over a culture period of 21 d. [ 145 ] Moreover, several studies have reported that mechanical strength of anisotropic nanofibers is significantly higher than that of the disordered nanofibers. [ 146 ] The anisotropic ECM ultrastructure within the fibrotic microenvironment is a critical cue in maintaining the myofibroblasts phenotypes in SSc. This was demonstrated by using a 3D model of either randomly oriented or aligned electrospun poly-caprolactone (PCL) nanofibers with adsorbed type I collagen. Guidance cues from aligned collagen fibers enhanced the fibrogenic potential of dermal fibroblasts by increasing cell migration, adhesion, and guidance signaling pathways. Arhgdib (Rho GDP-dissociation inhibitor 2) was one of the most upregulated genes following fibroblast culture on aligned fiber substrates, and siRNA knockdown of Arhgdib significantly reduced directed cell migration under aligned fiber culture conditions. [ 147 ] Another study utilized glass slides coated with aligned fibers to investigate alignment and migration of lung fibroblasts isolated from SSc patients. The results indicated that migration took place when lung fibroblasts were cultured on aligned collagen following stimulation with PDGF, but was not induced on the woven, randomly orientated collagen substrates. [ 148 ] In addition, heparin, which binds ligands including PDGF and stem cell factor (SCF), and imatinib, which blocks downstream tyrosine kinase receptors, both inhibited lung fibroblast migration individually. Importantly, the two drugs showed synergistic effect in SSc cells, supporting a possible pilot evaluation of combination therapy. [ 148 ] A recent study investigated the effect of collagen microarchitecture on myofibroblast differentiation and fibrosis. Adipose stromal cells (ASCs) were cultured on collagen gels consisting of networks with thin fibers and low porosity, or scaffolds with thicker collagen fibers with larger pores. Interestingly, ASCs contractility on collagen matrices with thicker fibers and larger pores resulted in collagen fiber densification and alignment in the direction of cell polarization and migration, increasing stiffness in a physiologically relevant regime of strain. The stiffening of local matrix, in turn, stimulated a contractile phenotype via a positive feedback loop, thereby modulating myofibroblast differentiation and fibrosis. [ 149 ] Due to the interaction between increased ECM alignment and the formation of fibrotic lesions, it is of considerable interest to further investigate the influence of patterned cell culture substrates on tissue mechanics and collagen alignment. 7. 3. Macromolecular Crowding The recapitulation of collagen matrix formation in vitro has proven challenging, partly due to the omission of important cofactors in the culture media and partially because of sub-optimal cell culture conditions. The omission of ascorbate in cell culture results in minimal production and deposition of collagen within the cell layer. Ascorbic acid is a crucial cosubstrate for the enzymes prolyl hydroxylase and lysyl hydroxylase, which are responsible for the posttranslational hydroxylation of prolyl and lysyl residues on collagen fibers. Hydroxylation of the prolyl residues renders the collagen triple helix thermostable and hydroxylation of the lysyl residues is responsible for the extracellular cross-linking of collagen fibers. [ 150 ] However, even with ascorbate supplementation, cells deposit only subphysiological amounts of secreted collagen I into their matrices. Indeed, in standard cell culture settings, the conversion of water-soluble procollagen to insoluble collagen is relatively slow, since the proteinases required for the enzymatic cleavage of procollagen are dispersed in dilute culture media. [ 151 ] To overcome this limitation, macromolecular crowding (MMC) has been introduced as a means to accelerate ECM deposition in vitro. The addition of macromolecules into the culture media emulates the naturally crowded in vivo milieu, and thus amplifies deposition of cell-secreted ECM. [ 152 ] The addition of macromolecules to the culture media also results in a more efficient volume occupancy, preventing the dispersion of active molecules, consequently accelerating the conversion of procollagen to collagen and ultimately deposition of the latter. [ 153 ] Moreover, use of MMC has been demonstrated to drive the molecular assembly of collagen fibrils in vitro and to stabilize the formed matrix through enzymatic crosslinking. [ 154 ] One application of MMC to produce the full deposition of collagen in fibroblast cultures has been the development of a valuable screening tool, the so-called “Scar in a Jar”, for antifibrotic compound screening. [ 155 ] The original scar in a jar model consisted of human fibroblasts cultured in the presence either of 500 kDa dextran sulphate (DxS) or a mixture of neutral 70 and 400 kDa Ficoll (Fc). Crowding of culture medium with dextran sulphate served as the rapid deposition mode as it yielded maximum granular deposition of collagen I by 48 h, whereas neutral mixed Ficoll served as the accelerated mode (Fc), which resulted in the deposition of a fibrillar collagen meshwork after 6 d of culture in comparison with non-crowded cultures. [ 156 ] This system has been used to evaluate the potential of antifibrotic compounds effective at the epigenetic, post-transcriptional/translational level. Another study utilized the principle of macromolecular crowding to create an ECM-rich in vitro hypertrophic scar model that more closely recapitulates “in vivo-like” conditions than customarily used monolayer culture systems. [ 157 ] This model was used to test the antifibrotic effect of a series of naphthalene derivatives derived from medicinal herbs on human dermal fibroblasts. Interestingly, shikonin and naphthazarin were shown to inhibit the transcription and translation of collagen in human scar-derived fibroblasts and induced apoptosis via the mitochondrial apoptosis pathways, suggesting their potential therapeutic value for the treatment of dermal fibrosis and scarring. [ 157 ] A similar study, combined macromolecular crowding with TGF-β to develop a robust, high throughput, phenotypic screening assay using pulmonary fibroblasts derived from patients affected by idiopathic pulmonary fibrosis. [ 158 ] The Scar-in-a-Jar offers a novel pathophysiologically relevant screening and evaluation tool for antifibrotic compounds interfering with different key steps in the collagen biosynthesis pathway. Overall, despite all these progresses in the modeling of fibrosis using MMC in vitro, the importance of the crowded extracellular niche is still underestimated, [ 159 ] likely due to the fact that optimal crowding agents remain still elusive and further studies are required to reveal and unravel their diverse effects on cell behavior and phenotype. Optimization of MMC protocols will contribute to the generation of models with high levels of biomimicry which can be used as an instrument to recreate SSc conditions. 7. 4. Dimensionality and 3D Architecture Cell-based assays have been crucial in drug discovery and provide a simple, rapid, and cost-effective tool to support screening of therapeutics before large-scale and cost-intensive animal testing. To date, the majority of cell-based assays are based on traditional 2D monolayer cultures grown on flat dishes optimized for cell attachment and growth. [ 160 ] However, 2D culture systems have multiple limitations, including loss of tissue-specific architecture, non-physiological mechanical and biochemical signals, and non-physiologic cell-to-cell and cell-to-matrix interactions. [ 161 ] The standard 2D environment may therefore provide misleading results regarding the predicted responses of cells to drug treatments ( Figure 5 ). [ 162 ] To overcome these important limitations, there has been tremendous progress in tissue engineering and regenerative medicine over the past few decades that has led to the development of a wide range of 3D cell culture systems. Indeed, multicellular organisms reside within a complex 3D environment, rich with multiple ECM components, several cell populations and soluble factors, which not only provide structural support to tissues and organs, but also physiological conditions that allow for optimal functionality and delineate specific microenvironments. [ 163 ] Table 4 summarizes several 3D scaffold-free and scaffold-based SSc models, including the therapeutic approaches used to evaluate each model’s feasibility for drug development and screening. 7. 4. 1. Self-Assembled Models Tissues are formed by building blocks that self-assemble into highly organized structures that enable and regulate their functions. [ 164 ] Tissue engineering by self-assembly (TESA) relies on the inherent capacity of cells to self-assemble into highly organized 3D tissue-like constructs and to produce their own ECM, without the need of an external scaffold. These assemblies can be created through self-assembled aggregation, cell sheets, tissue strands or direct bioprinting. [ 165 ] 3D cell aggregates allow for the fabrication of organotypic microtissues due to their multidimensional cell–cell interactions and communication. Multicellular spheroids are scaffold-free cellular models based on the spontaneous aggregation of cells into spherical compact clusters on nonadherent substrates. The complex cell interactions recapitulate spatial and functional characteristics of the native tissue modulating cell activities and signaling. [ 166 ] For this reasons, 3D multicellular spheroids have been utilized as in vitro models of fibrosis. Recently, 3D human fibroblasts spheroids have been created for the development of an in vitro fibrogenesis model. Fibroblast-based spheroids showed significantly higher expression levels of fibrotic genes (αSMA and collagen I) compared to 2D monolayer culture. Furthermore, since the absence of immune cell mediators was recognized as a likely limit to physiologic model behavior, hybrid spheroids were fabricated with fibroblasts and macrophages. The addition of macrophages to the fibroblasts spheroids at a ratio of 1:16 (macrophages-fibroblasts) resulted in an increase of fibroblast activation and myofibroblast differentiation. Similarly, more macrophages were polarized toward an inflammatory type M1 in this group with greater CCR7 and pSTAT1 expression. In addition, hybrid spheroids demonstrated high expression of fibrosis-related genes (collagen I, collagen III and TGF-β) and inflammatory genes (TNF, IL-1β and IL-6). This system thus represents a valuable in vitro fibrogenesis model for high-throughput antifibrosis therapy screening. [ 167 ] Building on the success of spheroids, researchers have also focused on organoids, which generally better replicate tissue morphology and organization, and embed multiple cell populations that are distributed physiologically. [ 168 ] Organoids are in vitro self-assembling 3D organ-like architectures grown from tissue-specific adult stem cells or pluripotent stem cells, such as embryonic stem cells (ESCs) and induced pluripotent stem cells (iPSCs). [ 169 ] Organoids have been engineered to model pathogenetic mechanisms that affect lung homeostasis and involve complex interactions between different cell types, such as those occurring in interstitial lung diseases, which is a frequent complication of systemic sclerosis. [ 170 ] For instance, lung organoids composed of iPSC-derived mesenchymal cells, treated with exogenous TGF-β1, demonstrated increased contraction and the development of fibroblastic foci by expressing of collagen 1 and α-SMA. [ 171 ] Organoids represent valuable tools for the screening of compounds with pro-regenerative, antifibrotic or tissue protective capabilities for precision medicine. However future applications of organoids may be limited by the fact that they lack mechanical cues, vasculature, and immune components, and may be prone to tumorigenicity in case of IPSCs. Another strategy to fabricate cellular 3D constructs consists of stacking of multiple cell sheets, which takes advantage of the cell-deposited ECM network that intertwines the different cell sheets. [ 172 ] Self-assembled reconstructed dermis equivalents have been created to study the fibrotic phenotype of SSc fibroblasts. Fibroblasts isolated from lesional or nonlesional skin of early- and late-stage SSc patients were grown for 35 d to form sheets and two fibroblast sheets were layered and maturated to generate a thick dermal-like layer. [ 173 ] The study showed that in this system, development of skin fibrosis resulted in progressive changes at the fibroblast level, from a normal phenotype to a sustained and autonomous fibrotic phenotype. Based on these results, the authors suggested that antifibrotic treatments of SSc could gain efficacy if they were tailored to disease duration and severity. Another study developed a tissue-engineered model of self-assembled reconstructed skin to mimic interactions between dermal and epidermal cells. [ 174 ] Four dermal fibroblasts layers were superimposed to form a reconstructed dermis and undifferentiated, differentiated and unsorted keratinocytes were seeded onto the dermal sheets. Results showed that undifferentiated keratinocytes inhibited dermal fibrosis through downregulation of TGF-β, induction of FGF-β (fibroblasts growth factor β) and initiation of desmosome formation, indicating that undifferentiated keratinocytes may be a promising option for fibrotic scar prevention. Despite the utility of self-assembled skin models for disease modeling, these systems lack the presence of critical components such as hair follicles, glands, tactile corpuscles, and subcutaneous adipose tissue, which play a critical role in physiological functioning. However, progress in biofabrication techniques will likely address the incorporation of such appendages. [ 175 ] 7. 4. 2. Scaffold-Based Models Scaffold-based models of 3D tissues consist of cells grown in the presence of support scaffolds consisting of either hydrogel-based or polymeric fiber-based materials. These scaffolds can be composed of natural, synthetic, or combinations of different polymers, and represent a 3D construct that is structurally, mechanically and functionally similar to the biological tissue. [ 176, 175 ] For the fabrication of human skin equivalents, dermal fibroblasts are generally cultured in collagen matrices to recreate a dermal-like layer, whereas keratinocytes are cultured on top of these dermal equivalents, followed by air–liquid interface culture to promote full epidermal maturation. [ 177 ] For example, bioengineered human skin equivalents were fabricated using both SSc-derived dermal fibroblasts and normal dermal fibroblasts. [ 178 ] Stromal equivalents were assembled by embedding dermal fibroblasts into collagen type I hydrogels. Full skin equivalents were fabricated by adding keratinocytes onto the collagen gel, after the maturation of the dermal layer. Results evidenced that SSc fibroblasts altered collagen architecture, as seen by a more mature and aligned fibrillar structure, enhanced stromal rigidity with increased collagen crosslinking, and upregulated LOXL-4 expression and innate immune signaling genes. Interestingly, knockdown of LOXL-4 suppressed rigidity, contraction and α-SMA expression in the SSc skin equivalents and TGF-β-induced ECM aggregation and collagen crosslinking in the stromal compartment. This skin-like tissue platform provided a suitable tool to test mechanisms that mediate skin fibrosis. Nevertheless, the lack of vascular and immune cells into the 3D skin-like tissues limits its potential. Bioengineered skin equivalents have also been utilized to develop skin-humanized mouse models to test the progression of SSc and to monitor the response to antifibrotic drugs in vivo. For instance, 3D bioengineered skin, based on plasma-derived hydrogels and containing human SSc-derived keratinocytes and fibroblasts, were grafted onto immunodeficient (SCID) mice. Results showed that the human skin-SCID mouse models closely replicate the SSc fibrotic phenotype in vivo up to 16 weeks. [ 179 ] Despite the relevance of cell populations derived from SSc patients, the fabrication of 3D tissue equivalents is limited by the availability of donor cells. In this regard, human iPSCs offer great potential to generate relevant cell types and can provide an alternative source of cells for tissue engineering purposes. [ 180 ] Currently, patient-specific iPSCs are generated by reprogramming of adult somatic cells by ectopic expression of pluripotency-associated transcription factors including OCT4, SOX2, KLF4, and c-MYC. [ 181 ] Cord blood mononuclear cells (CBMCs)-derived iPSCs have been used to generate dermal fibroblasts, which have been used to create 3D dermal equivalents. Treatment with TGF-ß1 activated iPSC-fibroblasts and increased their proliferation rate and ECM production. In addition, TGF-β1 treatment increased the thickness of the 3D iPSC-fibroblasts construct. Treatment with pirfenidone, a drug used to treat idiopathic pulmonary fibrosis, elicited an antifibrotic effect, attenuating the increase in dermal thickness and expression of fibrosis genes. These results suggest that the use of iPSC-derived fibroblasts in skin equivalents could be utilized for drug screening purposes. [ 182 ] A simple and versatile technique used for the fabrication of ECM-mimicking fibrous scaffolds is electrospinning. Electrospun fibrous scaffolds create nano- to microscale fibrillar network with interconnecting pores, resembling natural ECM in tissues, and facilitating the formation of artificial tissues in vitro. [ 183 ] Poly-caprolactone (PCL) electrospun scaffolds coated with bleomycin treated lung extracts, obtained from solubilization with a glass homogenizer, have been used as a model of idiopathic pulmonary fibrosis (IPF) to induce bone marrow-derived cells to differentiate into myofibroblast-like cells. [ 184 ] This approach combines biological extracts isolated from fibrotic lungs with synthetic nanofibers that serve as an ECM-like substrate to determine the effect of biochemical signals present in the fibrotic microenvironment. This model has the potential to help identify compounds that either mitigate or reverse the fibrotic differentiation of bone marrow -derived cells. Moreover, bone marrow-derived cells cultured on electrospun fibers with higher elastic modulus displayed increased fibrotic gene expression, demonstrating the importance of matrix modulus in cell differentiation. [ 184 ] The design of electrospun scaffolds with appropriate topographical features is critical to determine cell function and to foster desired cell differentiation. For example, a recent study showed that nanometer scale electrospun fibers upregulated the expression of α-SMA, TGF-β, and vimentin filaments in comparison to micrometer scale fibers. The size change of the electrospun fibers (from micrometers to nanometers) altered fibroblast differentiation and led to higher α-SMA expression and more contractile myofibroblasts. [ 185 ] Another study demonstrated that electrospun fiber diameter can modulate epithelial to mesenchymal transition (EMT). Indeed, epithelial cells grown on fibers with an average diameter of 5μm exhibited a downregulation of epithelial markers such E-cadherin and upregulation of mesenchymal markers such as vimentin. However, cells grown on fibers with an average diameter of 0. 5 μm grew as compact colonies with a stable epithelial phenotype. [ 186 ] Despite considerable advances in biomaterials design and development for the engineering of physiological-relevant tissue models, several challenges hinder the applicability of such models as part of daily pharmaceutical research. One such limitation is the scalability and manufacturing of complex, bio-mimetic and reproducible scaffold-based models in compliance with current good manufacturing practices (cGMPs), which result in cost-intensive processes, not comparable to 2D and self-assembled models. 7. 4. 3. Decellularized Matrix Bioscaffolds Given the complexity of ECM composition and structure, designing and fabricating a biomaterial-based scaffold that fully mimics the biochemical and structural properties of native tissue ECM is currently challenging. However, decellularization of whole tissues and organs by removing cellular components provides a useful method for harvesting an ECM which retains tissue-specific 3D morphology, biochemical, and biomechanical cues. [ 187, 163a ] A 3D model of the fibrotic lung microenvironment created from decellularized lung explants was used to determine whether the lung ECM of patients with scleroderma leads to the development of fibrocytes from peripheral blood mononuclear cells. Fibrocytes are collagen-producing leukocytes abundantly present in patients with SSc-related interstitial lung disease (ILD) via unknown mechanisms that have been associated with altered expression of neuroimmune proteins. Culture of control cells and patient-derived cells in lung scaffolds from patients with SSc-related ILD increased production of procollagen I, which was stimulated by enhanced stiffness and abnormal ECM composition. Moreover, enhanced detection of netrin-1 (a laminin-like protein that regulates cell–matrix interactions) expression in cells from patients with SSc-related ILD was observed, and antibody-mediated netrin-1 neutralization attenuated detection of collagen-producing leukocytes in all settings. [ 188 ] This study demonstrated the utility of decellularized platforms for disease modeling and potential drug discovery. Other studies provided insights into ECM-mediated positive-feedback loops using decellularized lung scaffold from patients with IPF. [ 189 ] In the absence of exogenous factors, IPF ECM alone can promote normal lung fibroblasts to become activated myofibroblasts, and once fibroblasts are activated, IPF ECM sets up a positive feedback loop capable of sustaining progressive fibrosis. Although successful decellularization has been achieved for many organs, standardized decellularization protocols still need to be defined with the final goal of advancing the creation of in vitro models. [ 190 ] Moreover, the development of decellularized skin matrices will allow further exploration in the pathogenesis of SSc dermal fibrosis. 7. 4. 4. Organ on Chip Recent advances in microfabrication and microfluidics have enabled the development of microengineered models of human organs—known as organs-on-chips—that have the potential to provide platforms to model organ level responses for drug discovery in an in vitro setting. [ 191 ] Important cues involved in the pathogenesis of fibrosis, such as mechanical strain, fluid flow, and hydrostatic pressures, can be integrated in such models. Recently, a 3D-bioengineered pulmonary fibrotic (Eng-PF) tissue was developed utilizing silk collagen-I hydrogels seeded with pulmonary fibroblasts, airway epithelial cells, and microvascular endothelial cells. Eng-PF tissue was cultured under tension, tethered along the longitudinal axis of a bioreactor plate, and had the capacity to be cyclically strained with perfusion ability. Eng-PF tissues were able to model myofibroblast differentiation and permit evaluation of antifibrotic drugs, such as pirfenidone and nintedanib. Further, Eng-PF tissues were used to model epithelial injury with the addition of bleomycin and cellular recruitment by perfusion of cells through a hydrogel microchannel. [ 192 ] Unlike other 3D tissue systems, lung-on-chip platforms are standalone devices, therefore scaling these devices for high throughput antifibrotic drug screening could be challenging. Organ-on-chip technologies have also been advanced in the production of skin equivalents. The main advantages supported by microfluidics are the presence of fluid flow and fine control of the microenvironment, which yield improved epidermal morphogenesis and differentiation, and enhanced barrier function. [ 193 ] Although a skin-on-a-chip model consisting of epidermal, dermal and endothelial cells has been used to stimulate skin inflammation and edema, [ 194 ] a skin chip-model that replicate dermal fibrosis still needs to be developed. 8. Missing Players: Vasculature and Hematopoietic Immune Cells Recent research has enabled the development of in vitro models which recapitulate different aspects of SSc, which could significantly contribute to early-stage drug discovery. However, current bioengineered systems often lack the presence of perfusable (micro-) vasculature and organ-specific immune responses, which need to be incorporated into the tissue models to achieve patho-physiologically relevant levels of function. It is known that patients affected by SSc present a series of vascular abnormalities. Among them are direct damage of vascular and perivascular cells, abnormal vasoreactivity, hypoxia, impaired angiogenesis, and platelet dysfunction. These vascular changes result in decreased capillary blood flow, and subsequently in clinically overt symptoms such as Raynaud’s syndrome and fingertip ulcers. [ 195 ] In addition, vascular damage contributes to the production of a cascade of soluble mediators that ultimately influence the onset and progression of tissue fibrosis. [ 196 ] Given the importance of the vasculature in SSc, the integration of vascular structures into in vitro cultured tissues is of paramount importance to provide realistic models of complex interactions, which modulates tissue remodeling and fibrosis in the tissue of interest. One common strategy relies on the spontaneous organization of endothelial cells to form vascular networks within biomaterials scaffold or in multicellular assemblies. [ 197 ] Initially, endothelial cells form a primitive network within an avascular tissue, which is similar to vasculogenesis. In order to maturate into a functional vascular network with sufficient vascular cell viability and function, perfusion bioreactors are required to mimic in vivo like flow rate. [ 198 ] For example, a recent study focused on the development of vascularized skin equivalents as an advanced model of human skin with a fully polarized epidermal layer, a stratified dermis and a functional vascular system with physiological perfusion pressures. [ 199 ] These models were induced to undergo fibrotic transformation and resembled key features of SSc skin, with accumulation of ECM, fibroblast to myofibroblast transition and aberrant activation of TGF-β signaling. In addition, treatment with nintedanib in a pharmacologically relevant dose exerted antifibrotic effects in vascularized human skin equivalents by attenuating TGF-β signaling, reducing fibroblast to myofibroblast transition and decreasing ECM deposition. [ 199 ] By incorporating a mature vascular network, this platform provided a pathophysiologically relevant human setting for the evaluation of antifibrotic drugs, potentially improving predictive value. Considering that the immune system plays a pivotal role in this disease, [ 200 ] another key element which is generally missing during the development of in vitro tissue analogues is the presence of relevant immune cells. Activation of both innate and adaptive immune responses leads to activation of fibroblasts, differentiation into myofibroblasts, ECM deposition and finally fibrosis. Monocytes, macrophages and dendritic cells all release soluble mediators that can directly affect fibroblast activation and tissue remodeling, or indirectly affect it by inducing the release of profibrotic factors by other cell types, including T cells and B cells. [ 18 ] Using cocultures of immune cells with fibroblasts, a number of valuable in vitro models have tried to approximate the role of innate/adaptive immunity activation in fibroblasts differentiation and organ fibrosis. For example, it was demonstrated that coculture of SSc fibroblasts and SSc plasma-differentiated macrophages (expressing high levels of CCL2, IL-6, and TGF-β) using transwell plates resulted in activation of signaling pathways involved in the regulation of inflammation and fibrosis, suggesting that therapeutic targeting of these cells may be beneficial in ameliorating SSc progression. [ 201 ] Another study demonstrated that B cells cocultured with human dermal fibroblasts (HDFs) derived from SSc patients are potent inducers of IL-6, CCL2, and TGF-β1, which enhance collagen production by fibroblasts. [ 202 ] The integration of key immune cell subsets is likely a needed strategy to improve the relevance of SSc in vitro models. However, the complex mechanisms by which the immune system orchestrates organs and tissues are challenging to replicate, especially in prolonged culture conditions. New approaches such as “on-a-chip” platforms can incorporate multiple cell types under controlled biochemical and biophysical conditions contributing to the formation and progression of SSc. [ 203 ] 9. Challenges and Future Directions Among autoimmune diseases, SSc is one of the most devastating pathologies, and its heterogeneity and complexity pose unique challenges for the discovery and development of effective therapeutic strategies. Although tremendous scientific progress has significantly increased the knowledge of the biological and molecular mechanism at the basis of SSc, numerous unanswered questions remain in the field of SSc therapeutics. Animal models represent indispensable for preclinical drug testing, even though none of these models faithfully recapitulates the full spectrum of SSc. For this reason, the development of in vitro models that replicate the complex and dynamic SSc milieu are essential to overcome some of the in vivo disease models’ shortcomings. As herein reviewed, a variety of biological, biochemical, and biophysical cues have integrated into in vitro platforms to mimic the pathophysiological signals typical of SSc, thereby improving their suitability for the identification and screening of effective drugs. However, in vitro models have not yet fully recovered the SSc phenotype to a level comparable with the human disease. Due to the complexity of SSc, it is likely unrealistic to create models that fully embrace all aspect of the pathology, including vascular component, immune system and organ fibrosis. Multifactorial approaches combining synergistic microenvironmental cues are one way forward to develop more complex in vitro systems. Advancements in biomaterial science will contribute to the development of new platforms that not only give structural dimensionality to the models, but will also modulate cell behavior by providing heterogeneous spatial organization and spatiotemporal controlled biological and mechanical signals. [ 204 ] Progress in iPSCs technologies have the potential to provide disease-relevant cells in a personalized manner and could facilitate the development of patient-specific SSc models for personalized medicine without requiring multiple tissue collections. [ 205 ] Emerging technologies that go beyond well-established ex vivo assays for the characterization of fibrotic hallmarks (SDS-PAGE, [ 206 ] quantification of hydroxyproline content, [ 206 ] Sircol collagen assay, [ 207 ] histological and immunohistological analysis), not only help to better understand the biology of the disease, but also lead to improved assessment of the therapeutic effects of potential drug candidates. Examples include assessment of ECM structure and stiffness by second harmonic generation (SHG) microscopy [ 208 ] and atomic force microscopy, [ 145 ] as well as noninvasive imaging of tissue and organ damage [ 209 ] by magnetic resonance imaging (MRI), [ 210 ] computed tomography (CT), [ 211 ] ultrasound (US), [ 212 ] and positron emission tomography (PET). [ 213 ] These imaging techniques can be readily utilized to measure treatment responses over time without the need to sacrifice experimental animals, thereby facilitating the clinical translation of antifibrotic therapies. Overall, despite the wide recognition of the utility and potential of 3D models in the drug development pipeline, there is not substantial evidence that such systems outweigh the data obtained from 2D models or the cost of testing in animals. Indeed, drug discovery programs are still far from being driven by primary screens in 3D and it is still premature to claim that 3D models improve the clinical success rates of drug candidates. In addition, the majority of models are conventionally not designed for automated analyses, being incompatible with high-content screening platforms, and thus are held back by routine applications into industrial settings. Therefore, future efforts should be directed not only to the standardization and validation of such models, but also to their miniaturization in order to enhance experimental efficiency with automation. While further advances are needed, in vitro models of SSc represent promising platforms for disease modeling and for drug discovery, increasing our understanding of the mechanisms underlying this devastating disease.
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10. 1002/adbi. 202000624
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Advanced Biology
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In Vitro Models for Studying Respiratory Host–Pathogen Interactions
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Abstract Respiratory diseases and lower respiratory tract infections are among the leading cause of death worldwide and, especially given the recent severe acute respiratory syndrome coronavirus‐2 pandemic, are of high and prevalent socio‐economic importance. In vitro models, which accurately represent the lung microenvironment, are of increasing significance given the ethical concerns around animal work and the lack of translation to human disease, as well as the lengthy time to market and the attrition rates associated with clinical trials. This review gives an overview of the biological and immunological components involved in regulating the respiratory epithelium system in health, disease, and infection. The evolution from 2D to 3D cell biology and to more advanced technological integrated models for studying respiratory host–pathogen interactions are reviewed and provide a reference point for understanding the in vitro modeling requirements. Finally, the current limitations and future perspectives for advancing this field are presented.
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1 Introduction Respiratory diseases are among the leading causes of death worldwide, [ 1 ] with more than 1 billion people suffering from long standing respiratory illness. [ 2 ] Among others, the most common conditions include asthma, [ 3, 4 ] chronic obstructive pulmonary disorder (COPD), [ 5 ] and idiopathic pulmonary fibrosis, [ 6 ] where dysregulation, immune‐hyperresponsiveness, and remodeling of the airway epithelium is evident. In addition to chronic respiratory disease, respiratory infection also contributes a substantial burden on society, especially lower respiratory tract infections which account for 4 million deaths per annum. [ 2 ] Respiratory pathogens can be highly contagious and those with underlying respiratory or immune disorders are particularly at risk of death. This area of research is highly topical given the recent Severe acute respiratory syndrome coronavirus‐2 (SARS‐CoV‐2) pandemic, which has reached over 110 749 023 confirmed cases and 2 455 131 deaths, as of February 21, 2021. [ 7 ] Thus, basic pulmonary drug research and biopharmaceutical development of respiratory therapeutics, antivirals and vaccines is of paramount importance. In order to deliver the most effective treatments, however, a fundamental understanding of human lung biology is required and with it, models which accurately represent the complexity found in vivo. This is especially important when coupled to the high drug attrition rates, time to market, and ethical concerns surrounding the use of animals in research, currently seen in drug R&D pipelines. [ 8, 9, 10 ] Other reviews exist with a specific focus on microfluidic [ 11 ] and in silico models for drug delivery, deposition, and pharmacokinetics in preclinical lung models [ 12 ] as well as ex vivo tissue engineering for lung transplantation applications. [ 13, 14 ] In this review, however, we discuss the function and cellular composition of the pulmonary epithelium barrier—the first line of respiratory defense. We then describe the pulmonary immune system, providing a primer on its response to common respiratory pathogen, and remodeling of the respiratory epithelium in disease (asthma and COPD). We will finish by presenting the most common in vitro models for studying host‐respiratory pathogen interactions, advances in technology integrated models and future perspectives for studying these complex systems. Altogether, this review should provide the user with a basic biological understanding of the respiratory epithelial barrier and immune components required to study respiratory host–pathogen interactions in vitro. Additionally, it may be used as a reference point for understanding the requirements, relative merits, and drawbacks of using a variety of currently available in vitro lung models, ranging from 2D to complex 3D cultures. In the context of this review, 2D culture is defined as the growth of cell monolayers on a flat substrate, for example, a petri dish or polymer membrane, while 3D culture is defined as a tissue‐specific microenvironment which allows cells to retain their in vivo 3D architecture and function, for example, spheroids, organoids and use of hydrogels, scaffolds, and bioreactors. 1. 1 The Respiratory System The human respiratory system is responsible for essential breathing processes and gas exchange. Furthermore, the pulmonary epithelium constitutes a unique interface with the outside environment, acting as a physical and immunological barrier against noxious stimuli and pathogens. Its homeostatic functions include the dynamic regulation of ion permeability, transport of essential nutrients and antimicrobial secretion. [ 15 ] The respiratory system can be divided into the upper (nasal cavity, pharynx, and larynx) and lower airways (trachea, bronchi, bronchioles, alveoli, and lung parenchyma) ( Figure 1 ). The lower airway can then be further sub‐divided into three zones, according to the cellular phenotypes present: The proximal airway (trachea and bronchi) (Figure 1A ), the bronchoalveolar duct junction (Figure 1B ) and the alveoli (Figure 1C ). The proximal airway consists of a mucus layer, a thin surfactant layer, a periciliary layer, and the epithelial layer. Mucus consists of water (97%) with small amounts of lipids, carbohydrates, and proteins. [ 16 ] The most abundant proteins are mucins, secreted by goblet cells or submucosal glands, which give the mucus a gel like consistency and overall negative charge. [ 16 ] These characteristics aid in capturing inhaled particles, toxins and pathogens, which are cleared from the respiratory tract via the coordinated beating of cilia (mucociliary clearance). Beneath the mucus, in contact with the cilia, is the periciliary fluid layer which contains anti‐microbials, anti‐virals, and anti‐fungals. [ 17 ] Surfactant is an amphiphilic layer located between the mucus and periciliary fluid, containing predominantly phospholipids and cholesterol, with its main role to reduce surface tension and increase respiratory compliance. Surfactant is secreted in small amounts by club (formerly Clara) cells in proximal airways, although its major source of secretion is from alveolar type 2 pneumocytes. [ 17 ] Figure 1 Cellular components of the lower airway pulmonary epithelium. A) The proximal (trachea, bronchi) airway epithelium consists of secretory club cells, ciliated cells, mucus producing goblet cells, basal stem cells, and pulmonary neuroendocrine cells. B) The distal portion of the lower airway consists of the bronchioles, the bronchoalveolar duct junction, and the alveoli. The bronchoalveolar duct is comprised of ciliated and club cells only. C) The alveolar epithelium consists of type 1 and type 2 pneumocytes. The blood circulation and immune cells also contribute to the defense mechanisms via interaction with the pulmonary epithelium. Image created using BioRender. com. In addition to the aforementioned cell types, the proximal epithelium also comprises basal (stem) cells and pulmonary neuroendocrine cells (PNEC). Basal cells are responsible for epithelial regeneration upon damage, while PNECs are involved in neuroendocrine, exocrine, and immune signaling. The epithelium not only regulates selective permeability, but also homeostatic levels of hydration by active transport through the epithelial sodium channel, the cystic fibrosis transmembrane conductance regulator channel, and the calcium activated chloride channel. [ 18 ] In contrast to proximal regions, the alveolar epithelium contains no ciliated cells nor does it secrete mucus, as this would reduce efficiency of gas exchange across the air‐blood barrier. [ 19 ] Instead, a surfactant fluid layer together with alveolar macrophages are responsible for the protection against inhaled irritants. [ 20, 21 ] The Alveolar epithelium consists of type 1 and type 2 pneumocytes, responsible for gas exchange and surfactant secretion, respectively. Type 2 cells also retain inducible progenitor cell properties and, if the alveolar epithelium is damaged, can differentiate into type 1 cells. [ 22 ] 1. 2 The Respiratory Immune System It is increasingly recognized that the respiratory immune system plays a fundamental role in maintaining epithelial barrier integrity and lung homeostasis, with disruption leading to the development of inflammation and disease. Additionally, the immune system contributes to barrier and protective functions through the continuous sampling of the airway lumen for non‐harmful, immunogenic, or pathogen derived antigens. [ 23 ] Table 1 summarizes the main airway epithelial and immune cell types responsible for epithelial barrier integrity and protection, with a brief description below. Table 1 The main airway epithelial and immune cell types responsible for epithelial barrier integrity and protection Airway epithelial cell Epithelial barrier cell function Location in respiratory tract Pathogen defense role Goblet cell Mucin production Proximal, distal airways and submucosal glands. Mucin directly binds/traps pathogen and cell debris; Initiates microbial phagocytosis. [ 42, 43 ] Clara cell Surfactant production Proximal and distal airways. Surfactant directly binds/traps pathogen and cell debris; activates immune cells; initiates opsonization for pathogen clearance; Initiates microbial phagocytosis. [ 44, 45 ] Ciliated cell Ciliary movement and clearance of mucus Proximal and distal airways. Involved in the Muco‐ciliary clearance mechanism and physical removal of cell debris and pathogens from respiratory tract. [ 46 ] Alveolar type 2 cell Surfactant production and inducible progenitor for type 1 alveoli cells Alveoli. Surfactant directly binds/traps pathogen and cell debris; activates immune cells; initiates opsonization for pathogen clearance; Initiates microbial phagocytosis. [ 44, 45 ] Airway Immune cell Immune component Location in respiratory tract Pathogen defense role Dendritic cells (DC) Innate immune system Conducting airways and alveoli. Send extensions trough mucosal epithelium to sample airway. Can migrate to regional lymph nodes, once activated. Local non‐specific inflammation; Detection of antigens; antigen presentation and priming of adaptive immune response. [ 35, 47 ] Neutrophil Innate immune system Conducting airways and alveoli. Phagocytosis; release of cytotoxic granules and neutrophil extracellular traps for pathogen entrapment; promotes recruitment of adaptive and innate immune system. [ 30 ] Natural killer (NK) cell Innate and adaptive immune system Conducting airways and alveoli. Directly binds infected cells and promotes lysis/apoptosis; releases cytotoxic granules; promotes adaptive immune response. [ 27 ] Macrophage Innate immune system Alveoli (90%) and conducting airways (10%). Quiescent macrophages attach to epithelial cells, activated macrophages circulate in airways. Quiescent macrophages suppress the overstimulation of immune system; activated macrophages secrete cytokines, stimulate dendritic cells and phagocytose cell debris and pathogens; can also present antigens in some cases. [ 20, 29 ] T‐cell Adaptive immune system. Naïve T‐cells can differentiate into regulatory, helper, cytotoxic or memory T‐cells Naive T‐cells located in lymph nodes and lymph tissue. Once activated, can circulate throughout airways and alveoli. Regulatory T‐cells suppress the overstimulation of immune system; Helper T‐cell, for example, CD4+T regulate the adaptive immune response, especially B‐cells and macrophages; cytotoxic T‐cell, for example, CD8+ bind and lyse infected cells; memory T‐cells remain and circulate after infection to ensure rapid response to reinfection. [ 35, 37, 48 ] B‐cell Adaptive immune system. Naïve B‐cells can differentiate into plasma cells or memory B‐cells Naive B‐cells located in lymph nodes and lymph tissue. Once activated, can circulate throughout airways and alveoli. Plasma cells secrete specific antibodies which neutralize pathogens or bind and lyse infected cells; memory B‐cells remain and circulate after infection to ensure rapid response to reinfection. [ 38, 49 ] John Wiley & Sons, Ltd. This article is being made freely available through PubMed Central as part of the COVID-19 public health emergency response. It can be used for unrestricted research re-use and analysis in any form or by any means with acknowledgement of the original source, for the duration of the public health emergency. As mentioned, airway epithelial cells (AECs) provide a physical barrier against the environment, but these cells also secrete a range of effector and regulatory molecules. These may take the form of mucins and surfactant proteins, which directly bind infectious agents and cell debris, [ 23 ] or reactive species, such as nitric oxide (NO), which may influence smooth muscle contraction [ 24 ] and activation of the adaptive immune response ( Figure 2 A, B ). [ 23 ] AECs and dendritic cells (DC) display a range of specialized receptors capable of detecting self from non‐self antigens. [ 25 ] Activation of these specific pattern‐recognition receptors initiates various immunogenic and pathogen clearance mechanisms including the early inflammatory response, recruitment of innate immune cells, and activation of the adaptive immune response. Figure 2 Respiratory immune cell activation in response to pathogen invasion. Airway epithelial cells (AECs) and dendritic cells (DCS) continually sample airway lumen for either airborne pathogens or allergens. Activation of specific pattern‐recognition receptors on the surface of DCs initiate an inflammatory cascade in the early stages of pathogen invasion, inducing chemokine, cytokine, and immunoregulatory compound, for example, nitric oxide (NO) production. A) Adaptive immune cells are also recruited to the site of infection and contribute to the inflammatory response as well as modulating the adaptive immune response. B) In the absence of mucus producing epithelial cells in the alveoli (which would otherwise slow down gas exchange), respiratory macrophages are the main resident immune cell type, performing a protective and phagocytic role. C) Antigen presenting DCs migrate to the lymph nodes, located throughout the proximal and distal lung regions, and prime naïve adaptive immune cells. Activated B‐ and T‐cells then migrate to the site of infection or remain in peripheral circulation as memory cells. Image created using BioRender. com. During initial stages of pathogen invasion, the early inflammatory response is driven by the production of inflammatory chemokines and cytokines alongside recruitment of neutrophils, dendritic cells, and natural killer (NK) cells to the site of infection (Figure 2A ). [ 26, 27 ] Phagocytic macrophages may also be recruited, however, the majority of respiratory macrophages reside in the alveoli (Figure 2B ) rather than the conducting airways. Neutrophils, DCs, and macrophages are capable of working synergistically upon pathogen infection to promote airway inflammation, cytokine secretion, and lysis of viral‐infected cells. [ 20, 28, 29 ] Additionally, these innate immune cells are involved in the modulation of the adaptive immune response via induction of T‐cells and enhancement of DC recruitment. [ 30, 31, 32 ] Induction of the adaptive immune response is also propagated through the ability of DCs to undergo a phenotypic change to present antigens. [ 33, 34 ] Indeed, in response to respiratory infection, antigen‐presenting DCs migrate to regional lymph nodes where they prime naive adaptive immune cells for differentiation and proliferation (Figure 2C ). [ 35 ] T‐cell populations mitigate pathogen invasion via mechanisms specific to cell phenotype, including regulatory, cytotoxic, helper, and memory T‐cell populations (Table 1 ). Briefly, regulatory T‐cells are responsible for homeostatic regulation of the adaptive immune system [ 23 ], while cytotoxic T‐cells directly bind and lyse infected cells. Memory T‐cell population remain in blood circulation, lymphoid or lung tissues, with lung specific memory T‐helper cells contributing to viral‐mediated immunity upon reinfection. [ 36 ] B‐cell populations, once primed, mitigate pathogen invasion via the production of specific antibodies, which induce lysis and apoptosis. [ 37 ] Specific memory B‐cell populations also remain as long‐lived plasma cells which persist in a quiescent state in many tissues. [ 38, 39 ] Adaptive memory immune cell populations decline over time, with the rate of decline dependent on pathogen type and environmental conditions, meaning the potential loss of immunity over time. [ 40, 41 ] It is also important to mention a unique immune component, specific to mucosal surfaces such as the lung and gut: the mucosal immune system, also referred to as mucosal associated lymphoid tissue. In the event that pathogens evade the physical cellular barriers of the respiratory system, mucosal tissue has unique innate and adaptive immune mechanisms, similar, but separate from the peripheral lymphoid system. [ 16, 23 ] Thus, the mucosal immune system provides an additional protective layer in respiratory infections. 2 Modelling the Respiratory System In Vitro 2. 1 Cell Types As with the study of other biological systems, murine models are the most extensively studied in respiratory homeostasis, pathology and immune regulation. Indeed rodent models offer a complete, functioning biological system. However, since the introduction of the 3Rs principles (reduce, replace, refine), originally proposed in 1959, [ 50 ] together with the cosmetic testing ban of 2013, [ 51 ] there have been increasing ethical concerns surrounding animal use for scientific research. Furthermore, rodent models often lack clinical translatability, with high drug attrition rates seen in many phase III clinical trials. [ 8, 9, 10 ] Thus, human derived in vitro models offer an alternative for bridging the translational gap and have been increasingly researched and developed in recent years. However, given the complexity and cellular heterogeneity of the repository epithelium throughout the airways, the specific pathogen, effector location, and diseased/healthy phenotype should inform the cell types chosen to model respiratory host–pathogen interactions. Many lung and immune‐derived cell lines are available for culture, with the most commonly used listed in Table 2. Among the lung derived cell types, Calu‐3 and A549 cell lines are most widely used. Calu‐3 cells are derived from the submucosal gland of a human cancer patient, express vast numbers of goblet cell markers, differentiate into multiple cell types when cultured at the air liquid interface (ALI), and are useful for studying mucus production and mucociliary dysfunction. A549s, although commonly used to model the alveolar epithelial barrier, are derived from type 2 pneumocytes. These cells are secretory in nature and do not contribute largely to barrier formation. [ 52 ] Thus, other alveolar cell models, that represent the barrier forming type 1 pneumocytes, would be better suited for permeability, diffusion or barrier disruption experiments. To date, only one cell line is available for modelling alveolar type 1 pneumocytes (hAELVi cells), [ 53 ] with other attempts mainly involving the isolation and culture of type 2 cells to give type 1‐like cell phenotypes. Important to note, however, is the derivation of cell lines from cancerous tissue and their phenotypic representation of limited cell types. Thus, the use of primary cells is preferable in representing different cell type populations, signaling interplay and the patient heterogeneity found in vivo. However, primary cells are in limited supply and are more difficult to culture. The respiratory immune system must also be represented for in vitro models to fully reflect the in vivo respiratory barrier environment. Among immune‐derived cell types, the most widely used are those obtained from peripheral blood monocytes such as macrophages [ 54 ] and dendritic cells. [ 55 ] It is also possible to obtain tissue specific, resident immune cells such as those adhered to the epithelium or parenchyma of lung biopsies, the most commonly derived being alveolar macrophages. [ 56 ] However, the process of isolating these cells is much more time consuming and complex compared with obtaining them from blood. Immune cell lines which are derived from bone barrow of cancer patients are also available and, depending on the culture and stimulant conditions, it is possible to direct their differentiation into multiple myeloid cell types. Although this pluripotency may be advantageous in obtaining and representing multiple immune cells types, the cancerous nature of their origins will likely not reflect healthy phenotypes found in vivo. The cell type(s) one chooses for modelling the lung depends on the specific application and experimental question to be addressed. Important points to consider are: Location along the respiratory track, cell phenotypes and populations to be represented, importance of epithelial barrier formation and type of cell secretion or immune‐cell signaling pathways under study. Table 2 Cell types used for in vitro respiratory models Name Cell type Cell origin Use in modeling specific cell types Respiratory system‐derived HNE [ 57 ] Primary cell Human primary nasal epithelial cells from patient brushings. Nasal epithelial cells. NHBE [ 58 ] Primary cell Primary human bronchial epithelial cells from patients. Bronchial epithelial cells. Calu‐3 [ 59 ] Cell line Human adenocarcinoma cell line from 25‐year‐old male patient. Bronchial epithelial cells. 16HBE140 [ 60 ] Cell line Human bronchial epithelial cell line from a 1‐year old male lung/heart transplant patient. Bronchial epithelial cells. A549 [ 61 ] Cell line Human adenocarcinoma cell line from 58‐year‐old male. Alveolar type 2 cells. hAELVi [ 53 ] Cell line Human alveolar epithelial cell line. Alveolar type 1 cells. hAEpC [ 62 ] Primary cell Isolation and culture of type 2 human carcinoma alveolar epithelial cells. Alveolar type 2 and type 1‐like cells. TT1 [ 63 ] Cell line Transduced human type 2 carcinoma cells line (type 1‐like phenotype). Alveolar type 1 cells. NCl‐H441 [ 64 ] Cell line Human type 2 carcinoma cell line. Alveolar type 2 cells. Immune system‐derived Use in modeling disease‐associated inflammatory pathways Macrophage [ 54 ] Primary cell Human peripheral blood monocytes. Macrophage induced phagocytosis and inflammation. Dendritic cell [ 55 ] Primary cell Human peripheral blood monocytes. Dendritic cell induced inflammation. Neutrophils [ 65 ] Primary cell Human peripheral blood. Neutrophil induced inflammation. Alveolar macrophage [ 56 ] Primary cell Human lung tissue or bronchoalveolar lavage fluid. Macrophage induced phagocytosis and inflammation. HL‐60 [ 66 ] Cell line Human acute promyelocytic leukemia cell line from a 36 year old women patient. Spontaneous and directed differentiation into neutrophilic, monocytic, eosinophilic, and macrophage phenotypes. THP‐1 [ 67 ] Cell line Human acute monocytic leukemia cell line from a 1 year old male patient. Spontaneous and directed differentiation into neutrophilic, monocytic, eosinophilic, and macrophage phenotypes. HMC‐1 [ 68 ] Cell line Human acute systemic macrocytosis cell line. Mast cell induced inflammation. LADR [ 69 ] Cell line Human acute systemic macrocytosis cell line. Mast cell induced inflammation. John Wiley & Sons, Ltd. This article is being made freely available through PubMed Central as part of the COVID-19 public health emergency response. It can be used for unrestricted research re-use and analysis in any form or by any means with acknowledgement of the original source, for the duration of the public health emergency. 2. 2 Traditional Model Systems for Studying Respiratory Host–Pathogen Interaction Having given an overview of the respiratory epithelial system, lung and immune cell types available, we now consider the range of traditional in vitro models available for studying the lung microenvironment, each with their own merits, drawbacks, and benefit–cost ratio. Of important consideration is the societal need for having representative, reproducible in vitro platforms for the efficient discovery of virulence mechanism and development of antiviral vaccines for any future novel pathogens. Table 3 gives an overview of the traditional lung models available which are also briefly discussed below. Table 3 Most common in vitro respiratory models to study host–pathogen interactions Model type Advantages Disadvantages In vitro example of host pathogen interaction Cell type(s) used Submerged cell line culture – Easy to culture. – 2–5 days culture period. – Less skill required. – Readily available/cheap. – Representative of one cell type only. – Usually a cancerous cell line. – 2D culture. – Not representative of air interface. Respiratory syncytial virus [ 70 ] Bronchial cell line (BEAS‐2B); Primary human nasal and bronchial epithelial cells. ALI monoculture – Representative of air interfaced condition found in vivo. – Permits the study of viral entrance and metabolic pathways apically and basally. – More expensive. – 3–4 weeks culture period with primary cells. – 2D architecture. SARS‐CoV [ 72 ] Primary human alveolar type II cells. SARS‐CoV [ 92 ] Calu‐3 cell line. ALI co‐culture – Most biomimetic static cell culture available. – Representative of multiple cell types and systems found in vivo. – 2. 5D architecture. – High level of skill needed to culture. – 4–6 weeks culture period with primary cells. Aspergillus (A. ) fumigatus [ 73 ] Human primary bronchial epithelial cells, small airway cells, human blood derived macrophages, and dendritic cells. Polymer scaffolds – Ability to house multiple cell types. – 3D architecture. – High level of skill and precision needed to slice and culture. – Difficulty in monitoring cells within structure. Influenza A [ 115 ] Human primary small epithelial cells. Papain (mimics air bourne allergen) [ 81 ] Calu‐3 epithelial cell line, MRC‐5 fibroblast cell line, blood‐derived dendritic cells. Organoids – Derived from stem cells. – Representative of the integrated tissue found in vivo. – 3D structure. – High level of skill needed to culture. – 3–5 weeks culture period. – Cant access/monitor apical and internal cell types without disrupting. Parainfluenza [ 86 ] Human embryonic stem cells. Respiratory syncytial virus [ 85 ] Human embryonic stem cells. Multiple emerging influenza virus [ 91 ] Tissue resident adult stem cells. Precision cut lung slices (PCLS) – Fully differentiated tissue. – Representative of the heterogeneous phenotypes of population. – 3D architecture. – Culture times are less than that of ALI culture. – High level of skill and precision needed to slice and culture. – Expensive and limited supply. Influenza [ 88 ] Healthy lung slices from cancer patients undergoing lung resection. Rhinovirus [ 89 ] Healthy and asthmatic lung slices from patient donors. LPS (mimics bacterial infection) [ 90 ] Lung slices from patients with a variety of medical conditions from the National Disease Research Interchange. Cartoon insets created using BioRender. com. John Wiley & Sons, Ltd. This article is being made freely available through PubMed Central as part of the COVID-19 public health emergency response. It can be used for unrestricted research re-use and analysis in any form or by any means with acknowledgement of the original source, for the duration of the public health emergency. 2. 2. 1 2D In Vitro Models Culturing a submerged cell line in 2D offers a relatively cheap and quick culture method (typically 3–5 days) which may be advantageous for high throughput screening and assay development. For example, a cell line may be cultured in a well plate, inoculated with an isolated virus of varying MOIs (multiplicity of infection) and plaque forming units measured from cell supernatant, [ 70 ] with the whole assay taking less than a week to perform. However, submerged cell culture does not reflect the native air interface of the respiratory system which can influence the differentiation and growth processes of cell culture. [ 58, 61, 71 ] Air‐interfaced cultures are most appropriate in this context, with the supply of nutrients both apically and basally during differentiation, and air lifting post‐differentiation, reflecting the environments and processes found in vivo. Additionally, ALI culture grown in a Transwell configuration provides a more in depth analysis of viral entrance. For example, it is possible to inoculate with infected serum, or using aerosol deposition atop the cell culture, mimicking entrance in vivo. Collecting cell supernatant both apically and basally, then permits the study and spatial identification of cell specific entrance with methods such as 3D immunofluorescence rendering and quantification as well as RNA extraction, viral plaque forming assays, and scanning electron microscopy. [ 72 ] Culturing primary cells in Transwell configurations, and the formation of a pseudostratified epithelium, permits patient‐ and disease‐specific studies of response to infection and therapeutics in 2. 5D. Multiple cell types, such as epithelial, endothelial, or immune cells can also be co‐cultured on either the basal or apical side on the Transwell filter insert, representing a more complex and complete model. Additionally, by using functional confocal microscopy and live capture video analysis, it is possible to obtain pathogen‐induced measures of immune cell recruitment, receptor entrance, and transmigration through the membrane and cell layers. [ 73 ] Other gold‐standard assays involve monitoring epithelial barrier integrity during pathogen challenge using trans‐epithelial electrical resistance (TEER) [ 74 ] or ionic conductance. [ 75 ] However, important to note, is that primary cells are of limited supply and require a much longer culture, inoculation, and treatment period than cell lines (typically 4–6 weeks). 2. 2. 2 Toward 3D In Vitro Models In contrast to 2D models, 3D models more accurately represent the physiological architecture found in vivo. For example, it is possible to provide structural, mechanical, and spatiotemporal cues to the biological system, factors known to guide developmental and differentiation processes. [ 76 ] Additionally, it is possible to recreate and study cell–cell and cell–extracellular matrix (ECM) interactions. An example of this, is the use of hydrogels in a range of tissue specific 3D models including lung organogenesis, [ 77 ] tumorigenesis, [ 78 ] and airway scaffolds. [ 79 ] Hydrogels may be natural or synthetic and can be chemically, mechanically, and physically tuned to their specific application. Synthetic hydrogels are made from materials such as polyethylene glycol, polylactic acid, poly(lactic‐co‐glycolic acid), polyvinyl alcohol, and polycarprolactone. Natural hydrogels are made from a combination of polysaccharides, such as alginate, hyaluronic acid, agarose, chitosan, dextran, and cellulose, and proteins such as collagen, gelatine, fibrin, and poly‐ l ‐lysine (PLL). Matrigel is an example of a natural hydrogel, derived from Engelbreth‐Holm‐Swarm mouse tumor basement membranes, that is widely used in tissue culture applications. The chosen materials, however, are based on trade‐offs between biocompatibility, biodegradability, homogeneity, and mechanical durability, each having their own advantage. 2. 2. 3 Polymer Scaffolds Polymer scaffolds are commonly fabricated using electrospinning [ 80, 81 ] or phase separation and freeze drying techniques. [ 79, 82, 83 ] Scaffolds are seeded with lung cells and/or supporting immune and fibroblast co‐cultures, providing a biomimetic lung architecture that permits cell movement and ECM‐interaction. Furthermore, compared to 2D culture models, the use of scaffolds can increase the viability, differentiation, and expression of phenotypic markers found in vivo. [ 79, 83 ] 3D scaffolds have been used to model respiratory infection [ 115 ] and immune response to allergens in lung disease, [ 81 ] with changes in epithelial barrier permeability, gene expression post‐inoculation observed. It is also possible to fix or lyse scaffolds for detailed microscopy analysis to observe any cell‐scaffold interaction. [ 79 ] Although 3D polymer scaffolds pave the way for more representative models of lung tissue, they still hold some limitations such as heterogeneity of scaffold pore size and static cell culturing conditions. 2. 2. 4 Organoids In contrast to cellular models, organoids represent a fully differentiated 3D tissue structure. Lung organoid models are derived from human inducible pluripotent stem cells (iPSCs), embryonic stem cells, or ex vivo adult stem cells (ASCs) and may be grown at ALI or embedded within hydrogels. [ 84 ] Although 3D hydrogel organoid models are largely applied to the study of developmental processes [ 77 ] and signaling networks involved in the evolution of lung cancer, [ 78 ] they have been increasingly used in the field of respiratory disease, virology, and drug toxicology testing. Pathogen inoculation proceeds by applying a viral solute or aerosol on top of Matrigel embedded lung organoids, [ 85 ] while disease phenotypes may be induced by stimulating organoids with disease‐associated cytokine cocktails or via genetic modification of stem cells. [ 85 ] In response to infection, it is possible to image, in real time, the entrance site and migration of infection both locally, within specific cell types, and globally throughout the entire lung. [ 85, 86, 87 ] Furthermore, organoid models offer the advantage of being able to study cell–ECM interaction, an important consideration when evaluating immune cell and pathogen interaction within the native organ. Organoids may also be co‐cultured with human endothelial cells, [ 77 ] improving biomimicry and providing an opportunity to model vasculature‐organ‐ECM interactions and virulence of pathogens. However, a limitation still remains within these complex 3D organoid models in the inability to access or monitor the apical or inner epithelium of the organoid. 2. 2. 5 Precision Cut Lung Slices Another 3D model, representative of the native lung tissue, is precision cut lung slices (PCLSs). In contrast to cell and organoid culturing methods that require lengthy culturing times for differentiation, PCLSs offer the advantage of retaining native tissue structure and specific macrophage populations. PCLSs have been used to study respiratory pathogen virulence [ 88, 89, 90 ] as well as respiratory diseases, inflammation, and response to novel drug candidates. [ 91 ] However, like human primary cells, human PCLSs are limited in supply and last in culture for an average of 7 days compared to that of 21–28 days for ALI culture. Thus, they are unsuitable for long term exposure studies. Having given an overview and progression of traditional models for studying host respiratory processes, we now consider, in more detail, the most common types of respiratory pathogen and how host–pathogen responses may be modeled in vitro. 2. 3 Respiratory Disease and Infection Many chronic autoimmune and lung diseases display aberrant immune and epithelial barrier function as hallmarks of their pathology. Here, we focus on asthma and COPD which, worldwide, have the highest prevalence among respiratory diseases. Thus, understanding their underlying biomechanisms, co‐morbidities, and vulnerabilities are of high socio‐economic and therapeutic importance. Although there is evidence for asthma‐COPD overlap syndrome, [ 93, 94 ] highlighting the complex and interconnected mechanisms underlying their pathology, here the diseases will be discussed predominately in isolation. 2. 3. 1 Asthma Asthma is characterized as a chronic inflammatory condition with concomitant remodeling of the proximal and distal airways. [ 95 ] Although heterogeneities and subtypes exist, asthma can be broadly classified as intrinsic (non‐allergic) or extrinsic (allergic), [ 96 ] with clinical presentation of exacerbation including shortness of breath, wheezing, cough and, in severe cases, airway obstruction and respiratory failure. [ 95 ] Many reviews exist on the molecular, immunological, and pathological mechanism leading to remodeling of the airway epithelium in asthma [ 3, 4, 95, 97 ] ; however, here we highlight models used to represent asthma in vitro. Asthma may be modeled with the use of biopsies, [ 98 ] PCLSs, [ 99 ] or primary asthma cells grown at ALI. [ 100 ] Co‐culture of asthmatic primary and immune cells may be used to study the crosstalk between the epithelium, immune system, and inflammatory signals in disease progression. [ 100 ] Additionally, tri‐culture with epithelial, endothelial, and immune cell components may be used to represent blood vessel compartments and signaling interplay between these cell types. Perfused culture systems are also useful in mimicking the native environment under flow, and have been increasingly used to studying lung inflammation, fibrotic remodeling, and response to therapeutics. [ 101, 102 ] Additionally, these systems may be used to study differences in healthy versus diseased airway response to environmental triggers and drugs. Indeed, application of vaporized cigarette smoke, under flow, revealed previously undiscovered disease specific molecular signatures, potentially useful for future biomarker, and drug target studies. [ 103 ] Asthmatic in vitro models successfully recapitulate aspects of the airway environment found in patients, for example, displaying fewer epithelial tight junction protein complexes, increased permeability, and increased sensitivity to environmental triggers such as cigarette smoke. [ 104 ] Additionally, the use of patient samples means that the heterogeneity in disease severity will be represented when subjecting these systems to infection or novel therapeutics. This is highly relevant, given that respiratory infection is a major cause of asthmatic exacerbations. [ 105 ] Indeed, a body of evidence exists which argues that viral infection during childhood contributes to the initial pathogenesis of asthma. [ 106, 107 ] Thus, modeling the interactions between the asthmatic pro‐inflammatory environment, environmental triggers, and pathogen‐specific virulence, are fundamental in understanding and treating asthmatic population with an increased vulnerability to infection. 2. 3. 2 COPD COPD is characterized as a progressive and chronic inflammatory disease, occurring in all parts of the lung including airways, pulmonary vasculature, and lung parenchyma. [ 108 ] COPD shares some commonalties with asthma, for example, airway remodeling, chronic inflammation, and enhanced immune recruitment; however, COPD has its own defining features. In contrast to asthma, the airway remodeling that occurs in COPD is fibrotic, fixed, and irreversible [ 5, 108 ] and, among environmental factors, smoking has the largest influence on disease progression. [ 109 ] Substantial epithelial and endothelial apoptosis is present [ 110 ] and, in advanced stages, COPD exacerbations can lead to hyperventilation, hypoxia, and respiratory failure. [ 111, 112 ] Reviews exist which describe the inflammatory and molecular mechanisms behind COPD pathology in detail [ 5, 108, 113 ] ; however, here ways to model COPD are highlighted. Often, patient biological samples are taken in the form of blood, bronchoalveolar lavage, [ 76, 114, 115, 116 ] biopsies, [ 117, 118 ] and cell brushings. [ 79, 80, 119 ] These are used to study levels of inflammatory markers and immune cell activity, which may also form the basis for patient stratification and treatment. It is also possible to study markers of mucociliary clearance, such as levels of ciliary metaplasia [ 80, 120 ] and ciliary beat frequency, [ 79 ] the reduction of which may increase patient vulnerability to infection and sputum production. An altered respiratory microbiota is also implicated in COPD pathophysiology, [ 121 ] with acute exacerbations linked to microbial–pathogen interactions and infection. [ 122, 123, 124 ] Mechanisms underlying these findings may be recapitulated using in vitro models of COPD, cultured with primary cells. For example, oxidative mechanisms, [ 125 ] and an enhanced inflammatory environment [ 126 ] have been shown to augment epithelial cytokine and specific recognition receptor expression in viral‐induced COPD exacerbations. Additionally, to account for the increased risk of smokers developing COPD and viral‐induced exacerbations, mechanisms behind smoke induced epigenetic changes in bronchial epithelium have also been explored, such as an increase in mesenchymal markers [ 127 ] and a decrease in antiviral cytokine expression. [ 128 ] These in vitro COPD models also serve the purpose of high throughput drug development for disease related complications, such as viral induced exacerbations, [ 127 ] highlighting the potential for personalized therapeutics based on disease heterogeneity and severity. So far we have considered the respiratory epithelial and immune systems in health and in diseases such as COPD and asthma. Next, we discuss specific pathogen virulence mechanisms, the altered host response to infection, and how this interplay may be modeled in vitro. 3 Respiratory Pathogens The presence and accumulation of pathogens within the respiratory system can perturb homeostasis by overcoming the epithelial barrier and eliciting an immune response. Despite the overlap of symptoms and clinical manifestations of respiratory infections, individual pathogen types and species have distinct modes of entrance and virulence ( Table 4 ), with the host environment and health status influencing severity and vulnerability to infection. Respiratory infection may be modelled in vitro by inoculating the cell culture system with isolated pathogen particles or with immunostimulants which mimic pathogen‐specific inflammatory processes. Important considerations in choosing a model pathogen is whether the cell model expresses the relevant pathogen‐specific receptor, the respiratory location which the pathogen infects, and the experimental readouts you wish to use to assess virulence. Here, a brief overview of the most prevalent viral, bacterial, and fungal respiratory pathogens are given, together with their application in in vitro culture systems for studying respiratory host–pathogen interaction. Table 4 The most common viral, bacterial, and fungal pathogens known to cause repository infection Pathogen Clinical symptoms (complications) Respiratory tract infected part Pathogen entrance mechanism Viral MERS‐CoV Fever, chills, sore throat, cough, shortness of breath, headache, vomiting, diarrhoea, myalgia (pneumonia, septic shock, severe acute respiratory distress syndrome, respiratory failure, multi‐organ failure). Upper and lower respiratory tract. Cell mediated membrane fusion or endocytosis via CD26 receptors. [ 131 ] SARS‐CoV Fever, chills, myalgia, shortness of breath (pneumonia, fibrosis, severe acute respiratory distress syndrome, respiratory failure). Upper and lower respiratory tract. Cell mediated membrane fusion or endocytosis via ACE2 receptors. [ 132 ] SARS‐CoV‐2 (COVID‐19) Fever, chills, cough, shortness of breath, sore throat, rhinorrhoea, temporary anosmia or ageusia (pneumonia, septic shock, severe acute respiratory distress syndrome, respiratory failure, multi‐organ failure). Upper and lower respiratory tract. Cell mediated membrane fusion or endocytosis via ACE2 receptors. [ 133 ] Seasonal influenza Fever, sore throat, cough, headache, rhinorrhoea, myalgia, headache, (laryngotracheobronchitis, bronchitis). Upper respiratory tract. Cell mediated membrane fusion via sialic acid containing receptors and protease cleavage. [ 136 ] Respiratory syncytial virus (RSV) Fever, sore throat, cough, headache, rhinorrhoea, shortness of breath, wheezing, (laryngotracheobronchitis, bronchitis). Lower respiratory tract. Cell mediated envelope fusion via nucleolin containing receptors. [ 137 ] Rhinovirus Sore throat, cough, rhinorrhoea (bronchitis). Upper respiratory tract. Cell mediated endocytosis via ICAM‐1, LDL or CDHR3 receptors. [ 169 ] Bacterial Streptococcus pneumoniae Fever, chills, cough, shortness of breath, chest pain, (pneumonia, septic shock, bacteraemia, meningitis). Forms part of upper respiratory tract flora but can migrate and cause infection in lower respiratory tract and/or spread systemically. Extracellular colonization; polysaccharide capsule promotes adherence and protection. [ 144 ] Haemophilus influenzae Fever, chills, cough, shortness of breath, chest pain, (pneumonia, bronchitis, septic shock, bacteraemia, meningitis). Forms part of upper respiratory tract flora but can migrate and cause infection in lower respiratory tract and/or spread systemically. Internalization by epithelial cells via micropinocytosis and rearrangement of epithelial cytoskeleton [ 141 ] ; internalization by macrophage and neutralizes lysosomes to prevent detection or lysis. [ 147 ] Mycobacterium tuberculosis Fever, chills, chest pain, cough, weight loss (meningitis, respiratory failure, multi‐organ failure). Lower respiratory tract and can spread systemically. Internalization by macrophages via phagocytosis and neutralizes lysosomes to prevent detection or lysis; able to survive indefinitely but erupts to cause infection when host is immunocompromised. [ 150, 151 ] Fungal Aspergillus (mold; most common species A. fumigatus) Fever, chills, shortness of breath, wheezing, headache, cough, (Rhinitis, bleeding of the lungs, systemic infection, and multi‐organ failure). Upper and lower respiratory tract can spread systemically. Can invade tissues by extending hyphae through endothelial and epithelial barriers. [ 142, 168 ] John Wiley & Sons, Ltd. This article is being made freely available through PubMed Central as part of the COVID-19 public health emergency response. It can be used for unrestricted research re-use and analysis in any form or by any means with acknowledgement of the original source, for the duration of the public health emergency. 3. 1 Viral Pathogens Viral infection may be replicated in vitro by incubating cell culture systems with an immunostimulant which mimics viral inflammatory processes such as polyinosinic:polycytidylic acid (Poly I:C), [ 129 ] or by isolating live viruses and administering them in serum or aerosol deposition. Viral isolation first requires sampling and collection from an infected biological specimen, such as a nasal swab, which is then grown in vitro by infecting cells (typically mammalian cells), as viral replication requires a host. Media from infected cells can then be collected and separated from cells via filtration or centrifugation, as a source or virus particles. [ 130 ] Common viral pathogens cultured in this way include Corona viruses, influenza, respiratory syncytial virus (RSV), and rhinoviruses which are listed in Table 4 and briefly discussed below. Corona viruses are classified into four types (alpha, beta, gamma, and delta) with Middle Eastern respiratory syndrome coronavirus (MERS‐CoV), severe acute respiratory syndrome Coronavirus (SARS‐CoV), and SARS‐CoV‐2; all of different lineages within the beta category. Corona viruses infect epithelial cells of the upper and lower respiratory tract via viral spike protein binding and cleavage by host cell proteases. MERS‐CoV enters via the CD26 receptor [ 131 ] while both SARS‐CoV [ 132 ] and SARS‐CoV‐2 [ 133 ] enter via the angiotensin converting enzyme 2 (ACE 2) receptor. In the case of novel pathogens, such as recently emerged corona viruses, it is essential to recapitulate infection in a representative in vitro model, to gain insight into the mechanisms of transmission, pathogenesis, and possible targets for vaccines. The specific mode of entrance and virulence have been studied in coronaviruses using cell lines, [ 92, 134 ] primary cells, [ 72, 134 ] and patient biopsies. [ 135 ] For example, the apical entrance of coronavirus in bronchiole epithelial cells, via the ACE 2 receptor, has been shown by protein co‐localization in confocal Z‐Stack immunofluorescence imaging. [ 92 ] Apical entrance and release of virions may also be demonstrated by sampling supernatant from apical and basolateral serum as well as via transmission electron microscopy imaging ( Figure 3 A ). Figure 3 Examples of in vitro models used to study entrance and virulence mechanisms of A) viral, B) bacterial, and C) fungal pathogens. A) Apical entry and release of severe acute respiratory syndrome‐associated coronavirus in polarized Calu‐3 lung epithelial cells. Above: Transmission electron microscopy of release of SARS‐CoV virons from the apical surface of polarized Calu‐3 cells. Below: Colocalization of ACE‐2 and viral antigen in infected Calu‐3 cells, both ACE‐2 (green) and viral antigen (red) could be detected in infected cells. Importantly, both ACE‐2 and viral antigen appeared to colocalize in infected cells (yellowish). Reproduced with permission. [ 92 ] Copyright 2005, ASM. B) Infection of primary human bronchial epithelial cells by Hemophilus influenzae. Above: Images collected by dual‐wavelength CLSM of cells infected for 3 h; colocalization of airway nuclei, bacteria (green) and vacuoles (red) can be seen in yellow, suggesting bacteria have been taken into the cells. Scale bar is 50 µm. Below: The series (A through F) demonstrates lamellipodia surrounding bacteria (black arrow) at the surface of a submerged airway cell culture after 4 h of infection. Reproduced with permission. [ 141 ] Copyright 1999, ASM. C) Polarized response of endothelial cells to invasion by Aspergillus fumigatus. A. fumigatus hyphae invade the abluminal and luminal surface of endothelial cells by different mechanisms. Above: Hyphae invading the abluminal surface of endothelial cells, Arrows indicate an endothelial cell that is being invaded by a hypha. Below: Hyphae invading the luminal surface, arrows indicate endothelial cell pseudopods. Hyphae are shown in green and microfilaments in red. Bars represent 5 µm. Reproduced with permission, [ 142 ] Copyright 2009, Wiley. The inset cartoon schematics represent the type of model chosen. Cartoon insets created using BioRender. com. Influenza viruses infect epithelial cells of the upper respiratory tract, via binding of viral hemagglutinin to sialic acid containing receptors of target cells. [ 136 ] The symptoms elicited following infection are due largely to the release of proinflammatory cytokine and chemokines for example interferons and tumor necrosis factor from viral‐infected cells. In vitro, emerging strains of influenza may be studied in order to elucidate replication and infectivity mechanisms as well as strain specific cytokine/chemokine profiles. [ 87, 88 ] RSV infects cells of the lower respiratory tract via binding of viral fusion glycoprotein with Nucleolin containing surface receptors of target cells. [ 137 ] RSV is easily transmitted and is a major cause of respiratory infection in children and infants. In vitro, the link between RSV virulence, airway hyperresponsiveness, and the production of specific cytokine profiles may be modeled using cell lines, [ 138 ] organoids, [ 85 ] or primary culture derived from pediatric patients populations. [ 139 ] RSV infection of lung specific immune cells has also be used to study cross‐talk between immune cell and epithelial components for both pathogen virulence and protection mechanisms. [ 70 ] Rhinoviruses are one of the most common causes of the common cold and exacerbations in lung disease such as asthma. Rhinoviruses have three species (A, B, and C) with infection occurring in epithelial cells of the upper respiratory tract. Rhinovirus induced asthmatic exacerbations may be modeled in vitro models by comparing healthy and asthma derived primary cells [ 125, 140 ] or lung slices [ 89 ] from patients and observing disease or patient specific inflammatory cytokine profiles as well as cell specific immune cell migration. Findings from in vitro studies such as these may then be replicated and correlated to in vivo investigations to assess translatability to the human condition, a principle factor in improving drug development and therapeutics in the clinical setting. 3. 2 Bacterial Pathogens Bacterial infection may be replicated in vitro by incubating cell culture systems with an immunostimulant which mimics bacterial inflammatory processes such as Lipopolysaccharides (LPS) or endotoxins. [ 143 ] In contrast to a live virus, bacteria do not require host cells for replication, rather, growth and isolation of specific strains may be acquired using selective agar or media. Common respiratory bacterial pathogens cultured in this way include Streptococcus pneumoniae, Mycobacterium tuberculosis, and Haemophilus influenzae (Table 4 ). S. pneumoniae commonly forms part of the upper respiratory tract flora and its presence is asymptomatic in most healthy individuals. However, under favorable environments or in compromised individuals, S. pneumoniae colonizes extracellular respiratory space, migrate to the lower respiratory tract and is the major cause of bacterial pneumonia in vulnerable patients. Virulence is associated with the release of invasion proteins such as pneumolysin, which contribute to host cell entrance and death via pore formation, toxin‐induced apoptosis or induction of host cell epigenetic changes. [ 144 ] S. pneumoniae infection is also shown to decrease mucocilary clearance mechanisms and induce epithelial autolysis in primary respiratory organoid and biopsy samples. [ 145 ] Cell line models have also been useful as a high‐throughput means for identifying novel targets and developing alternative treatments for resistant strains. [ 146 ] Similar to S. pneumoniae, H. influenzae may be present in the upper respiratory tract flora and is an opportunistic pathogen, causing infection in vulnerable or immunocompromised individuals by migrating to the lower respiratory tract and/or systemically. Virulence is caused by surface Lipooligosaccharides and lipoproteins, which when attached to the mucosal surface, exert disruptive effects on cilia function. [ 147 ] H. influenzae also produced proteases which help to evade macrophage induced lysis via mechanisms similar to that of M. tuberculosis. Virulence mechanism such as these have been studied in vitro by infecting cell lines. [ 148, 149 ] It has also been shown that infection occurs via the rearrangement of epithelial cytoskeletons and micropinocytosis, demonstrated by microvilli and lamellipodia extending and engaging with bacteria, and the presence of bacteria within vacuoles of epithelial cells, respectively (Figure 3B ). [ 141 ] M. tuberculosis infects the lower respiratory tract and is the causative agent of tuberculosis. Infection occurs via macrophagic phagocytosis and contaminant neutralization of lysosomes. M. Tuberculosis is able to lie dormant within these cells, erupt when the host is immunocompromised and even cause chronic infection. Virulence of M. Tuberculosis is associated with the production of toxins, such as tuberculosis necrotizing toxin, [ 150 ] encapsulation in a lipid containing coating, and participation in lysis‐evading mechanisms. [ 151 ] Replication mechanisms have been studied in human alveolar cell lines [ 152, 153, 154, 155 ] and in co‐culture with immune cell and ECM components. [ 156, 157 ] In addition to complications caused by primary bacterial infections, viral infection also increases the risk of developing a secondary bacterial infection, termed a bacterial superinfection. [ 47 ] Mechanisms behind this include viral‐induced desensitization of macrophages [ 158, 159 ] and an impaired neutrophil and monocyte response. [ 47 ] Viral‐induced epithelial damage may also facilitate the passage and colonization of bacterial pathogens within the respiratory tract and lung parenchyma. Furthermore, it is important to consider the respiratory microbiome in influencing a patient's susceptibility and response to infection. Many reviews exist which describe the complex interaction between the respiratory microbiome, epithelium, and immune system, [ 160, 161, 162, 163 ] but in brief, the microbiome is influenced by a range of early life experiences such as mode of delivery, environment, diet, and respiratory infection. Additionally, the presence of underlying disease, immunosuppression, or certain drug treatment may influence microbiota profile, potentially leading to an inflammatory environment. Under these conditions, commensal microbial species may become pathogenic, such as those mentioned above, for example, S. pneumoniae, H. influenzae. Conversely, commensal respiratory bacteria may also have a protective effect. For example, H. influenza, which is a common cause of respiratory infection in children, may offer specific protective roles against developing RSV. [ 164 ] Additionally, patient‐specific microbiota profiles have been linked to having protective affects against influenza infection and virulence. [ 165 ] Therefore, the interaction between commensal and pathogenic microbes, within the respiratory system, are an important consideration when assessing patient specific responses to infection and therapeutics. 3. 3 Fungal Pathogens Like bacteria, fungal species may live in symbiosis with a host and, although possible to inhale infectious fungal agents, most infections are of the opportunistic type, developing disease mainly in immunocompromised individuals. [ 166 ] Fungi replication occurs via spore spreading and, like bacteria, can be grown and isolated in vitro using selective agar or media. A. fumigatus is the most common respiratory fungal species and is associated with development of aspergillosis. Aspergillosis may take a variety of forms. Allergic aspergillosis occurs when patients experience an allergic reaction to fungal spores and is most common in patients with underlying inflammatory lung conditions asthma and cystic fibrosis. Acute invasive aspergillosis on the other hand, is the most severe form of the disease and occurs in immunocompromised patients when the infection spreads systemically to other organs. Virulence of A. fumigatus occurs through the production of toxins such as Aflatoxin and Gliotoxin which exert immunosuppressive effects including disrupting cilia function, inhibiting phagocytosis, and inducing apoptosis. [ 167 ] In vitro models of A. fumigatus have demonstrated Hyphae extensions are capable of penetrating pulmonary endothelial and epithelial cells as a mechanism of invasion. [ 142, 168 ] Additionally, hyphae invasion induces a polarized response in endothelial cells, such that luminal invasion occurs via endocytosis and the formation of pseudopods, whereas abluminal invasion occurs via the disruption of microfilaments (Figure 3C ). The evidence provided thus far encompasses studies which use traditional in vitro models of respiratory infection; however, in the hope of providing more relevant, biomimetic, and high throughput drug discovery platforms, a range of more advanced and technology integrated model systems are continuously being developed. These are discussed in detail below. 4 Advances in Technology Integrated Models for Studying Host Pathogen Interaction In parallel to the growing ethical concerns surrounding animal use in research and their lack of their clinical translatability, [ 8, 9, 10 ] there has been a surge in the development of technology integrated 3D in vitro models which better reflect the human in vivo lung condition. For example, it is possible to integrate previously static 2D, 2. 5D, and 3D models, for example, ALI co‐culture, organoids, etc. (as discussed in Section 4) with technological advances such as perfusion chambers [ 73, 170 ] and lung‐on‐chips. [ 101, 171, 172 ] Technology integrated biological systems have advanced knowledge surrounding the effect of culturing conditions and model architecture on relevant parameters such as cellular differentiation, immune cell recruitment, and cytokine profile, such that lung models are becoming increasingly, and more accurately, representative of the human condition. Here, we discuss the progression from perfusion bioreactor chambers to microfluidics and sensor integrated lung‐on‐chips, and how these have advanced our understanding of lung cell culture. 4. 1 Lung‐on‐Chip With the development of fluidics and commercially available perfusion chambers, it is possible to accelerate the speed of growth, differentiation, and development of 2D lung epithelial ALI models. Indeed, with perfusion systems, ciliogenesis, mucus production, and barrier formation are observed up to 14 days earlier when compared to static culture. [ 73 ] Systems such as these enable the fast‐track addition of immune co‐culture and pathogen infection studies, significantly shortening experimental protocol times without sacrificing the complexity of a 3D ALI model. In parallel, the revolutionary development of microfluidic organ‐on‐chip technology during the last decade permits the coupling of microfluidics with microsensor technologies. Indeed, in addition to applying effective shear stress and flow, which enhances cellular differentiation, [ 173 ] sample preparation, and delivery of nutrients, [ 101 ] it is possible to integrate on‐chip biosensors such as pH sensors, microscopes, and electrodes. [ 172, 174 ] In comparison to traditional culture systems, this enhances the speed of detection, breadth of readout data, device portability, and accelerates the point of care diagnostics. Organ‐on‐chips are commonly microfabricated from transparent and biocompatible polymers such as poly(dimethylsiloxane (PDMS) and poly(methlymethacrylate) (PMMA), via soft lithography‐based techniques, and consist of multiple layers of cell culture chambers integrated with microfluidic perfusion systems. A pivotal study carried out by Huh and colleagues in 2010 involved the development of a novel actuation system which mimicked breathing in the human lung. [ 171 ] This “breathing lung‐on‐chip” consisted of an upper and lower chamber, separated by a thin porous PDMS membrane. The upper compartment contained human alveolar epithelial cells cultured at air interface, while the lower compartment contained human microvascular endothelial cells co‐cultured with human neutrophils under dynamic flow. Together, these compartments represented the alveolar capillary lung unit which, under the application of a vacuum in adjacent chambers, underwent cyclically stretching representative of physiological breathing ( Figure 4 A ). Importantly, this study demonstrated the effect of breathing on enhancing inflammatory and immune signaling in primary co‐culture. For example, when exposed to air pollutant extracts, an increase in pro‐inflammatory adhesion molecules and reactive oxygen species was observed, when compared to static cell culture. Similar lung‐on‐chip models, which utilize an actuated micro diaphragm to induce mechanical breathing, also demonstrate a breathing‐induced increase in barrier permeability, metabolic activity, and wound healing in lung epithelial cells. [ 175 ] These examples highlight the importance of representing in vivo like breathing forces when assessing the extent of inflammation and immune cell activation that an airborne particle, pathogen, or drug candidate could have. Additionally, improvements in design and fabrication methods permit passive, rather than active, perfusion of chips. This reduces the need for additional external equipment and tubing, while improving reproducibility. [ 176 ] Modern chip technology is also becoming increasingly compatible with standard characterization methods such as TEER, enzyme linked absorbency assays (ELISA), and permeability assays, [ 176 ] making their integration into mainstream laboratories more amenable. Figure 4 Examples of advanced technology integrated in vitro models of the lung showing A) mechanical actuation, B) complex microfluidic airway design, C) compartmentalization of lung components and infectious agents and D) advanced electronic monitoring of ALI culture. A) Compartmentalized PDMS microchannels form the alveolar‐capillary barrier. The device recreates physiological breathing movements by applying vacuum to the side chambers and causing mechanical stretching of the PDMS membrane. Reproduced with permission. [ 171 ] Copyright 2010, AAAS. B) Anatomically inspired microfluidic acini‐on‐chip featuring an asymmetrical bipurification model of distal airways (blue arrows) and air‐ducts (red arrows). Reproduced with permission. [ 177 ] Copyright 2019, Wiley. C) A microbial culture insert is inoculated with A. fumigatus on the left and P. aeruginosa on the right, facilitating volatile factor contact between the microbial cultures and air‐exposed center lumens lined with bronchiolar epithelial cells. Scale bar is 250 µm. Reproduced with permission. [ 168 ] Copyright 2017, Nature Publishing Group. D) Effect of E Cigarette Emissions on Tracheal cells monitored at ALI using an organic electrochemical transistor. Integration of an ALI airway epithelium model into a flexible gate‐OECT platform for ALI resistance sensing, which conforms to the cell secreted mucus. Reproduced with permission. [ 178 ] Copyright 2019, Wiley. In another study, focusing on the effect of smoking induced respiratory inflammation and disease progression, a novel multi‐compartment robotic smoking machine was microengineered. [ 103 ] This replicated all major aspects of physiological breathing including mechanical inhalation of smoke, controlled respiration parameters (respiration cycle, puff time, and inter puff interval), and flow rate over an air interfaced lung‐on‐chip. Compared to traditional exposure protocols, which deposit cigarette smoke extracts on top of cell culture, this study applied whole cigarette smoke under physiologically flow. This novel protocol revealed novel disease specific molecular signatures, potentially useful for future biomarker and drug target studies. Additionally, this study gave a detailed insight into smoke‐induced changes in ciliary beat frequency distribution, which may be linked to reduced mucociliary clearance observed in smokers. With advances in fabrication methods, it is also possible to create complex 3D microchannel networks, within lab‐on‐chip systems. Indeed, Schnirman and colleagues [ 177 ] fabricated an anatomically inspired microfluidic model, mimicking the bifurcation networks of human alveolar tree structures which matched functional residue capacity values of pediatric populations (Figure 4B ). Although complex structural models such as these are technically challenging to implement, they are fundamental in replicating and simultaneously studying the full breadth of cell types present in all parts of the airway. Models such those described above more fully recapitulate the human condition compared to traditional cell models and, although difficult to implement, are essential in reducing animal research and drug attrition rates. Biomimetic 3D lung‐on‐chip models have also been increasingly applied to the study of pathogen invasion, disease‐associated inflammation, host–pathogen interaction, and therapeutic treatment of novel infectious agents. Unlike traditional in vitro models of pathogen invasion, which involve the direct measurement of pathogen‐induced effects upon epithelial layers, chip systems are able to physically compartmentalize and connect different cell and microbial populations. Thus, it is possible to study more complex interactions between the host and pathogen under physiological flow. It is also possible study the communication that occurs between the lung ECM, immune system, and circulating volatile compounds at ALI. Indeed, Barkal and colleagues [ 168 ] microengineered an innovative bronchiole‐on‐chip device which contained a central airway lumen and adjacent endothelial lumens connected via a fibroblast‐collagen matrix (Figure 4C ). A separate ”clickable” module, seeded with compartmentalized infectious microbials, was attached to the main lung unit. This was used to study pathogen‐derived volatiles on the respiratory epithelium. In this instance, co‐infection with the fungal and bacterial agents A. fumigatus and P. aeruginosa, respectively, was shown with hyphae extensions and leukocyte migration clearly observed at the site of infection (Figure 4C(III) ). Lung‐on‐chip devices have also been used to study lung epithelial permeability [ 172, 175 ] and single‐strain pathogen infection. [ 163, 171 ] Pathogen‐induced effects on lung epithelial permeability may be measured via TEER or passage of fluorescently labelled molecules through the epithelium. [ 172, 179 ] Additionally, cell effluent can be collected and assayed via ELISA or PCR for relative change in cytokine profile. [ 163 ] In the case of co‐culture, immune cell migration to the site of infection may be observed via high resolution and real‐time microscopy imaging. [ 171 ] As well as modelling inflammation and immune recruitment in healthy lungs, lung‐on‐chip devices are also used to model pathogen induced exacerbations in lung disease. [ 63, 64 ] Diseased phenotypes may be modeled by directly culturing primary cells from diseased patients or alternatively, inflammation can be induced by stimulating cells with inflammatory proteins or cytokines implicated in disease pathology. For example, allergenic asthma‐like lung inflammation can be induced with the cytokine IL‐13, which is known to induce airway‐hyperresponsiveness and goblet cell hyperplasia in vivo. [ 73, 158 ] These models can also be used as drug discovery platforms by applying novel therapeutics to microchannels and measuring effects on epithelial cell composition, function, and cytokine profile. Studies such as these, illustrate the importance of modelling complex aspects of the in vivo lung environment, such as physiological flow rate and breathing‐induced mechanical strain. Organ‐on‐chip technology paves the way for portable, multi‐parametric, and simultaneous assay platforms, which increasingly makes the study of respiratory pathogens in healthy and diseased human airways more efficient and accessible. 5 Future Directions Although substantial progress has been made in recent years toward 3D and technology integrated in vitro lung models, there remain some limitations or problems to address. For example, in the case of organoid or complex scaffold structures, there is limited capability in monitoring the cellular components found within the core of the 3D systems. Secondly, in the case of ALI cultures, the present gold standards for monitoring epithelial integrity, such as TEER, require the apical surface to be submerged in an electrolyte/media. This negates the advantages of ALI culturing method, as well as preventing any long term/real‐time TEER measurements of any ALI culture. Some novel innovations and future prospective, which address these limitations, are highlighted below. 5. 1 Conducting Polymer Scaffolds As mentioned above, the use of polymer scaffolds and hydrogels in 3D cell culture has proven advantages such as increased viability, differentiation, and the ability to study cell–ECM interactions. However, there also remain limitations in accurately assessing/monitoring the inner portions of these 3D structures. One solution to this is the fabrication of complex cell architectures within conducting polymers, permitting the electronic monitoring of enclosed cell populations. Interestingly, in a 3D tissue engineered tubular model, the fabrication of conducting polymer scaffolds demonstrated the ability to monitor cell adhesion, growth and migration in real‐time, via material‐integrated electronic sensing abilities. [ 180 ] This highlights the potential of scaffold systems to accurately monitor complex 3D architectures in a dynamic and mid‐throughput manner. One can see how this technology may be adapted or integrated into lung scaffolds for monitoring epithelial/endothelial permeability and immune cell adhesion/migration when performing pathogen challenge experiments. It is also possible to utilize hydrogels as biosensors by tuning them to detect pH, temperature, light, or electricity, [ 181 ] which can be particularly beneficial for use in microfluidic devices for creating on‐chip readout systems. The field of bioelectronics, discussed below, looks promising for future application in monitoring cell and tissue culture in a non‐invasive, label free, and real‐time manner. 5. 2 Advances in Flexible Electronics A technology capable of conforming and electronically monitoring a range of complex 3D architectures, lies in the field of bioelectronics. Indeed, parallel to the rise of biocompatible and wearable electronic sensors in medical and commercial settings, [ 182, 183 ] flexible electronics have also been implemented in a variety of in silico [ 184 ], in vivo, and in vitro research applications. [ 185 ] Of note are poly (3, 4‐ethlyenedioxythiophene) doped (p‐type) with poly (styrene sulfonate) (PEDOT:PSS)‐based electrodes or organic electrochemical transistors (OECTs). The detailed physical theory of OECT operation is explored elsewhere, [ 186, 187 ] but such technology has been integrated into a variety of biological formats, including Transwell ALI culture, [ 178 ] planar and microfluidic devices, [ 188, 189, 190, 191 ] PEDOT:PSS bio‐scaffolds, [ 192, 193 ] self‐rolling sensors, [ 194 ] and neuromorphic devices. [ 195 ] In vivo examples include bioresorbable patches [ 196, 197 ] and implantable electrocorticography devices for monitoring neuronal epileptiform discharge. [ 198, 199 ] In each of these applications, OECT devices have shown superior performance when compared to conventional electrode recordings, including lower operational voltages, increased signal‐to‐noise ratio (SNR), and increased biocompatibility. Furthermore, OECTs display high capacitance, low impedance, mechanical flexibility, chemical tunability, and optical transparency, making them ideal candidates for multiparametric sensing, simultaneous characterization with optical techniques and improved efficiency, and accuracy of data acquisition. [ 187, 200, 201 ] OECT devices have been used to study epithelial barrier formation and disruption, [ 202 ] stem cell differentiation, [ 203 ] and to detect analytes in human fluid samples for diagnostic purposes. [ 204 ] In the line of pathogen infection, the application of OECTs have been largely used to study food‐borne or bacterial infection of intestinal [ 205 ] and kidney cell lines. [ 206 ] In relation to the respiratory epithelium, OECTs have been applied to the study of E‐cigarette aerosol exposure on human tracheal barrier integrity in ALI cultures [ 178 ] (Figure 4D ) and conductivity of ion channels implicated in pulmonary disease. [ 207 ] Additionally, if biofunctionalized, OECTs, can achieve a high detection sensitivity of protein biomarkers, [ 208 ] cell surface glycans, [ 209 ] and human viruses [ 210 ] which demonstrates the capability of this technology in advancing host–pathogen interaction studies. 6 Conclusions Respiratory infection and related co‐morbidities are one of the leading causes of death worldwide, while also contributing a substantial socio‐economic burden. Given the recent SARS‐CoV‐2 pandemic, it has become increasingly evident that more efficient and biomimetic in vitro systems are needed to improve the efficacy, reproducibility, and translatability of therapeutics, antivirals, and vaccines. Here, we have given an overview of the biological and immunological components responsible for respiratory epithelial barrier integrity in health and disease. Furthermore, we have given an overview of the most common respiratory pathogens, as well as traditional 2D and more complex 3D in vitro models for studying host–respiratory pathogen interactions. Great improvements have been made in recent years in the fields of tissue engineering, material science, and biotechnology that have enabled the production of complex 3D models. For example, improvements in hydrogel composites have allowed for improved differentiation, proliferation, and longevity of cell/tissue culture. Developments in microfluidic and microfabrication techniques have also contributed crucial knowledge on the importance of mechanical, biochemical, and spatiotemporal cues in replicating an entire organ system. Additionally chip technology permits the integration of multiple biosensors in a compact design which offers advantages such as speed of processing, detection, breadth of readout data, and device portability. Finally, with the rise in the field of flexible electronic biosensors, which have the ability to physically conform to a range of complex 3D architecture, give multimodal, real time, and long term readouts, the future may see further integration of this technology with respiratory in vitro models. Conflict of Interest The authors declare no conflict of interest. Author Contribution S. L. B. , S. J. , and R. M. O, are responsible for the design of article contents. S. L. B, is responsible for the majority of the writing, design, and production of figures and tables. R. M. O. and J. S. provided feedback and guidance for the construction of the article. All authors have read the article and given approval.
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10. 1002/adbi. 202200267
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Advanced biology
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Bioengineering and Clinical Translation of Human Lung and its Components
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Clinical lung transplantation has rapidly established itself as the gold standard of treatment for end-stage lung diseases in a restricted group of patients since the first successful lung transplant occurred. Although significant progress has been made in lung transplantation, there are still numerous obstacles in the path of clinical success. The development of bioartificial lung grafts using patient-derived cells might serve as an alternative treatment modality; however, challenges include developing appropriate scaffold materials, advanced culture strategies for lung-specific multiple cell populations, and fully matured constructs to ensure increased transplant lifetime following implantation. This review highlights the development of tissue-engineered tracheal and lung equivalents over the course past two decades, key problems in lung transplantation in a clinical environment, the advancements made in scaffolds, bioprinting technologies, bioreactors, organoids, and organ-on-a-chip technologies. The review aims to fill the lacuna in existing literature towards a holistic bioartificial lung tissue, including trachea, capillaries, airways, bifurcating bronchioles, lung disease models and their clinical translation. Herein, the efforts are on bridging the application of lung tissue engineering methods in a clinical environment as we think that tissue engineering holds enormous promise for overcoming the challenges associated with clinical translation of bioengineered human lung and its components.
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No full text available
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10. 1002/adem. 201800166
| 2,018
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Advanced engineering materials
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Mechanical Properties of Graphene Foam and Graphene Foam – Tissue Composites
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Graphene foam (GF), a 3-dimensional derivative of graphene, has received much attention recently for applications in tissue engineering due to its unique mechanical, electrical, and thermal properties. Although GF is an appealing material for cartilage tissue engineering, the mechanical properties of GF – tissue composites under dynamic compressive loads have not yet been reported. The objective of this study was to measure the elastic and viscoelastic properties of GF and GF-tissue composites under unconfined compression when quasi-static and dynamic loads are applied at strain magnitudes below 20%. The mechanical tests demonstrate a 46% increase in the elastic modulus and a 29% increase in the equilibrium modulus after 28-days of cell culture as compared to GF soaked in tissue culture medium for 24h. There was no significant difference in the amount of stress relaxation, however, the phase shift demonstrated a significant increase between pure GF and GF that had been soaked in tissue culture medium for 24h. Furthermore, we have shown that ATDC5 chondrocyte progenitor cells are viable on graphene foam and have identified the cellular contribution to the mechanical strength and viscoelastic properties of GF – tissue composites, with important implications for cartilage tissue engineering.
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No full text available
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10. 1002/adem. 202000759
| 2,020
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Advanced Engineering Materials
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Rapid Fabrication of Sterile Medical Nasopharyngeal Swabs by Stereolithography for Widespread Testing in a Pandemic
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The 3D printing of nasopharyngeal swabs during the COVID‐19 pandemic presents a central case of how to efficiently address a break in the global supply chain of medical equipment. Herein a comprehensive study of swab design considerations for mass production by stereolithography is presented. The retention and comfort performance of a range of novel designs of 3D‐printed swabs are compared with the standard flocked‐head swab used in clinical environments. Sample retention of the 3D swab is governed by the volume, porosity density, and void fraction of the head as well as by the pore geometry. 3D‐printed swabs outperform conventional flock‐head swabs in terms of sample retention. It is argued that mechanically functional designs of the swab head, such as corkscrew‐shaped heads and negative Poisson ratio heads, maximize sample retention and improve patient comfort. In addition, available designs of swab shafts for an optimized sample collection procedure are characterized. The study is conducted in vitro, using artificial mucus, covering the full range of human mucus viscosities in a 3D‐printed model of a nasal cavity. The work sets the path for the resilient supply of widespread sterile testing equipment as a rapid response to the current and future pandemics.
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1 Introduction The challenge of the local and global supply chain to respond to potential catastrophes is very high, as shown by the Federal Emergency Management Agency (FEMA) Supply Chain Resilience Guide. [ 1 ] Responses to natural disasters, such as hurricanes, have shown that even with historical analyses that prefigure response strategies, there remain a lack of readiness, awareness, and organization between governmental relief and private‐sector supply networks in solving logistical bottlenecks, surges, and restoration. [ 2 ] The lack of preparedness is even greater when facing a pandemic, such as that of COVID‐19, as unprecedented catastrophe development means that we rely only on a small subset of data from past events whose scale is not comparable to the current event, such as past epidemics like SARS and MERS, to come up with ideal solutions. To meet supply chain demands, mass production pipelines need time to ramp up their efficiency to provide a reliable flow of provisions. To bridge this time gap and solve this tactical problem, local communities can support and provide to supply chain resilience strategies through classical rapid prototyping. As FEMA recommends, community lifelines can effectively be repurposed for “maintaining or restoring the most critical services or infrastructure. ” [ 1 ] In a matter of weeks, a local rapid short‐term response can be made possible through fast design and fabrication methods provided by rapid prototyping. Over a series of months, the industrial mass production methods can scale up to respond to a global demand of resources. For example, shortage in ventilators was addressed by enthusiast communities who developed cost‐effective, rapid prototyping‐based, open‐source designs of mechanized bag valve masks, [ 3, 4 ] all in a matter of few single weeks from the state‐of‐emergency announcement on March 13, 2020, immediately resulting in an National Institutes of Health approval pipeline. [ 5 ] Any person with basic mechanical engineering skills could implement and troubleshoot such a ventilator in a day, enabling one to tactically meet immediate local needs. Later, in a more global and strategical approach to the shortage of ventilators, Ford Motors adopted its production line for an approved third‐party commercially available design. Ford's production of ventilators started on April 20 first with 500 units, continued into May with 10 000 units, and has an expectation to produce 30 000 units per month moving forward. [ 6 ] It should be noted that to enable the maker communities to fill the gaps in supply chains of medical equipment effectively, an emergency approval process needs to be developed. While technology to close these gaps is available, the policy awaits optimization. The natural capability of rapid prototyping to incorporate quick change in functional designs over a series of iterations, along with the recently increased use of digitalized technology and 3D printing in biomedical devices and biocompatible material application, makes rapid prototyping the obvious short‐term response method of fabrication to meet supply chain needs. [ 7, 8 ] In the current pandemic, the need for early widespread testing of COVID‐19 has yielded a significant demand for nasopharyngeal swabs (NPS) and other testing materials. Additive manufacturing is increasingly used in biomedical applications in recent years. [ 7, 9, 10, 11 ] Specifically, in test swabs production, it is capable of meeting the market needs while significantly simplifying the supply chain. [ 12 ] This paper details 3D‐printed NPS as an alternative to existing mass produced equipment and compares our original designs to other open‐source, computer‐aided designed swabs developed by the online biomedical engineering community. 2 Comparative Design Review of NPS Many comparative studies have been conducted prior to the COVID‐19 pandemic to evaluate the performance of conventional specimen extraction methods in clinical studies. Efforts toward the development of new swab designs should base themselves first and foremost on the existing standards. Even so, imitation and innovation are not mutually exclusive processes. The different methods of specimen sampling, shown in Table 1, vary in terms of design parameters, including comfort, safety, accuracy, ease of self collection, and the ability to mimic these standard biomedical tools in rapid fabrication. The following discussion covers benefits and drawbacks of using nasal or nasopharyngeal wash, nasal aspirate (NA), oropharyngeal swab (OPS), nasal swab (NS), and NPS. The types of FDA‐approved NPS for SARS‐CoV‐2 sampling are shown in Figure 1. Table 1 Design constraints and objectives for the rapid fabrication of NPSs Parameters Conditions Constraints Dimensions Head diameter, D head = 3. 0 mm Neck diameter, D neck ≈ 1. 0 mm Handle diameter, D handle = 2. 5 mm Total swab Length, L swab ≥ 150. 0 mm Head length, L head ≥ 15. 0 mm Breakpoint from head, L score = 70. 0 mm Function Swab must be compatible with specimen preparation for RT‐PCR. Swabs should withstand storage under freezing temperatures. Swabs should be easily breakable at their score. Swabs should be autoclavable for complete sterilization. Swab should bend but not deform. Safety Swab features must not break off during use. Swabs should be sterilized before use. Residual uncured liquid resin from 3D‐printed swabs must be removed, through isopropyl alcohol or ethyl alcohol, and germicidal ultraviolet light, following exposure timescales that do not damage the material properties of the swabs. Designs should be optimized for mass production and limit random and systematic errors such that it does not affect the functionality and any potential harm to its user. Ecological concerns associated with fabrication, material, and disposal methods should be handled as appropriate. A sterile packaging solution, such as self‐sealing, autoclavable bags should complete the fabrication process. Objectives Tip surface area should be maximized. Surface features should be smooth and rounded, close to flocked swabs. The experience should be comfortable when inserting, swabbing around, and withdrawing the head. Easy to use for self‐collection. A minimum production yield of thousands of swabs per day. The swab should be easy and comfortable to break off into the transport vial. Length of the swab head submerged in vial transport solution is about 30 mm. © 2020 Wiley‐VCH GmbH Figure 1 NP swab designs. A) FDA‐approved NP swabs and their anatomy. (I) Puritan LA‐117. (II) Copan ESwab 481C. (III) Copan ESwab 482C. B) Copan FLOQSwabs 56380CS01 midturbinate swab anatomy. C) Scanning electron microscopy images of standard swab heads. (I) Spun fiber (scale bar = 200 μm). [ 13 ] (II) Nylon flock (scale bar = 200 μm). Reproduced with permission. [ 13 ] Copyright 2010, American Society for Microbiology. (III) Polyurethane foam (scale bar = 500 μm) Reproduced with permission. [ 14 ] Copyright 2009, John Wiley & Sons. D) Prototyped swab computer aided design (CAD) for 3D‐printed fabrication. (I–VIII) FAMES Lab designs. (IX–X) Abiogenix designs. (XI–XIII) Fanthom designs. (XIV) USF Health design. (XV–XIX) Wyss designs. (XX) Copan ESwab 481C swab shown as a reference. E) Heads of the swabs designs shown in panel (D) of this figure (Non‐FAMES Lab CAD imported from ref. [ 15 ]). Copyright © 2020 Wiley Periodicals, Inc 2. 1 Conventional Swabbing Tools This section offers a comparison of the different approved methods for the detection of the SARS‐Cov‐2 virus. Background information about each method is available in Supporting Information. Virus detection from NA samples was shown to be superior to NPS samples. [ 16 ] However, aspirate methods require additional suction equipment that does not make this tool cost effective, scalable, and easily implementable for widespread testing, unlike the range of different NS and NPS which are easy and safe and can be done anywhere without the need for additional devices. [ 17 ] One study showed that patients found the nasal wash (NW) more comfortable, in addition to being a more effective method than NPS in pathogen detection by culture, but with reverse transcription polymerase chain reaction (RT‐PCR), both methods were comparable. [ 18, 19, 20 ] On the contrary, another comparison between self‐administered foam NS versus staff‐collected NW showed that the former was comfortable, easy, and preferred by patients over the latter. [ 20 ] The adequate detection sensitivity when using NS for a wide range of respiratory viruses has been proven in previous clinical work, [ 17 ] as well as for NPS, [ 21 ] where it was shown that NPS sampling is more effective than OPS sampling, but NW is more sensitive than NPS sampling. Another study observed that although NPS was more reliable than OPS, it was not optimal and that the combination of the two had a better detection yield. [ 22, 23 ] However, data sensitivity must be balanced by the cost of additional swab fabrication, patient comfort, and compliance, hence why the additional use of OPS is unnecessary. A solution is to combine swab collection with saline spray, which showed to provide more sensitive results and was recommended as a better alternative to the high sensitivity of NW for conducting respiratory virus inspection. [ 20 ] Midturbinate swabs are designed to come into contact with a larger nasal surface area and for self‐collection, as shown in Figure 2. Compared with other swabs, midturbinate swabs have an additional collar at the handle to guide the maximum insertion depths of 55 mm, tolerated for adults. The study showed that self‐collected flocked midturbinate NSs were equivalent to staff‐collected NS or NPS and more efficient than rayon NPS. [ 24 ] Midturbinate swabs have shown to be less invasive and better suited for self‐collection due to the safety collar, as well as similarly comfortable compared with NSs and at least as sensitive as standard NPSs. [ 25, 26 ] The longer conical flocked head maximizes mucus contact, enabling better collection of specimens while offering a safe method of respiratory cell collection. [ 27 ] The major drawback of the collar in midturbinate swabs is in its rapid fabrication. Its structure takes more volumetric space, thus reducing the number of swabs that can be printed in a batch when using stereolithography (SLA). Under noncrisis conditions, this is not an issue, but under current conditions, fast‐paced production and item count are of the utmost priority. Figure 2 Swab 3D printing results. A) Printed selected swab designs. (I–VIII) FAMES Lab designs. (X) Abiogenix designs. (XIV) USF Health design. B) Close‐up picture of the printed head designs from panel (A). C) Pictures of the Dental SG Resin (Formlabs, Inc. ) surface under a microscope. (I) Neck of the swab. (II) Hexagonal geometry of the swab head. D) Spatial volume of each head design. E) Rigid shaft features and translucent color change through different fabrication stages of sterile swabs. (I) Swab freshly printed after an isopropyl alcohol wash to remove residual uncured resin (yellow). (II) Swab after thermal and UV curing to strengthen material properties (orange). (III) Swab after being autoclaved to sterilize the resin (white). Copyright © 2020 Wiley Periodicals, Inc Self‐collected swabs have been proven to be a reliable alternative to healthcare worker‐collected swabs, because self‐swabbing is feasible without any necessary prior training. [ 20, 24, 25 ] Specifically, self‐administered collection has proven to be useful in community‐oriented research. [ 28 ] This could be a key facilitator of widespread testing of quarantined individuals during a pandemic outbreak and allow healthcare workers to allocate time to urgent treatment rather than testing. Furthermore, patients seem to prefer self‐collection. [ 25 ] In a comparison of flocked nylon fiber swabs versus polyurethane foam swabs, as shown in Figure 1, it was found that the former had a larger interacting surface area, held more fluid volume, and could release the fluid more readily into the testing medium than the latter. However, the latter is better in performance when it comes to anterior nares swabbing. [ 29 ] Flocked nylon fiber swabs have been shown to collect more respiratory epithelial cells from the posterior nasopharynx. [ 30 ] Flocked swabs also have the advantage of preserving samples at its extremities, allowing for easy sample separation during specimen testing preparation. [ 25 ] Another study showed that different swab material and head or bristle designs impact sampling differently based on the collection site. This same study claimed however that the design of the swab seemed to matter less than the location of the virus in the respiratory tract. [ 23 ] This history of swab‐type comparisons provides benchmarks for specimen volume collection, respiratory virus detection, comfort of use, flexibility, fabrication and scalability, cost effectiveness, and widespread implementation, as shown in Table 1. A variety of materials, beyond the polymers discussed here, have been used for swab designs in the past, including wood, [ 31 ] which suggests that using nontraditional materials for swab fabrication, such as dental photopolymers, is not far‐fetched. Although dental photopolymers are not cost effective due to the chemical complexity of the polymer composite, their rapid production ability using affordable widespread precise equipment, in addition to their comparable functionality and efficiency, renders them an attractive possibility. 2. 2 Open‐Source Collaborative Designs Countless additive manufacturing companies across the world have joined the COVID‐19 response efforts through the fabrication of biomedical equipment such as swabs, ventilators, face shields, and more. Many universities have transitioned their courses online, and research has been limited to essential work, leading to the emergence of new collaborations between laboratories, manufacturers, and healthcare institutions, in part, yielding important considerations for swab design, fabrication methods, and validation protocols. These, combined with those that have resulted from our team's collaborators—namely, Indiana University (IU) Protolab, IU Health, Eskenazi Health, and Deaconess Health—and previous assessments of regular manufactured swabs discussed in the previous section, have been listed in terms of constraints and objectives in Table 1. 2. 3 Design Requirements Designs must adhere strictly to requirements and avoid mass production printing errors, as the nasopharyngeal space is very sensitive and prone to tissue damage, bleeding, and irritation. Institutions seeking to develop their own designs should use these parameters as a reference for their work and should seek local partners for production and testing. Methods for assembling swabs with more conventional polymers, such as polypropylene (PP), polytetrafluoroethylene (PTFE), polydimethylsiloxane (PDMS), and more, safe adhesives, and wicking substrates have been reviewed. [ 15 ] However, this fabrication process is hardly scalable for high throughput and requires significant labor if not automated, meaning that this solution should be left to contributors with existing industrial capabilities to that effect. The design fabrication should not depend on the labor of untrained students, researchers, and other members of the local community, even given the crisis conditions. While some have suggested that only immunocompromised patients need sterilized swabs, and that all other swabs can simply be disinfected, [ 15 ] this is too low of a standard, especially when dealing with such a novel and dangerous virus. Further, as swabs are produced in mass, making sterilization the base requirement, as opposed to making separate batches of sterilized and disinfected swabs, is more efficient and eliminates the risk of caregivers, giving immunocompromised patients the wrong swab. Safety and efficiency are not only important to those producing swabs, but also to the frontline healthcare workers dealing with an already overwhelming workload. 2. 4 Design Selection for Testing At the Fibers and Additive Manufacturing Laboratory (FAMES Lab), we developed a variety of designs that meet the constraints and objectives of Table 1, as shown in Figure 1. These were designed to optimize specimen collection. They are printed out of biocompatible dental resin by SLA and their assessment is covered in the next sections. We include here designs X and XIV from Abiogenix and the University of South Florida, respectively. These parallel design efforts have been picked due to their additional evaluation. The Abiogenix design (X) picked was one of three preferred by Stanford Health in their feedback. [ 15 ] The USF Health design (XIV) has gone through a series of validation processes and rapid clinical trials at Northwell Health and Tampa General Hospital, showing promising results in comparison with standard swabs. Figure 2 shows the accuracy level of the printing results of the selected designs and the FAMES Lab designs. FAMES Lab designs revolve around the inclusion of a hollow core to serve as a reservoir for carrying collected mucus when withdrawing a swab from the nasopharynx. In addition, FAMES Lab designs explore mechanisms and geometries such as those found in bubble wands, Archimedes screws, and negative Poisson ratios to optimize mucus collection and retention. We provide information here on swab prototype testing from open‐source collaborative work in an attempt to set a general protocol for swab prototyping and fabrication through additive manufacturing technology. 3 Rapid Prototyping and Fabrication Rapid prototyping creates functional systems or part representation before the release of a commercial product. This process usually involves additive manufacturing, an automated, simple method of producing an object with complex geometry with speed, precision, limited resources, and processing steps, often from a computer‐aided design. Rapid prototyping proves its efficacy also in times of crisis, such as the COVID‐19 outbreak, where biomedical equipment needs to be available to caregivers with haste, with little time to optimize the otherwise robust traditional manufacturing processes such as injection molding. Rapid prototyping equipment also has the advantage of being available in many fabrication spaces, from research laboratories to industrial centers to community‐operated collaborative workspaces, enabling the engagement of many toward the resolution of the present pandemic. SLA is used here as the additive manufacturing solution, as opposed to fused‐deposition modeling, selected laser sintering, laminated object manufacturing, or digital light processing due to the requirements for high accuracy and precision, excellent surface finish, biocompatibility, thermal resistance, flexibility, and functionality. 3. 1 Stereolithography SLA is one of the oldest methods of 3D printing, invented by Hideo Kodama in 1981 as an automated fabrication of 3D objects through layers of hardened polymers by ultraviolet exposure. [ 32 ] Rapid prototyping through SLA enables us to 3D print objects with arbitrarily complex geometries and with critical demands on precision and accuracy. SLA is a 3D vat polymerization printing process that involves the change of properties of light‐activated resin polymers when exposed to ultraviolet or visible light, namely photopolymerization. The polymer composition involves a variety of chemical components such as photoinitiators, absorbers, precursors, additives, and fillers. The hardened resin is a collection of low‐molecular‐weight monomers bonded by covalent bonds into a solid crosslinked unit. As polymerization is an ongoing reaction that is never fully complete, it often involves required postprocessing steps in the process. Photopolymers are slow to dissolve and instead swell and soften when absorbing a solvent. [ 33, 34, 35 ] Among light‐activated composite resin types, dental resins have a history of use in dentistry, making it well suited to invasive human specimen sampling. 3. 2 Autoclavable Dental Resins In dentistry, the use of biocompatible photopolymers as an oral restorative biomaterial dates back to 1960s. [ 36, 37 ] Dental resins are generally based off vinyl, polystyrene, or acrylic resins to be used as relining, die, impression, crown, and with other oral care applications. [ 38 ] Although biocompatibility resists precise definition in areas like tissue engineering, [ 39 ] in dentistry, an extended history of biocompatibility tests and protocols have been designed for local (mucosal and pulpal toxicity) and systemic (allergic, estrogenic, mutagenicity, and more) adverse reactions in vitro, in animals, in clinical studies, as well as occupational cyclic exposure. [ 40, 41, 42, 43, 44 ] One notable biocompatible acute toxicity assessment test is performed on fish embryos, where their stages of development are observed for chemical effects from exposure to the material over time. [ 45 ] Furthermore, the effects of the photo‐hardening transition from liquid to solid resin have also been studied for a variety of conditions, such as shrinkage, irradiation, or cure depth effects. [ 46, 47, 48, 49 ] Dental SG Resin (Formlabs, Inc. ) as a photoreactive resin is a light‐yellow translucent liquid and contains toxic elements. It has a viscosity between 800 and 1500 mPa s and an approximate composition of 90% methacrylic oligomers, [ 50 ] a standard monomer group used in dental resins, [ 41 ] and 3% phosphine oxides, [ 50 ] used for material color stability. [ 51 ] As the liquid resin holds toxic properties—causing possible irritation if ingested, inhaled, or through skin and eye contact—its curing process has to be followed strictly, especially for clinical applications. [ 52 ] Dental SG Resin has a type D durometer hardness value of 80 according to standard ISO 868:2003 and a Carpy impact strength of 12–14 kJ m −2 according to standard ISO 20795‐1:2013. In terms of flexibility, the material has a flexural strength greater or equal to 50 MPa and a flexural modulus greater or equal to 1500 MPa, according to standard ISO 20795‐1:2013. [ 53 ] However, material properties can vary depending on the print geometry, temperature, and orientation. To address this concern, three‐point bending tests done on swabs by Formlabs will be discussed in a later section of this paper. The autoclave was first invented by Denis Papin in 1679 and is now a standard for sterilization used in biological and medical research and industry, hospitals, mortuaries, and waste disposal. Its cycle of operation combines saturated steam loaded with latent heat, elevated moist heat up to 134 °C or dry heat up to 180 °C, and pressure and clean water to create a germicidal environment and induce protein denaturation, within its sealed chamber. This standard process must be used for cleaning surgical tools and any implantable or intrusive medical equipment, such as swabs. The cured dental resin must therefore be stable both against heat and humidity in the autoclave process. Dental SG Resin has been designed in its proprietary chemical composition for that purpose, unlike Dental LT Resin by Formlabs. However, it has been observed that once autoclaved, the brittleness of the material increases requiring a further look into the properties of autoclaved dental resins. 4 Preclinical Swab Design Testing In a preliminary step, certain observations of the mechanical performance of the selected swab designs were done, as shown in Figure 3. The roughness of the head designs was assessed to see what potential damage they could do to nasal tissue surfaces, by vigorously applying the design head onto a flat surface coated with wax. The head was first moved in an upward and downward direction and then rotated ten times. Although the head designs were not very rough overall, designs II, IV, and XIV showed the worst outcomes. The break‐off of the swab head into transport vials was assessed by snapping each of the four types of shafts present among the selected designs. Design X was the only one showing resistance which can lead to frustration for a caregiver swiftly packing up collected specimen into a vial. The different neck bending capabilities of the four shaft types were also tested both before and after sterilization. It was observed that sterilization seems to make the dental resin more brittle. A simple test, shown in Figure 3, suggests that the angle of curvature barely changes. Formlabs ran a comparative study on the tensile (ASTM D638) and torsional tests of swab Design XIV, Puritan swabs, and Copan swabs, which met the dimension requirement of Table 1, included in Supporting Information. Snapped swab heads holding a viscous xanthan gum and Milli Q water solution (4. 0 W V −1 ) were placed in 15 mL Falcon tubes in a centrifuge for 5 min at 3000 RPM and showed that collected samples were easy to extract from all head designs. Leaching of the artificial mucus was successful for every design, which is required for RT‐PCR sample preparation. This viscous solution will be further discussed in the next section as a good model to mimic mucus properties. In terms of fabrication, using Form 3B (Formlabs, Inc. ), a batch of 400 swabs can be printed with a reasonable large‐print software processing time. In a calculation of a mass print of swab design V, the 400 swabs, arranged in a square array, representing 3196 layers of 50 μm‐thick layers and a resin volume of 265. 23 mL, would require 69 h and 23 min to print, 2–3 h of post‐process, 45 min sterilization cycle, and cost about $113. 07 to produce, that is 28 cents per unit swab, taking into account material, equipment operation energy consumption, and the hourly labor of a makerspace technician. More information on the printing parameters are provided in Supporting Information. Figure 3 Preliminary mechanical performance observations. A) Scratching test of selected swab head designs I–VIII, X, and XIV on a layer of flat surface coated in wax at room temperature. B) Score break‐off of each type of shaft for the selected designs. C) Bending tests of each swab neck design groups before and after sterilization by autoclave. Copyright © 2020 Wiley Periodicals, Inc 4. 1 Artificial Mucus Xanthan gum is a commonly used biopolymer, acting as a thickening agent for applications such as food, pharmaceutical, and cosmetic formulations. [ 54 ] The xanthan gum solution prepared from Xanthomonas campestris (Sigma‐Aldrich, G1253) to mimic mucus with different viscosity properties will be hereinafter referred to as artificial mucus. Human mucus viscosity varies based on many factors, including the specific disease a patient is affected with, such as respiratory track diseases like rhinitis, chronic sinusitis, and chronic bronchitis. [ 55 ] A wide viscosity range was achieved by altering the percentage weight to volume of the xanthan gum powder, as shown in Figure 4, effectively covering a broad range of viscosity behaviors. Homogeneity of the artificial mucus highly depends on mixing method. Highly viscous artificial mucus which is above 1% W/V needs more mixing time until the gum is dissolved homogenously without any clumping. Xanthan gum powder should be mixed directly when exposed to Milli Q water. Xanthan gum powder may stick to the bottom of the beaker, though this can be prevented by a glass stirrer to collect all agglutinate powders from the bottom. Mixing with a magnetic stirrer caused air bubbles in the solution which were eliminated by the desiccator before the viscosity measurement of each concentration of the artificial mucus to obtain accurate viscosity data. Figure 4 First artificial mucus retention testing results. A) Viscosity measurements of different concentrations of xanthan gum and Milli Q water, designed as an artificial mucus for swab retention testing, covering the range of human mucus viscosity. B) Swab mass retention dipping test where swabs are weighed, dipped into a falcon tube containing artificial mucus, stirred three times, withdrawn, and weighed again. C) Dipping test results for the selected swab head designs. (I) Retention mass of the artificial mucus for designs I–VIII by the FAMES Lab, design X by Abiogenix, design XIV by USF Health, and design XXI which is the FLOQSwabs 503CS01 swab by Copan Diagnostics. (II) Normalized retention of the artificial mucus in terms of the volumetric space of the head designs. Copyright © 2020 Wiley Periodicals, Inc 4. 2 Mucus Retention Dipping Test Printed NPS were dipped 30 mm deep according to design requirements in 50 mL conical Falcon tubes. During each test, swabs were introduced without touching the inside of the Falcon tube, which was necessary to keep the contact area with the artificial mucus consistently sterile for each trial. Shaft design and the head of the volume did not affect the retention rate because the dipping motion was directed vertically, straight down within the confinement of the diameter of the 50 mL falcon tube, which was wide enough for the manipulation of the head volume of all NPS designs. This consistency is shown in general by the amount of small error bars (see Supporting Information) for each of the concentrations of the artificial mucus, resulting in a more coherent visual comparison of the different selected NPS designs. Five trials of the 11 dipping tests for the 10 different NPS swab designs were conducted across 7 different concentrations of artificial mucus, as shown in Figure 4. When overall results from all the artificial mucus concentrations were considered with the normalization by the total head volume of the NSP, the highest retention was found in design II and design VIII. These have a similar design that heavily relies on the surface tension of the collected liquid to hold itself in the hollow reservoir of the head. Design II has higher head volume than design VIII, enabling it to hold a greater volume of sputum. A dramatic increase in retention was observed with design VI and VII past a threshold off 2% W/V concentration of artificial mucus, as the latter's viscosity increases. The dimension of the honeycomb structure affected the amount of artificial mucus collected by the NPS head in design VI and VII, suggesting that the larger honeycomb gaps result in better sample collection. NPS design II was designed alike Archimedes screw involving a specific rotational direction which helps withdraw sputum as the head is rotated during sample collection. Due to the direction‐specific nature of this design, two dipping tests were administered clockwise and counterclockwise for design III. The results show that there is a slightly better retention rate when rotating the swab in the counterclockwise direction as expected. We have compared the performance of 3D‐printed designs to that of standard swabs (FLOQSwabs 503CS01, Copan Diagnostics). The result of those trials is shown in Figure 4, labeled as design XXI, where the flocked swab performed fine for low viscosities but surprisingly bad for higher viscosities, in terms of both average retention and its random variation, reflected by the larger error bars (see Supporting Information). 4. 3 Nasal Model for Swab Behavior Analysis The nasal model consists of a 3D‐printed support structure, a 10 mL narrow mouth Erlenmeyer flask, and a plastic tube with an inner diameter of 6. 35 mm, which was used to mimic the confined nose canal of the nasal cavity, as shown in Figure 5. The model simulated the nasal cavity, wherein the NPS has to be inserted to reach the nasopharynx, and was based on computed tomography measurements done on an asymptomatic adult‐sized nasal air space. [ 56 ] NPS designs with different shafts and different head volumes and the viscous behaviors of the artificial mucus both affected retention rates significantly. Due to the slight bend in the nasal pathway featured in the model, NSP with a rigid shaft design (designs V–VIII, X, and XIV) had a tendency to break while turning inside the 10 mL narrow‐mouth Erlenmeyer flask. The flexible shaft performed much better with the nasal model. Large NPS heads (designs I–V) were not very efficient in the confined space of the tube. During the swab withdrawal step, a significant amount of the collected artificial mucus, especially the mucus with low viscous behavior, was lost. This loss invited instability in retention measurements for each trial with artificial mucus with same viscosity. Figure 5 Second artificial mucus retention test results. A) Nasal model for swab retention test in a confined space. (I) Illustration of the anatomy that the nasal model is designed to represent. (II) 3D‐printed model where swabs are weighed, pushed through a tube, rotated for collection, withdrawn, and weighed again. B) Normalized shaft performance for the same design head within the confined space of the nasal model. C) Nasal model test results of artificial mucus retention for designs I–VIII by the FAMES Lab, design X by Abiogenix, design XIV by USF Health, and design XXI which is the FLOQSwabs 503CS01 swab by Copan Diagnostics. (I) Retention mass of the artificial mucus. (II) Normalized retention of the artificial mucus in terms of the volumetric space withdrawn through a tube for the length of the head. Copyright © 2020 Wiley Periodicals, Inc Taking these results into account and normalizing them in terms of volumetric flow when withdrawing the head out of the tube, the highest retention was found in design III (with both clockwise and counterclockwise directional usage), design IV, and design II. The volumetric flow equation is based on Stokes drag flow, shown in Equation ( 1 ), for which its derivation is provided in Supporting Information, where V out is the total volume withdrawn from the nasal model, D Tube is the tube diameter, D Head is the maximum head diameter, L Head is the length of the head, V out is the total volume withdrawn from the nasal model, and V Head is the volume of the swab head (see Figure 3 ). (1) V out = π ⋅ ( D Tube 2 − D Head 2 ) ⋅ L Head 8 + V Head Equation ( 1 ): Volumetric flow of the withdrawn head with collected specimen, to which swab performances in the nasal model are normalized. It is interesting to point out that in this test the clockwise direction of swab III performed better despite expectations. Design IV was created with a negative Poisson ratio, such that it would radially contract upon application of compressive stress along the head axis to maximize the patient's comfort during insertion and would radially expand upon application of tensile stress along the head axis to maximize sample retention during extraction. Indeed, the geometric shape of the head provided better retention among most of the other NSP designs. To address the lack of rigid shaft performance in the bent confined model, a comparison between the performance of a flexible handle and a rigid handle for the same head design II and VIII was conducted, as shown in Figure 5. The results show that the difficulty in maneuverability of the shaft barely affects the retention capabilities of the head designs. The trials ran for the flocked swab, design XXI, shown in Figure 5, had a similar outcome as the dipping test. The flexibility of the industrial swab made the collection as easy as it was for designs I–IV, which have a flexible shaft. However, its retention of artificial mucus with high viscosity was suboptimal, unlike its efficiency with lower viscosity behaviors which were adequate. All performances for both the dipping and nasal model tests are shown in Table 2. Table 2 Conclusion of key performance assessment points for each swab designs Design Dipping retention test Nasal model retention test Retention mass Normalized retention Strength (artificial mucus [g 100 mL −1 ]) Retention mass Normalized retention Strength (artificial mucus [g 100 mL −1 ]) I Average Mediocre 2. 0–4. 0 Average to good Average to good 0. 0–0. 5 II Best Second best 1. 5–4. 0 Good to Best Good to best 1. 0–2. 0 III Second best Average 1. 0–4. 0 Average to good Good 1. 0–1. 5 IV Average to good Mediocre to average 2. 0–4. 0 Second best Second best 1. 5–2. 0 V Second worst Average 2. 0 Mediocre Mediocre 0. 5–1. 0, 2. 0 VI Mediocre Average 2. 0 Worst Worst 1. 0–1. 5 VII Mediocre to average Average to good 1. 0–4. 0 Mediocre Mediocre 1. 0–1. 5 VIII Average Best 1. 5–4. 0 Average Mediocre 1. 0–1. 5 X Worst Average 1. 5–4. 0 Mediocre Mediocre 1. 0–4. 0 XIV Average Worst Average Average 1. 0–4. 0 XXI Mediocre to good Average to mediocre 0. 0–1. 5 Average to good Average 0. 0–1. 5 © 2020 Wiley‐VCH GmbH 5 Conclusion Some preliminary results have suggested that conventional swabs outperform new prototypes. [ 57 ] Our study highlights exactly which design factors may be responsible for these differences and thus offers insights into precisely how prototypes should be designed so as to improve upon the conventional swab design. In comparing a variety of original designs, we have demonstrated the efficiency of various surface porosity, geometries, and volumetric space optimizations. We have provided direction regarding the appropriate design for the shaft ability to snap easily swabs with collected specimen from their sacrificial handle, surface roughness comfort levels, and head and shaft dependence in confined airspaces. In our study we have provided a protocol for preclinical trials to observe mucus retention through an anatomical model and mimic the tenderness of human tissue in comfort level trials. In addition, we created a xanthan gum‐based artificial mucus that covers the range of human mucus viscosity and can be used for future applications as a semisynthetic mucus model. Rapid prototyping and fabrication by SLA is an attractive option to consider for NP swab production. Especially given the current crisis, rapid prototyping offers the capability for quick, precise design and manufacture, which will in turn enable widespread, reliable testing for COVID‐19. Widespread, reliable testing is necessary for the prompt diagnosis and treatment of the disease and ultimate management of the pandemic. Though not cost effective, dental photopolymers are an appealing material for a possible NP swab design given that they enable fast, precise fabrication and are conducive to the design constraints discussed in this paper. Importantly, the history of swab design and a comparison of existing designs provide a basis for future innovation. 6 Experimental Sections 6. 1 6. 1. 1 3D Printing A CAD model was printed by SLA in Dental SG Resin or Castable Wax, using a Form 3B printer (Formlabs, Inc. ). Parts were then washed for 15 min in an isopropanol bath to remove leftover uncured resin. The parts were left to dry for 1 h. Unless the component was printed in castable wax, the parts were then cured under both ultraviolet and blue light of wavelength 400–500 nm (Dulux L BL 18 W/71 and Dulux L BL UVA 18 W/78 lamps) and heated at 60 °C. Biomedical equipment that required sterilization were then placed in an autoclave (Steris Amsco 250LS) for a full sterilization cycle of 45 min, reaching a maximum temperature of 123. 6 °C. Artificial Mucus Model A total of 0. 1, 0. 3, 0. 5, 1, 1. 5, 2, and 4 g of xanthan gum from Xanthomonas campestris, purchased from Sigma‐Aldrich, were mixed with 100 ml Milli Q water to reach the final concentrations of 0. 1% W/V, 0. 3% W/V, 0. 5% W/V, 1% W/V, 1. 5% W/V, 2% W/V, and 4% W/V xanthan gum/ Milli Q solutions. Solutions were mixed in a 250 mL glass beaker using a magnetic stirrer at 300 RPM at room temperature until the xanthan gum powder was mixed homogenously without any clumping. A glass stirring rod was used to mix the agglutinated xanthan gum on the bottom of glass beaker. Beakers were closed with parafilm to eliminate solution evaporation. All solutions were kept resting at room temperature for at least 12 h to complete the hydration process. Air bubbles in all solutions were eliminated with a vacuum supplied by a desiccator. Solutions were collected at 50 ml falcon tubes. Viscosity Measurements Viscosity of the xanthan gum/Milli Q solutions were measured with a Black Pearl Rotational Rheometer, ATS RheoSystems. Measurements were carried out with a 25 mm concentric cylinder system within a range of 0. 1–300 RPM at 25 °C. A 15 mL sample volume was used for each measurement. Dipping Retention Test Swabs were dipped in 50 mL falcon tubes containing the artificial mucus. Test swabs were turned three times in the counterclockwise direction inside the falcon tubes. Design IV was specifically tested both clockwise and counterclockwise due to its design. The weight of each test swab was measured before and after being soaked with the artificial mucus using a Sartorius Secura Analytical Balance. The retention rate was measured with five trials for each concentration of the artificial mucus for each design. Nasal Model Retention Test The 3D design of the nasal structure was printed using Form 3B SLA printer (Formlabs, Inc. ). A clear polyurethane tube (inner diameter = 6. 35 mm, outer diameter = 9. 52 mm) was used to mimic the nasal canal, through the nasal cavity, to the nasopharynx. Around 10 mL of each artificial mucus solution with different viscosities was poured into a 10 mL Erlenmeyer flask at the end of the tube, to mimic the sputum collection area in the nasopharynx. The housing structure for the nasal model was designed by CAD and 3D printed in Castable Wax by SLA in Form 3B (Formlabs, Inc. ). The narrow mouth of the Erlenmeyer flask and the tube were sealed using parafilm to eliminate possible leaks. Test swabs were put into the plastic tube through the 10 mL Erlenmeyer flask and turned three times inside each artificial mucus concentration. The weight of each test swab was measured before and after being dipped in the artificial mucus using a Sartorius Secura Analytical Balance. The retention rate was measured with five trials for each concentration of artificial mucus for each design. Flocked Swab Washing For the flocked swab (FLOQSwabs 503CS01, Copan Diagnostics) trials, between each consecutive testing, the swab was washed and dried. The flocked swab was cleaned with three wash and dry cycles, followed by three dipping cleaning in 5 mL warm water in 15 mL Falcon tubes, where the swab head was soaked for 10 min. The wet head was then dried with hot air for fast evaporation at 100 °C. No deformation or melting of the flocked head was observed during the drying phase. This process was repeated as necessary, typically no more than three times, until the cotton swab was clean and weighed its general weight with a precision of ±0. 005 g. Conflict of Interest The authors declare no conflict of interest. Supporting information Supplementary Material Click here for additional data file.
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10. 1002/adem. 202300301
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Advanced engineering materials
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Incorporating the Antioxidant Fullerenol into Calcium Phosphate Bone Cements Increases Cellular Osteogenesis without Compromising Physical Cement Characteristics
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Herein, fullerenol (Ful), a highly water-soluble derivative of C 60 fullerene with demonstrated antioxidant activity, is incorporated into calcium phosphate cements (CPCs) to enhance their osteogenic ability. CPCs with added carboxymethyl cellulose/gelatin (CMC/Gel) are doped with biocompatible Ful particles at concentrations of 0. 02, 0. 04, and 0. 1 wt v% −1 and evaluated for Ful-mediated mechanical performance, antioxidant activity, and in vitro cellular osteogenesis. CMC/gel cements with the highest Ful concentration decrease setting times due to increased hydrogen bonding from Ful’s hydroxyl groups. In vitro studies of reactive oxygen species (ROS) scavenging with CMC/gel cements demonstrate potent antioxidant activity with Ful incorporation and cement scavenging capacity is highest for 0. 02 and 0. 04 wt v% −1 Ful. In vitro cytotoxicity studies reveal that 0. 02 and 0. 04 wt v% −1 Ful cements also protect cellular viability. Finally, increase of alkaline phosphatase (ALP) activity and expression of runt-related transcription factor 2 (Runx2) in MC3T3-E1 pre-osteoblast cells treated with low-dose Ful cements demonstrate Ful-mediated osteogenic differentiation. These results strongly indicate that the osteogenic abilities of Ful-loaded cements are correlated with their antioxidant activity levels. Overall, this study demonstrates exciting potential of Fullerenol as an antioxidant and proosteogenic additive for improving the performance of calcium phosphate cements in bone reconstruction procedures.
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1. Introduction Calcium phosphate cements (CPCs) have been extensively used as clinical bone substitutes for over 20 years. CPCs have a similar structure to natural bone and can easily be applied to osseous defects due to their moldability and low-temperature setting reaction. [ 1 - 4 ] They can produce apatite in physiological conditions and do not harm nearby tissues via heat generation as seen with highly exothermic polymethylmethacrylate (PMMA) cements. [ 5, 6 ] CPCs are commercially available in a pure form or as composites when combined with various naturally derived polymers such as carboxymethyl cellulose (CMC) and gelatin (gel). When combined with CPCs, CMC improves overall material handling by increasing viscosity [ 7, 8 ] and enhancing resistance to CPC wash-out. [ 8, 9 ] Gel was also reported to enhance the mechanical properties of CPCs in many studies. [ 10 - 12 ] Bigi et al. [ 10 ] showed that 15 wt% gel increased the compressive strength of apatite CPCs to a value five times greater than unmodified cements. In addition to the contribution of gel on CPC mechanical properties, gel also improves cement biocompatibility. [ 10, 13 ] Despite the improvement in CPC performance imparted by incorporating these polymers, further enhancements are still needed to heighten CPC osteogenic capabilities [ 8, 14 ] for deployment in clinical bone augmentation procedures. [ 15 ] Many materials have been suggested as cement additives to enhance the osteogenic ability of CPCs. The most heavily investigated candidates are therapeutic growth factor proteins including bone morphogenetic protein-2 (BMP-2) [ 16 ] and transforming growth factor-β (TGF-β). [ 17 ] However, these recombinant proteins are expensive and difficult to use across an innately heterogenous patient population that often requires variable dosing of these compounds. [ 18, 19 ] Moreover, disadvantages of BMP-2 use in the clinic such as ectopic bone formation and excessive inflammation [ 19, 20 ] have been reported. Bioactive ions such as Sr +2 [ 21 ] and Mg +2 [ 22 ] have also been investigated as proosteogenic additives to CPCs with some success. Additionally, recent studies revealed that CPCs loaded with magnetic nanoparticles and primed with electromagnetic fields enhanced osteogenic differentiation of human dental pulp stem cells and rat adipose-derived stem cells. [ 14, 23 ] Despite the many strategies investigated for improving the osteogenic potential of CPCs, demonstrated efficacy of these new methods in clinical bone reconstruction procedures has remained elusive. One underexplored strategy for increasing the osteogenic ability of CPCs is the incorporation of antioxidants to reduce local levels of reactive oxygen species (ROS) in the local tissue. Previous studies have shown that elevated levels of ROS cause oxidative stress and can force mesenchymal stem cells, osteoblasts, osteocytes, and MC3T3-E1 preosteoblast cells into apoptosis. [ 24 - 26 ] Moreover, other reports indicate that high ROS concentrations inhibit osteogenic differentiation of mesenchymal stem cells and MC3T3-E1 cells. [ 27 - 29 ] Trolox and selenium, antioxidants that have been extensively studied for bone tissue engineering applications, both enhance cellular osteogenic differentiation and reduce oxidative stress in cells seeded on bone scaffolds. [ 29, 30 ] These antioxidants were also effective ROS scavengers when incorporated into brushite CPCs, [ 31, 32 ] though their effects on osteogenic differentiation were not determined with these cements. Therefore, the potential for antioxidants to increase the osteogenic capability of CPCs is still unknown. Fullerenol (Ful) is a powerful ROS scavenger and a cutting-edge carbon nanomaterial due to its specific physical and chemical features. The electron-deficient positions on the surface of Ful neutralize ROS molecules by efficiently transferring the radicals’ unpaired electrons within the fullerene cage. [ 33, 34 ] Furthermore, Podolsky et al. [ 35 ] suggested that hydroxyl groups on Ful (molecular composition of C 60 (OH) 22–24 ) also participate in ROS scavenging, making Ful an even more potent antioxidant than unmodified fullerene. Ful also possesses low cytotoxicity, [ 36, 37 ] with previous work showing that Ful particles increased the viability of bone marrow macrophage cells [ 38 ] and enhanced the proliferation of adipose-derived stem cells. [ 37 ] Other studies have reported Ful’s ability to increase osteogenic differentiation of stem cells via upregulation of key osteogenic markers such as alkaline phosphatase (ALP), runt-related transcription factor 2 (Runx2), and osteocalcin (OCN). [ 36, 39 ] Hence, Ful is a strong candidate to enhance osteogenic properties of CPCs. Therefore, the aim of this study is to enhance the osteogenic ability of CMC/gel-incorporated apatite CPCs by incorporating Ful particles into these materials. For investigating the effect of Ful on the physical characteristics of CPCs, setting time, apatite formation, apatite morphology, compressive strength, and modulus were all investigated. Moreover, the ROS scavenging abilities of Ful-incorporated CPCs were evaluated. Finally, the osteogenic differentiation of pre-osteoblasts treated with Ful-doped CPCs was determined by immunofluorescence (IF) measurements of actin network formation, cellular production of ALP, and polymerase chain reaction (PCR) measurements of osteogenesis-associated gene expression. In short, this work is the first study to incorporate the antioxidant Ful into calcium phosphate cements and demonstrates the exciting potential of this novel nanomaterial as a pro-osteogenic additive. 2. Results and Discussion 2. 1. Characterization of Ball-Milled Powder Before preparing the cements, the size and crystallinity of cementing powder consisted of tetracalcium phosphate (TTCP) and dicalcium phosphate dihydrate (DCPD) were characterized via scanning electron microscopy (SEM) and X-ray diffraction (XRD). The size and crystallinity of TTCP/DCPD particles before and after ball milling are presented in Figure S1, Supporting Information. Figure S1A, B, Supporting Information, shows that smaller particle sizes in a narrower range were obtained via ball milling. In Figure S1C, D, Supporting Information, the shapes of TTCP and DCPD can be seen; TTCP particles have round edges and smooth surfaces due to their sintering procedure, while DCPD particles are smaller, irregularly shaped, and adhered to the larger TTCP. [ 40 ] According to ImageJ analysis ( Figure S1E, F, Supporting Information ), the average sizes of TTCP and DCPD particles were reduced to 5. 1 μm and 162 nm from 18. 8 μm and 464 nm via ball milling, respectively. In Figure S1G, Supporting Information, XRD spectra taken before and after ball milling indicate that only DCPD and TTCP phases were present after ball milling. 2. 2. Physical Characterization of Ful-Incorporated Cements To prepare Ful-incorporated cements, first, 1 wt v% −1 CMC and 1. 5 wt v% −1 Gel were dissolved in hardening agent solution and Ful was incorporated into CMC/gel solution at different concentrations. To determine the molecular interactions of Ful with CMC and gel, Fourier-Transform infrared spectroscopy (FTIR) was conducted. Figure 1 shows FTIR spectra for the liquid phases of control, CMC/gel, Ful0. 02, Ful0. 04, and Ful0. 1 cement formulations. The proposed structure after Ful conjugation to CMC/gel cements is presented in Figure 1A. In Figure 1B, FTIR spectrum of the control cement’s liquid phase shows the characteristic peaks at 989. 06, 1076. 22, 1636. 30, and 3315 cm −1. The peaks at 989. 06 and 1076. 22 cm −1 correspond to symmetric stretching vibration and asymmetric stretching vibration of the P─O bond, respectively. [ 41, 42 ] The peak at 1636. 30 cm −1 is assigned to the bending vibration of the H─O─H bond, and the peak at 3315. 00 cm −1 corresponds to the stretching vibration of the OH bond. [ 41, 43 ] Additions of CMC/gel and/or Ful to the cement liquid phase caused shifting in the OH band as shown in Figure 1C. With the CMC/gel addition, the OH band at 3315. 00 cm −1 shifted to a lower frequency of 3284. 46 cm −1 due to stretching vibrations of OH groups in the CMC and stretching vibrations of NH groups in gel. [ 44 ] Moreover, 0. 02 and 0. 04 wt v% −1 Ful incorporation to cement liquid phases leads to a further respective downshift to 3273. 65 and 3267. 49 cm −1 due to an increase of OH content and hydrogen bonding in these samples. [ 45, 46 ] Previous studies also investigated the interaction of Ful particles with amine groups of amino acids and demonstrated that Ful could form hydrogen bonds with these molecules via hydroxyl – amine associations. [ 47 - 49 ] Indeed, Dong et al. [ 48 ] observed a shift to a lower frequency in the NH band from an FTIR spectra of phenylalanine/Ful due to hydrogen bonding between amine group of phenylalanine and Ful hydroxyl units. Moreover, hydrogen bonding of antioxidants with carboxyl and hydroxyl groups of polysaccharides was reported by Lombo-Vidal et al. [ 50 ] In short, the demonstrated molecular associations between Ful particles and CMC/Gel cement components are well supported from previous studies. For using moldable materials in orthopedic applications, an initial setting time of 8 min and a final setting time of 15 min are recommended. [ 8, 51 ] Figure 1D demonstrates the results of setting time measurements. Though baseline CPCs showed initial/final setting times of 11. 4/14 min, CMC/Gel addition significantly increased the initial and final setting times to 20. 8 and 27. 9 min due to immobilization of TTCP and DCPD particles from the viscosity increase. [ 8, 52 ] The addition of 0. 02 and 0. 04 wt v% −1 Ful did not change initial and final setting times of CMC/gel, though 0. 1 wt v% −1 Ful significantly decreased the CMC/gel initial setting time to 16. 2 min and final setting time to 18. 2 min, closer to clinically recommended time frames. These measurements reveal that high concentrations of Ful accelerated the hardening of CMC/Gel cements. This increased cement reactivity is potentially mediated by an increase of the common ion effect via hydrogen bonding of Ful particles to HPO 4 −2 [ 53 ] since the interaction of hydroxyl groups with phosphates has been previously suggested. [ 47 ] Therefore, these data indicate that HPO 4 −2 sites on Ful0. 1 facilitated binding of Ca +2 ions to promote apatite nucleation in these Ful-doped cement formulations. 2. 3. Cement Phase Analysis and pH Measurements in PBS Though Ful incorporation into CPC cements decreased their setting time, it did not significantly impact the conversion rate of TTCP and DCPD particles into calcium-deficient hydroxyapatite (CDHA) domains when incubating the cements in PBS, as shown in Figure 2. Figure 2A - C shows XRD spectra at hours 1, 3, and 24 of incubation which reveal the cements’ crystal phases under physiological conditions. It can be deduced that DCPD consumption was completed and apatite started to form in less than 1 h in all cements based on the appearance of characteristic apatite peaks 25. 88, 32. 05°, and 32. 50°. [ 54 ] However, there were still residual amounts of TTCP (major peaks at 29. 23° and 29. 81° as shown in Figure S1G, Supporting Information ) at 1 h that were consumed within 3 h so that cements wholly comprised apatite, as shown in Figure 2B. Indeed, the height of the apatite peaks increased between 1 and 24 h to further confirm this conversion. This conversion rate for apatite formation from TTCP and DCPD was one of the fastest reported in the literature as presented in Table S1, Supporting Information, [ 54 - 56 ] and this rapid apatite formation can likely be attributed to the utilization of smaller-sized TTCP and DCPD particles and usage of sodium phosphate solution instead of distilled water. [ 57 ] Finally, no differences were observed for the time needed for conversion of TTCP and DCPD to CDHA with Ful incorporation into cements ( Figure 2A - C ). Figure 2D shows the pH change of physiological solutions in which the cements were incubated, further confirming the TTCP/DCPD conversion kinetics from the XRD analyses in Figure 2A - C. The pH of all solutions increased in the first 3 h but decreased drastically afterward due to the absence of basic TTCP, consumption of phosphate from the physiological solution, and resulting apatite growth. [ 57, 58 ] Control cements reached their minimum pH at 168 h (day 7) while CMC/gel, Ful0. 02, Ful0. 04, and Ful0. 1 cements reached their minimum pH at 96 h (day 4). These behaviors are possibly due to greater amount of ─COO − ions from the CMC/gel which can act as nucleation sites for Ca +2 ions. [ 11, 43 ] The minimum pH level for all cement solutions was between 6. 8 and 7. 0, and after they reached their minimum, pH increased and kept constant between 7. 0 and 7. 4. Finally, SEM imaging indicated that all cement formulations possessed nanosized platelet-like crystals [ 59, 60 ] with no obvious differences between groups being observed ( Figure S2, Supporting Information ). 2. 4. Cement Phase and Morphological Analysis and pH Measurements in Simulated Body Fluid (SBF) Phase analyses of cements were also carried out after incubating cement samples in simulated body fluid (SBF) and comparing against as-prepared cements. The XRD spectrum of as-prepared cements ( Figure 3A ) shows that the crystalline phase of these materials features high levels of apatite even without an SBF incubation. Figure 3B, C shows similar results obtained after the incubation in SBF for 3 or 24 h, demonstrating increasing apatite formation and disappearance of characteristic TTCP peaks. CDHA formation was also determined via energy-dispersive X-ray (EDX) analysis of cements incubated in SBF for 24 h as shown in Figure S3, Supporting Information. The incubated cements achieved Ca/P ratios of 1. 40–1. 50, which is the Ca/P ratio of CDHA ( Figure S3, Supporting Information ). [ 61 ] Moreover, the pH change of SBF-incubated cements over time ( Figure 3D ) demonstrated a similar trend with the pH profile in PBS ( Figure 2D ). Finally, SEM images of cements before incubation in SBF ( Figure 3E ) and after SBF treatment ( Figure 3F ) show that all groups included nanosized platelet-like crystals [ 59, 60 ] indicative of apatite formation ( Figure 3E, F ). Importantly, as-prepared cements ( Figure 3E ) did not display obvious differences between PBS-incubated ( Figure S2, Supporting Information ) or SBF-incubated ( Figure 3F ) materials 2. 5. Cement Compression Testing Cancellous bone has compressive strength values between 2 and 16 MPa and compressive moduli in the range of 120–1100 MPa, [ 62, 63 ] and synthesizing a CPC with proximate mechanical properties is critical for providing mechanical stability with cancellous bone implants. [ 64, 65 ] As shown in Figure 4, CMC/gel, Ful0. 02, Ful0. 04, and Ful0. 1 cements (model samples shown in Figure 4A ) were mechanically characterized through compression testing (representative stress/strain traces shown in Figure 4B ). Control cements achieved compressive strength and elastic modulus values of 2. 04 and 145. 55 MPa, respectively, as shown in Figure 4C, D. In alignment with previous reports, [ 9, 10, 12, 13, 66 ] solely adding CMC to cements did not alter compressive strength or modulus compared to unmodified control cement samples; however, CMC/gel significantly increased the compressive strength and modulus compared to control samples. Finally, Ful addition did not significantly affect the compressive strength and modulus values of CMC/Gel samples, and all formulations eclipsed the benchmark values of 2 MPa for strength and 150 MPa for modulus. These collective data indicate that Ful-doped cements are mechanically suitable for treatment of nonload-bearing bone reconstructions. 2. 6. Measurement of In Vitro ROS Scavenging Ability To assess the antioxidant capacity of Ful-loaded CPCs, a DPPH radical inhibition assay was performed on cement samples. Figure 5A demonstrates the color change of DPPH solution after 24 h treatment with Ful-incorporated cements. DPPH’s color change from purple color to yellow is a marker of reduction of DPPH radicals [ 67 ] by Ful-incorporated cements, while no change was observed in DPPH solutions treated with control and CMC/gel cements. Quantitative analysis of DPPH activity reduction by Ful-incorporated cements is plotted in Figure 5B, with Ful0. 02 and Ful0. 04 samples demonstrating 46. 58% and 51. 41% DPPH inhibition within 24 h. However, DPPH was only inhibited by 29. 28% for the Ful0. 1 samples potentially due to Ful aggregation, as previously described by Roy et al. [ 68 ] for these particles at high concentrations. [ 68, 69 ] Regarding the significant role of hydroxyl groups both in aggregation and in radical scavenging, [ 35 ] it is suggested that Ful particle aggregation diminishes the presentation of active sites on the surface of Ful particles and therefore reduces their interaction with radicals. [ 68, 69 ] 2. 7. In Vitro Cytotoxicity and Cement Release Kinetics of Ful Particles In vitro cytotoxicity evaluations of Ful-loaded cement extracts collected from 1, 3, and 5-day incubations were obtained in L-929 cells, as shown in Figure 6A. Though derived from fibroblasts, L-929 cells were chosen for initial cytotoxicity screening of these cements to align with ISO 10 993-5 standard protocols. The day 3 extracts showed that the cell viability increased with medium-dose Ful particles. Moreover, images of extract-treated L-929 with live–dead staining ( Figure 6B ) agree with the quantitative viability results in Figure 6A as Ful0. 02-and Ful0. 04-treated cells show more coverage than the other treatment groups. To be able to understand the reason behind the trend of percent cell viability, both Ful particle and calcium ion release studies were conducted. In vitro release kinetics of Ful particles from Ful-loaded cements are shown in Figure 6C, demonstrating that Ful nearly completely released from the cements within 24 h, though since Ful particles tend to accumulate on the surface of these cements; burst discharge from CPC materials is common. [ 31, 32 ] Similarly quick release kinetics of antioxidant molecules from CPCs were also reported for trolox [ 31 ] and selenium. [ 32 ] As expected, the concentration of cumulative Ful released from the cements was proportional to the initial Ful concentration in each sample. Since the amount of Ful released was the same across the studied time points in Figure 6C, the increase in cell viability on day 3 could be due to the protective effects of Ful against toxicity from excess CPC-released calcium ions, which can be harmful to cells. [ 70 ] Excess quantities of calcium ions can compromise cellular viability, [ 71 ] and in vitro release kinetics of Ca ions from cement formulations demonstrated that less than 12 mg L −1 of free calcium was released from control or Ful0. 04 cements over 5 days ( Figure 6E ) but did increase between the first and fifth days of incubation. Since CPC residues are known to promote in vivo tissue inflammation and increase in vitro ROS levels, [ 72 - 74 ] we concluded that the higher viability levels for cells treated with Ful0. 02 and Ful0. 04 extracts were likely due to Ful’s antioxidant capacity as similarly demonstrated in other reports. [ 37, 75, 76 ] The cell protective effect of Ful is correlated with its ROS scavenging ability, and as suggested by Hao et al. , [ 37 ] Ful increases the expression of MAPK-related proteins p38 and ERK to suppress ROS-induced toxicity. Ful0. 1 cements demonstrated lower levels of ROS scavenging than Ful0. 02 and Ful0. 04 cements. Finally, all cement fifth-day extracts were still biocompatible; however, the viability percentage of cells incubated in Ful0. 1 fifth-day extracts was significantly lower than cells incubated with CMC/gel fifth-day extracts. This result closely correlates with lower ROS scavenging ability of Ful0. 1 than Ful0. 02 and Ful0. 04. Hence, it can be assumed that the high antioxidant capacity of Ful0. 02 and Ful0. 04 formulations facilitated the increased survival of L-929 cells. 2. 8. In Vitro Cellular Osteogenic Differentiation Mediated by Ful-Incorporated Cements To determine Ful’s osteogenic potential when delivered to cells from CMC/Gel cements, IF staining, ALP production measurements, and gene expression studies were conducted with MC3T3-E1 pre-osteoblasts in vitro. The effects of Ful-loaded cements on cellular osteogenic differentiation are shown in Figure 7. First, MC3T3-E1 cells were incubated with varying concentrations of Ful particles to understand Ful’s impact on cell viability ( Figure 7A ). Crucially, the total dose of Ful released from cements (below 0. 025 mg mL −1 in Figure 6C ) was lower than 0. 500 mg mL −1 since Ful’s concentration appears to (nonsignificantly) begin reducing cellular viability, as shown in Figure 7A. However, no Ful concentration up to 1 mg mL −1 was found to significantly reduce viability of MC3T3-E1 cells over 24 h as in agreement with previous studies listed in Table S2, Supporting Information. [ 37, 38, 77 - 81 ] When the concentration of Ful reached 0. 25 mg mL −1, it was discovered that it increased the viability of MC3T3-E1 cells over 24 h, as shown in Figure 7A. Literature also emphasizes that Ful is a highly biocompatible material, [ 37, 38, 77 - 81 ] and previous assessments with Ful concentrations up to 10 mg mL −1 were shown to be non-cytotoxic using human skin fibroblasts, [ 77, 78 ] murine macrophage cells, [ 80 ] and human epidermal keratinocytes. [ 79 ] MC3T3-E1 cells induced to differentiate in osteogenic media with respective cement extracts for 7 days were assessed for actin fiber formation ( Figure 7B ). As shown in these microscopic images, the actin fibers of cells treated with Ful0. 04 extracts were denser and more visible than the ones found on other cement groups. The quantitative analysis of actin fluorescence intensity ( Figure 7C ) also shows that Ful0. 04 significantly increased the actin density of MC3T3-E1 cells. F-actin organization/polymerization is a known hallmark of osteogenic differentiation, [ 82, 83 ] indicating that Ful0. 04 extracts had the strongest osteogenic impact on MC3T3-E1 cells compared with the other cement extract treatments. In contrast, the actin density of cells treated with Ful0. 1 media was significantly lower than the actin density of Ful0. 04-treated cells, indicating that Ful particles are likely aggregating at this higher concentration. [ 68, 69 ] As shown in Figure 7D, quantification of cellular ALP production following cement extract treatment largely mirrored the results of IF staining. Ful0. 04 extract media significantly increased ALP activity in MC3T3-E1 cells compared to the cells treated with control, CMC/gel, and Ful0. 1 cement extracts. This significant ALP increase with Ful0. 04 treatment was seen at both 7 and 14 days, and ALP production increased over time for all cement extract treatment groups. Moreover, gene expression of the osteogenesis marker Runx2 also increased between 7 and 14 days for all cement treatment groups but tripled with low-dose Ful treatment, as shown in Figure 7E. However, cells treated with the high-dose Ful0. 1 cement extracts did not significantly increase Runx2 expression compared to cells treated with CMC/gel extracts ( Figure 7E ). Finally, a nonsignificant increase of COL1 expression was detected for low-dose Ful cement and it significantly decreased for high-dose Ful cement ( Figure 7F ). Previous studies of metabolic changes during the osteogenic differentiation process have conclusively shown that ROS levels are suppressed with antioxidant enzymes during osteogenesis. [ 84 ] Despite a somewhat unclear mechanism, numerous in vitro and in vivo studies also suggest that antioxidants can accelerate osteogenic differentiation by reducing ROS concentrations. [ 29, 36, 85, 86 ] In addition to the previously suggested antioxidants, the specific effect of Ful on osteogenic differentiation has also been studied. [ 36, 39 ] Liu et al. [ 39 ] showed that Ful particles upregulated Runx2 and OCN levels, while Yang et al. [ 36 ] reported that ALP and OCN levels, along with cellular mineralization, were all enhanced following treatment with Ful particles. Moreover, it was suggested that Ful enhanced the expression of FoxO1, which is linked to Runx2, and is responsible for protecting against ROS in bone tissue. [ 36 ] Guided by these previous studies, it can reasonably be deduced that the antioxidant capacity of ROS-scavenging particles correlates with their osteogenic ability. This coupling of antioxidant activity and cellular osteogenesis is further supported by the data in this study ( Figure 5 and 7 ), and future analyses with these materials will determine the exact mechanisms relating Ful’s antioxidant activity to bone formation since ROS interacts with several key osteogenic pathways, including Wnt, FOXO, Hedgehog, and MAPK/ERK. [ 27, 36, 39 ] Further in vivo studies with Ful-loaded CPCs in conventional rodent models such as calvarial defects [ 21, 22 ] or femoral segmental defects [ 87 ] are also required to demonstrate the efficacy of Ful in bone healing. However, this study provides strong initial evidence of the benefits Ful can impart to CMC/Gel cements for application in bone reconstruction procedures. 3. Conclusion In this work, the antioxidant molecule Fullerenol was efficiently integrated into a calcium phosphate cementing agent and did not alter apatite formation time, apatite morphology, or compressive mechanical properties. Low Ful concentrations did not impact cement hardening rates, while high Ful concentrations did significantly decrease cement setting times. Crucially, Ful-loaded cements displayed potent antioxidant activity as measured by DPPH radical scavenging assays and correspondingly increased the osteogenic differentiation of MC3T3-E1 pre-osteoblast cells as determined from actin network staining, ALP production measurements, and PCR gene expression analyses. These findings also show that at high concentrations, Ful displayed reduced antioxidant activity (likely due to Ful aggregation) and similarly promoted lower levels of cellular osteogenesis. Overall, these collective data indicate that Ful delivered from bone cements promotes earlier and more potent cellular osteogenic differentiation than unmodified CPCs without compromising physical cement characteristics. 4. Experimental Section Ball Milling Process and Characterization of Powder: Tetracalcium phosphate (TTCP, Hitemco Medical, USA) and dicalcium phosphate dihydrate (DCPD, Sigma-Aldrich, United States) were utilized to synthesize calcium-deficient hydroxyapatite (CDHA) cements. A mixture of TTCP and DCPD in equivalent masses was ground and sieved through a 75 μm sieve. TTCP/DCPD particles were ball milled (BM_S38003, MSE Teknoloji) with 3 mm yttrium-stabilized zirconia (MSE Teknoloji) for 48 h in absolute ethanol (Merck, Germany) at 120 rpm. [ 88, 89 ] The weight ratio of powder/ethanol/ball was 1/2/1. 30. Absolute ethanol was selected to leave the powder mixture undissolved, and a milling speed of 120 rpm was chosen from calculating 75% of critical speed. [ 90 ] After ball milling, the slurry was dried at 50 °C and grounded into powder. SEM images of particles were taken with a Philips-FEI XL30 in secondary electron mode, and the average particle size was determined via ImageJ. Finally, XRD spectra were obtained using a Rigaku D/MAX-Ultima+/PC equipped with CuKα radiation and step angle of 0. 02. The analyzed powder was kept in a vacuum desiccator until its use. Preparation of Cements: The composition of cements is presented in Table 1. First, the hardening agent solution was prepared. Briefly, a 5. 68 wt v% −1 Na 2 HPO 4 (Merck, Germany) solution was produced in distilled water and adjusted to pH of 7. 4 via dropwise titration with 1 m HCl (Merck, Germany). Control cements were obtained by mixing a hardening agent solution with ball-milled TTCP/DCPD powder. For CMC/gel solution, 1. 5 wt v% −1 gel from bovine skin (Sigma-Aldrich, United States) was dissolved in hardening agent solution at 60 °C for 15 min before adding 1 wt v% −1 CMC (Sigma-Aldrich, United States). After dissolution of CMC at 90 °C, the solution was cooled at room temperature. CMC/gel cements were prepared via mixing CMC/gel solution with TTCP/DCPD powder as previously described. [ 12, 91 ] Ful0. 02, Ful0. 04, and Ful0. 1 solutions were obtained by dissolving 0. 02, 0. 04, or 0. 1 wt v% −1 C 60 (OH) n · mH 2 O Ful ( n > 40, m > 8, Sigma-Aldrich, United States) in CMC/gel solutions. Ful handling and incorporation into the solutions were all done in an enclosed glovebox or fume hood. Gum-like consistency of the powder/liquid mix was obtained using a powder/liquid weight ratio of 1. 25 in all cements. Structural Analysis: Functional groups in the respective cements’ liquid phases were analyzed via FTIR (Nicolet FTIR Instruments, Thermo Fisher Scientific). Spectrum from 4000 cm −1 to 550 cm −1 was recorded with 32 scans. [ 92 ] Setting Time Measurements: Setting times of cement samples were measured by the Gilmore test according to standard C266-20 from the American Society for Testing and Materials. [ 93 ] Respective specimens were fabricated at 5 mm height and 10 mm diameter, immersed in phosphate buffered saline (PBS, Sigma-Aldrich, United States) and kept at 37 °C and 100% relative humidity. The initial and final setting times were determined using a Gilmore apparatus (Utest) by recording the time at which the light needle ( m = 113. 4 ± 0. 5 g, d = 2. 12 ± 0. 05 mm) and heavy needle ( m = 453. 6 ± 0. 5 g, d = 1. 06 ± 0. 05 mm) could not make an indentation in the respective sample. Phase Analysis and pH Measurements in PBS: For phase analysis, cements were immersed in PBS at 37 °C for 1, 3, and 24 h. At these time points, the cements were taken from the incubator, removed from PBS, and then immediately frozen at −80 °C. After samples were frozen, they were lyophilized over 24 h before analyzing with XRD. XRD spectra of specimens at 2 θ = 20°–50° were obtained using a Rigaku D/MAX-Ultima+/PC equipped with CuKα radiation. Step angle was 0. 02°. [ 94 ] For pH measurements, cement specimens were incubated in PBS at 37 °C and pH of physiological solution were measured via a pH meter (MP225, Mettler Toledo) at specific time points between 1 h and 28 days. [ 94 ] During each incubation, PBS was refreshed every 2 days. Phase Analysis and pH Measurements in SBF: SBF was prepared according to the protocol described in Kokubo et al. [ 95 ] and cements were incubated in SBF for 3 and 24 h. At these time points, cements were taken from incubator, removed from SBF, and taken to −80 °C freezer with as-prepared cements. Subsequently, they were lyophilized for 24 h. XRD spectra of the cements were obtained via Bruker D8 Advance XRD. For pH measurements, cements were incubated in SBF and measured for pH changes over time using a pH meter (MP225, Mettler Toledo) at specific time points between 1 h and 14 days. SBF was refreshed every 2 days. Imaging of Surface Morphology and Apatite Formation: For imaging of surface morphology and apatite formation, cement samples were kept in either PBS or SBF at 37 °C for 24 h. [ 96 ] After 24 h, the aqueous media were removed and specimens were lyophilized. Lyophilized samples were sputter coated with platinum, and cross-sectional images of fractured surfaces were taken with a Philips-FEI XL30 SEM in secondary electron mode in 100x–100, 000x magnification. Cross-sectional images of as-prepared and incubated cements were collected with a Thermo Scientific Quatro S Environmental SEM in secondary-electron mode via Everhart–Thornley detector (ETD) detector to increase image magnification to 160 000× without the need for sample sputter coating. For elemental analysis, EDX was conducted on the same device using a gaseous secondary electron detector (GSED) detector to determine cements’ elemental Ca/P ratio. Compression Test: For the mechanical analysis of the cements, samples were molded with 10 mm diameter and 20 mm length. Before testing, the cements were incubated at 37 °C for 24 h in PBS before confirming height and diameter values for the respective samples. Compression tests were performed on wet samples using a Zwick roell z100 testing system with a 100 kN load cell and a crosshead speed of 0. 5 mm min −1. [ 97 ] Compressive strength and elastic modulus values were calculated from the processed stress/strain data. In Vitro ROS Scavenging Ability Test: In Vitro: ROS scavenging potential of Ful-containing cements and CMC/Gel were determined using a 1, 1-Diphenyl-2-picrylhydrazine (DPPH) assay. [ 98, 99 ] DPPH solutions was prepared by dissolving DPPH (31. 62 mg, Sigma-Aldrich, United States) in 80:20 v/v% solution of ethanol (400 mL, Thermo Fisher Scientific, United States) and deionized water. Respective cement samples (10 mg) were placed into DPPH solution (2 mL) and shook at 37 °C in the dark. At 1, 12, and 24 h, samples of DPPH (100 μL) were removed and absorbance was measured on a Tecan MPlex microplate reader at 517 nm. Absorbance values of DPPH solutions incubated with Ful containing cements ( A Ful ) were compared to absorbance readings of DPPH solution incubated with CMC/gel cement as shown in Equation (1) below. (1) % DPPH Amount = 1 − ( A Ful − A CMC ∕ Gel A CMC ∕ Gel ) × 100 In Vitro Ful Release from CPCs: For determining the amount of Ful released from cements, Ful particles were labeled with fluorescein 5-isothiocyanate (FITC, Thermo Fisher Scientific, United States) via an isothiocyanate-hydroxyl conjugation. First, Ful (1 mg) was dissolved in anhydrous dimethyl sulfoxide (2 mL, Sigma, Germany). Then, FITC (3 mg) was added into solution and stirred for 16 h at room temperature in the dark. To extract the FITC-labeled Ful, acetone (13 mL, Thermo Fisher Scientific, United States) was added to the reaction solution to precipitate out Ful-FITC while leaving unreacted FITC in the supernatant. The precipitated reaction mix was then spun down at 3000×g for 5 min before removing the supernatant and air drying the remaining Ful-FITC. As described above, cements were then prepared with the Ful-FITC along with control and CMC/gel cements. Cement samples ( n = 3 samples per group, 30 mg) were placed into individual wells in 24-well tissue culture plates containing PBS (1 mL). Then, well plates were put on a shaker and kept in the dark. On days 1, 3, and 5, the fluorescence emissions of releaseate media were measured at FITC’s excitation wavelength at 495 nm and emission wavelength at 520 nm via microplate reader. [ 100 ] Fluorescence of pure PBS was subtracted from the releaseate fluorescence readings. Finally, Ful-FITC release kinetics were determined using a standard curve prepared from FITC solutions in serial dilutions of PBS. In Vitro Cell Culture Studies: Murine L-929 fibroblasts (American Type Culture Collection (ATCC), United States) and murine MC3T3-E1 pre-osteoblasts (Subclone 4, ATCC, United States) were the two immortalized cell lines used for in vitro biological evaluations. L-929s were cultured in Dulbecco’s modified Eagle’s Medium low glucose (Gibco, United States) supplemented with 10% fetal bovine serum (Gibco, United States) and 1% penicillin-streptomycin (Gibco, United States). MC3T3-E1 cells were cultured in either growth medium (alpha-MEM without ascorbic acid (Gibco, United States), 10% fetal bovine serum, and 1% penicillin-streptomycin) or osteogenic induction medium (growth medium supplemented with 10 −8 m dexamethasone (Thermo Fisher Scientific, United States), 0. 01 mol L −1 L-ascorbic acid (Thermo Fisher Scientific, United States), and 50 mg mL −1 β-glycerophosphate (Thermo Fisher Scientific, United States)). All cells were maintained in a 95% humidified incubator with 5% CO 2 at 37 °C. In Vitro Cytotoxicity Evaluation: A cytotoxicity test was conducted according to ISO 10 993-5. [ 101 ] Cement samples (30 mg) were immersed in sterile PBS (1 mL, Sigma-Aldrich, United States), and the extracts were aspirated on days 1, 3, and 5. Then, the extracts were sterile filtered through 0. 22 μm polyvinylidene fluoride (PVDF) hydrophilic membranes and combined with growth medium in a 1:10 ratio. When cultured L-929 cells reached confluency, they were seeded in a 96-well plate at a concentration of 5 × 10 3 cells/well. [ 102 ] After overnight incubation, old media were removed, and cells were treated with PBS (100 μL) with or without the respective cement extracts for 24 h. Metabolic activity was determined via the MTS assay (Promega, United States). The absorbance was measured at 490 nm using a microplate reader (Varioskan Lux, Thermo Fisher Scientific). The results were normalized to the absorbances of the nontreated group. Live–Dead Staining: Live–dead staining was performed on the cells treated with day three cement extracts. First, the staining solution was prepared by adding calcein (0. 5 μL, Thermo Fisher Scientific, United States) and ethidium bromide (2 μL, Thermo Fisher Scientific, United States) to 1 mL D-PBS (Gibco, United States). Then, the media on the cells was removed from the wells, cells were rinsed with D-PBS, and staining solution was added to each cell well before incubating at 37°C for 45 min. Finally, the images were obtained via fluorescence microscopy using FITC and Texas Red imaging channels (Axiovert A1, Zeiss). [ 103 ] Cell Viability Test with Ful Particles: MC3T3-E1 cells were cultured to confluency before seeding at a density of 5000 cells/well in a 96-well plate for 24 h. Subsequently, the cells were treated with Ful 100 μL in cell culture media in serial dilutions including 1000, 500, 250, 125, 62. 5, and 0 μg mL −1 for 24 h. [ 37 ] Finally, the solutions in each well were aspirated, and the CellTiterGlo (Promega, United States) solution was added. After incubation for 10 min at room temperature, the bioluminescence of the wells was measured at 515 nm with a microplate reader (Tecan MPlex). Luminescence values for Ful-treated wells were normalized to the nontreated group. In Vitro Osteogenic Differentiation Studies: For osteogenic differentiation studies, cement formulations were incubated in MC3T3-E1 growth medium for 1 week before extracts were collected. [ 21 ] To prepare the osteoblastic induction medium for each cement, these extracts were filtered through sterile 0. 22 μm PVDF filters and combined with the osteoblastic induction medium in a 1:1 ratio. Cells were cultured in growth medium for 24 h. On the second day, half the medium in each well was removed before adding 1 mL of respective cement extract media samples. The osteoblastic induction medium was changed every 2 or 3 days. Immunofluorescence Staining: Cytoskeletal differences in treated MC3T3-E1 cells were observed on day 7 post-treatment. The culture medium was removed, cells were washed with PBS, and the cells were fixed with 3. 7% formaldehyde (Riedel-De-Haën, Germany) for 30 min at room temperature. Fixed cells were permeabilized for 5 min with 0. 1% Triton-X 100 (Biobasic, Canada) before being incubated in blocking solution for 10 min. Cell cytoskeletal actin was visualized by treating the cells with Alexa Fluor 594 phalloidin (Thermo Fisher Scientific, United States) for 60 min. Subsequently, the samples were rinsed with PBS to remove unbound conjugates. Moreover, cell nuclei were stained with 4′, 6-diamidino-2-phenylindole (DAPI, Sigma, Germany) for 15 min. Finally, the samples were rinsed thoroughly with PBS before adding Prolong Diamond Antifade Mountant (Thermo Fisher Scientific, United States) and imaged under the fluorescence microscope (VertA1, Zeiss, Germany) using the TRITC and DAPI channels. Finally, the fluorescence intensity of cells was measured via ImageJ. Alkaline Phosphatase (ALP) Activity Test: For the ALP activity test, MC3T3-E1 cells were seeded at a density of 10 000 cells cm −2 in 6-well plates (2 mL media per well) and cultured in osteoblastic induction medium for either 7 or 14 days. At these time points, medium from the cells was first aspirated before briefly rinsing with D-PBS and trypsinizing. The trypsinized cells were added to individual microcentrifuge tubes (1. 5 mL) and lysed in cold D-PBS via vortexing. Lysates were processed to measure ALP activity according to the manufacturer’s instructions (Abcam, United States). Briefly, the lysates were resuspended in assay buffer (500 μL) and centrifuged at 4 °C for 15 min at 16 000 rpm. Precipitations were removed, and supernatants were transferred to a new tube and kept cold during the following steps. To measure ALP activity, 80 μL supernatant samples were put into wells in 96-well tissue culture plates in serial dilutions before adding p-nitrophenyl phosphate solution (50 μL) to each well. [ 104 ] After incubation for 60 min in the dark, NaOH solution (20 μL) was added, and absorbance at 405 nm was measured via a microplate reader (Tecan MPlex). The absorbance of the background was also considered and subtracted from the absorbance of the sample wells. Finally, the p-nitrophenol (pNP) concentration of each well was calculated using the prepared standard curve, and ALP activity was calculated as shown below. (2) ALP activity ( U ∕ m L ) = ( A ∕ V ) ∕ T where A is the amount of pNP generated in samples calculated from the standard curve (μmol), V is the volume of sample added in assay well (mL), T is the reaction time (minutes), and units are glycine units. A Bradford assay was also utilized to calculate protein amounts for normalizing ALP activity. Briefly, Coomassie reagent (250 μL, Bioworld, United States) was added to supernatant (5 μL) and incubated for 10 min at room temperature. Finally, absorbance was measured at 595 nm using a microplate reader (Tecan MPlex), and protein amount was calculated according to the prepared standard curve calculated with solutions of bovine serum albumin protein. PCR: MC3T3-E1 cells were seeded in a 12-well plate at 20 000 cells per well and incubated for 24 h to allow cells to adhere. After incubation, cells were treated with 1 mL of a 1:1 volumetric mix of osteoblastic induction medium and filtered extracts from respective cement formations ( n = 3 wells per treatment). Cells were allowed to differentiate for 7 and 14 days, with media being replaced with fresh osteogenic media on days 4, 7, and 10. On days 7 and 14, cells were washed and lysed before extracting and purifying RNA using a PureLink RNA Mini Kit (Invitrogen, Waltham, MA), following the manufacturer’s protocol. Extracted RNA was quantified for quality and concentration using a NanoQuant plate on a Tecan MPlex microplate reader. cDNA was synthesized using an iScript Reverse Transcription Supermix for RT-qPCR (Bio-Rad, Hercules, CA). Quantitative real-time PCR (qRT-PCR) was performed using iTaq Universal SYBR Green Supermix (Bio-Rad, Hercules, CA). Relative expression levels of runt-related transcription factor 2 (RUNX2) and Collagen Type 1 (COL1) were normalized to glyceraldehyde 3-phosphate dehydrogenase (GAPDH) using the ΔΔCt method. Primer sequences are given in Table 2. Statistical Analysis: Unless otherwise stated, all experiments employed at least n = 3 samples per group. Results were reported as mean±standard error (SEM) and p < 0. 05 was considered statistically significant. Statistical analyses were performed using IBM Statistics 25. Normality of data were confirmed via Shapiro-Wilk testing. For ALP activity on 14 th day and setting time analysis, the significance of difference was determined via Kruskal–Wallis with Dunn’s test for pairwise comparisons. For the other analyses, one-way ANOVA followed by Tukey posthoc testing was utilized. GraphPad Prism 9 was utilized to plot the results and single asterisk (*), double asterisk (**), triple asterisk (***), and quadruple asterisk (****) were used to represent p < 0. 05, p < 0. 01, p < 0. 001, and p < 0. 0001 respectively. Supplementary Material supplement
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10. 1002/adfm. 201202685
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Advanced functional materials
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Charge-Tunable Silk-Tropoelastin Protein Alloys That Control Neuron Cell Responses
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Tunable protein composites are important for constructing extracellular matrix mimics of human tissues with control of biochemical, structural, and mechanical properties. Molecular interaction mechanisms between silk fibroin protein and recombinant human tropoelastin, based on charge, are utilized to generate a new group of multifunctional protein alloys (mixtures of silk and tropoelastin) with different net charges. These new biomaterials are then utilized as a biomaterial platform to control neuron cell response. With a +38 net charge in water, tropoelastin molecules provide extraordinary elasticity and selective interactions with cell surface integrins. In contrast, negatively charged silk fibroin protein (net charge −36) provides remarkable toughness and stiffness with morphologic stability in material formats via autoclaving-induced beta-sheet crystal physical crosslinks. The combination of these properties in alloy format extends the versatility of both structural proteins, providing a new biomaterial platform. The alloys with weak positive charges (silk/tropoelastin mass ratio 75/25, net charge around +16) significantly improved the formation of neuronal networks and maintained cell viability of rat cortical neurons after 10 days in vitro. The data point to these protein alloys as an alternative to commonly used poly-L-lysine (PLL) coatings or other charged synthetic polymers, particularly with regard to the versatility of material formats (e. g. , gels, sponges, films, fibers). The results also provide a practical example of physically designed protein materials with control of net charge to direct biological outcomes, in this case for neuronal tissue engineering.
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No full text available
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10. 1002/adfm. 201301275
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Advanced functional materials
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Iron Oxide-labeled Collagen Scaffolds for Non-invasive MR Imaging in Tissue Engineering
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Non-invasive imaging holds significant potential for implementation in tissue engineering. It can e. g. be used to monitor the localization and function of tissue-engineered implants, as well as their resorption and remodelling. Thus far, however, the vast majority of efforts in this area of research have focused on the use of ultrasmall super-paramagnetic iron oxide (USPIO) nanoparticle-labeled cells, colonizing the scaffolds, to indirectly image the implant material. Reasoning that directly labeling scaffold materials might be more beneficial (enabling imaging also in case of non-cellularized implants), more informative (enabling the non-invasive visualization and quantification of scaffold degradation) and more easy to translate into the clinic (since cell-free materials are less complex from a regulatory point-of-view), we here prepared three different types of USPIO nanoparticles, and incorporated them both passively and actively (via chemical conjugation; during collagen crosslinking) into collagen-based scaffold materials. We furthermore optimized the amount of USPIO incorporated into the scaffolds, correlated the amount of entrapped USPIO with MR signal intensity, showed that the labeled scaffolds are highly biocompatible, demonstrated that scaffold degradation can be visualized using MRI and provided initial proof-of-principle for the in vivo visualization of the scaffolds. Consequently, USPIO-labeled scaffold materials seem to be highly suitable for image-guided tissue engineering applications.
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No full text available
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10. 1002/adfm. 201302859
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Advanced functional materials
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Multilayered Inorganic Microparticles for Tunable Dual Growth Factor Delivery
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There is an increasing need to control the type, quantity, and timing of growth factors released during tissue healing. Sophisticated delivery systems offering the ability to deliver multiple growth factors with independently tunable kinetics are highly desirable. Here, a multilayered, mineral coated micro-particle (MCMs) platform that can serve as an adaptable dual growth factor delivery system is developed. Bone morphogenetic protein-2 (BMP-2) and vascular endothelial growth factor (VEGF) are bound to the mineral coatings with high binding efficiencies of up to 80%. BMP-2 is firstly bound onto a 1 st mineral coating layer; then VEGF is bound onto a 2 nd mineral coating layer. The release of BMP-2 is sustained over a period of 50 days while the release of VEGF is a typical two-phase release with rapid release in the first 14 days and more sustained release for the following 36 days. Notably, the release behaviors of both growth factors can be independently tailored by changing the intrinsic properties of the mineral coatings. Furthermore, the release of BMP-2 can be tuned by changing the thickness of the 2 nd layer. This injectable microparticle based delivery platform with tunable growth factor release has immense potential for applications in tissue engineering and regenerative medicine.
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No full text available
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10. 1002/adfm. 201302901
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Advanced functional materials
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Arrayed Hollow Channels in Silk-based Scaffolds Provide Functional Outcomes for Engineering Critically-sized Tissue Constructs
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In the field of regenerative medicine there is a need for scaffolds that support large, critically-sized tissue formation. Major limitations in reaching this goal are the delivery of oxygen and nutrients throughout the bulk of the engineered tissue as well as host tissue integration and vascularization upon implantation. To address these limitations we previously reported the development of a porous scaffold platform made from biodegradable silk protein that contains an array of vascular-like structures that extend through the bulk of the scaffold. Here we report that the hollow channels play a pivotal role in enhancing cell infiltration, delivering oxygen and nutrients to the scaffold bulk, and promoting in vivo host tissue integration and vascularization. The unique features of this protein biomaterial system, including the vascular structures and tunable material properties, render this scaffold a robust and versatile tool for implementation in a variety of tissue engineering, regenerative medicine and disease modeling applications.
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10. 1002/adfm. 201303400
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Advanced functional materials
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Deconstructing the Effects of Matrix Elasticity and Geometry in Mesenchymal Stem Cell Lineage Commitment
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A wide variety of environmental factors including physical and biochemical signals are responsible for stem cell behavior and function. In particular, matrix elasticity and cell shape have been shown to determine stem cell function, yet little is known about the interplay between how these physical cues control cell differentiation. For the first time, by using ultraviolet (UV) lithography to pattern poly(ethylene) glycol (PEG) hydrogels we are able to manufacture microenvironments capable of parsing the effects of matrix elasticity, cell shape, and cell size in order to explore the relationship between matrix elasticity and cell shape in mesenchymal stem cell (MSC) lineage commitment. Our data shows that cells cultured on 1, 000 μm 2 circles, squares, and rectangles were primarily adipogenic lineage regardless of matrix elasticity, while cells cultured on 2, 500 and 5, 000 μm 2 shapes more heavily depended on shape and elasticity for lineage specification. We further went on to characterize how modifying the cell cytoskeleton through pharmacological inhibitors can modify cell behavior. By showing MSC lineage commitment relationships due to physical signals, this study highlights the importance of cell shape and matrix elasticity in further understanding stem cell behavior for future tissue engineering strategies.
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10. 1002/adfm. 201303460
| 2,014
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Advanced functional materials
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Highly Aligned Nanofibrous Scaffold Derived from Decellularized Human Fibroblasts
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Native tissues are endowed with a highly organized nanofibrous extracellular matrix (ECM) that directs cellular distribution and function. The objective of this study is to create a purely natural, uniform, and highly aligned nanofibrous ECM scaffold for potential tissue engineering applications. Synthetic nanogratings (130 nm in depth) were used to direct the growth of human dermal fibroblasts for up to 8 weeks, resulting in a uniform 70 μm–thick fibroblast cell sheet with highly aligned cells and ECM nanofibers. A natural ECM scaffold with uniformly aligned nanofibers of 78 ± 9 nm in diameter was generated after removing the cellular components from the detached fibroblast sheet. The elastic modulus of the scaffold was well maintained after the decellularization process because of the preservation of elastin fibers. Reseeding human mesenchymal stem cells (hMSCs) showed the excellent capacity of the scaffold in directing and supporting cell alignment and proliferation along the underlying fibers. The scaffold’s biocompatibility was further examined by an in vitro inflammation assay with seeded macrophages. The aligned ECM scaffold induced a significantly lower immune response compared to its unaligned counterpart, as detected by the pro-inflammatory cytokines secreted from macrophages. The aligned nanofibrous ECM scaffold holds great potential in engineering organized tissues.
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10. 1002/adfm. 201303547
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Advanced Functional Materials
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Synthesis and Characterization of Gelatin-Based Magnetic Hydrogels
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A simple preparation of thermoreversible gelatin-based ferrogels in water provides a constant structure defined by the crosslinking degree for gelatin contents between 6 and 18 wt%. The possibility of varying magnetite nanoparticle concentration between 20 and 70 wt% is also reported. Simulation studies hint at the suitability of collagen to bind iron and hydroxide ions, suggesting that collagen acts as a nucleation seed to iron hydroxide aggregation, and thus the intergrowth of collagen and magnetite nanoparticles already at the precursor stage. The detailed structure of the individual ferrogel components is characterized by small-angle neutron scattering (SANS) using contrast matching. The magnetite structure characterization is supplemented by small-angle X-ray scattering and microscopy only visualizing magnetite. SANS shows an unchanged gelatin structure of average mesh size larger than the nanoparticles with respect to gel concentration while the magnetite nanoparticles size of around 10 nm seems to be limited by the gel mesh size. Swelling measurements underline that magnetite acts as additional crosslinker and therefore varying the magnetic and mechanical properties of the ferrogels. Overall, the simple and variable synthesis protocol, the cheap and easy accessibility of the components as well as the biocompatibility of the gelatin-based materials suggest them for a number of applications including actuators.
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1 Introduction Living organisms are able to produce highly sophisticated materials. 1 Biominerals are organic-inorganic hybrid materials abundant in Nature. 1 They are formed under highly controlled conditions, 2 show complex morphologies and are very often hierarchically structured. 3 On each hierarchical level, the optimum structure is realized 3 and consequently, much improved physical properties are obtained. An amazing biomineral is the chiton tooth. These animals scratch algae from rocks, which requires that their teeth are very hard and wear resistant. In fact, chiton teeth show the highest hardness and stiffness among the known biominerals being for example three times harder than human tooth enamel. 4 The teeth themselves are hardened by the inclusion of magnetite nanoparticles (15–20 nm) into a protein-polysaccharide gel matrix. The high nanoparticle content makes the hybrid material very hard and wear resistant even enabling to scratch corals and stones. Chiton teeth are not the only biomineral which is formed in a gel matrix as a template. Nacre, another biomineral known for its exceptional fracture toughness is also synthesized inside a gel, in this case a silk hydrogel. 5 Such syntheses could be successfully mimicked resulting in inclusion of the hydrogel inside a single crystal, 6 which caused an increased mechanical stability, just as found in biominerals. 7 Results like this imply that mineralization of inorganic components inside an aqueous hydrogel is an advantageous strategy towards advanced hybrid materials synthesis. Even if the gels are not as highly mineralized as in case of Biominerals, they can exhibit interesting properties since the viscoelasticity of the hydrogel and the swelling behaviour can be combined. In addition, many hydrogels are biocompatible which is an issue in medical applications. 8 – 10 Variation of the mineral content can change the counterplay between the two components in the hydrogel and thus the properties of the hybrid material. It is therefore not astonishing that a large number of studies on hydrogels filled with inorganic nanoparticles were reported including hydrogels filled with metal, 11 – 13 semiconductor 14, 15 and magnetic nanoparticles. 16 – 20 The properties of these gels can be adjusted over a large range by simple variation of the organic-inorganic content allowing their use in applications as diverse as catalysis, 21 – 23 switchable electronics, 24, 25 tissue engineering, 26, 27 drug delivery, 28 – 31 wastewater treatment, 32 – 34 hyperthermia cancer therapy 35, 36 and soft actuators. 37, 38 However, the synthesis of biocompatible magnetic hydrogels with high and adjustable magnetite content was not yet reported. These are, moreover, especially interesting because they can be moulded into any shape, can be used for medical applications and can be addressed by external magnetic fields making them interesting for applications as actuators, switches etc. Here we report the synthesis of biocompatible and thermoreversible gelatin gels which are mineralized with magnetite. A large variation of the organic/inorganic ratio is possible, magnetite loads of up to 70 wt% can be realized, which influences the materials structure, as shown by small-angle neutron scattering (SANS) using contrast variation to reveal the details of the organic, as well as of the inorganic structure. 2 Results and Discussion 2. 1 Gelatin Hydrogel Gelatin, which is derived from partial hydrolysis of native collagen, can be considered as a polydisperse copolymer with a broad molar mass distribution. The gelatin Type B used in our experiments has a molar mass of approximately 25–50 kDa. Native collagen forms a triple helical structure which is stabilized by interchain hydrogen bonds. 39 At temperatures above the gelation temperature (T gel ) gelatin forms a homogeneous solution in water, which transforms below T gel and above the overlap concentration of ∼0. 5 wt% (in H 2 O) to a thermoreversible physical gel. 40 During the cooling process the gelatin undergoes conformational changes, the so called coil to helix transition. Most of the gelatin chains form a three-dimensional interconnected network of chains reverted back from a random coil to a triple-helical structure. 39 The structure of a gelatin hydrogel at room temperature in D 2 O was determined by SANS. The scattering pattern plotted in Figure 1 were measured at classical SANS and USANS diffractometers delivering scattering at very small Q of the order 10 −3 up to 3 nm −1. USANS instruments have to use optical devices such as refraction lenses or mirrors. 41 There are two distinct Q-regimes which are well described by the solid line representing the best fit of the data using the two levels Beaucage expression. 42 The Beaucage expression is given according to 1 representing a combination of Guinier's and Porod's laws describing the scattering at low and large Q, respectively. More quantitatively both approximations are valid for the parameter u = R g Q smaller or larger than 1, u representing the product of radius of gyration R g and scattering vector Q (defined below). Guinier's law has the shape of a Gaussian function whereas for Q larger than 1/R g (u>1) a power law according to is often observed, which in case of α = 4 represents the famous Porod law of compact particles with a sharp surface. 43 Figure 1 SANS macroscopic cross-section dΣ/dΩ versus scattering vector Q for 18 wt% gelatin in D 2 O (T = 20 °C). At low Q (< 0. 02 nm −1 ) USANS data are also presented after rescaling. The solid line represents a fit of the two levels Beaucage equation. For 18 wt% gelatin in D 2 O the contribution from large inhomogeneities are dominant in the low Q regime (Q < 0. 2 nm −1 ) representing a two phase random medium of about R g = 140 nm radius of gyration and a power law exponent of α = 3 in the intermediate Q regime between 0. 02 and 0. 2 nm −1. These data indicate the formation of large-scale gel networks of mass fractal characteristic and of about 350 nm diameter. The fractal characteristic was concluded from the α = 3 exponent and the diameter was evaluated from R g, assuming a spherical shape. 43 The second relevant Q-regime between 0. 2 nm −1 and 1. 5 nm −1 shows a power law exponent of ∼1. 3 indicating the rod-like structure of gelatin triple helix bundles. The third relevant Q-regime beyond 1. 5 nm −1 shows Q −4 Porod law behavior, valid for three dimensional compact structures thereby indicating a shorter helix axis of about d ∼ 1/(Q = 1. 5 nm −1 ) = 0. 7 nm. The length of the helix bundle is roughly estimated as L ∼ 4. 1 nm from R g ∼ 1. 2 nm and d according to R g 2 = L 2 /12 + d 2 /8. 43 Moreover, the transition between the two power law regimes at Q c of approximately 0. 29 nm −1 allows us to estimate the average gelatin mesh size of 2π/Q c ∼ 22 nm. Furthermore, the amplitude of the Porod regime delivers the total surface area per unit volume of the gel according to P 4 = 0. 583 cm −1 nm −4 (P 4 = 2π × N × S × Δρ 2, N particle number density, S particle surface, and Δρ = 4. 71×10 10 cm −2 scattering contrast). 43 For 18 wt% gelatin this value is calculated as N × S ∼ 4. 2 × 10 5 cm −1 (∼31. 6 m 2 /g). These parameters show that the gel of the present study is a good candidate for growing nanoparticles within the mesh of the biopolymers; the high porosity represents a good medium for the growth of nanoparticles without aggregation. SANS data from solutions of varying gelatin concentration are shown in Figure S3a (see Supporting Information). The scattering below Q ∼ 0. 1 nm −1 shows an accumulation of a network structure of similar size and number density proportional to gel concentration. In the intermediate Q-regime we observe a lowering of the power law exponent from ∼2 to ∼1 with increasing gel concentration. The trend to α = 1, i. e. scattering from rod like particles, indicates an enhanced amount of triple helix bundles, which is not accompanied by a significant change of average mesh size (Table S1 ). In conclusion, the choice of appropriate gelatin concentration will achieve both high efficiency for ion transport for optimal iron mineralization as well as high mechanical strength. The latter parameter is known to increase with the amount of gelatin. 2. 2 Ferrogel Synthesis Magnetic field-sensitive gels are called ferrogels. They can be synthesized through various procedures such as blending, in situ co-precipitation or grafting method. 44 Here, we report an in situ mineralization protocol designed for the preparation of ferrogels consisting of biodegradable polymer gelatin and magnetic iron oxide nanoparticles. A three-step process was applied as schematically represented in Figure 2. In a first step, gelatin hydrogels were prepared at different biopolymer concentrations, ranging from 6 to 18 wt% to allow for different mesh sizes in the gelatin gels through concentration dependent variation of the crosslinking degree. These hydrogels were soaked in a solution of Fe II (0. 1 mol L −1 ) and Fe III (0. 2 mol L −1 ) ions with a molar ratio of ferrous to ferric ions of 1:2 until they reached the swelling equilibrium. In a third step magnetite was formed inside the gelatin network after immersing the gel into a NaOH (0. 1 mol L −1 ) solution, which did not affect the gel properties. The porous polymer network structure of the hydrogel in combination with the carbonyl, amine and anionic groups of the gelatin molecules binds the metal cations 45 (see Section 2. 7) and thereby acts as a template for co-precipitation of magnetic nanoparticles according to the following reaction: 2 Figure 2 Schematic representation of the ferrogel synthesis. a) Unloaded gelatin hydrogel, b) hydrogel loaded with ferrous and ferric ions, and c) magnetic nanoparticles distributed inside the hydrogel after in situ co-precipitation with NaOH. The loading of iron ions into the gelatin hydrogel can be followed visually by a change of color from white (native hydrogel) to bright orange (Figure 2 ). The intensity of the color also provides a measure of the concentration of iron present in the gel matrix. Indeed, after several washing steps with water, the iron loaded gel does not change its color, indicating strong binding of the iron ions to the gelatin network. When sodium hydroxide is added to the iron-containing hydrogels the color changed rapidly from bright orange to black denoting the formation of magnetite inside the gelatin matrix. We could repeat the three-step protocol several times in order to control the mineral content. Similar approaches were reported 46, 47 for example Breulmann et al. 16 synthesized magnetite nanoparticles in situ in a polystyrene-polyacrylate copolymer gel to form an inorganic-organic nanocomposite with magnetic and elastic properties. By contrast Reddy et al. 48 precipitated magnetite within a polyacrylamide/gelatin hydrogel matrix in order to produce a biocompatible magnetic hybrid material. The final amount of magnetite nanoparticles inside the biopolymer matrix was determined by thermogravimetric analyses (TGA) in an oxygen atmosphere. From the thermograms, onset temperature and completion of degradation temperature can be identified (see Figure S2 in the Supporting Information). All dry ferrogels investigated show two stages of weight loss. First, there is a minor change in mass between 80 and 180 °C due to removal of moisture from the sample. In a second stage (200–400 °C), the gelatin part of the biopolymer is completely decomposed. Thus, the remaining mass represents the amount of iron oxide originally distributed in the ferrogel. In comparison, for pure magnetite there was no weight loss noted in the interval of 200–800 °C. Our experiments show that after one reaction cycle of magnetite incorporation, mineral contents of ca. 20–30 wt% can be realized. Higher nanoparticle loads of up to 70 wt% are also possible, but require repetition of the reaction cycles with repeated Fe-ion loading and mineralization cycles. In comparison to previous investigations which reported magnetite amounts of up to 30 wt%, 49 the mineral loading in the present study is significantly higher. 2. 3 Crystal Morphology and Size Transmission electron microscopy (TEM) images Figure 3 ) reveal that the applied in situ co-precipitation method led to the formation of magnetic nanoparticles with a mean diameter of about 10 ± 5. 3 nm inside a 10 wt% gel matrix, all synthesized nanoparticles showed similar mean diameters irrespectively of the gelatin concentration. These particles show spherical morphology and a well-developed crystallinity, which is supported by selected-area electron diffraction (SAED) and X-ray diffraction (XRD) analysis (see Figure S1 in the Supporting Information). Notably, the magnetic nanoparticles do not show any uncontrolled aggregation which might be due to colloidal stabilization by gelatin. The arrangement of the crystallites along the gelatin triple helices (see Section 2. 7) can be attributed to the strong protein ion interaction which leads to the alignment along the biopolymer fibers as shown in Figure 3. Gelatin Type B shows an isoelectric point of 4. 7–5. 2, 45 whereas the iron oxide particles in water have a point of zero charge around 7. 50 The measured pH value of the ferrogels after synthesis is around 6, which provides an attractive interaction between the positively charged nanoparticles and the negatively charged gelatin molecules. Figure 3 shows mirco-cuts of an embedded ferrogel sample (10 wt% gelatin after 6 reaction cycles), where a homogenous distribution of the colloidal stable magnetic nanoparticles inside the hydrogel is evident. Interestingly, the TEM data suggest the presence of two distinct particle populations, that is, ca. 10 nm crystalline nanoparticles coexist with small clusters in the size range of 4. 0 ± 1. 1 nm. Moreover, TEM observations of ferrogel samples after different numbers of reaction cycles show that repetition of the precipitation procedure does not influence the size or shape of the nanoparticles but just their number which is in agreement with earlier findings. 16 The SAED pattern in Figure 3 b exhibits diffraction peaks that can be indexed to both magnetite (Fe 3 O 4 ) and maghemite (Fe 2 O 3 ). Due to the similar diffraction patterns of these two magnetic mineral phases, it is not possible to unequivocally distinguish between their crystal structures in ED and XRD analyses. The results obtained by SAED are in agreement with the data collected by XRD. All synthesized hybrid materials display an XRD pattern typical for magnetite or maghemite, with no other impurities being detected. Systematic HR-TEM studies of iron oxide nanoparticles showed lattice spaces attributed to magnetite. The HR-TEM image (Figure 3 a) shows spacings of 0. 48 and 0. 24 nm, which correspond to the (111) and (222) plane of magnetite, respectively. This implies that all collected XRD and ED data represent the inverse spinel mineral magnetite. Figure 3 TEM images of a) and b) ultramicro-cuts of an embedded ferrogel at 10 wt% gelatin concentration after 6 reaction cycles (RC) at different magnifications. Morphological aspects of the dried gelatin hydrogels and the magnetic hybrid gels were further investigated by scanning electron microscopy (SEM). In this study, SEM was used to visualize differences in the gelatin network structure before and after magnetite loading. The prepared samples were dried at room temperature, which leads to a decrease in the sample volume and results in the formation of a contracted porous hybrid material. Cross-sections of dried gelatin and ferrogel samples can be seen in Figure 4. Comparing the pore structure with and without loaded magnetite suggests that the dried ferrogel samples exhibit smaller pore sizes and therefore a denser gel network. Figure 4 b clearly illustrates the homogenous distribution of aggregated magnetite nanoparticles inside the ferrogel at the micrometer scale. Figure 4 Morphology and pore size of two dried hydrogels a) without and b) with magnetite incorporated. 2. 4 Small Angle Scattering Study on the Hybrid Structure In order to further clarify the structure of the gelatin–nanoparticle hybrid material, we performed SANS contrast variation experiments, which allowed to independently explore the inorganic nanoparticle structure as well as the gelatin gel network. Contrast variation SANS experiments became a standard method as pointed out and applied also in cognate disciplines such as biomineralization. 51, 52 ] Figure 5 displays two SANS scattering patterns of a ferrogel dissolved in pure D 2 O and in an aqueous mixture of 28 vol% D 2 O matching the scattering of magnetite and gelatin, respectively. In D 2 O, when gelatin is the only visible part, the scattering at small Q (Q < 0. 2 nm −1 ) delivers a radius of gyration of about 110 nm which is of similar size (140 nm) as found for pure gelatin (Figure 1 ), suggesting scattering from gelatin and the absence of magnetite aggregation as it was also concluded from TEM. The large Q-regime shows stronger scattering than pure gelatin suggesting an increased triple-helix to coil ratio; the fitting shows a slight increase of R g to 1. 7 nm due to a larger amount of triple-helix structure. Figure 5 SANS scattering pattern of the ferrogel in pure D 2 O and in a mixed D 2 O/H 2 O solvent of 28 vol% D 2 O and 72 vol% H 2 O. The solid lines represent the fitting of the Beaucage expression. The form factor of the magnetite is plotted as dashed dotted line. The open circles in Figure 5 represent the scattering of the ferrogel in a 28 vol% D 2 O aqueous solution, which matches the gelatin scattering and visualizes the magnetite nanoparticles. 53 There is weakly enhanced scattering in the small Q-range, which might be caused from non-perfect matching of gelatin or of very small amount of aggregated magnetite. The scattering in the intermediate Q-range is caused from individual nanoparticles of R g = 10. 4 ± 1. 2 nm showing a Q −2 power law at intermediate Q which might indicate a larger size distribution. The diameter D of the magnetite particles can be estimated as D ∼ 27 ± 4 nm (R g = D/2. 58). 43 This value is larger than obtained from TEM (diameter D = 10 nm), but of similar size as the mesh of the biopolymer gel (∼22 nm). The bigger particle diameter obtained by SANS could also result from the dense and chain-like packing of the nanoparticles partially observed along the triple helices between the crosslinking points. This means, that magnetite does not destroy the fractal structure of the ferrogel representing a three-dimensional interconnected network as also seen from the Q −2. 4 power law of the “gelatin” scattering in Figure 5. Moreover, the results suggest that the gel matrix determines the size of the magnetite as it prevents the nanoparticles from further growth as well as from aggregation. TEM also observed smaller particles of 4 ± 1. 1 nm diameter which is consistent with the transition to Porod behavior at Q P = 1. 4 nm −1 (D ∼ 2π/Q p = 4. 4 nm). The scattering of magnetite in Figure 5. shows some slight correlation of the particles which means the observation of a structure factor. The effect of spatial correlation of the magnetite particles becomes more transparent from the X-ray scattering experiments discussed in Figure 6. Figure 5 showing scattering patterns of 12 and 18 wt% ferrogels in wet and dry conditions. In X-ray scattering the contribution of the gel matrix is less than 5%, which means that the scattering is dominated by magnetite. Figure 6 SAXS intensity dΣ/dΩ(Q) versus scattering vector Q for a 18 and 12 wt% wet and dry ferrogel. In all cases the structure factor S(Q) was not negligible. Therefore, the data were fitted with the product of Equation S2 in the SI (structure factor) and the Beaucage equation (form factor) as shown by the solid lines. The dashed dotted lines represent the form factor of the particles. The experimental data are described by Equation 3 representing a product of form factor of 3 with dΣ B /dΩ(Q) and the structure factor S(Q) describing the correlation between the magnetite particles. dΣ B /dΩ(Q) expressed by Equation 1, was also applied for the other SANS data. The fits depicted as solid lines describe the data sufficiently well as already demonstrated for the SANS data in Figure 5. The dashed dotted lines represent the form factor (dΣ B /dΩ(Q)) of magnetite. The expression of the structure factor is given in Equation S3 (SI) as derived on basis of the hard sphere potential in Equation S2. The parameters of this analysis are compiled in Table S2 and S3 for small-angle X-ray scattering (SAXS) and SANS, respectively. For the 18 wt% wet ferrogel we find particles of 10 nm from SAXS and SANS whereas a slightly larger negative exponent of 2. 73 (instead of 2) for the SAXS data. In contrast to SANS no transition to Q −4 was observed for these solutions from SAXS. The 12 wt% wet ferrogel shows a smaller R g of 8. 7 nm and Q −3 power law at larger Q. The dry samples show much stronger scattering and correlation between magnetite because of their enhanced dense packing. At intermediate Q one has Q −3 which at Q = 2. 3 nm −1 transforms to Q −4 power law. The Q −3 behavior suggests a composite mass fractal structure of magnetite. The correlation peak of the dry samples at Q m ∼ 0. 4 nm −1 provides an average distance of the scattering particles of Λ ∼ 10 nm 54 (Guinier; Λ = 1. 23π/Q m ), which is almost the same as R g of the particles. This means that in the dry ferrogel the nano-particles are closely packed. In summary the SANS and SAXS data suggest that the gelatin content has no significant influence on the size of the nanoparticles, i. e. the gel matrix of different concentrations has similar average mesh size which is of similar size as the nanoparticles. Parameters are shown in Table S1, S2, and S3. 2. 5 Magnetic Measurements The magnetic properties of the composite materials were characterized by using a superconducting quantum interference device (SQUID) magnetometer. Figure 7 a shows the magnetization (M) of a representative dried ferrogel sample (10 wt% gelatin in the hydrated gelified state, 60. 4 wt. -% mineral content in the dried ferrogel state) as a function of the applied field (H) at 293 K and 2 K. At T = 2 K the magnetization curve shows typical ferrimagnetic hysteresis due to magnetic anisotropy. At 293 K there is no hysteresis observed, which is typical for superparamagnetic materials, 55 and consistent with the small size of the nanoparticles. The field-cooled (FC) and zero-field-cooled (ZFC) magnetizations of the magnetite-gelatin composites were also measured (Figure 7 b). The maximum of the ZFC curve corresponds to the blocking temperature (TB). 56 Values obtained for TB as well as for the saturation magnetization (Ms) at 5000 Oe are listed in Table 1 for several representative samples. The studies show a blocking temperature of around 120 K which also confirms the superparamagnetic behavior of the nanoparticles. The higher TB found for the ferrogel at a lower gelatin concentration (cf. Table 1 ) might be due to stronger dipolar interactions and dense particle packing in the dry ferrogel state. As can be seen in Table 1 the saturation magnetization (Ms) of the composite materials is lower than that of bulk magnetite (92 emu/g) as well as maghemite (56 emu/g). 57 This effect has been observed in many previous studies, and it was proposed that with decreasing particle size, the growing degree of spin disorder at the surface causes the decrease in Ms. 58 It has also been reported that defects on the particle surface can influence the magnetic properties. 58 The obtained magnetic measurement data show the phenomena of superparamagnetism for the designed composite materials, the same result is also observed for magnetite nanoparticles synthesized by a co-precipitation method in water. 57, 59 – 62 From these observations we conclude that the gelatin network has no effect on the magnetic properties of magnetite nanoparticles synthesized inside the gel matrix. Table 1 Saturation magnetization (MS, measured at 5000 Oe) and blocking temperature (TB) values of selected ferrogel samples with varying gelatin concentrations after 8 reaction cycles gelatin conc. (in hydrogel state) [wt%] magnetite cont. (in dry ferrogel) [wt%] TB [K] MS 2 K [emu/g] MS 293 K [emu/g] 8 61. 6 150 36. 85 27. 70 10 60. 2 134 36. 00 26. 26 12 60. 0 130 38. 43 28. 19 14 52. 8 126 36. 16 26. 40 bulk magnetite 92 bulk maghemite 57 56 Figure 7 Magnetic properties of the synthesized hybrid materials. a) Magnetization curves of a dried ferrogel at 2 K and 293 K. Inset: Enlargement of the low field region showing the different coercive fields for the NPs at 2 and 293 K. b) ZFC-FC curves as a function of temperature. 2. 6 Swelling Studies Swelling studies were conducted in order to probe structural changes in the gelatin network upon incorporation of the magnetite nanoparticles. To that end, the water uptake of dried gel pieces and ferrogels was measured gravimetrically, until swelling equilibrium was reached. The swelling degree (Sd) of the investigated gels is given by the following equation: 4 wherein Ws represents the weight of the swollen hydrogel after swelling equilibrium was reached and Wd is the dry weight of the as-prepared xerogels. The swelling experiments were performed with samples containing 6 to 18 wt% gelatin and after 1, 3, and 6 mineralization reaction cycles (RC). Figure 8 shows the swelling behavior of representative ferrogel samples with different mineral content compared to plain gelatin reference samples. It is evident that the ferrogel after 1 RC already shows a more pronounced increase in the degree of swelling compared to the neat gelatin reference. This unexpected effect might be due to the incorporation of the positively charged iron oxide nanoparticles into the polymer matrix, which can increase the osmotic pressure and therefore increase the swelling propensity of the ferrogel. On the other hand, we observe that as the amount of magnetic nanoparticles in the matrix is further increased (i. e. after 3 and 6 reaction cycles), the ferrogels show a systematically decreasing swelling tendency. This result can be attributed to an attractive interaction between the iron oxide nanoparticles and the gelatin polymer matrix, potentially involving the carboxylate groups of gelatin, which can act as iron binding sites. Hence the small crystallites can act as points of crosslinking and therefore strengthen the gelatin hydrogel structure, leading to an effective decrease of the swelling degree and thus in the gravimetric water uptake. These observations are in line with the results obtained from SANS and SAXS studies. In summary, these experiments have shown that the introduction of nanoparticles into the gelatin matrix has a pronounced effect on its swelling behavior. Therefore we conclude that the structure of the gelatin hydrogel changes with varying content of magnetic nanoparticles inside the matrix. Figure 8 Degree of hydrogel swelling plotted as a function of the swelling time at 25 °C for different samples with a gelatin concentration of 10 wt%. The equilibrium swelling degrees Sd (%) for the plotted samples are 779. 2 ± 9. 6 (gelatin), 1531. 4 ± 62. 0 (RC 1), 684. 65 ± 80. 84 (RC 3) and 195. 64 ± 0. 26 (RC 6). 2. 7 Simulation Studies We performed molecular simulation studies of Fe 2+ /Fe 3+ and hydroxide ion association to a triple helical (Gly-Hyp-Pro)n peptide to characterize the interplay of collagen and inorganic nanoparticle formation on the molecular scale. To allow direct comparison, the collagen fragment and the simulation method is chosen in full analogy to earlier studies on calcium and phosphate ion association to collagen. 63 From this, favorable association sites for both Fe II (OH) 2 and Fe III (OH) 3 ion clusters were identified. Figure 9 illustrates representative constellations as observed for each species. It is noteworthy, that both precursors to magnetite bind to collagen via hydrogen bonds and salt bridges without distorting the triple helix. Instead, Fe(OH) x binds to carbonyl and hydroxyl groups which oxygen atoms tend to complete an octahedral coordination polyhedral for either Fe 3+ and Fe 2+ association. The close interplay of Fe(OH) x motifs and collagen as observed from molecular simulation hints at the suitability of collagen to bind iron and hydroxide ions (with the later only forming stable bonds in combination with iron ions). From this we conclude that collagen acts as a nucleation seed to iron hydroxide aggregation, and thus intergrowth of collagen and magnetite nanoparticles already at the precursor stage. Moreover, the TEM micrographs of the final magnetite-collagen composites indicate a structural alignment of the nanoparticles (Figure 3 ), which we attribute to magnetite nucleation along collagen fibers. This interplay of organic and inorganic components could give rise to hierarchical composites as observed for calcium phosphate–collagen based biominerals. 64 Figure 9 (left) Representative structure for Fe III (OH) 3 coordination by collagen. Note that three carbonyl/hydroxyl groups are providing O·Fe salt bridges via one short (2. 3 Å) and two weaker (2. 6 Å) contacts. (right) Fe II (OH) 2 cluster coordination by collagen leading to distorted/incomplete octahedral coordination of Fe II (the number of coordinating water molecules from the solvent varies from 0 to 2). Atom colors: Fe (yellow), O (red/green for solvent), H (white), N(blue) and C(grey). 3 Conclusions We have reported a simple synthesis procedure to produce ferrogels with a biocompatible gelatin gel matrix and magnetite nanoparticels. The repetition of the reaction cycles (RC) allows variation of the mineral content between 20 wt% (1 RC) to 70 wt% (8 RC) to form a highly mineralized inorganic-organic hybrid material. Since gelatin gels are thermoreversible, they can be moulded into any shape prior to mineralization which is a big advantage concerning applications. Once they are mineralized, their melting following a temperature decrease is significantly hindered likely due to the introduction of additional crosslinks introduced by the interaction of the magnetite nanoparticles with the gelatin matrix. The structure of the ferrogels was characterized with respect to gelatin as well as magnetite nanoparticles using SANS contrast matching, which is able to individually access the structure of each individual compound over the entire colloidal range as well as SAXS only visualizing magnetite. SANS shows with respect to gel concentration an unchanged gelatin structure of average mesh size larger than the nanoparticles. The size of the nanoparticles seems to be limited by the gel mesh size and independent of gelatin concentration between 6 and 18 wt%. SANS shows no aggregation of magnetite in agreement with TEM. Magnetite particles itself show spatial correlations in SANS and SAXS due to excluded volume interaction, which particularly become strong for the dry samples. The corresponding structure factor is described on basis of hard core interaction. The structural parameters of the gelatin hydrogel are compiled in Table S1, which were determined from the USANS and SANS data in Figure 1 and S3a. Increasing the gel concentration between 6 and 30 wt% forms hydrogel networks of enhanced radii of gyration between 156 and 191 nm. The scattering of the network shows a Q −3 power law indicating a mass fractal structure of the network. 43 The scattering at large Q delivers a mesh size of about 20 nm for all gel samples. A similar SANS study on hydrogels is found in a recent publication. 65 SQUID measurements showed that the NPs are superparamagnetic as expected for this particle size and have a similar blocking temperature as compared to pure magnetite in this size range. The saturation magnetization of the synthesized NPs is lower than that for bulk magnetite which is likely a result of surface defects of the nanoparticles. Swelling measurements showed that the adsorption of the magnetic nanoparticles onto the polymeric matrix influences and limits the swelling behavior of the ferrogels. This supports the finding that the gelatin gels lose their thermoreversible properties after magnetite inclusion as additional crosslinker. The degree of swelling can consequently be controlled by the amount of mineral inside the biopolymer. Therefore we can vary the mechanical properties of the ferrogels, which is topic of a subsequent study. Overall, we have succeeded in the simple preparation of gelatin-based ferrogels with a constant gel structure, but the possibility of a largely varying mineral content. Since the mineral particles are not washed out and also, the gelatin based ferrogels do not dissolve anymore even in an excess of water due to their additional electrostatic crosslinking by the nanoparticles, they have promising applications as biocompatible actuators which can be driven by external magnetic fields as can be seen in Figure 10. Attraction of ferrogel with a) no magnetic field and b) external magnetic field (ca. 1T) 4 Experimental Section Chemicals : The following commercially available chemicals were purchased and applied in the synthesis without further purification: FeCl 2 ·4H 2 O (Sigma-Aldrich), FeCl 3 ·6H 2 O (Sigma-Aldich), 0. 1 M NaOH solution (Merck), Gelatin Type B (∼225 Bloom, Sigma-Aldrich), 4-chloro- m -cresol (Fluka), Methanol (VWR). For the preparation of the reactant solutions double-distilled and deionized (Milli-Q) water was used. All solutions were degassed with argon before usage. Gel Synthesis : Different amounts of gelatin were allowed to swell in water for 24 hours at 6 °C. Homogeneous solutions were prepared by heating these gels for 2 hours at 50 °C. In each case, 2 mL of solution is filled in disposable base molds (30 mm × 24 mm × 5 mm) and allowed to form a gel there. To avoid decomposition by bacteria, a 5 wt% solution of 4-chloro- m -chresol in methanol was added (0. 15 mL per 1 g of dry gelatin). Synthesis of Ferrogel : In situ mineralization of magnetite nanoparticles in gelatin hydrogel was carried out via co-precipitation of FeCl 2 and FeCl 3. Each gelatin hydrogel sample was introduced into a solution, containing FeCl 2 (0. 1 M) and FeCl 3 (0. 2 M), where it was left for 96 hours at 6 °C. The iron (II) and iron (III)-loaded gels were washed with water and placed in 0. 1 M NaOH solution for 150 min. Transmission Electron Microscopy (TEM) : For TEM and HR-TEM analysis, a Zeiss Libra 120 operating at 120 keV and a JEOL JEM-2200FS operating at 200 keV were used, respectively. For material characterization, two distinct sample preparation techniques were applied. On the one hand, a drop of a diluted dispersion of magnetic nanoparticles extracted from the hydrogel was placed on a Formvar coated copper grid and left to dry on a filter paper. On the other hand, the grid was dipped inside the hydrogel matrix and aliquots were blotted using a filter paper. For micro-cut preparation, the ethanol dehaydrated ferrogel samples were embedded in LR white Resin (Medium Grade) and cut with a Leica EM Trim. Scanning Electron Microscopy (SEM) : For SEM analysis a Zeiss Neon 40 EsB operating in high vacuum was used. An InLens and SE detector was used for signal collection and an acceleration voltage of 2 kV was chosen for recording the images. The specimens were coated with a thin layer of gold in order to avoid charging effects. Thermogravimetric Analysis (TGA) : The mineral content of the hydrogels was determined by means of TGA (Netzsch, Selb, Germany). Measurements were carried out at a heating rate of 5 K/min under a constant oxygen flow. Samples were scanned from 293 K to 1273 K. Small-Angle Neutron Scattering SANS and USANS : Neutron experiments were carried out at, respectively, the KWS1 and KWS 3 diffractometers of JCNS outstation at FRM II in Garching, Germany. 66 Three configurations were used at KWS 1, namely the sample-to-detector (SD) distances of 2, 8 and 20 m, the corresponding collimation length of 8 and 20 m, and a wavelength of 0. 7 nm (Δλ/λ = 10%). These settings allowed covering a Q-range from 0. 02 to 3. 5 nm −1. The scattering vector Q is defined as with the scattering angle δ and the wavelength λ. A two-dimensional local sensitive detector was used to detect neutrons scattered from sample solutions. Gel solutions were filled in rectangular quartz cells with path-length of 1 or 2 mm. Plexiglas was used as secondary standard to calibrate the scattering intensity in absolute units. The data correction and calibration were performed using the software QtiKWS (V. Pipich (2013)). 67 In order to cover the broader length scale of the network structure, ultra-small-angle Neutron scattering (USANS) experiments were carried out at the KWS3 diffractometer using parabolic mirror as an optical element, and covering the smaller Q range from 0. 001 to 0. 02 nm −1. 66 Small-Angle X-ray Scattering (SAXS) : SAXS experiments were carried out at a HECUS S3-Micro small-angle X-ray scattering instrument. The instrument uses Cu Kα radiation (0. 154 nm) produced in a sealed tube. Gel samples were placed in Hilgenberg quartz capillaries with an outside diameter of 1 mm and wall thickness of 0. 01 mm. The scattered intensity was corrected with the transmission of the samples calculated considering the absorption of the sample and that of the capillary. The dry gel samples were cut to a thin film with a thickness of 1 mm and measured directly. The scattered X-rays are detected with a two-dimensional multiwire area detector and afterwards converted to one-dimensional scattering by radial averaging and represented as a function of momentum transfer vector Q similar to the SANS experiments. Magnetization Studies Using Superconducting Quantum Interference Device (SQUID) : Magnetization measurements were carried out by using a quantum design SQUID 5 T magnetic properties measurement system (MPMS). For measurements, dried ferrogels were introduced into gelatin capsules and magnetization loop measurements at 2 K and 293 K were performed. In addition zero-field-cooled and field-cooled curves were obtained by applying 0. 01 T and heating or cooling the sample. Swelling Studies : For the drying process the gel samples were left at room temperature for at least 5 days. Air-dried samples of different concentrations of gelatin hydrogel, iron-loaded hydrogel and ferrogels with different numbers of reaction cycles were weighed in the dry state. The samples were left for swelling in 30 ml Milli Q water in a closed vessel at RT. Before weighing, the excess water of each sample was removed with a filter paper. All samples were weighed after a certain amount of time until equilibrium swelling was reached. Molecular Simulation : a series of Fe III (OH) x (OH 2 ) 4-x and Fe II (OH) y (OH 2 ) 8-y clusters were pre-modeled from ab-initio calculations in vacuum. For all clusters, the high-spin constellation was identified as preferred by several electron volts. Imposing overall charge neutrality (i. e. x+y = 3+2) we found the neutral Fe III (OH) 3 · H 2 O and Fe II (OH) 2 · 6 H 2 O clusters as energetically preferred. Docking to collagen was modeled in aqueous solution using empirical force-fields. 68 Investigation of biologically-designed metal-specific chelators for potential metal recovery and waste remediation applications, 69 and the Kawska-Zahn docking procedure were described previously. 70 Along this line, ion clusters are initially docked to collagen in absence of water. Such putative association complexes are then immersed in aqueous solution (periodic simulation cell comprising more than 15 000 water molecules) and subjected to relaxation from 100 ps molecular dynamics runs at room temperature and ambient pressure. To account for the manifold of possible arrangements intrinsic to the systems complexity, a series of 100 independent docking runs were performed for each ionic species. The resulting structures were then classified in terms of hydrogen bonds and O·Fe distances to discriminate the representative configurations of the Fe III (OH) 3 · O (collagen) and Fe II (OH) 2 · O (collagen or H 2 O) coordination constellations as shown in Figure 9.
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10. 1002/adfm. 201303655
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Advanced functional materials
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Composite Living Fibers for Creating Tissue Constructs Using Textile Techniques
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The fabrication of cell-laden structures with anisotropic mechanical properties while having a precise control over the distribution of different cell types within the constructs is important for many tissue engineering applications. Automated textile technologies for making fabrics allow simultaneous control over the color pattern and directional mechanical properties. The use of textile techniques in tissue engineering, however, demands the presence of cell-laden fibers that can withstand the mechanical stresses during the assembly process. Here, the concept of composite living fibers (CLFs) in which a core of load bearing synthetic polymer is coated by a hydrogel layer containing cells or microparticles is introduced. The core thread is drawn sequentially through reservoirs containing a cell-laden prepolymer and a crosslinking reagent. The thickness of the hydrogel layer increases linearly with to the drawing speed and the prepolymer viscosity. CLFs are fabricated and assembled using regular textile processes including weaving, knitting, braiding, winding, and embroidering, to form cell-laden structures. Cellular viability and metabolic activity are preserved during CLF fabrication and assembly, demonstrating the feasibility of using these processes for engineering functional 3D tissue constructs.
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No full text available
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10. 1002/adfm. 201400274
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Advanced functional materials
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Digital Plasmonic Patterning for Localized Tuning of Hydrogel Stiffness
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The mechanical properties of the extracellular matrix (ECM) can dictate cell fate in biological systems. In tissue engineering, varying the stiffness of hydrogels—water-swollen polymeric networks that act as ECM substrates—has previously been demonstrated to control cell migration, proliferation, and differentiation. Here, “digital plasmonic patterning” (DPP) is developed to mechanically alter a hydrogel encapsulated with gold nanorods using a near-infrared laser, according to a digital (computer-generated) pattern. DPP can provide orders of magnitude changes in stiffness, and can be tuned by laser intensity and speed of writing. In vitro cellular experiments using A7R5 smooth muscle cells confirm cell migration and alignment according to these patterns, making DPP a useful technique for mechanically patterning hydrogels for various biomedical applications.
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No full text available
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10. 1002/adfm. 201400526
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Advanced functional materials
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Highly tunable elastomeric silk biomaterials
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Elastomeric, fully degradable and biocompatible biomaterials are rare, with current options presenting significant limitations in terms of ease of functionalization and tunable mechanical and degradation properties. We report a new method for covalently crosslinking tyrosine residues in silk proteins, via horseradish peroxidase and hydrogen peroxide, to generate highly elastic hydrogels with tunable properties. The tunable mechanical properties, gelation kinetics and swelling properties of these new protein polymers, in addition to their ability to withstand shear strains on the order of 100%, compressive strains greater than 70% and display stiffness between 200 – 10, 000 Pa, covering a significant portion of the properties of native soft tissues. Molecular weight and solvent composition allowed control of material mechanical properties over several orders of magnitude while maintaining high resilience and resistance to fatigue. Encapsulation of human bone marrow derived mesenchymal stem cells (hMSC) showed long term survival and exhibited cell-matrix interactions reflective of both silk concentration and gelation conditions. Further biocompatibility of these materials were demonstrated with in vivo evaluation. These new protein-based elastomeric and degradable hydrogels represent an exciting new biomaterials option, with a unique combination of properties, for tissue engineering and regenerative medicine.
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No full text available
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10. 1002/adfm. 201401300
| 2,014
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Advanced functional materials
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Layer-by-layer assembly of 3D tissue constructs with functionalized graphene
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Carbon-based nanomaterials have been considered as promising candidates to mimic certain structure and function of native extracellular matrix materials for tissue engineering. Significant progress has been made in fabricating carbon nanoparticle-incorporated cell culture substrates, but limited studies have been reported on the development of three-dimensional (3D) tissue constructs using these nanomaterials. Here, we present a novel approach to engineer 3D multi-layered constructs using layer-by-layer (LbL) assembly of cells separated with self-assembled graphene oxide (GO)-based thin films. The GO-based structures are shown to serve as cell adhesive sheets that effectively facilitate the formation of multi-layer cell constructs with interlayer connectivity. By controlling the amount of GO deposited in forming the thin films, the thickness of the multi-layer tissue constructs could be tuned with high cell viability. Specifically, this approach could be useful for creating dense and tightly connected cardiac tissues through the co-culture of cardiomyocytes and other cell types. In this work, we demonstrated the fabrication of stand-alone multi-layer cardiac tissues with strong spontaneous beating behavior and programmable pumping properties. Therefore, this LbL-based cell construct fabrication approach, utilizing GO thin films formed directly on cell surfaces, has great potential in engineering 3D tissue structures with improved organization, electrophysiological function, and mechanical integrity.
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No full text available
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10. 1002/adfm. 201500875
| 2,016
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Advanced functional materials
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Protein Corona Influences Cell–Biomaterial Interactions in Nanostructured Tissue Engineering Scaffolds
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Biomaterials are extensively used to restore damaged tissues, in the forms of implants (e. g. tissue engineered scaffolds) or biomedical devices (e. g. pacemakers). Once in contact with the physiological environment, nanostructured biomaterials undergo modifications as a result of endogenous proteins binding to their surface. The formation of this macromolecular coating complex, known as ‘protein corona’, onto the surface of nanoparticles and its effect on cell-particle interactions are currently under intense investigation. In striking contrast, protein corona constructs within nanostructured porous tissue engineering scaffolds remain poorly characterized. As organismal systems are highly dynamic, it is conceivable that the formation of distinct protein corona on implanted scaffolds might itself modulate cell-extracellular matrix interactions. Here, we report that corona complexes formed onto the fibrils of engineered collagen scaffolds display specific, distinct, and reproducible compositions that are a signature of the tissue microenvironment as well as being indicative of the subject's health condition. Protein corona formed on collagen matrices modulated cellular secretome in a context-specific manner ex-vivo, demonstrating their role in regulating scaffold-cellular interactions. Together, these findings underscore the importance of custom-designing personalized nanostructured biomaterials, according to the biological milieu and disease state. We propose the use of protein corona as in situ biosensor of temporal and local biomarkers.
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No full text available
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10. 1002/adfm. 201501277
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Advanced Functional Materials
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Modular and Versatile Spatial Functionalization of Tissue Engineering Scaffolds through Fiber‐Initiated Controlled Radical Polymerization
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Native tissues are typically heterogeneous and hierarchically organized, and generating scaffolds that can mimic these properties is critical for tissue engineering applications. By uniquely combining controlled radical polymerization (CRP), end‐functionalization of polymers, and advanced electrospinning techniques, a modular and versatile approach is introduced to generate scaffolds with spatially organized functionality. Poly‐ε‐caprolactone is end functionalized with either a polymerization‐initiating group or a cell‐binding peptide motif cyclic Arg‐Gly‐Asp‐Ser (cRGDS), and are each sequentially electrospun to produce zonally discrete bilayers within a continuous fiber scaffold. The polymerization‐initiating group is then used to graft an antifouling polymer bottlebrush based on poly(ethylene glycol) from the fiber surface using CRP exclusively within one bilayer of the scaffold. The ability to include additional multifunctionality during CRP is showcased by integrating a biotinylated monomer unit into the polymerization step allowing postmodification of the scaffold with streptavidin‐coupled moieties. These combined processing techniques result in an effective bilayered and dual‐functionality scaffold with a cell‐adhesive surface and an opposing antifouling non‐cell‐adhesive surface in zonally specific regions across the thickness of the scaffold, demonstrated through fluorescent labelling and cell adhesion studies. This modular and versatile approach combines strategies to produce scaffolds with tailorable properties for many applications in tissue engineering and regenerative medicine.
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1 Introduction The ability to design surface properties of scaffolds to direct cell and protein binding is key in tissue engineering (TE) and regenerative medicine. Following injury or scaffold implantation there are many scenarios where cell ingrowth or protein fouling onto a tissue or scaffold surface may lead to undesirable outcomes such as problematic scar tissue. This is particularly key at gliding tissue interfaces. Scaffolds that exhibit effective opposing cell‐adhesive and antifouling sides would revolutionize the outcome of injuries and operations at such biological interfaces through their ability to support tissue healing (cell‐adhesive surface) and to reduce cell‐ingrowth and protein adsorption (antifouling surface) in a spatially controlled manner. Injuries or operations to gliding surfaces such as musculoskeletal tissues or involving the peritoneal tissues of the abdomen or pelvis are often complicated with undesirable adhesions. These adhesions are bands of scar tissue that directly result from cellular ingrowth and bridging between previously gliding surfaces resulting in restriction of movement that causes pain and compromised function. Similarly, protein deposition at vascular interfaces such as on artificial blood vessels or at sites of injury can lead to blood clotting that can cause disastrous consequences. This work presents a modular and versatile scaffold system that has been specifically designed to allow facile modification for specific application. A porous structure manufactured with a processing technique that can add to this versatility is required; this can be achieved with electrospinning. Electrospinning is an ideal technique for producing 3D networks of fibers of tunable size, orientation, composition, and density that mimic the properties of native extracellular matrix (ECM) 1, 2, 3 and can generate scaffolds with spatially arranged functionalization through layering of various polymers during electrospinning. 4, 5, 6, 7, 8, 9 Developing a controlled method that allows multiple zonally arranged functional groups within a continuous scaffold allows for the production of a hierarchical structure that can modulate cell behavior within each functional zone. Electrospinning and the use of fibers also allows for high density surface functionalization that can dramatically change the surface properties whilst maintaining the spatial control of presentation and fiber morphology. Controlled radical polymerization (CRP) techniques are facile and versatile methods for providing surface functionalization with a wide range of monomers. These have not been fully exploited in scaffold functionalization and yet are extremely powerful methods for preparing antifouling surfaces. Atom transfer radical polymerization (ATRP) and reversible addition‐fragmentation chain transfer (RAFT) have been used to attach polymers to a variety of surfaces to generate surface‐derived functionality. 10, 11, 12, 13 These polymerization methods are versatile and offer excellent control of chain length, architecture, reaction kinetics, and they can add a vast array of functionality as a large number of monomers may be incorporated, which dictates the final material properties. 10, 14 In this work we have employed the versatility of surface‐initiated CRP to create an antifouling polymer bottlebrush on one side of an electrospun poly‐ ε ‐caprolactone (PCL) scaffold with a cell‐adhesive peptide on the opposing surface across a diameter of a few hundred micrometers. The antifouling surface is generated by grafting a high density, antifouling, biocompatible polymer bottlebrush based on poly(ethylene glycol) (PEG) from initiators presented on the PCL scaffold surface. PEG is an antifouling polymer that has been used to mediate cell and protein adhesion to surfaces and has been used in devices approved for implantation into the body. 15 The ability to incorporate additional (bio)functionalities within the antifouling brushes is showcased with the addition of a biotinylated monomer, which can provide a versatile handle for the post hoc addition of various streptavidin‐coupled moieties. To maximize the antifouling ability by creating a dense hydrated polymer layer, we elected to use the “grafting from” approach, whereby the initiating group is attached to the surface and the polymer grows out from it. This avoids the steric hindrance and resultant low density that is found in a “grafting to” approach. 16 Second, we selected the oligomeric monomer of PEG, poly(ethylene glycol) methyl ether methacrylate (OEGMA) to generate a pOEGMA bottlebrush structure that leads to a vastly higher density of PEG being displayed on the surface for superior performance. Polymer brush growth from electrospun fibers has typically been achieved using ATRP polymerization of a variety of different monomers. In most approaches that are directed towards biomedical applications, the initiating group is incorporated as a post electrospinning modification before polymerization has been undertaken. 17, 18, 19, 20, 21, 22 Our strategy offers significant advantages to this by incorporating the initiator as an end‐group to the polymer prior to electrospinning to allow precise control over the spatial position of the functional groups without disrupting the fiber architecture. This approach has been used previously for the polymerization of styrene, 23 2‐hydroxyethyl methacrylate, 16 and N, N ‐isopropylacrylamide, 24 but not in a biomedical application. The selection of the initiating group and electrospinning arrangement further allows surface enrichment of the fibers with initiating groups through electrostatic attraction. 23 Second, a modified form of ATRP, namely, activators regenerated by electron transfer atom transfer radical polymerization (ARGET ATRP), was employed that avoids the high concentrations of potentially toxic transition metal catalyst and organic solvents used in conventional ATRP. 25 This is critical for scaffolds designed for biomedical use as high levels of contamination can be difficult to thoroughly remove from a bulk scaffold. Finally, in contrast to conventional ATRP, ARGET is less sensitive to small amounts of oxygen contamination, offering increased ease of use. 26 To our knowledge this is the first use of controlled radical polymerization to polymerize pOEGMA from a PCL surface. Our approach of pre‐functionalizing PCL with end‐functional groups before electrospinning offers a facile and versatile method for maintaining scaffold architecture, functionality, and material properties whilst having precise control over the spatial location and molecular weight of the grafted polymer. It also allows control over the density of functional groups by simply changing the concentration of the PCL conjugates. Specifically, we modified one batch of PCL with the initiating group for polymerization and a separate batch with the canonical adhesion peptide sequence Arg‐Gly‐Asp‐Ser (RGDS). The polymer conjugates were sequentially electrospun as a bilayer to achieve spatial control of the functionality and surface properties. This work builds from our recent work using sequential electrospinning to form opposing gradients of two different peptides in a PCL scaffold, which directed the specific binding and spatial organization of biopolymers (glycosaminoglycans, GAG) within the scaffold. 27 These techniques provide a new and versatile platform for the preparation of multi‐functional TE scaffolds to address unmet clinical need in orthopedic, plastic, reconstructive, and general surgery. 2 Results and Discussion 2. 1 Production of Polymer Bottlebrush through Surface‐Initiated Polymerization ARGET ATRP reaction kinetics were first established and optimized in solution; polymerizations were conducted in water/IPA (1:1 v/v) in order to prevent dissolution of the PCL. A molar ratio of 150:1:1:0. 15 OEGMA : initiator : Cu(II)Cl 2 : sodium ascorbate at 30°C resulted in reproducible >70% conversion within 2 h. Good control was achieved as evidenced by pseudo first‐order kinetics and the low molecular weight dispersity of the polymers ( M w / M n < 1. 3). Surface‐initiated ARGET ATRP of pOEGMA was successfully performed from 2D silicon surfaces, silicon wafers, optimized and characterized, before progressing to 3D electrospun scaffolds to ensure reproducible results. The initiating group α‐bromoisobutyryl bromide (BiBB) was attached to the surfaces using 3‐(aminopropyl)triethoxysilane (APTES) before pOEGMA was grafted from the wafer. A free sacrificial initiator, ethyl‐α‐bromoisobutyrate (EBiB), was used in solution to aid control of the polymerization and to allow analysis of the free polymer as a surrogate for the surface bound polymer; this has been shown to be a reliable tool for controlling the M n and M w / M n for the polymers grown from surfaces within the same reaction vessel. 28 Ellipsometry and atomic force microscopy (AFM) confirmed the presence of a polymer brush layer with a dry thickness of ≈6. 2 nm (±0. 038 nm, MSE 3. 548) ( Figure 1 B, C and Figure S1, Supporting Information), and X‐ray photoelectron spectroscopy (XPS) confirmed this layer to be organic with the expected changes in the ratio of the C—O bond (Table S1, Supporting Information). Further evidence for the successful polymerization of OEGMA from the surface is given by the increased wettability after polymerization with a change in water contact angle from 63. 6° ± 2. 3° on Si‐APTES‐Ini functionalized wafers to 36. 3° ± 5. 9° (Figure 1 D). Figure 1 Demonstration and characterization of surface‐initiated polymer brush growth from functionalized 2D silicon surfaces. A) XPS analysis of pOEGMA grafting from a silicon wafer functionalized with APTES‐Ini with controls (dashed bottom trace), silicon functionalized with APTES‐Ini that underwent polymerization with no reducing agent, ascorbic acid (AScA, dotted middle trace), and pOEMGA grafting from a silicon wafer (top trace, left). Conversion by 1 H‐NMR (X) is included above each trace. Si‐APTES‐pOEGMA sample with the C1s peaks fitted (right). B) AFM scratch test and C) representative profile of pOEGMA grafted from silicon wafers. D) Water contact angle measurement of a silicon wafer functionalized with (i) APTES‐Ini and (ii) following grafting of pOEGMA. 2. 2 Surface‐Initiated Polymerization from 3D Electrospun Fibers To undertake surface‐initiated polymerization from 3D electrospun fibrous scaffolds we commenced by modifying a PCL‐diol ( M w 14 000 Da) with the initiating group (BiBB) to produce a polymerization initiating region at either end of the PCL polymer chain (PCL‐Ini). PCL was selected to form the bulk of the scaffold due to its bioresorbability, good handling properties, electrospun fiber morphology, suitable degradation rate, ease of chemical modification, and its current use in Food and Drug Administration (FDA)‐approved devices. 29 To ensure that the initiating region of the PCL conjugates was presented on the fiber surface, we set up the electrospinner with the cathode at the spinnerette to convey a positive charge to the surface of the polymer solution. The alkyl‐bromide group present within BiBB can become electronegative due to its polarity, 23 dragging it electrostatically to the surface of the polymer solution stream and leading to surface enrichment of the initiating group on the fiber. The PCL‐Ini was subsequently electrospun in combination with a high molecular weight PCL to form functionalized fibrous scaffolds which were then imaged by scanning electron microscopy (SEM) to validate consistent fiber morphology ( Figure 2 A). The addition of up to 17% (w/w) of the PCL‐Ini did not significantly alter the electrospinning process or fiber formation (Figure S2, Supporting Information). pOEGMA bottlebrushes were grafted from the electrospun fibers using the parameters optimized in the 2D silicon system (Figure 2 B). As with the 2D silicon surfaces, a sacrificial initiator was included in order to target a degree of polymerization (DP) of 150. A typical polymerization achieved a ≈75% conversion (by 1 H‐NMR) and M n 45 000, M w 53 000 with a dispersity ( M w / M n ) of 1. 18 (from size exclusion chromatography, SEC, analysis of the free polymer). Figure 2 Grafting of pOEGMA bottlebrushes from prefunctionalized electrospun scaffolds to create an antifouling, non‐cell‐adhesive surface as part of a dual functional scaffold. A) Representative SEM micrograph of electrospun PCL‐pOEGMA fibres. B) ARGET ATRP reaction scheme for polymerization of pOEGMA from the PCL‐Ini fibers with inset schematic images of PCL‐Ini following electrospinning (left) and following polymerization of pOEGMA from the fiber surface (right). C) Schematic outlining the bifunctional scaffold structure produced using layered electrospinning with postprocessing polymerization to create an antifouling PCL‐pOEGMA surface with an opposing cell binding PCL‐cRGDS surface. XPS confirmed successful grafting of pOEGMA from functionalized electrospun scaffolds as a large increase in C—O signal is observed following polymerization for both the 17% PCL‐Ini (w/w) and the 9% PCL‐Ini (w/w) ( Figure 3 A). This is in contrast to the control scaffold lacking in initiating groups that shows minimal increase in the C—O signal confirming successful covalent attachment of polymer to the scaffolds through the initiating groups (Table S1, Supporting Information). A dramatic change in the water contract angle adds further evidence for successful pOEGMA grafting. Hydrophobic electrospun PCL/PCL‐Ini scaffold surfaces, with contact angles of 113. 5° ± 7. 8°, become hydrophilic after polymerization, with the water droplet immediately completely wetting the surface (Figure 3 B). Figure 3 Demonstration and characterization of surface‐initiated polymer brush growth from functionalized 3D electrospun scaffolds. A) High resolution C1s core‐level spectra of pOEGMA grafting from electrospun scaffolds with 9% and 17% (w/w) PCL‐Ini before (left) and after (right) grafting of pOEGMA. Conversion by 1 H‐NMR (X) is inset. B) Water contact angle measurement of (i) electrospun PCL/PCL‐Ini and (ii) PCL‐pOEGMA with inset schematics. C) Antifouling ability of PCL and PCL‐pOEGMA electrospun scaffolds was compared using fluorescently labelled proteins and GAGs, expressed as μg cm −2 of electrospun scaffold, ** p < 0. 005. D) PCL (i) and PCL‐p(OEGMA‐ co ‐biotin) (ii) fibers labeled with fluorescein‐streptavidin and imaged using confocal microscopy. To visualize the polymer brush and to demonstrate the versatility of the system we included a biotinylated monomer unit (biotinylated PEG monomer 3, Supporting Information) that could be co‐polymerized into the bottlebrush. The resultant random co‐polymer of PCL‐pOEGMA‐ co ‐biotinylated PEG (PCL‐p(OEGMA‐ co ‐biotin)) allows labeling with streptavidin‐conjugated probes. Following histological sectioning of the scaffolds, labelling with fluorescein‐streptavidin, and imaging with confocal microscopy, the fluorescent signal was clearly visualized on the electrospun fibers demonstrating that the polymer brush is grafted from the fiber surface (Figure 3 D). Histological sections of the scaffolds imaged with wide field fluorescent microscopy further demonstrated the polymer brush evenly distributed throughout the cross section of the scaffold (Figure S3, Supporting Information) and further imaging of electrospun scaffolds following grafting of PCL‐p(OEGMA‐ co ‐biotin) confirmed the covalent attachment of the pOEGMA to the fiber surfaces. A control scaffold of electrospun PCL without the initiating group was present in the same reaction vessel as PCL‐Ini (17% w/w) scaffolds and demonstrated no fluorescent signal following washing, and incubation with fluorescein‐streptavidin. The successful surface grafting of pOEGMA is further supported by the difference observed in water contact angle between electrospun scaffolds with and without initiating groups that underwent polymerization. The control scaffolds remained hydrophobic while the pOEGMA functionalized scaffolds were highly hydrophilic (Figure S4, Supporting Information). 2. 3 pOEGMA Surface Functionalization for Antifouling and Prevention of Cell Adhesion pOEGMA is known to have antifouling properties and we looked to demonstrate this property from the functionalized scaffold. 30 17% (w/w) PCL‐pOEGMA scaffolds having achieved a conversion of >72% within the same reaction vessel were tested for antifouling and resistance to cell adhesion. pOEGMA grafted scaffolds demonstrated excellent antifouling ability and resistance to common ECM protein absorption when compared to PCL scaffolds. This was established using a panel of fluorescently labelled proteins and GAGs. These biomolecules were chosen as they represent a range of molecular weights, charge, and hydrophilicity, and several are known to modulate the binding and activity of other biomolecules such as growth factors. Adsorption of these molecules to a surface would likely lead to increased biomolecule deposition and ultimately, cell adhesion. pOEGMA scaffolds dramatically outperformed native PCL for both bovine serum albumin (BSA) and fibronectin showing a 3. 6 fold decrease in binding for BSA and greater than ten fold decrease for fibronectin (Figure 3 C). Relative to BSA, adsorption of hyaluronic acid (HA), heparin, and chondroitin sulphate (CS) for both PCL and pOEGMA scaffolds was negligible. In preparation for the creation of the dual functionality scaffold we prepared a cell adhesive PCL to compare to the PCL‐pOEGMA. We selected the canonical peptide motif RGDS sequence as a model cell‐adhesive biomolecule that promotes cell adhesion through integrin binding. 31 Fibroblasts, of which tenocytes are a specialized form, are known to bind to RGDS. 32 PCL was conjugated to a cyclized form of the RGDS (cRGDS), the natural presentation of the ligand in fibronectin. 33 PCL was conjugated with the cRGDS (Supporting Information) and electrospun into a scaffold using the standardized protocol. Bovine tenocytes were seeded onto both the PCL‐cRGDS and PCL‐pOEGMA scaffolds and cultured for 7 d to assess how the different surfaces supported cell adhesion and survival. Tenocytes seeded on the PCL‐cRGDS scaffolds adhered well, formed a confluent cell layer, and exhibited spread morphology as demonstrated by confocal imaging of the scaffold surface. Conversely, the cells seeded on the PCL‐pOEGMA scaffolds were few in number and those found were more rounded in morphology, indicating poor attachment ( Figure 4 A). The lack of robust attachment to the scaffold, as indicated by the rounded morphology, may have resulted in the surface cells being washed away while the remaining cells were trapped within the fibrous structure. These observations were compared for the whole scaffolds using a colorimetric assay for cellular metabolic activity based on the reduction of the tetrazolium dye 3‐(4, 5‐dimethylthiazol‐2‐yl)‐2, 5‐diphenyltetrazolium bromide (MTT) to approximate the relative number of cells (Supporting Information). A large reduction (p < 0. 0001) in overall metabolic activity was observed at day 7 between PCL‐cRGDS and PCL‐pOEGMA scaffolds implying a reduced cell number on the PCL‐pOEGMA scaffolds (Figure 4 B). The estimated cell numbers are somewhat higher than the appearance of the scaffolds by confocal microscopy would suggest. This reflects the presence of a small number of rounded cellular aggregates on the PCL‐pOEGMA surface, indicative of preferential cell–cell interactions over cell–surface interactions, in contrast to the densely populated spread cell morphology seen on the PCL‐cRGDS surface. Together, the estimated cell numbers and confocal microscopy findings show consistently different cellular adhesion between the PCL‐pOEGMA and PCL‐cRGDS surfaces. This is preserved in the bi‐functional scaffold, as evidenced by fluorescence microscopy, in accordance with our design ( Figure 5 C, D). Figure 4 Cell‐adhesive and non‐cell‐adhesive properties of functionalized electrospun scaffolds. A) Representative confocal microscopy images of bovine tenocytes cultured for 7 d on electropun PCL‐cRGDS (i) and PCL‐pOEGMA scaffolds (ii). Cell nuclei stained with draq5 (purple) and actin with phalloidin (green). B) Metabolic activity of bovine tenocytes cultured on scaffolds for 7 d was assessed by MTT assay. Estimated cell number is stated for each bar. *** Significant difference ( p < 0. 0001), error bars represent standard deviation. Figure 5 Dual functionality scaffolds demonstrated by fluorescent labelling of functionalities and cell adhesion. Fluorescence microscopy images of cross sections of bi‐functional scaffolds formed with opposing PCL‐Ini and PCL‐cRGDS surfaces. Post‐processing polymerization was used to produce a PCL‐p(OEGMA‐ co ‐biotin) surface. A, B) Overlaid fluorescence images of histological cross‐sections labelled with fluorescein‐streptavidin (green) on the p(OEGMA‐ co ‐biotin) and with Cy5 (red) for the PCL‐cRGDS showing well defined spatial resolution. Insets (left) show the brightfield and single channel fluorescence images with 100 μm scale bars. C, D) Bovine tenocytes were seeded on fresh scaffolds and cultured for 7 d before being stained with DAPI (blue) for cell nuclei and imaged with fluorescent microscopy. The PCL‐pOEGMA surface is shown in the upper image (C) and the opposing PCL‐cRGDS surface is shown in the lower image (D). 2. 4 Spatial Control of Polymer Brush Leading to a Dual Functionality Scaffold Having demonstrated the antifouling property of the PCL‐pOEGMA surface and cell‐adhesive property of the PCL‐cRGDS surface, we progressed to immobilizing them within a single construct. We sequentially electrospun the two functionalized polymers, PCL‐Ini and PCL‐cRGDS, to produce discrete sections within the same electrospun construct. To confirm the presence and spatial location of the functionalized PCL, we used specific fluorescent labels to tag corresponding functional groups within the pOEGMA or cRGDS sections. The PCL‐pOEGMA side included the biotinylated monomer previously described, to produce a PCL‐p(OEGMA‐ co ‐biotin) brush that could be labeled with streptavidin‐fluorescein. To visualize the cRGDS, we modified a Cy5 dye with an amine, which can react with the carboxylic acids found only on the aspartic acid (D) residue side chains exposed on the PCL‐cRGDS fibers (Figure S13, Supporting Information). Histological sections of the dual functional scaffolds were labeled to illustrate these discrete fiber locations and show successful spatial control of the different functionalities within the substance of a single electrospun construct (Figure 5 A, B). To further interrogate the dual functionality scaffolds, bovine tenocytes were seeded on both surfaces of the scaffolds and cultured for 7 d. The scaffolds were then stained with 4′, 6‐diamidino‐2‐phenylindole (DAPI) to label the cell nuclei on both sides of each scaffold and fluorescent imaging was performed by imaging one side, before turning it over and imaging the opposite side. The desirable functionality of the opposing surfaces is preserved despite the bilayer processing. The PCL‐cRGDS surface supports a high cell density whereas very few cells are seen on the opposing pOEGMA surface (Figure 5 C, D). Excellent spatial control of functionalized polymers is demonstrated by the sharp transition of layers seen with fluorescent labelling. Sequential electrospinning of end‐group functionalized polymers is shown to be a highly effective method for spatial control within a construct. This could be used to generate more complex architecture including multilayering. 3 Conclusion In conclusion, an effective dual functionality scaffold with a cell adhesive surface and an opposing antifouling, non‐cell adhesive surface has been successfully produced using a combination of end functionalization of PCL with a polymerization initiating group (BiBB) and a cell binding motif (cRGDS) that may be sequentially electrospun to produce a scaffold with zonally arranged functional groups. Post‐processing with ARGET ATRP is a facile and highly versatile method to produce a surface‐initiated pOEGMA bottlebrush that elicits a high performance antifouling and cell resistant coating. We have demonstrated the versatility of the polymerization for the addition of multifunctionality through the use of a biotinylated monomer unit in a single processing step. This modular and versatile approach could be used as a platform scaffold for multiple applications in tissue engineering. 4 Experimental Section All chemical reagents were purchased from Sigma Aldrich (UK) and deuterated solvents for 1 H‐NMR from Merck (Darmstadt, Germany) unless specifically noted. ARGET ATRP of OEGMA in Solution and from Surfaces : pOEGMA bottlebrushes were produced in solution and from 2D silicon and 3D electrospun functionalized surfaces. When producing brushes from surfaces, a sacrificial initiator, ethyl 2‐bromoisobutyrate (EBiB), was used in solution to enhance control of the polymerization and allow for surrogate characterization of the polymerization on the surface. 14, 34 The amount of initiator present on the 2D silicon and 3D electrospun surfaces was estimated and if greater than 10% of the sacrificial initiator mass, the mass of sacrificial initiator was reduced proportionally to maintain the reaction conditions. In a typical experiment OEGMA (340 mg, 0. 7 mmol, with a M w of the poly(ethylene glycol unit) of 400 Da, Polysciences, Germany), activated by the removal of inhibitors, was introduced into a round bottom flask with copper (II) chloride (Cu(II)Cl 2, 0. 64 mg, 0. 0047 mmol), tris[(2‐pyridyl)methyl]amine (TPMA, 1. 37 mg, 0. 0047 mmol), and EBiB (0. 92 mg, 0. 0047 mmol) in 3 mL of 50:50 isopropyl alcohol (IPA)/H 2 O. Following through mixing 2 mL was transferred to a test tube that was empty or containing functionalized 2D silicon or 3D electrospun surfaces. It was sealed, introduced into an ice bath, and degassed with argon for 20 min before ascorbic acid (AscA 0. 08 mg, 0. 00047 mmol) was added from a degassed solution using a gas tight syringe. The vessel was transferred to a pre‐warmed heat block at 30 °C and allowed to react for 2 h at which time the reaction was quenched by bubbling oxygen through the reaction mixture. Conversion was calculated by 1 H‐NMR (400 Hz, d 4‐MeOD) and molecular weights determined by SEC using a GPCMax VE 2001 (Viscotek). The SEC was run with an eluent of N, N ‐dimethylformamide (DMF) with 0. 075% (w/v) at a flow rate of 0. 7 mL min −1 over two Polymer Standards Service (PSS) Gram DMF columns at 35 °C. Molecular weights were determined relative to pMMA standards (Agilent Technologies, UK) without correction. Prior to the measurements the copper was removed using heavy metal chelating beads (Cuprisorb, Fish Fish Fish, UK) and filtered through a 0. 22 μm syringe mounted polytetrafluoroethylene filter. Scaffolds and silicon wafers functionalized with polymer brushes were removed from solutions and thoroughly rinsed three times in 100% ethanol before washing in Milli‐Q H 2 O for 24 h with three intermittent water changes. Any remaining copper was removed through the addition of Cuprisorb heavy metal chelating beads in the final washing step before drying in a vacuum desiccator at room temperature. Functionalization of Silicon Wafers with APTES and BiBB for 2D Silicon Polymerization : P‐doped silicon wafers (University Wafer, Boston, USA) were prepared for functionalization with sonication in acetone for 3 min, rinsing with Milli‐Q H 2 O and immersion in piranha solution (1:3, hydrogen peroxide: concentrated sulfuric acid) for 1 h, rinsing twice with Milli‐Q H 2 O, and drying under a stream of nitrogen. APTES was deposited on the surface of the silicon wafers using vapor deposition using a modified protocol from the literature; cleaned silicon wafers were laid in a glass petri dish and a glass vial containing 10 mL of anhydrous hexane and 0. 25 mL of APTES was placed in the center of the dish. 35 The petri dish was placed in a desiccator to which a vacuum was applied and maintained for 90 min. Wafers were subsequently removed and inserted into test tubes, sealed with septa and parafilm, degassed with nitrogen and to each a degassed solution of 5 mL of anhydrous hexane, 0. 125 mL of BiBB, and 0. 165 mL of anhydrous triethylamine (TEA) was introduced and allowed to react at room temperature for 1 h. Following this the wafers were removed and washed with hexane, ethanol and Milli‐Q H 2 O and then dried under a stream of nitrogen. Wafers were stored in a vacuum desiccator and protected from light until used. Water Contact Angle Measurement : Water contact angles on 2D silicon wafer surfaces and electrospun scaffolds were measured with a Kr˝uss Easy Drop DSA 100 (Hamburg, Germany) and the associated DSA1 v 1. 9 software, a drop size of 7 μL, and at room temperature. Brush Thickness Measurements with AFM and Spectroscopic Ellipsometry : Dry thickness measurements were undertaken with AFM and ellipsometry on the same 2D silicon samples to establish the thickness of the grafted pOEGMA layer. AFM measurements were taken using an Agilent Technologies 5500 atomic force microscope with a silicon probe, tip radius <10 nm, force constant 40 nN m −1. A scratch test was performed whereby the AFM probe tip was moved towards the sample, contact was made, and the force increased until the underlying silicon wafer was contacted. The tip was then moved laterally to create a full‐thickness scratch. The AFM was then used in tapping mode to create a depth profile across the scratch area. Using the associated software (Pico Image) thickness measurements of the polymer layer were made. Ellipsometry measurements were performed at room temperature using a SOPRA GESP 5 variable angle spectroscopic ellipsometer. The data were recorded through incidence angles of 65°–75° with respect to the substrate normal, across a wavelength range from 900–1600 nm (20 nm steps). Cauchy model fits were used to analyze the ellipsometric data. Good agreement was found between the ellipsometrically deduced brush thicknesses and AFM. Production of Functionalized Electrospun Scaffolds and Fiber‐Initiated Controlled Radical Polymerization–Synthesis of PCL‐Ini 1 : PCL‐diol ( M w 14 000 Da) was functionalized with 2‐bromoisobutyryl bromide (BiBB) using a protocol adapted from the literature to produce PCL‐Ini 1. 24 In a typical run, 5 g of 14 kDa PCL‐diol was introduced to 200 mL of anhydrous tetrahydrofuran in a round bottom flask and stirred at room temperature in a sealed nitrogen atmosphere. Following dissolution of the PCL, 2. 9 mL TEA and 50 mg of 4‐dimethylaminopyridine were added and the vessel immersed in an ice bath. After 15 min of cooling, 265 μL of BiBB was introduced dropwise, the ice bath removed, and the reaction stirred overnight at room temperature. The mixture was filtered, and the filtrate was collected, reduced through vacuum rotary evaporation, and precipitated into cold diethyl ether. The precipitate was isolated by vacuum filtration and dried. The product was confirmed using 1 H‐NMR (400 MHz, deuterated CDCl 3 ) and was stored until needed in a vacuum desiccator, protected from light. 1 H NMR (400 MHz, CDCl 3 ) δ ppm: 4. 24 − 4. 20 (m, 4H) 4. 05 ( t, J = 6. 7 Hz, 240H), 3. 78 − 3. 71 (m, 4H), 2. 30 ( t, J = 7. 5 Hz, 244H), 1. 92 (s, 0H), 1. 73 − 1. 55 (m, 480H), 1. 46 − 1. 32 (m, 252H) (see Figure S5, Supporting Information, for peak assignments). Electrospinning PCL and PCL Conjugates : Two different PCL‐Ini:PCL ratios were electrospun into scaffolds; 12 mg mL −1 (9% w/w) PCL‐Ini or 24 mg mL −1 (17% w/w) PCL‐Ini was added to 12% (w/v) PCL ( M n 70 000–90 000 Da) in 1, 1, 1, 3, 3, 3‐hexafluoro‐2‐propanol (HFIP) and mixed overnight. Solutions were transferred to plastic syringes, loaded onto a programmable syringe pump (Kd Scientific, UK) and extruded at a rate of 2 mL h −1 through a blunt 18‐gauge stainless steel needle charged with +16 kV (Glassman, Bramley, Hampshire, UK). The needle was placed at distance of 11 cm from a grounded 10 × 10 cm plate for small master scaffolds or a mandrel rotating at 0. 33 m s −1 for large master scaffolds, each coated with aluminum foil. No difference was seen in fiber alignment or morphology between collectors. All scaffolds were electrospun under the same conditions, stored in a vacuum desiccator, and protected from light until needed. XPS Analysis of the 2D Silicon and 3D Electrospun Polymer Brush Functionalized Surfaces : XPS was used to characterize the surface of both the 2D silicon and 3D electrospun samples. The spectra were measured using a Thermo Fisher K‐Alpha XPS System (Thermo Fisher Scientific Inc. ) with a monochromatic Al‐Kα (energy = 1486. 71 eV) X‐ray source. Samples were positioned at the electron take‐off angle normal to the surface with respect to the analyzer. Survey spectra were measured over a range of 0–1400 eV and recorded for each sample, then followed by high resolution spectra for C1s and O1s. A low energy electron/ion flood gun was used to ensure effective surface charge compensation. XPS spectra were calibrated to the adventitious C1s signal (285. 0 eV). Curve fitting was carried out using Thermo Avantage Software (v. 5. 948) using a Shirley background. Peak areas were normalized within Thermo Avantage using atomic sensitivity factors for the Al Kα anode (“AlWagner” library) 36 and from these areas the carbon composition and elemental ratios were determined. Modification of the Polymerization Protocol for Use with the Biotinoylated PEG Monomer 3 : For samples requiring fluorescent labeling, minor modifications were needed to mitigate any possible chelation of the copper catalyst by additional groups in the reaction mixture. The reaction was prepared as previously outlined but with 5 mol% of the OEGMA replaced with biotinylated PEG monomer 3 (Supporting Information), and the reaction was left to proceed overnight. Samples were washed and prepared as previously described. Histological Sectioning, Labeling, and Fluorescent Imaging of the Scaffolds : Scaffolds for histological section and analysis were embedded in polyester wax (VWR, UK) in a method modified from Steedman et al. 37 Scaffolds were embedded following incubation in a series of wax solutions maintained at 42 °C, first 30 min in 1:1 (v/v) polyester wax and ethanol, followed by two rounds of pure polyester wax for 30 min and 1 h, respectively. The scaffolds were then embedded vertically in polyester wax and allowed to set. Sections were cut at 10 μm onto untreated glass slides, dried, dewaxed with ethanol, and affixed at either end with a drop of inert adhesive. Sections were blocked with 1% (w/w) BSA and 0. 1% (w/v) tween 30 in phosphate buffered saline (PBS) for 30 min, stained for 15 min with fluorescein‐streptavidin (Vector Labs, UK) diluted to 1 μg mL −1 in 1% (w/w) BSA in PBS at pH 8. 4, and washed three times in PBS. Slides were coverslipped for confocal microscopy with FluorSave fluorescent mounting media (Millipore, UK). Standard fluorescent imaging was performed on a Leica inverted optical microscope fitted with an Olympus DP70 digital camera. Confocal imaging was performed on a Leica SP5 inverted confocal microscope and images processed using GIMP 2. 1. Measurement of Protein Adsorption onto PCL and pOEGMA Scaffolds : Fluorescently labeled rhodamine‐heparin (rhod‐hep, M W 18 kDa), rhodamine‐CS (rhod‐CS, M W 50 kDa), and fluorescein‐HA (fluor‐HA, M W 1500 kDa) were purchased (Creative PEGWorks, Winston Salem, USA). Rhodamine‐fibronectin (rhod‐fib) and rhotamine‐BSA (rhod‐BSA) were synthesized prior to experimentation (Supporting Information). Stock solutions of labeled protein were 50 μg mL −1 in PBS for all proteins with the exception of rhod‐BSA which was 10 μg mL −1 in PBS. Circular PCL‐Ini scaffolds, 6 mm in diameter, with and without pOEGMA brush functionalization were immersed in 70% (v/v) ethanol and washed with PBS three times to ensure uniform hydration. Excess liquid was removed and the scaffolds introduced into high return 1. 5 mL centrifuge tubes for incubation with 200 μL protein solution (test samples) or PBS (control) at 37 °C for 18 h. Scaffolds were then washed in PBS in 28 mL light protected glass vials overnight to remove any unbound protein. The fluorescent signal in the scaffolds was then quantified on a Perkin Elmer Envision multimode detector (Germany) at an excitation wavelength of 550 nm and emission of 580 nm for rhodamine labeled proteins and at an excitation wavelength of 490 nm and emission of 520 nm for fluorescein labeled proteins. Cell Adhesion Testing of PCL‐cRGDS and PCL‐pOEGMA Scaffolds : Bovine tenocytes were isolated through primary cell culture from three independent animals all of which were a maximum of 2 years in age (Supporting Information). Cell adhesion experiments were performed using a protocol modified from the literature. 38 In brief, 24‐well plates were coated with two‐component silicon elastomer (Sygard 184, Dow Corning) prepared in a 10:1 ratio and cured for 48 h. Each independent experiment utilized seven PCL‐pOEGMA scaffolds of 6 mm diameter, that had previously been functionalized with a pOEGMA brush and seven 6 mm diameter PCL‐cRGDS scaffolds. Scaffolds and stainless steel insect pins (0. 15 mm, Watkins and Doncaster, UK) were immersed into 70% (v/v) ethanol for 15 min before washing in sterile PBS supplemented with 1% (w/v) anti/anti three times. The silicone‐coated well plate was thoroughly sprayed with 70% (v/v) ethanol before a scaffold was inserted into each well and fixed with an insect pin through the center of the scaffold. A row of empty wells was left as a control. The plate and scaffolds were sterilized under UV light in the cell culture hood for 8 h. Scaffolds were washed with sterile PBS immediately before cell seeding. Bovine tenocytes were prepared in a single cell suspension at a concentration of 5 × 10 5 cells mL −1 from which 50 μL (2. 5 × 10 4 cells) were seeded onto each scaffold. After allowing 2 h for cell attachment at 37 °C, 5% CO 2, and 100% relative humidity, a further 1 mL of normal growth media (NGM) was gently added to each well. Seeded scaffolds were cultured for 7 d, with the media replaced after the fourth day. On the seventh day the media was aspirated from the wells and the scaffolds were washed with sterile PBS. Two scaffolds of each type were prepared for imaging, the remaining five scaffolds of each type were used for quantification of cellular metabolism using an MTT assay. This was compared to a standard curve produced from a cell ladder to estimate cell number. Scaffolds reserved for imaging were fixed in 4% (w/v) paraformaldehyde (PFA) for 15 min, washed, and stored at 4 °C in PBS until used. Prior to imaging, scaffolds were blocked in 1% (w/v) BSA and 0. 1% (w/v) tween 20 in PBS for 30 min. A solution of 5 × 10 −6 m draq5 (Thermo Scientific, UK) to stain cell nuclei and phalloidin (Alexa Fluor 488, diluted 1:400, Life Technologies) to stain actin was diluted in 1% (w/v) BSA in PBS and was incubated with the scaffolds under light protection for 20 min before being washed with further PBS. Samples were imaged in PBS by inverted confocal microscopy as previously described. Production of Dual Functionality Scaffolds and Characterization : Dual functionality scaffolds were fabricated as a continuous scaffold using layered electrospinning. Two separate solutions of functionalized PCL were prepared in HFIP as above with the addition of 1 mg mL −1 PCL‐cRGDS to one, and 17% (w/w) PCL‐Ini to the other. These were sequentially electrospun in accordance with the protocol described above. The PCL‐cRGDS solution (2 mL) was electrospun first, followed by PCL‐Ini solution. This was allowed to run for a further 30 min to ensure coverage of the scaffold. Post‐processing polymerization of pOEGMA and pOEGMA‐ co ‐biotin was performed as described above. Sectioning, Labeling, and Imaging of Dual Functionality Scaffolds : Dual functionality scaffolds with a PCL‐p(OEGMA‐ co ‐biotin) surface were embedded in polyester wax and blocked out in a vertical orientation before sectioning at 10 μm onto glass slides, and dewaxed with ethanol as previously described. The scaffolds were immobilized onto glass coverslips with an inert adhesive (Aquarium Sealant, King British, Gainsbrough, UK) that was allowed to cure overnight on the bench. The scaffolds were then incubated in 0. 2% (w/v) tween 20/0. 2% (w/v) triton X in PBS for 60 min before being washed in Milli‐Q H 2 O and the excess blotted away with filter paper. To label the cRGDS moiety, an amino‐Cy5 dye 7 was synthesized (Supporting Information), diluted to 0. 1 × 10 −3 m in 20 × 10 −3 m sodium borate buffer solution at pH 9, and combined with 1‐ethyl‐3(3‐dimethylaminopropyl)carbodiimide (EDC, 2 × 10 −3 m ) and N ‐hydroxysuccinimide (NHS, 2 × 10 −3 m ). This solution was applied to the immobilized scaffolds and the reaction proceeded on the bench at room temperature for 30 min before the immobilized scaffolds were washed with 20 × 10 −3 m sodium borate buffer, Milli‐Q H 2 O, 0. 2% (w/v) tween 20/0. 2% (w/v) triton X solution, Milli‐Q H 2 O, 50% (v/v) IPA, 100% IPA, 50% (v/v) IPA, and further Milli‐Q H 2 O. The scaffolds were then blocked and labeled with fluorescein‐streptavidin as previously described, light protected, and imaged immediately. Sections were imaged on an upright Olympus BX51 epifluorescent microscope equipped with an Olympus DP70 camera. Images were obtained in three channels: Bright field, FITC to image the fluorescent brushes, and TxRed to image the Cy5 labeling of cRGDS. Cell Adhesion Assessment and Imaging of Dual Functionality Scaffolds : Cell adhesion was assessed on the dual functionality scaffolds using a minor variation to the above protocol. Scaffolds and cells were prepared in an identical manner. When seeded, 25 000 cells were seeded on one side of the scaffold, incubated for 30 min to allow for cell attachment, gently turned over with sterile forceps, and a further 25 000 cells were seeded on the opposing side. The scaffold was then immobilized in a silicone‐coated well plate ensuring that there was free space for media below the scaffold (between the scaffold and the silicon coating). The plate was then returned to the incubator for 2 h for further cell attachment before NGM was gently added as per protocol. After 7 d the dual functional scaffolds were fixed in 4% (w/v) PFA and blocked in 1% (w/v) BSA and 0. 1% (w/v) tween 20 as described above. 4′, 6‐diamidino‐2‐phenylindole (DAPI, diluted 1:5000, Sigma) and phalloidin (diluted 1:200, Alexa Fluor 568, Invitrogen) were diluted in 1% (w/v) BSA solution and incubated with the scaffolds for 15 min. After washing in PBS the scaffolds were incubated in fluorescein‐streptavidin (diluted 1:500, Vector Laboratories) in PBS at pH 8. 4 for a further 15 min. After washing in PBS the scaffolds were mounted on glass slides with Fluorosave mounting media (Calbiochem, VWR) and coverslipped. Slides were protected from light before being imaged on both the confocal and inverted optical microscopes as described above. Statistical Analysis : All experimental test groups had a sample size of at least n = 3 for biochemical analysis. All cell‐related work was repeated with bovine tenocytes from three different animals. Data are presented as mean +/− standard deviation (SD). Statistical significance was determined by students T‐tests using Excel software, with a significance accepted where p ‐value < 0. 05. Supporting information As a service to our authors and readers, this journal provides supporting information supplied by the authors. Such materials are peer reviewed and may be re‐organized for online delivery, but are not copy‐edited or typeset. Technical support issues arising from supporting information (other than missing files) should be addressed to the authors. Supplementary Click here for additional data file.
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10. 1002/adfm. 201504160
| 2,016
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Advanced functional materials
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Silk Biomaterials with Vascularization Capacity
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Functional vascularization is critical for the clinical regeneration of complex tissues such as kidney, liver or bone. The immobilization or delivery of growth factors has been explored to improve vascularization capacity of tissue engineered constructs, however, the use of growth factors has inherent problems such as the loss of signaling capability and the risk of complications such as immunological responses and cancer. Here, a new method of preparing water-insoluble silk protein scaffolds with vascularization capacity using an all aqueous process is reported. Acid was added temporally to tune the self-assembly of silk in lyophilization process, resulting in water insoluble scaffold formation directly. These biomaterials are mainly noncrystalline, offering improved cell proliferation than previously reported silk materials. These systems also have appropriate softer mechanical property that could provide physical cues to promote cell differentiation into endothelial cells, and enhance neovascularization and tissue ingrowth in vivo without the addition of growth factors. Therefore, silk-based degradable scaffolds represent an exciting biomaterial option, with vascularization capacity for soft tissue engineering and regenerative medicine.
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10. 1002/adfm. 201605352
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Advanced functional materials
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Gold Nanocomposite Bioink for Printing 3D Cardiac Constructs
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Bioprinting is the most convenient microfabrication method to create biomimetic three-dimensional (3D) cardiac tissue constructs, which can be used to regenerate damaged tissue and provide platforms for drug screening. However, existing bioinks, which are usually composed of polymeric biomaterials, are poorly conductive and delay efficient electrical coupling between adjacent cardiac cells. To solve this problem, we developed a gold nanorod (GNR) incorporated gelatin methacryloyl (GelMA)-based bioink for printing 3D functional cardiac tissue constructs. The GNR concentration was adjusted to create a proper microenvironment for the spreading and organization of cardiac cells. At optimized concentration of GNR, the nanocomposite bioink had a low viscosity, similar to pristine inks, which allowed for the easy integration of cells at high densities. As a result, rapid deposition of cell-laden fibers at a high resolution was possible, while reducing shear stress on the encapsulated cells. In the printed GNR constructs, cardiac cells showed improved cell adhesion and organization when compared to the constructs without GNRs. Furthermore, the incorporated GNRs bridged the electrically resistant pore walls of polymers, improved the cell-to-cell coupling, and promoted synchronized contraction of the bioprinted constructs. Given its advantageous properties, this gold nanocomposite bioink may find wide application in cardiac tissue engineering.
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No full text available
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10. 1002/adfm. 201606273
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Advanced functional materials
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Self-assembled Hydrogel Fiber Bundles from Oppositely Charged Polyelectrolytes Mimic Micro-/nanoscale Hierarchy of Collagen
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Fiber bundles are present in many tissues throughout the body. In most cases, collagen subunits spontaneously self-assemble into a fibrilar structure that provides ductility to bone and constitutes the basis of muscle contraction. Translating these natural architectural features into a biomimetic scaffold still remains a great challenge. Here, we propose a simple strategy to engineer biomimetic fiber bundles that replicate the self-assembly and hierarchy of natural collagen fibers. The electrostatic interaction of methacrylated gellan gum (MeGG) with a countercharged chitosan (CHT) polymer led to the complexation of the polyelectrolytes. When directed through a polydimethylsiloxane (PDMS) channel, the polyelectrolytes formed a hierarchical fibrous hydrogel demonstrating nano-scale periodic light/dark bands similar to D-periodic bands in native collagen and aligned parallel fibrils at micro-scale. Importantly, collagen-mimicking hydrogel fibers exhibited robust mechanical properties (MPa scale) at a single fiber bundle level and enabled encapsulation of cells inside the fibers under cell-friendly mild conditions. Presence of carboxyl- (in gellan gum) or amino- (in chitosan) functionalities further enabled controlled peptide functionalization such as RGD for biochemical mimicry (cell adhesion sites) of native collagen. This biomimetic aligned fibrous hydrogel system can potentially be used as a scaffold for tissue engineering as well as a drug/gene delivery vehicle.
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10. 1002/adfm. 201606614
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Advanced functional materials
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Elastomeric Fibrous Hybrid Scaffold Supports In Vitro and In Vivo Tissue Formation
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Biomimetic materials with biomechanical properties resembling those of native tissues while providing an environment for cell growth and tissue formation, are vital for tissue engineering (TE). Mechanical anisotropy is an important property of native cardiovascular tissues and directly influences tissue function. This study reports fabrication of anisotropic cell-seeded constructs while retaining control over the construct’s architecture and distribution of cells. Newly synthesized poly-4-hydroxybutyrate (P4HB) is fabricated with a dry spinning technique to create anelastomeric fibrous scaffold that allows control of fiber diameter, porosity, and rate ofdegradation. To allow cell and tissue ingrowth, hybrid scaffolds with mesenchymalstem cells (MSCs) encapsulated in a photocrosslinkable hydrogel were developed. Culturing the cellularized scaffolds in a cyclic stretch/flexure bioreactor resulted in tissue formation and confirmed the scaffold’s performance under mechanical stimulation. In vivo experiments showed that the hybrid scaffold is capable of withstanding physiological pressures when implanted as a patch in the pulmonary artery. Aligned tissue formation occurred on the scaffold luminal surface without macroscopic thrombus formation. This combination of a novel, anisotropic fibrous scaffold and a tunable native-like hydrogel for cellular encapsulation promoted formation of 3D tissue and provides a biologically functional composite scaffold for soft-tissue engineering applications.
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10. 1002/adfm. 201701183
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Advanced functional materials
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Engineered Axonal Tracts as “Living Electrodes” for Synaptic-Based Modulation of Neural Circuitry
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Brain-computer interface and neuromodulation strategies relying on penetrating non-organic electrodes/optrodes are limited by an inflammatory foreign body response that ultimately diminishes performance. A novel “biohybrid” strategy is advanced, whereby living neurons, biomaterials, and microelectrode/optical technology are used together to provide a biologically-based vehicle to probe and modulate nervous-system activity. Microtissue engineering techniques are employed to create axon-based “living electrodes”, which are columnar microstructures comprised of neuronal population(s) projecting long axonal tracts within the lumen of a hydrogel designed to chaperone delivery into the brain. Upon microinjection, the axonal segment penetrates to prescribed depth for synaptic integration with local host neurons, with the perikaryal segment remaining externalized below conforming electrical-optical arrays. In this paradigm, only the biological component ultimately remains in the brain, potentially attenuating a chronic foreign-body response. Axon-based living electrodes are constructed using multiple neuronal subtypes, each with differential capacity to stimulate, inhibit, and/or modulate neural circuitry based on specificity uniquely afforded by synaptic integration, yet ultimately computer controlled by optical/electrical components on the brain surface. Current efforts are assessing the efficacy of this biohybrid interface for targeted, synaptic-based neuromodulation, and the specificity, spatial density and long-term fidelity versus conventional microelectronic or optical substrates alone.
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10. 1002/adfm. 201701713
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Advanced functional materials
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Tissue-Engineered Peripheral Nerve Interfaces
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Research on neural interfaces has historically concentrated on development of systems for the brain; however, there is increasing interest in peripheral nerve interfaces (PNIs) that could provide benefit when peripheral nerve function is compromised, such as for amputees. Efforts focus on designing scalable and high-performance sensory and motor peripheral nervous system interfaces. Current PNIs face several design challenges such as undersampling of signals from the thousands of axons, nerve-fiber selectivity, and device–tissue integration. To improve PNIs, several researchers have turned to tissue engineering. Peripheral nerve tissue engineering has focused on designing regeneration scaffolds that mimic normal nerve extracellular matrix composition, provide advanced microarchitecture to stimulate cell migration, and have mechanical properties like the native nerve. By combining PNIs with tissue engineering, the goal is to promote natural axon regeneration into the devices to facilitate close contact with electrodes; in contrast, traditional PNIs rely on insertion or placement of electrodes into or around existing nerves, or do not utilize materials to actively facilitate axon regeneration. This review presents the state-of-the-art of PNIs and nerve tissue engineering, highlights recent approaches to combine neural-interface technology and tissue engineering, and addresses the remaining challenges with foreign-body response.
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10. 1002/adfm. 201705563
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Advanced functional materials
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Integration of Phase-Change Materials with Electrospun Fibers for Promoting Neurite Outgrowth under Controlled Release
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We report a temperature-regulated system for the controlled release of nerve growth factor (NGF) to promote neurite outgrowth. The system is based upon microparticles fabricated using coaxial electrospray, with the outer solution containing a phase-change material (PCM) and the inner solution encompassing payload(s). When the temperature is kept below the melting point of the PCM, there is no release due to the extremely slow diffusion through a solid matrix. Upon increasing the temperature to slightly pass the melting point, the encapsulated payload(s) can be readily released from the melted PCM. By leveraging the reversibility of the phase transition, the payload(s) can be released in a pulsatile mode through on/off heating cycles. The controlled release system is evaluated for potential use in neural tissue engineering by sandwiching the microparticles, co-loaded with NGF and a near-infrared dye, between two layers of electrospun fibers to form a tri-layer construct. Upon photothermal heating with a near-infrared laser, the NGF is released with well-preserved bioactivity to promote neurite outgrowth. By choosing different combinations of PCM, biological effector, and scaffolding material, this controlled release system can be applied to a wide variety of biomedical applications.
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10. 1002/adfm. 201706046
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Advanced functional materials
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Bioorthogonal Strategies for Engineering Extracellular Matrices
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Hydrogels are commonly used as engineered extracellular matrix (ECM) mimics in applications ranging from tissue engineering to in vitro disease models. Ideal mechanisms used to crosslink ECM-mimicking hydrogels do not interfere with the biology of the system. However, most common hydrogel crosslinking chemistries exhibit some form of cross-reactivity. The field of bio-orthogonal chemistry has arisen to address the need for highly specific and robust reactions in biological contexts. Accordingly, bio-orthogonal crosslinking strategies have been incorporated into hydrogel design, allowing for gentle and efficient encapsulation of cells in various hydrogel materials. Furthermore, the selective nature of bio-orthogonal chemistries can permit dynamic modification of hydrogel materials in the presence of live cells and other biomolecules to alter matrix mechanical properties and biochemistry on demand. In this review, we provide an overview of bio-orthogonal strategies used to prepare cell-encapsulating hydrogels and highlight the potential applications of bio-orthogonal chemistries in the design of dynamic engineered ECMs.
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10. 1002/adfm. 201707107
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Advanced functional materials
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Functionally Graded, Bone- and Tendon-Like Polyurethane for Rotator Cuff Repair
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Critical considerations in engineering biomaterials for rotator cuff repair include bone-tendon-like mechanical properties to support physiological loading and biophysicochemical attributes that stabilize the repair site over the long-term. In this study, UV-crosslinkable polyurethane based on quadrol (Q), hexamethylene diisocyante (H), and methacrylic anhydride (M; QHM polymers), which are free of solvent, catalyst, and photoinitiator, is developed. Mechanical characterization studies demonstrate that QHM polymers possesses phototunable bone- and tendon-like tensile and compressive properties (12–74 MPa tensile strength, 0. 6–2. 7 GPa tensile modulus, 58–121 MPa compressive strength, and 1. 5–3. 0 GPa compressive modulus), including the capability to withstand 10 000 cycles of physiological tensile loading and reduce stress concentrations via stiffness gradients. Biophysicochemical studies demonstrate that QHM polymers have clinically favorable attributes vital to rotator cuff repair stability, including slow degradation profiles (5–30% mass loss after 8 weeks) with little-to-no cytotoxicity in vitro, exceptional suture retention ex vivo (2. 79–3. 56-fold less suture migration relative to a clinically available graft), and competent tensile properties (similar ultimate load but higher normalized tensile stiffness relative to a clinically available graft) as well as good biocompatibility for augmenting rat supraspinatus tendon repair in vivo. This work demonstrates functionally graded, bone-tendon-like biomaterials for interfacial tissue engineering.
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10. 1002/adfm. 201800618
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Advanced functional materials
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Auxetic Cardiac Patches with Tunable Mechanical and Conductive Properties toward Treating Myocardial Infarction
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An auxetic conductive cardiac patch (AuxCP) for the treatment of myocardial infarction (MI) is introduced. The auxetic design gives the patch a negative Poisson’s ratio, providing it with the ability to conform to the demanding mechanics of the heart. The conductivity allows the patch to interface with electroresponsive tissues such as the heart. Excimer laser microablation is used to micropattern a re-entrant honeycomb (bow-tie) design into a chitosan-polyaniline composite. It is shown that the bow-tie design can produce patches with a wide range in mechanical strength and anisotropy, which can be tuned to match native heart tissue. Further, the auxetic patches are conductive and cytocompatible with murine neonatal cardiomyocytes in vitro. Ex vivo studies demonstrate that the auxetic patches have no detrimental effect on the electrophysiology of both healthy and MI rat hearts and conform better to native heart movements than unpatterned patches of the same material. Finally, the AuxCP applied in a rat MI model results in no detrimental effect on cardiac function and negligible fibrotic response after two weeks in vivo. This approach represents a versatile and robust platform for cardiac biomaterial design and could therefore lead to a promising treatment for MI.
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1 Introduction Cardiovascular diseases (CVDs) are the leading causes of death and disability worldwide. [ 1 ] Biomaterials and regenerative therapies represent exciting possible solutions. [ 2 – 8 ] Herein, we investigate a new biomaterial design for cardiac patches to treat myocardial infarction (MI), one of the largest contributors to CVDs. MI are caused by an occlusion of one or more of the coronary arteries, resulting in the myocardial tissue becoming ischemic and reducing the heart’s ability to pump blood around the body. [ 9 ] Cardiac patches seek to strengthen the tissue, supply cells or growth factors to vitalize the tissue, and bridge electrical and/or mechanical stimulation across the infarct to maintain and improve cardiac function. For a cardiac patch to be successful in triggering regeneration of the myocardium, the biomaterial design will benefit from considering and optimizing the cytocompatibility, electrical conductivity, and mechanical properties. [ 2 – 8 ] It is widely accepted that the mechanical properties of biomaterials used for treating MI are extremely important. [ 2 – 8, 10 ] Typically, it is believed that the mechanical properties of a cardiac patch should match those of healthy native heart tissue. However, the Young’s modulus of the native human heart varies from 0. 02 to 0. 50 MPa depending on whether the heart is in systole or diastole, with infarct tissue being even stiffer. [ 7, 8 ] This range is broad and the optimum mechanical properties for such an application are debatable. [ 7, 8 ] Further research into the regeneration of infarcted cardiac tissue using cardiac patches with controllable mechanical properties will be instrumental in the development of future strategies for treating MI. [ 7, 8 ] Another important mechanical property which is often overlooked is the Poisson’s ratio. The Poisson’s ratio ( ν ) describes the effect on transversal strain ( ε T ) when a material is under longitudinal tension or compression, creating longitudinal strain ( ε L ), and is defined by Equation (1) (1) ν = − ε T ε L Most materials typically have a positive Poisson’s ratio and as a result, when stretched longitudinally, they contract transversally. However, auxetic materials are defined as having a negative Poisson’s ratio and therefore expand in multiple directions simultaneously. This results in the improvement of other unique properties, such as shear resistance, indentation resistance, and synclastic curvature, all of which are interesting properties in a cardiac patch. [ 11 – 13 ] Auxetic micropatterning is presented here as a unique way of incorporating these properties into cardiac patches. There has been very little investigation into the use of auxetic materials or the importance of the Poisson’s ratio in the field of biomaterials, particularly for treating MI. [ 14 – 17 ] In addition to mechanical strength, electrical signals are vital to the heart’s contractions and function. Due to their inherent electroactive nature, conductive materials are of interest in biomaterial design for electroresponsive tissues such as the heart and the central nervous system. Doped-conjugated polymers could be particularly useful in cardiac patches due to their potential ability to improve the electrical pathway which is damaged in the infarct tissue after an MI. [ 18 – 23 ] Electroactive biomaterials were recently shown to favorably influence the electrophysiology of infarcted animal cardiac tissue ex vivo[ 24 ] and in vivo. [ 19 ] In contrast, the existing insulating biomaterials have been shown to further hinder signal propagation through the infarct tissue. [ 25 ] Previously, our group developed a conductive cardiac patch with exceptional electrical stability, which was shown to increase the conduction velocity (CV) across the damaged electrical pathway of the infarct region in a rat MI model. [ 24 ] However, the Young’s modulus of this biomaterial is significantly greater (6. 73 ± 1. 1 MPa)[ 24 ] than that reported for native human heart tissues (0. 02–0. 50 MPa). [ 7, 8 ] Building on our previous work, [ 24 ] here we incorporate mechanical and topological anisotropy in an auxetic design to develop a conductive cardiac patch that can better comply mechanically with the heart. We introduce the ability to control and tune the effective stiffness and anisotropy of auxetic, conductive cardiac patches and we investigate their impact on cardiac function. 2 Results and Discussion 2. 1 Design and Fabrication of Auxetic Cardiac Patches There are many ways in which auxetic behavior can be imparted into a material. [ 13 ] The design used here is known as the re-entrant honeycomb or “bow-tie” deformation mechanism ( Figure 1A ). [ 13, 26 ] A particularly interesting feature of this bowtie geometry is that it is anisotropic, that is, it is stiffer in one direction than the other, which is similar to native heart tissue. The anisotropic ratio of effective stiffness of native hearts has been reported over a wide range of values (1. 9–3. 9), [ 27 – 30 ] depending on species, age, and health. However, it is consistently reported that the heart is stiffer in the circumferential (transverse) direction than in the longitudinal direction. [ 27 – 30 ] It is assumed that cardiac biomaterials should match native anisotropic mechanical properties to not impede heart function. [ 27 – 30 ] In this study, the stiffer direction (1-direction) of the auxetic cardiac patch (AuxCP) is aligned with the stiffer circumferential (transverse) direction of the heart, while the less stiff AuxCP direction (2-direction) is aligned with the less stiff longitudinal direction of the heart ( Figure 1B ). When this bow-tie pattern is stretched in one direction, the diagonal ribs flex and open causing expansion in the other direction, resulting in a negative Poisson’s ratio ( Figure 1C, D ). The Poisson’s ratio is negative within the first 5–15% strain (depending on original bow-tie dimensions), a representative example is illustrated in Figure S1 ( Supporting Information ). The material used is a conductive composite previously developed by our group. [ 24 ] Briefly, this composite consists of an interconnected network of polyaniline and phytic acid grown on a chitosan surface. Ammonium persulfate is used to trigger the polymerization of the aniline in situ, while the phytic acid cross-links both the chitosan and polyaniline. The phytic acid also has a second role as the dopant, resulting in a highly stable, electrically conductive thin film with the rich dark green color associated to the emeraldine acid oxidation state of polyaniline. The fabrication of the bow-tie geometry was achieved through excimer laser microablation in the chitosan films. The process can be easily modified to create patches with varying bow-tie geometries and sizes ( Figure 1A ). After micropatterning the chitosan, the patches are coated with polyaniline and their precision is maintained to within 5% of the original microablated pattern dimensions (Figure S2, Supporting Information ). Unlike other micropatterning methods such as photolithography, this technique is not limited to photocurable materials, [ 7 ] opening the door to the micropatterning of a wide array of biomaterials, particularly conductive biomaterials. In addition, the fabrication process was optimized to produce 40 samples (6 × 6 bow-tie repeat units) in under an hour while maintaining the precision of the patterning, producing cardiac patches considerably faster than can be achieved with many other rapid prototyping methods, such as common 3D printing techniques. [ 17 ] Therefore, this technique has great potential for mass production and distribution. Moreover, this is an extremely robust and versatile technique for 2D microfabrication of biomaterials that allows tissue specific tailoring. 2. 2 Characterization of Tunable Mechanical Properties Theoretical models based on the 2D bow-tie pattern, previously established in the literature, [ 26 ] were used to predict the mechanical properties of the micropatterned patches. The equations use the Young’s modulus of the bulk material and the various bow-tie dimensions ( Figure 1A ) to calculate the resultant effective stiffness ( E ) values and the anisotropic ratio of effective stiffness ( E 1 / E 2 ). Tensile tests proved that varying the bow-tie dimensions A, B, θ, and R, can tune E 1, E 2, and E 1 /E 2 as predicted by the theoretical model[ 26 ] ( Figure 2A–H ). A full table of information about all nine different sample sets can be found in Table S1 ( Supporting Information ), including all bow-tie dimensions and measured mechanical properties. (Representative stress– strain curves produced from these tensile tests are also shown in Figure S3, Supporting Information. ) Specifically, by increasing the length of dimension A from 320 to 480 μm, E 1 / E 2 decreases from 5. 71 to 1. 99 ( Figure 2A, B ). By increasing the length of dimension B from 240 to 360 μm, E 1 / E 2 increases from 1. 79 to 5. 11 ( Figure 2C, D ). By increasing angle θ from 40° to 80°, E 1 / E 2 increases from 0. 85 to 3. 48 ( Figure 2E, F ). Hence, by varying any of these three dimensions, we can tune the cardiac patch to have an anisotropic ratio of effective stiffness matching the reported ratio for native heart tissue (rat heart E trans / E long : 1. 9 to 3. 9). [ 27 – 30 ] This technique provides greater flexibility and control over a wide range of anisotropic mechanical properties, allowing the material to be tailored to match patient needs with varying ages or health conditions. This is in contrast to Engelmayr et al. who achieved a fixed anisotropy of effective stiffness of 2. 7 with their nonconductive, nonauxetic micropattered cardiac patch. [ 29 ] Varying R has little effect on E 1 / E 2 ( Figure 2G, H ). However, when decreasing R from 150 to 50 μm, the magnitudes of E 1 and E 2 decrease from 2. 77 ± 0. 5 to 0. 39 ± 0. 2 MPa and from 1. 10 ± 0. 1 to 0. 14 ± 0. 04 MPa, respectively, providing an additional level of control in mechanical properties. As well as matching the anisotropic ratio of effective stiffness, this method allows us to match the effective stiffness of the cardiac patch to native human heart tissue (0. 02–0. 50 MPa). [ 7, 8 ] Ultimate tensile strength (UTS) and strain at failure are in the range of 0. 06 ± 0. 03 to 1. 53 ± 0. 9 MPa and 27 ± 6 to 96 ± 29%, respectively, which are comparable to reported values for native human heart tissue. [ 7, 31 ] Poisson’s ratio values were negative for all but one of the sets of bow-tie dimensions. The exception was for the patch with an angle between the ribs of 80°, which leaves little room for it to expand laterally when stretched (Poisson’s ratio of ν 12 : 0. 00 ± 0. 00 and ν 21 : 0. 08 ± 0. 1). The majority of the patches were found to have Poisson’s ratios in the range of −1. 45 ± 0. 2 to −0. 15 ± 0. 1 (Table S1, Supporting Information ). The Poisson’s ratio can also be tuned to a desired value by tuning the bow-tie dimensions, as has been demonstrated before with a nonconductive poly(ethylene glycol) material by Zhang et al. [ 15 ] They developed a biomaterial patterned with the same re-entrant honeycomb geometry with a tuned Poisson’s ratio for the purpose of probing how cells sense and respond to these local subtle differences in mechanical properties. [ 15 ] However, their two-photon fabrication technique is restricted to photocurable materials, whereas the excimer laser microablation fabrication techniques used here can be used for almost any biomaterial. As a result of the wide range of effective stiffness (human heart: 0. 02–0. 50 MPa)[ 7, 8 ] and anisotropy (rat heart: 1. 9–3. 9), [ 27 – 30 ] reported for native healthy heart tissues, the ideal mechanical properties for a cardiac patch are still under debate. Our new approach provides a promising technique to probe the appropriate biomaterial mechanical properties for treating infarcted hearts. The biomaterials can be tuned to a wide range of desired values of effective stiffness and anisotropy while maintaining the bulk properties. Moreover, this control over the mechanical properties is particularly exciting as it can be extended to other tissues and applications where the tissues are expanded under stress, such as arterial stents, [ 32 ] esophageal stents, [ 14 ] and osteochondral implants[ 33, 34 ] among others. 2. 3 Characterization of Conductivity The AuxCPs maintain a similar level of conductivity compared to the unpatterned cardiac patches (UnpatCPs) previously produced, [ 24 ] with values in the 10 −2 S cm −1 range (UnpatCP: 13. 6 ± 2. 9 × 10 −2 S cm −1, AuxCP 1-dir: 9. 3 ± 1. 5 × 10 −2 S cm −1 and AuxCP 2-dir: 2. 4 ± 0. 9 × 10 −2 S cm −1 ; Figure 3 ). Interestingly, the conductivity of the AuxCP is also anisotropic. The data show that conductivity is significantly higher in the 1-direction than the 2-direction. Similarly, the results also show that the UnpatCP has a significantly higher conductivity than the AuxCP in the 2-direction. The stated conductivity values are calculated by taking the length between the two electrodes as the current travel distance. While this is true for the UnpatCP, it is more complex for the AuxCP. By following the pattern in each direction, it can be seen that the distance the current travels is longer than the distance between the two electrodes, with the longest route being in the 2-direction. The inverse of these calculated effective distances correlates well with the trend in conductivities, as can be seen in Figure 3. Overall the patterning has a significant yet small effect on the conductivity of the patches, as they maintain conductivities in the 10 −2 S cm −1 range. The conductivity values measured for the AuxCPs are comparable with other polyaniline-based composites used for cardiac tissue engineering from literature. For example, the polyaniline-poly(glycerol sebacate) composite produced by Qazi et al. [ 35 ] has conductivities in the 10 −3 S cm −1 range with 15 v/v% polyaniline and in the 10 −2 S cm −1 range with 30 v/v% polyaniline. [ 35 ] However, we previously reported that our chitosan-polyaniline composite has superior electrical stability over other such polyaniline-based composites, [ 24 ] a vital characteristic for tissue engineering applications. Moreover, the electrical conductivities reported here for the AuxCPs fall within the range believed to be relevant for tissue engineering applications (10 −4 –10 S cm −1 ). [ 20 ] However, the ideal electrical conductivity for cardiac biomaterials is still undefined and is an area that would benefit from further investigation. 2. 4 Neonatal Rat Ventricular Myocytes Cultured on Auxetic Cardiac Patches In vitro cell culture experiments were carried out to demonstrate the cytocompatibility of the patches, prior to ex vivo and in vivo experimentation. Both neonatal rat ventricular myocytes (NRVM) and fibroblasts were cultured on AuxCPs for 3 d (Figure S4A, B, Supporting Information ). The cells were well dispersed along the pattern and presented the elongated phenotype expected of cardiomyocytes. Moreover, the apparent striation pattern, positive for α -actinin, confirms cardiomyocyte phenotype (Figure S4B, Supporting Information ). Some NRVMs can be seen to take on a more circular morphology and form aggregates, suggesting dead cells. However, a cell proliferation assay (see the Experimental Section, Supporting Information, for a description) demonstrates that cells on both the AuxCPs and UnpatCPs have significantly higher levels of cell metabolic activity compared to the glass control (Figure S4C, Supporting Information ). Hence, both the AuxCPs and UnpatCPs can be considered cytocompatible for both cardiac fibroblasts and cardiomyocytes. 2. 5 Ex Vivo Investigation of Cardiac Electrophysiology After optimizing the auxetic mechanical properties, characterizing the conductivity and demonstrating the cytocompatibility of these cardiac patches, we investigated the impact of the patches on the electrophysiological and mechanical function of the heart. First, to probe the effect of the AuxCPs electrical conduction, we carried out ex vivo experiments on both ultrathin myocardial slices and whole rat hearts[ 24 ] using AuxCP-1 (Table S1, Supporting Information ; setup number 1). The results suggested that AuxCP-1 had an anisotropy in conductivity between directions 1 and 2 ( Figure 3 ) and was therefore an interesting candidate for investigating the impact of its electrical conductivity when attached to the heart tissue in two possible orientations (1-dir: where the more conductive 1-direction of the AuxCP is positioned parallel to the longitudinal axis of the cardiac tissue; 2-dir: where the more conductive 1-direction of the AuxCP is positioned perpendicular to the longitudinal axis of the cardiac tissue). 2. 5. 1 Impact of Auxetic Cardiac Patches on Myocardial Slices Upon electrical field stimulation, we observed a reduction in contractility when placing the AuxCPs on myocardial slices, compared to the slices on their own. This is reported as the relative contractility, that is, 100% without a patch to; Slice + 1-dir AuxCP: 82. 6 ± 5. 0% and Slice + 2-dir AuxCP: 86. 0 ± 4. 1%; Figure 4Ai. However, we did not observe a significant difference between the two possible AuxCP orientations. In both cases, we observed a reduced contractility of the slices. We hypothesize that this effect is due to the electrical conductivity of the AuxCP. To this end, we conducted the same experiment with nonconductive chitosan patches, consisting of the same auxetic micropattern yet without the polyaniline coating. Contractility upon electrical stimulation of the myocardial slices with this chitosan patch was comparable to slices without any patch. This suggests that the presence of the doped-polyaniline is critical to the observed reduction in contractility, which could be attributed to the electroactive nature of the polyaniline-chitosan composite patch compared to the chitosanonly patch. This is comparable with our previous work, which found that the UnpatCPs reduced the contractility of the slices while the chitosan-only patches of the same architecture did not influence the contractility of the slices. [ 24 ] Interestingly, however, this previous work also showed a size-dependent effect, with the larger (25 mm 2 ) UnpatCPs having the greatest effect on contractility compared to the smaller patches (17 and 9 mm 2 ), where no reduction in contractility was observed. [ 24 ] In the experiments reported here, we used AuxCPs of similar dimensions (21. 1 mm 2 ). Taking the patterning into consideration, the total surface area of the AuxCPs, and therefore the material placed on the myocardial slices, is considerably less (10. 0 mm 2 ). Contrary to our expectations and our previous work, this reduced surface area still affected the contractility of the myocardial slices. We thus propose that the contractility of myocardial slices upon electrical stimulation is influenced not only by the size, weight, and conductivity of a patch, but also by its geometry and auxetic properties. The ex vivo evaluations described here further underpin the importance of considering various aspects in cardiac patch design, and highlight the complexity in trying to elucidate the role of these different parameters from one another experimentally. To gain further insight into the true effect of the AuxCPs on cardiac function, further ex vivo, as well as in vivo experiments have been conducted and their descriptions follow. A microelectrode array (MEA) system[ 24 ] was also used to examine the AuxCPs on the myocardial slices, which measures signal propagation through the tissue upon electrical field stimulation, allowing activation time maps to be produced along with the corresponding calculated peripheral CVs ( Figure 4Aii ). Because CV is faster along the fibers, [ 24 ] the slice is stimulated in two orientations, longitudinal in-line with the fibers and transverse across the fibers. The peripheral CVs for both slice orientations (longitudinal and transverse) are unaffected by the addition of the AuxCPs in either direction (1- or 2-direction) and there is no significant difference seen for the chitosan patches ( Figure 4Aii ). Previously, we have reported a reduced CV for the myocardial slices with UnpatCPs both in the transversal and longitudinal directions on the MEA system. [ 24 ] Here, we show that the AuxCPs do not restrict the CV neither in the transversal nor in the longitudinal directions on the MEA system, which could be due to the significantly smaller surface area of the AuxCPs (10 mm 2 ) compared to the UnpatCPs (25 mm 2 ). However, this would contradict the results seen for the force transducer contractility measurements. The MEA system stimulates and reads the response specifically from the bottom side of the tissue while the patch is placed on the top side of the tissue. This is different from the force transducer contractility measurements, which measures an overall effect on the tissue. Hence, it is also possible that the effect of the AuxCPs on CV is restricted to the surface of the tissue and does not pass all the way through the tissue (≈15 cell layers) to the MEA dish electrodes. 2. 5. 2 Impact of Auxetic Cardiac Patches on Whole Hearts Optical mapping experiments[ 24 ] were conducted to investigate the effect of the AuxCPs on the electrophysiology of whole hearts. The optical mapping measurements were taken before and after application of the AuxCPs. The optical measurements were processed to produce activation time maps and conduction velocity maps ( Figure 4B ). As expected, prior to application of the patches, the CV is significantly reduced in the MI hearts compared to the healthy hearts ( Figure 4Ci ). Similar to the heart slice measurements, the patch was placed in two different orientations. The orientation of the patch had an impact on the CV for healthy hearts. The patch oriented in the 1-direction significantly reduced the CV (from 100% without a patch to 82. 9 ± 5. 5, 84. 9 ± 7. 7, and 87. 5 ± 8. 9% for the local basal, local apical, and base to apex measurements, respectively). The patch oriented in the 2-direction had no significant impact on the CV of healthy hearts ( Figure 4Cii ). In contrast, with MI hearts neither orientation of the patch had a significant impact on CV ( Figure 4Ciii ). The results for the UnpatCPs and nonconductive unpatterned chitosan patches were published previously[ 24 ] and so for animal welfare reasons these experiments were not repeated. We previously showed a significant decrease in CV in healthy hearts and a significant increase in CV in MI hearts when the UnpatCPs were applied, and no significant difference in either case for the chitosan patches. [ 24 ] This whole heart CV data combined with the myocardial slice CV data suggests that the UnpatCPs could influence the electrophysiological signaling of the heart. [ 24 ] In this study, we show that the AuxCPs do not induce a significant difference in CV for both healthy and MI whole hearts and healthy slices, when oriented in the correct direction (2-dir: less stiff patch direction aligned with less stiff longitudinal direction of the heart). Furthermore, an additional AuxCP (AuxCP-10) was designed with a new set of bow-tie dimensions to maximize surface area (AuxCP-10 = 18. 7 mm 2, AuxCP-1 = 10. 0 mm 2 ), and this design did not influence the CV of either healthy or MI hearts in both 1- and 2-directions (Table S2 and Figure S5, Supporting Information ). Therefore, this demonstrates that the AuxCPs do not have a detrimental effect on the electrophysiology of both healthy and infarcted hearts over a wide range of dimensions and surface areas. Nevertheless, these preclinical experiments cannot fully predict the effect of AuxCPs on human heart electrophysiology and function. As auxetic behavior is independent of scale, larger AuxCPs will maintain similar mechanical properties. In addition, the excimer laser microablation process lends itself well to scale-up; with the ability for high-throughput production; and patch size only limited by the XY stage dimensions (currently ≈150 × 150 mm) and thus theoretically suitable for up to human heart dimensions. However, left ventricular wall thickness and heart rate are significantly different between rats and humans, and the effects of patch application in larger animals and human heart samples require further investigation to inform future clinical applications of this technology. The increase in CV previously observed when the UnpatCPs were applied to MI hearts is an interesting and promising result. [ 24 ] Despite the observed reduction in CV for ex-vivo healthy hearts upon application of the UnpatCPs, in vivo implantation of the patch in healthy hearts showed that it did not influence the proarrhythmic state of the heart under stress. Nonetheless, from both studies it appears that an ideal case would be to tune the conductive properties of the cardiac patches so that they increase the CV in the infarct area closer to the values expected of healthy heart tissue. This could conceivably be achieved with a deeper understanding of how to manipulate cardiac electrophysiology, along with further optimization of pattern dimensions and the material’s conductive properties. Currently, the details of the molecular mechanism by which doped-conjugated polymer based materials affect cardiac electrophysiology are not fully understood and this is an interesting question worth further investigation in future studies. 2. 6 Ex Vivo Investigation of Mechanical Integration The main purpose of micropatterning the cardiac patch was to improve its mechanical conformability to native heart tissue. In order to investigate and quantify this, a bespoke, custom-made device was built, in which dissected sections of the left ventricle could be placed into grips, and cyclically stretched at specific distances and frequencies ( Figure 5A ). For these experiments AuxCP-2 was used (Table S1, Supporting Information ; setup number 2), as this has an anisotropy of E 1 / E 2 = 2. 79, which falls in the middle of the mechanical anisotropy range previously reported for native heart tissue (rat heart E trans / E long : 1. 9 to 3. 9). [ 27 – 30 ] The patches were always oriented to match the mechanical anisotropy of the patch to the tissue (i. e. , the less stiff direction of the patch (2-direction) parallel to the less stiff longitudinal axis of the heart tissue). The patches were attached to the tissue by the previously described laser photoadhesion technique. [ 24 ] Longitudinal strains, transverse strains, and the ratio of the two strains were measured from digital microscope videos of the tissue as it was stretched with either an AuxCP or UnpatCP attached, and normalized to the tissue before the patch was attached ( Figure 5B, C ). When analyzing strains across the entire tissue (≈12 × 12 mm initial tissue size) there was no significant difference in the transverse and longitudinal strains for both AuxCPs and UnpatCPs compared to the tissue without a patch ( Figure 5B ), suggesting that patches do not interfere with global tissue mechanics. However, strain measurements across the patch area (≈6 × 6 mm initial patch size) demonstrated significant decreases in both longitudinal and transverse strains for UnpatCPs compared to both the tissue without a patch and with AuxCPs. Further, UnpatCPs had a normalized ratio of strains of −0. 70 ± 0. 6 (Tran/Long), demonstrating a dramatic reversal in mechanics. We observed a greater conformability of the AuxCPs to native tissue movement compared to the UnpatCPs. The AuxCPs significantly decreased strains across the patch area compared to the tissue without a patch. However, interestingly, although strains were reduced with the AuxCPs, they maintained a similar normalized ratio of strains as tissue without a patch (1. 20 ± 0. 3 and 1 ± 0, respectively), suggesting that AuxCPs stretch and conform to the native tissue movements while providing mechanical support ( Figure 5C ). 2. 7 In Vivo Investigation of Cardiac Function after Myocardial Infarction We used a rat MI model to investigate the effect of the AuxCPs on cardiac function. Once again, we used AuxCP-2 oriented to match the mechanical anisotropy of the patch to the hearts anisotropy ( Figure 1B ). The AuxCP was attached to the heart by the previously described laser photoadhesion technique, [ 24 ] immediately after induction of MI through permanent ligation of the left anterior descending coronary artery (LAD) (MI AuxCP, N = 6). We used this model as an initial proof of principle study, although a more clinically relevant scenario would be to apply the patch after the MI has resolved and scar has formed. For animal welfare reasons a second thoracotomy, which would be required for such experiments, was avoided at this stage. However, a more clinically relevant chronic heart failure model will be important in future experiments. While some progress has been made in this field to develop a less invasive surgical technique for the application of a cardiac patch to the heart after an MI, [ 28 ] currently this would still require a second thoracotomy, which is highly challenging in rodents. The AuxCPs remained intact and attached to the hearts two weeks after the surgeries ( Figure 6A ). A control group also underwent the same surgery, with an induced MI and application of the laser, but without an AuxCP applied (MI control, N = 4). In addition, a healthy control group underwent the same surgery, with the application of the laser, but without inducing an MI and without applying the AuxCP (sham, N = 6). Histological analysis shows no significant difference between the MI controls and the MI AuxCP hearts, suggesting negligible fibrotic response to the cardiac biomaterial two weeks after the surgeries ( Figure 6B, C ). Cardiac structure and function was assessed in vivo at 1 and 14 d after the surgery using ultrasound. The MI control and MI AuxCP groups showed reduced cardiac function compared to the healthy sham controls as expected. Both MI groups show similar cardiac function at 1 and 14 d after the surgeries for fractional shortening, left ventricular (LV) volume at end-diastole, LV volume at end-systole and mitral valve (MV) E / A peak ratio, suggesting that the AuxCPs have no detrimental effect on cardiac function ( Figure 6D–G ). However, by ultrasound measurements we also show that while the LV mass is similar for all groups at 1 day, by 14 d a significant increase in LV mass is seen for the MI control group which is not seen for the MI AuxCP group. A possible explanation may be that the patch reduces wall stress and attenuates hypertrophy ( Figure 6H ), as has been shown previously with other cardiac patches. [ 36 ] However, further evaluation of cardiomyocyte size and gene expression would be required to elucidate the role of the AuxCPs in attenuating hypertrophy. Greater benefit to cardiac function may have been observed over a longer follow-up period or under a chronic heart failure model, giving more time for LV remodeling to occur, which will be important considerations in future in vivo experiments for these AuxCPs. Nevertheless, these results suggest that the AuxCPs integrate well with native heart tissue in vivo and have no detrimental effect on cardiac function, nor do they induce a significant fibrotic response. This approach to biomaterial design creates an ideal platform for future incorporation of therapeutic molecules for delivery to the epicardium. 3 Conclusion We can create auxetic micropatterned cardiac patches by excimer laser microablation with mechanical properties tuned to match those of native heart tissue while maintaining the bulk properties of the material. We showed that the AuxCPs are conductive (≈10 −2 S cm −1 ) and cytocompatible with cardiomyocytes. Ex vivo experiments demonstrated that the AuxCPs have no detrimental effect on the cardiac electrophysiology of both healthy and MI hearts. Further new ex vivo experiments showed that the AuxCPs stretch and conform to match the movements of native heart tissue, unlike the UnpatCPs. Finally, the AuxCPs integrated with native heart tissue without detrimental effect on cardiac function in a rat MI model over two weeks in vivo. This study demonstrates a conductive cardiac patch improved through the use of an auxetic design, which can be tuned to match the mechanical demands of the heart and is a promising cardiac patch. Further, there are still many unanswered questions in the cardiac field as to the appropriate mechanics and conductivity of patches and our auxetic patch design is a promising tool for investigating optimal properties due to its high degree of tunability. Hence, the findings presented here may lead to improved biomaterial designs toward treating myocardial infarction. 4 Experimental Section Extended experimental descriptions can be found in the Supporting Information. All chemical reagents were purchased from Sigma Aldrich (UK) or VWR (UK) unless specifically noted. All animal procedures were carried out in accordance with the UK Home Office Animals (Scientific Procedures) Act 1986 and Directive 2010/63/EU of the European Parliament on the protection of animals used for scientific purposes. Fabrication of Auxetic Cardiac Patches—Fabrication and Physical Characterization of Chitosan-Polyaniline Films The chitosan-polyaniline films were fabricated by the method previously described by Mawad et al. [ 24 ] For the AuxCPs, chitosan films were micropatterned with a re-entrant honeycomb (bow-tie) pattern, before coating with polyaniline and phytic acid. Fabrication of Auxetic Cardiac Patches—Micropatterning by Excimer Laser Microablation Chitosan films on glass slides were placed on the XYZ stage (accuracy ± 1 μm) of a custom-built excimer laser processing system with an LPX305i Excimer Laser (Lambda-Physik) charged with krypton fluoride gas (emitting 25 ns pulses of light at 248 nm). The light pulses passed through various optics in order to illuminate the part of the sample to be ablated, as described previously. [ 37 ] The metal mask (0. 15 mm thick brass) used in this work contained the enlarged (by five times) re-entrant honeycomb geometry (12. 5 × 12. 5 mm 2 ), which was then focused on the sample through a 5× projection lens (2. 5 × 2. 5 mm 2 ). The programmable XYZ sample stage moved the sample by a precise distance for the ablation to be repeated; hence, fabrication of larger patch areas of a repeating pattern was possible. The optimum conditions for ablating the chitosan films were identified as 600 pulses at 350 mJ cm −2 delivered at 40 Hz. Characterization of Tunable Mechanical Properties—Mechanical Characterization Tensile tests were carried out on patches with 6 × 6 repeat units in both 1- and 2-directions of the patch (directions annotated in Figure 1A ). Precise length and width measurements were known from the fabrication technique. Thickness was determined by scanning electron microscopy, operated at 10 kV (Figure S6, Supporting Information, 29. 12 ± 6. 8 μm, N = 3, n = 5). Patches were wetted with deionized water and mounted using custom made stainless steel grips on an Electroforce 3200 mechanical tester with a 250 g load cell (TA Instruments, New Castle, DE) controlled by WinTest software (Ver. 7). Patches were strained to failure at a rate of 0. 1% strain s −1, assuming quasi-static loading. The effective stiffnesses ( E ) were determined by taking the slope of a regression within the initial linear region of the stress–strain curve up to 10% strain. The anisotropic ratio of effective stiffnesses ( E 1 / E 2 ) was calculated by dividing the mean of E 1 ( N = 10) by the mean of E 2 ( N = 10). UTS was measured as the maximum stress reached and the strain-at-failure was taken as the strain at the UTS point. Ex Vivo Investigation of Cardiac Electrophysiology—Myocardial Slice Measurements Live rat left ventricular myocardial slices were prepared according to the protocol described by Watson et al. [ 38 ] Myocardial slice force transducer and microelectrode array measurements were recorded as previously described by Mawad et al. [ 24 ] Ex Vivo Investigation of Cardiac Electrophysiology—MI Surgical Procedure and Ex Vivo Optical Mapping Experiments Adult male S-D rats (250–350 g) were subjected to MI induced by ligation of the LAD and optical mapping was performed as described previously. [ 24 ] Data were analyzed using a bespoke MATLAB script as described previously. [ 24, 39 ] Ex Vivo Investigation of Mechanical Integration A custom-built device was used to cyclically load rat left ventricles with and without an AuxCP or UnpatCP attached. The heart was rapidly explanted and rinsed free of blood in ice-cold oxygenated Krebs–Henseleit solution (composition in mM: 119 NaCl, 4. 7 KCl, 0. 94 MgSO 4, 1 CaCl 2, 1. 2 KH 2 PO 4, 25 NaHCO 3, 11. 5 glucose, and equilibrated with 95% O 2 + 5% CO 2 ), containing heparin (12 U mL −1 ) and 30 × 10 −3 m 2, 3-butanedione monoxime (BDM). BDM is a myosin ATPase inhibitor used to inhibit contraction and reduce tissue damage during dissection and cyclic stretching. The right ventricle was removed, and the septum bisected to open the left ventricle. The papillary muscles were cut, and remaining septum removed to leave the left ventricular free wall intact and open flat. The apical and basal ends of the left ventricular free wall were glued via the epicardial surface onto a plastic insert for the biomechanical rig. The biomechanical rig device consists of a linear actuator which can be set to cycle at specific distances, rates, and frequencies. The tissue was marked with ink in a grid formation, for image analysis. The tissue was cycled ±1 mm under tension at 1 Hz and filmed with a Dino-lite Edge AM4815ZT digital microscope (Brunel Microscopes Ltd. , Wiltshire, UK). The AuxCP or UnpatCP (≈10 × 10 mm) was then secured to the epicardium using the laser technology described and shown to be safe for tissue applications previously. [ 24 ] Briefly, the polyaniline free border of the patch (1–2 mm), containing chitosan and Rose Bengal (0. 1 w/v%), was irradiated by a green diode-pumped solid state laser (532 nm; CNI Lasers, China). The laser is set with a power of 170 mW in a continuous wave with a beam spot diameter on the tissue of 6 mm. The patch border was spot-irradiated for a total time of 4 min. The AuxCPs were always oriented so the less stiff 2-direction was aligned with the fibers. The cyclic loading was then repeated. Image processing software Fiji (free download under https://fiji. sc/ ) was used to calculate transverse and longitudinal strains. The strain values were then normalized to the corresponding tissue section before the patch was applied. In Vivo Investigation of Cardiac Function after Myocardial Infarction—MI Surgical Procedure and Patch Implantation Adult male S-D rats (250–350 g) were subjected to MI induced by ligation of the LAD, see the Supporting Information for further details. The patches were secured in place using laser technology described above. [ 24 ] COVA + CARD (Biom’up, France), a cardiac postoperative adhesion preventing membrane, was applied to the outside of the AuxCP (or on the epicardium in the case of the control groups), and sutured in 3–4 places, to the pericardium and thymus. Statistical Analysis Mechanical properties and conductivity measurements are expressed as mean ± standard deviation (SD). In vitro, ex vivo, and in vivo measurements are expressed as mean ± standard error of the mean (SE). The nonparametric Mann–Whitney post hoc test was used for comparison (OriginPro 9. 1, OriginLab Corporation). Values of p < 0. 05 were considered significant. Supplementary Material Supporting Information is available from the Wiley Online Library or from the author. Supporting information
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10. 1002/adfm. 201801850
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Advanced functional materials
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3D Printed Stem-Cell Derived Neural Progenitors Generate Spinal Cord Scaffolds
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A bioengineered spinal cord is fabricated via extrusion-based multi-material 3D bioprinting, in which clusters of induced pluripotent stem cell (iPSC)-derived spinal neuronal progenitor cells (sNPCs) and oligodendrocyte progenitor cells (OPCs) are placed in precise positions within 3D printed biocompatible scaffolds during assembly. The location of a cluster of cells, of a single type or multiple types, is controlled using a point-dispensing printing method with a 200 μm center-to-center spacing within 150 μm wide channels. The bioprinted sNPCs differentiate and extend axons throughout microscale scaffold channels, and the activity of these neuronal networks is confirmed by physiological spontaneous calcium flux studies. Successful bioprinting of OPCs in combination with sNPCs demonstrates a multicellular neural tissue engineering approach, where the ability to direct the patterning and combination of transplanted neuronal and glial cells can be beneficial in rebuilding functional axonal connections across areas of central nervous system (CNS) tissue damage. This platform can be used to prepare novel biomimetic, hydrogel-based scaffolds modeling complex CNS tissue architecture in vitro and harnessed to develop new clinical approaches to treat neurological diseases, including spinal cord injury.
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10. 1002/adfm. 201906330
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Advanced functional materials
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3D Printed Cartilage-Like Tissue Constructs with Spatially Controlled Mechanical Properties
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Developing biomimetic cartilaginous tissues that support locomotion while maintaining chondrogenic behavior is a major challenge in the tissue engineering field. Specifically, while locomotive forces demand tissues with strong mechanical properties, chondrogenesis requires a soft microenvironment. To address this challenge, 3D cartilage-like tissue is bioprinted using two biomaterials with different mechanical properties: a hard biomaterial to reflect the macromechanical properties of native cartilage, and a soft biomaterial to create a chondrogenic microenvironment. To this end, a hard biomaterial (MPa order compressive modulus) composed of an interpenetrating polymer network (IPN) of polyethylene glycol (PEG) and alginate hydrogel is developed as an extracellular matrix (ECM) with self-healing properties, but low diffusive capacity. Within this bath supplemented with thrombin, fibrinogen containing human mesenchymal stem cell (hMSC) spheroids is bioprinted forming fibrin, as the soft biomaterial (kPa order compressive modulus) to simulate cartilage’s pericellular matrix and allow a fast diffusion of nutrients. The bioprinted hMSC spheroids improve viability and chondrogenic-like behavior without adversely affecting the macromechanical properties of the tissue. Therefore, the ability to print locally soft and cell stimulating microenvironments inside of a mechanically robust hydrogel is demonstrated, thereby uncoupling the micro- and macromechanical properties of the 3D printed tissues such as cartilage.
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No full text available
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10. 1002/adfm. 201907102
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Advanced functional materials
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Creating Physicochemical Gradients in Modular Microporous Annealed Particle Hydrogels via a Microfluidic Method
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Microporous annealed particle (MAP) hydrogels are an attractive platform for engineering biomaterials with controlled heterogeneity. Here, we introduce a microfluidic method to create physicochemical gradients within poly(ethylene glycol) based MAP hydrogels. By combining microfluidic mixing and droplet generator modules, microgels with varying properties were produced by adjusting the relative flow rates between two precursor solutions and collected layer-by-layer in a syringe. Subsequently, the microgels were injected out of the syringe and then annealed with thiol-ene click chemistry. Fluorescence intensity measurements of constructs annealed in vitro and after mock implantation into a tissue defect showed that a continuous gradient profile was achieved and maintained after injection, indicating utility for in situ hydrogel formation. The effects of physicochemical property gradients on human mesenchymal stem cells (hMSCs) were also studied. Microgel stiffness was studied first, and the hMSCs exhibited increased spreading and proliferation as stiffness increased along the gradient. Microgel degradability was also studied, revealing a critical degradability threshold above which the hMSCs spread robustly and below which they were isolated and exhibited reduced spreading. This method of generating spatial gradients in MAP hydrogels could be further used to gain new insights into cell-material interactions, which could be leveraged for tissue engineering applications. A new droplet microfluidic approach to obtain microporous annealed particle hydrogels with physicochemical gradients is presented. Gradient formation is achieved by precisely controlling the mixing of two precursor solutions, and the gradient can be maintained after injection. This approach can be leveraged to produce new materials for tissue repair and to gain unique insights on cell-material interactions.
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10. 1002/adfm. 201909553
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Advanced functional materials
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Engineering liver microtissues for disease modeling and regenerative medicine
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The burden of liver diseases is increasing worldwide, accounting for two million deaths annually. In the past decade, tremendous progress has been made in the basic and translational research of liver tissue engineering. Liver microtissues are small, three-dimensional hepatocyte cultures that recapitulate liver physiology and have been used in biomedical research and regenerative medicine. This review summarizes recent advances, challenges, and future directions in liver microtissue research. Cellular engineering approaches are used to sustain primary hepatocytes or produce hepatocytes derived from pluripotent stem cells and other adult tissues. Three-dimensional microtissues are generated by scaffold-free assembly or scaffold-assisted methods such as macroencapsulation, droplet microfluidics, and bioprinting. Optimization of the hepatic microenvironment entails incorporating the appropriate cell composition for enhanced cell-cell interactions and niche-specific signals, and creating scaffolds with desired chemical, mechanical and physical properties. Perfusion-based culture systems such as bioreactors and microfluidic systems are used to achieve efficient exchange of nutrients and soluble factors. Taken together, systematic optimization of liver microtissues is a multidisciplinary effort focused on creating liver cultures and on-chip models with greater structural complexity and physiological relevance for use in liver disease research, therapeutic development, and regenerative medicine.
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10. 1002/adfm. 201909556
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Advanced functional materials
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Recent Advances in Smart Biomaterials for the Detection and Treatment of Autoimmune Diseases
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Autoimmune diseases are a group of debilitating illnesses that are often idiopathic in nature. The steady rise in the prevalence of these conditions warrants new approaches for diagnosis and treatment. Stimuli-responsive biomaterials also known as “smart”, “intelligent” or “recognitive” biomaterials are widely studied for their applications in drug delivery, biosensing and tissue engineering due to their ability to produce thermal, optical, chemical, or structural changes upon interacting with the biological environment. This critical analysis highlights studies within the last decade that harness the recognitive capabilities of these biomaterials towards the development of novel detection and treatment options for autoimmune diseases.
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10. 1002/adfm. 201909802
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Advanced functional materials
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Metabolic Labeling to Probe the Spatiotemporal Accumulation of Matrix at the Chondrocyte–Hydrogel Interface
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Hydrogels are engineered with biochemical and biophysical signals to recreate aspects of the native microenvironment and to control cellular functions such as differentiation and matrix deposition. This deposited matrix accumulates within the pericellular space and likely affects the interactions between encapsulated cells and the engineered hydrogel; however, there has been little work to study the spatiotemporal evolution of matrix at this interface. To address this, metabolic labeling is employed to visualize the temporal and spatial positioning of nascent proteins and proteoglycans deposited by chondrocytes. Within covalently crosslinked hyaluronic acid hydrogels, chondrocytes deposit nascent proteins and proteoglycans in the pericellular space within 1 d after encapsulation. The accumulation of this matrix, as measured by an increase in matrix thickness during culture, depends on the initial hydrogel crosslink density with decreased thicknesses for more crosslinked hydrogels. Encapsulated fluorescent beads are used to monitor the hydrogel location and indicate that the emerging nascent matrix physically displaces the hydrogel from the cell membrane with extended culture. These findings suggest that secreted matrix increasingly masks the presentation of engineered hydrogel cues and may have implications for the design of hydrogels in tissue engineering and regenerative medicine.
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No full text available
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10. 1002/adfm. 201910250
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Advanced Functional Materials
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Bioprinting Neural Systems to Model Central Nervous System Diseases
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Abstract To date, pharmaceutical progresses in central nervous system (CNS) diseases are clearly hampered by the lack of suitable disease models. Indeed, animal models do not faithfully represent human neurodegenerative processes and human in vitro 2D cell culture systems cannot recapitulate the in vivo complexity of neural systems. The search for valuable models of neurodegenerative diseases has recently been revived by the addition of 3D culture that allows to re‐create the in vivo microenvironment including the interactions among different neural cell types and the surrounding extracellular matrix (ECM) components. In this review, the new challenges in the field of CNS diseases in vitro 3D modeling are discussed, focusing on the implementation of bioprinting approaches enabling positional control on the generation of the 3D microenvironments. The focus is specifically on the choice of the optimal materials to simulate the ECM brain compartment and the biofabrication technologies needed to shape the cellular components within a microenvironment that significantly represents brain biochemical and biophysical parameters.
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1 Introduction 1. 1 State of the Art of Neural In Vitro Models Central nervous system (CNS) diseases, such as Parkinson's disease (PD) and Alzheimer's disease (AD), involve progressive dysfunction of distinct neuronal populations. Notwithstanding the growing amount of studies on CNS diseases, our understanding of brain functioning is still largely unsatisfactory and we have specifically strong limitations in the in vivo study of the human brain. On the other side, limited source of human neural cells used to be the main obstacle for in vitro studies. For this reason animal models have been extensively explored as alternate to elucidate mechanisms of human nervous diseases, [ 1 ] but they cannot answer the questions that are related to the specific individual variability, sensitivity, and complexity of the human brain. [ 2 ] Therefore, besides animal models, we strongly need to employ in vitro models of human brain tissue that could substantially help the development of new therapeutic tools. To this aim biomaterial engineering, biofabrication and cell reprogramming technologies can help to design a new set of bioinspired systems that will be used in the near future to model all human brain physiology, pathology, and pharmacology. 3D culture has emerged as a new tool in the field of cell biology and physiology to study cell–cell and cell–extracellular microenvironment interactions in a more in vivo like situation. Cells behave rather differently in a 3D system than in a traditional 2D system in terms of morphology, viability, proliferation, differentiation, and gene expression profile that result to be closer to the in vivo situation. [ 3 ] Moreover, in vitro 3D models serve as a platform in a more convenient, economic and high‐throughput way to investigate toxicology and drug discovery. [ 1, 4 ] The use of spheroid‐ or organoid‐based methods has given rise to sophisticated in vitro tissue models such as brain, [ 4, 5 ] which present, to certain extent, either key structures or key functions that closely resemble neural tissues. Spheroid‐ or organoid‐based neural tissue models have the advantage to exploit internally intrinsic signaling and cues to drive spontaneous differentiation of embryonic stem cells (ESCs)/induced pluripotent stem cells (iPSCs)/neural stem cells (NSCs)/neural progenitor cells (NPCs) into a mixed population of neurons and glia. [ 6 ] In addition, cells within a self‐assembled structure secrete their own extracellular matrix (ECM), which closely resembles the desired matrix stiffness and composition of the in vivo microenvironment. [ 7 ] Notably, it has recently been shown that brain organoids can also be polarized giving rise to neural phenotypes that resemble specific CNS territories. [ 8 ] Limitations of the brain organoid system include reproducibility; indeed, the consistency of these models is poorly controlled in terms of size and cell viability. Moreover, most of the spheroid‐ or organoid‐based models fail to grow beyond 2–3 mm and longer than a few months since necrosis is always found in the core area due to insufficient waste and nutrients exchange. [ 9 ] An additional vasculature induction may allow reaching bigger sizes and longer cultivation times. [ 10 ] 1. 2 Challenges in Neural In Vitro Modeling Notwithstanding the recent progresses in the development of neural 3D models, many hurdles are still hindering their implementation in the biomedical field. Indeed, CNS likely bears the most complex tissue architecture of our body and spontaneous growth in 3D matrix leads to the generation of tissue‐like structures that can only recapitulate the early neurodevelopment stages that are therefore not helpful to model most neurodegenerative diseases. Therefore, in order to model complex neural structures (i. e. , neuronal pathways), we need to have precise positional control of the neural cells and of the microenvironment complexity. [ 11 ] Bioprinting emerges as a powerful tool to impose positional control in 3D neural models. Indeed, this technology can bring together biomaterials, bioactive factors, and cells to fabricate 3D cellular structures that mimic in vivo neural architecture characteristics. [ 12 ] Bioink development has a pivotal role in the progress of the bioprinting field. [ 13 ] Currently, for neural tissue engineering, natural hydrogels such as Matrigel and collagen still perform best in forming 3D cellular networks while keeping the highest cell viability. [ 14 ] However, the batch‐to‐batch inconsistencies of natural hydrogels and their not‐xeno‐free condition limit their use in standardized protocols. [ 15 ] A possible alternative is based on synthetic hydrogels that can be modified to support neural cell survival and maintenance by finely tuning mechanical and chemical properties. [ 16 ] In addition to functional bioinks, bioprinting allows spatial control of printed biomaterials as well as cells. Therefore, printing brain‐like structures, which comprise different neural cell types in different layers, has become a potential approach to generate complex neural tissues. [ 17 ] Moreover, with bioprinting techniques, a gradient of either mechanical or biochemical properties can easily be achieved. [ 18 ] This feature of bioprinting greatly favors research in the context of neuron axon outgrowth guidance and neural circuits modeling. So far, several bioprinting techniques have been used for neural tissue modeling and a variety of neural tissue‐like models have been constructed. [ 19 ] In this review, we will summarize the field of in vitro neural tissue engineering: current status of bioprinting building blocks and technologies; existing drawbacks of bioprinting and possible solutions; and potential future applications that could lead to the generation of complex neural in vitro models for biomedical applications ( Figure 1 ). Figure 1 Schematic representation of personalized in vitro bioprinted neural tissue models. A) Derivation of neural cells from somatic cells of an individual through direct cell reprogramming or induced pluripotent stem cell (iPSC) approach. B) Bioprinting with patient‐specific neural cells to generate neural tissue models conveying features of neuronal circuit and blood–brain barrier formation. C) Intended applications with bioprinted neural tissue models. 2 Bioprinting Technologies for Neural Cells Herein, we summarize the concept of bioprinting, categorize each technique based on their signature printing modality as well as their main dis‐/advantages ( Table 1 ), and give an outlook on the latest advances of the past 5 years in creating neural models. For further reading on the topic of biofabrication including bioprinting, we refer the reader to several recent reviews focusing specifically on biofabrication and bioprinting, [ 20 ] their application in in vitro models, [ 21 ] and on the challenges toward the bioprinting of functional living tissues. [ 22 ] Table 1 Comprehensive overview of common bioprinting techniques and its applicability for neural modeling from the last 5 years Printing modalities Reported printed features width Cell type Viability readout Advantages and salient achievements Limitations and challenges Nozzle‐based Inkjet 50–80 μm [ 26 ] Rat (adult) glial and retinal cells [ 26 ] Trypan blue: ≈69% (postprinting) [ 26 ] Delivery rate: ≈7 mm 3 min −1[ 27 ] Fabrication rate: 1–10k droplets s −1[ 28 ] Bioink viscosity range: 2–20 mPa s [ 29 ] Compatible with other biofabrication strategies involvingextrusion of thermoplastics [ 30 ] Limited structural integrity and shape fidelity [ 27, 31 ] Reported cell numbers varied per deposited drop [ 26 ] Low cell densities (<1 × 10 6 ; 2100 cells mm −1 (physiological) vs 20 cells mm −1 (printed)) [ 30 ] ≈300 μm [ 30 ] P2–4 rat primary retinal ganglion cells (RGCs) [ 30 ] Calcein AM/SYTOX: 1. 2‐fold increased survivability (printed vs nonprinted) if growth medium was added to the bioink formulation (postprinting) [ 30 ] Extrusion‐based 200 μm [ 32 ] hiPSC‐derived spinal progenitor cells (sNPCs) and oligodendrocyte progenitor cells (OPCs) [ 32 ] Calcein AM/EthD:5, 15, 30 min of exposure to nonhumidified environment: ≈98%, ≈45%, ≈0% [ 32 ] Wide range of viscosity (30 mPa s to 60 kPa s) [ 33 ] Achieved in situ reprogramming and differentiation [ 34 ] High cell densities (8 × 10 7 cells mL −1 ) [ 34a ] Shown formation of intercellular connections in neuronal networks [ 35 ] High density of cell aggregates are printable, microfluidic‐chip enabled complex prints [ 36 ] Easy accessibility: RepRap hardware [ 37 ] Embedding bioprinting strategies (e. g. , FRESH) [ 38 ] Poor structural integrity and shape fidelity in low‐viscosity inks, and potential cell damage from shear forces while extruding [ 27, 31 ] Slow printing speed may lead to gel dehydration thus requiring strategies to control the printing environment [ 32 ] ≈410 μm (nozzle diameter) [ 34b ] Adult fibroblasts [ 34b ] VB‐48/propidium Iodide (PI): ≈65% (24 h post printing) CCK‐8: ≈85% (recovery after day 7) [ 34b ] ≈200 μm (nozzle diameter) [ 34a ] hiPSC [ 34a ] PrestoBlue: proliferation over 9 days of culture [ 34a ] ≈100 μm (nozzle‐ϕ) [ 35 ] hiPSC‐derived neural precurs or cells (NPCs) [ 35 ] Calcein AM/EthD: ≈80% (postprinting) [ 35 ] ≈400 μm (anticipated by aggregate size) [ 36 ] hiPSC‐derived neural aggregates [ 36 ] Calcein AM/EthD: ≈95% (day10), ≈65% (day 15) Guava ViaCount (FACS): ≈90% (day 6) [ 36 ] Optic‐based Laser‐induced forward transfer (LIFT) ≈200 μm [ 39 ] E15 rat primary dorsal root ganglia (DRG) [ 39 ] Live‐Dye /PI: ≈85% (24 h post printing) [ 39 ] Successful prints performed with hyaluronic acid and Matrigel [ 40 ] Neurite growths reported [ 39 ] Proven in situ differentiation of bioprinted cells [ 34 ] Low cell density: ≈80 cells per drop [ 39 ] Limited manufacturer diversity might affect device accessibility [ 41 ] Not assessed [ 40 ] hiPSC [ 40 ] Trypan Blue: ≈82% (2–3 h post printing) [ 40 ] Stereolitho‐graphy (SLA) ≈190 μm [ 42 ] Mouse NSCs (NE‐4C) [ 42 ] Calcein AM/PI: ≈100–70%, 40–120 mW laser power) [ 42 ] ≈5 μm micrometer‐scale resolution a chievable [ 43 ] Custom devices a vailable [ 44 ] Combined with 3D printing using PCL fibers [ 42 ] Used with conductive graphene‐loaded bioinks [ 45 ] Potential cell damage due to UV light exposure Laser output affects cell survival [ 42 ] Potential toxicity from photosensitive resins and initiators [ 27, 31 ] ≈1k μm [ 45 ] Mouse NSCs [ 45 ] CCK‐8: Proliferation over 5 days of culture [ 45 ] Digital light processing (DLP) 50–100 μm [ 46 ] No current report on the bioprinting of cells of neuronal lineage Delivery rate: ≈20 mm 3 min −1[ 27 ] High Resolutions <100 μm achievable [ 46 ] Easily accessible: Commercial projectors [ 46a ] Biocompatible polyethylene glycol diacrylate (PEGDA) and Gel‐MA resins already available [ 47 ] Potential cell damage due to UV light exposure Potential toxicity from photosensitive resins and initiators [ 27, 31 ] Two‐photon polymerization (2PP) <1 μm [ 48 ] Highest lateral resolution ≈100 nm [ 48b ] Suitable for nano‐ and micropatterning potential Costly laser‐based equipment [ 48b ] Potential toxicity from initiators [ 49 ] John Wiley & Sons, Ltd. Conceptually the process of bioprinting encompasses the controlled and automated spatial deposition of a cell‐enriched suspension or hydrogel formulation often referred to as the “bioink. ” [ 23 ] These printable materials are used to generate 3D biological objects starting from a customizable CAD design or from a medical image. [ 24 ] Dedicated softwares convert the information in the CAD design into the instructions for the printer, which dispenses or patterns the bioink in a layer‐by‐layer fashion. [ 25 ] 2. 1 Nozzle‐Based Techniques The signature of this family of techniques is the directed deposition from a nozzle of cell suspensions, cell aggregates, or hydrogels, either as cell‐free biomaterial inks or including living cells ( Figure 2 ). When hydrogels with low mechanical properties are used, which is a typical situation the field of neural tissue engineering, these printed structures are stabilized via crosslinking mechanisms that depend on the chemistry of the ink (i. e. , ionic, pH, light‐ or enzymatically induced). The brain is the softest tissue in the human body, [ 50 ] which makes neural cells extraordinary sensitive to mechanical stimuli and their environment. [ 26, 51 ] Hence, optimizing the microenvironment is an ongoing quest especially in tailoring the hydrogel composition. [ 16 ] Particularly, stability and shape fidelity post printing are daunting challenges when using bioinks suitable for neuronal patterning, which often should display marked viscoelasticity and low stiffness. In neural tissue engineering, elastic moduli in the range 0. 1–50 kPa have shown effects on neurite sprouting, morphology as well as differentiation, and preserving neuronal stemness, [ 51b ] whereas the native compressive stiffness of brain tissue has been reported to vary regionally from 0. 3 to 2. 7 kPa. [ 52 ] Printing with hydrogels displaying similar properties may result in undesired deformation of the constructs postprinting. Additionally, even the printing process can result in decreased cell viability due to shear force during extrusion from the nozzle [ 27, 53 ] or dehydration. [ 32 ] To overcome these challenges, viable strategies are the use of spheroids, [ 54a ] more robust cell types, or stem cells that can be differentiated into neurons postprinting, [ 40 ] even via in situ reprogramming. [ 34b ] Likewise, cell viability can also be improved by accurately designing shear‐thinning bioinks, as well as bioinks that display low‐shear flow profiles, [ 44 ] or via the selection of printing methods that permit to easily pattern low‐viscosity inks for soft tissue engineering. [ 38, 55 ] Figure 2 Common bioprinting techniques can be categorized into nozzle‐based or optical‐based printing methods. Nozzle‐driven printing can generate either filaments or droplets, based on the modality of dispensing of the material. Filament geometries are produced by mechanical forces (applied via piston, pneumatic devices, or screw‐driven extruders) while droplet geometries are obtained by localized heating or pressure increase within a nozzle. Another drop‐on‐demand technique (laser‐induced forward transfer, LIFT) relies on a pulsed laser transferring energy in the form of heat to a water‐based bioink, propelling it in the form of a droplet onto a moving collector plate. Additionally, optical bioprinting methods can be based on localized crosslinking of a photosensitive, cell‐laden resin, either by means of laser scanning (SLA) or digital light projection (DLP) driven by a micromirror array. Inks with low viscosity in the range of 2–20 mPa s have been reported to be processed via inkjet bioprinting. [ 29 ] This method relies on the formation of bioink droplets which are collected onto a plate. The formation of droplets is either controlled by an electrical‐controlled thermal expansion or pressure‐inducing piezoelectric crystal located in the tip of the nozzle. [ 56 ] Since the method permits to process prevalently low‐viscous materials and low cell concentrations to prevent nozzle clogging, it features an excellent viability despite the short exposure of cells to potentially harmful high temperatures or pressure while forming the droplet. Furthermore, the high speed of fabrication is beneficial for a high survivability of the cells. [ 16, 27 ] Besides early efforts of facilitating inkjet printing for hippocampal, cortical, and motor neuron bioprinting, [ 57 ] only a few researchers have picked up the method for neural applications. Recent reports also described the successful printing of several adult rat CNS‐derived cells. [ 26 ] Although the viability of the printed retinal ganglion neurons and glia decreased to 69% after printing, the cells preserved their phenotype and showed neurite outgrowth, a promising initial step for the establishment of neural constructs. The most commonly used nozzle‐based printing method is extrusion‐based bioprinting. This technique builds their constructs by using a continuous filament or strand of cell‐laden polymers. Inside a reservoir, the flow of the material is induced either by a turning screw, a piston, or on pneumatic actuation. [ 27, 52, 56 ] In contrast to droplet‐based approaches, extrusion bioprinting permits us to work with higher‐viscosity materials, [ 13 ] facilitating printing with high shape fidelity. Paired with a wide range of materials, precise control of crosslinking mechanisms, and tuning of the rheological properties of the ink, this method has emerged as a dominant method of choice when it comes to bioprinting, also thanks to its accessible hardware and the presence of several open source, low‐cost bioprinting devices. [ 37, 38 ] The main drawbacks of this technology are related to its low fabrication speed combined with an inverse relationship between resolution and cell viability. [ 27 ] The opportunities and limitations of extrusion bioprinting for neural application were showcased by Joung et al. [ 32 ] In the pursuit of treating patients suffering from spinal cord injury with a functional 3D‐bioprinted transplantable model, the authors struggled with a limited printing time window. Bioprinting under air exposure dried out the neural progenitor cells within 30 min, resulting in a 100% death rate. Hence, for optimal viability outcome, the constructs had to be printed within 15 min and subsequently submerged in medium. [ 32 ] In light of such limitations, latest advances in using “embedded” extrusion bioprinting may push the boundaries for neural bioprinting. [ 37, 38, 55 ] In such a scenario, the bioink is extruded directly into a supportive bath made of a defined hydrogel microstructure, which behaves as a plastic fluid such as gelatine and alginate. [ 38, 59 ] This approach prevents the collapse of 3D‐printed structure, then allowing us to use bioinks based on collagen and fibrin blends with alginate, which can be processed as a solution and crosslinked in the bath forming gels with an elastic modulus of ≈50 kPa. [ 38 ] This approach resulted in an improved cellular viability (i. e. , 86% for mouse embryonic stem cells (mESCs)). [ 37 ] Second, embedded bioprinting techniques would avoid potential risks of dehydration that could occur during bioprinting of large constructs, in the case long printing time windows are necessary. It should be noted, however, that such a potential challenge could also be readily overcome in conventional bioprinting setups (i. e. , by integrating humidity control features in the printing hardware and environment). 2. 2 Optic‐Based Techniques Although nozzle‐based techniques are still considered the most common bioprinting techniques, photon‐ or optical‐based bioprinting has steadily been gaining momentum over the past decade as an alternative bioprinting approach. The main feature of these bioprinting approaches is the use of laser or projected light for a controlled deposition of cell‐enriched polymers (laser‐induced forward transfer (LIFT)) or the spatially selective curing of photosensitive resins by either two‐photon effect, a guided single light beam (stereolithography (SLA)), or projection of the entire pattern at once (digital light pattering (DLP)) per layer (Figure 2 ). [ 27 ] Analogous to inkjet printing, cell‐laden droplets of minute volume can be produced via LIFT. [ 60 ] In this method, a focusing optical system directs a laser beam onto a donor ribbon (glass slide) coated with an energy‐absorbing layer (EAL) as well as bioink. The absorbed laser energy will induce the formation of a water vapor bubble at the EAL–bioink interface that will propel a droplet of bioink onto a collecting plate, which moves in the x – y plane. Recently, Koch et al. showed the potential of LIFT to print human iPSCs (hiPSCs) [ 40 ] in various biomaterials (i. e. , hyaluronic acid). While confirming that hiPSCs are sensitive to environmental conditions as well, the authors could exclude that the laser‐printing procedure is negatively affecting the pluripotent state or viability. By utilizing a laser beam, the relative size of the droplets can reach the order of the picoliters, and thereby has a higher resolution in contrast to nozzle‐based approaches, besides avoiding risks for cell viability that could be given by shear forces. Despite these advantages, the hardware for LIFT has so far been generally expensive, which limits the diffusion of this technology across different research centers. [ 60 ] Different from LIFT, SLA‐based approaches feature a focused laser beam used to directly crosslink photosensitive, hydrogel‐based resins [ 41 ] like gelatin methacryloyl (GelMA) [ 61 ] or SilkMA. [ 62 ] Usually these resins crosslink in the spectral range of UV (365 nm) to visible light [ 46a ] and can encapsulate cells within the printed resin layer, under a photoexposure window that does not harm the embedded cells and their genetic material. [ 63 ] Depending on the printer arrangement, the crosslinked pattern is attached to a movable stage, which subsequently lowers or raises to build objects in a layer‐by‐layer fashion. [ 27 ] While the x – y resolution of this approach depends on the laser guidance and beam diameter, the z ‐resolution relies on the step size of the platform. Voxel resolution in the range of 5 μm has been achieved with custom‐made printers. [ 43 ] In a combined approach with polycaprolactone (PCL) and gelatin‐coated electrospun fibers, Lee et al. [ 64 ] printed 3D biomimetic neural tissue equivalents with directed and extended neurite growth from primary cortical neurons. Among guided laser beam technologies, two‐photon polymerization (2PP) printing has not been reported for bioprinting of neurons yet, to the best of our knowledge. On one hand, given that neural cells, particularly primary neurons, are particularly sensitive to environmental stress, photoinitiation‐based toxicity could be a possible limitation. Some initiators used in two‐photon bioprinting have been shown to be potentially internalized by cells and have resulted in cytotoxicity upon light exposure. This challenge can be solved by designing macromolecular initiators that cannot easily overcome the cell membrane, [ 49 ] and such systems could potentially be beneficial also in the bioprinting of structures embedding neurons. Furthermore, this technology may find applications in neuron printing in the coming future, given the unique advantages that could be provided by the ability of 2PP to resolve features in the order of tens of nanometers. [ 42 ] Harnessing such potential, and considering the known ability of neurons to respond to nanoscale elements also in 3D, [ 65 ] may thus help introducing nano‐ and micropatterned geometrical features in the proximity of printed hydrogels embedding cells that could be used to guide neurite outgrowth network formation. [ 65 ] Importantly, topographical features around the cells (e. g. , to guide axonal alignment or cell migration) could be produced either using the two‐photon process not only for printing but also for spatially controlled subtractive manufacturing. A proof of concept of this possibility was showcased by Lunzer et al. [ 66 ] In this study, the fabrication of horseshoe‐shaped microchannels with ≈20 μm diameter in close proximity of adipose‐derived mesenchymal stem cells embedded within a hyaluronic‐acid‐based hydrogel was reported. By modulating the extent of degradation via altering the laser power (30–100 mW), differences in the ability of the cells to invade the produced microchannels were observed. In order to overcome the need of a guided and focused laser beam, the bioprinting field has moved toward the projection of dynamic photomasks or patterns onto resins. [ 61, 67 ] Core component of this DLP approach is the use of digital mirror devices (DMDs), which most standard projectors feature. In principle, the DMD is an array of micromirrors, which are controlled by a computer to either reflect or deflect the light toward the projection plane. Each mirror thereby represents a single pixel in the projected pattern. Combined with a platform moving in the z ‐direction, similarly to what is used in SLA bioprinting, the layers are crosslinked by exposure of adequate light energy. In contrast to SLA or 2PP, projecting entire patterns at once gains the advantage of speed on the cost of losing x – y resolution, depending on the mirror density of the DMD. Industry‐grade DMDs reach resolutions in the range of <50 μm, whereas more cost‐efficient DMDs range up to 100–200 μm. [ 61, 67 ] In combination with easy obtainable light emitting diodes (LEDs) or bulbs with fixed wavelength emission, DLP bioprinting does not necessarily need to involve laser light. To date, only a limited amount of studies applied this technique in neural tissue engineering, even though open source DLP devices [ 67 ] and DLP‐ready and biocompatible polyvinyl alcohol–methyl acrylate [ 61 ] /polyethylene glycol (PEG) [ 47 ] + GelMA bioink formulations have been proposed. Similar to 2PP, scientists have rather used DLP in the context of fabricating scaffolds for neural cell seeding approaches. In particular, PEG‐ and GelMA‐based inks were recently used by Koffler et al. [ 47 ] to create a biomimetic scaffold for spinal cord injury and subsequent seeding of NPCs. Upon engrafting the construct in vivo for 6 months in rats, partially to fully recover of spinal cord functionality and motile function of the animals was observed. [ 47 ] 2. 3 Bioassembly Strategies Other biological fields continue to move into the direction of creating multicellular systems like organoids which match the critical need for a cellular‐mimetic microenvironment. Arguably a fusion of organoid technology together with creating reproducible conditions and complex 3D architectures using 3D bioprinting and patterning technologies will lead the field toward a future of multicellular and geometrical complex neural systems. Right on the edge of culturing 3D organoids or spheroids in a regular dish, the idea of patterning assembled cellular blocks envisions to create complex cellular geometries without the need of long‐lasting scaffolds or hydrogels therefore deemed as “scaffold‐free” approaches. [ 27 ] In comparison with the above‐mentioned bioprinting techniques, bioassembly methods facilitate clustering of blocks with high cell density that can closely recapitulate the in vivo physiological conditions. The idea behind this approach ranges from precisely positioning spheroids in needle arrays [ 68 ] to infusing compact nonfused spheroid baths with sacrificial ink. [ 69 ] In the latter case, dense cell‐made slurries (≈10 8 cells mL −1 ) display a viscoplastic fluid behavior, and can act as a support medium within which an extrusion bioprinter can deposit a sacrificial ink consisting of gelatine. Once heated to 37 °C the sacrificial ink can be liquefied, leaving vascular‐like channels within the dense cell assembly that can be used to actively perfuse nutrients. [ 69 ] Among several cell‐based structures, this method was applied also to print vessels within assembled cerebral organoids, potentially offering new opportunities to increase the size of in vitro brain models. Preformed assembly blocks with high neuron content have also been produced in the form of fibrous units produced by coaxial extrusion, [ 70 ] elongated single neuron cultures in micromolds produced by soft lithography, [ 71 ] or more complex cell patterns produced within millimeter‐sized molds. [ 72 ] These basic elements are assembled together into selected geometries via micromanipulation or solely by cell self‐assembly, forming larger electrophysiologically functional networks that, although used so far only for in vivo transplantation, [ 73 ] may offer a powerful tool for in vitro disease studies. 2. 4 Field‐Based Cell Assembly and Bioprinting Strategies While bioassembly approaches take advantage of the self‐organization capacity of cells and organoids, other techniques artificially mediate cell aggregation making use of exogenously applied force fields based on magnetic, acoustic, and optical manipulation of cells, [ 74 ] materials, or growth factors. The potential of field‐based approaches in neuron assembly was initially demonstrated using an acoustic sorting method termed bioacoustic levitation (BAL), applied to fabricate patterns that resemble the cerebral layering. [ 75 ] HUES54‐derived neuronal progenitor cells were mixed within fibrin and subsequently exposed to ultrasonic acoustic standing waves. In dependence of the acoustic frequency, the method creates pressure fields, which manipulate the position of the cells within the gel. Within seconds this will lead to a sorting into a layered fashion, which upon culturing over a period of days (7 days) leads to intra‐ and interlayered neurite elongations. [ 75 ] Magnetic field can potentially also be utilized, provided that magnetic nanoparticles or magnetosensitive ions (i. e. , Gd 3+ in solution) are supplied to a cell suspension. [ 76 ] More recently, the use of optical fields to sculpt in 3D cell‐laden hydrogels in a layerless fashion has led to the development of volumetric bioprinting (VBP). [ 77 ] Applying the principles of optical tomography in reverse, this technique casts a series of tomographic projections onto a rotating cell‐laden bioink. The cumulative light dose given by the convergence of the projections in specific voxels within the volume results in the polymerization of large cell‐laden constructs within seconds. [ 77 ] Such a rapid printing approach could solve current limitations in neuron bioprinting given by the long processing time experienced in extrusion printing by these labile cells. [ 32 ] 3 Bioinks, Biomaterial Inks, and Cell Types for the Biofabrication of Neural Pathways 3D culture environments are a fundamental component for establishing models of multiple pathways and systems within the CNS that accurately capture the complex interactions between different neurons and stromal cell subtypes. As the first step, accurately designed biomaterials are needed to successfully guide cell response and facilitate differentiation into fully mature neurons and to promote their subsequent long‐term survival. Biofabrication technologies allow controlling the spatial patterning of multiple cell types and materials through automated processes, including bioprinting and bioassembly. [ 53 ] While cells are fundamental components of any biofabrication strategy, the building blocks in bioprinting are either carrier solutions or materials embedding cells, termed bioinks, or cell‐free biomaterial inks, on which living cells are then incorporated post printing. [ 78 ] Despite the high potential of bioprinting in capturing salient features of the 3D architecture of neural systems, the field of bioprinting of brain and CNS structures is comparatively young, compared to that of printing other tissues like skin, [ 79 ] cartilage, [ 80 ] bone, [ 81 ] vascular structures, [ 82 ] and cardiac muscle, [ 83 ] among others. Nonetheless, several notable examples of neural tissue inks have been reported in the literature, highlighting the challenges in identifying suitable biomaterials for bioprinting neurons and NPCs and their supporting milieu. Additionally, important lessons can be derived from neural models and implantable grafts produced through more conventional tissue engineering technologies. Based on such background information, this section will highlight the most recent advances in bioinks/biomaterial inks ( Table 2 ) for biofabrication of neural networks, as well as key biomaterial‐based platforms and neural cell types that could lead to significant advances once implemented in automated biofabrication processes. Table 2 Materials used as bioink components for the biofabrication of functional constructs capturing CNS functions Main ink components Printing method Main findings and neuron functions observed Ref. Cell culture media Inkjet Preserved high viability postprinting and ability to form neurite outgrowth comparable to nonprinted cells [ 26 ] Dulbecco's phosphate‐buffered saline solution and fibrin Inkjet Preserved viability post‐printing. Healthy electrophysiological activity, as measured with patch‐clamp tests. [ 57b ] Polyethylene glycol diacrylate (PEGDA) and gelatin methacryloyl (GelMA) Digital light projection Cells were seeded onto printed scaffolds, fabricated to mimic the geometry of the rat and human spinal cord Cell‐laden scaffolds implanted in rodents showed axonal regeneration and partially restored impaired locomotor functions [ 47 ] Carboxymethyl‐chitosan and agarose blend Extrusion Printed cells were able to differentiate into both neurons and neuroglia. Formation of synaptic contacts in 3D [ 95 ] RGD‐modified gellan gum Coaxial extrusion Printing of a multilayered construct, with cells in the first and last layers. Neurons sprout processes through the middle layers, mimicking the brain cortex [ 17 ] GelMA Extrusion A compartment with glioblastoma multiforme (GBM) cells was encased in a printed macrophage‐laden gel. GBM cells recruit macrophages and trigger their differentiation to tumor‐associated macrophages. Co‐culture boosts the invasion of GMB in the surrounding gel [ 100 ] Dopamine‐functionalized GelMA Extrusion Dopamine functionalization did not boost proliferation. Enhanced differentiation of NSCs into maturing neurons [ 102 ] Fibrinogen, RGD–alginate, hyaluronan blend Extrusion Alignment of Schwan cells along the main axis of the bioprinted filament, via shear‐induced alignment of fibrin nanofibers [ 103 ] gelatin‐fibrin blend, Matrigel Alginate/methyl‐cellulose blend (supporting) Extrusion Porous channels printed with the supportive ink and filled with Matrigel to form a pattern of alternated segments containing either NPCs or OPCs. Axonal sprouting and NPC maturation along the channel. No observed OPC maturation or axon myelination [ 32 ] Decellularized brain ECM Extrusion Observed insurgence of chemoradiation and temozolomide resistance in cells within bioprinted cultures [ 107 ] Silk fibroin Extrusion in suspended nanoclay bath Contextual differentiation of neuronal cells and myoblasts. Formation of synaptic contact with acetylcholine and glutamic acid stimulation of human myocytes [ 104 ] John Wiley & Sons, Ltd. 3. 1 Low Moduli Matrices, High Shape Fidelity Printing, and the Shape‐to‐Function Relationship Among the many classes of biomaterials investigated to act as artificial surrogates of the extracellular matrix, hydrogels are a prevalent choice as bioinks for biofabrication, due to their water‐rich environment suitable for cell embedding and for the subsequent postprinting culture. Importantly, hydrogels permit the efficient exchange of nutrients and catabolites to sustain the metabolic demands of neurons. Thanks to their high water content, hydrogels can also be designed to closely resemble the extracellular environment of the brain and its typically low mechanical properties (0. 1–2 kPa). Healthy human brain displays shear stiffness values between 2 and 4 kPa, [ 84 ] and such a value can decrease, together with a loss of viscoelasticity, as hallmark of many neurodegenerative conditions, such as AD and PD, multiple sclerosis, and amyotrophic lateral sclerosis diseases. [ 85 ] Additionally, mechanobiological response to matrix stiffness is a powerful regulator of both progenitor and differentiated cells in the nervous system, [ 86 ] and hydrogels with elastic moduli as low as in the range of 10–100 Pa have been used in vitro to facilitate neuronal cell sprouting. [ 87 ] Hence, using hydrogels able to reproduce such stiffness ranges is important, not only to model the ECM interactions in healthy and pathological conditions, but also to facilitate cell migration, axonal sprouting, and growth, interneuron and neuron–stromal cell interactions. These characteristics are paramount in defining the functionality of brain circuits, and that would otherwise be hampered in stiffer, highly crosslinked hydrogels. Such design requirement renders printing 3D structures with high shape fidelity particularly challenging. [ 13, 88 ] Ideal hydrogels for biofabrication should rapidly gelate into shape stable, stiff structures postprinting. However, soft hydrogel structures, beneficial for 3D cell culture, can easily deform postprinting, either under the effect of gravity or flow due to surface tension and poor elastic properties. [ 89 ] Advances in biofabrication have brought forward new methods to tune the rheological properties of soft inks or to modify the printing environment to ensure high shape fidelity (i. e. , by printing with microfluidics printheads [ 90 ] or by extrusion in supporting baths, [ 55a ] among other strategies), and such methods, of interest also for brain tissue bioprinting, have been extensively reviewed elsewhere, together with key strategies to facilitate the maturation of bioprinted constructs into functional tissues. [ 22 ] Finally, although mechanical and rheological properties play a paramount role both in guiding cell fate [ 51, 91 ] and determining printability, ideal hydrogels should present biochemical properties and cell adhesive domains relevant for the desired cell subsets to be cultured, as well as permitting the transmission of electrical signals across the biofabricated neural networks. 3. 2 Natural‐Origin Hydrogels Hydrogels derived from extracellular components can readily be recognized by cells and often offer ligands that promote cell attachment, migration, and proliferation. Additionally, they can generally be remodeled and degraded into metabolizable compounds, such as amino acids or carbohydrates. Brain ECM is prevalently composed by proteoglycans and glycosaminoglycans (GAGs) such as chondroitin sulfate and hyaluronan and link proteins, particularly in the perineural nets around neuronal processes and in the interstitial space. [ 92 ] Proteoglycans and GAGs are expressed by glial cells and neurons, and regulate a wide array of phenomena including plasticity and inflammatory processes. [ 93 ] Different from most tissues, the brain ECM has limited collagen content, but basal lamina components including laminins, fibronectin, and collagen type IV are present especially at the interface between the CNS and the blood vessels. [ 92 ] In order to mimic glycan component of GAGs, many natural‐origin hydrogels derived from polysaccharides, often used in tissue engineering, such as alginate, chitosan, cellulose, and agarose, have been used as bioink components for patterning neuronal cell‐laden 3D constructs. Differentiation of hESCs into precursors of midbrain dopamine neurons was achieved in 3D cultures in alginate, [ 94 ] a material commonly used as bioink due to its shear‐thinning behavior and rapid gelation kinetics in the presence of divalent cations, and tested also as the substrate for bioprinting neurons in the context of peripheral nerve repair. [ 94 ] Blend bioinks composed of carboxymethyl chitosan and agarose were used to print hNSCs and supported cell function, as demonstrated by differentiating the cells in situ into neurons and neuroglia. Culture and maturation of the printed system resulted in the establishment of an interconnected neural network, synaptic activity, spontaneous neural activity, as well as a bicuculline‐induced increase in calcium response. [ 95 ] Likewise alginate, carboxymethyl chitosan, and agarose blends were used to print iPSCs and subsequently trigger their neural differentiation. High cell viability and pluripotency maintenance were observed. [ 34a ] Importantly, cells that were initially dispersed homogeneously into the hydrogel matrix migrated throughout the construct over time and formed interconnected networks of aggregates within 9 days of culture. Gellan gum, a polysaccharide of bacterial origin, which undergoes ionic crosslinking in the presence of cations, was modified to carry the fibronectin‐derived integrin ligand arginylglycylaspartic acid and used to encapsulate primary cortical neurons. [ 17 ] Using a coaxial needle, this bioink was printed as core material in the inner bore of the nozzle, while ensheathed in an outer flow of its crosslinker (a solution with cations). Printed filaments were stack in multiple layers to print a layered structure mimicking the brain cortex, in which primary cortical neurons originating from BALB/cArcAusb mice' embryos could project their axons across the different layers. [ 17 ] Besides polysaccharides, other ECM constituents like collagen, laminin, fibrin, and fibronectin can also be used in bioinks as their native composition typically presents functional domains and nanoarchitectures that facilitate adhesion, guidance of neural cell morphology, and proliferation. [ 18, 96 ] Collagen has been used extensively as a coating in 2D cultures as well as a scaffold material for the culture of neurons and glial cells in a 3D environment, [ 97 ] with even reports of extended cultures and prolonged survival of differentiated neurons up to 73 days. [ 97b ] Gelatins, hydrolyzed forms of collagen, are also common alternatives in bioprinting, also due to the ease of production of such material from wastes of the food and leather industries. Given the instability at physiological temperature and rapid degradation of pristine gelatin, crosslinkable versions of this material are required for long‐term cell culture. Photo‐crosslinkable, methacryloyl‐modified gelatin (GelMA) rapidly became widespread cell carriers for bioprinting due to their rapid crosslinking kinetics and versatility in obtaining hydrogels with a wide array of mechanical properties, [ 98 ] and recently, attention is being attracted also by thiol–ene crosslinkable gelatin variants. [ 99 ] In the framework of recapitulating in vitro neural systems, GelMA has been used for modeling the microenvironment of brain tumors, specifically how the interplay between tumor‐associated macrophages and glioblastoma cells promotes tumor invasiveness into a bioprinted mini brain. [ 100 ] Dopamine‐functionalized GelMA was also used to print mouse NSCs as well as human‐derived glioblastoma cells. [ 101 ] Although the functionalized hydrogel did not induce any noticeable effect on cell proliferation, an improved differentiation of the neural cells and enhanced expression of the neural markers β‐III tubulin (TUBB3) and microtubule‐associated protein 2 (MAP2) were observed, together with the formation of an interconnected neural network. [ 102 ] Interestingly, some ECM proteins, such as collagen and fibrin, typically self‐organize into micro‐ and nanoscale fibrillar structures that can facilitate the spreading and elongation of neural processes. In combination with extrusion bioprinting, such materials can be used to hierarchically organize printed cells, guiding their alignment not only by confining them within a printed filament, but also via the (sub)micrometer fibrillary elements that build up the bulk of the hydrogel. This concept was recently exemplified via printing Schwan cells in fibrin. [ 103 ] Controlling the printhead velocity and extrusion pressure, it is even possible to align such ECM fibers along the printing direction, which, in turn, align the printed cells within a preferential direction, [ 103 ] a condition desirable to create defined pathways within engineered neural networks. In addition to native and brain‐derived components, ECM proteins of nonmammalian origin such as silk fibroin (SF) have been used as successful bioink component to print neural cell lines. [ 104 ] Finally, despite such availability of different natural materials to establish fully chemically defined bioinks, complex mixtures of ECM components, such as Matrigel, still remain a preferred culture substrate for neurons. 3D cultures in Matrigel were shown to support long‐term survival of neural cells up to 5 months [ 105 ] as well as to permit the culture of multicellular structures like midbrain‐mimetic organoids. [ 106 ] Importantly, in a bioprinted gelatin–fibrin structure with combined iPSC‐derived cells, ventral NPCs and oligodendrocytes, maturation, and axonal sprouting along printed channels was observed only when Matrigel was added to the bioink, underscoring the impact of the diverse factors that compose such materials. [ 32 ] However, since Matrigel is derived from the basal membrane of mouse sarcoma, its use results in large batch‐to‐batch variability and often inconsistencies in experiments performed across different batches, questioning its efficacy and pharmacological relevance as substrate to generate reliable 3D models for studying neurological diseases. Alternatively, CNS decellularized extracellular matrix extracts could also be obtained from human or animal cadaveric donor tissues, reconstituted to form printable hydrogels. Such a type of bioink has recently been proposed to culture glioblastoma cells from human tissue biopsies in order to perform drug efficacy tests to identify the best treatment for each donor, establishing a bioprinted platform for personalized medicine. [ 107 ] 3. 3 Synthetic and Hybrid Hydrogels Given the inherent batch variability within natural origin materials, different types of synthetic, hydrogel‐forming polymers have been studied as 3D culture platforms and bioinks for neural tissue engineering. These are often used in combination with biologically derived cues such as ECM‐derived proteins, GAGs, growth factors, or biofunctional peptides to form natural/synthetic hybrid systems and to improve their biological performance and promote the bioactivity of the embedded cells. A key advantage of fully chemically defined systems and of highly tunable synthetic materials is the ability to accurately control the physicochemical properties of the hydrogel together with the controlled introduction of bioactive cues to enable the rational design of engineered matrices to guide stem cell differentiation. [ 108 ] Despite these advantages, a limited number of synthetic materials have been used for brain tissue bioprinting, and most reports describe more conventional tissue engineering strategies. Nonetheless, important lessons can be learnt from such studies. Amidst synthetic hydrogels, systems based on polyethylene glycol and its derivatives remain among the most explored. For instance, PEG–hyaluronan hydrogels, modified with matrix metalloproteinases cleavable crosslinkers, have been used to study the response of glioblastoma cell lines to increasing values of stiffness of artificial extracellular matrices, [ 109 ] and star‐shaped PEG–heparin hydrogels modified with RGD peptides have successfully been employed to culture primary nerve cells as well as NSCs. [ 110 ] Tyramine‐modified polyvinyl alcohol, modified with silk sericin and gelatin to improve its biofunctionality, was used successfully to culture a neuronal cell line (PC12) and support the formation and outgrowth of dendritic processes. [ 111 ] Polyurethane‐based hydrogels have been proposed as injectable systems for treating traumatic brain injuries (as tested in vivo in a zebrafish model), and the possibility of bioprinting NSCs within such materials was proven. [ 112 ] Specific classes of synthetic polymers have also been studied due to their excellent conductive properties, such as the case of polypyrroles (PPy). Besides excellent electrical properties, PPy have favorable cell and tissue compatibility. However, due to poor solubility and degradation profile, the use of PPy requires their combination with other materials, such as silk fibroin. [ 113 ] Johnson et al. integrated electrical, topographical, and chemical cues into a tissue scaffold in order to promote neuron regeneration. [ 114 ] A PPy‐coated SF (PPy/SF) conductive composite scaffold was fabricated with an electrical conductivity of 1 × 10 −5 –1 × 10 −3 S cm −1, nanoscale fibers, and without cytotoxic properties, allowing the culture of Schwann cells. [ 115 ] Synthetic peptides have also been investigated as biomimetic hydrogel‐forming materials, as these can be designed in their entirety to embed, together with cells, bioactive moieties or domains able to guide the differentiation and maturation of neural cell subtypes. Few peptide‐based bioinks are currently available for bioprinting [ 116 ] although limited work has been performed in the field of neural tissue bioprinting, and most data on peptide‐based hydrogels and neurons are based on conventional tissue engineering strategies. Among notable examples, amphiphilic peptides are able to self‐assemble into nanofibrillar hydrogel structures upon exposure to physiological temperature. These hydrogels show domains that can capture specific growth factors or even mimic the bioactivity of such potent biomolecules that have been proposed for neural tissue repair. [ 117 ] Another example is Puramatrix, a commercially available synthetic matrix based on the acetyl—(Arg—Ala—Asp—Ala)₄—CONH₂ peptide hydrosol. Under physiological salt conditions, it self‐assembles into a 3D hydrogel with a nanometer scale fibrous structure and has been shown to support dorsal root ganglia outgrowth in a 3D culture system combined with a PEG‐based hydrogel. [ 118 ] 3. 4 Neural Cell Types Used in Bioprinting Approaches Several cell types have been broadly used for neural tissue printing such as primary cells, immortalized cell lines, and iPSCs ( Table 3 ). Depending on what purpose researchers wish to serve, it is important to consider which cell type should be used. Table 3 Main neural cell types used in bioprinting approaches Category Starting cell type during printing Terminal neural phenotype obtained Ref. Primary cells Mouse cortical neurons Neurons (TUBB3+); Astrocytes (GFAP+) [ 17 ] Rat retinal ganglion cells Neurons (TUBB3+); Glial cells (Vimentin+) [ 26 ] Rat hippocampal and cortical cells Neurons (TUBB3+) [ 57b ] Mouse NSCs Neurons (TUBB3+) [ 102 ] Rat Schwann cells Schwann cells (S100b+) [ 103, 115 ] Mouse NSCs Neurons (TUBB3+); Astrocytes (GFAP+) [ 112 ] Rat astrocytes and neurons Neurons (MAP2+); Astrocytes (GFAP+) [ 124 ] Porcine Schwann cells Schwann cells (S100b+) [ 125 ] Rat NSCs Astrocytes (GFAP+) [ 126 ] Mouse NSCs Labeled with PKH26 dye [ 112 ] Rat NPCs Neurons (MAP2+); Astrocytes (GFAP+); Oligodendrocytes (Olig2+); Schwann cells (S100b+) [ 47 ] Rat superior cervical ganglia (SCG) sensory neurons and hippocampal neurons Neurons (Tau+) [ 127 ] Rat NSCs Neurons (NF‐H+); Astrocytes (GFAP+) [ 128 ] Established cell lines Mouse NSC (C17. 2) Morphology under a bright field microscope [ 18b ] Human NPCs (NT2) Neurons (TUBB3+) [ 57b ] Mouse NSCs (NE‐4C) Neurons (TUBB3+) [ 64 ] Human NSCs (ReNcell CX) Neurons (TUBB3+); GABAergic neurons (TUBB3+/GABA+/GAD+); Oligodendrocytes (OLIGO2+); Astrocytes (GFAP+) [ 95 ] Mouse glioblastoma (GL261) Glioblastoma (GFAP+/Chil1+) [ 129 ] Rat Schwann cells (S16Y); Rat neuronal cell line (PC‐12); Human glioblastoma (D54‐MG) Not mentioned [ 101 ] Human glioblastoma (U‐87 MG) Glioblastoma (F‐actin+) [ 107 ] Human NPCs (ReNcell VM) Neurons (MAP2+); Astrocytes (GFAP+) [ 130 ] Mouse neuroblastoma (NG108‐15) Morphology under a bright field microscope [ 125 ] Rat neuronal cell line (PC‐12) Neurons (TUBB3+) [ 131, 132 ] Human glioma (U87) Neurons (TUBB3+) [ 133 ] Mouse NPCs (NE‐4C) Neurons (TUBB3+); Astrocytes (GFAP+) [ 134 ] Human neuroblastoma (SH‐SY5Y) SH‐SY5Y (NFH+) [ 135 ] iPSCs hiPSCs Neurons (TUBB3+); [ 40 ] hiPSCs Ventral midbrain dopaminergic neurons (TUBB3+/TH+/FOXA2+/LMX1A+) [ 54a ] hiPSC‐derived organ building blocks (OBBs) Neurons (TUBB3+) [ 69 ] hiPSC‐derived spinal neuronal progenitor cells (sNPCs) Neurons (TUBB3+) [ 32 ] miPSC‐derived oligodendrocyte progenitor cells (OPCs) Oligodendrocytes (labeled with enhanced green fluorescent protein or mCherry) [ 32 ] hiPSCs Neurons (MAP2+); GABA neurons (GABA+) Astrocytes (GFAP+) [ 34a ] hiPSC‐derived neuronal and glial precursor cells Neurons (MAP2+); Astrocytes (GFAP+) [ 35 ] hiPSC‐derived NSCs NSCs (Nestin+/SOX2+/SOX1+/PAX6+) [ 135 ] hiPSC‐derived neural aggregates Neurons (TUBB3+) [ 36 ] hiPSC‐derived NPCs Spinal cord motor neurons (TUBB3+/ChaT+); Astrocytes (GFAP+) [ 136 ] John Wiley & Sons, Ltd. Primary neural cells can keep their authenticity so that any experiment readouts coming out of these cells would mostly reflect on a real in vivo situation. Either adult or embryonic neural cells have been isolated to print. [ 17, 26 ] Drawbacks come as, on the one hand, animal dissection is intrinsically laborious, and these cells can only be passaged in vitro for limited times, thus hampering the 3D printing process, which always requires large amounts of cells. On the other hand, regarding usage of human neural cells, the cell sources are much more restricted and ethical issues are inevitable. The use of immortalized neural cell lines addresses the problem that primary cells can be expanded in vitro for limited time. Some human neural cell lines such as glioma‐/glioblastoma‐derived ones are also established to serve human tissue engineering purposes. Such cell lines can be passaged for a number of times while maintaining the ability to differentiate into neurons and glial cells when exposed to differentiation medium. [ 119 ] However, one general problem about immortalized cells is that immortalization to some extent modifies the chromatin leading to a loss of certain functions of the cells. [ 120 ] This could greatly limit the use of immortalized cells when a lot of functional tests are needed during the experiments. Human ESCs/NSCs/NPCs can be a potential solution to the problems shared by the previous cell types but their availability is still limited. On the other side, the dawn of iPSC technology has made it possible to derive large amounts of human neural cells in vitro. [ 121 ] This has largely bypassed the ethical issues of neural cells obtained from either human embryos or embryonic stem cells. Recent advances in cell reprogramming even allows for direct conversion of somatic cells into neurons or glial cells. [ 122 ] Moreover, with the iPSC approach, patient‐specific neural models can be derived, making the development of patient‐specific models as well as personalized medicine possible. [ 123 ] 4 Future Directions in Advanced Neural In Vitro Modeling The electrophysiological activity of neurons and specifically the transmission of electric and biochemical signals across specific neural networks is intimately intertwined with the geometrical and architectural compositions of these networks, which connect different neural subsets. On one hand, biofabrication technologies can, depending on their resolution, enable to pick and place or dispense the required cells, multicellular building blocks, and cell‐laden hydrogels that are needed to mimic the in vivo relative positioning and interconnectivity of healthy and diseased neural pathways. On the other hand, events and signals that influence NSC fate decisions or the maturation of differentiated cells are affected by a wide array of physicochemical cues that should be applied synergistically to cell printing and assembly in 3D. Combining knowledge from material science, brain and neural cell biology, and electrophysiology with biofabrication technologies can pave the way to the generation of neuronal systems in vitro that can capture the complexity of their native counterparts. 4. 1 Axon and Dendritic Process Guidance Besides key considerations on cell sources and biomaterials as building blocks, the guidance of cellular processes responsible for the interconnectivity of neural networks is paramount in the generation of functional models. Indeed, neurite‐outgrowth factors have demonstrated potential applications in peripheral nerve injuries, providing a basis for their use in 3D neural circuit models. [ 137 ] When building extensive neural networks in vitro, the guidance of axons in an efficient and long‐lasting manner will be essential. [ 138 ] On the other hand, since axon length influences sensitivity and is potentially involved in disease progression, ideally, the axons between nuclei should be of the same length as in physiological conditions. As discussed before, many different factors influence cell differentiation and axon protrusion. Different methods have been employed in an attempt to mimic this environment in vitro including those of chemical, physical, and electrical nature. 4. 1. 1 Chemical Methods Biochemical cues are usually patterned onto a flat surface using microcontact printing (μCP). The potential bioactive molecules can be divided into two classes of neurite‐outgrowth factors: ECM constituents and neurotrophins. In a 3D setting, biologically active molecules could be encapsulated and incorporated into a hydrogel before printing. Tang‐Schomer showed that the delivery of ECM components promoted axon growth. [ 137 ] Laminin and fibronectin showed increased axon growth in comparison to neurotrophic factors like brain‐derived neurotrophic factor, glial cell‐derived neurotrophic factor, nerve growth factor, and neurotrophin‐3 (1 × 10 −6 m ). Neurons have shown extreme sensitivity to subtle changes in the gradient steepness and concentration of insulin‐like growth factor 1. [ 139 ] These results suggest that cells are highly sensitive to these gradients, but not much research using different factors has been done. For an in‐depth review on structural axon guidance, the reader is referred to the study of Seiradake et al. [ 140 ] Another powerful factor in axon guidance is the addition of embedded or tethered biological cues to the permissive 3D scaffold when exogenous differentiation is deemed essential. In the case of dopamine neurons, it has previously been shown that the combination of neurotrophic factors and GAG‐based matrices can improve maturation and neurite outgrowth. [ 141 ] Developmental‐inspired approaches will be probably be the key for future neural pathway models. Specifically, several axon guidance molecules are known to drive neurite outgrowth during CNS development. [ 142 ] Interestingly, in the case of dopamine neurons, it has already been shown that mouse primary dopamine neurons outgrowth can be guided by a focal source of netrin‐1 by interacting with the deleted in colorectal cancer receptor expressed on the dopamine neuron cones. [ 143 ] Therefore, overexpression into target neurons of axon guidance molecules such as netrin‐1 could be used to create a gradient to drive neuronal connections in vitro. [ 144 ] 4. 1. 2 Topography and Mechanical Cues Different physical restraints like microgrooves and micropillars have been used in axon guidance. Axons and dendrites have shown the ability to climb over microgrooves up to 600 nm high. [ 145 ] Micropillars, however, contributed to proliferation and differentiation in vitro. Neurite length seemed to be the longest on the pillars with the smallest interspacing with a width of 2 μm. [ 146 ] Parallel and aligned electrospun fibers do not form a physical restraint but provide contact guidance for cell spreading, migration, and axonal growth, [ 147 ] as shown with polylactide electrospun mats implanted in vivo in mice brains. [ 148 ] Such nanofibrous mats closely mimic the topography of acellular nerve matrix, only with larger diameters of microchannels to promote nerve infiltration. In addition, it is possible to introduce nanotopographical elements like grooves and pores onto electrospun fibers, for instance, via the addition of different combinations of solvents into the polymer solution. This approach has been shown to improve cell alignment along the main direction of the fibers via mechanosensing, as shown by the enhanced shuttling of yes‐associated protein from the cytoplasm to the nucleus. [ 149 ] In addition, the use of nanofibers made from materials displaying a low elastic modulus, such as fibrin, is preferred when aiming to promote neurite outgrowth. [ 150 ] However, electrospinning does not typically allow for the incorporation of cells [ 151 ] and it has most often been used to form 2D culture surfaces. Nonetheless, the potential of electrospun nanoscale structures to align cells can be translated to 3D settings. For this purpose, microfragments of electrospun nanofibers with embedded paramagnetic ferric oxide nanoparticles have been suspended together with cells into a fibrin precursor solution. [ 152 ] In such a setup, the application of a magnetic field can induce alignment of the microfibers, which can be stabilized upon gelation of the fibrin and that in turn was shown to guide the alignment of neurons derived from the dorsal root ganglia. [ 152 ] Interestingly, such an improved cell alignment leads to the propagation of calcium signals across the direction of cell alignment. Similar effects can be obtained using magnetoresponsive rod‐shaped nanogels. [ 153 ] Depending on the hydrogel of choice, these composite systems can also be used as injectable cell delivery vehicles [ 154 ] and therefore may be readily translated as components for bioprintable inks. [ 155 ] 4. 1. 3 Electrical Stimulation and Electroconductive Elements The use of exogenous electric signaling significantly promotes axon growth and creates large‐scale axon alignment in 3D. [ 156 ] Neurite outgrowth has been known to naturally orientate toward the negative pole in a direct current electric field. [ 157 ] A different form of electrostimulation is using an alternating current (AC), attracting the growing tips of axons. Axons orient themselves perpendicular to the adjacent electrode. [ 137 ] Axon length reached 1296. 1 ± 49. 8 μm after 4 days of 2 Hz stimulation. [ 137 ] Although promising, this is still far from biomedical relevant neuronal pathways such as the 4 cm estimated nigrostriatal dopamine axons in the human brain. Indeed, in the CNS, the axons can be far away from the cell bodies, encountering very different kinds of biophysical environments that change over time, making it difficult to mimic such dynamic environment in vitro. [ 145, 158 ] Regardless of the method at hand, axonal dispersion and off‐target reinnervation remain a challenge. [ 159 ] To enhance the control over the electrical properties of the engineered neural networks, additive manufacturing technologies also provide the opportunity to fabricate composite structures embedding conductive elements. [ 160 ] Conductive polymers, such as polyaniline, poly(3, 4‐ethylenedioxythiophene) (PEDOT) and polystyrene sulfonate (PSS) are often combined with hydrogels suitable for cell culture to create 3D substrates for neuronal cell culture. [ 161 ] For instance, oligoaniline‐doped chitosan was successfully used to induce the differentiation of olfactory ecto‐mesenchymal stem cells toward dopaminergic neurons, [ 162 ] which are key actors in PD. Likewise, the embedding in biocompatible hydrogels of nanoconductive elements, such as graphene, [ 163 ] black phosphorus, [ 164 ] carbon nanotubes, [ 165 ] and nanoclay, [ 166 ] has been reported. Importantly, conductive bioinks for bioprinting have already been proposed and promisingly tested in vitro and in vivo for skeletal and cardiac muscle cultures. [ 166 ] A possible limitation for the use of nanoconductive materials is linked to the difficulties in their processability. This issue could partially be overcome by using a mix of blended materials such as polyaniline/poly(lactic acid) that showed to be compatible with the generation of electrospun fibers that support neuron outgrowth. [ 167 ] Another challenge in the application of nanoconductive materials is their limited elasticity that hinders their capacity to mimic in vivo mechanical properties. To address this limitation, it is possible to design an elastic material (i. e. , PCL) with interspersed blocks of conductive polymers. [ 168 ] Finally, it has to be considered that conductive polymers and nanomaterials could exert cytotoxic effects on neural cells, and such negative effects can depend on the chemistry, size, and geometry of the material, as shown, for instance, in studies involving specific formulations of carbon nanotubes and graphene. [ 169 ] Overall, the combination of the described approaches with the increasing work reported in the literature on injectable conductive hydrogels for neuron embedding pave the way toward the generation of 3D‐patterned neural networks with enhanced electric signal transmission. 4. 2 Toward Modeling Neural Environment Complexity Though representing as one of the most promising methods in neural tissue engineering, bioprinting approaches still need to be adapted to implement several aspects that contribute to neural complexity. Glial cells such as astrocytes and microglia are important elements in CNS providing support to neurons and maintaining brain homeostasis, which are crucial to generate a neural tissue model that aims to recreate the neuronal microenvironment. Printing with NSCs or NPCs normally leads to the generation of a mixed population of neurons and glia if no lineage‐specific patterning cues are included in the culture medium to drive a certain neural fate. [ 32, 34, 95, 112, 124, 130 ] A possible solution to this issue would be a multistep modeling strategy where, in the first place, NSCs/NPCs are printed and differentiated into neurons followed by subsequent printing of astrocytes as well as microglia in the culture system to further support neurons. This approach could allow the combination of different neural types but would not be helpful to also implement different neuronal types that could generate specific neural circuits. For example, a challenging, yet important goal in PD modeling, would be to recreate the in vitro nigrostriatal pathway and the neural circuitry of the basal ganglia, in which it is interconnected ( Figure 3 ). In order to recreate several interconnected neuronal clusters of dopamine (DA), gamma‐aminobutyric acid (GABA), and glutamate neurons, we can conceive to take advantage of cell reprogramming techniques to impose specific neuronal and neural phenotypes by overexpressing key transcription factors (TFs) active during CNS development. [ 170 ] Indeed, using doxycycline‐inducible viral vectors we can hypothesize to first bioprint ESCs/iPSCs transduced with different reprogramming TF cocktails to generate different neuronal or neural phenotypes and then to induce the differentiation simultaneously by adding doxycycline to the medium. The combination of bioprinting and cell reprogramming would have the advantage to simplify the problems due to the sensitivity of specific neural cells to bioprinting, allowing a simultaneous assembly of complex neural systems. Figure 3 The circuitry of the basal ganglia. The nigrostriatal pathway is the main neuronal pathway degenerated in Parkinson's disease and is composed by midbrain substantia nigra (SN) dopamine neurons that project toward GABA medium spiny neurons in the striatum (STR). The nigrostriatal pathway is itself embedded in a broader circuitry: the basal ganglia. Within basal ganglia, each neuronal cluster can exert excitatory (with glutamate neurons), inhibitory (with GABAergic neurons), or modulatory (with dopamine neurons) action. SNc: substantia nigra pars compacta; SNr: substantia nigra pars reticulata; STN: subthalamic nucleus; GPe: globus pallidus external; and GPi: globus pallidus internal. Another challenge is to print vasculatures with a size equivalent to brain capillaries. Under same perfusion volume, capillaries have higher efficiency in nutrient and waste exchange. [ 171 ] Different groups have achieved printing vasculatures with diameters as long as hundreds of micrometers. [ 172 ] Building vasculature on a much smaller scale indeed requires higher resolution of printing processes where the mechanical properties of bioink and the size of the printer nozzle should be finely tuned. Endothelial cells embedded in 3D gel can also result in spontaneous angiogenesis without any intended manipulation. [ 173 ] The generation of capillaries in bioprinting neural tissue models remains a crucial achievement in order to simulate physiological interaction between neural cells and the vascular system. In addition, finding an universal medium to support both neurons and vasculatures within the printed model is also a hard task. Serum interferes with neuron behavior in many ways; thus, a serum‐free environment is favorable for neuron health in the long run, whereas serum is crucial for endothelial cells which form the vasculature. [ 174 ] Besides, other factors such as (vascular endothelial growth factor, fibroblast growth factor, heparin, and hydrocortisone are also necessary for maintaining vasculature functions, thus raising questions whether these factors are toxic to neural cells. Instead of finding a compromise in the universal medium composition, several studies have tried to feed neural cells and endothelial cells separately using each one's own medium. [ 173, 175 ] A possible approach is to create another channel loaded with a serum‐free medium to nourish surrounding neural tissues while keeping the serum‐containing medium in the endothelial tubes to support vasculatures. Finally, another development step that needs to be embedded in advanced in vitro neural models is the possibility of measuring physiological parameters such as the electric activity. Extracellular electrodes are capable of stimulating and recording nerve fibers without disruption and can therefore provide a more physiologically accurate reading. Multiple electrode methods have been used over time; however, the amount of stimulation and recording sites has always been a challenge. Microelectrode arrays (MEAs) have overcome this problem, giving them a significant advantage over traditional methods. It has mostly been used in describing spatiotemporal dynamics. By bioprinting onto MEAs, different neural cells can be stimulated and recorded at the same time. Samhaber et al. present a new method for precision patterning of neurons on MEAs, utilizing μCP combined with a custom‐made device to position patterns on MEAs with high precision. This technique is called accurate positioning microcontact printing (AP‐μCP). Survival of cells on these arrays depends on the preparations. Close contact between the cells and array has to be established. [ 176 ] A limitation of this method is that nutrient and oxygen diffusion is blocked from one side, [ 176 ] but we clearly need to take into consideration how to implement physiological recordings in the next‐generation biofabricated neural systems. 4. 3 Final Remarks Though representing as one of the most promising methods in neural tissue engineering, bioprinting approaches still fall short in several aspects, which remain to be addressed in the coming years. These aspects include first of all implementation of in vitro axon guidance that is crucial to recreate in vitro neuronal pathways rather than randomly connected neurons. Although first studies show the integration of fiber‐guided axonal growth, it still is questionable, if the integration of fibers is in the spatial range of directed neurite growth. Directed neurite growth and thus guided intercellular connectivity would be of particular of interest to improve the reliability of drug discovery platforms for diseases where specific neuronal pathways such as the nigrostriatal one in PD are degenerated. Other key points are the optimization of printing parameters to increase neural cell viability; the identification of bioink compositions, which properly mimic natural brain ECM biomacromolecules to better support the assembly of different neural cells; the incorporation of perfusion systems to achieve long‐term cell survival; and the controlled generation of neural circuitry and systems to monitor their electrical activity. In conclusion, notwithstanding all the described approaches, bioprinting is still a technique within a large portfolio of techniques and cannot per se answer all the needs in the field of in vitro neural modeling. Indeed, it has to be considered that other biological fields continue to move into the direction of creating multicellular systems like organoids matching the critical need for a cellular‐defined microenvironment. Moreover, bioprinting technologies further require scalability and creative combination with additional biofabrication methods as well as low‐cost hardware and easy accessibility to advance in the field of neural applications. Arguably, a fusion of organoid technology together with creating reproducible 3D architectures and conditions could help bioprinting to move toward the generation of complex multicellular neural systems. All these action points will be instrumental to move toward reliable in vitro modeling of human neural tissues, therefore helping to limit the usage of improper animal models and at the same time paving the way for the improvement of biomedical research in the field of neurological diseases. Conflict of Interest The authors declare no conflict of interest.
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10. 1002/adfm. 201910442
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Advanced functional materials
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Controlled Apoptosis of Stromal Cells to Engineer Human Microlivers
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Engineered tissue models comprise a variety of multiplexed ensembles in which combinations of epithelial, stromal, and immune cells give rise to physiologic function. Engineering spatiotemporal control of cell-cell and cell-matrix interactions within these 3D multicellular tissues would represent a significant advance for tissue engineering. In this work, a new method, entitled CAMEO (Controlled Apoptosis in Multicellular tissues for Engineered Organogenesis) enables the non-invasive triggering of controlled apoptosis to eliminate genetically-engineered cells from a pre-established culture. Using this approach, the contribution of stromal cells to the phenotypic stability of primary human hepatocytes is examined. 3D hepatic microtissues, in which fibroblasts can enhance phenotypic stability and accelerate aggregation into spheroids, were found to rely only transiently on fibroblast interaction to support multiple axes of liver function, such as protein secretion and drug detoxification. Due to its modularity, CAMEO has the promise to be readily extendable to other applications that are tied to the complexity of 3D tissue biology, from understanding in vitro organoid models to building artificial tissue grafts.
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10. 1002/adfm. 202000543
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Advanced functional materials
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Bi-layered Tubular Microfiber Scaffolds as Functional Templates for Engineering Human Intestinal Smooth Muscle Tissue
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Designing biomimetic scaffolds with in vivo -like microenvironments using biomaterials is an essential component of successful tissue engineering approaches. The intestinal smooth muscle layers exhibit a complex tubular structure consisting of two concentric muscle layers in which the inner circular layer is orthogonally oriented to the outer longitudinal layer. Here, we present a three-dimensional (3D) bi-layered tubular scaffold based on flexible, mechanically robust and well aligned silk protein microfibers to mimic native human intestinal smooth muscle structure. The scaffolds were seeded with primary human intestinal smooth muscle cells to replicate human intestinal muscle tissues in vitro. Characterization of the tissue constructs revealed good biocompatibility and support for cell alignment and elongation in the different scaffold layers to enhance cell differentiation and functions. Furthermore, the engineered smooth muscle constructs supported oriented neurite outgrowth, a requisite step to achieve functional innervation. These results suggested these microfiber scaffolds as functional templates for in vitro regeneration of human intestinal smooth muscle systems. The scaffolding provides a crucial step toward engineering functional human intestinal tissue in vitro, as well as for the engineering of many other types of smooth muscles in terms of their similar phenotypes. Such utility may lead to a better understanding of smooth muscle associated diseases and treatments.
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No full text available
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10. 1002/adfm. 202000639
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Advanced Functional Materials
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Spatiotemporally Controlled Photoresponsive Hydrogels: Design and Predictive Modeling from Processing through Application
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Abstract Photoresponsive hydrogels (PRHs) are soft materials whose mechanical and chemical properties can be tuned spatially and temporally with relative ease. Both photo‐crosslinkable and photodegradable hydrogels find utility in a range of biomedical applications that require tissue‐like properties or programmable responses. Progress in engineering with PRHs is facilitated by the development of theoretical tools that enable optimization of their photochemistry, polymer matrices, nanofillers, and architecture. This review brings together models and design principles that enable key applications of PRHs in tissue engineering, drug delivery, and soft robotics, and highlights ongoing challenges in both modeling and application.
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No full text available
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10. 1002/adfm. 202000893
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Advanced functional materials
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A Chemically Defined Hydrogel for Human Liver Organoid Culture
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End-stage liver diseases are an increasing health burden, and liver transplantations are currently the only curative treatment option. Due to a lack of donor livers, alternative treatments are urgently needed. Human liver organoids are very promising for regenerative medicine; however, organoids are currently cultured in Matrigel, which is extracted from the extracellular matrix of the Engelbreth-Holm-Swarm mouse sarcoma. Matrigel is poorly defined, suffers from high batch-to-batch variability and is of xenogeneic origin, which limits the clinical application of organoids. Here, a novel hydrogel based on polyisocyanopeptides (PIC) and laminin-111 is described for human liver organoid cultures. PIC is a synthetic polymer that can form a hydrogel with thermosensitive properties, making it easy to handle and very attractive for clinical applications. Organoids in an optimized PIC hydrogel proliferate at rates comparable to those observed with Matrigel; proliferation rates are stiffness-dependent, with lower stiffnesses being optimal for organoid proliferation. Moreover, organoids can be efficiently differentiated toward a hepatocyte-like phenotype with key liver functions. This proliferation and differentiation potential maintain over at least 14 passages. The results indicate that PIC is very promising for human liver organoid culture and has the potential to be used in a variety of clinical applications including cell therapy and tissue engineering.
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1 Introduction The liver is the largest internal organ in the body and is responsible for crucial metabolic functions. Upon acute damage, the liver has a great regenerative capacity and can restore up to 70% of its mass. [ 1 ] However, injury caused by chronic diseases such as viral infections or alcoholic fatty liver disease lead to a gradual decrease in liver function, eventually resulting in end-stage liver disease. [ 2 ] For end-stage liver disease, the only curative treatment option is organ transplantation. [ 3 ] Unfortunately, the demand for transplantable donor livers far exceeds availability, resulting in approximately 20% of patients dying while on the waitlist. [ 4 ] In addition, the experience of waiting for a donor organ is a great burden to the patient and the stress of not knowing if a donor organ will become available in time can negatively impact prognosis. [ 5 ] Alternatives for donor livers are thus urgently needed. [ 6 ] Liver organoids hold great promise for regenerative approaches, for example, cell therapy or as cellular building blocks for liver tissue engineering. [ 7 ] Organoids are 3D miniature versions of their organ of origin, which contain most, if not all, cell types that are present in vivo. [ 8 ] Human liver organoids can be initiated from a variety of cells, including embryonic stem cells, [ 9 ] induced pluripotent stem cells (PSCs), [ 10 ] multipotent adult tissue-resident stem cells, [ 11, 12 ] and primary hepatocytes. [ 13 ] In this paper, we focus on organoids established from leucine-rich repeat-containing G-protein coupled receptor 5 positive (LGR5 + ) adult liver stem cells. These organoids can be expanded infinitely in culture, remain genetically stable, and can be differentiated into the hepatocyte [ 11, 12 ] or cholangiocyte [ 14 ] lineages; as such, they reflect important functional and structural aspects of the liver. Transplantation experiments have shown that organoids were able to engraft into the damaged parenchyma of mice and rats, where they became functional hepatocytes, although the repopulation efficiency remained low. [ 11, 12, 15 ] Another benefit is the ability to genetically modify these organoids in vitro using the CRISPR/Cas system, and in theory, enable the correction of relatively simple genetic defects of a patient prior to autologous cell transplantation. [ 16 ] The current gold standard for organoid culture is critically dependent on the use of Matrigel as a 3D matrix. Matrigel is a gelatinous protein mixture extracted from the extracellular matrix (ECM) of the Engelbreth-Holm-Swarm mouse sarcoma, which is propagated in mice. [ 17 ] Matrigel is widely used because it is extremely bioactive and therefore supports proliferation of a wide variety of organoid types. [ 18 ] Organoids cultured in Matrigel, however, are unsuitable for clinical use due to the murine tumor origin of Matrigel. Another disadvantage of Matrigel is very high (up to 50%) batch-to-batch variations, which may influence reproducibility of organoid experiments. [ 19 ] In this paper, we sought to develop a synthetic, defined and easy-to-handle Matrigel alternative for the expansion and differentiation of human liver organoids. In 2016, Gjorevski et al. published the first landmark paper in the field of synthetic hydrogels for organoid culture. [ 20 ] Murine intestinal organoid formation and differentiation was successfully achieved in an enzymatically-crosslinked polyethylene glycol (PEG)-based hydrogel when modified with laminin or Arg-Gly-Asp (RGD) peptides. This group recently published a preprint on bioRxiv, in which they applied the same PEG-based hydrogel for the culture of murine and human liver organoids. [ 21 ] Moreover, in 2018, Broguiere et al. published a defined (although not synthetic) hydrogel based on a combination of a thrombin-crosslinked fibrin gel, nanocellulose and laminin-111/entactin complex. [ 22 ] The gel was as efficient as Matrigel for the expansion and differentiation of murine intestinal organoids and also seemed to be applicable for the expansion of human organoids from a variety of different organs. However, both PEG and fibrin-based hydrogels must be chemically or enzymatically cross-linked to achieve gelation, which complicates the retrieval of organoids from the hydrogel and may be a disadvantage for certain applications. In this study, we used hydrogels based on polyisocyanopeptides (PIC), a synthetic nonimmunogenic material, for the expansion and differentiation of human liver organoids. [ 23 ] Because PIC-based hydrogels exhibit thermoreversible gelation, they are convenient to work with, provide gentle conditions for organoids during passaging and have a unique range of applications, such as the use of the hydrogel for bioprinting without a necessity for enzymatic or light-induced cross-linking. [ 24 ] PIC is a free-flowing liquid below 16 °C; [ 25 ] when the temperature rises above 16 °C, the liquid becomes a viscous hydrogel within minutes. This flexible behavior facilitates simple cell recovery protocols that merely rely on changing the temperature. It is also possible to modify the PIC backbone with a variety of desired molecules and bioactive epitopes. [ 26 ] Most importantly, PIC is a bioinert material and has already been applied by others in vivo without evoking any adverse immune response. [ 27 ] We therefore expect that culturing organoids in a PIC hydrogel will enable their use for organoid-based clinical applications such as cell therapy and tissue engineering approaches. 2. Results and Discussion 2. 1 An Optimized PIC Hydrogel Supports Liver Organoid Expansion We first developed a PIC hydrogel that supports organoid formation and proliferation. PIC itself does not contain any bioactive components to support cell attachment or induce proliferation. Accordingly, we observed marginal organoid formation, with little to no proliferation when single organoid cells were seeded in plain PIC and cultured for 7 days in organoid expansion medium (EM) ( Figure 1a ). To stimulate proliferative pathways involved in cell attachment to the ECM, we tested a PIC hydrogel that contained covalently-attached RGD sequences, which can be recognized by cellular integrins. [ 20 ] However, the incorporation of RGD motifs in PIC was not sufficient to induce organoid proliferation ( Figure 1a ). Our findings contrast with the recently published preprint by Sorrentino et al. , where the addition of RGD motifs to a PEG hydrogel seemed sufficient for human liver organoid culture; our dissimilar results may be explained by a difference in RGD concentrations. In our commercially purchased hydrogel, RGD peptides are covalently coupled to the PIC at a concentration of ≈0. 2 m M, whereas concentrations of 1–2 m M have been used in earlier publications. [ 21 ] We continued to add 3 mg mL −1 lamininentactin complex (LEC) to the plain PIC gel, which resulted in efficient organoid formation and proliferation that seemed comparable to the Matrigel controls ( Figure 1a ). LEC is the main component of Matrigel and has previously been shown to promote formation and proliferation of murine intestinal organoids. [ 22 ] We then examined the effects of the mechanical properties of PIC hydrogels on liver organoid proliferation. We used PIC at two different molecular weights, 1 kDa PIC (1k PIC) and 5 kDa PIC (5k PIC), and modulated the hydrogels by varying PIC concentrations, resulting in a higher stiffness and lower porosity with increasing concentrations. Light microscopy pictures and Alamar blue assays (ABAs) confirmed that lower PIC concentrations led to a significantly increased organoid proliferation ( Figure 1b ). After determining the optimal PIC concentration (1 mg mL −1 ), we continued to optimize the concentration of LEC in PIC. In both the 1k PIC and 5k PIC hydrogel backbones, we observed a concentration-dependent increase in organoid proliferation ( Figure 1c ), with highest proliferation rates at 3 mg mL −1 LEC in a 1k PIC hydrogel ( Figure 1d, e ). Taken together, a 1k PIC hydrogel at a concentration of 1 mg mL −1 PIC, supplemented with 3 mg mL −1 LEC, provided an environment that optimally supported organoid formation and proliferation with an efficiency comparable to Matrigel. 2. 2. Soft PIC Hydrogels Enhance Organoid Expansion To gain more insights into the different PIC hydrogels, we analyzed their mechanical properties. As the gelation of both Matrigel and PIC hydrogels is temperature-dependent, we placed the different hydrogel premixes on a temperature-controlled plate and followed their gelation in time. The plate was heated from 4 to 37 °C at a speed of +7 °C min −1 and then maintained at a temperature of 37 °C for 10 min. Our analysis of storage modulus G′ showed that for all premixes, gelation started during the heating period and reached a plateau after 10 min at 37 °C ( Figure 2a ). We observed that the stiffness of the 5k PIC (≈83 Pa) was comparable to that of Matrigel (≈71 Pa), whereas the stiffness of all other PIC hydrogels (with and without LEC) was lower. The 5k PIC hydrogels demonstrated a higher stiffness (≈83 Pa) compared to 1k PIC hydrogels (≈18 Pa) of the same concentration and the addition of LEC resulted in a decreased stiffness for both the 1k (≈12 Pa) and 5k (≈38 Pa) PIC hydrogels ( Figure 2a, b ), which was most likely caused by sterical interference of PIC network formation by the large laminin molecules. By contrast, increasing LEC concentrations did not significantly affect hydrogel stiffnesses of the same PIC concentration, suggesting that the LEC concentration-dependent organoid proliferation was caused by biological rather than mechanical effects (unpublished data). As such, the softest PIC hydrogel (1k PIC hydrogel at a concentration of 1 mg mL −1 PIC, supplemented with 3 mg mL −1 LEC) best supported liver organoid expansion. This corresponds with our observations that organoids proliferate better in 70% Matrigel (diluted with culture media and thus, softer) compared to 100% Matrigel (unpublished data). Remarkably, Sorrentino et al. observed that, for their PEG hydrogels, a much higher stiffness of 1. 3 kPa optimally supported colony formation of human liver organoids, whereas a lower stiffness of 0. 3 kPa was much less efficient. [ 21 ] The authors, however, did not quantify proliferation of the liver organoids in their various hydrogels. For serial passaging and long-term expansion of organoids in 3D culture, the thermoreversible character of PIC hydrogels is one of the most significant advantages over other hydrogels. Using a rheometer, we compared the thermoreversible character of PIC hydrogels with Matrigel through a controlled temperature change. We first heated the hydrogel premixes from 4 to 37 °C, allowing them to form a hydrogel. Subsequently, we cooled the hydrogels down to 4 °C at a speed of 7 °C min −1 and then kept the samples at 4 °C for another 10 min to see whether they would become liquid again. Our results show that the optimized PIC-LEC hydrogel promoted a faster gel/sol transition than Matrigel ( Figure 2c ). After approximately 6 min of heating, the PIC-LEC hydrogels turned into a gel, which immediately turned into solution again upon cooling ( Figure 2c ). In contrast, Matrigel still had a stiffness of ≈50 Pa after a 15 min cooling period ( Figure 2c ). This corresponds with our observations that PIC hydrogels are more easily manipulated and thereby easier to handle during organoid passaging, giving them a major advantage over Matrigel. We also characterized the dynamic properties of PIC hydrogels after gelation, since this may influence organoid shape and size. We investigated the modulus change of the hydrogels in a frequency range from 10 to 0. 1 rad s −1. Both Matrigel and PIC hydrogels were stable in the measured frequency range and behaved as a viscoelastic solid ( Figure S1, Supporting Information ). Interestingly, the addition of LEC did not influence the dynamic properties of the PIC hydrogels, indicating that the dynamic character of the PIC-LEC hydrogels did not play a key role in supporting organoid expansion. 2. 3. Organoids in PIC Retain a Stem/Progenitor Phenotype and Are Highly Proliferative To further characterize liver organoids in the optimized PIC-LEC hydrogel, we seeded single organoid cells in EM conditions and cultured them for 14 days. Single organoid cells seeded in Matrigel were used as a control. Immunofluorescent analysis showed that the cells in all hydrogels retained an epithelial E-cadherin (ECAD)-positive phenotype and were highly proliferative ( Figure 3a ). To further interrogate this proliferative phenotype, we performed mRNA sequencing on organoids from two different donors that were cultured in 1k PIC-LEC, 5k PIC-LEC, or Matrigel for 14 days. A Pearson’s correlation test showed that all hydrogel conditions (Matrigel, 1k PIC, and 5k PIC) displayed a similar correlation coefficient for each donor ( Figure 3b ). We then analyzed the expression of a few well-known liver stem/progenitor cell markers (LGR5, SOX9, EPCAM, PROM1, AXIN2), proliferation markers (KI67, PCNA), and hepatocyte markers (ABCC2, ALB, CYP3A4), and compared these with gene expression in cryopreserved hepatocytes and human liver tissue. As expected, stem cell markers and proliferation markers were highly expressed in all hydrogel conditions in both donors, whereas hepatocyte markers were low ( Figure 3c ). 2. 4 Organoids in PIC can be Differentiated into Functional Hepatocyte-like Cells It has previously been shown that human liver organoids can be differentiated toward hepatocyte-like cells when cultured in differentiation medium (DM), which contains N- [(3, 5-difluorophenyl)acetyl]- L -alanyl-2-phenyl]glycine-1, 1-dimethylethyl ester (DAPT), fibroblast growth factor 19 (FGF19), dexamethasone and bone morphogenetic protein 7 (BMP-7), and is void of inducers of proliferation, such as Rspondin-1 and forskollin (FSK). [ 12, 28 ] In order to compare the differentiation potential of human liver organoids cultured in PIC-LEC to Matrigel, we differentiated organoids toward hepatocyte-like cells. Light microscopy demonstrated that both Matrigel and PIC-LEC organoids acquired a more dense and dark morphology at day 8 of differentiation compared to expansion conditions ( Figure 4a ). Analysis of the organoids by quantitative real-time PCR (qRT-PCR) after 8 days of differentiation confirmed that the stem cell marker LGR5 was downregulated in both hydrogels ( Figure 4b ). In contrast, several hepatocyte markers such as cytochrome p450 3A4 (CYP3A4), albumin (ALB), and multidrug resistance-associated protein 2 (MRP2) were upregulated and reached comparable levels in PIC-LEC and Matrigel ( Figure 4b ). We also assessed the functionality of the differentiated organoids in both hydrogels with several assays. Important functions of hepatocytes include the production of serum proteins, vectorial uptake, and secretion of several compounds, and urea cycle activity. [ 29 ] To determine the production of serum proteins, we quantified the intracellular ALB concentration of the organoids at day 8 of differentiation. We measured comparable concentrations of ALB in PIC-LEC and Matrigel ( Figure 4c ). To determine vectorial transmembrane transport, we exposed organoids in both hydrogels to rhodamine123 (Rh123), a fluorescent compound that is actively secreted from the apical membrane of hepatocytes by the transporter multidrug resistance gene 1 (MDR1). [ 30 ] Organoids in both hydrogels accumulated fluorescence inside their lumens ( Figure 4d ). To confirm that this accumulation was MDR1-specific, organoids were pretreated with the competitive MDR1 inhibitor verapamil. This resulted in an accumulation of Rh123 in the cytoplasm of the cells, whereas no fluorescence was observed in the lumen of the organoids, confirming the MDR1-specific transport of Rh123. Finally, we tested organoids in both hydrogels for their capacity to eliminate ammonium from the media, a measure for urea cycle activity. Ammonia reacts with α - ketoglutaric acid to form L -glutamate, which is catalyzed by L -glutamate dehydrogenase (GLDH). [ 31 ] The intracellular concentration of GLDH was comparable in organoids from both hydrogels ( Figure 4e ), along with ammonia elimination from the media ( Figure 4f ). 2. 5 Long-Term Expansion of Organoids in PIC One of the main advantages of organoid cultures is that they can be serially passaged, allowing for seemingly unlimited expansion. This continued expansion is fueled by the LGR5 + stem cells in liver organoid cultures. [ 32 ] In order to test if this stem cell phenotype and organoid proliferation capacity can be retained in PIC-LEC over several passages, we cultured human liver organoids from two donors for 14 passages with weekly passaging. Light microscopy images confirmed that the organoids retained their cystic morphology through all passages and remained highly comparable in PIC-LEC and Matrigel ( Figure 5a ). qRT-PCR showed that the expression of LGR5 in organoids cultured in PIC-LEC or Matrigel was stable over the course of 14 weeks (4 months) ( Figure 5b ). We then assessed the differentiation potential of organoids that had been cultured in PIC-LEC for 2, 6, and 10 passages (P2, P6, and P10). After those passages in EM, in either PIC-LEC or Matrigel, organoids from two donors were differentiated in DM for 8 days and subsequently analyzed by mRNA sequencing (P6) and qRT-PCR (P2, P6, and P10). At P6, the 100 most significant differentially-expressed genes in DM versus EM displayed an identical pattern in PIC-LEC compared to Matrigel ( Figure 5c ). Gene ontology analysis [ 33 ] of these genes revealed significant enrichment for processes related to liver functions, such as lipid and alcohol metabolism ( Tables S1 and S2, Supporting Information ). Analysis by qRT-PCR confirmed that the gene expression of several hepatocyte markers in differentiated organoids was upregulated compared to their respective EM controls in all hydrogels and that this upregulation was reproducible at P2, P6, and P10 ( Figure 5d ), indicating that organoids retain their differentiation potential over several passages in PIC-LEC. 2. 6 Human Recombinant Laminin-111 Can Substitute LEC The hydrogel we have developed and extensively characterized for human liver organoid culture contains a chemically well-defined PIC component supplemented by a bioactive LEC, and the combination has several advantages over Matrigel. However, the LEC originates from mouse sarcoma and is thus not xenofree and not suitable for clinical applications. Therefore, we examined the possibility of using human recombinant laminin-111 (hrlaminin-111) as a substitute for LEC. The experiments with LEC showed a concentration-dependent proliferation of organoids in LEC at concentrations of 1, 2, and 3 mg mL −1 ( Figure 1c ). We prepared PIC hydrogels with the same concentrations of hrlaminin-111 and seeded single organoid cells in those gels. For two of the three analyzed donors, we observed a comparable concentration-dependent proliferation in the PIC-LEC and PIC-hrlaminin-111 ( Figure 6a ). For the third donor however, proliferation in the PIC-hrlaminin-111 seemed lower compared to PIC-LEC ( Figure S2, Supporting Information ). To further characterize the phenotype and proliferative potential of organoids cultured in PIC-LEC and PIC-hrlaminin-111, we conducted immunofluorescent stainings for ECAD, keratin 19 (K19), Ki67, and PCNA, which confirmed that the cells in all hydrogels retained an epithelial ECAD-positive phenotype and were highly proliferative ( Figure 6b ). These data indicate that, although less efficient, it seems possible to replace LEC by hrlaminin-111 and possibly other human recombinant ECM proteins as the bioactive components in PIC, rendering the PIC completely synthetic and applicable for clinical applications such as cell therapy or tissue engineering. 3 Conclusion In this study, we developed and characterized a novel synthetic hydrogel for the expansion and differentiation of human adult stem cell-derived liver organoids. PIC hydrogel alone was not sufficient to support organoid growth, but PIC mixed with LEC supported organoid formation and proliferation ( Figure 1 ). As no covalent reaction can occur between PIC and LEC in the absence of catalysts, the interaction of PIC and LEC is most probably non-covalent self-assembly, but the exact interaction of PIC and LEC remains to be determined. The optimized PIC-LEC hydrogel supported organoid proliferation at rates that were similar to Matrigel ( Figure 1 ), with lower stiffnesses most favorable for organoid proliferation. We showed that the organoids were highly proliferative in PIC-LEC when cultured in EM ( Figure 3 ) and could be efficiently differentiated toward functional hepatocyte-like cells when cultured in DM ( Figure 4 ). Importantly, the stem cell phenotype and proliferation and differentiation capacity of the organoids could be maintained in PIC-LEC over several passages, enabling their seemingly unlimited expansion and subsequent maturation ( Figure 5 ). Finally, we showed that the LEC in the PIC-LEC gels could be replaced by hrlaminin-111, resulting in a completely synthetic hydrogel for the expansion of human liver organoids. Future studies will have to confirm the suitability of this synthetic hydrogel for differentiation of human liver organoids toward hepatocytes and cholangiocytes. Moreover, hrlaminin-111 might be replaced or complemented by other ECM components to improve the differentiation to either hepatocytes or cholangiocytes. [ 34 ] It will also be interesting to replace full-length ECM proteins by peptides derived from laminin-111 and other ECM components. For PEG hydrogels, it has been shown that high concentrations of covalently-linked A55 peptide (GGFLKYTVSYDI) and AG73 peptide (RKRLQVQLSIRT), which are both located on the α -1 laminin chain, supported intestinal organoid proliferation and survival that approached levels in Matrigel. [ 20 ] Together with the recent preprint by Sorrentino et al. , this is the first chemically-defined and synthetic hydrogel that supports human liver organoid expansion and differentiation at rates comparable to Matrigel. [ 21 ] Both hydrogels, PEG-based and PIC-based, have their own advantages. An enzymatically or chemically crosslinked PEG hydrogel was the first synthetic hydrogel applied for intestinal organoid culture derived from multipotent or PSCs. [ 20, 35 ] ECM components such as fibronectin, laminin-111, collagen IV, hyaluronic acid, and perlecan were added to the hydrogel, and a minimized PEG-RGD hydrogel was sufficient to support organoid formation and proliferation. This hydrogel system was recently adapted to liver organoids. [ 21 ] Proliferation rates of intestinal or liver organoids in the PEG hydrogels were not quantified however, so it remains to be determined how they compare to our PIC hydrogel. In a very elegant approach, the authors also modified the PEG hydrogels to be hydrolytically or protease-degradable, as such mimicking the dynamic character of the in vivo environment. [ 20, 35 ] Similar modifications may be applied to the PIC hydrogels by applying a copper-free click chemistry as previously published. [ 36 ] The most striking difference between the PEG and PIC hydrogels is the mode of gelation. Whereas the chemically- or enzymatically-crosslinked PEG hydrogel is advantageous for applications where a controlled stiffness over a prolonged period of time is necessary, the thermoreversible properties of the PIC hydrogel allow for easier cell recovery during organoid culture and make it advantageous for certain practical purposes such as bioprinting and in vivo cell therapy. [ 37 ] PEG-based hydrogels have already been applied as delivery vehicles in vivo, where they supported localized engraftment of human intestinal organoids derived from PSCs in colonic mucosal wounds and enhanced wound closure. [ 35 ] In order to facilitate gelation in situ, a custom made-device was applied, in which the hydrogel precursor solution and the crosslinking solution only met in the tubing during injection. [ 35 ] PIC hydrogels have also been applied in vivo for wound healing studies and subcutaneous cell transplantations without causing any adverse effects. [ 26, 27 ] The thermo-responsive properties of the PIC hydrogels made in vivo applications very convenient, and the PIC gelated within 1 min upon contact with the warm skin. [ 27 ] These results highlight the great potential for in vivo applications of organoids cultured in PIC. Of note, we recently published a protocol for the large-scale expansion of human liver organoids in suspension culture, [ 28 ] and foresee that a combination of this large-scale expansion method and our well-defined synthetic hydrogel will pave the way for clinical applications of human liver organoids in the near future. [ 38 ] 4 Experimental Section Hydrogel Preparation and Characterization Noviogel (PIC) was purchased from Sopachem (1k-PIC-P, 1k-PIC-RGD, 5k-PIC-P, and 5k-PIC-RGD, followed by catalog numbers: NCN01, NCN03, NCN02, and NCN04, respectively). Stock solutions of 5 mg mL −1 were made by adding 3 mL of advanced DMEM/F12 (AD, Gibco) to each bottle of PIC. To optimize the PIC hydrogels, both 1k and 5k PIC were tested in three concentrations: 1, 1. 75, and 2. 5 mg mL −1. Where indicated, LEC (Corning) was added to each PIC formulation at concentrations of 1, 2, or 3 mg mL −1 to test the dose effects. When LEC was replaced with human recombinant laminin-111 (hrlaminin-111, Biolamina, LN111-050) the same concentrations were used (1, 2, or 3 mg mL −1 ). Working Solutions of Thermoresponsive Materials PIC, LEC, and Matrigel (Corning, 356 237) were divided into aliquots to reduce frequent freeze-thaw cycles and stored at −20 °C. To make hydrogels, materials were thawed and prepared on ice before gelation. Once plated, hydrogels were allowed to solidify for 15–30 min at 37 °C. Rheological measurements were conducted on a discovery HR-2 rheometer (TA instruments) to test the mechanical properties of hydrogels made from PIC, PIC-LEC, or Matrigel. A temperature-controlled plate (20 mm parallel plate, aluminium-105381) was used for all measurements. Settings included a geometry diameter of 20 mm and a measuring gap of 200 μm. For hydrogel formation measurements, 150 μL of each hydrogel premix was deposited onto the plate and conditioned for 2 min at 4 °C. The plate was then heated (strain 1%, angular frequency 1. 0 rad s −1 ) from 4 to 37 °C at a speed of 7 °C min −1, and then continuously measured at 37 °C for another 10 min. Following the hydrogel formation step, frequency sweep tests were carried out and data was collected in the frequency range of 10 rad s −1 to 0. 1 rad s −1. For thermoreversible characterization, hydrogel formation was carried out as described above. After remaining at 37 °C for 10 min, the plate was cooled from 37 to 4 °C at a rate of 7 °C min −1. When the temperature of the plate reached 4 °C, it was held for another 10 min to observe if the hydrogels could become fluidic again. Human Liver Organoid Establishment Liver biopsies were obtained during liver transplantation at the Erasmus Medical Center Rotterdam and in accordance with the ethical standard of the Helsinki Declaration of 1975. Use of the tissue for research purposes was approved by the Medical Ethical Council of the Erasmus Medical Center and informed consent by the liver transplant recipient was given (MEC-2014-060). To establish human liver organoid lines, liver biopsies were cut into small pieces, followed by the enzymatic digestion with type II collagenase (0. 125 mg mL −1, Gibco) and dispase (0. 125 mg mL −1, Gibco) in DMEM GlutaMAX (Gibco) containing 1% fetal calf serum (FCS, Gibco). The supernatant was collected every hour. Tissue digestion followed by supernatant collection was performed three times. Collected single cells were washed in DMEM GlutaMAX (Gibco) containing 1% FCS (Gibco) and centrifuged for 5 min at 1500 rpm. The cells were resuspended in Matrigel (Corning) at a concentration of ≈400 cells per μL. Cells were seeded in droplets (50 μL) in non-attaching 24-well plates (M9312, Greiner, Merck). EM was added after ≈15 min incubation at 37 °C, 5% CO 2. Organoid Expansion and Differentiation in PIC and Matrigel Droplets For human liver organoid expansion, previously defined EM was used. [ 9 ] EM was based on advanced DMEM/F12 (AD, Gibco) containing 1% GlutaMax (Gibco), HEPES (10 m M, Gibco), and 1% penicillin-streptomycin (Gibco, Dublin, Ireland). EM was supplemented with 10% Rspondin-1 conditioned medium (the Rspo1-Fc-expressing cell line was kindly provided by Calvin J. Kuo), 1% N2 supplement (Invitrogen), 2% B27 supplement without vitamin A (Invitrogen), N-acetylcysteine (1. 25 m M, Sigma-Aldrich), nicotinamide (10 m M, Sigma-Aldrich), epidermal growth factor (50 ng mL −1, EGF, Invitrogen), hepatocyte growth factor (25 ng mL −1, HGF, Peprotech), fibroblast growth factor 10 (0. 1 μg mL −1, FGF10, Peprotech), recombinant human (Leu15)-gastrin I (GAS, 10 nM, Sigma-Aldrich), FSK (10 μM, FSK), and A83-01 (5 μM, Tocris Bioscience). During the optimization of PIC hydrogels, single cells were prepared from organoids by trypsinizing with TrypLE Express Enzyme (12604-013, Gibco). Single cells were seeded in droplets (50 μL) at a concentration of 300–1000 cells per μL, in temperature and humidity-balanced 24-well plates (M9312, Greiner, Merck). After seeding, plates were incubated at 37 °C for 15–30 min to facilitate hydrogel gelation and thereafter, EM (500 μL) containing Y-27632 (10 μM, SelleckChem) was added to each well. After three days of culture, Y-27632 (SelleckChem) was no longer used. Medium was refreshed every 2–3 days; organoids were passaged weekly at a 1:3–1:4 split ratio by mechanical dissociation method. All cultures were incubated at 37 °C, 5% CO 2. For liver organoid differentiation, previously-defined differentiation medium (DM) was used. DM was based on advanced DMEM/F12 (AD, Gibco) containing 1% GlutaMax (Gibco), HEPES (10 m M, Gibco), and 1% penicillin-streptomycin (Gibco), and supplemented with 1% N2 supplement (Invitrogen), 2% B27 supplement without vitamin A (Invitrogen), N-acetylcysteine (1. 25 m M, Sigma-Aldrich), EGF (50 ng mL −1, Invitrogen), HGF (25 ng mL −1, Peprotech), FGF19 (100 ng mL −1, FGF19, Peprotech), recombinant human GAS (10 nM, Sigma-Aldrich), A83-01 (500 nM, Tocris Bioscience), BMP-7 (25 ng mL −1, Peprotech), 30 μM dexamethasone (Sigma-Aldrich), and DAPT (10 μM, Selleckchem). Prior to switching to DM from EM, liver organoids were primed in EM containing BMP-7 (25 ng mL −1, Peprotech) for three days. Medium was then switched to DM. Medium was refreshed every 2–3 days for 8 days. Measurement of Cell Proliferation ABA was used to measure the cell proliferation. For ABA, the stock solution (Dal1100, Life Technologies Europe BV) was diluted 1:10 in DMEM/F12 without phenol red (21 041, Gibco), sterilized through a 0. 22 μm filter, and pre-warmed to 37 °C. EM was removed from each well and replaced with pre-warmed ABA solution. Organoids were incubated in ABA solution for 90 min at 37 °C, 5% CO 2. After incubation, ABA solution was transferred to a new 24-well plate (Greiner) after which, fresh pre-warmed EM was added to the old plate. For ABA measurement, Fluoroskan Ascent FL (Thermo Fisher Scientific) was used. The wavelength of excitation and emission were 544 and 577 nm, respectively. ABAs were conducted on days 0, 1, 4, and 7 after single-cell seeding. Data analysis was normalized to day 0; results were graphed as fold-changes using GraphPad Prism (GraphPad Software). Cryopreserved Hepatocyte Culture LiverPool Cryoplateable Hepatocytes (10-donor, mixed gender) were provided by Bioreclamation IVT. Hepatocytes were cultured in a collagen sandwich according to the manufacturer’s instructions, and supplemented with the recommended InVitroGRO CP media (Bioreclamation IVT). For mRNA sequencing, hepatocytes were harvested after 4 h of sandwich culture. RNA Isolation and qRT-PCR The RNeasy Mini Kit (Qiagen, Hilden, Germany) was used to isolate RNA from tissues, hepatocytes, and organoids following the manufacturer’s instructions. RNA quality and quantity was measured with DS-11 spectrophotometer (DeNovix). Complementary DNA (cDNA) was synthesized with the iScript cDNA synthesis kit (Bio-Rad) following the manufacturer’s instructions. qRT-PCR was used to determine relative expression of target genes using validated primers ( Table S3, Supporting Information ) using the SYBR Green method (Bio-Rad). Normalization was carried out using reference genes Hypoxanthine-guanine phosphoribosyltransferase and ribosomal protein L19. Rhodamine123 Transport Assay Liver organoids were differentiated in PIC-LEC or Matrigel droplets for 8 days as previously described. For Rh123 transport assays, organoids were pretreated with DM containing verapamil (10 μM, Sigma-Aldrich) or DMSO for 30 min. Organoids were then removed from PIC-LEC or Matrigel and resuspended in DM containing rh123 (100 μM, Sigma-Aldrich) and incubated at 37 °C for 10 min. Fluorescence was visualized by an EVOS FL cell imaging system (Life Technologies). Ammonium Elimination Assay For ammonium elimination assays, liver organoids were differentiated in PIC-LEC or Matrigel droplets for 8 days as previously described. Organoids were incubated with DM supplemented with NH 4 Cl (2 m M ) for 24 h. After 24 h, media samples were harvested and stored at −20 °C. Afterward, Tryple-Express (Gibco) was added to each well and organoids were trypsinized for cell counting. Cell counts were carried out using the TC20 automated cell counter (Bio-Rad). Viable cells were determined using trypan blue exclusion assay. Ammonium concentrations were measured with the urea/ammonia Assay Kit (Megazyme). As a control, DM containing NH 4 Cl (2 m M ) was incubated for 24 h without cells. Ammonia elimination rates were normalized to live cell numbers. GLDH Expression and Albumin Production To quantify the intracellular levels of GLDH and ALB, liver organoids were differentiated in PIC-LEC or Matrigel droplets for 8 days, as previously described. Organoids were provided with fresh DM 24 h before being lysed in MilliQ water. GLDH and ALB were measured in the cell lysates using a DxC-600 Beckman chemistry analyzer (Beckman Coulter). Values were normalized to total protein concentrations. Microscopy and Immunofluorescence Analysis Imaging of the organoids was performed using an EVOS FL cell imaging system (Life Technologies) and an Olympus BX51 microscope in combination with an Olympus DP73 camera. Detailed information of applied antigen retrieval methods, antibodies, dilutions, and incubation times are listed in Table S4, Supporting Information. Bright field images were taken to track organoid morphology throughout expansion in different hydrogel formulations. Images were also taken to compare the morphology of organoids in EM and DM. For immunofluorescent (IF) staining, organoids were fixed with 4% neutral buffered formalin containing 0. 1% eosin and incubated at 37 °C overnight. Fixed samples were dehydrated and embedded in paraffin or stored in 70% ethanol at 4 °C for up to 1 month; 4 μm thick paraffin sections were prepared for IF staining. To start the IF staining procedure, the paraffin sections were first heated at 62 °C for 15 min and dewaxed by xylene, followed by rehydration in gradient ethanol concentrations from 100% to 70%. Then, sample sections were incubated in antigen retrieval solution for 30 min at 98 °C. After balancing to room temperature, sample sections were treated with NH 4 Cl solution (20 m M ) for 10 min to reduce background autofluorescence and blocked with 10% goat serum for 1 h to avoid non-specific antibody binding. Next, primary antibodies against ECAD, Ki67, PCNA, and K19 were added to the sections and incubated overnight at 4 °C. After being washed with tris-buffered saline containing Tween 20 three times, sample sections were incubated with secondary antibodies (5 μM), including mouse anti-rabbit Alexa Fluor 488 (Molecular Probes), mouse anti-rabbit Alexa Fluor647 (Molecular Probes), rabbit anti-mouse Alexa Fluor488 (Molecular Probes), and rabbit anti-mouse Alexa Fluor647 (Molecular Probes). Nuclei were stained with DAPI (0. 5 μg mL −1, Sigma Aldrich). Whole Transcriptome Sequencing and Analysis For whole transcriptome sequencing, RNA was isolated from organoids using the RNeasy Mini Kit (Qiagen) according to the manufacturer’s instructions. Library preparation, 75 bp single-end sequencing on Illumina NextSeq500 (Illumina) and mapping raw reads was performed as previously described. [ 39 ] The raw files have been uploaded to Gene Expression Omnibus under the accession GSE143223. Data was merged with previously sequenced LiverPool Cryoplateable Hepatocytes (GSE123498) [ 28 ] and normalized. Differentially-expressed genes were identified using the DESeq2 package with standard settings. [ 40 ] Human liver tissue whole transcriptome sequencing data (Liver_FASEB) was obtained from an online database. [ 41 ] Heatmaps were generated using gplots. ToppFun was used for functional enrichment analysis based on functional annotations and protein interactions networks. [ 33 ] Human Recombinant Laminin-111 Condensation To obtain high-concentration human recombinant laminin-111 (hrlaminin-111), commercial hrlaminin-111 (Biolamina) was condensed with an Amicon Ultra-0. 5 centrifugal filter unit (Millipore, Merck) following the manufacturer’s instructions. hrlaminin-111 was concentrated from 100 μg mL −1 to ≈5 mg mL −1, as measured by BCA assay with a Pierce BCA protein assay kit (Thermo Scientific). Afterward, the condensed hrlaminin-111 replaced LEC to make a chemically-defined PIC hydrogel for liver organoid culture. Statistical Analyses ABA results ( Figure 1d, e ) and rheological results ( Figure 2 ) were analyzed in an Excel datasheet and converted into graphs using Graphpad Prism 8. qRT-PCR results ( Figures 4b and 5b ), ALB secretion ( Figure 4c ), GLDH expression ( Figure 4e ), and ammonium elimination ( Figure 4f ) were analyzed using a Tukey’s multiple comparisons test by two-way ANON multiple comparisons. The p -values are indicated in the respective figures. Supplementary Material Supporting Information is available from the Wiley Online Library or from the author. Supporting information
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10. 1002/adfm. 202003710
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Advanced functional materials
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An Electroactive Oligo-EDOT Platform for Neural Tissue Engineering
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The unique electrochemical properties of the conductive polymer poly(3, 4-ethylenedioxythiophene):polystyrene sulfonate (PEDOT:PSS) make it an attractive material for use in neural tissue engineering applications. However, inadequate mechanical properties, and difficulties in processing and lack of biodegradability have hindered progress in this field. Here, the functionality of PEDOT:PSS for neural tissue engineering is improved by incorporating 3, 4-ethylenedioxythiophene (EDOT) oligomers, synthesized using a novel end-capping strategy, into block co-polymers. By exploiting end-functionalized oligoEDOT constructs as macroinitiators for the polymerization of poly(caprolactone), a block co-polymer is produced that is electroactive, processable, and bio-compatible. By combining these properties, electroactive fibrous mats are produced for neuronal culture via solution electrospinning and melt electrospinning writing. Importantly, it is also shown that neurite length and branching of neural stem cells can be enhanced on the materials under electrical stimulation, demonstrating the promise of these scaffolds for neural tissue engineering.
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1 Introduction To investigate fundamental questions in human brain development and disease, tissue engineering approaches have employed 3D polymer based biomaterial scaffolds as guidance cues to grow 3D neuronal cultures in conditions resembling the in vivo environment. [ 1 – 5 ] Yet, fully recapitulating all functional aspects of neuronal networks within an in vitro system still presents significant challenges, due to the complex physiological environment of the brain. An ideal biomaterial for complex functional neural cultures must be able to provide both precise architectural control to guide neural growth and effective electrical communication to interface with electrically excitable neural cells. [ 6 ] There is therefore a pressing need for new 3D electroactive biomaterials able to mediate neural tissue engineering. Advances in bio-fabrication technologies have led to promising developments in specialized 3D architectures for neural tissue engineering, including 3D scaffolds with microscale or nanoscale topography to guide cellular interactions. [ 7 – 11 ] Techniques such as electrospinning have been used to create biomimetic fibrous matrices which have successfully been applied for the growth of neural networks, as suitably aligned nanoscale fibers can provide directionality for the growth of neurons. [ 9, 12 – 17 ] More recently, melt electrospinning writing (MEW) has emerged as a promising method to deposit highly defined fibrous scaffolds. [ 18 ] This technique has so far been minimally used in neural tissue engineering, yet has numerous advantages including reproducible and precise control over fiber diameter and scaffold architecture. MEW therefore provides exciting opportunities for specific cell guidance and the creation of networks with precise alignment and directionality. [ 19 ] The absence of solvents during fabrication by MEW also minimizes toxicity issues during cell culture. Incorporating electroactive functionality into neural scaffolds has been shown to aid recapitulation of in vivo biological environments. [ 20, 21 ] Neurons are connected to each other via synapses, which are used to translate changes in membrane potential into biochemical messages that are exchanged between neighboring neurons and propagate ongoing signals within the network. Although the membrane potential is determined by ion exchange, exogenous electrical stimulation can be a useful tool to interface with living neurons, and in previous studies it has been shown to enhance neuronal differentiation and neurite outgrowth. [ 20, 22, 23 ] Conjugated polymer (CP)-based materials have therefore recently gained significant attention in the field of neural tissue engineering due to their capacity to conduct charge through the transport of both electrons and ions. [ 24 – 26 ] The versatility and flexibility of polymer synthesis also enables the production of varied scaffold architectures, in contrast to traditional metal electrodes, bridging the communication gap between biology and electronics. [ 27 – 31 ] When compared to metal electrodes, CPs have been shown to be beneficial for neural tissue engineering due to their ability to modulate material stiffness and impedance to better match that of neural tissue. [ 32 ] The lower stiffness of CPs compared to metal electrodes has been shown to improve long-term contact with neuronal cells and the electrode, by reducing the mechanical mismatch and reactive tissue response to stiff metal devices, prolonging electrical stimulation and extending beneficial effect on neuronal growth. [ 24, 32, 33 ] This can potentially be further improved by incorporating CPs into conjugated hydrogel (CH) materials, as hydrated hydrogel networks provide an ideal substrate for 3D cell culture, recreating the environment of native soft tissue, and allowing the diffusion of nutrients and signaling biomacromolecules. [ 34 – 36 ] Electroactive hydrogels have been shown to permit electrical stimulation of cells in 3D and to mediate biological signaling. [ 27, 37 – 39 ] However, continued issues with low conductivity have hindered translation of CHs into biomedical applications. [ 24, 40, 41 ] In the last decade, poly(3, 4-ethylenedioxythiophene):polystyrene sulfonate (PEDOT:PSS) has dominated the CP field, due to its excellent electrical properties, chemical and thermal stability and low oxidation potential. [ 24, 25 ] However, despite the promise of CPs as attractive materials for in vitro models of functional brain tissue, challenges with their use continue to hinder widespread application in tissue engineering. Difficulties in functionalization, lack of biodegradability, poor material reproducibility due to low solubility and processability, and inadequate mechanical properties are particularly prominent. [ 25, 42 ] Recently, electroactive oligomers have emerged as a promising alternative to full length CPs. [ 42, 43 ] Typically they share similar electrical properties after doping to the corresponding conducting polymer formed from doping of the parent CP, yet have more potential to install versatile chemical functionality for further derivatization. [ 44, 45 ] The use of conjugated oligomers has several desirable advantages. First, they provide a precisely defined structure that is in stark contrast to the polydisperse nature of CPs. This provides improved control and homogeneity of material structure and function, which is essential to elucidate the underlying mechanisms of signal transduction at the material-cell interface. Second, the greatly improved solubility and processability offered by oligomeric structures provides opportunities to exploit many emerging techniques for advanced biomaterial preparation, and for synergistically regulating physical, topographical, and electrical cues within a single scaffold. [ 45, 46 ] Finally, unlike full length CPs, short oligomers, typically <10 monomers, can be consumed by macrophages, allowing the production of materials that are both electroactive and biodegradable. The benefits of degradable scaffolds for both in vivo and in vitro applications have been highlighted by a number of recent reports demonstrating the benefits of material remodeling and degradation by growing cells on tissue development, as well as for the long term tolerance of hybrid materials. [ 45, 47 – 49, 52, 53 ] In 2002, Schmidt and colleagues first demonstrated the feasibility of incorporating pyrrole oligomers into a fully biodegradable polymer, via degradable ester linkages. [ 47 ] Recent studies from this group reported an oligoaniline based electroactive polymer, with simple and scalable synthesis and purification, that could be used for electrochemically triggered delivery of anti-inflammatory drugs. [ 50, 51 ] Aniline oligomers have also been explored for use in electroactive biodegradable scaffolds for soft tissue repair, including a dopant-free conjugated elastomer polymer, generated by chemically linking conjugated aniline oligomers, biodegradable poly(caprolactone) (PCL) segments and a dopant component. [ 52 ] Polyester-based scaffolds are particularly attractive due to the tunability of polymer degradation rate based on monomer composition. [ 36, 53 ] However, it has so far proven difficult to apply such an approach to PEDOT due to the reported chemical instability of oligoEDOTs and the lack of free functional groups for further functionalization. Our group have recently offered a potential solution to this problem, by reporting the synthesis of precisely defined, glyoxyl-capped, bifunctional oligoEDOT constructs that are stable, possess tunable properties, and provide diverse reactive handles for further derivatization. [ 54 ] In the present study, we utilized our oligoEDOT constructs to provide a modular platform for biomaterial synthesis. End functionalized oligoEDOTs were used as macroinitiators for the synthesis of PCL block co-polymers and as crosslinkers for hydrogelation. We show that solvent electrospinning and MEW can be used to fabricate fibrous scaffolds, with defined nanotopography. This represents a major advantage of our approach, as these methods cannot typically be applied directly to PEDOT-based materials, due to their poor processability. Moreover, we demonstrate that oligoEDOT-PCL is a permissive substrate for neuronal cell culture, and that electrical stimulation enhances neurite length and branching of neuronal cells. 2 Results and Discussion 2. 1 Synthesis of End-Functionalized OligoEDOTs To construct end-functionalized oligoEDOTs we adopted a synthetic approach recently reported by our group. [ 54 ] The iterative process consists of thiophene glyoxylation, bromination, chain extension, and oligomer cross-coupling, and yields end-functionalized EDOT oligomers of defined chain lengths ( Figure S1, Supporting Information ). We previously demonstrated that a variety of functional end-capping handles could be easily incorporated, enabling the generation of a diverse range of functionalized alkoxy-thiophene monomers and oligomers. [ 54 ] We reasoned that amino derivatives would provide a versatile reactive handle for further derivatization, while incorporation of a short triethylene-glycol linker would provide flexibility and enhanced solubility, facilitating processing and subsequent modification ( Figure 1a ). After treatment of EDOT with oxalyl chloride, the intermediate glyoxylyl chloride 1, was reacted with a mono-Boc (tert-butyloxycarbonyl) protected, diaminotriethylene-glycol linker 2, to generate the N -protected, mono-functionalized EDOT monomer 3 ( Figure S1, Supporting Information ). Subsequent bromination yielded di-functional monomer 4, which could undergo subsequent chain extension and oligomer couplings via palladium-catalyzed direct arylation ( Figure S1, Supporting Information ). The use of direct arylation, over alternative strategies such as Stille or Kumada couplings, limited potential problems with poor functional group compatibility and residual catalyst toxicity. [ 55 ] Through this strategy, we were able to generate a range of protected amine-functionalized oligoEDOTs in an iterative manner, with defined chain lengths ( n = 2–5, 5 – 8 ) ( Figure S1, Supporting Information ). 2. 2 oligoedot-pcl characterization To prepare constructs suitable for further processing and scaffold fabrication, we selected PCL as a suitable co-polymer, due to its biodegradable properties and well-established use in tissue engineering. Boc-protected dimer ( 5 ), tetramer ( 7 ), and pentamer ( 8 ) oligoEDOTs were deprotected under acidic conditions to yield the free amines ( 9, 11, and 12 respectively), ( Figure S1, Supporting Information ). Ring-opening polymerization of ε -caprolactone was then undertaken using the amino-oligoEDOT as a macroinitiator ( Figures 1b and 2a ). The new ABA block co-polymer structures (subsequently named oligoEDOT-PCL 16 a–c) were synthesized with a total molecular weight of ≈25 kDa (i. e. , two PCL chains of ≈12 kDa off a central oligoEDOT core of ≈1kDa, weight dictated by monomer:oligoEDOT ratio). This molecular weight was chosen to provide sufficient PCL for improved processability while maximizing the electroactivity of the oligoEDOT block. To investigate the optoelectronic properties of the oligoEDOT-PCL constructs, UV–vis spectra were recorded on thin films ( Figure S2, Supporting Information ). Chemically synthesized PEDOT in the neutral (undoped) state typically absorbs in the visible region from 400 to 600 nm. [ 56 ] Accordingly, the maximum absorbance in samples of the three oligoEDOT-PCL films were between 450 and 500 nm, with the absorbance spectra red-shifted with increasing oligomer chain length as expected with increasing conjugation length ( Figure S2, Supporting Information ). The optical band gap ( E g, opt ) was calculated from the onset of absorption for the three oligoEDOT-PCL films, which ranged from 2. 48 eV for diEDOT-PCL to 2. 03 eV for pentaEDOT-PCL, as further detailed in Table S1, Supporting Information. These observations are consistent with the optical spectra range of the parent amino-oligoEDOTs and the structures recently reported by us. [ 54 ] This confirmed that incorporation of the oligoEDOT into a PCL ABA block co-polymer did not significantly alter the optical properties of the EDOT-block. Next, we conducted cyclic voltammetry to elucidate the electrochemical properties of dimer, tetramer and pentamer oligoEDOT-PCL films in organic and aqueous solutions ( Figure 2b and Figure S3, Supporting Information ). TetraEDOT-PCL 16 films provided the most stable electrochemical behavior over several samples. We therefore used tetraEDOT-PCL for more in-depth thin film characterization, scaffold preparation, and neuronal cell culture. The tetraEDOT-PCL film exhibited a broad anodic (0. 8–1. 1 V) and a broad cathodic (0. 4–1 V) peak, with tetrabutylammonium hexafluorophosphate (TBAPF 6 ) as the supporting electrolyte at a scan rate of 0. 1 V s −1, demonstrating redox activity corresponding to the formation of a radical cation ( Figure 2b ). To further examine the electronic properties of the oligoEDOT-PCL films, we applied Kelvin probe force microscopy (KPFM) to investigate changes in surface potential at the localized micro-scale. Such experiments provide a more accurate reflection of the local electrical environment experienced by individual cells than bulk material measurements. Figure 2c shows a heterogenous surface potential map of the tetraEDOT-PCL 16 b polymer film, with a local surface potential, V CPD, obtained from the contact potential difference (CPD) between the conductive tip and sample. Distinct regions of lower and higher work functions were evident on the surface of the film. Figure 2d illustrates a line scan along the black dotted line, depicting the difference in work function between the bright and dark areas. These results suggest nanoscale aggregation of the conjugated oligoEDOT on the PCL co-polymer surface. Subsequently, we used atomic force microscopy (AFM) scanning to further examine the effects of polymerizing PCL with conjugated EDOT oligomers on surface topography ( Figure 2e ). TetraEDOT-PCL films were compared with pure PCL films with a comparable molecular weight ≈25 kDa, and high molecular weight PCL ≈c75 kDa. High molecular weight PCL films displayed lamellar structures, which aggregated into compact spherules of ≈1. 5–2 m in size, consistent with previous reports ( Figure 2e ). [ 57 ] Image scans of low molecular weight PCL films showed a similar surface morphology, with evidence of more striated lamellar, indicating increased crystallinity, consistent with the reduced molecular weight. AFM scans of tetraEDOT-PCL revealed distinct differences in the surface morphology compared to both pure PCL samples, as individual fibrils of ≈40 nm were resolved on the polymer surface. These image scans demonstrate that the inclusion of EDOT oligomers influences the surface morphology. Additionally, by blending oligoEDOT-PCL films with higher molecular weight PCL at different blend ratios, we were able to tune the morphological properties of the films ( Figure S4, Supporting Information ). Polymer blending has emerged as a versatile technique to adjust material properties, and has previously been used to combine certain favorable material properties. [ 58 ] By blending oligoEDOT-PCL with high molecular weight PCL, we demonstrate the possibility to adjust the mechanical properties of the material according to the application and scaffold type. It is important to note that the addition of increasing volumes of insulating PCL will also negatively influence the electrical properties of the material. Further studies are required to elucidate the PCL:oligoEDOT ratio window in which both beneficial electroactivity and enhanced material processability are maintained. 2. 3 OligoEDOT Crosslinked Hydrogels We next wanted to test whether our oligoEDOTs could be effectively used as hydrogel crosslinkers, to demonstrate the versatility of our approach for materials preparation. We therefore set out to undertake the gelation of complementary polyethylene glycol (PEG) macromers ( Figure 1c ). While amine salts 9 – 12 were able to provide partial water solubility, upon buffering to physiological pHs the oligomers quickly became insoluble at the concentrations required for hydrogelation. Attempts to form hydrogels with 8-arm PEG succinimide esters therefore yielded highly heterogeneous, mechanically weak gels encapsulating areas of precipitated oligomer ( Figure S5, Supporting Information ). We therefore sought to make use of the synthetic versatility of our amino-oligoEDOT series to introduce water solubilizing functionalities. Thiol groups provide convenient reactive handles for hydrogelation via a number of mechanisms, including nucleophilic Michael addition and radical thiol-ene reactions. [ 58 ] These reactions often proceed with high specificity and rapid reaction kinetics, leading to their widespread use in the biomaterial community. Moreover, it has been widely shown that simple changes in gelation conditions such as pH and temperature, or alterations to thiol or alkene chemistry, can dramatically influence gel properties providing tunability. [ 60 – 62 ] Amino-triEDOT 10 was therefore derivatized with cysteine 13 in an attempt to both enhance water solubility and increase hydrogelation efficiency ( Figure S6, Supporting Information ). We attempted gelation with 8-arm PEG-maleimide, however water solubility continued to limit efficiency and gel homogeneity. A Glu-Glu-Cys tripeptide 14 was therefore ligated at the oligomer termini to provide additional charge at physiological pH ( Figure S6, Supporting Information ). Functionalized oligomer 15 was found to be water soluble over a wide pH range, even at the high concentrations necessary for hydrogelation. Mechanically stable hydrogels (5% by polymer weight) were immediately formed upon mixing with 8-arm PEG-maleimide in PBS ( Figure S5, Supporting Information ). A slight improvement in gel homogeneity could be achieved by undertaking gelation at pH 6, due to an increase in gelation time that allowed efficient mixing. Gel stability was high, as monitored through leaching of the triEDOT crosslinker 15 out of the gel ( Figure S7, Supporting Information ). Although a slight pH dependence on gel swelling ratio (relative to the lyophilized gels) was observed, the effect was minimal and likely due to the presence of the EEC tripeptide ( Figure S8, Supporting Information ). The same peptide-modified oligomer strategy was then applied to the formation of tetra- and penta-EDOT crosslinked gels. However, the increase in oligomer length was found to be sufficient to reduce water solubility and prevent gelation, with oligomer precipitation resulting instead. For this reason, further hydrogel processing and characterization was not pursued, and we instead chose to focus on the generation of fibrous oligoEDOT-based scaffolds as described in subsequent sections. These results highlight the importance of overcoming the hydrophobicity of extended π -systems to allow the synthesis of electroactive hydrogels. [ 27 ] Although this hydrophobicity IS limiting in the work presented here, the high modularity of our oligoEDOT derivatization strategy offers a potential means to address this difficulty. As demonstrated above, water solubilizing chains can be easily ligated to oligomer building blocks. Furthermore, we have previously demonstrated that our oligoEDOT synthesis strategy is applicable to alternative dialkoxythiophene monomers. [ 54 ] The use of reported water soluble EDOT derivatives, bearing carboxyl or sulfonate group is particularly attractive. [ 63 – 65 ] It is therefore likely that optimization of this process will allow the future synthesis of tetra- and penta-oligomers with sufficient solubility for gelation. 2. 4 OligoEDOT-PCL Films Support Neural Stem Cell Growth and Differentiation To investigate the suitability of oligoEDOT-PCL films as substrates for neural scaffolds, human iPSC-derived neural stem cells (NSCs) were seeded onto spin-cast films. [ 66 ] Seeded scaffolds were cultured in fibroblast growth factor-2 (FGF2) containing media to promote NSC proliferation, or media without any mitogens to promote differentiation (Basal Medium) ( Figure 3a ). [ 67 ] After 7 days of culture, cells were stained for β III-tubulin (a marker of differentiated neurons), nestin (a marker of neuronal progenitors and stem cells) and ki-67 (a proliferation marker). The addition of FGF2 media showed that cells were proliferating and promoted the expression of ki-67 ( Figure 3a ). The total cell count on the oligoEDOT-PCL films in basal and FGF2 media showed no significant difference to the PCL scaffolds or glass controls ( Figure 3b ). Immunostaining revealed a significant increase in β III-tubulin positive cells cultured in basal media on all substrates, with no significant difference being observed between NSCs cultured on oligoEDOT-PCL films compared to the control groups ( Figure 3c ). Together, these results demonstrate that oligoEDOT-PCL films are biocompatible and are a suitable substrate for NSC differentiation and proliferation. 2. 5 Scaffold Preparation of OligoEDOT-PCL Electrospun fibrous membranes have been used extensively in tissue engineering applications, due to the architectural resemblance of micrometer diameter fibers to the fibrils found in native extracellular matrix. [ 68 ] To generate electroactive 3D fibrous scaffolds from our oligoEDOT-PCL constructs, membranes were prepared using solvent electrospinning, and MEW. Typically, these methods cannot be applied directly to PEDOT due to its poor solubility and processability. Our block co-polymer approach therefore offers an opportunity to address these limitations and construct electroactive, fibrous, EDOT-based scaffolds for neuronal culture. For solution electrospinning, tetraEDOT-PCL films were blended at a 50% ratio with high molecular weight PCL, and polymer solutions were spun at a 20% w/v concentration in a 9:1 mixture of CHCl 3 : CH 3 OH, following our previously reported protocol. [ 58 ] This resulted in the deposition of pink colored fibrous mats after electrospinning ( Figure S9, Supporting Information ). Fiber diameters were measured by scanning electron microscopy (SEM), with an average diameter of ≈400 nm ( Figure 4a ). In order to assess the suitability of the produced mats for neuronal culture, NSCs were seeded on electrospun tetraEDOT-PCL scaffolds and cultured in basal medium to promote differentiation for 24 h prior to fixing ( Figure 4b ). Immunostaining for β III-tubulin and nestin indicated that NSCs adhered to the membrane and underwent neuronal differentiation ( Figure S10, Supporting Information ). Interestingly, the oligoEDOT-PCL fibers were found to be fluorescent in the red channel ( Figure S10, Supporting Information ). We reasoned that this may be due to the shear forces or exposure to high voltages during electrospinning, affecting the molecular properties of the polymer. Current studies in our group are trying to elucidate the physical processes that occur to cause this change in oligomer emission. Recent studies have reported that low pore size in electrospun scaffolds can act as a barrier to cell penetration. [ 69, 70 ] Studying cell behavior in complex 3D environments can also be challenging due to randomly orientated fibers and heterogenous architectures. [ 70 ] To investigate the suitability of oligoEDOT-PCL for the production of scaffolds with defined microarchitectures, we also used MEW to construct a 5 cm × 5 cm 3D lattice, which comprised of 10 stacked layers, with an approximate spacing of 100 μm between grids and a fiber diameter of 5 μm ( Figure 4c ). Interestingly, the scaffold color here was gold, further suggesting that factors during the fabrication process, such as shear stress and voltage, may have an effect on the molecular packing of the oligoEDOT-PCL polymer ( Figure S11, Supporting Information ). We examined cell adherence by seeding NSCs on oligoEDOT-PCL scaffolds for 24 h in differentiation media, and SEM imaging revealed that NSCs adhered to individual fibers in the scaffold ( Figure 4d ). Consistent with previous studies, neurites within our scaffolds preferentially extended in the same directions as the fibers ( Figure 4d, e ). [ 9, 13 – 15, 17 ] 2. 6 Electrical Stimulation Enhances Neurite Outgrowth The key advantage of CP-based scaffolds for neural tissue engineering are the opportunities they provide to electrically stimulate cells. Previous studies have shown that switching the redox state of CPs by applying a voltage can change neuronal behavior, including cell adherence, proliferation, and differentiation, though the precise mechanism is not yet fully understood. [ 20, 71, 72 ] However, the poor mechanical properties of conventional CPs can limit the ability to fabricate complex scaffolds and sustain long term electrical stability. Our oligoEDOT platform provides an exciting opportunity to overcome this challenge. To investigate the effects of electrical stimulation on NSCs cultured on oligoEDOT-PCL films, we examined neurite length and branching in NSCs after applying a pulsed direct current (DC) ( Figure 5a ). To prepare the films, tetraEDOT-PCL 16b films were deposited by spin coating directly on to indium tin oxide (ITO) glass, a conductive substrate serving as the working electrode. We chose to use PCL films spin coated on ITO glass as the non-conductive control to eliminate possible contributions from electrochemical processes in the cell culture medium. For the conductive control, bare ITO was used, which has previously been used as a substrate to stimulate neural cells in vitro and has been shown to evoke an electrical response in cultured neurons. [ 73 ] An electrode cell assembly, constructed in-house, was used for in vitro NSC stimulation ( Figure S12, Supporting Information ), and NSCs were seeded in the chambers for 24 h prior to stimulation. A platinum wire counter electrode was placed inside the cell culture medium in the micro chambers, at a distance of 1 cm from the oligoEDOT-PCL films. To better mimic physiological conditions of neural network activity, we used pulsed electrical stimulation. We also reasoned that this would avoid a build-up of charge, thereby allowing long term electrical stimulation, limiting any adverse effects on cell viability. Trains of 1 ms pulses of 600 mV at 1 Hz were applied for 24 h, followed by fixing and immunostaining of cells. We observed an increase in mean NSC neurite length following stimulation on tetraEDOT-PCL films (142. 1 ± 10. 4 μm) compared to unstimulated tetraEDOT-PCL films (111. 4 ± 8. 7 μm) ( Figure 5b and Figure S13, Supporting Information ). Similarly, neurite length increased when NSCs were stimulated on ITO glass control substrates, compared to the unstimulated group (140. 1 ± 10. 7 and 109. 9 ± 8. 06 μm respectively) ( Figure 5b and Figure S13, Supporting Information ). Surprisingly, neurite length was found to decrease upon stimulation of NSCs on the PCL negative control group (60. 5 ± 5. 2 and 82. 4 ± 6. 45 μm) for stimulated and unstimulated groups respectively, emphasizing the important role of the oligoEDOT block in promoting neurite extension. Neurite branching was similarly increased on stimulated oligoEDOT-PCL films and ITO control substrates, compared to their respective unstimulated groups ( Figure 5c ). These results are consistent with previous studies which have reported an increase in neurite outgrowth following electrical stimulation on conventional CPs. [ 20, 74 – 77 ] Our substrates, exhibiting greatly improved processability and more versatile material properties, therefore offer an exciting alternative to pure CP scaffolds for stimulating neuronal cultures. It has previously been proposed that cellular changes in response to electrical stimulation are initiated at the cell surface, altering cell surface receptors and protein adsorption, or modulating the growth cone morphology. [ 78, 79 ] In our study, it is also plausible that electrical stimulation causes changes in the redox states of the EDOT oligomer, changing the bulk properties and surface tension of the polymer, thereby causing changes in NSC behavior. Further investigations into the origins of these effects are currently underway. Critically, such studies are facilitated by the precisely defined molecular structure provided by our oligoEDOT strategy. 3 Conclusion In a search for electroactive materials suitable for the fabrication of complex architectures in tissue engineering, we have developed a new electroactive ABA block co-polymer, oligoEDOT-PCL. We have achieved a route to synergistically apply electrical cues and create controlled topographies by combining the electroactive properties of oligoEDOT structures, with the favorable processability of PCL. The combination of these features is critical to achieving more complex architectures for in vitro models of the developing brain, and we are currently exploring the application of this material to develop highly defined scaffolds to guide 3D neuronal growth. Future work should focus on harnessing the well-defined molecular structure offered by our oligoEDOT synthesis strategy for precise control over chemical functionalization, and the redox active properties for controlled delivery of soluble growth factors and charged small molecules. This will provide the opportunity to better mimic native tissue and provide spatio-temporally controlled chemical guidance cues, such as patterning factors. Significantly, this study demonstrates that oligoEDOT based biomaterials have potential for neural tissue engineering, by providing a modular platform for biomaterial synthesis, thereby improving processability and the potential to generate complex 3D architectures for tissue engineering scaffolds. 4 Experimental Section Details of oligoEDOT synthesis and their PCL co-polymers, electrical stimulation device design and setup are included in the Supporting Information. Preparation of OligoEDOT-PCL FILMS : Thin films were prepared by spin coating using oligoEDOT-PCL polymer solutions. Namely, diEDOT-PCL, tetraEDOT-PCL, pentaEDOT-PCL and control PCL polymer solutions, including high M w PCL (70–90 kDa) and low M w PCL (24 kDa) were spin coated at 6 w/v %. Polymer solutions were prepared in CHCl 3, sonicated for 15 min and left overnight to aid polymer dissolution. For UV–vis, microscope glass slides (10 mm × 10 mm) were cut with a diamond tip. The glass slides were cleaned by sonicating for 15 min each in acetone and isopropanol, dried under a nitrogen stream and treated with oxygen plasma before spin coating. 100 μL of polymer solution was spin coated on glass substrates for 20 s at 1500 rpm and placed in a fume hood overnight. UV–Vis Spectroscopy : Absorbance spectra were recorded with a Shimadzu UV-1601 UV/vis spectrophotometer, in the range between 300 and 900 nm. The spectra diEDOT-PCL, tetraEDOT-PCL and pentaEDOT-PCL films were measured, and film samples were attached with blue tac to the outside of the cuvette holder. In order to calculate the optical band gap energy, the longest absorption wavelength λ onset was used, according to the following equation. [ 80 ] (1) E g = 1242 / λ onset Cyclic Voltammetry : Cyclic voltammograms were conducted using an Autolab PGSTAT101 potentiostat. Recordings were carried out using a conventional three-electrode setup, using an isolated Ag/AgCl reference, platinum counter electrode, and a glassy carbon working electrode. Films were deposited on the glassy carbon electrode via drop casting. The measurements were carried out in 0. 5 M tetrabutylammonium hexafluorophosphate dissolved in propylene carbonate as the supporting electrolyte. Samples were measured over the potential range of −0. 45 to 1. 20 V at a scan rate of 100 mV s −1. SEM : OligoEDOT-PCL solution electrospun and MEW scaffolds were mounted on a SEM pin mount, and sputter coated with a thin gold layer. Images were obtained using a JSM 6010LA SEM (JEOL). NSCs on oligoEDOT-PCL scaffolds were pre-fixed with 3. 7% (w/v) paraformaldehyde and washed in 0. 1 M sodium cacodylate buffer for 5 min. NSCs were then fixed in 2. 5% (w/v) glutaraldehyde in 0. 1 M cacodylate buffer for 1 h at room temperature, and washed 2 } 5 min in 0. 1 M sodium cacodylate buffer. For osmium tetroxide staining, a solution of 1% (v/v) OsO 4 was prepared in 0. 1 M sodium cacodylate buffer, and samples were incubated for 1 h. Samples were rinsed with milli-Q water twice for 5 min. Samples were then serially dehydrated in graded series of ethanol, as described, and treated with hexamethyldizilazane for 5 min. Samples were then attached to SEM pin mount and coated with 24 nm chromium in a sputter coater (Q150T S Quorum). Images were obtained using a Zeiss Sigma 300. AFM : AFM measurements were performed on an Agilent 5500 AFM. Topographical images were obtained in tapping mode using Micromasch HQ-NSC cantilevers (nominal spring constant of 40 N m −1 ) at a resonant frequency of 260 kHz. Images were taken at 40, 10, and 2 μm scale, and were processed using Gwyddion software. KPFM : KPFM was used to obtain information about the surface potential of the oligoEDOT-PCL films, and was performed using an Asylum MFP-3d microscope. Images were obtained using Nanosensors PPP-EFM cantilevers (nominal spring constant of 2. 8 N m −1 ), coated in Pt/Ir with an applied DC bias of 0. 2 V. CPD on the sample was calculated using the following equation (2) CPD = ϕ tip − ϕ sample e where e is the electron charge. OligoEDOT-PCL samples were spincoated on ITO glass. Cell Culture : Human episomal iPSC line (Epi-hiPSC) (Thermo Fisher Scientific, U. K. ), were maintained in feeder-free conditions on Matrigel substrate with Essential 8 media (Thermo Fisher Scientific) and passaged when they reached 80–90% confluence with EDTA solution (0. 5 m M EDTA/PBS). Neural induction was based on a previously published protocol. [ 62 ] iPSCs were differentiated into neuroectoderm when they reached 80–90% confluence by dual SMAD signaling inhibition using neural induction medium [(Advanced DMEM/F-12 medium (Thermo Fisher Scientific), 0. 2% (v/v) B27 Supplement (Invitrogen), 1% (v/v) N2 supplement (Invitrogen, U. K. ), 1% (v/v) penicillin/streptomycin (Invitrogen), dorsomorphin (2 μ M ; Calbiochem, U. K. ), 1% (v/v) GlutaMAX (Invitrogen) supplemented with SB431542 (10 μ M ; Tocris, U. K. ), and N-acetylcysteine (1 m M ; Sigma-Aldrich)] for 7 d, as previously described. [ 62 ] NSCs were then passaged and plated on laminin-coated plates NSCR base medium [ 62 ] (DMEM/F-12 medium (Thermo Fisher Scientific), 1% (v/v) N2 supplement, 0. 2% B27 supplement, 1% (v/v) nonessential amino acids, 1% (v/v) penicillin/streptomycin and 1% (v/v) GlutaMax solution (all from Invitrogen), and B-27 medium (Neurobasal medium (Thermo fisher scientific), 2% (v/v) B27 supplement, 1% (v/v) nonessential amino acids, 1% (v/v) penicillin/streptomycin and 1% (v/v) GlutaMAX solution). After 3–5 days, iPSC derived NSCs formed neural rosette structures, and NSCs were then maintained in F20 medium (NSCR neural maintenance base medium supplemented with 20 ng mL −1 of FGF2 [PeproTech]) to promote expansion of NSCs. NSCs were typically passaged every 5 days on matrigel coated plates, and the medium was changed every 48 h, until cells reached 80–90% confluence. FILM Preparation for Cell Culture and Cell Culture Device Setup : TetraEDOT-PCL was blended with high molecular weight PCL ( M w ≈ 75 000) at a 50:50 ratio and dissolved at 6% (w/v) in chloroform. Polymer films were prepared by spin coating on cover glass as described above and transferred into coverslip bottom 24 well plates (ibidi). Silicone O-rings (10. 77 mm diameter) were autoclaved, placed over glass coverslips in the 24 well plates, followed by three washes in sterile PBS and sterilized by cell culture-grade UV-light irradiation for 30 min. 500 μL of matrigel (at a dilution of 1:120 in DMEM/F12 medium) was added to each well and incubated for 30 min at 37°C in a cell culture incubator. For electrical stimulation experiments, a custom-made cell culture device was designed using teflon microchambers, which were autoclaved prior to use. Polymer films were prepared by spin coating 6% (w/v) TetraEDOT-PCL on ITO glass electrodes, as described, and assembled in the microchambers in a cell culture hood. The wells were washed three times in sterile PBS and sterilized by cell culture-grade UV light irradiation for 30 min followed by 500 μL of matrigel (at a dilution of 1:120 in DMEM/F12 medium). NSCs were seeded on the polymer film samples at a density of 10 × 10 4 per well using a dry plating method and left to incubate at 37°C for 20 min. For NSC proliferation and differentiation experiments, F20 media and 50:50 media was added, respectively, and medium was changed every 48 h. For electrical stimulation experiments, 50:50 media was added to promote neuronal differentiation. Electrical Stimulation of NSCs on OligoEDOT-PCL Films : After 24 h of NSC seeding for cell attachment, a counter electrode platinum wire was suspended into the microchamber well, parallel to the oligoEDOT-PCL film seeded with NSCs. The platinum wire and ITO were connected to an eDaq potentiostat (EA163) and trains of 1 ms electrical pulses of 600 mV at 1 Hz were applied for a period of 24 h. After stimulation, NSCs were fixed with 3. 7% (v/v) paraformaldehyde for 15 min followed by immunostaining. Immunostaining : For NSC proliferation and differentiation experiments on oligoEDOT-PCL films, cells were washed with PBS and fixed with 3. 7% (v/v) paraformaldehyde (Sigma-Aldrich) for 15 min at room temperature, and then washed three times with sterile PBS. OligoEDOT-PCL films were permeabilized with 0. 2% (v/v) Triton X-100 (Sigma-Aldrich) in PBS for 10 min, and then blocked with 3% (v/v) goat serum (Sigma-Aldrich) for 30 min, followed with primary antibodies, nestin (1:500; Millipore, U. K. ), β III-tubulin (1:1000; Sigma-Aldrich), and Ki67 (1:1000; Abcam), followed with DAPI (Sigma-Aldrich) and secondary antibodies (Alexa Fluor dyes; Thermo Fisher Scientific) for 1 h. For electrical stimulation experiments, cells were fixed, permeabilized and blocked as described, followed with primary antibodies, nestin (1:500; Millipore, U. K. ), β III-tubulin (1:1000; Sigma-Aldrich), followed with DAPI (Sigma-Aldrich) and secondary antibodies (Alexa Fluor dyes; Thermo Fisher Scientific) for 1 h. For electrical stimulation experiments, stained samples on ITO glass were mounted with coverslip slides with FluorSave Reagent (Millipore) and stored at 4°C. Images of NSCs on polymer films were obtained with Inverted Widefield Microscope (Zeiss Axio Observer), and Z stacks of ≈30 slices were obtained. Images on electrospun fibrous mats were acquired with a Zeiss LSM 510 inverted confocal microscope. Imaging Analysis and Statistical Analysis : Images analysis was performed with ImageJ software. Z stack images were deconvolved using Huygens software. Neural stem cell (NSC) proliferation and differentiation was analyzed by counting the percentage of β III-tubulin + cells over the total number of cells, using the cell counter plugin. Neurite outgrowth was analyzed using the Neurite tracings plugin. Neurite outgrowth was analyzed by measuring individual neurites for each soma, using the Neurite tracings plugin in Fiji. Neurite branching was quantified using the cell counter plugin in Fiji. 3 or 4 random images were taken for each sample. For statistical analysis all experiments were conducted at least four times. Biocompatibility experiments were conducted with a minimum of 2 biological replicates and 3 technical replicates, and a minimum of 15 images were analyzed for each condition. Neurite length and branching experiments were conducted with a minimum of 3 biological replicates and 2 technical replicates, except where replicates without quantifiable cells were excluded from the analysis. A minimum of 12 images were analyzed for each condition. One-way ANOVA with post hoc Tukey’s test was used, and a p -value of <0. 05 was considered statistically significant. Solution Electrospinning : TetraEDOT-PCL was blended with high molecular weight PCL ( M w ≈ 75 000 Da) at a 50:50 ratio and dissolved at 20% (w/v) in a 9:1 mixture of chloroform and methanol. OligoEDOT-PCL polymer solutions were processed into fibrous scaffolds using a custom build electrospinning device, as previous described. [ 57 ] Scaffolds were spun at 12 kV, 20 μL min −1 flow rate, with a distance of 15 cm between the needle and collector plate. Fiber diameters were measured using ImageJ, based on SEM. MEW : TetraEDOT-PCL was blended with PCL ( M w ≈ 75 000 Da) at a 50:50 ratio, by dissolving in CHCl 3, and precipitating in diethyl ether. The blended polymer was loaded into a 3 cc polypropylene syringe (Nordson #7 012 074) and heated overnight at 75°C in a vacuum oven to remove bubbles. A 23G needle (Nordson #7 018 302) was installed, and the syringe was inserted into the spinneret. After equilibrating for 30 min at 70°C, scaffolds were printed using a 10 kV accelerating voltage, 10 mm collector distance, axis velocity of 1500 mm min −1, feeding air pressure of 1. 0 Bar, and heating temperatures between 60 and 80°C. Printing was controlled by MACH 3 CNC software (ARTSOFT, Livermore Falls, USA), and dimensions were specified by G-code. Each scaffold consisted of 10 stacked layers. Scaffolds were detached from the collector plate using a drop of ethanol and moved to a petri-dish. For cell culture experiments, scaffolds were sterilized by cell culture-grade UV light irradiation for 30 min. Supplementary Material Supporting Information is available from the Wiley Online Library or from the author. Supplementary Information
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10. 1002/adfm. 202005010
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Advanced functional materials
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Reversible Functionalization of Clickable Polyacrylamide Gels with Protein and Graft Copolymers
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Modular strategies to fabricate gels with tailorable chemical functionalities are relevant to applications spanning from biomedicine to analytical chemistry. Here, the properties of clickable poly(acrylamide-co-propargyl acrylate) (pAPA) hydrogels are modified via sequential in-gel copper-catalyzed azide-alkyne cycloaddition (CuAAC) reactions. Under optimized conditions, each in-gel CuAAC reaction proceeds with rate constants of ~0. 003 s −1, ensuring uniform modifications for gels < 200 μm thick. Using the modular functionalization approach and a cleavable disulfide linker, pAPA gels were modified with benzophenone and acrylate groups. Benzophenone groups allow gel functionalization with unmodified proteins using photoactivation. Acrylate groups enabled copolymer grafting onto the gels. To release the functionalized unit, pAPA gels were treated with disulfide reducing agents, which triggered ~50 % release of immobilized protein and grafted copolymers. The molecular mass of grafted copolymers (~6. 2 kDa) was estimated by monitoring the release process, expanding the tools available to characterize copolymers grafted onto hydrogels. Investigation of the efficiency of in-gel CuAAC reactions revealed limitations of the sequential modification approach, as well as guidelines to convert a pAPA gel with a single functional group into a gel with three distinct functionalities. Taken together, we see this modular framework to engineer multifunctional hydrogels as benefiting applications of hydrogels in drug delivery, tissue engineering, and separation science.
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No full text available
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10. 1002/adfm. 202006796
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Advanced functional materials
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Bone-on-a-chip: microfluidic technologies and microphysiologic models of bone tissue
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Bone is an active organ that continuously undergoes an orchestrated process of remodeling throughout life. Bone tissue is uniquely capable of adapting to loading, hormonal, and other changes happening in the body, as well as repairing bone that becomes damaged to maintain tissue integrity. On the other hand, diseases such as osteoporosis and metastatic cancers disrupt normal bone homeostasis leading to compromised function. Historically, our ability to investigate processes related to either physiologic or diseased bone tissue has been limited by traditional models that fail to emulate the complexity of native bone. Organ-on-a-chip models are based on technological advances in tissue engineering and microfluidics, enabling the reproduction of key features specific to tissue microenvironments within a microfabricated device. Compared to conventional in-vitro and in-vivo bone models, microfluidic models, and especially organs-on-a-chip platforms, provide more biomimetic tissue culture conditions, with increased predictive power for clinical assays. In this review, we will report microfluidic and organ-on-a-chip technologies designed for understanding the biology of bone as well as bone-related diseases and treatments. Finally, we discuss the limitations of the current models and point toward future directions for microfluidics and organ-on-a-chip technologies in bone research.
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No full text available
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10. 1002/adfm. 202006967
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Advanced functional materials
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Additive Manufacturing of Material Scaffolds for Bone Regeneration: Toward Application in the Clinics
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Additive manufacturing (AM) allows the fabrication of customized bone scaffolds in terms of shape, pore size, material type and mechanical properties. Combined with the possibility to obtain a precise 3D image of the bone defects using computed tomography or magnetic resonance imaging, it is now possible to manufacture implants for patient-specific bone regeneration. This paper reviews the state-of-the-art of the different materials and AM techniques used for the fabrication of 3D-printed scaffolds in the field of bone tissue engineering. Their advantages and drawbacks are highlighted. For materials, specific criteria, were extracted from a literature study: biomimetism to native bone, mechanical properties, biodegradability, ability to be imaged (implantation and follow-up period), histological performances and sterilization process. AM techniques can be classified in three major categories: extrusion-based, powder-based and liquid-base. Their price, ease of use and space requirement are analyzed. Different combinations of materials/AM techniques appear to be the most relevant depending on the targeted clinical applications (implantation site, presence of mechanical constraints, temporary or permanent implant). Finally, some barriers impeding the translation to human clinics are identified, notably the sterilization process.
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No full text available
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10. 1002/adfm. 202007199
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Advanced functional materials
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Engineering new microvascular networks on-chip: ingredients, assembly, and best practices
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Tissue engineered grafts show great potential as regenerative implants for diseased or injured tissues within the human body. However, these grafts suffer from poor nutrient perfusion and waste transport, thus decreasing their viability post-transplantation. Graft vascularization is therefore a major area of focus within tissue engineering because biologically relevant conduits for nutrient and oxygen perfusion can improve viability post-implantation. Many researchers utilize microphysiological systems as testing platforms for potential grafts due to an ability to integrate vascular networks as well as biological characteristics such as fluid perfusion, 3D architecture, compartmentalization of tissue-specific materials, and biophysical and biochemical cues. While many methods of vascularizing these systems exist, microvascular self-assembly has great potential for bench-to-clinic translation as it relies on naturally occurring physiological events. In this review, we highlight the past decade of literature and critically discuss the most important and tunable components yielding a self-assembled vascular network on chip: endothelial cell source, tissue-specific supporting cells, biomaterial scaffolds, biochemical cues, and biophysical forces. This article discusses the bioengineered systems of angiogenesis, vasculogenesis, and lymphangiogenesis, and includes a brief overview of multicellular systems. We conclude with future avenues of research to guide the next generation of vascularized microfluidic models and future tissue engineered grafts.
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No full text available
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10. 1002/adfm. 202009946
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Advanced functional materials
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Single-Cell Microgels for Diagnostics and Therapeutics
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Cell encapsulation within hydrogel droplets is transforming what is feasible in multiple fields of biomedical science such as tissue engineering and regenerative medicine, in vitro modeling, and cell-based therapies. Recent advances have allowed researchers to miniaturize material encapsulation complexes down to single-cell scales, where each complex, termed a single-cell microgel, contains only one cell surrounded by a hydrogel matrix while remaining <100 μm in size. With this achievement, studies requiring single-cell resolution are now possible, similar to those done using liquid droplet encapsulation. Of particular note, applications involving long-term in vitro cultures, modular bioinks, high-throughput screenings, and formation of 3D cellular microenvironments can be tuned independently to suit the needs of individual cells and experimental goals. In this progress report, an overview of established materials and techniques used to fabricate single-cell microgels, as well as insight into potential alternatives is provided. This focused review is concluded by discussing applications that have already benefited from single-cell microgel technologies, as well as prospective applications on the cusp of achieving important new capabilities.
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No full text available
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10. 1002/adfm. 202100015
| 2,021
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Advanced Functional Materials
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Development of Fluorine‐Free Tantalum Carbide MXene Hybrid Structure as a Biocompatible Material for Supercapacitor Electrodes
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Abstract The application of nontoxic 2D transition‐metal carbides (MXenes) has recently gained ground in bioelectronics. In group‐4 transition metals, tantalum possesses enhanced biological and physical properties compared to other MXene counterparts. However, the application of tantalum carbide for bioelectrodes has not yet been explored. Here, fluorine‐free exfoliation and functionalization of tantalum carbide MAX‐phase to synthesize a novel Ta 4 C 3 T x MXene‐tantalum oxide (TTO) hybrid structure through an innovative, facile, and inexpensive protocol is demonstrated. Additionally, the application of TTO composite as an efficient biocompatible material for supercapacitor electrodes is reported. The TTO electrode displays long‐term stability over 10 000 cycles with capacitance retention of over 90% and volumetric capacitance of 447 F cm −3 (194 F g −1 ) at 1 mV s −1. Furthermore, TTO shows excellent biocompatibility with human‐induced pluripotent stem cells‐derived cardiomyocytes, neural progenitor cells, fibroblasts, and mesenchymal stem cells. More importantly, the electrochemical data show that TTO outperforms most of the previously reported biomaterials‐based supercapacitors in terms of gravimetric/volumetric energy and power densities. Therefore, TTO hybrid structure may open a gateway as a bioelectrode material with high energy‐storage performance for size‐sensitive applications.
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1 Introduction With ongoing rapid development in the area of bioelectronics, small and light weight biocompatible electrodes are in high demand for biomedical implantable devices such as, cardiac pacemakers, neurostimulators, and cochlear implants. [ 1, 2, 3, 4, 5 ] The supercapacitors (SCs) with promising properties, such as, high power density, unlimited cycle life, eco‐friendliness, and low‐temperature charging ability, can become the ideal energy storage devices for medical electronics. Furthermore, implantable SCs can also be used for monitoring biological and electrophysiological information inside the human and mammalian body. Therefore, an ideal SC should be biocompatible, miniature in size, and it should possess high energy and power densities. Additionally, the electrochemical stability of biocompatible SC‐based systems under physiological conditions is another advantage. However, current SCs consisting of biocompatible electrode materials are not able to meet all the aforementioned requirements in a single package. To date, several studies have reported the fabrication of biocompatible energy storage devices using advanced carbon‐based materials such as, graphene nanosheets, carbon nanotubes, and fibers with a focus on maximizing the electric double layer (EDL) capacitive properties. [ 6, 7, 8 ] However, bioelectrical applications of these materials are limited in terms of the energy surface support including low energy density per unit volume or mass. [ 9 ] It is important to further note that these anisotropic carbon structures implicate decreased biocompatibility in terms of cellular growth, proliferation, and differentiation. [ 10 ] In this regard, recently reported 2D transition metal carbides and nitrides (MXenes) are considered among booming materials because of their application in multiple fields. [ 11, 12, 13, 14 ] The main energy‐storage characteristics of the MXene nanosheets are excellent volumetric capacitance and energy density, which are vital for size‐sensitive applications. MXene materials possess hydrophilic surfaces and are selectively etched from their MAX phase structures, where “M” represents one of the early transition metals (e. g. , titanium, tantalum, niobium, or zirconium), “A” denotes one of the A‐group elements (e. g. , aluminum, silicon, or phosphorus) and “X” denotes either carbon or nitrogen. MXene structures possess tremendous physicochemical, electrical, optical and biological properties which have enabled them to be exquisitely used in various electrical and biological applications. [ 15, 16, 17 ] Given this, MXene is a suitable class of material for preparing electrodes for energy storage due to its superior metallic conductivity, electrochemical stability and higher volumetric capacitance compared to the conventional carbon‐based electrode materials. [ 13, 18 ] In previous studies, the exfoliation of different forms of MXene such as titanium carbide (Ti 3 C 2 T x ) and niobium carbide (Nb 2 CT x ) has been reported for application in lithium‐ion (Li‐ion) batteries, capacitors and regenerative medicine. [ 19, 20, 21 ] Recently, a new composition of MXene nanosheets, tantalum carbide (Ta 4 C 3 T x ) has been reported. [ 22, 23 ] Tantalum (Ta)‐based materials are well‐known due to their excellent bio‐functionality when compared to other MXene counterparts such as titanium (Ti)‐ and niobium (Nb)‐based composites. [ 24, 25, 26, 27 ] Moreover, the physical properties of Ta, including density, electrical conductivity and mechanical Young's modulus are higher in comparison to Ti and Nb. In fact, the exfoliated tantalum carbide MXene nanosheets have been recently reported as efficient nanoplatforms for cancer therapy. [ 28 ] However, in most of the MXene based previous reports, hydrofluoric acid (HF) was used as an etchant to remove aluminum (Al) from the MAX phase. Unfortunately, HF is highly corrosive and causes significant burns and toxicity upon contact, ingestion, or inhalation. More importantly, it leads to the formation of fluorine bonds in the end product that substantially decreases the electrochemical activity of MXene‐based electrodes. Additionally, the application of fluorine‐containing etchants leads to interaction between the remaining Al from the MAX phase structure and inert fluorine terminals, which is environmentally harmful during application of MXene‐based materials. Therefore, fluorine containing etchants could potentially decrease the biocompatibility and volumetric capacitance properties of MXene‐based energy storage electrodes. Furthermore, functionalization of oxygen groups in the structure of 2D MXene nanosheets is reported to enhance the electrochemical performance of the composite. [ 18 ] The available literature on oxidized Ti 3 C 2 T x and Nb 2 CT x MXene composites revealed that the formation of crystalline transition metal oxide particles up to a few hundred nanometers in size enhanced the surface activity of nanosheets. [ 13, 29, 30 ] Therefore, it is conceptualized that the formation of oxygen‐containing functional groups on the surface of Ta 4 C 3 T x will promote its electrochemical properties when used as supercapacitor electrode. Furthermore, growth of metal oxide crystals on the surface of MXene prevents restacking of exfoliated MXene due to van der Waals interaction; therefore, functionalization of oxygen‐containing groups preserves the electrochemical performance of delaminated MXene. In the current study, we present, for the first time, oxidized fluorine‐free exfoliation of tantalum carbide MAX phase to synthesize a new Ta 4 C 3 T x MXene‐tantalum oxide (TTO) hybrid structure. Our study demonstrates that TTO has excellent potential to be used as a bioelectrode material for long‐term supercapacitor applications. This new electrode has long‐term electrochemical stability, excellent volumetric capacitance, and high energy/power densities and charging rate. Our findings confirm that the TTO hybrid structure is highly biocompatible with different human cell types. The energy density of the TTO electrode outperforms almost all the existing biomaterial‐based electrodes. In addition, volumetric capacitance of the TTO is significantly higher than the majority of previously reported organic/inorganic biocompatible electrodes. This novel TTO nanostructure may act as a favorite electrode material for future applications in size‐sensitive biomedical energy storage devices. 2 Results and Discussion 2. 1 Fabrication of Ta 4 C 3 T x MXene‐Tantalum Oxide Hybrid Structure We employed an innovative fluorine‐free etching method to prepare TTO hybrid nanostructure from raw and bulk material. Recently, application of an alkaline‐induced method for removal of Al from the Ti 3 AlC 2 MAX phase to synthesize Ti 3 C 2 T x MXene was reported. [ 31 ] However, the main challenge in etching the Al layer from MAX phase in alkaline media is the blocked/slow kinetic reactions due to the formation of unwanted oxide/hydroxide layers on the MXene surface. [ 32 ] To address this, we utilized a modified alkaline‐based etching method to prepare the TTO via a two‐step acidic/alkaline (HCl/KOH) treatment. Briefly, Ta 4 AlC 3 MAX phase powder was treated sequentially with 6 m hydrochloric acid (HCl) solution and 6 m potassium hydroxide (KOH) solution to synthesize an exfoliated nanocomposite. This led to formation of multilayered oxidized Ta 4 C 3 T x nanosheets anchored with Ta‐oxide particles. These were subsequently subjected to thermal treatment at 220 °C under moderate air heating for further functionalization and oxidation, resulting in the formation of final TTO hybrid structure. The step‐by‐step schematic model for the synthesis and functionalization of the mixed‐dimensional TTO nanocomposite is shown in Figure 1 a. Figure 1 Schematic model and stoichiometry of TTO hybrid structure. Illustration of a) step‐by‐step schematic and b) mechanism of reaction for the fluorine‐free conversion of the Ta 4 AlC 3 MAX phase to surface‐modified Ta 4 C 3 T x MXene nanosheets decorated with tantalum oxide nanoparticles. In our proposed mechanism of reaction (Figure 1b ), the Al‐atoms on the edge and outer surface of Ta 4 AlC 3 MAX phase were rapidly chlorinated using HCl through the production of soluble aluminum chloride (AlCl 3 ). [ 31 ] Therefore, a considerable amount of surface Al atoms are thought to be removed during this step. Subsequent etching treatment in KOH solution further led to lower levels of insoluble aluminum hydroxide [Al(OH) 3 ] and aluminum oxide hydroxide [AlO(OH)] on the surface of the material when compared to the classic alkaline protocol. The lattice‐like features of the Ta‐layers limit the transformation of insoluble Al‐based compounds to soluble aluminate [Al(OH) 4 − ]. [ 31, 32 ] However, the hybrid protocol employed in this study allows continued exfoliation and secondary crystal nucleation of the tantalum carbide material to give rise to oxidized Ta 4 C 3 T x MXene nanosheets. Furthermore, ongoing oxidation of exposed inner Al and Ta atoms by OH − resulted in a higher degree of functionalization with —OH and = O groups. 2. 2 Characterization of Ta 4 C 3 T x MXene‐Tantalum Oxide Hybrid Structure The scanning electron microscopic (SEM) images of the functionalized Ta 4 C 3 T x MXene nanosheets prior to thermal treatment are presented in Figure 2 a. The well‐exfoliated MXene nanosheets are strewn with a considerable number of tantalum oxide nanoparticle clusters. Subsequent thermal treatment at 220 °C for 2 h led to significantly enhanced exfoliation and functionalization of MXene nanosheets with Ta‐oxide nanoparticles, which is described here as a TTO hybrid structure (Figure 2b and Figure S1, Supporting Information). Furthermore, field‐emission SEM images of the oxidized Ta 4 C 3 T x MXene and TTO hybrid structure revealed a slight decrease in the wall‐to‐wall interlayer space of MXene nanosheets after thermal treatment (Figure S2 a, b, Supporting Information). Figure 2 Morphology and microstructural characterization of the synthesized TTO hybrid structure. Scanning electron microscopic (SEM) images of the functionalized a) Ta 4 C 3 T x MXene and b) TTO nanostructure samples after heat treatment at 220 °C for 2 h. Field‐emission SEM observations of TTO hybrid structure reveal effective delamination of layers, anchored by a tantalum oxide particle array. The thermal treatment further improved the oxidization of Ta 4 C 3 T x MXene layers. TEM images of the oxidized c) Ta 4 C 3 T x MXene (inset) and d) TTO (inset) composites. The images show successful organization of TTO hybrid structures. The thermal treatment led to further improvement and well‐defined distribution of Ta‐oxide nanoparticles. High‐resolution TEM images confirmed the presence of two different lattices with d‐spacing of 0. 261 and 0. 338 nm, which is attributed to Ta 4 C 3 T x MXene layers and tantalum oxide composites. e) XPS narrow scan spectra of Ta 4f, O 1s, Al 2p corresponding to Ta 4 AlC 3 MAX phase and oxidized TTO samples after thermal treatment at 220 °C for 2 h, confirming proper extraction of Al from the MAX phase structure with effective exfoliation of MXene nanosheets. The XPS fittings further demonstrate that exfoliated Ta 4 C 3 T x MXene sheets were successfully composited with Ta 2 O 5 ‐TaO 2 particles. The high‐resolution transmission electron microscopy (TEM) of oxidized Ta 4 C 3 T x MXene and the TTO hybrid structure revealed that individual nanoparticles are ≈5 nm in diameter and cluster to form larger decorations seen in the SEM images (Figure 2c, d ). The fast Fourier transform (FFT) analysis of TTO hybrid structure showed d‐spacing lattices of 0. 261 and 0. 368 nm, corresponding to oxidized Ta 4 C 3 T x MXene and tantalum oxide (Ta 2 O 5 ) crystals respectively (Figure 2d and inset). These lattice parameters are in good agreement with previously reported literature on stable Ta 4 C 3 T x MXene and Ta 2 O 5. [ 28, 33, 34, 35 ] This FFT analysis is also congruent with published literature on the coexistence of bulky tantalum carbide (TaC) and Ta 2 O 5, which reported lattice spacing of ≈0. 26 and 0. 38 nm, corresponding to the TaC (111) and Ta 2 O 5 (001) planes, respectively. [ 33, 34, 35 ] The selected area electron diffraction (SAED) patterns of TTO hybrid structure displayed a higher degree of crystalline structure with well‐defined hexagonal planes compared to oxidized Ta 4 AlC 3 MXene samples (Figure S2 c, d, Supporting Information). Together these observations provide robust evidence that the innovative fluorine‐free exfoliation and functionalization protocol employed in the current study has worked successfully to synthesize TTO nanostructure from the Ta 4 AlC 3 MAX phase (Figure S3, Supporting Information). The physicochemical properties of materials were further evaluated by X‐ray photoelectron spectroscopy (XPS), energy‐dispersive X‐ray spectroscopy (EDS), SAED, and X‐ray diffraction (XRD) analysis. XPS was used to characterize the structural transformation of Ta 4 AlC 3 MAX phase to oxidized Ta 4 C 3 T x MXene and TTO hybrid structure. Comparison of survey spectra showed a significant change in the elemental composition of materials during the synthesis process (Figure S4 a, Supporting Information). In particular, characteristic Al 2p peaks of Ta 4 AlC 3 MAX phase were significantly decreased in the TTO hybrid structure. High resolution XPS spectra of TTO hybrid structure showed well‐defined characteristics of Ta 4 C 3 T x MXene. A comparison between Al 2p spectra of the Ta 4 AlC 3 MAX phase and TTO confirmed effective elimination of Al layer, with more than 84% of elemental Al removed during the hybrid synthesis process (Figure 2e, Figure S5 and Table S1, Supporting Information). This degree of exfoliation is highly comparable with the widely‐used HF etching method for exfoliation. [ 22 ] This is confirmed by the spectra of Ta 4f, which revealed two main peaks of TaC (4f 5/2 and 4f 7/2) at binding energies of 18 and 24 eV attributed to exfoliated Ta 4 C 3 T x MXene nanosheets. The Ta 4f, O 1s, and C 1s spectra demonstrated a change in the degree of exfoliation and functionalization when converting Ta 4 AlC 3 MAX phase to TTO hybrid structure (Figure 2e, Figure S4 b–d and Table S1, Supporting Information). Ta 4 C 3 O x decreased from 41. 7% to 27. 6%, and Ta 4 C 3 (OH) x increased from 30. 8% to 31. 6%. Additionally, these spectra revealed an approximately 20% decrease in the surface atomic ratio of Ta‐C bonds and an increase of 10. 43% in the surface atomic ratio of Ta‐O bonds in the TTO hybrid structure. The TTO hybrid structure contains two lateral species of Ta 4+ and Ta 5+ as the main tantalum oxide crystals of Ta 2 O 5 and TaO 2 at the binding energy of 22 to 27 eV in the Ta 4f spectrum. Taken together, these findings confirm the formation of tantalum oxide during the synthetic process. Finally, the successful synthesis of TTO hybrid structure using the fluorine‐free process was further corroborated by EDS and XRD. The EDS elemental analysis demonstrated successful extraction of Al from the structure of Ta 4 AlC 3 MAX phase with a decrease in the atomic percentage of Al from 20. 57% to 11. 31% (Figures S6 and S7, Supporting Information). The average weight percentage of Al similarly decreased from 12. 64% to 4. 43%. Concurrently, the atomic percentage of oxygen increased from 20. 03% to 30. 05%. The histogram also confirmed the absence of F, Cl, and K in the final composition of the TTO hybrid structure. The XRD pattern of the Ta 4 AlC 3 bulk material was typical with standard peaks at their expected 2θ values (Figure S8, Supporting Information). [ 22 ] In agreement with the previous observations, peaks originating from the aluminum‐containing MAX phase were significantly decreased after fluorine‐free etching and exfoliation by the HCl/KOH process. One of the Ta 4 AlC 3 peaks at around 16. 5° 2θ was entirely removed in the TTO hybrid structure. Additionally a minor contamination peak ascribed to Ta 2 C with the reflection at 2θ ≈ 50° was absent in the XRD spectrum of TTO hybrid structure. [ 36 ] Furthermore, a newly emerged (002) peak at around 7° 2θ corresponds to an aluminum‐etched tantalum carbide‐tantalum oxide material with the enlarged lattice parameters. Additional downshifts are also observed in the XRD spectra of the TTO hybrid structure due to increased carbon lattice spacing after the acid/alkaline treatment. Finally, characteristic small peaks corresponding to tantalum oxide particles anchored on the MXene surface were also observed in the structure of the TTO. Together these findings support the successful synthesis and functionalization of layered TTO hybrid structure using a fluorine‐free exfoliation and functionalization protocol. 2. 3 Specific Surface Area of Ta 4 C 3 T x MXene‐Tantalum Oxide Hybrid Structure The surface area of carbon‐based nanomaterials is an important determinant of their electrochemical properties. The specific surface area of Ta 4 AlC 3 MAX phase, oxidized Ta 4 C 3 T x MXene, and TTO hybrid structure was determined using Brunauer–Emmett–Teller (BET) nitrogen adsorption isotherms ( Figure 3 a ). The specific surface area of Ta 4 AlC 3 MAX phase, oxidized Ta 4 C 3 T x MXene, and TTO hybrid structures were 1. 29, 41. 79, and 51. 02 m 2 g −1 respectively. There was a 40‐fold increase (approximately) in the surface area from Ta 4 AlC 3 MAX phase to oxidized Ta 4 C 3 T x MXene, which is a result of Al etching and formation of a porous MXene structure. [ 37, 38, 39 ] The specific surface area of the TTO hybrid structure is ≈20% higher than that of the oxidized Ta 4 C 3 T x MXene and can be attributed to higher levels of Ta‐oxide nanoparticles on the surface of the oxidized MXene material. This high surface area can be readily detected by assessing the optical properties of the TTO hybrid structure. The aqueous suspension of TTO hybrid structure (50 µg mL −1 ) exhibited high degrees of autoflorescence at several wavelengths across the visible spectrum (Figure S9, Supporting Information). Figure 3 Specific surface area measurement using Brunauer–Emmett–Teller analysis. a) N 2 adsorption‐desorption isotherm curves of the Ta 4 AlC 3 MAX phase, oxidized Ta 4 C 3 T x MXene, and TTO hybrid structure. The BET data depicted that specific surface area of the materials was 1. 29, 41. 79, and 51. 02 m 2 g −1, respectively. b) Pore size distribution of the MAX phase, oxidized MXene, and TTO hybrid structure. As shown, the average pore diameter of the MAX phase was decreased about four‐fold in TTO nanostructure. Consistent with Barrett‐Joyner‐Halenda theory, the total pore volume was increased by ≈14‐fold in the TTO hybrid structure when compared with Ta 4 AlC 3 MAX phase. The average pore diameter, however, decreased from 81. 15 nm in Ta 4 AlC 3 MAX phase to 32. 25 nm in oxidized Ta 4 C 3 T x MXene and 24. 42 nm in TTO hybrid structure (Figure 3b ). The dramatic increase in specific surface area during the synthesis process can thus be explained by a significant increase in the overall porosity of the TTO hybrid structure through formation of new micro‐ and mesopores during the hybrid acid/alkali and thermal treatments. 2. 4 Electrochemical Properties of Ta 4 C 3 T x MXene‐Tantalum Oxide Hybrid Electrode The electrochemical properties of the TTO hybrid structure were characterized by two‐electrode system. The TTO hybrid structure and oxidized Ta 4 C 3 T x MXene based electrodes were fabricated using a 8:1:2 weight ratio of MXene material, Super P carbon black and polyvinylidene fluoride (PVDF). The cyclic voltammetry (CV) and galvanostatic charge/discharge (GCD) were measured in the presence of a polyvinyl alcohol/phosphoric acid (PVA/H 3 PO 4 ) solid electrolyte. The CV profiles of the oxidized Ta 4 C 3 T x MXene electrode and the TTO hybrid structure electrode were quasi‐rectangular at scan rates ranging from 1 to 100 mV s −1 with near mirror symmetry among all CV profiles, indicating that the majority of capacitance is associated with the electric double layer capacitance (EDLC) mechanism ( Figure 4 a, b ). Additionally, GCD curves of the oxidized Ta 4 C 3 T x MXene electrode and the TTO hybrid structure electrode feature nearly triangular shapes with extremely low internal resistance at the beginning of the discharge curve (Figure 4c, d ). This reflects the pseudocapacitive nature of the oxidized MXene and TTO hybrid structure electrodes. [ 40 ] In particular, the TTO hybrid structure exhibited significantly greater capacitance when compared with the oxidized Ta 4 C 3 T x MXene material. The observed specific capacitance values for TTO were more than twofold higher than oxidized Ta 4 C 3 T x MXene material at the same scan rate or specific current (Figure 4e, f ). In fact, the new TTO hybrid structure electrode showed higher volumetric capacitance than most of the recently reported biomaterials‐based electrodes (Table S2, Supporting Information). Figure 4 Electrical and electrochemical measurements of fabricated TTO hybrid structure electrode. a) Cyclic voltammetry curves of the oxidized Ta 4 C 3 T x MXene electrode and b) TTO hybrid structure electrode at different scan rates in PVA/H 3 PO 4 solid electrolyte after 10 000 cycles of the two‐electrode experiment. c) The galvanostatic charge/discharge (GCD) curves of the oxidized Ta 4 C 3 T x MXene electrode and d) the TTO hybrid structure electrode. e) Specific capacitance for both electrodes and volumetric capacitance of TTO hybrid structure electrode at different scan rates and f) different specific currents. g) The Nyquist plot of the oxidized Ta 4 C 3 T x MXene electrode and TTO hybrid structure electrode. The inset shows the electrical equivalent circuit. h) CV curve of the TTO hybrid structure electrode at 100 mV s −1. The pink and blue areas show the direct contributions of the capacitive and diffusion mechanisms respectively. i) Capacity contribution from capacitive and diffusion‐controlled kinetic processes at different scan rates for the TTO hybrid structure electrodes. The electrochemical impedance spectroscopy analysis was used to investigate the ion‐diffusion/transport resistance of the TTO electrode in the frequency range of 0. 01 Hz to 200 kHz at open‐circuit‐potential measurements. The impedance spectra of the oxidized Ta 4 C 3 T x MXene electrode and the TTO electrode form a small arc and a spike at the higher and lower frequency regions respectively. The TTO hybrid structure electrode clearly showed lower electrolyte resistance than the oxidized Ta 4 C 3 T x MXene electrode (Figure 4g ). This may be due to the higher concentration of tantalum oxide nanoparticles on the surface of the TTO. As a result, in the equivalent electrical circuit model of the TTO hybrid structure electrode, there is a Warburg impedance element (W) representing linear diffusion under semi‐infinite conditions. Furthermore, the solution resistance ( R s ) signifies the electrolyte resistance and a constant phase‐element ( C dl ) obtained by EDLC of the TTO (Figure 4g and inset). The electrochemical mechanism measurements of TTO hybrid structure were further investigated based on the following equation: (1) i = k 1 ν + k 2 ν 1 / 2 This equation includes two scan rate‐related terms at a fixed scan rate and potential. The k 1 ν term is attributed to the current density contributed by the fast‐kinetics process in both electric double‐layer capacitance and Faraday pseudocapacitance. [ 41 ] The k 2 v 1/2 term is related to the current density contributed by a slow diffusion‐controlled process. [ 41 ] Both constants, k 1 and k 2 are obtained from a different form of the above equation in a log‐log plot, as shown below. (2) i / ν 1 / 2 = k 1 ν 1 / 2 + k 2 The contribution of the fast‐kinetics process for various sweep rates is shown in Figure 4h, i and Figure S10, Supporting Information. The decoupling result for the TTO hybrid structure electrode at 100 mV s −1 offers a remarkable fast‐kinetics contribution of 87. 0% (Figure 4h ). Furthermore, while the fast capacitance remained unaffected by scan rate, the contribution of slow capacitance increases at lower scan rates (Figure 4i ). This effect is attributable to the oxygen‐containing terminal groups on the surface of the TTO hybrid structure, which facilitates electrochemical Faradaic reactions to result in a greater contribution of diffusion‐controlled mechanisms at these lower scan rates. These findings were confirmed with measurements from a three‐electrode system using the same electrolyte against an Ag/AgCl reference electrode. The results further elucidated the pseudocapacitance effect of Ta 2 O 5 and other oxygen‐containing functional groups in acidic electrolytes. The TTO hybrid structure electrode exhibited EDLC behavior with quasi‐pseudocapacitance behavior (Figure S11 a, Supporting Information). The oxidized surface of TTO may significantly increase the wettability and enlarge the ion‐accessible surface area, facilitating rapid ion diffusion/transportation into the internal pores. The possible redox reaction of Ta 4 C 3 T x MXene in an acidic medium is depicted in Equation ( 3 ). [ 42 ] (3) Ta 4 C 3 O x OH y + δ H + + δ e − ↔ Ta 4 C 3 O x − δ OH y − δ The possible redox reaction of Ta 2 O 5 on Ta 4 C 3 T x MXene in an acidic medium is shown in Equation ( 4 ). [ 43 ] (4) Ta 2 O 5 + β H + + β e − ↔ H β Ta 2 O 5 The normalized capacitance value for the two‐electrode system included contributions of two TTO hybrid structure electrodes, while the three‐electrode system used a single TTO hybrid structure electrode. Therefore, the values derived from three‐electrode measurements should be approximately two times of the values obtained from two‐electrode measurements (Figure 4e ; Figure S11 b, Supporting Information). [ 38 ] Thus, the obtained results from the three‐electrode and two‐electrode systems of the TTO hybrid structure are in agreement with each other. 2. 5 Ragone Plot of Ta 4 C 3 T x MXene‐Tantalum Oxide Electrode Supercapacitor Next, we wanted to evaluate and compare the performance of the TTO hybrid structure electrode with literature‐reported organic/inorganic energy storage materials for bio‐implantable applications. Ideally, the best bioelectrode materials for implantable supercapacitors should possess excellent energy and power density in a single product, while having low individual component toxicity in case of damage and uncontrolled failure. In particular, currently used lithium‐ion batteries and their toxic electrolytes in cardiac pacemakers, neurostimulators, cochlear implants, and spinal cord stimulators have been reported to seriously jeopardize patient safety in cases of premature failure. [ 44, 45 ] In the current study, Ragone plots are presented to compare the volumetric performance of the TTO hybrid structure supercapacitor with several previously‐reported biocompatible supercapacitors and MXene‐based electrode supercapacitors ( Figure 5 a, b ). The TTO hybrid structure electrodes were packaged into a symmetric supercapacitor using copper current collectors and a PVA/H 3 PO 4 solid electrolyte. Our data confirm that the TTO hybrid structure supercapacitor was superior to almost all other bioelectrode materials, including EDL capacitors, biophilized graphene oxide modified protein electrode supercapacitor, aluminum electrolytic capacitors, and lithium‐ion thin film batteries (Figure 5a and Table S2, Supporting Information). [ 46 ] Additionally, the fluorine‐free TTO supercapacitor has competitive energy and power densities with other previously published MXene‐based supercapacitors (Figure 5b ). Importantly, its performance also exceeds that of many currently reported non‐MXene carbon‐based electrodes, including graphene/CNT nanocomposites (165 F cm −3 ), [ 47 ] graphene‐based electrodes (260 F cm −3 ), [ 48 ] activated carbons (60–100 F cm −3 ), [ 49, 50 ] and carbide‐derived carbons (180 F cm −3 ). [ 51, 52 ] Finally, the TTO electrode possesses excellent areal efficiency when compared with other recently reported supercapacitor electrode materials (Figure S12, Supporting Information). Figure 5 Comparison of TTO supercapacitor with some of the previously reported organic and inorganic electrode materials. a) Ragone plots comparing the performance of TTO with electrical double layer (EDL) capacitors (35 and 50 mF, 300 µF/3 V), graphene oxide modified protein electrode supercapacitor, aluminum electrolytic capacitor (12 µA h/3. 3 V) and lithium‐ion thin film battery (LiTF, 500 µA h/5 V). [ 45, 46, 47, 48, 49, 50, 51, 52 ] Energy density/power density of the TTO is significantly higher than above‐mentioned electrodes. b) Ragone plots comparing energy and power densities of the TTO hybrid structure supercapacitor to PANI/Ti 3 C 2 T x, [ 42 ] RuO 2 /MXene yarn, [ 43 ] Mo 1. 33 C/PEDOT:PSS, [ 43 ] Ti 3 C 2 Tx/RGO, [ 53 ] and MXene/NiCo‐LDHs. [ 54 ] c) Supercapacitor cycling stability, volumetric capacitance retention, and charge–discharge cycle at a current density of 1 A g −1. d) The image shows an LED powered by the TTO supercapacitor electrodes. Zoom‐view panel shows the schematic view of TTO‐based solid‐state supercapacitor containing PVA/H 3 PO 4 gel electrolyte. The picture demonstrates that TTO supercapacitor was able to successfully power the LED. In addition to its excellent capacitance properties, the TTO displayed excellent long‐term stability over 10 000 cycles with capacitance retention maintained over 90% of the initial performance (Figure 5c ). The capacitance retention was stabilized after 1500 cycles, with only an additional 2. 4% decrease in capacitance over the subsequent 8500 cycles. Together, these data confirm that the TTO hybrid structure supercapacitor synthesized in the current study possesses long cycle stability, excellent volumetric capacitance and high charge/discharge rate performance. As a proof‐of‐concept, the energy storage performance of the TTO hybrid structure supercapacitor was functionally assessed by connecting it to a light‐emitting diode (LED); the experiment demonstrated that the symmetric TTO was able to successfully power the LED (Figure 5d ). 2. 6 Biocompatibility of Ta 4 C 3 T x MXene‐Tantalum Oxide We also investigated the biocompatibility of oxidized TTO electrode with human induced pluripotent stem cells (hiPSC)‐derived cardiomyocytes, neural progenitor cells (NPCs), and fibroblasts. The hiPSC‐derived cells were obtained using our established differentiation protocols (Figure S13, Supporting Information). [ 55 ] When MXene‐based materials (at a concentration 50 µg mL −1 ) were cocultured with these cells for 24 h, assessment of cytotoxicity using the WST‐1 assay showed that all forms of the Ta 4 C 3 T x MXenes were compatible with cardiomyocytes, NPCs, and fibroblasts ( Figure 6 a ). Figure 6 Assessment of biocompatibility of the Ta 4 AlC 3 MAX phase and Ta 4 C 3 T x MXene‐tantalum oxides materials with human cells. a) The MAX phase, oxidized Ta 4 C 3 T x, and TTO materials were cocultured with human iPSC‐derived‐ fibroblasts, cardiomyocytes, and neural progenitor cells (NPCs) for 24 h. WST‐1 proliferation assay was performed to evaluate cytocompatibility of materials. Our data demonstrate that MXene was compatible with all three cell types, as coculture with biomaterial did not affect cellular proliferation compared to control group. b) Cytotoxicity evaluation of the MAX phase, oxidized Ta 4 C 3 T x, and TTO hybrid structure was assessed by LDH release after coculturing with human MSC for 24 h. LDH data show no significant difference among different MXene groups and the control group. c) LIVE/DEAD assay was performed using the fluorescent dye to assess biocompatibility of human MSC with the material. After coculture with different forms of MXene, MSC were stained with Calcein (for live cells, green) and EthD‐1 (for dead cells, red). Images were captured using Nikon Ti‐2 fluorescent microscope. No significant difference in viability between different groups was detected. ( n =3–4 per group). (“ns” = statistically no significant difference, * = p < 0. 05 and ** = p < 0. 01). Additionally, TTO‐based bio‐electrodes may also be beneficial in the development and post‐delivery monitoring of functional cell‐based tissue constructs. Bone marrow‐derived mesenchymal stem cells (MSC), a commonly used cell type in tissue engineering, were found to be biocompatible with all forms of Ta 4 C 3 T x MXene‐based samples used in this study. When cocultured with materials for 72 h, assessment of cytotoxicity by LDH assay showed excellent residual viability of cells in all groups as there were no significant differences in cytotoxicity between different MXene groups and the control group (Figure 6b ). On the other hand, MAX phase, due to the presence of Al, demonstrated lower biocompatibility to the cells. Representative images captured using a fluorescence‐based live/dead assay also confirmed this finding (Figure 6c ). These data confirm that oxidized TTO electrode is biocompatible and it can be used for implantable bioelectronic devices and tissue engineering applications. 3 Conclusion This study reported the first fluorine‐free synthesis and application of Ta 4 C 3 T x MXene‐tantalum oxides hybrid structure material for energy storage applications. The TTO‐based electrode showed excellent volumetric capacitance compared to previously reported biocompatible electrodes. Furthermore, the TTO hybrid structure is highly biocompatible with different types of human cells, which is highly beneficial for future applications in bioelectronics and biosensors. Finally, when assembled into a symmetric supercapacitor, the TTO hybrid structure material possessed high energy/power densities and long‐term cyclability. The stability of TTO electrodes was estimated to be over 10 000 cycles. 4 Experimental Section Fluorine‐Free Synthesis of Oxidized Ta 4 C 3 T x and Ta 4 C 3 T x MXene‐Tantalum Oxide Hybrid Structure Ta 4 C 3 T x MXene nanosheets were partially exfoliated using hydrochloric acid (HCl). To do so, Ta 4 AlC 3 MAX Phase powder was incubated in 6 m solution of HCl in water at 37 °C for 72 h in a shaking incubator at 260 rpm. The precipitates were collected after washing with ultrapure distilled water by spinning at 5000 rpm for 5 min each. The precipitates were freeze dried for 48 h and subsequently air dried at 60 °C. The complete etching, exfoliation, and surface modification of the obtained material (dry powder) was achieved by treating it with potassium hydroxide (KOH, 6 m ) at room temperature for 90 h. The edge exfoliation of specimens was obtained by centrifugation at 5000 rpm followed by several washing steps and vacuum lyophilization (−80 and −54 °C) for 48 h to avoid uncontrolled oxidization. The powder was then double‐dried in an atmospheric oven at 50 °C for 48 h. The resultant nanocomposite obtained at this step was labeled as oxidized Ta 4 C 3 T x at room temperature (22 °C). For further functionalization and oxidation, the treated Ta 4 C 3 T x nanosheets were subjected to thermal treatment at 220 °C for 2 h under moderate air heating and labeled as TTO hybrid structure (220 °C). Physicochemical Characterization The structural properties of materials were characterized using an FEI Nova NanoSEM 450 (Thermo Fisher Scientific), FEI Talos F200X S/TEM (Thermo Fisher Scientific), Thermo Nicolet Nexus 870, and Kratos Axis Ultra XPS at the Manitoba Institute of Materials (MIM), University of Manitoba, Winnipeg, Canada. The SEM samples were mounted on pin stubs using carbon tape and coated with a gold‐palladium (Au‐Pd) coating to enable high magnifications. XRD peaks of powdered samples were collected in the range from 5 to 80° 2‐theta using continuous scan mode with a scan rate of 3° min −1 and report interval of 0. 05°. The measurement of specific surface area of the materials was determined by the BET analysis. Electrode Fabrication, Electrical, and Electrochemical Measurements The TTO electrodes were fabricated using the following procedures. Each TTO hybrid structure electrode was synthesized using 8:1:2 ratio of TTO hybrid structure (160 mg), Super P carbon black (20 mg), and PVDF (40 mg) in N ‐methyl‐2‐pyrrolidone solvent. The slurry prepared by mixing these components was brushed on a carbon paper and pressed after drying in a vacuum oven at 70 °C for 24 h. The capacitance properties of the prepared TTO electrodes were characterized by using two‐ and three electrode systems at room temperature. Cyclic voltammetry and the constant current charge–discharge measurements were performed on Autolab electrochemical workstation (PGSTAT302 N model) and CH Instrument 640E Bipotentiostat. The specific capacitance for cyclic voltammetry‐based measurement was calculated according to the following equation: (5) C = 1 2 v m Δ V ∫ I d V where C, I, ν, m, and Δ V are the specific capacitance, current, scan rate (V s −1 ), weight of electrode, and scanning potential window, respectively. The constant current charge–discharge test was performed for specific capacitance. Values were calculated using the following equation: (6) C = I Δ t m Δ V where I, Δ t, and Δ V are respectively discharge current, discharge time, and discharge potential window. [ 38 ] The energy density and power density of the device were calculated as E = C (Δ V ) 2 /7. 2 and P = 3600 E /Δ t, where E and P, are energy and power densities, respectively. Using the density of the packed electrolyte (2. 3 × 10 −3 kg cm −3 ), the volumetric energy density ( E vol ), and volumetric power density ( P vol ) were calculated using the following equations: (7) E vol W h c m − 3 = E × ρ (8) P vol = 3600 × E vol Δ t The CV and GCD measurements were performed with all solid‐state two‐electrode system in the presence of PVA and H 3 PO 4 gel electrolyte. For this, a gel electrolyte solution was prepared by mixing 10 g of PVA and 10 g of H 3 PO 4 in 100 mL of deionized water at 85 °C. The solution was incubated in an oven at 40 °C for a week to solidify. A thin layer of the PVA/H 3 PO 4 electrolyte was sandwiched between two active electrodes of the same size and mass and subjected to a hot‐pressing step at 10. 9 psi. The copper (Cu) foils were attached to the other side of the TTO electrode to be used as the current collector. For Ragone plots the total mass of the packaged TTO‐based supercapacitor (959 mg) was used, including TTO hybrid structure electrodes, electrolyte, carbon blacks, and PVDF to calculate and evaluate the total energy or power densities. The capacitance of TTO‐based electrodes was further characterized by using a three‐electrode system at room temperature with phosphoric acid (H 3 PO 4 ) solution as electrolyte. The platinum (Pt) and silver/silver chloride (Ag/AgCl) were used in the experiment as the counter electrode and the reference electrode, respectively. Density Measurement The gravimetric capacitance of the TTO electrode is converted to volumetric capacitance by Archimedes’ Principle. The density of TTO electrode was calculated using the following equation: (9) ρ = W a W ax ρ ax where W a, W ax, ρ, and ρ ax are the weight of the sample in air, weight of the sample in the auxiliary liquid of known density, the density of the sample and auxiliary liquid, respectively. Assessment of Energy Storage Performance of Ta 4 C 3 T x MXene‐Tantalum Oxide Supercapacitor The energy storage performance of TTO was assessed using a light‐emitting diode (LED). The TTO electrode in the supercapacitor was charged and was connected in an LED output and the performance was observed. Induced Pluripotent Stem Cells Generation, Culture, and Differentiation hiPSCs were generated from peripheral blood mononuclear cells (PBMC) isolated from human blood (collected from healthy individuals). [ 21 ] All protocols were approved by the University of Manitoba Health Research Ethics Board (B2015:025, HS18974). To reprogram PBMCs toward iPSCs a commercial reprogramming kit (CytoTune‐iPS 2. 0 Sendai Reprogramming Kit) was used (A16517, ThermoFisher Scientific, US). The detailed procedure is described in previously published studies. [ 56, 57 ] The hiPSCs were cultured in TeSR‐E8 (0 5990, STEMCELL Technologies) on Geltrex (A1413302, Gibco) and allowed to differentiate toward fibroblast, cardiomyocytes (CMs) and NPCs using our previously published protocols. [ 55 ] Briefly, embryoid bodies (EBs) were prepared in suspension in low attachment plates (174 932, Thermo Scientific) and plated onto gelatin‐coated plates on Day 8 (PMEF‐CFL‐P1, EMD Millipore). They were allowed to differentiate spontaneously toward fibroblasts, which were manually dissected from the culture plate and characterized by staining for HSP47 (ab77609, Abcam) and FSP (ab11333, Abcam) (Figure S13, Supporting Information). The cell populations were enriched over several passages to ensure a pure fibroblast population. The hiPSCs were differentiated to cardiomyocytes using following protocol: iPSCs (>passage 20) were passaged onto Geltrex‐coated plates using Versene Solution (15 040 066, Gibco) and grown till the cells reached ≈85% confluency. The medium was replaced with CDM3, consisting of RPMI 1640 (61 870 036, Gibco) supplemented with 500 µg mL −1 recombinant human albumin (A9731, Sigma‐Aldrich), and 213 µg mL −1 L‐ascorbic acid 2‐phosphate (A8960, Sigma‐Aldrich). The culture medium was replaced on alternate days (48 h). At days 0–2, the medium was supplemented with 6 µM of the glycogen synthase kinase 3‐β inhibitor CHIR99021 (SML1046, Sigma‐Aldrich). On day 2, the medium was changed to CDM3 supplemented with 2 µM of the Wnt inhibitor‐ Wnt‐C59 (5. 00496. 0001, CalBiochem). Day 4 onward, the cells were cultured in medium without the inhibitors. The beating cells were observed from day 7. At day 10, medium was replaced with RPMI 1640 without glucose (11 879 020, Gibco), 500 µg mL −1 recombinant human albumin, and 213 µg mL −1 L‐ascorbic acid 2‐phosphate supplemented with 4 mM L‐lactic acid (71 720, Sigma‐Aldrich) for metabolic enrichment of cardiomyocytes. The cardiomyocytes were characterized by immunofluorescence staining for sarcomeric alpha actin (ab9465, Abcam) and MYH6 (ab50967, Abcam). The differentiation of iPSCs toward NPCs was carried out using EB method by initiating the treatment with the TGF‐beta/Smad inhibitor SB 431 542 (16‐141, Torcis) for 2 days. The EBs were then plated on polyornithine‐coated plates on day 5 (A004, Merck Millipore). The neural rosettes were visually identified at day 7–10. After that, the rosettes were excised and grown on polyornithine‐coated plates in STEMPRO NSC SFM kit (A1050901, Gibco). The NPC characterization was carried out by immunostaining for NESTIN (sc‐23927, Santa Cruz Biotechnology) and PAX6 (sc‐81649, Santa Cruz Biotechnology). Human Mesenchymal Stem Cells Culture Human bone marrow derived MSC were purchased from Lonza (PT 2501, CA10064‐080) and cultured in low‐glucose DMEM (10 567 014, Gibco) according to previously published protocols. [ 58 ] Cell Proliferation Assay Human iPSC‐derived fibroblasts, cardiomyocytes, and NPCs were plated on 96‐well plates and cocultured with or without the raw MAX phase and oxidized TTO composites at a concentration of 50 µg mL −1 for 24 h. Then, the cell proliferation was assessed using the WST‐1 proliferation kit (K301, BioVision). Assessment of Cytotoxicity To assess cytotoxicity, human MSC were cultured with different forms of MXene for 24 h at a concentration of 50 µg mL −1. To evaluate cytotoxicity, LDH release from damaged cells (if any) was measured in the supernatant using a Cytotoxicity Detection Kit (MK401, Takara Bio). Assessment of Cellular Viability using LIVE/DEAD Assay To assess the effect of TTO MXene on cell viability, LIVE/DEAD assay was performed. Briefly, human MSC (2 × 10 5 ) were plated on 96‐well plates and cocultured with or without the MAX phase and TTO composites for 72 h. The cells were then stained using a LIVE/DEAD Viability/Cytotoxicity Kit (L3224, Invitrogen) for 30 min and then visualized using Nikon Eclipse Ti‐2 fluorescence microscope. Calcein was detected using the GFP Filter (Ex480/Em535) and EthD‐1 was detected using the TRITC Filter (Ex540/Em605). Statistical Analysis Data were reported as mean ± SD unless otherwise specified. Comparison of data between multiple groups was performed using one‐way analysis of variance (ANOVA) followed by Tukey's post‐hoc multiple comparison test, and analysis between two groups was made using an unpaired Student's t ‐test (two‐tailed). Statistical analysis was performed using GraphPad Prism 8. 0. 1 (San Diego, USA). Statistical significance was defined as p < 0. 05. Conflict of Interest The authors declare no conflict of interest. Author Contributions The study was conceptualized and designed by A. R. , A. A. , and S. D. A. R. , A. A. , G. L. S. , W. Y. , and Y. C. carried out the experiments and acquired the data. A. R. , A. A. , W. Y. , A. A. P. , and S. D. interpreted the data and performed statistical analysis. A. R. , A. A. , W. Y. , and S. D. designed the figures. A. R. , A. A. , and S. D. drafted the manuscript. All authors have read and approved the final manuscript. Supporting information Supporting Information Click here for additional data file.
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10. 1002/adfm. 202100850
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Advanced functional materials
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Self-Oxygenation of Tissues Orchestrates Full-Thickness Vascularization of Living Implants
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Bioengineering of tissues and organs has the potential to generate functional replacement organs. However, achieving the full-thickness vascularization that is required for long-term survival of living implants has remained a grand challenge, especially for clinically sized implants. During the pre-vascular phase, implanted engineered tissues are forced to metabolically rely on the diffusion of nutrients from adjacent host-tissue, which for larger living implants results in anoxia, cell death, and ultimately implant failure. Here it is reported that this challenge can be addressed by engineering self-oxygenating tissues, which is achieved via the incorporation of hydrophobic oxygen-generating micromaterials into engineered tissues. Self-oxygenation of tissues transforms anoxic stresses into hypoxic stimulation in a homogenous and tissue size-independent manner. The in situ elevation of oxygen tension enables the sustained production of high quantities of angiogenic factors by implanted cells, which are offered a metabolically protected pro-angiogenic microenvironment. Numerical simulations predict that self-oxygenation of living tissues will effectively orchestrate rapid full-thickness vascularization of implanted tissues, which is empirically confirmed via in vivo experimentation. Self-oxygenation of tissues thus represents a novel, effective, and widely applicable strategy to enable the vascularization living implants, which is expected to advance organ transplantation and regenerative medicine applications.
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1. Introduction The bioengineering of tissues and organs has the potential to deliver functional replacement organs. [ 1, 2 ] However, the creation of viable clinically sized implants has remained a grand challenge. Current tissue engineering strategies, although suitable for the development of small living tissues that remain viable upon implantation, are incompatible with maintaining the viability of larger implants. Specifically, while small implants can rely on host-to-implant diffusion of oxygen and nutrients, large implants inevitably suffer from nutrient diffusion limitations that induce anoxia-induced cell death and sub-sequent implant failure. [ 3, 4 ] Various strategies have been explored to accelerate the implant’s access to the host’s oxygen, metabolites, and nutrients. These strategies commonly rely on endowing implants with angiogenic growth factors, (pro)angiogenic cells, and/or bioprinting of vascular structures to establish functional anastomosis. [ 5 – 7 ] Although successful in small implants, these angiogenic and vasculogenic strategies are ineffective for larger implants due to the required fundamental biological processes that impose absolute limits on the rate in which functional blood vessels can be established within implants, which amongst other includes the migration speed and proliferation rate of endothelial cells. Moreover, incorporation of angiogenic cells or growth factors will further increase the burden of an already metabolically deprived tissue by having to support additional cells as well as the angiogenic process. [ 8, 9 ] Furthermore, before achieving functional anastomosis, implanted engineered tissues will be rapidly depleted of oxygen resulting in hypoxic (<5% dissolved oxygen) followed by anoxic (<0. 5% dissolved oxygen) microenvironments. [ 10, 11 ] While acute hypoxia is conducive for regenerative responses such as angiogenesis, anoxia inhibits these regenerative processes and causes cell death. [ 12 ] A living implant’s ability to vascularize and survive is therefore inversely correlated with the duration and intensity of its metabolic deprivation, which currently is dictated predominantly by the implant’s size. [ 13 ] Consequently, the historically explored metabolism-dependent pro-angiogenic and vasculogenic strategies have not been able to offer an effective solution to orchestrate full-thickness vascularization of clinically sized living implants. [ 14 ] We hypothesized that the incorporation of a controlled oxygen release system into a living implant would facilitate its vascularization by transforming the cytotoxic anoxia into pro-angiogenic hypoxia. Advantageously, the proposed metabolic support is only transiently required as it is designed to bridge the implant’s prevascular period. To achieve this feat, a system for the controlled and prolonged release of oxygen is needed. Several oxygen releasing and generating strategies have already been developed, for example, hemoglobin and myoglobin sub-stitutes, polymer-based oxygen carriers, perfluorocarbons, and solid peroxides. [ 15 – 19 ] These compounds and materials have the ability to bind, dissolve, or generate physiologically relevant quantities of oxygen, but typically release their payloads within minutes to hours. In contrast, alleviating the oxygen deprivation during the prevascular phase of large implants demands release periods of multiple days to weeks. [ 20, 21 ] Interestingly, week-long oxygen generation could potentially be achieved by encapsulating a solid peroxide such as calcium peroxide (CaO 2 ; CPO) in hydrophobic bulk material. [ 22 ] Specifically, a material’s hydrophobic nature could be used to limit the exposure of encapsulated solid peroxides to water molecules, which effectively grants control over the hydrolysis rate of solid peroxides and hence oxygen release. Unfortunately, encapsulating cells in hydrophobic materials associates with poor outcomes in terms of cell survival and tissue formation. In short, the development of biocompatible long-term oxygen generating biomaterial capable of sustaining the survival and function of clinically sized tissues has remained a challenge. Here, we report the development of novel micromaterials that function as hydrophobic oxygen-generators (HOGs), which can be readily integrated within engineered tissues. HOGs allowed for safe and long-term oxygen release that granted control over the in situ oxygen tension by alleviating anoxic stress throughout the entire volume of the engineered tissues. Most notably, the incorporation of HOGs orchestrated the rapid vascularization of implanted living tissues. Specifically, metabolic support enabled the survival of encapsulated stem cells, which facilitated the production of high quantities of vascular endothelial growth factor (VEGF). Uniquely, we here report that mildly increasing the oxygen tension in implanted tissues drives full-thickness vascularization of living implants. HOGs therefore represent a promising solution to maintain the survival of clinically sized engineered tissues and facilitate their functional integration within the receiving host. 2. Results 2. 1. Fabrication and Characterization of HOGs The stoichiometry of CPO’s oxygen generating ability is described by the following chemical equations: (1) CaO 2 + 2 H 2 O → Ca ( OH ) 2 + H 2 O 2 (2) 2 H 2 O 2 → O 2 + 2 H 2 O From these equations, it is evident that oxygen release from CPO can thus be rate limited by controlling the exposure of CPO to water molecules. We, therefore, theorized that the encapsulation of CPO in hydrophobic microparticles—named HOGs—would result in a controlled and long-term oxygen release system ( Figure 1A ). HOG microparticles were produced using a controllable and scalable water-in-oil-in-water double emulsion synthesis method. Pristine polycaprolactone (PCL) micromaterials were synthesized as a non-oxygen generating control. Size analysis revealed that PCL MPs, pristine CPO, and HOGs associated with a diameter of 2. 3 ± 0. 9, 1. 2 ± 0. 2, and 4. 6 ± 1. 2 μm, respectively ( Figure 1B ). Varying the final CPO concentration within HOGs did not measurably affect the size of the synthesized microparticles. With an encapsulation efficiency of >90%, CPO in concentrations up to 20% (wt. /wt. ) did not significantly affect the encapsulation efficiency and CPO loading capacity in HOGs ( Figure S2, Supporting Information ). Micromorphological analysis of CPO, PCL, and HOGs by scanning electron microscopy (SEM) confirmed the spherical and intact nature of PCL and HOG micromaterials ( Figure 1C – E ). Energy dispersive X-ray spectrometry (EDS) based elemental mapping suggested relatively higher oxygen content in HOGs as compared to pristine PCL micromaterials ( Figure 1F – H ). Incubating HOGs under standard culture conditions for twelve days resulted in a notable decline in relative oxygen content in HOGs, which suggested the release of oxygen from the HOGs into the culture medium ( Figure S3, Supporting Information ). Moreover, elemental mapping confirmed the presence of calcium in HOGs, which corroborated the presence of CPO in HOGs as pristine PCL micromaterials did not contain any detectable levels of calcium ( Figure 1I - K ). Indeed, Alizarin Red S stained HOGs and not pristine PCL microparticles ( Figure 1L ). The hydrophobic nature of PCL was confirmed via water contact angle (WCA) measurements, which was diminished in a dose-dependent manner following CPO incorporation ( Figure 1M and Figure S4, Supporting Information ). The pristine PCL polymer is hydrophobic in nature. PCL sheets exhibited WCA of 94 ± 4° and water drops could remain on the surface of PCL for a long time ( Figure S4, Supporting Information EDX). The incorporation of CPO in PCL polymer showed a significant decrease in WCA pointing toward reduced hydrophobicity and thus increased hydrophilicity. As the hydrophobicity of PCL was inversely correlated with the concentration of CPO particles, the presence of CPO on the surface of HOGs was assumed. To confirm that CPO was efficiently encapsulated within the PCL microparticles, HOGs were cut open using a focused ion beam (FIB) and imaged using SEM, which revealed homogeneously distributed darker regions within the HOGs that matched CPO’s particle size ( Figure 1N ). High-resolution elemental mapping of FIB-SEM cut HOGs confirmed that these regions were CPO nanoparticles based on the spatial distribution patterns of calcium, oxygen, and carbon ( Figure 1N ). Indeed, the vast majority of CPO was detected within the bulk of the HOGs. 2. 2. Engineering Self-Oxygenating Hydrogels using HOGs HOGs were used to create a novel class of self-oxygenating hydrogels. Specifically, HOGs containing various CPO concentrations (e. g. , 0%, 1%, 2. 5%, 5%, and 10% (wt. /vol. )) were mixed in gelatin methacryloyl (GelMA) polymer solutions resulting in self-oxygenating hydrogel, which were photocrosslinked in molds to form 4 mm wide and 1 mm thick cylindrical multiphase polymeric microcomposites. As GelMA is composed of gelatin, an amphiphilic polymer that has shown excellent properties as a stabilizer for drug delivery and tissue engineering applications it was expected to allow for homogeneous dispersion of HOGs with GelMA hydrogels. [ 23 ] Realizing a homogenous distribution of HOGs in the hydrophilic bulk material is essential to achieve consistent and equally distributed oxygenation of engineered materials. Homogeneous distribution CPO or HOGs particles was demonstrated in freeze-dried GelMA hydrogel constructs using SEM analysis ( Figure 2A – C ), which was corroborated using EDS based spatial mapping of calcium ( Figure 2D ). Incorporation of pristine CPO dose-dependently increased the swelling of the hydrogel and decreased its overall elastic modulus ( Figure 2E and Figure S5, Supporting Information ). The higher swelling ratio of GelMA hydrogel containing CPO, as compared to those loaded by same amount of HOGs ( Figure S5, Supporting Information ), suggests that the hydrophilic nature of CPO and the porosities of hydrogel structure had an effect on the absorption and retention of aqueous media, respectively. The porosities of GelMA with 20% CPO (wt. /vol. ) and GelMA with 20% HOGs (wt. /vol. ) were analyzed for Figure 2B, C by ImageJ software and calculated to be 83 ± 2. 6% and 74 ± 0. 8%, respectively. This could potentially be explained by the intense burst release of O 2 producing micro-bubbles that cause a more porous structure in case of CPO containing hydrogel as compared to a hydrogel containing HOGs. This might also partially explain the difference in the mechanical properties between these hydrogels ( Figure 2E ). This was most likely mediated via the quenching of the photoexcited photoinitiator via the fast-paced generation of oxygen of non-encapsulated CPO. [ 24 ] Indeed, incorporation of HOGs minimized the swelling compared to the GelMA hydrogel with CPO only, albeit, without decreasing the hydrogel’s elastic modulus. In fact, the elastic modulus increased in a dose-dependent manner becoming over an order of magnitude higher for 10% HOG-GelMA as compared to pristine GelMA. Previous reports of multiphase polymeric microcomposites in which polymeric micromaterials where incorporated in hydrogel systems have made similar observations. [ 25, 26 ] Moreover, this avoidance of decrease in elastic modulus suggested that HOGs indeed mitigated the initial bulk release of oxygen by reducing the rate of the hydrolytic conversion of CPO into calcium hydroxide and hydrogen peroxide. To validate this, hydrogen peroxide release kinetics of CPO-GelMA and HOG-GelMA were determined for up to twelve days. These release kinetics are in line with CPO being exposed directly to water and are in agreement with previous literature. [ 17 ] In contrast, HOG-GelMA hydrogels released notably lower levels of hydrogen peroxide at initial stages, which slowly and stably declined over time ( Figure 2G ). HOGs thus offer cytoprotection from the oxidative stresses by partially shielding cells from hydrogen peroxide-induced toxicity. As the hydrogen peroxide is subsequently converted to oxygen and water, the oxygen concentration in CPO-GelMA and HOG-GelMA hydrogels was determined in a time-resolved manner inside a continually purged modular hypoxia incubator. While CPO-GelMA hydrogels associated with a short-lived oxygen generation, HOG-GelMA hydrogels were characterized by a sustained release of oxygen for nearly two weeks ( Figure 2H, I ). Although the CPO or HOG concentration did not have a major effect on the release duration of hydrogen peroxide or oxygen, it positively correlated with the released quantities. Consequently, varying the CPO concentration allowed for tuning of the oxygen concentration without affecting its release duration. Overcoming diffusion limits for survival of any cellularized implanted tissue, [ 27 ] the observed oxygen release kinetics of HOGs-GelMA were seen to be in line with those required to bridge the prevascular phase survival of living implants. [ 28 – 30 ] 2. 3. Self-Oxygenating Tissues Survive and Express High Levels of VEGF under Anoxia The ability of hydrophobic micromaterials to protect cells from the cytotoxic effects of the generated hydrogen peroxide was investigated. Self-oxygenating tissues were produced by encapsulating human mesenchymal stem cells (hMSCs) in GelMA hydrogels containing up to 10% of CPO or HOGs. It was observed that under normoxic conditions the addition of CPO to GelMA resulted in massive cell death, even at concentrations as low as 2. 5% CPO ( Figure S7A, B, Supporting Information ). In contrast, self-oxygenating tissues based on HOG incorporation demonstrated only a minor and transient decrease in cell survival, even at concentrations of HOGs as high as 10% ( Figure S7C, D ). This vast improvement in cell survival confirmed HOGs’ cytoprotective capacity, which was most likely achieved by rate-limiting the buildup of hydrogen peroxide. We next determined whether HOGs also offered cytoprotection in terms of cell survival under anoxic culture conditions by metabolically supporting the encapsulated cells. Control tissues cultured under anoxic culture conditions rapidly and progressively reduced the percent of viable cells: after a single day only <30% of cells survived, which further declined to <10% after 12 days of culture ( Figure 3 ). CPO-based self-oxygenating tissues associated with further decreased cell survival rates, which was most likely caused by the rapid generation of hydrogen peroxide. In fact, virtually none of the encapsulated hMSCs remained viable after six days of anoxic culture ( Figure S8, Supporting Information ). Importantly, self-oxygenating tissues based on HOG-GelMA formulations enabled ≈80% cells to remain viable over a period of 12 days under the hostile anoxic culture conditions, even at concentrations as low as 2. 5% HOGs ( Figure 3A, B ). This survival rate of all engineered tissues was near-identical regardless of anoxic or normoxic culture condition, which indicated that HOGs could effectively avoid anoxia-induced cell death throughout the entire volume of self-oxygenating tissues. We then reasoned that the ability of HOGs to engineer consistently hypoxic microenvironments could be leveraged to stimulate encapsulated cells to express physiologically relevant levels of angiogenic factors such as VEGF. To test this hypothesis, self-oxygenating tissues composed of hMSCs encapsulated in 2. 5% HOG-GelMA cultured under anoxic conditions were compared to HOG-free engineered tissues cultured under various oxygen tensions (e. g. , normoxic, hypoxic, and anoxic). As expected, the culture’s oxygen concentration was inversely correlated with the total amount of secreted VEGF ( Figure 3C ). However, the total amount of VEGF secreted by control (e. g. , non-oxygenating) tissues cultured under anoxic conditions progressively dwindled over time to levels even below those produced by control tissues cultured under normoxic conditions. In contrast, self-oxygenating tissues consistently produced high levels of VEGF over a period of six days. This remarkable difference between self-oxygenating tissues and control tissues could be explained by a progressive cell loss in control tissues caused by the anoxic culture conditions ( Figure 3D ). Indeed, the amount of VEGF secreted per cell had remained consistently high in control tissues despite the cell loss ( Figure 3E ), which was in line with previous reported literature. [ 31, 32 ] However, the angiogenic behavior of the implant is not determined by the per cell production of angiogenic factors, but rather via the formation of chemotactic gradients based on the total amount of angiogenic factors. Consequently, we hypothesized that implanted self-oxygenating tissues would demonstrate improved angiogenic behavior as compared to their conventional counterparts. 2. 4. HOG Mediated Self-Oxygenation of Engineered Tissues Orchestrates Full-Thickness Vascularization In Vivo To predict the in vivo behavior of self-oxygenating tissues, numerical simulations were performed to model their behavior following their implantation. Specifically, a previously developed set of partial differential equations of the taxis-diffusion-reaction type were used to describe various key cellular processes including oxygen tension fluctuations, angiogenic factor production, angiogenesis, cell proliferation, and cell death. [ 33 ] This model has previously been used to corroborate that oxygen tension was a key factor in terms of implant failure in clinically sized tissues. [ 34 ] Here, we have adapted the numerical model by incorporating the approximated oxygen generation kinetics of 1%, 2%, and 4% of CPO or HOGs wt. /v. into the simulated tissue implant ( Figure S9, Supporting Information ). In the control tissues, the simulated oxygen tension rapidly dropped to anoxic levels, which predicted poor VEGF production and angiogenesis ( Figure 4A ). Although CPO incorporation elevated the oxygen tension substantially during the first 24 h following implantation, the simulation predicted a neglectable improvement in implant vascularization due to oxygen depletion after three days of implantation. In contrast, simulated self-oxygenating tissue implants effectively prevented the formation of anoxic regions, which maintained cell survival and resulted in potent VEGF production. This associated with improved and accelerated vascularization that enabled full-thickness vascularization of the simulated self-oxygenating tissues. In addition, increasing the HOG concentration to 4% was predicted to have little additional effects while reducing the HOG concentration to 1% would mitigate the beneficial effect ( Figure S10A, B, Supporting Information ). To empirically validate these simulations driven predictions, engineered tissues composed 50 mm 3 GelMA hydrogels containing 2×10 6 cells mL −1 hMSCs and 0%, 1%, or 2% HOGs were subcutaneously implanted in rats for seven days. The macroscopic analysis revealed a striking difference in gross appearance between the explanted tissues. Where control tissues were characterized by a dark color that was reminiscent of necrotic tissue, self-oxygenating tissues possessed a healthy-looking light-pink appearance with complex patterns that suggested vascular structures ( Figure 4B ). Whole explant live/dead analysis confirmed that while control tissues were mostly void of living cells, self-oxygenating tissue were still populated by numerous living cells and appeared to indeed possess intricate networks of vessels. To study host–graft interaction, midsagittal sections of the explants were analyzed using histology and immunohistochemistry. Hematoxylin and eosin (H&E) ( Figure 4C ) and Masson’s trichrome (TRI) ( Figure 4D ) staining confirmed that the core of control tissues was void of cells. Moreover, control tissues demonstrated low levels of biomaterial degradation, absence of vessels, and contained cellular infiltration in their peripheral regions, which were covered in a fibrous capsule. In marked contrast, self-oxygenating tissues contained a high number of cells, lacked a fibrous capsule, and possessed an abundance of vessel-like structures throughout the entire implant. We postulated that these distinct differences between conventional and self-oxygenating tissues could be explained by the retention or loss of hMSCs and their associated bioactivity, immunoregulatory, and tissue remodeling properties. Indeed, Human nuclear antigen (HNA) staining confirmed that the hMSCs of self-oxygenating tissue remained present in the implant and were predominantly located around the vessels ( Figure 4E ). In addition, while control tissues displayed an intense C-reactive protein (CRP) staining at their peripheral regions, self-oxygenating tissues presented notably milder level. This reduction in local inflammation could potentially have contributed to the improved survival and retention of the implanted hMSCs. Tunel staining confirmed that locally inflamed regions indeed associated with increased levels of apoptosis, which was present at a notably lower level in self-oxygenating tissues containing 2% of HOGs ( Figure 4F ). Self-oxygenation of tissues also appeared to contribute to the integration with host tissue as the observed vessels were HNA-and CD31+, which suggested ingrowth of the host’s vascular system ( Figure 4G ). While control tissues exclusively presented CD31+ vessels in their peripheral regions, self-oxygenating tissues contained CD31+ vessels throughout their volume. Semi-quantification of stained tissue sections corroborated the inverse correlation between the HOG concentration and the hydrogel degradation as well as the number of apoptotic cells ( Figure 4H, I ). In line with these findings, the HOG concentration positively correlated with an increase in the number of CD31+ vessels, especially those located at the core of the implant ( Figure 4J ). HOGs also contributed to blood vessel maturation within the living implant in terms of vessel diameter ( Figure 4K ). HOGs thus enabled full-thickness implant vascularization, which was most likely mediated via the VEGF secreted by metabolically supported hMSCs. 3. Discussion We here report on HOG micromaterial mediated self-oxygenation of engineered tissues, which enabled the survival and full-thickness vascularization of implanted tissues. Specifically, HOGs alleviated anoxic stress throughout the bulk of engineered tissues over a prolonged period of time, which allowed the tissue implant to survive independently of the oxygen that diffused from the host organism. Encapsulating hydrophilic CPO in hydrophobic PCL endowed the resulting HOGs with an amphiphilic nature, albeit, with dominant hydrophobic characteristics owing to their PCL bulk. This composite based micromaterial design extended the release of oxygen-derived from CPO hydrolysis by several folds as compared to pristine CPO. While the hydrophobic nature of PCL slows down hydrolysis, PCL’s relatively high oxygen permeability coefficient allows for the near-instant release of generated oxygen to surrounding cells and tissues. [ 35 ] HOGs generated sufficient amounts of oxygen in the implanted tissues to transform their own milieu into a hypoxic microenvironment, which enabled control over the living implant’s fate. HOG hydrophobicity can thus be tuned to control CPO hydrolysis and hence determine the rate and duration of oxygen release to cater to the physiologic needs of the tissue of choice. Matching oxygen generation with oxygen consumption is essential to orchestrate implant vascularization as both anoxic and normoxic microenvironments are detrimental to implant vascularization due to massive cell death and low-level production of angiogenic growth factors, respectively. In contrast, we aimed for the engineering of hypoxic microenvironments that allowed for continued cell survival and production of high levels of angiogenic factors, which was achieved using HOGs and associates with rapid full-thickness implant vascularization. Self-oxygenation of engineered tissues via CPO is achieved by generating oxygen via the intermediate generation of hydrogen peroxide, which can induce cytotoxicity. Fortunately, PCL has been shown to shield cells from H 2 O 2 insults up to 150 μm, which represents concentrations many times higher than the levels observed in our study. CPO is known to adversely impact cell survival when present at a concentration above 3% wt. /vol. [ 36 ] This is most likely caused by the resulting change in pH, an increase in calcium ions, or an increase in free radicals. We observed this phenomenon with CPO, but not with HOGs, which remained cytocompatible at substantially higher concentrations. A unique feature of this work is the demonstration that self-oxygenation of tissues orchestrates full-thickness vascularization following implantation via in situ control over implant metabolism. This metabolism supporting strategy distinguishes this work from previous research, which has focused on strategies that are associated with an increase in metabolic burden (e. g. , incorporation of angiogenic cells, vasculogenic cells, or growth factors). [ 37 ] Although metabolically burdening approaches can be effective in small implants, they are ineffective for larger implants. Specifically, implanted clinically sized solid tissues will inherently develop large anoxic and metabolically deprived regions, which not only induce the death of implanted cells, but also thwart the host’s angiogenesis processes. We here introduce the concept of homogenous self-oxygenation of living implants to offer both short-term (e. g. , cell survival) and long-term (full-thickness implant vascularization) protection by metabolically supporting implanted living tissues using HOGs to engineer stably hypoxic microevironments that produce high levels of angiogenic factors such as VEGF. Although the vascularization observed in this study was primarily achieved via angiogenesis, HOGs’ ability to sustain cell survival while driving the production of high levels of VEGF also has the potential to improve implant vascularization for tissue engineering strategies that aim to utilize vasculogenesis. Elevating the local oxygen tension of an implant to hypoxic levels for a prolonged period of time via self-oxygenation thus allows for the successful bridging of an implant’s prevascular period, which has remained a key hurdle in translating engineered tissue into a clinical reality. In particular, self-oxygenation driven vascularization of tissues is particularly suited for applications that require the implantation of a living implant of voluminous sizes with intense and/or intense vascularization of the implant. Such applications include, but are not limited to, engineered tissues to address critically sized bone defects, volumetric muscle loss, kidney failure, chronic liver disorders, and various cardiac pathologies. While a working understanding of how metabolic programming can steer cell fate has been gained in recent years, tools to accurately control the metabolism (e. g. , by controlling the oxygen tension for prolonged period of time) in vivo have remained scarce. [ 38, 39 ] HOGs, therefore, represent a novel enabling tool to steer the behavior of implanted living tissues. Moreover, this work underlines the relevance of including metabolic parameters into the design space of engineered tissues, which has remained an underexplored domain. Indeed, in situ control over the oxygen tension can guide the formation and maturation of engineered tissues via metabolic programming of cell fate. [ 40, 41 ] It is of note that the current study is limited to the use of PCL, however, it is anticipated that alternative hydrophobic materials will also offer extended oxygen generation profiles owing to their ability to limit calcium peroxide hydrolysis mediated homolytic cleavage. The current study has explored the use of HOGs to create self-oxygenating tissues to orchestrate angiogenesis and enable full-thickness vascularization of living implants. However, HOGs could offer potential solutions for numerous additional challenges. Besides supporting the survival of implanted cells, HOGs can also steer the function and fate of implanted cells and host cells that are in close proximity to the implant via metabolic programming by controlling the local oxygen tension. Moreover, HOGs ability to control the oxygen tension in vivo could also be explored to treat pathologies that are adversely affected by prolonged hypoxia such as diabetic ulcers, tumors, and asphyxial cardiac arrest. 4. Conclusion In this study, we demonstrated that oxygen generating micromaterials can be leveraged to improve the vascularization of living implants. Specifically, microencapsulation of oxygen releasing material such as solid peroxides allows for prolonged oxygen release duration while simultaneously minimizing the cytotoxicity of oxygen generation process. We demonstrated that pristine CPO in cellularized hydrogels is toxic due to the nature of CPO hydrolysis that associates with an intense rapid burst release of hydrogen peroxide and oxygen, which lasts only for a few days. In contrast, microencapsulation of solid peroxides within a hydrophobic material lowered the hydrogen peroxide concentration and extended the duration of oxygen generation. HOGs were demonstrated to allow for self-oxygenation of tissues to alleviate the detrimental anoxic stresses within the cellular constructs by creating a pro-angiogenic hypoxic microenvironment, which associated with production of high levels of VEGF over a prolonged period of time owing to continued cellular survival. In vivo experimentation demonstrated that self-oxygenation of engineered living tissues potently orchestrated rapid full-thickness vascularization and offered control over the in vivo oxygen tension for prolonged periods of time, which thus offered a method to metabolically program the fate and behavior of both implant and host. HOGs therefore represent a promising stepping-stone toward the development of clinically sized tissues for regenerative applications. 5. Experimental Section Preparation of CPO-Laden PCL Microparticles: A double emulsion synthesis method of water-in-oil-in-water (w/o/w) was used to prepare CPO-laden PCL microparticles. Briefly, 3 mL of 10% (wt. /vol. ) PCL (Mw = 80 000, Sigma Aldrich) solution was prepared in dichloromethane (DCM, CH 2 Cl 2, purity ≥99. 8%, Sigma Aldrich). To obtain 0%, 5%, 10%, 15%, 20%, and 30% (wt. /wt. ) CPO in PCL microparticles, the predetermined concentration of CPO (CaO 2, Sigma Aldrich) in 1 mL of ethanol (CH 3 CH 2 OH, purity ≥99%, Sigma Aldrich) was added. The solution was emulsified using ultrasonication (Qsonica sonicators, Newtown borough, CT, USA) for 3 min with a 1 s on/off pulse at 30% amplitude. Subsequently, 10 mL of 3% (wt. /vol. ) poly(vinyl alcohol) (Mw = 89 000–98 000, Sigma Aldrich) in distilled water was added to the PCL-CPO solution and ultrasonicated for additional 5 min with a 1 s on/off pulse at 30% amplitude at room temperature. After ultrasonication, the solution was stirred to dry off the solvents at room temperature for 24 h and concentrated by centrifugation (Eppendorf 5702, Germany) at 10, 000 rpm for 5 min. The microparticle pellet was washed three times with Dulbecco’s phosphate-buffered saline (DPBS, Gibco, Carlsbad, CA, USA) to remove any residual additives. The resultant pellet was lyophilized and stored in a dry and cold place until further use. As control, PCL particles were prepared without CPO employing the same protocol. Encapsulation Efficiency/Loading Capacity of CPO in HOGs: Encapsulation efficiency and loading capacity of CPO in HOGs was determined as a percentage of the amount of CPO lost after every washing step to the initial CPO concentration used. Every wash performed was kept at 60 °C for 24 h to remove water, ethanol, and additives. Any residual PCL was removed by adding 3 mL DCM to the dried residue and stirred on a magnetic stirrer at 500 rpm for 30 min and filtered through a poly(vinylidene fluoride) membrane filter (durapore membrane, 0. 22 μm pore size, USA). The mass of remained non-encapsulated CPO was measured. Percent encapsulation efficiency and percent loading capacity for CPO in HOGs were calculated as follows: (3) Encapsulation Efficiency ( % ) = Total CPO added - CPO lost in washing Total CPO added × 100 (4) CPO loading capacity = Encapsulated CPO Final mass of HOCs × 100 Wettability of the Surface of CPO Containing PCL Microparticles: 5 mg of pristine PCL or CPO containing PCL microparticles were used to cover a stainless-steel surface. A 20 μL water droplet was placed on the specimen surface, which was microscopically imaged after 30 s and analyzed using ImageJ software (NIH, Bethesda, MD, USA) to quantify the contact angle between the surface and the water droplet. Particle Size Distribution of PCL, CPO, and HOGs: For size distribution, 1 mg of CPO powder, PCL microparticles, or HOGs were dispersed separately in 10 mL of ethanol and vortexed for 5 min to obtain a homogenous solution. 2 mL of each solution was analyzed on particle size distribution using dynamic lighting scattering (DLS, Malvern Nano ZS ZEN3600). Microscopic and Elemental Analyses of Microparticles and Hydrogels: The microstructure of the lyophilized samples of PCL, HOGs, and GelMA scaffolds with and without oxygen-generating agents (e. g. , CPO or HOGs) was evaluated using high-resolution SEM (Zeiss MERLIN HR-SEM) and analyzed on microstructure and porosity using the image-based analysis software ImageJ. To enhance contrast and reduce charging effects, the samples were coated with a 5 nm layer of Pt/Pd alloy SEM analysis. Ultra-high resolution field emission SEM elemental mapping in the form of EDS analysis was performed on intact or FIB (Nova 600 Nanolab DualBeam) edged samples to evaluate the elemental distribution of carbon, calcium, and oxygen. Synthesis of GelMA: Gelatin methacryloyl (GelMA) was synthesized as previously reported. [ 42 ] Briefly, 10% (wt. /vol. ) of powdered gelatin (Sigma Aldrich, Type A, 300 bloom from porcine skin) was dissolved in DPBS (Gibco, USA) and stirred at 60 °C until its complete dissolution. To the mixture, 8% (vol. /vol. ) of methacrylic anhydride (Sigma-Aldrich) was gradually added while constantly stirring the mixture, representing a high methacryloyl substitution. [ 43 ] The reaction was carried at constant stirring for 1 h at 60 °C and 500 rpm. Two volumes of pre-heated (60 °C) DPBS were added to the solution. This was followed by dialysis of the solution mixture using 12–14 kDa cutoff dialysis membranes (Thermo Fisher Scientific) against deionized water. Dialysis was carried out for 1 week to remove any salts and unreacted methacrylic anhydride, after which the solution was filtered, and placed at −80 °C to freeze before lyophilizing. At the end of lyophilization, freeze-dried GelMA was obtained in the form of white porous foam. To determine the degree of methacrylation in the synthesized GelMA, fluoraldehyde o-phthaldialdehyde assay (OPA) was performed. [ 44 ] Briefly, GelMA at 0. 5 mg mL −1 was dissolved in DPBS. To this fluoraldehyde OPA reagent was added at a 1/1 (vol. /vol. ) ratio and allowed to react for 1 min to finish the reaction. Using a microplate reader, fluorescence intensity of the resulting mixture for conversion of amine groups in gelatin to methacrylated gelatin was read at 340 nm/455 nm. Gelatin as standard ( I standard ) and DPBS as blank ( I blank ) were used to calculate the degree of methyacrylation (DoM) by following the equation: (5) DoM = 1 − ( I sample − I blank ) / ( I standard − I blank ) Mechanical Properties OMP Encapsulated GelMA Hydrogels: The mechanical properties of particle-encapsulated GelMA hydrogels were conducted by a rotational parallel-plate rheometer (AR-G2 rheometer, TA instruments, USA) at room temperature. Disk shaped samples (n = 3 per group) with a thickness of 1. 5 mm were placed between two 20 mm diameter plates and subjected to a frequency sweep at 0. 1–10 Hz. The experiments were performed in a displacement-control mode for 1% shear strain. Storage (G′) and loss (G″) moduli were calculated against the applied frequency ( Figure S6, Supporting Information ). Expansion of hMSCs: Bone marrow-derived Human Mesenchymal Stem Cells (hMSCs) that were purchased from Lonza (Allendale, NJ), were cultured in low glucose-DMEM media containing 10% fetal bovine serum, 100 IU mL −1 penicillin-streptomycin, and 1 ng mL −1 basic fibroblast growth factor (bFGF) in a humidified incubator at 37 °C with 5% CO 2. Cell culture media was refreshed every third day. hMSCs were used for experiments between their 3rd and 5th passage. Fabrication of Self-Oxygenating Tissues: 10% (wt. /vol. ) GelMA was dissolved in DPBS solution at room temperature. Pristine CPO and HOGs were added to the GelMA solution at different concentrations (1%, 2. 5%, 5%, 10%, and 20% (wt. /vol. )). Additionally, 0. 25% (wt. /vol. ) photoinitiator (Irgacure 2959, Ciba Specialty Chemicals) in DPBS was added to all polymer solutions to allow for UV-initiated photocrosslinking. The resulting solutions were poured into 8 mm wide and 1 mm thick cylindrical molds of polydimethoxysilane (Sigma Aldrich). The solutions were then exposed to UV light (60 s at 2. 5 mW cm −2 ) to form disk-shaped self-oxygenating hydrogels. GelMA hydrogels without CPO or HOGs were used as controls. To create self-oxygenating tissues, GelMA hydrogels containing various concentrations of CPO and HOGs were prepared using low glucose-DMEM cell culture medium supplemented with 0. 25% (w/v) photoinitiator. Trypsinized hMSCs were suspended in CPO-GelMA and HOG-GelMA prepolymer solutions at a cell density of 3 × 10 6 cells mL −1. Cell-laden hydrogel precursor solutions (50 μL) were pipetted onto 3-(trimethoxysilyl) propyl methacrylate coated glass and crosslinked via UV light exposure (850 mW power with 8. 5 cm working distance) for 40 s. The hydrogels were cultured in a bi-weekly refreshed DMEM-low glucose medium under normoxic (21% O 2, 5% CO 2, and 74% N 2 ) and hypoxic (1% O 2, 5% CO 2, and 94% N 2 ) culture conditions. Water Uptake of Hydrogel Constructs: The weight of GelMA, CPO-GelMA, and HOG-GelMA was determined and represented as W 0. Each sample was immersed in DPBS at 37 °C for 24 h. The hydrogel was gently removed from the DPBS solution and after blotting extra DPBS with paper tissue, it was weighed (W t ). The swelling ratio (S. R. ) of each sample was calculated according to the following equation: (6) S. R. = ( W t − W 0 ) / W 0 × 100 W 0 is the initial weight of the hydrogel and W t the weight of the hydrogel after 24 h of incubation. Quantification of Hydrogen Peroxide and Oxygen Release: All experiments in hypoxia condition were performed by a hypoxic chamber (Stem Cell Technology, USA), with the O 2 concentration controlled by an electrode-based ISO-OXY 2 oxygen sensor (World Precision Instruments) and maintained at 2–3%. The closed-system of oxygen chamber could not exchange any material, oxygen, or matter with the surroundings ( Figure S1, Supporting Information ). GelMA hydrogel (1 mL) constructs with different amounts of CPO and HOGs (0%, 1%, 2. 5%, 5%, 10%, and 20% (wt. /vol. ) were individually placed into the wells of a 24-well plate containing 3 mL of deoxygenated DPBS for up to 12 days. 50 μL of media was collected daily and quantitated on its hydrogen peroxide concentration using an Amplex red hydrogen peroxide assay kit (Thermo Fisher Scientific, USA) in the presence of peroxidase from Horseradish type VI (HRP, Sigma Aldrich, 250 U mg −1, Sigma-Aldrich) according to the manufacturer’s protocol. To this, the mixture of 100 μ m Amplex red and 0. 25 U mL −1 HRP in distilled water was prepared. Since Amplex red reagent is not able to detect exact concentration of hydrogen peroxide below 10 μ m, plasma-treated CPO- and HOGs-laden hydrogels were diluted 200 times before adding to reagent. 50 μL of the Amplex Red/HRP reagent was mixed with 200 μL of the 200×-diluted plasma-treated samples in a 96-well plate and incubated for 1 h. The resulting fluorescence signal was determined out using a Tecan well-plate reader (BioTek instruments, Ink, USA) using λ ex = 560/20 nm for excitation wavelengths and λ em = 590/20 nm for emission wavelengths. To determine the oxygen generation profiles, 24 well plate containing GelMA hydrogel constructs with different amounts of CPO and HOGs (0%, 1%, 2. 5%, 5%, 10%, and 20% (wt. /vol. ) in the GelMA solution) were placed in a near-airtight glovebox (Stem Cell Technology, USA) that was continuously purged with nitrogen gas ( Figure S1, Supporting Information ). The O 2 concentration of the media within the 24 well plates was measured using an electrode-based ISO-OXY 2 oxygen sensor (World Precision Instruments) for up to twelve days. Cell Survival in CPO-GelMA and HOGs-GelMA Hydrogels: To evaluate cell survival, 1 × 10 6 hMSCs were encapsulated in 1 mL of 10% (w/v) GelMA hydrogels. Cell viability was evaluated by staining hMSCs encapsulated in GelMA hydrogel using Live/Dead assay (Invitrogen, USA) according to the manufacturer’s instructions after 1, 3, 6, and 12 days of culture. Stained cells were visualized and microphotographed using fluorescence microscopy (Zeiss, Axio Observer A. 1). The metabolic activity of encapsulated cells was assessed using PrestoBlue assay (Invitrogen, USA) according to the manufacturer’s protocol and measured using a BioTek Synergy 2 plate reader. Simulations for Oxygen Release: Numerical simulations were performed based on a previously validated multiscale in silico model of bone fracture healing. [ 45 – 46 ] The hybrid framework combines partial differential equations at the tissue level to model the key processes of tissue formation and implant integration. The hybrid framework is based on an agent-based description at the cellular level that simulates developing vasculature with discrete endothelial cells, including the intracellular Dll4-Notch signaling in every endothelial cell. [ 46 ] The development of discrete vascular trees, which serve as a nutrient source, is determined by sprouting, vascular growth followed by anastomosis at the cellular level. The sprouting of host vasculature is modeled by capturing the intracellular levels of VEGFR-2, active VEGFR-2, effective active VEGFR-2, Notch1, active Notch1, effective active Notch1, Dll4, and actin. The rules that capture the lateral inhibition mechanism during tip cell selection were adapted from a previously developed agent-based model. [ 47 ] At the tissue level, the tissue formation and implant integration is described as a spatiotemporal variation of 10 continuous variables, which was previously described for simulation of the bone fracture healing process. [ 45 ] In short, the considered cell types (i. e. , progenitor cells and fibroblasts) can migrate, proliferate, and secrete growth factors (i. e. , VEGF). Tissues, nutrients, and blood vessels are modeled in separate spaces and can thus “co-exist” in the same location. Here, we adapted the previously developed model (36) by changing the geometry, the boundary conditions, and the VEGF-production rate of the fibroblasts (non-dimensional of 10 in comparison to 1 previously) to emulate the oxygen-beads set-up. CPO and HOGs were modeled as oxygen sources, adapting the generation rate to approximate the experimentally measured oxygen generation kinetics of the CPO or HOGs: (7) B e a d r e l e a s e = G n ⋅ H n 6 n 6 + H n 6 ⋅ e − τ. t with non-dimensional values G n = 2. 2, H n = 0. 11, and τ = 0. 5 for CPO and τ = 0. 05 for HOGs. The beads were randomly positioned in the simulation domain at 1%, 2%, or 4% of the simulation area and the angiogenic sprouts could grow into the simulated tissue implant from all sides ( Figure S7, Supporting Information ). Progenitor cells and fibroblasts could also migrate into the tissue implant from the surrounding tissues (non-dimensional Dirichlet boundary condition of 0. 01 during the first 3 days). The implant was initialized at 3. 7% oxygen tension, fibrous matrix (non-dimensional IC mf = 0. 1), and progenitor cells at 10% of 10% of the carrying capacity (non-dimensional ICmsc = 0. 1). All simulations were run multiple times to account for the stochastic bead placement and were checked for consistency. The simulation code is available upon request. VEGF and DNA Quantifications for Angiogenesis: To determine the amount of VEGF secreted by the engineered tissues, GelMA hydrogels containing 3·× 10 6 cells/mL of hMSCs and 2. 5% HOGs were cultured under anoxic conditions while GelMA hydrogels without HOGs were cultured under normoxic, hypoxic, and anoxic conditions. All hydrogels were cultured in DMEM low-glucose media containing 10% fetal bovine serum and 100 IU mL −1 penicillin-streptomycin for six days. After 1, 3, and 6 days of culture, the media’s VEGF levels were quantitated using a Human VEGF ELISA Kit (BioAim Scientific Inc. , Cat. No: 1 010 015) according to manufacturer’s instructions. The amount of VEGF was normalized for the hydrogel’s DNA content, which was quantified using the CyQuant kit (Invitrogen, Cat. No: C7026) according to the manufacturer’s instructions. Subcutaneously Implantation of Self-Oxygenation Tissues: To evaluate the in vivo behavior of self-oxygenating tissues, hydrogels composed of various concentrations (0%, 1%, or 2%) of HOGs, 2 × 10 6 cells mL −1 of hMSCs, and 10% (w/v) GelMA were fabricated. Hemispherical hydrogels (1 mm thickness and 8 mm diameter) were subcutaneously implanted in the backs on 12-week old nude rats (Charles River) based on a previously established protocol (2017N000114) by the Institute’s Committee on Animal Care. Briefly, anesthesia was induced and maintained with isoflurane in spontaneously breathing rats. After seven days, the implants were surgically extracted and frozen in OCT compound. The frozen samples were sectioned in 5 μm using a cryomicrotome. Sections were stained with TUNEL assay, DAPI based cell quantification, CRP, and HNA to assess survival and hypoxic stress of hMSCs. Also, vascular network formation was analyzed by staining for CD31. Evaluation of sample integration and degradation was performed by H&E and Masson trichrome stain. All stained slides were examined using a Nikon inverted microscope (TS2-LS, Tokyo, Japan). To evaluate the number and diameter of vessels in the explants, three power field images (40×) were taken alongside the borders and inside of the hydrogels. Representative images were then evaluated by two blinded independent analysts using the software Image J (National Institutes of Health, United States). The number of vessels was counted by using the following inclusion criteria: Only the presence of a visible lumen associated with positively labeled CD31 cells was counted as vessels, vessels presenting in the outer border of the high-power field were not counted, structures positively stained with CD31 but without visible lumen were discarded. Staining artifacts were identified by CD31 staining not associated with DAPI co-staining of endothelial cell nuclei; these artifacts were also discarded. The number of vessels per high power field was expressed by the mean number of vessels counted by the two analysts in each representative image. Additionally, to measure vessel diameter, the transversal diameter of counted vessels was measured by using the software Image J, and the diameter of vessels was expressed by the mean of vessel diameter measured by the two analysts in each of the representative images. Statistical Analysis: For statistical relevance, all experiments were performed using, at least, triplicates. All data has been presented by mean ± standard deviation (SD) or standard error mean in each experiment. ANOVA, t -test, and Duncan’s new multiple-range test was used for statistical analysis where appropriate. Supplementary Material supinfo
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10. 1002/adfm. 202108495
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Advanced Functional Materials
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Conversion of 2D MXene to Multi‐Low‐Dimensional GerMXene Superlattice Heterostructure
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Abstract Integration of 2D structures into other low‐dimensional materials results in the development of distinct van der Waals heterostructures (vdWHSs) with enhanced properties. However, obtaining 2D–1D–0D vdWHSs of technologically useful next generation materials, transition‐metal carbide MXene and monoelemental Xene nanosheets in a single superlattice heterostructure is still challenging. Here, the fabrication of a new multidimensional superlattice heterostructure “GerMXene” from exfoliated M 3 X 2 T x MXene and hydrogenated germanane (GeH) crystals, is reported. Direct experimental evidence for conversion of hydrothermally activated titanium carbide MXene (A‐MXene) to GerMXene heterostructure through the rapid and spontaneous formation of titanium germanide (TiGe 2 and Ti 6 Ge 5 ) bonds, is provided. The obtained GerMXene heterostructure possesses enhanced surface properties, aqueous dispersibility, and Dirac signature of embedded GeH nanosheets as well as quantum dots. GerMXene exhibits functional bioactivity, electrical conductivity, and negative surface charge, paving ways for its applications in biomedical field, electronics, and energy storage.
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1 Introduction Recently, there has been growing interest in the development of van der Waals heterostructures (vdWHSs) for applications in biomedical field, electronics and energy storage. [ 1, 2 ] The strategies to synthesize vdWHSs have not been limited to the 2D nanosheets rather there have been significant efforts to tailor the microstructure and properties of different materials. [ 3, 4, 5, 6 ] In fact, several 2D–0D, 2D–1D, and 2D–3D vdWHSs are obtained through chemical interaction of vertical 2D materials with other materials of different dimensionality. [ 7, 8 ] As a result, the obtained mixed‐dimensional vdWHSs provides enhanced surface and material properties to these heterostructures compared to their parent materials. [ 1, 9, 10, 11, 12 ] Ever since the discovery of 2D transition metal carbides “MXene” in 2011, interdisciplinary attention has been paid to the structural and electronic properties of these materials. [ 13, 14, 15, 16 ] The bulked MXene nanosheets (M n +1 X n T x, n = 1–3) which are synthesized by selective etching of their MAX phases, have high electron density at Fermi level with unique physicochemical, electrical and biological properties. [ 17, 18, 19, 20, 21, 22, 23 ] However, increasing concerns related to potential phase‐transformation of MXene nanosheets in aqueous dispersions under ambient conditions cannot be ignored. [ 24, 25, 26 ] Additionally, van der Waals forces available between MXene flakes tend to restack and aggregate the layers, hindering its long‐term applications. [ 27 ] These concerns have significantly limited the growth in the field. Therefore, new strategies to synthesize multidimensional MXene superlattice vdWHSs are urgently needed. Furthermore, the recent development of atomically thin monoelemental 2D materials (Xenes) has expanded the capacity in the field. [ 28, 29, 30, 31, 32, 33 ] These crystalline materials include one of the metalloids of groups 13–16 in their structures and have gained strength due to unique structural and electronic properties. Additionally, the intrinsic passivation of Xenes gives them the capability to stack low‐dimensional materials and construct hybrid or heterostructures. In particular, germanane (GeH) is one of the recently synthesized 2D monoelemental crystals. [ 34, 35, 36 ] GeH is a hydrogen‐terminated multilayer allotrope of germanium and is shown to possess high electron mobility, direct bandgap, active optical properties and quantum spin Hall effect in kagome lattice. [ 31, 37, 38 ] GeH also offers considerable electrical conductivity in layered van der Waals materials. [ 34, 39 ] However, the colloidal dispersibility of GeH nanosheets in aqueous media needs to be significantly improved while retaining its desirable properties. [ 30 ] The fabrication of 2D vdWHSs with distinct geometries has been reported for graphene, boron nitride and transition‐metal dichalcogenides. [ 1, 3, 4, 5, 10, 11, 12, 40, 41 ] However, development of multidimensional superlattice vdWHSs of wonder materials‐ MXenes and Xenes has not been explored yet. The current study reports a versatile strategy to synthesize a unique heterostructure “GerMXene” using 2D MXene and GeH materials. We have employed autoclave treatment method for conversion of Ti 3 C 2 T x MXene nanosheets to a multilow‐dimensional superlattice material. The uniqueness of this study resides in two important findings; first is activation of 2D Ti 3 C 2 T x nanosheets to form activated MXene (A‐MXene) composite displaying the morphology and specific characteristics of MXene nanosheets, quantum dots as well as stable surface titanium oxide nanoparticles in a single material. This has not been reported so far for any MXene in literature. Second, the embedding of GeH derived quantum dots in the atomic structure of A‐MXene to fabricate a superlattice GerMXene heterostructure. As‐synthesized GerMXene exhibits high surface area with concentrated functional groups and negative surface charge, which contributes to its excellent dispersibility and stability in aqueous media, optical absorption and bioactivity. A schematic model depicting the synthesis and atomic structure of multidimensional GerMXene heterostructure is shown in Figure 1 and Figure S1 (Supporting Information). Figure 1 Schematic illustration of the synthesis of GerMXene superlattice heterostructure. a) Schematic model showing the conversion of 2D titanium carbide MXene (Ti 3 C 2 T x ) nanosheets to multilow‐dimensional activated MXene (A‐MXene) crystals. A‐MXene material includes the microstructural characteristics of Ti 3 C 2 T x nanosheets, quantum dots as well as stable surface titanium oxides. The synthesis of hydrogenated 2D germanane (GeH) from germanium bulks and subsequent conversion to 0D GeH quantum dots. b) Illustration of synthesis model of quantum manipulation and functionalization of M 3 X 2 T x MXene nanosheets to form a new multilow‐dimensional superlattice GerMXene heterostructure. In the picture black, gray, pink, beige, light blue and dark blue represent carbon, titanium, aluminum, hydrogen, the lowest and the highest germanium atoms, respectively. Furthermore, the functional groups available on the surface of GerMXene are coded in red and green. 2 Results and Discussion 2. 1 Development of A‐MXene Complex (Ti 3 C 2 O/—OH/—O(OH)/—(OH) 2 /—F/Ti Oxides) The aqueous dispersion of multi, oligo and mono‐layered Ti 3 C 2 T x MXene nanosheets, followed by steam treatment at 121 °C at a pressure of 134 kPa for 30 min led to formation of A‐MXene crystals. The transmission electron microscopy (TEM) and selected area electron diffraction (SAED) images of the A‐MXene revealed the formation of additional crystal lattice distortions in the microstructure of Ti 3 C 2 T x MXene nanosheets ( Figure 2 a, b ; Figure S2, Supporting Information). Our synthesis protocol resulted in the direct conversion of 2D Ti 3 C 2 T x MXene to a unique composite material with enhanced surface properties compared to its pristine nanosheets. Additionally, A‐MXene displayed the microstructural characteristics of Ti 3 C 2 T x nanosheets, quantum dots and surface titanium oxide nanoparticles in a single material. Our TEM/SAED data of Ti 3 C 2 T x MXene nanosheets displayed single lattice fringes with d‐spacing of ≈2. 70 Å assigned to the hexagonal crystalline planes (‐1100, 01–10, and 10–10) of MXene nanosheets (Figure 2a, inset). These findings are in line with the previous reports on Ti 3 C 2 T x MXene nanosheets. [ 13, 42 ] The SAED pattern of A‐MXene confirmed that the basal planes’ hexagonal structure of Ti 3 C 2 T x nanosheets remained unchanged after hydrothermal treatment (Figure 2b, inset). However, the TEM/SAED data of A‐MXene confirmed the emergence of Ti 3 C 2 T x quantum dots and surface titanium (II) and titanium (III) oxide nanoparticles into MXene nanosheets through secondary nucleation using hydrothermal process (Figure 2b, inset). Additionally, the crystalline pattern of A‐MXene displayed diffusion effect of titanium on the surface of MXene nanocrystals with polycrystalline diffraction rings corresponding to the (101), (103), (200), (211), (208), and (215) planes. The proposed chemical reactions for the formation of A‐MXene are presented in Equations S1–S4 in the Supporting Information. Taken together, the A‐MXene heterostructure that was obtained in the current study demonstrated a unique morphology of Ti 3 C 2 T x nanosheets which was anchored by different particles of 2D–1D–0D dimensionality. Figure 2 Synthesis and microstructural characterization of multidimensional GerMXene superlattice heterostructure. TEM images and corresponding SAED patterns (insets) of different composites. a) 2D Ti 3 C 2 T x MXene nanosheets. b) Crystalline A‐MXene complex. c) 2D‐0D hydrogenated GeH nanosheets and derived quantum dots. d) Multidimensional GerMXene superlattice. High‐resolution TEM images displayed the lattice of distinct crystals formed in the structure of GerMXene material. Our data confirmed the successful synthesis of GerMXene with unique morphology and microstructure of 2D MXene in a multidimensional material, including Ti 3 C 2 T x nanosheets and quantum dots, GeH nanosheets and quantum dots, and surface titanium oxide nanoparticles. e) These data clearly showed the lateral characteristic of different crystals in the structure of GerMXene with d‐spacing ranged from 1. 780 to 3. 590 Å. f) The XPS narrow scan spectra of Ti 2p and Ge 3d corresponding to Ti 3 C 2 T x MXene, GeH, and GerMXene. XPS data confirmed an exact reaction between titanium and germanium (Ge=Ti=Ge) to form GerMXene superlattice heterostructure. Further, our XPS analysis clearly showed the emergence of new Ti–Ge peaks in the spectra of Ti 2p and Ge 3d at the binding energy of 452–472 and 27–39 eV, respectively. These findings confirm that GeH quantum dots derived from hydrogenated GeH nanosheets are embedded in the structure of A‐MXene and form stable titanium germanide heterostructure. As described above, the A‐MXene material contains significant amounts of Ti 3 C 2 T x quantum dots and titanium (II) and titanium (III) oxide nanoparticles, formed and distributed on the MXene surface. The frequency histograms and interquartile range (IQR) analysis of A‐MXene showed that the medians of particles are 4. 88 nm in diameter (IQR: 3. 675) and 205 nm (IQR: 220), respectively (Figure S3, Supporting Information). Notably, our data showed that this specific morphology of A‐MXene was obtained after hydrothermal treatment and bath sonication of Ti 3 C 2 T x dispersions. We also characterized the aqueous colloids of exfoliated Ti 3 C 2 T x sheets before and after hydrothermal treatment and after sonication (Figure 2b ; Figures S4 and S5, Supporting Information). Interestingly, our scanning electron microscopy (SEM) and energy‐dispersive X‐ray spectroscopy (EDS) data showed a remarkable difference between microstructure of these samples. In particular, the SEM and EDS images of A‐MXene showed an expansion of Ti 3 C 2 T x layers with high secondary crystallization and formation of unique particle geometry into MXene nanosheets (Figure S5, Supporting Information). However, spontaneous formation of Ti 3 C 2 T x quantum dots and surface titanium oxide particles was relatively lower in the sample that was initially treated by sonication. This outcome might be due to the effect of hydrothermal treatment to further intercalate and oxidize MXene nanosheets before mechanical ultrasonic vibration. Further, we performed X‐ray diffraction (XRD) analysis of Ti 3 C 2 T x MXene powder and its aqueous colloids. The XRD patterns demonstrated that dispersion of Ti 3 C 2 T x MXene sheets in aqueous media resulted in relative oxidation and hydrolysis of nanosheets to form stable titanium oxides. We also characterized the crystalline pattern of A‐MXene to determine its phase behavior when dispersed in water. Furthermore, we assessed the phase comparison of A‐MXene with pure Ti 3 C 2 T x MXene quantum dots. Interestingly, the XRD spectra of A‐MXene confirmed the presence (002) peak at around 2‐theta = 7° as the main characteristic of Ti 3 C 2 T x. Furthermore, additional peaks were identified in these samples corresponding to the formation of stable phases of titanium oxides on Ti 3 C 2 T x MXene surface. Notably, the XRD patterns of A‐MXene depicted removal of aluminum from the structure of Ti 3 C 2 T x MXene nanosheets and quantum dots (Figure S6, Supporting Information). 2. 2 Development of GeH Quantum Dots Germanene nanosheets possess relatively weak π‐bonding between germanium atoms, therefore functionalization methods are used to improve its structural and electronic properties. These functionalizations include embedding hydrogen, fluorine and chlorine groups into the structure of germanene. These modifications such as hydrogenation of germanene to GeH increase buckling of the end product. However, dispersibility of GeH flakes in aqueous media to get a uniform and stable colloidal solution is still very challenging to achieve. Because 2D materials with bigger flake size are more prone to wrinkling and precipitation. On the other hand, quantum dots because of their sphericity and small size are easy to disperse and form a colloidal suspension. Therefore, in the current study, we converted 2D GeH sheets to quantum dots with lower dimensionality. The TEM analysis of hydrogenated GeH nanocrystals revealed a distribution of quantum dots with an average diameter of ≈4. 45 nm (SD: 0. 770) into its planar structure (Figure S7, Supporting Information). The SAED pattern of GeH was characterized using reciprocal lattice points grown on (1120) and (0002) planes with an angle of 90° (Figure 2c, inset). Besides, we performed the XRD characterization of GeH nanosheets and quantum dots. As shown in Figure S8 in the Supporting Information, the XRD pattern of hydrogenated GeH nanosheets is almost similar to the pattern recorded for aqueous GeH quantum dots. At ≈2‐theta = 16°, a high‐intensity peak of GeH that has been identified as the dominant peak in the XRD spectrum of nanosheets and quantum dots corresponds to (002) planes. These data are in line with the previous reports on GeH material. Additionally, the XRD pattern of GeH exhibited a relatively small peak at ≈2‐theta = 28° corresponding to (100) plane. However, few minor peaks at 2‐theta of ≈27°, 34°, 45°, and 54° of GeH nanosheets are absent in the XRD spectra of quantum dots. These peaks belong to germanium traces that were significantly eliminated from the XRD of GeH quantum dots due to purification during synthesis process. Further physicochemical characterization including SEM, TEM, STEM, EDS, and elemental mapping of 2D GeH material and its derived quantum dots are presented in Figure S9 in the Supporting Information. 2. 3 Development of Multidimensional GerMXene Superlattice (Ge=Ti=Ge) The A‐MXene and GeH complexes were incorporated to synthesize a novel multidimensional GerMXene structure by spontaneous reaction of titanium and germanium. This synthesis process efficiently produced a superlattice heterostructure through rapid and spontaneous secondary nucleation. As shown in our TEM and fast Fourier transform (FFT) images, several crystalline lattices were observed in the structure of GerMXene, which were attributed to Ti 3 C 2 T x MXene nanosheets, quantum dots as well as GeH nanosheets, quantum dots and surface titanium oxide nanoparticles (Figure 2d and inset). The multidimensional GerMXene possesses defined characteristics of its metal carbide, metal oxide and hydrogenated metalloid components with unique raspberry‐like nanoparticles anchored on its surface. Additionally, the FFT pattern of GerMXene material implied multiple circular cases, generally growing in a hexagonal (√3 × √3) superstructure. Furthermore, GerMXene heterostructure benefited from multiple Dirac points of embedded hydrogenated GeH quantum dots in its composition. However, this configuration appears to be less circular in shape and closer to trigonal positions, far away from Dirac points when interacting with the nearest conical bands. [ 43, 44, 45 ] Besides, our high‐resolution TEM images demonstrated distinct crystalline lattices in the range of 1. 78–3. 79 Å, formed during the synthesis of GerMXene (Figure 2e ). Further, TEM analyses of multidimensional GerMXene heterostructure are presented in Figure S10 in the Supporting Information. Next, the elemental chemical state of GerMXene heterostructure was characterized by X‐ray photoelectron spectroscopy (XPS). In particular, we intended to show the differences in the chemical composition of MXene and GerMXene materials. As described earlier, the Ti 3 C 2 T x nanosheets were significantly affected by the hydrothermal treatment and covalent bond formation between MXene quantum dots and surface titanium oxide particles. This finding is further confirmed by XPS narrow scan analysis (Figure 2f ). The Ti 2p spectrum of Ti 3 C 2 T x MXene identified the Ti–C, Ti 2p (1/2, 3/2), and Ti (II, III) peaks at binding energies of 457. 16–464. 61 eV, confirming the successful synthesis of MXene nanosheets. However, the XPS spectra of GerMXene exhibited a significant difference in its surface chemistry compared to MXene sample (Figure 2f ). In the Ge 3d narrow scan region of GeH crystals, dominant peaks of Ge–O, Ge–Ge, and sp 3 ‐like Ge–Ge were seen at the binding energies of 30. 33–36. 29 eV. Interestingly, the XPS Ti 2p and Ge 3d analysis of GerMXene showed new titanium germanide peaks at the binding energy of 455. 24, 461. 21, and 31. 80 eV, respectively. This finding was also confirmed by our XRD data of GerMXene. A new peak was detected at 2‐theta around 22° which corresponds to (111) plane of TiGe 2 bonds ( Figure 3 a, b ). Notably, the other dominant peaks of titanium germanide, including (311), (113), (002) (313), and (600) at 2‐theta of 35–50° were covered by high‐intensity Ti 3 C 2 T x peaks. These exciting data are in line with our proposed chemical reactions as the first experimental evidence for synthesis of quantum manipulated GerMXene heterostructure. Notably, the peak intensity in the XPS Ge 3d spectra of GerMXene is significantly higher than Ge 3d of GeH material. This phenomenon is a further confirmation of successful fabrication of GerMXene with a high concentration of GeH composition in its atomic structure. The details of XPS peak positions and quantifications are listed in Table S1 in the Supporting Information. Figure 3 Structural and morphological characterization of GerMXene. a, b) The XRD analysis of multidimensional GerMXene showing structural characteristics of Ti 3 C 2 T X MXene, hydrogenated GeH and stable titanium oxides. Furthermore, the XRD pattern confirmed the formation of (111) peak at 2‐theta of ≈23° corresponding to titanium germanide bond. The XRD data of GerMXene is in agreement with the XPS analysis of this sample. c–h) SEM images of GerMXene heterostructure at different magnifications. Morphology of the material revealed the conversion of exfoliated 2D Ti 3 C 2 T x MXene nanosheets to a new multidimensional heterostructure anchored by unique raspberry‐like nanoparticles. In particular, GerMXene superstructure includes significant amounts of MXene nanosheets, MXene quantum dots, germanane nanosheets, germanane quantum dots, and stable titanium oxide nanoparticles in a single material. i, j) Scanning TEM and EDS images of crystalline GerMXene superlattice showing its multidimensional architecture and chemical composition. Our EDS mapping showed the elemental distribution of GerMXene. This analysis confirmed that titanium, germanium, carbon, oxygen, and fluorine are the main components of GerMXene. We further characterized the obtained unique surface topography of GerMXene superlattice heterostructure by SEM (Figure 3c–g ). The surface functionalized nanosheets were anchored by the heterogeneous distribution of Ti 3 C 2 T x, GeH and titanium oxide particles. Our SEM data depicted that raspberry‐like nanoparticles grown on the surface of GerMXene have an average diameter of 211. 264 nm (SD: 53. 912). Additionally, it was observed that the median interlayer distance of nanosheets in the MXene was significantly decreased from 140 nm (IQR: 270) to 64. 34 nm (IQR: 325. 99) in GerMXene material (Figure S11, Supporting Information). The scanning TEM (STEM) image and corresponding EDS mappings of GerMXene displayed a bright crystalline surface of material and its compositional elements (Figure 3i, j ). The morphological details of GerMXene were further characterized by SEM images at different magnifications as presented in Figure S12 in the Supporting Information. 2. 4 Chemical Reactions and Phase Analysis of Model Next, we described the possible chemical reaction and phase analysis for the formation of GerMXene. The unique GerMXene heterostructure may have grown due to the spontaneous and intrinsic tendency of titanium and germanium to form titanium germanide (TiGe 2 and Ti 6 Ge 5 ) bonds. These covalent reactions take place at room temperature and the resultant buckled material possesses excellent dispersibility in aqueous solutions. Hence, we described a reaction model between titanium and metalloid germanium that can occur with a small free energy change, which is possible through two chemical procedures. First, typical diffusion‐controlled and nucleation reactions allow the crystallization of Ti 6 Ge 5 structure at low temperatures. After this reaction, the resultant titanium germanide bonds will interact with the remaining germanium atoms to construct TiGe 2 bond due to direct primary or secondary nucleation mechanism (Equations S5–S8, Supporting Information). Our proposed reactions for the formation of titanium germanide are effective at room temperature and stable at higher temperatures of up to 650 °C. This is further supported by recent computational simulations based theoretical investigations. [ 46, 47 ] Therefore, the experimental model in the current study and theoretical evidence in literature strongly suggest the occurrence of a possible reaction between group 4 or 5 transition metals and group IVA metalloids. Next, to enhance the scope of the current study, we performed literature search to explore the possibility of occurrence of similar reactions with other MXenes and metals of similar structures such as tantalum‐germanide (TaGe 2 ), niobium‐germanide (NbGe 2 ), titanium silicide (TiSi 2 ), vanadium silicide (V 6 Si 5 ), vanadium tantalum silicide (Ta 2 V 4 Si 5 ) and niobium chromium silicide (Nb 4 Cr 4 Si 5 ). We found that computational simulations based theoretical studies (Figures S13–S16, Supporting Information) have reported that the dissimilarity values for the formation of these structures are between 0. 18 and 0. 45 that highlights the potential of our experimental protocols in the current study to synthesize new materials with other MXenes. [ 48 ] In case of GerMXene, TiGe 2 crystal structure is in the orthorhombic Fddd space‐group and the spread Ti–Ge or Ge–Ti and Ge–Ge bond distances are in the range of 2. 65–2. 93 Å and 2. 66–2. 92 Å, respectively. Additionally, the formation of Ti 6 Ge 5 crystals appears in the orthorhombic Ibam space‐group with three titanium sites. First, due to the spreading of Ti–Ge or Ge–Ti bonds at 2. 63–2. 95 Å, each titanium atom binds to seven atoms of germanium to generate a TiGe 7 pentagonal pyramid structure. Similarly, at the other sites, titanium bonds with the other seven and six germanium atoms. Also, in the first inequivalent site of germanium, it readily bonds with eight titanium atoms. In the second and third sites, the germanium atoms with a length of 2. 62 Å bond with other nine and seven titanium atoms, respectively. [ 49, 50, 51 ] The calculations for crystal model, formation energy, band structure and density of states along with their phase diagram, average absorption and elasticity are provided in Figures S13–S16 in the Supporting Information. The orthorhombic TiGe 2 crystal represents Fddd space group (Hermann–Mauguin symbol ) with an approximate density of 6. 55 g·cm −3, formation energy (ΔH f ) of —0. 423 eV, decomposition energy of 0. 013 eV, energy‐above Hull of 0. 000 eV per atom and the defined lattice parameters of volume is 97. 874 Å 3 ( a = 6. 202, b = 5. 111, c = 5. 030). [ 47, 52, 53 ] For Ti 6 Ge 5 structure with a lattice of 354. 398 Å 3, the similar orthorhombic crystal with a space group Ibam (72) and a density of 6. 09 g·cm −3 could form at formation energy, decomposition energy of 0. 042 eV, energy above Hull of –0. 666 and 0. 011 eV per atom, respectively. Taken together, the data in the current study provide an experimental reaction model to synthesize new MXene‐based heterostructures. 2. 5 Bioactivity and Biocompatibility of GerMXene Next, we determined the biocompatibility and potential of GerMXene for tissue engineering and regenerative medicine applications. Recently, there has been growing interest in the application of carbon nanomaterials containing scaffolds to mimic the native extracellular matrix (ECM) to facilitate cellular attachment, proliferation and signaling. [ 54, 55 ] In this regard, we added colloidal suspensions of GerMXene (100 µg·mL −1 ) into clinically approved chitosan hydrogel to synthesize 3D injectable scaffolds. The detailed preparation and physicochemical characterization of chitosan and GerMXene–chitosan hydrogels is described in the method section. We characterized the microstructural properties of pristine and composite scaffolds by SEM, TEM, EDS, and Fourier‐transform infrared (FTIR) spectroscopy. As shown in Figure 4 a, the hydrogels are uniformly crosslinked with the median pore size of 51. 55µm (IQR: 85. 23) in chitosan samples. Notably, an increase in the pore size of composite hydrogels to 80. 89 µm was observed after incorporation of GerMXene crystals into the chitosan network. These findings were further confirmed by SEM/TEM images and EDS elemental mapping of GerMXene–chitosan samples (Figure 4a ; Figures S17 and S18, Supporting Information). Furthermore, our FTIR data showed the emergence of additional peaks in the structure of GerMXene–chitosan with high amount of surface functional groups compared to chitosan samples. The FT‐IR spectra of GerMXene–chitosan presented the surface functional groups including —OH, COOH, C=O, Ti—O, Ti—C, Ti—F, Ge—H, Ge—H 2, and —NH bonds at a wavelength range of 400–4000 cm −1 (Figure S19, Supporting Information). Figure 4 Characterization of GerMXene–chitosan and chitosan scaffolds and assessment of biocompatibility with H9C2 cells. a) The SEM/TEM/SAED images and EDS mapping of the crosslinked GerMXene–chitosan hydrogel was performed. The backscatter images clearly showed a distribution of GerMXene in the polymeric matrix of chitosan. The SAED pattern depicted crystalline rings of GerMXene–chitosan. Furthermore, the backscatter TEM images confirmed a successful synthesis of multidimensional GerMXene–chitosan heterostructure. Also, EDS mapping further displayed an elemental composition of GerMxene–chitosan. b–e) The biocompatibility and bioactivity of b, c) chitosan hydrogel and d, e) GerMXene–chitosan composite hydrogels was assessed after co‐culture with H9C2 cells for 24 and 120 h ( n = 5 per group). After the co‐culture live/dead assay was performed in H9C2 cells. The cells were stained with green Calcein AM for live cells and red EthD‐1 for detecting dead cells. There was a significant increase in the proliferation of H9C2 cells after co‐culture with GerMXene–chitosan on day 5. Imaging of the cells was carried out using Nikon Ti‐2 fluorescent microscope. The cytotoxicity evaluation of chitosan and GerMXene–chitosan hydrogels was performed after co‐culture with H9C2 cells for 24 h using LDH assay. Our data revealed no significant increase in LDH release after co‐culture with the materials for 24 h ( n = 5‐8 per group). (“ns” = statistically no significant difference, *** = p < 0. 001, and **** p < 0. 0001). To evaluate the biocompatibility and bioactive properties of GerMXene, the chitosan and GerMXene–chitosan hydrogels were co‐cultured with H9C2 cells (cardiomyocyte cell line) for 24 and 120 h. Prior to that, we assessed the biocompatibility of aqueous colloids of A‐MXene, GeH quantum dots and GerMXene with H9C2 cells at a concertation of 100 µg·mL −1. Our data showed no significant cytotoxic effects of the tested materials in H9C2 cells after 24 h of co‐culture (Figure S20, Supporting Information). Interestingly we found that after 24 h of co‐culture of A‐MXene, GeH and GerMXene dispersions with H9C2 cells, these materials were able to spontaneously enter into the cells without any uptake enhancing techniques (Figure S21, Supporting Information). These findings highlight the bioactivity and ability of GerMXene to transit from blood into specific tissues during future in vivo applications for nanomedicine‐based therapies. Furthermore, our data demonstrated that incorporation of GerMXene into the chitosan scaffolds significantly enhanced the cellular attachment and proliferation within the chitosan matrix. These properties are highly desirable for future applications of GerMXene for cell therapy mediated tissue repair. We assessed the survival of H9C2 cells after co‐culture with GerMXene for 120 h, interestingly our data demonstrated a significant increase in the population of H9C2 cells in GerMXene–chitosan hydrogels. There was a remarkable improvement in cell survival and proliferation after 120 h of culture (Figure 4b–g ; Figure S22, Supporting Information). Furthermore, the assessment of lactate dehydrogenase (LDH) release confirmed that the material was not cytotoxic to H9C2 cells after 24 h of co‐culture with GerMXene colloids at 100 µg·mL −1. Measurement of biodegradation properties of chitosan and GerMXene–chitosan hydrogels demonstrated that the incorporation of GerMXene into hydrogel networks did not affect the biodegradability of chitosan; rather it significantly improved the ability of chitosan to support cell survival and growth ( Figure 5 a ). This functionality is further supported by the pH analysis of freeze‐dried chitosan and GerMXene–chitosan solutions after hydrogels were soaked in culture media for 21 days at 37 °C (Figure 5b ). Furthermore, our data showed that the addition of GerMXene did not cause any significant changes in the acidity or basicity of MilliQ water, highlighting its suitability for use in the aqueous applications (Figure 5c ). Together, these findings suggest the potential of GerMXene as a bioactive material for tissue engineering and regenerative nanomedicine applications. Figure 5 Measurement of biodegradation, pH, surface charge, and electrical conductivity of the synthesized GerMXene hydrogel. a, b) The biodegradability of freeze‐dried GerMXene–chitosan and chitosan hydrogels was measured after 20 days. The results revealed an approximate degradation of 60% in chitosan hydrogels in the presence of lysozyme ( n = 3). Interestingly, the addition of GerMXene into the hydrogel networks did not affect the biodegradability of chitosan hydrogel. c) pH measurement of GerMXene at a concentration of 1000 µg mL −1 ( n = 3). Our dada displayed no significant changes in the acidity or basicity of MilliQ water in the presence of GerMXene nanocrystals. d) Zeta potential measurement of aqueous GerMXene suspensions at concentration of 100 µg mL −1, compared to A‐MXene and germanane samples ( n = 10). Our data confirmed a negative surface charge behavior of GerMXene at −10 to −25 mV. e, f) The electrical conductivity of aqueous suspension of GerMXene at concentration of 100 µg mL −1 compared to A‐MXene and GeH quantum dots ( n = 3). 2. 6 Zeta Potential, Electrical Conductivity, and Surface Area of GerMXene In the next experiment, we measured the surface charge of colloidal suspensions of GerMXene at a concertation of 100 µg·mL −1. Our zeta potential (ζ) measurements demonstrated that aqueous GerMXene dispersions have a negative surface charge in the range of −15 to −25 mV (Figure 5d ). These data coincide with our measurement on surface charge of synthesized A‐MXene (−10 to −25 mV) and is in agreement with the reports in literature on surface charge properties of 2D MXene nanosheets. [ 56, 57 ] However, our ζ analysis of aqueous GeH colloids showed a significantly lower negative surface charge compared to GerMXene at a similar concentration (+5 to −5 mV). The difference between surface charge of GerMXene and GeH could be due to differences in the surface functional groups. Furthermore, we evaluated the electrical conductivity of aqueous GerMXene (Figure 5e, f ). The data showed an average electrical conductivity of ≈125 µS·cm −1 of GerMXene samples, which is comparable with A‐MXene (≈175 µS·cm −1 ) and significantly higher than GeH quantum dots (≈25 µS·cm −1 ). Notably, the electrical conductivity of GerMXene remained mostly unchanged after two months. The surface properties of 2D MXene nanomaterials play an important role in their application in different fields. To characterize the surface properties of GerMXene, we measured specific surface area and porosity by Brunauer‐Emmett‐Teller (BET) nitrogen adsorption‐desorption isotherms and Barrett‐Joyner‐Halenda (BJH) method. [ 23, 58 ] Our BET data demonstrate that GerMXene has a significantly higher surface area compared Ti 3 C 2 T x MXene nanosheets (Figure S23, Supporting Information). Interestingly, the specific surface area of 2D MXene increased from 7. 11 to 91. 53 m 2 g −1 in the GerMXene sample. Therefore, conversion of Ti 3 C 2 T x MXene to GerMXene heterostructure contributed to an approximately 13‐fold increase in the surface area of this material. This structural change is mainly attributed to multidimensionality of GerMXene structure and formation of new particles on the surface of MXene sheets during synthesis. In particular, GerMXene exhibits type‐IV isotherms with H3‐type hysteresis loops characteristic of mesoporous materials in the relative pressure ( P /P 0 ) range of around 0. 5–1. 0 (Figure S23a, Supporting Information). Notably, our BET analysis is concisely in agreement with the morphological characterization of GerMXene heterostructure (Figure 2d, e, Figure 3c–h ; Figure S12, Supporting Information). Furthermore, the total pore volume of GerMXene is also increased from 0. 442 cc g −1 in MXene to 0. 616 cc g −1 for GerMXene (Figure S23b, Supporting Information). The employed method, therefore, has effectively produced a new multidimensional superlattice with excellent surface properties. 2. 7 Optical, Thermal, and Structural Stability of GerMXene We assessed the optical properties of GerMXene suspensions by ultraviolet‐visible (UV‐Vis) spectroscopy and found well defined lateral carbon structure of Ti 3 C 2 T x MXene at a wavelength range of 280–900 nm. Furthermore, the UV‐Vis analysis of GerMXene depicted no significant changes in the optical absorption of GerMXene after 60 days of synthesis (Figure S24a, Supporting Information). As shown in the optical micrographs, GerMXene suspensions remained stable without significant agglomeration or stacking at different concentrations (Figures S24b and S25, Supporting Information). The enhanced stability of GerMXene heterostructure can be attributed to the covalent and vdW bonds present in its structure. The thermal stability of GerMXene was evaluated using thermogravimetric analysis (TGA). After heating for up to 1000 °C under nitrogen atmosphere, GerMXene crystals showed excellent thermophysical stability (Figure S26, Supporting Information). The TGA analysis confirmed no significant changes in the weight percentage of GerMXene material at this temperature with a char residue higher than 97%. However, as expected, at a temperature over 600 °C, a slight decomposition of surface functional groups was observed in the TGA curve of GerMXene powder, resulted in a minor mass‐loss (≈2%). We further characterized the structural stability of GerMXene in aqueous colloidal suspensions at a concentration of 1000 µg·mL −1 by centrifugation at 1500 rpm for 15 min. Our data displayed no significant changes in the morphology and microstructure of GerMXene before and after centrifugation (Figure S27, Supporting Information). We further increased the spinning speed to 3000 rpm for 15 min, our SEM data confirmed that the multidimensional raspberry‐like structure of GerMXene was intact even after high‐speed centrifugation. Furthermore, the stability of GerMXene solids was tested at different temperatures in the range of 4–70 °C. As demonstrated in Figure S28 in the Supporting Information, no significant changes were observed in the structure of GerMXene at tested temperatures. Together, these data support the successful synthesis of a new multidimensional heterostructure nanomaterial with unique morphology and excellent stability for future applications in multiple fields. 3 Conclusion In summary, we have reported a versatile synthesis strategy for converting 2D transition‐metal carbide (M 3 X 2 T x ) MXene sheets to multidimensional GerMXene superlattice heterostructure. The newborn GerMXene possesses a unique microstructure with enhanced surface properties compared to its parent materials, MXene and Xene. Also, GerMXene offers excellent biocompatibility and bioactivity with improved electrical, structural, and optical properties for application in multiple fields. 4 Experimental Section Synthesis of A‐MXene Complex The Ti 3 C 2 T x nanosheets (Laizhou Kai Kai Ceramic Materials Co. Inc. ) were dispersed in Milli‐Q water and subjected to autoclave treatment at 121 °C for 30 min. The mixture was allowed to cool at room temperature. The surface modification of multidimensional MXene crystals was further achieved by treating it with ultrasonic treatment at 4 °C for 60–90 min. The obtained material was sterilized and stored at 4 °C for future experiments. Synthesis of 0D GeH Quantum Dots Quantum‐sized dots of GeH were prepared through sonication in an ice bath using 2D hydrogenated GeH (906026, Sigma Aldrich, Canada). Briefly, 0. 1 g of GeH flakes were dispersed in 100 mL of isopropyl alcohol (IPA) solution and stirred for one hour for further exfoliation. The mixture was sonicated at 4 °C for 24 h. The dispersion solution was then centrifuged at 3500 rpm for 30 min and thoroughly washed with ultrapure distilled water to obtain GeH crystals in the supernatant. The obtained suspension of GeH nanocrystals was further sonicated at 4 °C for 15–20 days until a desired size of GeH quantum dots was obtained. The colloidal solution was subsequently sterilized and stored at 4 °C for further use. Synthesis of Multidimensional GerMXene The formation of GerMXene heterostructure was achieved through a fast and spontaneous reaction of titanium and germanium at room temperature. Briefly, the aqueous colloidal suspensions of A‐MXene and GeH quantum dots at concentrations of 1000 µg·mL −1 (in the ratio of 1:1) were mixed gently. The obtained dispersion was sterilized and stored at room temperature for further characterization and future experiments. Synthesis of GerMXene–Chitosan Hydrogels Chitosan (CS) was used as a base material to prepare hydrogel scaffolds following the in‐house protocols 22. Briefly, 1 g of UV‐exposed low‐molecular‐weight CS powder (Sigma Aldrich, Canada) was dissolved in 40 mL of 0. 1M acetic acid and centrifuged for 10 min at 3000 rpm. Next, the solutions containing β‐glycerophosphate (Calbiochem, 1 g·mL −1 ) and hydroxyethyl cellulose (0. 025 g·mL −1 ) were slowly added to the hydrogel under constant stirring at 4 °C. The hydrogel composites were incubated at 37 °C for 30–60 min to obtain viscous polymeric solutions. Physicochemical Characterization The physicochemical properties of materials were characterized by scanning electron microscopy (FEI Nova NanoSEM 450, Thermo Fisher Scientific), transmission electron microscopy (FEI Talos F200X S/TEM, Thermo Fisher Scientific), X‐ray photoelectron spectroscopy (Kratos Axis Ultra XPS) and Fourier‐transform infrared spectroscopy (Thermo Nicolet Nexus 870) at the Manitoba Institute of Materials (MIM), the University of Manitoba. The ultraviolet‐visible spectroscopy and microscopic measurements were carried out by Cytation5 Cell Imaging Multi‐Mode‐Reader (BioTek Instruments) and fluorescence microscope (Nikon Eclipse Ti‐2). The electrical conductivity of material solutions was measured by a DuraProb 4‐Electrode. The pH of the material solution was measured by Thermo Scientific Portable Meter. The surface charge of aqueous GerMXene colloids at a concentration of 100 µg·mL −1 was assessed using Nanobrook ZetaPALS (Brookhaven Instruments). Surface Area, Zeta Potential, and Thermogravimetric Measurements The specific surface area of synthesized materials was assessed by Brunauer‐Emmett‐Teller (BET) nitrogen adsorption‐desorption isotherms and Barrett‐Joyner‐Halenda (BJH) method. The surface‐charge of aqueous Ti 3 C 2 T x MXene nanosheets, germanane quantum dots, and GerMXene colloidal suspensions at a concentration of 100 µg·mL −1 was measured using a Brookhaven Nanobrook ZetaPALS Instrument. For thermogravimetric analysis (TGA), a Q‐600 SDT TA‐Instrument was used at a heating rate of 10 °C·min −1 in nitrogen (100 mL min). The temperature settings for TGA measurements were as following; it was gradually raised to 100 °C (10 °C/minute), kept constant for 10 min, and then increased to 1000 °C. Biodegradability Measurement of Hydrogels To measure the biodegradability of prepared hydrogels, equal weights (Mi) of freeze‐dried GerMXene–chitosan and chitosan hydrogels were soaked in DMEM containing lysozyme (Sigma Aldrich, CA) at a concentration of 500 µg·mL −1 and incubated at 37 °C for 3 weeks. At the end of the incubation period, the CS and composite hydrogels were freeze‐dried to measure the final weight (Mf). The biodegradation rate of the hydrogels was calculated by following equation (1) Biodegradability % = M i − M f M i × 100 Biocompatibility Assessment To evaluate the biocompatibility of prepared materials live/dead assay was conducted in H9C2 cells. Briefly, the H9C2 cells were plated in 96 well plates, the synthesized materials were subsequently added to the cultured cells and incubated for 24 and 120 h. The assessment of cell viability was performed using a LIVE/DEAD Viability Kit (L3224, Thermo Fisher Scientific, USA). The quantification was performed using Cytation 5 Cell Imaging Multi‐Mode Reader (BioTek Instruments, USA). The microscopic images were captured using Nikon Eclipse Ti‐2 Fluorescence Microscope (Nikon Instruments Inc. , USA). To assess cytotoxicity of synthesized materials toward H9C2 cells, the cells were cultured with different forms of materials at a concentration of 100 µg·mL −1 for 24 h. The lactate dehydrogenase (LDH) release from the damaged cells was measured in supernatants using a commercial kit (MK401, Takara Bio). Cellular Uptake Assessment of GerMXene To understand the interaction of GerMXene with H9C2 cells, aqueous suspensions of materials at a concentration of 100 µg·mL −1 were co‐cultured with H9C2 cells for 24 h. After the co‐culture the cells were fixed with paraformaldehyde (4%) and mounted using Diamond Antifade reagent (Prolong) containing DAPI (Thermo Fisher Scientific, USA). The images were then captured by a fluorescence microscope (Nikon Eclipse Ti‐2). Statistical Analysis All data in the study were reported as mean ± standard deviation. The statistical comparison between multiple groups was performed using one‐way analysis of ANOVA, followed by Tukey's post‐hoc multiple comparison test and Student's t‐test. In all biological data statistical significance is determined as p < 0. 05. Conflict of Interest The authors declare no conflict of interest. Author Contributions The study was conceptualized and designed by A. R. , A. A. , and S. D. A. R. , A. A. , and W. Y. carried out experiments and acquired data. A. R. , A. A. , W. Y. , H. E. , and S. D. interpreted the data and performed statistical and formal analysis. A. R. , A. A. , W. Y. , and S. D. designed the figures. A. R. , A. A. , and S. D. drafted the manuscript. All authors have read and approved the final manuscript. Supporting information Supporting Information Click here for additional data file.
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10. 1002/adhm. 201200382
| 2,017
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Advanced healthcare materials
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Nanoscale Topography and Chemistry Affect Embryonic Stem Cell Self-Renewal and Early Differentiation
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Adherent cells respond to a wide range of substrate cues, including chemistry, topography, hydrophobicity, and surface energy. The cell-substrate interface is therefore an important design parameter in regenerative medicine and tissue engineering applications, where substrate cues are used to influence cell behavior. Thin films comprising 4. 5 nm (average diameter) gold nanoparticles coated with a mixture of two alkanethiols can confer hemispherical topography and specific chemistry to bulk substrates. The behavior of murine embryonic stem cells (ESCs) on the thin films can then be compared with their behavior on self-assembled monolayers of the same alkanethiols on vapor-deposited gold, which lack the topographical features. Cells cultured both with and without differentiation inhibitors are characterized by immunofluorescence for Oct4 and qPCR for Fgf5, Foxa2, Nanog, Pou5f1, and Sox2. Nanoscale chemistry and topography are found to influence stem cell differentiation, particularly the early differentiation markers, Fgf5 and Foxa2. Nanoscale topography also affects Oct4 localization, whereas the chemical composition of the substrate does not have an effect. It is demonstrated for the first time that ESCs can sense topographical features established by 4. 5 nm particles, and these findings suggest that nanoscale chemistry and topography can act synergistically to influence stem cell differentiation. This study furthers the understanding of the effects of these substrate properties, improving our ability to design materials to control stem cell fate.
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No full text available
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10. 1002/adhm. 201300063
| 2,014
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Advanced healthcare materials
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Combinatorial biomatrix/cell-based therapies for restoration of host tissue architecture and function
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This Progress Report reviews recent advances in the utility of extracellular matrix (ECM)-mimic biomaterials in presenting and delivering therapeutic cells to promote tissue healing. This overview gives a brief introduction of different cell types being used in regenerative medicine and tissue engineering while addressing critical issues that must be overcome before cell-based approaches can be routinely employed in the clinic. A selection of 5 commonly used cell-associated, biomaterial platforms (collagen, hyaluronic acid, fibrin, alginate, and poly(ethylene glycol)) are reviewed for treatment of a number of acute injury or diseases with emphasis on animal models and clinical trials. This article concludes with current challenges and future perspectives regarding foreign body host response to biomaterials and immunological reactions to allogeneic or xenogeneic cells, vascularization and angiogenesis, matching mechanical strength and anisotropy of native tissues, as well as other non-technical issues regarding the clinical translation of biomatrix/cell-based therapies.
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No full text available
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10. 1002/adhm. 201300505
| 2,014
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Advanced healthcare materials
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Electrospun PGS: PCL Microfibers Align Human Valvular Interstitial Cells and Provide Tunable Scaffold Anisotropy
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Tissue engineered heart valves (TEHV) could be useful in the repair of congenital or acquired valvular diseases due to their potential for growth and remodeling. The development of biomimetic scaffolds is a major challenge in heart valve tissue engineering. One of the most important structural characteristics of mature heart valve leaflets is their intrinsic anisotropy, which is derived from the microstructure of aligned collagen fibers in the extracellular matrix (ECM). In the present study, we used a directional electrospinning technique to fabricate fibrous poly-(glycerol sebacate):poly(caprolactone) (PGS:PCL) scaffolds containing aligned fibers, which resembled native heart valve leaflet ECM networks. In addition, the anisotropic mechanical characteristics of fabricated scaffolds were tuned by changing the ratio of PGS:PCL to mimic the native heart valve’s mechanical properties. Primary human valvular interstitial cells (VICs) attached and aligned along the anisotropic axes of all PGS:PCL scaffolds with various mechanical properties. The cells were also biochemically active in producing heart valve-associated collagen, vimentin, and smooth muscle actin as determined by gene expression. The fibrous PGS:PCL scaffolds seeded with human VICs mimicked the structure and mechanical properties of native valve leaflet tissues and would potentially be suitable for the replacement of heart valves in diverse patient populations.
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No full text available
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10. 1002/adhm. 201300620
| 2,014
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Advanced healthcare materials
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Role of Extracellular Matrix Signaling Cues in Modulating Cell Fate Commitment for Cardiovascular Tissue Engineering
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It is generally agreed that engineered cardiovascular tissues require cellular interactions with the local milieu. Within the microenvironment, the extracellular matrix (ECM) is an important support structure that provides dynamic signaling cues in part through its chemical, physical, and mechanical properties. In response to ECM factors, cells activate biochemical and mechanotransduction pathways that modulate their survival, growth, migration, differentiation, and function. This review describes the role of ECM chemical composition, spatial patterning, and mechanical stimulation in the specification of cardiovascular lineages, with a focus on stem cell differentiation, direct transdifferentiation, and endothelial-to-mesenchymal transition. The translational application of ECMs will be discussed in the context of cardiovascular tissue engineering and regenerative medicine.
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No full text available
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10. 1002/adhm. 201300678
| 2,014
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Advanced healthcare materials
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Confinement and Deformation of Single Cells and Their Nuclei Inside Size-Adapted Microtubes
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No abstract available
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The mechanical properties of the microenvironment of cells, like substrate rigidity and topography, have a considerable impact on various aspects of cell fate, such as proliferation, [ 1 – 3 ] apoptosis, [ 4 ] and differentiation. [ 5, 6 ] For example, on 2D predefined adhesion sites, the cells need a minimum area for spreading to survive [ 4 ] and asymmetrical patterns manipulate cell orientation during mitosis. [ 3 ] Human mesenchymal stem cells cultured on a nanograting respond to the nanotopographical cues through a significant extension of the cell nucleus along the axis of the grating. [ 7 ] Human embryonic and mesenchymal stem cells that orient along nanopatterned groves and ridges are promoted to differentiate preferentially along the neural lineage so that the topographic features of the substrate not only induce the differentiation process but also affect the specific functions of the resulting cells. [ 5, 6 ] The mechanisms by which physical cues impact cell functioning are just starting to be unraveled [ 8 ] and their investigation is complicated by the 3D character of physiological tissues, which is challenging to recapitulate in vitro. [ 9 ] 3D cell culturing scaffolds are therefore of particular interest for in vitro tissue engineering applications because they mimic the 3D aspects of the in vivo extracellular environment in contrast to the traditional and frequently used monolayer cell culture approach. Examples are fiber meshworks, sponges, and hydrogels of several biocompatible materials. [ 10, 11 ] The analysis of cells grown inside these structures shows that the physical properties of the 3D matrices evoke a cell response that differs from the one on 2D substrates, such as morphology changes [ 10, 12 ] and the formation of modified adhesion complexes. [ 13 ] It is the subject of current research how cells transform the physical properties of their microenvironments into a biochemical response, a process termed mechanotransduction. [ 14, 15 ] The investigation of the responses that the cells generate by a continuous probing and sensing of the environment has been facilitated by the emergence of microtechnologies. [ 16, 17 ] Advanced microfabrication methods have enabled the reproducible and parallel fabrication of patterned cell culture scaffolds by selectively controlling and manipulating the mechanical features of the substrate. [ 12 ] A non-invasive shape change of cells can, for example, be achieved by the topographical patterning of the cell culture substrates with micro-sized adhesion areas [ 8, 18, 19 ] or microgrooves. [ 15 ] These micropatterned substrates allow for the analysis of single cells by, for example, microscopic means. [ 15, 19, 20 ] The observed morphology change of the attached cells has been shown to influence the shape of the cell nucleus and to impact cell functioning. There is, for instance, accumulating evidence that changes in the morphology of the cell nucleus can influence its gene expression pattern by affecting the non-random positioning of chromosomes. [ 8, 15, 18, 21 – 23 ] However, the fabrication of topographically structured substrates that can confine cells in more than one dimensionality remains challenging. The usage of 3D biomaterials scaffolds suffers the drawback of decreased single-cell resolution due to the increased volume depth, while microstructured substrates usually provide an asymmetric polarization of matrix adhesions at the basal side of the cell. A suitable technique that circumvents these restrictions is rolled-up nanotechnology on polymers, [ 24 ] which can generate transparent and biocompatible silicon oxide/silicon dioxide (SiO/SiO 2 ) microtubes with different diameters that readily serve as cell culturing scaffolds. Inside the microtubes, cell growth is restricted in two dimensions (lateral and vertical) while the behavior of single cells can be easily observed. These microctubes have been shown to support the growth of various types of cells such as yeast [ 25 ] and HeLa cells, [ 26 ] as well as guide neuron extensions (e. g. axons), [ 27 ] and have been successfully employed to study mitotic processes in confined spaces. [ 28 ] Here, we report the confinement and microscopic analysis of single human osteosarcoma U2OS cells growing inside tailor-made microtubes. The aim is to study the effects of varying extents of spatial confinement, given by biofunctionalized silicon oxide (glass) microtubes, on different cell characteristics. Specifically, we employed the microtube system to investigate the effects of constricted cell growth on the morphology and integrity of the cell nucleus, the stiffest and largest organelle of the cell, and on two geometrically demanding processes—cell growth and cell division. We show that the U2OS cells were successfully confined inside the transparent structures. Depending on the tube diameter as a measure of the confinement level, we observed a distinct cell elongation that strongly affected the morphology of the cell nucleus. By quantifying the amount of DNA damage foci present in the cells, we found no correlation with the size of the confining structures indicating that the changes in nuclear morphology had no major effects on nuclear DNA integrity. In contrast, while the confined cells were able to divide inside microtubes of a wide diameter range, the majority of cells did not survive mitosis (cell division) inside the smallest diameter range of microtubes. Collectively, these findings demonstrate the applicability of the rolled-up microtubes as versatile, 3D biocompatible, and adjustable cell culturing scaffolds for various materials–cell investigations while mimicking the in vivo confinement of cells in their physiological 3D environment. To assess the effects of varying levels of spatial confinement on cellular shape and integrity, we tuned the diameters of the microtubes during the fabrication process [ 29 ] and adjusted them to fit the sizes of U2OS cells and their cell nuclei (see Figure 1a ). We determined the nuclear dimensions of freely growing U2OS cells to be 16 ± 2 μm in width, 23 ± 3 μm in length, and 4 ± 1 μm in height ( n = 60). Therefore, we fabricated different SiO/SiO 2 scaffolds with microtube diameters ranging from 4 μm (the average nucleus height of free growing cells) to 25 μm (a little larger than the average nucleus length of free growing cells). This range was chosen to ensure a significant confinement of cells that grow inside microtubes of smaller diameters as well as not significantly constricting the cells inside microtubes of larger diameters. A biofunctionalization step of the sample surface with fibronectin, a protein of the in vivo extracellular matrix, promoted cell growth on the silicon oxide substrate and inside the microtubes. U2OS cells were seeded onto each sample at high density to maximize the number of cells migrating into the microtubes within 2 d of incubation. The optical transparency of the silicon oxide material of the microtubes enabled the high-resolution microscopic observation of single cells, which entered the microtubes and grew inside the SiO/SiO 2 scaffolds. A representative movie that demonstrates how the U2OS cells first extended cell membrane outgrowths into the microtube, before translocating the main cell body into the structure, is provided in the Supporting Information (see Movie S1 ). The imaging of the samples demonstrated that the cells were able to grow and survive inside the confinement for a wide range of microtube sizes down to a minimum diameter of 5 μm (see Figure 1 ), below which the tubes did not sustain cell growth anymore. Fixation of the cells, staining of the DNA by DAPI and bright-field microscopy, as well as fluorescence imaging, revealed the shape change of the cell nucleus in dependence on the microtube diameter. In order to quantify this effect, we acquired z -stacks (image series of different focal planes) of confocal images at the positions of interest with a Zeiss LSM 700 inverted microscope. The widths and lengths of the cell nuclei were measured using the maximum intensity projections of each z -stack. The nucleus length was defined as the longest dimension of the DAPI-stained area, the nucleus width as the widest dilation orthogonally to it. The cut view of the z -stack projection then revealed the average nuclear heights (see Figure 1 b, c ). By forming the nuclear aspect ratios (nucleus length over width and nucleus length over height, see Figure 1d ), we accounted for any differences in nuclear volumes that arise from the progression through the cell cycle for instance due to increased protein and DNA synthesis in preparation for mitosis. [ 30 ] Strikingly, the nuclear aspect ratios revealed two microtube diameter ranges, which show different effects on the nuclear dimensions. For diameters ranging from 17 to 8 μm, the aspect ratios increased only slightly with decreasing microtube diameters, implying that the cells adapted to the reduction in available space without a profound effect on nuclear shape. The cell nucleus height and length remained fairly constant in this microtube diameter range, while the nucleus width slightly decreased with decreasing microtube diameter (data provided in Table S1 and depicted in Figure S1, Supporting Information ). At a tube diameter of 8 μm the aspect ratios of the nuclear width and height reached identical values reflecting an elongated, rod-like shaped nucleus, which is in sharp contrast to the rather flat and spread-out morphology of unconfined U2OS cell nuclei. For microtubes smaller than this 8 μm threshold, the dependence of the aspect ratios on the microtube diameter became prominent. Any reduction of the microtube diameter below this threshold resulted in a linear and equal decrease of the nucleus width and height from 6 ± 1 to 4 ± 1 μm (see Table S1, Supporting Information ) while the nucleus length increased to compensate for the restriction and to conserve the volume of the cell nucleus. When compared to free growing cells, it becomes evident that the U2OS cells inside the microtubes possessed nuclei that were generally more slender and not as flat as the cells on the planar substrate. The narrowed shape of the cell arises due to the lateral restriction of the cell dilation by the microtube wall. Surprisingly, the nucleus height was substantially increased inside the microtubes, which cannot only be accounted for by the 2D confinement, as the cell nuclei dimensions were hardly affected in the microtube diameter range from 8 to 17 μm. The functionalized microtubes offer a cell culture environment where the cells can form adhesions to the substrate all around the cell body so that a transition from a planar and spread to a more 3D morphology occurs. The dimensions of the cell nuclei were therefore directly influenced by the increased dimensionality of the cell culture scaffold. In microtubes with diameters smaller than 8 μm, the cells have to elongate profoundly to be able to squeeze into the microtube. The decrease of the nucleus width and height below a value of 6 ± 1 μm can only be compensated for by a considerable increase of the nucleus length indicating a significant remodeling of the cell nucleus content. These results are in line with the recent finding that the nucleus is a large cell organelle that can undergo remarkable deformation to migrate through small pores in 3D scaffolds. It is assumed that the changes in the nuclear shape are transient and reflect the interplay of forces imposed by the geometry of the cell environment and the intracellular counterforces. [ 31 ] Therefore, the microtube cell culture scaffold is suitable to determine the spatial limit for the “self-imposed” nuclear deformation of cells. This spatial limit for U2OS cells can be found at around 4 μm, as these cells do not readily grow into microtubes of this or smaller diameters. This corresponds to the fact that the confinement inside these small microtubes would force both the widths and the heights of the confined nuclei to acquire values lower than the average 4 μm height of nuclei in U2OS cells grown on planar substrates. Next, we evaluated if the observed changes in nuclear morphology, especially inside the narrowest microtubes below 8 μm, could influence the integrity of DNA, which is tightly packed and harbored inside the nucleus to protect and maintain genome stability. To do so, we used GFP-53BP1 U2OS cells [ 32 ] that were modi fied to stably express a fluorescently labeled marker of the DNA damage response pathway (p53-binding protein 1, 53BP1). [ 33 ] We quantified the numbers of GFP-53BP1 fluorescence foci (spots of higher light intensity) inside the U2OS nuclei as markers of DNA lesions to assess the amount of DNA damage depending on the diameter of the microtube (see Figure 2 ). We detected no correlation between the two parameters (compare Figure 2a with b–d and see Figure 2e ) demonstrating that the quantity of DNA lesions in asynchronous cell cultures was overall comparable between confined and freely growing cells. However, at this stage, we cannot exclude that certain changes in DNA integrity may still occur in selected cell cycle stages. This finding confirms the applicability of the microtube structures as cell culturing scaffolds that can directly manipulate the shape of whole cells and their nuclei without grossly affecting the integrity of DNA even when the cells are grown inside the narrowest microtubes of diameters between 5 and 8 μm. To assess whether the tubes impacted on the fate of U2OS cells, for instance, by affecting their growth and proliferation—two processes that depend on nuclear function—we studied the occurrence of mitosis and the survival of U2OS cells in microtubes of varying diameters (see Figure 3a, b ). We performed live-cell imaging experiments for at least 20 h on cells grown either on planar surfaces (reference cells) or inside differently sized microtubes to analyze the survival rates of confined and unconfined cells. Cells undergoing apoptosis were clearly discriminable due to extensive blebbing and dissolution of the GFP-BP1 fluorescence signal from the cell nucleus. The majority of cells (80%, Figure 3c ) survived throughout the observation period in microtubes of diameters larger than 8 μm, compared to almost 100% in unconfined conditions. This behavior changed considerably for cells that grew in microtubes with diameters smaller than 8 μm where half of the cells died. One process that could affect the survival of cells under confinement is mitosis. We therefore discriminated in our analysis between dividing and non-dividing cells and kept monitoring the cell fate of any arising daughter cells for an additional time period of at least 3. 5 h after mitosis ( Figure 3d ). The analysis revealed that around half of the cells underwent a first mitotic event within the observation period in both unconfined cells and cells grown inside microtubes >8 μm (total heights of left bars in Figure 3d ). However, an increased proportion of the cells died inside microtubes >8 μm (hatched and striped bar fractions in Figure 3d ) and more cells died after the occurrence of mitosis than without a cell division inside microtubes of this range. Still, one third of the confined cells were able to divide and survive inside confinements with a minimum diameter of 8 μm compared to 53% in free cells. In contrast, only 18% of the cells confined in tubes <8 μm divided and half of these died in the 3. 5 h after mitosis allowing only 9% of the daughter cells to survive mitosis inside these highly constricting cavities. In summary, although the cells and their nuclei elongate and adopt the shape of the topographic structures, the entry of cells into mitosis and the long-time survival of the confined cells are affected. Moreover, while cells are flexible and can adjust to a topographic confinement to some extent, [ 34 – 37 ] our findings demonstrate that they require a minimum space for survival and division. We conclude that extended confinement and substantial squeezing of the cell nucleus can impair the normal progression of the cell cycle. Hence, our results demonstrate that the microtube structures serve as versatile and biocompatible 3D cell culturing scaffolds. Their dimensions are designed during the fabrication process to confine and to noticeably change the cell shape and hence the morphology of the cell nucleus. The on-chip glass microtubes serve as an easily controllable tool to mechanically manipulate single cells and to test the maximum deformability of their nuclei without grossly perturbing the integrity of their DNA. The tunability of the microtube diameters allows for the adaption of the confinement to the specific cell type under investigation. The microtube cell culture system can be further employed to study the dynamics of cell nucleus deformation and remodeling, as well as the effect of a defined and long-term nucleus deformation on various cell responses. Examples are changes in chromosome positioning and protein expression levels. Another interesting follow-up study could investigate the influence of the biofunctionalization on the cell response to the increased dimensionality of the cell culture system by substituting fibronectin with other proteins of the native extracellular matrix environment. Taken together, these and further investigations using the microtube cell culturing system described above will help increase our understanding of the molecular processes involved in the mechanotransduction of extracellular signals within the 3D spatial and mechanical configuration of tissues. Experimental Section Microtube Sample Fabrication The fabrication of the transparent silicon monoxide/silicon dioxide (SiO/SiO 2 ) microtube samples is described in detail elsewhere. [ 24, 29 ] Briefly, a layer of ARP-3510 photoresist (Allresist GmbH) was spincoated at 3500 rpm on 18 mm × 18 mm cover glass substrates (high-performance, thickness no. 1½, Zeiss). The polymer film was patterned by conventional photolithography to produce squares of 100 μm in width and of a length varying between 100 and 500 μm. For the creation of the SiO/SiO 2 bilayer film 5 nm of SiO and between 20 and 100 nm of SiO 2 are then deposited in a 30° glancing angle electron beam deposition step at deposition rates of 5 or 0. 5 Å s −1, respectively. The glancing angle ensures the maintenance of an uncovered region behind the polymer squares (ballistic shadow effect). This window defines the starting point for the selective dissolving of the photoresist film in dimethyl sulfoxide (Sigma–Aldrich) so that the gradual release of the prestressed silicon oxide film leads to the self-assembly of the rolled-up microtube structures. The samples were subsequently dried by critical point drying (CPD 030 Critical Point Dryer, Bal-Tec AG) and coated with an 18-nm thick aluminum oxide (Al 2 O 3 ) film by atomic layer deposition (Savannah 100, Cambridge NanoTech Inc. ) to avoid the collapse of the thin structures. The cover slides were then functionalized with fibronectin to promote cell adhesion on the sample surface. Therefore, the microtube samples were immersed overnight in a solution of octadecanylphosphonic acid (50 μmol; Aldrich) in toluene (Sigma–Aldrich) and rinsed with toluene, acetone (Technic France), and deionized water. To enable the cell culturing on the microtube samples, the cover glasses were glued to 3. 5 cm plastic petri dishes with a 1. 4 cm hole in the bottom (MatTek Corporation) using a two-components glue (picodent twinsil) before incubating the microtube structures with a 1× DPBS solution (Gibco) containing N -(3-dimethylaminopropyl)- N ′-ethylcambodiimide hydrochloride (0. 1 m, Sigma–Aldrich), N -hydroxylsulfosuccinimide (0. 025 m ; Aldrich), and fibronectin (0. 02 mg mL −1 ; Sigma) for 4 h at 37 °C. The microtube samples were carefully rinsed with 1× DPBS solution and stored at 4 °C until used. Cell Culture Human osteosarcoma U2OS GFP-53BP1 cells have previously been described and were subcultured according to standard adherent mammalian tissue culture protocols. [ 31 ] They were grown in high-glucose DMEM (Sigma–Aldrich) supplemented with FBS (10%, Sigma–Aldrich), pyruvate (1 × 10 −3 m, Gibco), penicillin (100 U mL −1, Gibco), streptomycin (100 μg mL −1, Gibco), and geneticin (0. 5 mg mL −1, Gibco) and were maintained in a humidified incubator (37 °C, 5% CO 2 ). For the experiments, the cells were detached using trypsin–EDTA solution (0. 25%; Sigma) and resuspended in fresh medium for seeding on the microtube samples. To each 3. 5 cm Petridish, 10 6 cells in 3 mL medium were carefully added and allowed to settle before growing them for 2 more days in the humidified incubator to yield a nearly confluent monolayer. For the live-cell imaging experiments, the cell culture medium was exchanged to 3 mL of prewarmed phenolred free medium (constituents as listed above). Immunocytochemistry Cells were fixed with 2% paraformaldehyde (Sigma–Aldrich) in DPBS for 15 min, rinsed once with DPBS, and permeabilized with 0. 1% Triton X-100 in DPBS for 15 min. After washing three times with DPBS for 5 min each and removal of the liquid, a small drop of 4′, 6-Diamidin-2-phenylindol (DAPI, Invitrogen) containing mounting medium (Vector Laboratories) was applied and a 1. 3-mm round coverslip (VWR) carefully placed on top of the microtube sample. The edges were sealed with conventional nail polish and the sample was stored at 4 °C until being imaged. Imaging and Analysis Optical images of the fixed samples were taken with a Zeiss LSM 700 inverse confocal laser scanning microscope (40× objective, water immersion, NA = 1. 2) employing the software ZEN 2010. Live-cell imaging was performed using a Zeiss Axio Observer. Z1 inverse microscope equipped with a 37 °C heated stage and CO 2 chamber (40× objective, oil immersion, NA = 1. 1). The software Axio Vision Rel. 4. 8 (Carl Zeiss, Inc. ) was used for the image acquisition. The acquired images and image series were processed and analyzed with Fiji (image processing package, distribution of ImageJ). The particle analyzer function was used to quantify the number of DNA damage foci in each cell nucleus (minimal focus size: 0. 1 μm 2 ). Data are presented as mean ± standard deviation. Supplementary Material Supplementary Information Supplementary video
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10. 1002/adhm. 201400065
| 2,014
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Advanced Healthcare Materials
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Assessing Cellular Response to Functionalized α-Helical Peptide Hydrogels
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No abstract available
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For applications in 2D and 3D cell cultures and tissue engineering, there is a need to develop biocompatible scaffolds that support cell and tissue growth therefore mimicking the biochemical and morphological properties of the natural extracellular matrix (ECM). In order to support cellular growth, the scaffold must provide mechanical stability, promote cellular attachment, proliferation and differentiation, permit diffusion of gases, nutrients and waste and allow control of the degradation rate of the temporary support while minimizing cytotoxic side effects in vivo. 1 Hydrogels have been extensively investigated and used clinically for cell support in vitro and in vivo in regenerative medicine: their underlying structure mimics the interconnected fibrous network of the ECM; 2 the hydrated and porous nature of the gels allows diffusion of nutrients into the scaffold and waste to diffuse out 3, 4 ; and bioactive molecules can be incorporated into the fabric of the gels via passive uptake, direct incorporation during material synthesis, or conjugation after synthesis and/or assembly. 5 – 8 While natural and ex vivo materials such as agarose, 9 alginate, 10 carrageenan, 11 gelatin, 12 collagen, 13 and Matrigel 14 are common current choices for such scaffolds due to their availability and established cellular responses, there is often a lack of control over their formation, degradation, mechanical properties, and chemical modification. Furthermore, ex vivo scaffolds, such as collagens and Matrigel, 15 show batch-to-batch variation and can potentially introduce disease. Synthetic scaffolds, such as poly(hydroxyethylmethacrylate), 16 poly(vinyl alcohol), 17 and polypeptide-based protein anchors 18 address some of these issues, and provide partially favorable environments for 2D and 3D cell cultures. However, their reduced complexity often fails to mirror native tissue and some degradation by-products can cause unwanted cellular responses. 19 The advantages and limitations of various tissue engineering scaffolds are reviewed by Chan and Leong. 20 In principle, bottom-up scaffolds generated and engineered via biomolecular design allow the desired traits from the natural and synthetic scaffolds to be combined into one construct, and so create platforms for guiding cell growth and inducing specific biological responses. With this in mind, a number of peptide-based systems have been reported that utilize amyloid-like assemblies, 21 – 23 α-helical assemblies, 24 – 26 and peptide amphiphiles 27 – 30 as building blocks. A challenge in this area is to build complexity and control into these systems, ideally in a modular or pick-and-mix way; some of the systems reported to date lend themselves better to this ambition than others. 31 Using a bottom-up design approach, we have reported a two-component peptide system for making hydrogels, termed hSAFs (hydrogelating self-assembling fibers). 32 The peptides (hSAF-p1 and hSAF-p2) are designed de novo using principles for peptide self-assembly. When mixed the two peptides form coiled-coil α-helical fibrous structures, which subsequently interact to form percolated gels. These gels support 2D cell culture. Here, we show that the original two-peptide hSAF system can be supplemented with other components to bring cell-binding functions to the system, hence building up complexity and functionality. To achieve this functionalization, we developed a variant of hSAF-p1 harboring an azide moiety ( Figure 1 A, B, Figure 1, Supporting Information). This peptide, hSAF-p1(N 3 ), was mixed with hSAF-p2 and after overnight gelation an alkyne-bearing peptide containing the cell adhesion motif Arg-Gly-Asp-Ser (alk-RGDS) was added and appended to the hydrogel via copper-catalyzed azide-alkyne cycloaddition (CuAAC; hereafter referred to as the “click reaction”) by overnight reaction in the presence of Cu(I) ( Figure B ). 33 Alk-RGDS was used in this study as it promotes cellular attachment via integrin binding. 34 The use of RGD to promote cellular adhesion in other peptide-based fibrous and hydrogel systems has been reported. 25, 35, 36 We argue here that we gain added utility and control over assembly and functionalization using a modular, dual-peptide system, that is, the α-helical de novo-designed hSAFs. The RGDS-decorated hSAF assemblies were α-helical to an extent comparable to the parent system (see Figure 2, Supporting Information); and electron microscopy (EM) showed that the decoration and subsequent washing procedure did not perturb the gel structure ( Figure 2 A–D). For this work, we incorporated azidonorleucine at the N -terminus of hSAF-p1, although successful decoration was also achieved by substitution at the C -terminus (Figure 3, Supporting Information). Gel formation with predecorated p1(N 3 ) and p2 was not successful (Figure 4, Supporting Information). Analysis of the decorated gels by high-performance liquid chromatography (HPLC) showed that the RGDS functionality extended entirely through 2 mm thick gels (Figure 5, Supporting Information); and an absorbance-based copper assay showed that the copper used to drive the reaction was successfully removed by subsequent washing (Figure 6, Supporting Information). Figure 1 Peptide sequences, a schematic of the click reaction, and the half-moon model. A) Peptide sequences used for this study. Key: z, azido norleucine; Prop, propiolate. B) The gel was formed using an N -terminally azido-modified hSAF-p1. Decoration was achieved by performing a click reaction with alk-RGDS on azide-containing gels catalyzed by CuSO 4 with ascorbic acid (AA). C) Side-by-side gel formation in 24-well cell-culture plates allowed a direct comparison of cellular behavior on undecorated hSAF- and RGDS-decorated hSAF gels. Key: undecorated hSAF gel, gray; and RGDS-decorated hSAF gel, blue. Figure 2 Fiber morphology and gel structure. Transmission electron images for the A) undecorated hSAF and C) RGDS-decorated hSAF fibers. Average fiber diameters were 13 ± 5 nm for hSAF-undecorated fibers and 17 ± 4 nm for RGDS-decorated hSAF fibers. B, D) Scanning electron images showing interconnected fibers forming porous hydrogels of similar morphology D) with and B) without alk-RGDS. The gels are self-supporting (insets). Scale bars on (A, C) equal 200 nm while scale bars on (B, D) equal 1 μm. To assess and compare cellular responses of undecorated and RGDS-decorated hSAFs, we constructed a “half-moon model” ( Figure 1 C), in which the two gels were prepared side-by-side in the same tissue-culture well, a similar model to that recently presented by Chan et al. 37 As a model for neuronal differentiation, we seeded PC12 cells 38 on both sides of the half-moon hSAF gels. The experiments were followed by light and fluorescence microscopy ( Figure 3 A–F), and after 14 days cell morphology indicated that cells had attached to both the undecorated hSAF- and RGDS-decorated hSAF sides. However, the number of cells attached to the latter appeared considerably greater ( Figure 3 A, D). Figure 3 Response of PC12 cells to hydrogels. A, D) Light microscopy images showing PC12 attachment, and elongated cell morphology, to undecorated hSAF- and RGDS-decorated hSAF gels after 14 d. B, E) Representative fluorescent images for DAPI-stained cells on undecorated hSAF- and RGDS-decorated hSAF gels. C, F) Viable cells on undecorated hSAF- and RGDS-decorated hSAF gels indicated by calcein-AM staining. G) Proliferation of PC12 cells on gels and TCP over 14 d as judged by MTT assays. H) DNA quantification using Hoechst dye for PC12 cells on the gels and TCP over 14 d. I) PC12 differentiation, J) number of neurite-like processes, and K) lengths of processes as a function of time. Due to a high proliferation rate, individual cell processes were difficult to identify at day 14 on Matrigel. Dashed lines represent the projections for Matrigel assuming that the underlying trend from the early time points continues. Key: undecorated hSAF gel, gray; RGDS-decorated hSAF gel, blue; Matrigel, red; and TCP, green. The proliferative activity of the cells growing on the RGDS-decorated side of the gels was ≈50% greater than that of cells on the undecorated side. The higher rates of metabolic activity and proliferation 39 on the decorated hSAF side were similar to those observed for PC12 cells seeded on commercially available Matrigel (Figure 3G, H; Figures 7 and 8, Supporting Information). The above-mentioned experiments were conducted without neural growth factor (NGF), which terminates mitosis and induces primary neural outgrowth in PC12 cells. 40 In parallel experiments, the introduction of NGF promoted cell differentiation, as defined by the presence of neurite-like extensions, by day 3 on both sides of the gel ( Figure 3 I). Again the degree of differentiation was higher on the hSAF-decorated side: (11 ± 4. 6)% cells showed processes on this side, compared with (6 ± 2. 6)% on the undecorated side); and the mean number of neural projections extending from the cell body was more than twice as high on the decorated versus undecorated hSAF gel by day 14 ( Figure 3 J). However, the projection length did not vary significantly between the decorated hSAF and undecorated hSAF halves of the gel ( Figure 3 K). An assessment of PC12 cells on gels with alk-RGDS attached via the C -terminus of hSAF-p1 showed that the effects were similar to those observed with the ligand attached via the N -terminus (Figure 9, Supporting Information). To test the specificity of the peptide–cell interactions, we compared hSAF gels decorated with alk-RGDS and alk-RGES. The latter reduces the efficacy of cell attachment considerably compared with alk-RGDS-based sequences. 41 We found this to be the case in the hSAF system: changing the aspartic acid (D) to a glutamic acid (E) reduced cellular attachment by approximately 50% (Figure 10, Supporting Information). The above studies used hSAF gels with every hSAF-p1(N 3 ) decorated. It would be advantageous to reduce this percentage to reduce reagent costs and to allow combinations of functionalities to be added via addition of cocktails of modifiers. To begin testing this, we prepared gels with 1%, 10%, and 100% hSAF-p1(N 3 ) in hSAF-p1 and performed the click reaction with alk-RGDS. The cellular responses to undecorated hSAF and 1% incorporation were similar. However, the behavior of the 10% and 100% decoration was also similar, showing that considerably less reagent can be used (Figure 11, Supporting Information). Finally, a separate assessment of 3T3 fibroblast cells with RGDS-functionalized hSAFs showed that although the attachment of the cells appeared greater on RGDS-decorated gels than on the undecorated gels, the proliferative activity of the cells on RGDS-decorated hSAF gels was comparable to that on tissue-culture-treated poly(styrene) (TCP; Figures 12–14, Supporting Information). Thus, not all cells respond significantly to our hSAF gel system. In summary, we have conjugated a cell-adhesion motif to a rationally designed self-assembling peptide hydrogel system, resulting in stable functional scaffolds suitable for cell culture. Utilization of a “half-moon” protocol allows functionalized and non-functionalized gels to be compared directly in the same tissue-culture well. The morphology, viability, and proliferative activity of PC12 cells seeded on the scaffold surface were demonstrated over 14 days, showing enhanced cellular growth and differentiation on RGDS-modified hSAF gels, highlighting the potential for adding cell-specific motifs to more closely mimic ECM biochemistry. This novel functionalized system offers complex functional scaffolds with tight control over morphology and biochemistry, and with the potential to engineer cell cultures, cell therapy delivery systems, and tissue matrices that closely reflect the in vivo environment and thereby enhance cell performance. Experimental Section Scaffold Formation : Peptides were synthesized using standard solid-phase peptide synthesis protocols on a CEM “Liberty” microwave-assisted peptide synthesizer. Peptides were purified by reversed-phase HPLC and their masses confirmed by MALDI-TOF mass spectrometry. Typically, hSAF gels were prepared by mixing separate 1 × 10 −3 m stock solutions for each parent peptide (hSAF-p1 and hSAF-p2), which were made up in 20 × 10 −3 m MOPS (3-( N -morpholino)propanesulfonic acid) buffer at pH 7. 4. This gave final solutions of 0. 5 × 10 −3 m in each peptide. These were left on ice for 5 min followed by 30 min incubation at 20 °C, resulting in gels, which we refer to as 0. 5 × 10 −3 m gels. (n. b. , For the C -terminally modified peptide, the stock solutions were prepared at 2 × 10 −3 m, giving “1 × 10 −3 m gels”. ) For decoration experiments, hSAF-p1 was substituted for hSAF-p1(N 3 ). After, gel formation was performed by addition of 2 × 10 −3 m alk-RGDS and CuSO 4 and ascorbic acid each at 4 × 10 −3 m final concentration at 20 °C overnight. The gel was then washed with 10 × 10 −3 m ethylenediaminetetraacetic acid (EDTA) buffer, phosphate buffered saline (PBS), and supplemented-Dulbecco's Modified Eagle Medium (S-DMEM). The presence of remaining copper after decoration was assessed by bicinchoninic acid assay (see Figure 6, Supporting Information). The extent of clicked alk-RGDS was analyzed by analytical HPLC followed, with peak identity confirmed by mass spectrometry. Half-moon gels were formed in 24-well cell-culture plates using sterile glass coverslips as temporary separators for the undecorated hSAF- and RGDS-decorated hSAF gels. Biophysical Measurements : Peptide secondary structure was determined via circular dichroism spectroscopy using a Jasco J-810 CD spectrometer. Fiber morphology was visualized using a JEM 1200 EX MKI transmission electron microscope with a MegaViewII digital camera. Gel scaffold morphology was determined by fixing the sample with glutaraldehyde, removing the moisture via a critical point drying method and imaging using a Jeol JSM-633OF field-emission scanning electron microscope. Cell Studies : PC12 cells, kindly gifted by Prof. Jeremy Henley at the University of Bristol, were seeded onto gels. Cellular morphology was assessed using a light microscope. For live cell imaging, the cells were stained with calcein-AM, their nuclei highlighted with DAPI (4′, 6-diamidino-2-phenylindole) and imaged using a Leica DM IRBE inverted epifluorescence microscope. The metabolic activity, and therefore the proliferation rate, of the cells was evaluated by an MTT (3-(4, 5-dimethylthiazol-2-yl)-2, 5-diphenyltetrazolium bromide) absorbance assay (see §1. 14, Supporting Information). These data were supported by a DNA quantification assay (see §1. 15, Supporting Information). Differentiated PC12 cells were imaged using light microscopy, and ImageJ was used to count the number of differentiated cells (where differentiation is defined as one or more neural extension being longer than the major diameter of the cell body), the number of extensions per cell and the length of extensions. All quantitative data are presented in the format “mean ± standard error of the mean. ” Significant differences between comparable groups were determined by analysis of variance (ANOVA) with post hoc Tukey–Kramer honestly significant difference (HSD). The significance level was set at p < 0. 05. Author Contributions E. A. , M. A. B. , N. M. , and D. N. W. conceived the project. All authors designed the various experiments. E. A. , K. L. H. , N. M. , A. R. T. , and A. W. made the peptides and performed the biophysical work. N. M. and E. A. conducted the cell-culture experiments. M. A. B. and D. N. W. supervised the work. N. M. , A. W. , and D. N. W. wrote the paper.
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10. 1002/adhm. 201400095
| 2,014
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Advanced healthcare materials
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Biphasic Ferrogels for Triggered Drug and Cell Delivery
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Ferrogels are an attractive material for many biomedical applications due to their ability to deliver a wide variety of therapeutic drugs on-demand. However, typical ferrogels have yet to be optimized for use in cell-based therapies, as they possess limited ability to harbor and release viable cells. Previously, we have demonstrated an active porous scaffold that exhibits large deformations under moderate magnetic fields, resulting in enhanced biological agent release. However, at small device sizes optimal for implantation (e. g. , 2 mm thickness), these monophasic ferrogels no longer achieve significant deformation due to a reduced body force. In this study, we present a new biphasic ferrogel containing an iron oxide gradient capable of large deformations and triggered release even at small gel dimensions. Biphasic ferrogels demonstrate increased porosity, enhanced mechanical properties, and potentially increased biocompatibility due to their reduced iron oxide content. With their ability to deliver drugs and cells on-demand, it is expected that these ferrogels will have wide utility in the fields of tissue engineering and regenerative medicine.
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No full text available
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10. 1002/adhm. 201400277
| 2,015
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Advanced healthcare materials
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An Enzyme-sensitive PEG Hydrogel Based on Aggrecan Catabolism for Cartilage Tissue Engineering
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A new cartilage-specific degradable hydrogel based on photoclickable thiol-ene PEG hydrogels is presented. The hydrogel crosslinks are composed of the peptide, CRDTEGE-ARGSVIDRC, derived from the aggrecanase-cleavable site in aggrecan. This new hydrogel is evaluated for use in cartilage tissue engineering by encapsulating bovine chondrocytes from different cell sources (skeletally immature (juvenile) and mature (adult) donors and adult cells stimulated with pro-inflammatory lipopolysaccharide (LPS)) and culturing for 12 weeks. Regardless of cell source, a two-fold decrease in compressive modulus is observed by 12 weeks, but without significant hydrogel swelling indicating limited bulk degradation. For juvenile cells, a connected matrix rich in aggrecan and collagen II, but minimal collagens I and X is observed. For adult cells, less matrix, but similar quality, is deposited. Aggrecanase activity is elevated, although without accelerating bulk hydrogel degradation. LPS further decreased matrix production, but did not affect aggrecanase activity. In contrast, matrix deposition in the non-degradable hydrogels consisted of aggrecan and collagens I, II and X, indicative of hypertrophic cartilage. Lastly, no inflammatory response in chondrocytes is observed by the aggrecanase-sensitive hydrogels. Overall, we demonstrate that this new aggrecanase-sensitive hydrogel, which is degradable by chondrocytes and promotes a hyaline-like engineered cartilage, is promising for cartilage regeneration.
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No full text available
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10. 1002/adhm. 201400458
| 2,015
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Advanced healthcare materials
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Spatial Control of Cell Gene Expression by siRNA Gradients in Biodegradable Hydrogels
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The extracellular environment exposes cells to numerous biochemical and physical signals that regulate their behavior. Strategies for generating continuous gradients of signals in biomaterials may allow for spatial control and patterning of cell behavior, and ultimately aid in the engineering of complex tissues. Short interfering RNA (siRNA) can regulate gene expression by silencing specific mRNA molecules post-transcriptionally, which may be valuable when presented in a continuous gradient for regenerative or therapeutic applications. Here, a biodegradable hydrogel system containing a gradient of siRNA is presented, and its capacity to regulate protein expression of encapsulated cells in a spatially continuous manner is demonstrated. Photocrosslinkable dextran hydrogels containing a gradient of siRNA have been successfully fabricated using a dual programmable syringe pump system, and differential gene silencing in incorporated cells that is sustained over time has been shown using green fluorescent protein as a reporter. This platform technology may be applied in tissue engineering to spatially control biologically relevant cellular processes.
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No full text available
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10. 1002/adhm. 201400695
| 2,015
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Advanced healthcare materials
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Development of a cellularly degradable PEG hydrogel to promote articular cartilage extracellular matrix deposition
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Healing articular cartilage remains a significant clinical challenge because of its limited self-healing capacity. While delivery of autologous chondrocytes to cartilage defects has received growing interest, combining cell-based therapies with scaffolds that capture aspects of native tissue and promote cell-mediated remodeling could improve outcomes. Currently, scaffold-based therapies with encapsulated chondrocytes permit matrix production; however, resorption of the scaffold does not match the rate of production by cells leading to generally low ECM outputs. Here, a PEG norbornene hydrogel was functionalized with thiolated TGF-β1 and crosslinked by an MMP-degradable peptide. Chondrocytes were co-encapsulated with a smaller population of MSCs, with the goal of stimulating matrix production and increasing bulk mechanical properties of the scaffold. Interestingly, the co-encapsulated cells cleaved the MMP-degradable target sequence more readily than either cell population alone. Relative to non-degradable gels, cellularly-degraded materials showed significantly increased GAG and collagen deposition over just 14 days of culture, while maintaining high levels of viability and producing a more diffuse matrix. These results indicate the potential of an enzymatically-degradable, peptide-functionalized PEG hydrogel to locally influence and promote cartilage matrix production over a short period. Scaffolds that permit cell-mediated remodeling may be useful in designing treatment options for cartilage tissue engineering applications.
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No full text available
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10. 1002/adhm. 201400724
| 2,015
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Advanced Healthcare Materials
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Going Beyond Compromises in Multifunctionality of Biomaterials
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No abstract available
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Design, synthesis, processing, and testing, or perhaps better altogether exploration, of biomaterials, which are intended to substitute native tissue, have generally been approached from one of two distinct perspectives ( Figure 1 A). Researchers with a background in chemistry or biology have been inspired by the natural surroundings of cells, the extracellular matrix (ECM). The ECM is a complex nanostructured system of fiber and network forming macromolecules as well as soluble compounds embedded in a hydrogel. The fiber and network forming macromolecules such as collagen and elastin enable the elastic deformability and recoverability of tissues. Water storage in the hydrogel is ruled by polysaccharide and proteoglycan components, such as glycosaminoglycans, for example, hyaluronic acid. The hydrogel furthermore counteracts the contraction by the fibers and the elastic network. At the same time, it allows the diffusion of gasses, ions, nutrients, and metabolites necessary for the supply of and communication between the cells. Anchoring of cells to the matrix as well as of the different macromolecular components is generally ruled by non-covalent, specific adhesion such as the interaction of the RGD sequence in, for example, fibronectin and integrins in cell membranes. The matrix is built up and degraded through hydrolytic as well as enzyme- and cell-mediated events, which in vivo leads to a continuous remodeling and renewal. In an attempt to learn from nature the macromolecular components of the ECM can be used or emulated, and selected functionalities of the ECM can be mimicked. Taking the ECM structure and functions as blueprint led to a much improved understanding of the interplay between cells and materials. However, materials designed in this way have rarely been advanced to technical or clinical applications. Examples for approaches in this field are the coating of polymers or metals with extracellular matrix extracts produced from sarcoma cells which are harvested and decellularized, [ 1 ] or the production of ECM on an artificial surface by cells which are removed prior to application. Such approaches suffer from an incomplete knowledge and limited control of the actual composition and chemical structure, as well as batch-to-batch variability of the biotechnologically produced ECM. Alternatively, full tissues are decellularized and used as guiding structure in tissue engineering. [ 2 ] From the perspective of the medical device design aiming at a specific clinical need, especially biomedical engineers have concentrated on formulating requirements, and used “from the shelf” materials to reach their goals. Figure 1 A) The extracellular matrix (ECM) is the natural and self-produced environment of cells. Its structure and functions are explored to gain a fundamental understanding. But the overall complexity of the ECM cannot (yet) be mimicked to enable multifunctional devices. On the other hand, a specific application can give rise to formulate and prioritize functions. However, addressing the prioritized functions with readily available materials often goes hand in hand with compromises for properties and functions of lower importance. Bridging of the two approaches demands novel strategies. Figures reproduced with permission: left, [ 3 ] Copyright 2011, IOS Press; right, [ 4 ] Copyright 2010, WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim. B–D) Functions of biomaterials. B) Degradability of materials is complex in vivo, as hydrolysis, enzymatic degradation, mechanical load, and cell-mediated processes contribute to degradation also of materials intended for long-term application. The rate of degradation will be influenced by individual preconditions. C) Control of release can be realized through diffusion or degradation control. D) Biomaterials mimicking different aspects of the ECM structure. Open porous and interconnective 3D structures (top) allow migration of cells through pores of sufficient size. Nanofiber meshes (middle) resembling the collagen fiber network of the ECM, with fiber diameters typically being in the range of 500 nm–2 μm (photo reproduced with permission;[ 4 ] Copyright 2010, WILEY-VCH Verlag GmbH & Co. KGaA). Hydrogels (bottom) can only be infiltrated by cells if cell-mediated degradation can take place (photo reproduced with permission;[ 5 ] Copyright 2010 of The Royal Society of Chemistry ( http://pubs. rsc. org/en/content/articlelanding/2010/jm/c0jm00883d ). Though, with the latter strategy, devices which are nowadays firmly established in the clinic were developed, the application of engineering plastics originally not intended for clinical use is often connected with compromises regarding the matching of properties and functions of the materials with the requirements of the application. For device design, one function of the material is frequently prioritized to understand, employ, expand, or tailor, while other functions, judged as of lower importance, are accepted as is. An example for a prioritized function is the structural function, which could be realized by existing engineering plastics, for example, for hip implants. An early example for the design of a material with one function (see Figure 1 B–D) is the tailoring of degradation rate of synthetic polymers, which was approached through changing comonomer types and ratio as well as molecular weight distribution of, for example, copolyesters. [ 6 ] Copolyesters such as poly(lactide- co -glycolide) (PLGA) or ε -caprolactone-based copolymers could be adjusted in their degradation rate in a time frame of weeks to years. The desire to change the release profile of bioactives from polymer matrices actually triggered the investigation of polymers with different degradation behavior: on the one hand, bulk degrading materials such as (co)polyesters and, on the other hand, surface degrading polymers such as polyanhydrides or poly(ortho esters). [ 7 ] However, the elastic properties of these classical degradable polymers are not comparable to the elastic properties of soft tissues. In some cases, such as in the use of PLGA as matrix of drug delivery systems or surgical sutures, such compromise might be acceptable, while the elastic properties of (co)polyesters are unsuitable for their application as soft tissue implant, for example, for augmentation. The successful application of polymer-based implants in the clinic rapidly stimulated the generation of ideas for novel applications. For each application, a characteristic combination of properties and functions is required. If more of these novel applications are to be realized, this also means that a larger variety of polymers will have to be available fulfilling these diverse requirement profiles. An essential function of biomaterials is their biocompatibility. [ 8 ] The concept of biocompatibility first of all implies the non-toxicity of a material (though, in fact some materials should be cytotoxic under certain boundary conditions, such as for cancer cells after specific cellular uptake). This non-toxicity of eluates of the material as well as measured in direct cell-material interactions is based on the fact that components, starting materials, catalysts, or degradation products are not released in toxic amounts and that the direct contact does not impair cellular function. Toxicity tests are typically performed with cells from cell lines, for instance, L929 mouse fibroblasts. A more specific evaluation covering toxicity as well as cell compatibility comprises the investigation with one or few specific cell types associated with a potential application of the materials. The histocompatibility of the materials with, for example, specific soft or hard tissues can be evaluated in vivo. Furthermore, the evaluation of hemocompatibility of materials[ 9 ] is of relevance for all implants in contact with blood as well as for extracorporeal devices such as heart-lung-machines or apheresis devices. Finally, the biofunctionality of a device is evaluated in vivo. The above discussed examples for aspects of biocompatibility and the corresponding plethora of toxicity and biofunctionality tests shows that the material function biocompatibility has to be considered with differentiated views. A material's performance has to correspond to each facet of biofunctionality, which increases the complexity of multifunctionality considerably. So, how can the high expectations towards multifunctionality be fulfilled? Approaches to biofunctional materials have on the one hand to be concentrated on realizing biochemical cues found in nature to be provided by a material. In addition to incorporation of RGD-based cell adhesion sites, options include provision of enzyme-sensitive moieties allowing cell-mediated degradation, or loading with bioactive macromolecules such as growth factors or cytokines. By increasing the complexity of the system, it was hoped that the biological performance of the biomaterial is improved. [ 10 ] However, contrasting the success in interesting pilot studies in vitro, the translation into, for example, clinical applications is seriously hampered by the accompanying increase of complexity of synthesis, shown by the increase of synthetic steps necessary to create such systems. The provision of protein cues released from the matrix has been shown to be risk-associated in vivo because of potential overshooting reactions and/or ectopic biological effects. [ 11 ] For example, application of bone morphogenic proteins has in some cases resulted in ectopic or excess bone formation as well as potential increase in cancer risk. Potential reasons might be unsuitable levels of the protein released, and/or insufficient feedback and deactivation of the active compounds as natural inhibiting and control mechanisms are not released at the same time. This might be addressed by applying cells as local factor release systems, for example, as suggested for muscle stem cell implantation for VEGF release in ischemic hearts, [ 12 ] or by cell-material constructs, in which the living cells communicate with their biological environment and release only the appropriate biological signals in the needed amount. A fundamentally different approach is the use of physical cues such as substrate elasticity[ 13 ] and geometry, for example, of 2D surface patterns to direct stem cell fate. [ 14 ] Early examples included the matrix-elasticity directing stem cell differentiation, [ 15 ] as well as the importance of pore size in scaffolds for good tissue integration. [ 16 ] Recently, the influence of the geometry of 3D microstructured microwells on human adipose-derived mesenchymal stem cells morphology, migration, and proliferation, [ 17 ] but also on differentiation and gene expression exemplarily highlights the interdependence of biological biomaterial studies and progress in biomaterial design and processing. The realization of multifunctional biomaterials as enabling technology for novel biomedical applications requires strategies, which go beyond compromises. Integrative approaches might be a way to achieve this goal. An interdisciplinary research team of chemists, material scientists, biologists, engineers, and physicians has to address one specific application and develop the right material for this purpose. This requires patience from the people nearer to the application as such material invention and exploration might take some time as well as more application motivated thinking in fundamental research. Different concepts for efficiently integrating multiple functions, which are independently from each other, need to be explored. This includes a separation of functions by being recalled not simultaneously but successively, or realizing different functions on different length scales or in different material phases. Finally, integrative processes might play an important role in the future by integrating chemical synthesis and processing in one step procedures. This concept deserves more attention to bridge the gap between the “learning from nature” and “device designer” communities, as in this way the potential drawbacks of multi-step syntheses so far devised for increasing the complexity of a system might be circumvented and translation will be facilitated.
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10. 1002/adhm. 201400760
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Advanced healthcare materials
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Biomaterials for Bone Regenerative Engineering
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Strategies for bone tissue regeneration have been continuously evolving for the last 25 years since the introduction of the “tissue engineering” concept. The convergence of the life, physical, and engineering sciences has brought in several advanced technologies available to tissue engineers and scientists. This resulted in the creation of a new multidisciplinary field termed as “regenerative engineering”. In this article, the role of biomaterials in bone regenerative engineering is systematically reviewed to elucidate the new design criteria for the next generation of biomaterials for bone regenerative engineering. We highlight the exemplary design of biomaterials harnessing various materials characteristics towards successful bone defect repair and regeneration. In particular, we concentrate our attention on the attempts of incorporating advanced materials science, stem cell technologies, and developmental biology into biomaterials design to engineer and develop the next generation bone grafts.
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No full text available
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10. 1002/adhm. 201400762
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Advanced Healthcare Materials
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Generation and Assessment of Functional Biomaterial Scaffolds for Applications in Cardiovascular Tissue Engineering and Regenerative Medicine
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Current clinically applicable tissue and organ replacement therapies are limited in the field of cardiovascular regenerative medicine. The available options do not regenerate damaged tissues and organs, and, in the majority of the cases, show insufficient restoration of tissue function. To date, anticoagulant drug‐free heart valve replacements or growing valves for pediatric patients, hemocompatible and thrombus‐free vascular substitutes that are smaller than 6 mm, and stem cell‐recruiting delivery systems that induce myocardial regeneration are still only visions of researchers and medical professionals worldwide and far from being the standard of clinical treatment. The design of functional off‐the‐shelf biomaterials as well as automatable and up‐scalable biomaterial processing methods are the focus of current research endeavors and of great interest for fields of tissue engineering and regenerative medicine. Here, various approaches that aim to overcome the current limitations are reviewed, focusing on biomaterials design and generation methods for myocardium, heart valves, and blood vessels. Furthermore, novel contact‐ and marker‐free biomaterial and extracellular matrix assessment methods are highlighted.
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1 Introduction Cardiovascular diseases (CVDs) such as coronary heart disease, rheumatic heart disease, congenital heart disease, myocardial infarction (MI) and strokes remain the number one cause of death in western countries. Despite the development of new therapies and the approach of risk‐reducing strategies to prevent CVDs, the World Health Organization estimates an increase from 17. 3 million deaths in 2008 to 23. 3 million by 2030. 1 To date, there are no effective therapies to fully restore cardiovascular organ function after damage, and due to an insufficient number of organ donors, 2, 3 the design of regenerative strategies such as stem‐cell‐based therapies and in vitro‐engineered tissues are in great demand. One focus in the field of regenerative medicine is the generation and improvement of functional biomaterials. These materials can be bioactive; releasing drugs, proteins, growth factors and extracellular matrix (ECM) components, or they can have an improved mechanical functionality. Today, there are several different types of biomaterials used to induce tissue regeneration or remodeling, and to support tissue function or even replace damaged structures. These biomaterials include metals, metal alloys and also synthetic and natural occurring polymers. 4 Enormous progress has been made in generating and designing biomaterials for tissue repair. Highly promising are the so‐called biomimetic materials, which mimic ECM architecture and provide potentially controllable in vivo‐like microenvironments for cells. 5 In vivo, cell behavior includes survival or controlled‐death, migration, proliferation and differentiation, which can be all controlled or impacted by the three‐dimensional (3D) ECM. 5, 6, 7 Accordingly, physical, chemical and biological cues from the ECM must be investigated and subsequently recapitulated. In this article, we provide an overview of functional biomaterials that are currently investigated for the use in the fields of cardiovascular tissue engineering and regenerative medicine, mainly focusing on myocardial repair, heart valves and blood vessels. Furthermore, promising contact‐ and marker‐free assessment methods are discussed for quality and pre‐implantation control or biomaterial assessment. 2 Biomaterials Used for Myocardial Repair Multiple clinical and preclinical trials have investigated the effect of stem cell injections to induce cardiac regeneration after MI. 8, 9, 10, 11, 12 Although the injection of a cell suspension showed promising results regarding improvement of cardiac function, there exist major obstacles such as poor cell retention and hence consistent efficacy cannot be achieved. 13 To increase cell survival and to enable a localized delivery of cells without mechanical washout, biomaterials are used ( Table 1 ). Table 1 Overview of biomaterials and fabrication methods used for myocardial repair Biomaterial Fabrication Method Cells/Molecules included Application and Results Ref. PEGylated fibrinogen Hydrogel formation Cardiomyocytes, Embryonic stem cells in vivo: improved fractional shortening 15 Chitosan Hydrogel formation Brown adipose derived stem cells in vivo: enhanced cardiomyocyte differentiation and increased angiogenesis 16 Chitosan Hydrogel formation Adipose‐derived mesenchymal stem cells in vivo: increased cell survival 17 Fibrin glue Hydrogel formation Adipose‐derived stem cells in vivo: improved heart function 18 Alginate‐RGD Hydrogel formation Human mesenchymal stem cells in vivo: reduced infarct size and improved cell survival 20 PEG‐vinylsulfone Hydrogel formation ‐ in vitro: directs differentiation of pluripotent cardioprogenitors 23 Alginate Hydrogel formation B16‐F10 cells shape memory gels in vivo: enhanced cell survival 24 Collagen type I Hydrogel formation miR‐29B in vivo and in vitro: reduced wound contraction, improved ECM remodeling 27 HEMA – hyaluronic acid Hydrogel formation Neuregulin‐1β in vivo: enhanced left ventricular ejection fraction 28 RADA16 Self‐assembling hydrogel VEGF and heparin in vivo: enhanced cardiac function 29 Poly(ethylene argininylaspartate diglyceride) Self‐assembling hydrogel sonic hedgehoc in vitro: upregulation of growth factors in a cardiac fibroblast cell culture 30 PEG Hydrogel formation SDF1‐GPVI in vitro: release studies in vivo: cell immobilization via SDF‐1‐GPVI 3, 32 ECM Cell sheets ‐ in vivo: improved cardiac function and neovascularization 35 Myocardial ECM Hydrogel formation ‐ in vivo (large animal): increased cardiac function and reduced infarct fibrosis 36 Xylan/PVA Electrospinning ‐ in vivo: increased cardiac cell proliferation 38 Pullan‐Dextran Salt‐leaching ‐ in vitro: hECFC culture 39 Hyaluronic acid Hydrogel formation ‐ in vivo: increased wall thickness 40 Alginate Hydrogel formation ‐ Clinical trial: preserved LV function 45 PU‐aniline pentamer with PCL Molding oligoanilines in vitro: cytocompatibility and conductivity tests 49 PLGA‐Gelatin Electrospinning ‐ in vitro: integration of cardiomyocytes 50 PGS Molding ‐ matches physical (mechanical) properties of the heart 51 PCL Electrospinning ‐ in vitro: cardiomyocyte attachment and beating on day 3 52 PCU or PLGA Electrospinning on a textile‐template ‐ in vitro: adhesion and proliferation of H9C2 cardiacmyoblasts cell line and beating on day 10 on PCU 58 PCL‐Gelatin Electrospinning Mesenchymal stem cells in vivo: reduced scar size and microvessel formation 59 PGS‐Fibrinogen Electrospinning VEGF and mesenchymal stem cells in vivo (large animal): improved ejection fraction 60 PEUU Phase separation ‐ in vivo: improved cardiac remodeling and contractile function 63 PLLA (on a PCLA sponge) Knitting Vascular smooth muscle cells in vivo: improved LV function 64 Alginate Hydrogel formation gold nanoparticles in vitro: improved electrical communication between neonatal ventricular myocytes 65 PLA Electrospinning carbon nanotubes in vitro: mesenchymal stem cell differentiation to cardiomyocytes 66 PANI‐PLGA Electrospinning ‐ in vitro: synchronized beating of cardiomyocytes 67 Algisyl‐LVR Hydrogel formation ‐ Clinical trial: improved LV function 70 John Wiley & Sons, Ltd. 2. 1 Injectable Hydrogels Immobilizing cells in a 3D hydrogel is a common approach. 14 Depending on the required gelation time, cell type, site of injury and desired time of degradation, various material types and different gelation mechanisms can be considered. The hydrogels used in preclinical studies include for example synthetic Polyethylene glycol (PEG) 15 or natural occurring chitosan, 16, 17 fibrin, 18 collagens 19 and alginate. 20 Wang et al. demonstrated cardiomyocyte differentiation and enhanced cell survival by injecting brown adipose tissue‐derived stem cells that were immobilized in a thermo‐responsive chitosan gel. 16 Further gelation mechanisms described for cardiac applications are photopolymerization, 15 Ca 2+, 20 pH 21 or phosphate 16, 17 dependent gelation, or assembly due to physical interactions like in the case of fibrin. 18 Kraehenbuehl et al. developed a PEG‐vinylsulfone hydrogel system that undergoes Michael‐type addition reaction in situ for crosslinking. 22 Combined with cells and bioactive factors, these gels were able to preserve the contractile function of cardiomyocytes and resulted in decreased infarct sizes. 23 An advantage of in situ‐forming hydrogels is the possibility of a minimal invasive delivery, utilizing suitable catheters. For this approach, the hydrogel should form immediately after injection to the site of injury, whereas catheter blockage due to premature gelation has to be avoided. 13 In order to overcome this problem, Bencherif et al. described a method to generate a preformed but still injectable hydrogel with shape memory effect. 24 Here, alginate served as the base material, which was metharcylated to allow later radical polymerization. The at –20 °C polymerized, nanoporous, preformed alginate hydrogel can be pushed through a needle or catheter. As soon as the applied shear force is removed, the hydrogel returns to its original preformed shape. 24 For applications in the heart, hydrogels can either be seeded with cells, or they can be loaded with bioactive agents in order to induce (stem) cell migration or differentiation, and thus tissue regeneration ( Figure 1 ). However, despite promising results designing injectable cell‐loaded hydrogels, the amount of cells that effectively differentiated into cardiomyocytes in these studies was insufficient. 13 Nevertheless, a cell‐free strategy is highly attractive in regards of regulatory aspects, and since this approach is not limited to a specific patient, an off‐the‐shelf up‐scalable production is possible. Recent in vitro and in vivo studies focused on the application of bioactive drug‐releasing hydrogels. Currently used and entrapped bioactive molecules include small molecules like prostaglandins, 25 RNA, 26, 27 growth factors 28, 29 or various other factors such as sonic hedgehog 30 and bone morphogenetic protein‐2. 31 Figure 1 Biomaterials for cardiac applications. Injectable hydrogels but also cardiac patches are used as treatment options for cardiac damage. Both materials can either be cell‐seeded or loaded with bioactive molecules such as RNA, small molecules, growth factors or proteins. Improvement in wound healing and increased ECM remodeling has been described when encapsulating micro RNA (miR)‐29B in a collagen hydrogels. 27 In this system, a collagen solution was incubated with a four‐arm PEG terminated succinimidyl glutarate in order to induce gelation. 27 Another approach aimed for myocardial repair and regeneration after MI damage by the sustained release of stromal derived factor‐1 (SDF‐1)‐GPVI from a photopolymerizable PEG‐based hydrogel in order to attract circulating endothelial progenitor cells (EPCs). 32 In this study, the released protein was able to bind extracellular occurring collagen with the GPVI domain, and EPCs that expressed CXCR4 were recruited by SDF‐1. 33 Due to the non‐fouling properties of PEG, the encapsulated proteins do not bind to the PEG‐molecules and can therefore be easily released from the gel. 34 Furthermore, drug delivery systems can be designed by using self‐assembling peptides. Guo et al. introduced a heparin‐binding domain to the self‐assembling peptide RADA16. The heparin‐binding vascular endothelial growth factor (VEGF) was entrapped in the biomaterial by first binding heparin to the newly designed RADA16 peptide before adding VEGF. 29 In vivo MI studies confirmed effectiveness of the generated biomaterial, since microvessel formation, cardiomyocyte survival as well as an increased cardiac function was observed. 29 An enhanced left ventricular ejection fraction (EF) in a model of ischemic cardiomyopathy was described in a study, where neuregulin‐1b growth factor was encapsulated in a hydrogel consisting of hyaluronic acid macromers with HEMA group modifications. 28 Very interesting results have been also achieved when injecting natural occurring biomaterials like ECM or selected ECM components. 35, 36 It has been speculated that these materials are advantageous due to the presence of native structures and architecture; however, chemically or enzymatically treated tissues lack of biochemically important matricellular proteins and proteoglycans, as well as microfibrillar structures. 37 Furthermore, natural occurring polysaccharides such as xylan, 38 dextran, 39 pullan, 39 chitosan, 16 hyaluronan, 40 and alginates 41, 42, 43 have been investigated for myocardial applications. Existing patents related to polysaccharide‐based strategies are nicely summarized by Silva et al. 44 A successful first‐in‐man clinical trial has been performed with alginate (IK‐5001). Interestingly, the alginate hydrogel alone preserved left ventricular function and left ventricular ejection fraction. 45 When designing hydrogels for cardiac applications, not only hydrogel composition and ingredients impact in vivo performance, but also the mechanical strength and stiffness of the hydrogel are important parameters. MI for example leads to wall thinning 46 and therefore a hydrogel is needed to mechanically support the damaged zone. However, it should be considered that hydrogels with too much stiffness can cause diastolic dysfunctions. 47 Therefore, it is important to specify adequate mechanical properties. In addition, it is known that cells, including stem cells, respond to their microenvironment, and hydrogel stiffness has been shown to play a major role in stem cell differentiation. 5, 48 Mesenchymal stem cells (MSCs) were seeded onto hydrogels with different stiffness. Neural differentiation occurred when the cells were cultured on a 0. 1 kPa hydrogel, whereas 11 kPa resulted in myogenic tissue, and 34 kPa were necessary to induce osteogenic differentiation. 5 To date, no best‐practice guidelines for optimal parameters with regards to therapeutic efficacy such as concentrations of bioactive agents, degradation rate and stiffness have been defined for biomaterials design. 13 Future studies must focus much more on optimizing the biomaterials design by considering mechanical aspects. Although injectable hydrogels are currently under intense investigation, there are also other biomaterials‐based strategies available to support cardiac repair. For example, solid preformed scaffolds, so‐called patches, can be applied epicardially onto the damaged heart. Just like hydrogels, these scaffolds can either be seeded with cells or they can be loaded with bioactive molecules (Figure 1 ). Patches do not only have a therapeutic effect by releasing bioactive molecules or delivering cells, they also provide mechanical support and are able to reduce dilation. 49 However, certain requirements need to be considered when engineering a cardiac patch: The patch should be stable but flexible and provide mechanical strength; however, too stiff materials can induce diastolic dysfunctions and can lead to non‐contractile constructs. 50 In most cases, it is desired that the patch degrades after sufficient remodeling. 49, 50 Although the selection of an adequate material is a crucial issue, many studies do not consider the mechanical properties of the chosen material. The stiffness or E‐modulus of human myocardium has been described with 0. 02–0. 5 MPa, and the tensile strength was specified with 3–15 kPa. 51 Prabhakaran et al. generated a nanofibrous composite scaffold of poly( dl ‐lactide‐co‐glycolide) (PLGA) and gelatin via electrospinning. The authors impressively showed that the stiffness of the material could be manipulated by introducing the natural occurring polymer gelatin. Gelatin decreased the E‐modulus and furthermore enabled in vitro integration of cardiomyocytes. 50 In addition, nanofibrous poly‐ε‐caprolactone (PCL) was identified as a beneficial scaffold for cardiomyocyte attachment. 52 The range of material stiffness can also be influenced by synthesizing the polymer at different temperatures. This phenomenon has been proven with molded poly (glycerol sebacate) (PGS) foils, which were polymerized at 110 °C, 120 °C and 130 °C. 51 It is important to notice that scaffold or patch characteristics highly vary between multiple manufacturing methods that are used for scaffold fabrication. 2. 2 Electrospinning to Generate a Cardiac Patch Electrospinning for example is a well described method for creating fiber‐containing and highly porous scaffolds. 53 The principle of electrospinning is schematically displayed in Figure 2. Briefly, a polymer is dissolved in a solvent and pumped through a syringe. The droplet on the tip of the syringe first forms a cone when a high voltage is applied. As soon as the electrical field is increased, a fiber ejects from the droplet, which travels in spinning motions to the counter electrode where the solvent evaporates and a fibrous scaffold is formed. 54, 55, 56 By adjusting parameters such as polymer concentration and molecular weight, solvent volatility and conductivity, electrode distance, needle size, flow rate and voltage, the fiber shape, fiber diameter and pore size can be influenced. 57 Nanofibrous scaffolds of electrospun PCL have been shown to be beneficial for cardiac applications. 52 Furthermore, electrospun polycarbonate urethane scaffolds exhibited comparable mechanical properties such as the native myocardium. 58 In vivo data indicated that a cell‐seeded electrospun PCL‐gelatin patch can provide sufficient mechanical strength to the myocardium, promotes angiogenesis and decreases scar size after MI. 59 In addition, large animal in vivo studies confirmed that an electrospun PGS‐fibrinogen patch modified with VEGF can lead to an improved ejection fraction. 60 With electrospinning it is not only possible to mimic the structural and mechanical properties of the native heart muscle, our group has shown that it is also possible to introduce crucial biochemical cues like proteoglycans to the electrospun fiber. 61 Figure 2 Electrospinning of patches for regenerative medicine and tissue engineering applications. A) General electrospinning set up: A polymer solution is pumped through a nozzle and forms a drop on the tip. The ejected fiber travels to the collector in spinning motions. Since the solvent evaporates, a randomly oriented solid fiber mat is deposed on the collector. B) Scanning electron microscopic image of an electrospun scaffold. C) The droplet on the needle tip forms a cone in an electrical field. As soon as the electrical field strength exceeds the surface energy of the droplet, a thin fiber is ejected. 2. 3 Further Patch‐forming Approaches Another method used to generate cardiac patches is thermally induced phase separation. With this technique it is possible to create well‐defined interconnected micropores utilizing simple equipment. 62 To improve cardiac remodeling, polyester urethane urea (PEUU) has been processed into patches using phase separation techniques. 63 Furthermore, Matsubyashi et al. engineered a muscle graft by knitting. Here, the authors reinforced a caprolactone‐co‐ l ‐lactide (PCLA) sponge with knitted poly‐ l ‐lactide (PLLA) in order to increase cardiac function. 64 Another interesting approach is the generation of conductive biomaterials that support the cardiomyocytes' ability to contract. Several groups intensively studied these highly functional biomaterials in order to engineer an optimal cardiac patch. 49, 65, 66, 67 Accordingly, a contracting patch could be obtained when using conductive polymers like polyaniline (PANI), 67 or by manipulating non‐conductive polymers with conducting elements such as carbon nanotubes 66 or biocompatible gold nanoparticles. 65 In terms of cytocompatibility, conducting polymers including PANI, 49 polypyrrole, 68 and polythiophene 69 have been successfully tested with various cell types, resulting in enhanced proliferation, adhesion and differentiation. Although these materials seem to be adequate for cardiac patch generation, a key limiting factor is that they are non‐degradable. In order to address this issue, low molecular weight oligoanilines have been examined. 49 Molded films were prepared by synthesizing a biodegradable polyurethane‐containing aniline pentamer, which was then blended with PCL. 49 Despite promising in vitro studies, these materials need to be further investigated in appropriate in vivo models. In regards to clinical translation, the innocuousness of the degradation products has to be carefully monitored. 2. 4 Future Directions for Usage of Biomaterials for Myocardial Repair To date, various methods to treat the injured myocardium have been studied. Cardiac patches showed encouraging results; however, the surgical procedure, which is required for implantation represents a drawback of this approach. In addition, conflicting results regarding the success of cardiac patches were generated, so there is a need for further long‐term in vivo studies. Moreover, since there are difficulties of finding the right cell source for myocardial tissue engineering, injectable cell‐free biomaterials may be more suitable for clinical use. It has been shown that injectable biomaterials provide mechanical strength 70 and therefore prevent negative remodeling. 47 Since too stiff materials lead to diastolic dysfunctions, it is necessary to carefully examine in future studies the mechanical properties needed for sufficient myocardial treatment. Furthermore, the question of, how the materials can be safely applied to patients needs to be answered. Alginate and fibrin are promising injectable biomaterials, which have already underwent clinical trials. However, the long‐term effects are not examined and in our opinion further investigations need to be performed. 3 Biomaterials Used for Heart Valve Tissue Engineering Four valves regulate the blood flow in the normal heart: the semilunar pulmonary and aortic valves, and the atrioventricular tricuspid and mitral valves. 71 These valves enable a unidirectional blood flow through the heart by opening and closing approximately 100 000 times a day. 72, 73 This translates in more than 3 billion opening and closing movements over an average life cycle, where the heart valve leaflets are exposed to mechanical forces such as flow, tension and flexure. 71 Accordingly, engineering a material that withstands these loads but maintains the hemodynamics of the heart is quite challenging. Native heart valve leaflets have a highly sophisticated histoarchitecture in order to enable a lifelong performance. In detail, three layers form a heart valve leaflet: the fibrosa, which is mainly composed of collagens, the glycosaminoglycan‐rich spongiosa and the elastic‐fiber containing ventricularis. 74, 75, 76 It is hypothesized that the complex architecture of a heart valve leaflet is formed and maintained mainly by two different cell types, the ECM‐producing valvular interstitial cells (VICs), which are located predominantly in the spongiosa layer, and the valvular endothelial cells (VECs) that form a confluent endothelium as a barrier to the blood stream. 72, 74 The current gold standard to treat diseased heart valves is the replacement using mechanical or biological valves. Mechanical valves are advantageous since they have an excellent durability compared to bioprosthetic valves, which are less durable and in addition prone to degradation; however, the risk of thromboembolism is higher with mechanical valves. 77 A bioprosthetic substitute needs a glutaraldehyde fixation prior to implantation, a process that stiffens the matrix and inhibits repopulation. 77 None of these replacement options offers the potential of growth or remodeling after implantation. 72, 74, 75 A tissue‐engineered valve would be a promising possibility to address these limitations; however, most of the currently designed tissue‐engineered valves were pre‐clinically or clinically unsuccessful due to problems with calcification or fibrosis. 75 A non‐thrombogenic and non‐calcific prosthesis with adequate mechanical properties and sufficient strength that maintains the native valve hemodynamics has so far not been designed. 72 There are two different approaches to create a tissue‐engineered heart valve substitute. One is the traditional idea of seeding autologous cells onto a 3D scaffold in vitro prior to implantation. The other approach involves tissue regeneration and relies on implanting a cell‐free construct, which enables material‐guided reseeding in vivo. 72 It has been proposed that the most adequate substrate in terms of mechanical and functional behavior is a decellularized xeno‐ or homograft. 78, 79, 80 These substrates already exhibit the complex native valvular structure and architecture, but their restricted availability (homograft) as well as their risk of transferring zoonoses are limiting factors and major drawbacks of these constructs. 80 Synthetic and natural polymers have been demonstrated to be an alternative for homografts and xenografts ( Table 2 ). The techniques used to generate a valve‐shaped 3D scaffold range from simple molding methods like freeze molding collagen type I and elastin to form a two‐layered leaflet, 81 to complex multi‐step approaches. Salt leaching is an interesting method to generate porous pre‐formed heart valve substitutes by adding sodium chloride crystals to a polymeric solution. 82 After the polymer hardens, the salt crystals can be washed out and a porous scaffold is formed. Polyhydroxyalkanoate (PHA) for example has been described to be favorable to fabricate a tri‐leaflet heart valve with salt leaching, which can be successfully seeded with vascular cells. 82 Another in vitro study described poly (glycerol sebacate) (PGS) and PCL as a suitable polymer combination to support the growth of MSCs and VICs, and in addition mimics the anisotropic mechanical properties of native heart valve leaflets. 83 In order to create such a material, the combination of different fabrication methods was necessary. First, PGS was molded, and in a second step PCL was electrospun onto both sides of the PGS sheet resulting in a three‐layered scaffold with good mechanical properties. 83 Combining different biomaterials but also using variable fabrication techniques in order to generate an optimal and complex cytocompatible scaffold is an increasing trend in current studies. Weber et al. created a tubular fibrin gel loaded with umbilical artery cells. They further strengthened the construct with a textile co‐scaffold. 84 In contrast to the molded fibrin, the textile was a thermostabilized wrap‐knitted PET scaffold. With this technique it was possible to reproduce the valvular geometry and the mechanical strength. 84 Table 2 Overview of biomaterials and fabrication methods suitable for heart valve tissue engineering Biomaterial Fabrication Method Cells/Molecules included Application and Results Ref. ECM Decellularization Endothelial cells and myofibroblasts in vitro: matrix characterization and reseeding in a bioreactor 76 Homograft Decellularization ‐ Clinical trial: excellent function, no thrombus formation 79 Elastin and collagen Molding ‐ in vitro: bi‐layered material characterization and cell‐matrix interaction studies 81 PHA Salt leaching ‐ in vitro: viable ECM formation in a bioreactor 82 PGS‐PCL Micromolding – Electrospinning ‐ in vitro: 3‐layered construct supported growth of VICs and MSCs, ECM deposition 83 Fibrin gel and PET mesh Hydrogel formation and knitting Umbilical artery smooth muscle cells/myofibroblasts in vitro: enhanced mechanical properties and tissue formation in a bioreactor 84 PEG‐PLA Elektrospinning ‐ in vitro: biomimicking scaffold, cytocompatible with VICs and VECs 85 ECM Cell sheets Human fibroblasts in vitro: matrix characterization 86 ECM Decellularization Endothelial progenitor cell‐derived endothelial cells OR CD 133 antibody in vivo (large animal): CD 133‐conjugated leaflets exhibited a progressive recellularization across the entire leaflet, no calcification 87 PEG Hydrogel formation and micropatterning RGDS peptide in vitro: controllable morphology and activation of VICs via micropatterns 88 PGA mesh‐P4HB ‐ECM Decellularization after ECM production with vascular derived cells Mesenchymal stem cells in vitro: mechanical and biochemical characterization in vivo (primate): moderate valvular insufficiency, rapid cellular repopulation 89, 90 ECM Decellularization Umbilical cord‐derived endothelial cells in vitro: complete recellularization in a bioreactor (Mitral valve) 92 Fibrin gel and PET mesh Hydrogel formation and knitting Umbilical vein smooth muscle cells/fibroblasts in vitro: tissue development in a bioreactor, recapitualtes the native structure (Mitral valve) 95 John Wiley & Sons, Ltd. The idea of mimicking nature is a popular approach when aiming for an optimized heart valve substitute. The first step towards a bio‐inspired tissue is to obtain comprehensive data that describes the tissue's ultrastructure including cells and ECM as well as biomechanical properties. As soon as these critical parameters are defined, an adequate material and material generation technique can be selected. In a recently published study we identified the mechanical, structural and biochemical properties of a native heart valve leaflet. Based on this blueprint, we generated a scaffold by electrospinning a blend of PLA and photocrosslinkable PEG. 85 In this study, the E‐modulus was determined in every single leaflet layer—the fibrosa, spongiosa and ventricularis—by applying atomic force microscopy (AFM) on unfixed cryosections. 85 Another material fabrication strategy is to utilize ECM‐producing cells as scaffold producers. Fibroblasts for example produce a significant amount of ECM in the form of sheets, when cultured in vitro. 86 It has been shown that these ECM‐sheets can be produced, and subsequently layered in order to increase strength and to shape a valve. 86 Advantageous in this case is that the ECM‐sheets can be produced patient‐tailored using autologous cells. In order to induce a variety of cellular responses, scaffolds can be functionalized with bioactive molecules. It was reported that substrates, which were functionalized with a conjugated anti‐CD133 antibody, attracted autologous cells and showed complete cellular ingrowth after implantation in a sheep model. 87 In vitro studies using RGD‐modified micro‐patterned PEG hydrogels demonstrated VIC activation. 88 The results of this study indicate that both, biochemical as well as topographical modifications are important modulators when designing a heart valve replacement. Dijkman et al. described the fabrication of a non‐woven PGA‐coated scaffold utilizing poly‐4‐hydroxybutyrate (P4HB). The scaffold was seeded with ovine vascular cells using fibrin as a cell carrier. 89 Interestingly, exposing the cell‐material construct to dynamic strains by applying increasing transvalvular pressure differences enabled tissue maturation. The authors further described the possibility to minimally invasive deliver such an off‐the‐shelf valve. 89 Later studies with the PGA‐P4HB valve showed remarkable cellular repopulation after implantation in primates. 90 Such off‐the‐shelf cell‐free medical products are quite advantageous in terms of reaching clinical approval, since they are potentially more cost‐efficient and available at any time. The current trend of producing off‐the‐shelf materials and to combine it with minimally invasive implantation strategies is of high interest for the field of regenerative medicine. To date, the majority of studies that reported on tissue engineering of heart valves focused on the semilunar (pulmonary and aortic) valves. Since it has been described that it is essential to mimic the shape of the mitral valve, 91 methods such as decellularizing atrioventricular heart valves have been implemented in order to achieve or maintain a proper function. 92 In addition to the mechanical characteristics, the ECM as well as the residing cells of the mitral valve were intensively investigated. 93, 94 Moreira et al. demonstrated an elegant way to fabricate a textile‐reinforced mitral valve including annulus, asymmetric leaflets (anterior and posterior), and chordae tendineae. With this artificial mitral valve it is possible to maintain the native valve function and hemodynamics. 95 3. 1 Bioreactors for Valve Tissue Maturation Bioreactors have been used in the field of tissue engineering to mimic the biophysical signals that are present in the native organo‐physiological environment. Depending on the organ or the cell type, various forces such as strain, pressure, torsion or flow occur in a tissue. Cellular behavior like migration, proliferation and differentiation are highly dependent on external forces and can be regulated by inducing physical stimuli with bioreactor systems. 96 This has been shown to be also advantageous when generating a tissue‐engineered heart valve in vitro. 75 For example, an increased ECM production by vascular cells was observed on a porous scaffold within a pulsatile flow bioreactor. 82 An increased collagen and elastin production but also a significantly improved recellularization of the heart valve has been described under pulsatile conditions. 78 These results indicate that not only the 3D scaffold but also the forces that are applied on the scaffold are important in order to obtain a highly functional tissue‐engineered construct. 3. 2 Future Directions for Utilization of Biomaterials for Heart Valve Tissue Engineering Applications Decellularized human valves have already been successfully used in clinical trials. 79 Despite promising in vitro and in vivo experimental results, no tissue‐engineered heart valve construct has found its way into clinical reality. Similar to what we have described for myocardial tissue repair attempts, finding an adequate cell source is the major challenge. Interesting results were obtained by culturing MSCs on synthetic materials; 83 however, it seems like there is a trend to generate cell‐free off‐the‐shelf substrates. In our opinion, it is important to mimic the mechanical, structural and biochemical properties of a native heart valve in order to overcome issues like limited tissue growth in pediatric patients. To prevent calcification and to enable cell attachment after implantation, it will be necessary to further investigate the biomaterial‐tissue‐interface and to screen for suitable surface modifications. As reviewed here, many in vitro studies were successfully performed showing sufficient cell attachment, cytocompatibility and mechanical graft stability; however, long‐term in vivo studies are needed in order to determine the true performance of the tissue‐engineered valves. 4 Biomaterials Used for Blood Vessel Engineering Blood vessels are responsible for the nutrient and oxygen supply of all organs. As a reason of lifestyle or genetic abnormalities, vascular diseases such as artherosclerosis or aneurysms occur. Current treatments to address these problems are the transplantation of autologous grafts from other regions of the body, 97 or the implantation of stents and synthetic grafts. 97, 98 Clinically applied synthetic polymer substitutes are made of expanded‐polytetrafluoroethylene (ePTFE), Dacron or polyurethane. 99, 100 Vessels with a diameter larger than 6 mm can be successfully replaced with these materials. However, due to thrombus formation in vessels of smaller sizes, there is an ongoing need for vascular alternatives. 97 To date, there are different strategies pursued in order to engineer a functional blood vessel. These strategies are either tissue‐engineered cell‐seeded scaffolds or bioactive, cell‐free approaches ( Table 3 ). Table 3 Overview of biomaterials and fabrication methods used to engineer blood vessel substitutes Biomaterial Fabrication Method Cells/Molecules included Application and Results Refs ECM Cell sheets Fibroblasts in vivo (small animal and primate): anti‐thrombogenic, mechanically stable, tissue integration Clinical trials 101, 106 PGA mesh with ECM woven structure with decellularized ECM Endothelial progenitor cells in vivo (large animal): resistance to clotting and intimal hyperplasia 102 ECM Cell sheets from human induced pluripotent stem cells ‐ in vitro: smooth muscle cell differentiation and collagenous matrix generation in a bioreactor 103 PGA‐PCLLA Solvent casting Heparin with VEGF OR CD34 antibody in vivo: increased endothelial cell attachment 107 PLLA‐PLCL Phase separation Heparin in vitro: improved anticoagulation properties in vivo: neovascularization (subcutaneous) 108 Elastin‐like protein‐ collagen I Hydrogel formation ‐ In vitro: mechanical characterization in vivo: limited early inflammatory response 109 PET‐PGA woven ‐ in vivo: mechanical integration with the aorta 110 PCL reinforced with PET freez‐dried tube reinforced with knitted PET ‐ Increased mechanical properties 111 PGA woven Mononuclear cells Clinical trial: no clacification or infection but stenosis in 4 cases 112 PLA PCL‐collagen I – ECM Elektrospinning Cell sheet Smooth muscle cells in vitro: improved cell viability 114 Agarose Bioprinting Mouse embryonic fibroblasts Development of a self‐supporting structure 115 John Wiley & Sons, Ltd. Tissue‐engineered cell‐seeded constructs have been intensively studied in order to replace a vein or artery. The inner layer of a blood vessel, the so‐called intima, is lined with endothelial cells. 80 A major drawback in engineering small diameter blood vessels is thrombus formation and inflammation due to non‐sufficient endothelialization. 80, 101 Therefore, either the seeding technique needs to be improved or more functional materials have to be developed. One approach to generate more functional implants is the culture of cell‐seeded scaffolds under defined and highly physiological conditions in a bioreactor system prior to implantation. 102, 103 Regarding the biomaterial tube itself; many different materials have been described. 80, 104 An interesting idea is the generation of native ECM sheets produced by smooth muscle cells (SMCs) from another species 105 or from human induced‐pluripotent stem cell‐derived SMCs. 103 The engineered vessels were then decellularized, leaving behind a robust ECM graft. In a final step, the decellularized grafts were seeded with endothelial cells of the graft recipient. 102, 103 L′Heureux et al. published a similar ECM‐sheet technique but utilized human fibroblasts. This in vitro‐engineered blood vessel showed complete integration after implantation into nude rats. 101 The approach has already been transferred into clinics, where a successful implantation of the tissue‐engineered blood vessel was performed. 106 However, when aiming to generate a graft for bypass surgery, the time it takes to produce autologous tissue‐engineered implants is rather long. Therefore, off‐the‐shelf small diameter blood vessels are more favorable. To achieve in vivo endothelialization of these acellular grafts, scaffold functionalization with chemokines and growth factors like SDF‐1 and VEGF can be performed. 98, 107 Figure 3 summarizes various possibilities to introduce bioactive agents to a scaffold. Wang et al. covalently linked heparin as an anticoagulant to a thermally induced phase separated scaffold made of PLLA and poly (l‐lactide‐co‐ε‐caprolactone) (PLCL). Interestingly, the scaffold enabled endothelial cell attachment, but showed low protein attachment, which is important to avoid stenosis, thrombus or neointima formation. 108 Another strategy for designing an acellular but functional scaffold for blood vessel replacement is the application of structural ECM components, such as collagens and elastin. 109 In addition to biochemical aspects, vascular grafts need to fulfill mechanical requirements in order to perform in vivo without complications. Due to the high blood pressure especially in arteries, both, biomaterial and scaffold design techniques need to be carefully selected. Woven core sheath fibers made of polyethylene terephthalate (PET) and polyglycolic acid (PGA) were used to enhance the aortic wall strength for aneurism repair. 110 In this study, the PGA component was replaced over time with native tissue in vivo, hereby enabling integration of the remaining scaffold, which helped stabilizing the vessel. 110 In another study, mechanical properties such as tensile strength, E‐modulus, compression recovery and radial compliance were significantly improved by reinforcing a PCL scaffold with a knitted PET fabric. 111 Woven PLA and PGA grafts seeded with mononuclear cells were implanted in patients without signs of calcification or infection; however, in four cases stenosis occurred. 112 Besides molding, weaving, phase separation and knitting, electrospinning has been used to generate vascular grafts. By utilizing a rotating mandrel as the collector, tubular scaffolds can be electrospun. 113 Also the combination of electrospinning and cell sheet engineering has been described as a promising vascular graft generation technique. 114 Another scaffold generation method that has been studied intensively for generating macro‐ and microvascular structures is the computer‐controlled layer‐by‐layer deposition technique named printing. 115 It is distinguished between three fabrication techniques: laser‐based, printer‐based and nozzle‐based. 116 Laser‐based systems such as stereolithography and multiphoton polymerization are used to generate scaffolds from a bath of photosensitive polymers, enabling vessel sizes down to the sub‐micrometer range. 117 Printer‐based or drop‐on‐demand systems include thermal and piezoelectric inkjet printers, which are also able to directly print cells. 118 The third technique is direct writing of polymers using pressure‐assisted nozzle‐based systems. 119 Depending on the system, various natural and synthetic biomaterials can be processed into a 3D tubular construct. 116 Figure 3 Schematic depiction of scaffold design strategies in combination with bioactive cues. The polymeric scaffold is shown in grey. Bioactive molecules ( pink ) can be delivered by A) encapsulation, B) physical adsorption on the surface, C) chemical binding ( orange ) to the surface or (C) by blending two materials. 4. 1 Elastic Fiber Generation in Tissue‐Engineered Vascular Grafts Although considerable progress has been made in the field of blood vessel tissue engineering, there is still a major limitation for all of these constructs – they lack of the presence of functional elastic fibers. Elastic fibers are responsible for tissue elasticity and resilience, especially in load bearing tissues such as heart valves, skin and also blood vessels. 120 The elastin content in blood vessels is around 30–50%. 121 Successful elastic fiber generation in tissue‐engineered constructs has been described in animal models; 101 however, mature human cell‐based elastic fibers have so far not been realized. There have been a few reports using antibody staining, western blots or PCRs to detect elastin/tropoelastin protein or gene expression. 122 However, the elastin protein is not the only protein forming a functional elastic fiber. Mecham et al. thoroughly investigated the processes that lead to normal elastic fiber assembly and found that cells secrete the soluble precursor tropoelastin that subsequently agglomerates in the extracellular space. 123, 124 Simultaneously, fibrillin containing micro‐fibrils are assembled, to which in a next step the tropoelastin agglomerates are placed. Finally, lysyl oxidase (LOX) crosslinks the deposited elastin molecules to form the elastic fiber. 123, 124, 125 To enable the process of elastogenesis, further proteins such as fibulins, elastin microfibril interface located protein 1 (EMILIN‐1), fibronectin, latent TGF‐beta binding protein 4 (LTBP‐4) are necessary. 123, 126 The process of elastogenesis is highly complex and still not completely understood; however, it has been shown that these events already play a major role in early human cardiovascular development. 127 Therefore, when aiming for a nature‐mimicking highly functional biomaterial scaffold, elastic fibers and the elastic fiber‐associated proteins must be considered. 4. 2 Future Directions for the Use of Biomaterials in Vascular Graft Design The most challenging part in blood vessel tissue engineering is the generation of small diameter grafts (<6 mm). Today, decellularized human vascular grafts are commercially available; however, the limiting factor is the lack of donors. The cell sheet technique seems to be quite promising, 106 but long production times as well as high costs are unfavorable. Vascular graft design based on the use of solely synthetic biomaterials can be limited due to low hemocompatibility or unfavorable biomechanical profiles, either too stiff or too weak, which can potentially lead to graft leakage or dilation. In order to address these problems, we believe that mimicking the biological and mechanical properties of the native vessel will be important next steps towards a tissue‐engineered blood vessel. The use of hybrid substrates manufactured from synthetic materials and natural ECM proteins is highly promising. Further studies will be required to improve our knowledge of normal and pathological vascular ECM development and remodeling. 5 Monitoring of Biomaterials Histological, histochemical and immunohistological methods are currently the gold standard for analyzing detailed 3D architectures of cells and ECM components. 128, 129 Although, this information is of high value to evaluate failure or success of an experimental approach, histological processing is not suitable for online process monitoring, as it requires the sacrificing of samples at each time point of the experiment. Since in vitro‐generated advanced therapy medicinal products (ATMPs) or biomaterials are designated to be implanted into a patient, non‐destructive methods are required to control their quality, structure and composition. Implementation of techniques that allow the monitoring of the progress of tissue formation and maturation time‐dependently in vitro is a goal not only for the field of tissue engineering. 128 Clinically approved imaging technologies that visualize tissues and organs within the human body to diagnose, treat and monitor disease stages represent a promising approach that could be adapted to be employed for the screening of tissue‐engineered materials. 128, 130 Computer tomography (CT), magnetic resonance imaging (MRI) and ultrasound (US) are imaging technologies that can provide depth‐resolved information of structural tissue features and are therefore of high interest for the label‐free assessment of in vitro‐engineered ATMPs and biomaterials. 131 In addition, novel optical techniques employing advanced laser systems to excite specifically light scattering, multiphoton absorption or intrinsic fluorescence were already adapted to provide valuable information of 3D tissue compositions and architectures. 132 In the following, we will discuss principles, advantages and limitations of optical techniques suitable to assess biomaterials and ATMPs in a label‐free fashion ( Table 4 ). Table 4 In vivo and in vitro applications of established and novel optical techniques Technique Energy Source of contrast Penetration depth In vivo application In vitro application Ref. CT X‐rays Absorption Throughout the human body Bone fractures Tumors Mineralization in cell cultures, Scaffold structure 128, 130, 133, 134 US imaging Ultrasonic waves Reflection, attenuation 1–3 cm Fibrosis Tumors ECM remodeling 128, 137, 138, 139 MRI Radio‐frequency pulses Relaxation times of dipolar molecules Throughout the human body Presurgical imaging ‐ 128, 144, 145 OCT Non‐coherent light Reflection 1–3 mm Cardiology Tumors Ophthalmic Scaffold structure ECM remodeling 128, 129, 132, 151, 152 Raman microspectroscopy NIR laser beam Molecular vibrations 100–300 μm Gastric neoplasia Skin cancer Cell and ECM identification and monitoring 162, 163, 164, 165, 166 Mutiphoton absorption SHG FLIM Pulsed NIR laser beam Auto‐fluorescence and decay times of intrinsic fluorophores 200 μm Skin cancer Fibers and fibrils in tissues, metabolic profiling, Scaffold structures 63, 132, 180, 187, 189, 190, 194, 195 John Wiley & Sons, Ltd. 5. 1 Adaption of Clinically Approved Optical Techniques for In Vitro Applications Electromagnetic radiation is employed for imaging of the human organs or even the whole body; measurements of spatially resolved absorption, refraction, attenuation or scattering of the incident energy are then transferred into image contrast depicting physical tissue properties. 128, 130, 131, 132 In clinics, X‐rays, light, US or electromagnetic fields are employed as radiation source. 128, 130, 131 X‐rays, the oldest imaging modality, utilize an electronic beam to irradiate tissues. 130 From its historical development till today, X‐ray beams were predominantly employed for displaying bone tissues and fractures since mineralized tissues highly absorb the electronic beam resulting in excellent image contrast. 130 The computational reconstitution of 2D and 3D X‐ray absorber maps, denoted as CT, vastly increases image contrast and can for instance display small tumors and pulmonary alveoli. 130 By miniaturizing the electronic beam, μ‐CT can achieve spatial resolutions up to nanoscale levels. 128, 130 For in vitro applications, μ‐CT has been successfully implemented to monitor and quantify cell‐mediated mineralization processes, spatially resolved and even under dynamic conditions in a bioreactor system. 133 Additionally, μ‐CT was employed on manufactured scaffold materials to provide quantitative measures of fibers, pore sizes and porosity. 128, 134 However, strongly different attenuation times of X‐rays are difficult to resolve; hence, multiple biomaterial composites such as natural ECM deposition within porous scaffolds could not be distinguished using μ‐CT. 135 In summary, μ‐CT is neither optimal to resolve soft tissue architectures nor is it capable to distinguish biochemical ECM components. 128, 131, 136 In contrast, US imaging employs acoustic waves and provides good soft tissue contrast. It is clinically established to detect fibrotic tissues and tumors. 128, 130 Dependent on the acoustic wave frequency, spatial resolutions can exceed 1 μm, although this showed a negative impact on the penetration depth. 128, 130 Nevertheless, ultrasound is considered a promising technology to monitor ECM formation and remodeling in tissue‐engineered constructs. 128, 137, 138, 139 Cell death events, biomaterial degradation and tissue development have been shown to affect the output US image. 128, 138, 139, 140 However, treatment of cell cultures with low‐intensity ultrasonic waves has been indicated to promote cell proliferation and differentiation in cultured stem cells. 141, 142 Therefore, continuous US imaging during cell and tissue culture could impact the quality of tissue‐engineered constructs. Moreover, ultrasonic imaging is highly sensitive to air bubbles, speckled‐noise and therefore error‐prone for dynamic culture systems. 128, 143 MRI excites nuclear magnetic resonance effects and transfers these into an image contrast. Resonant radiofrequency pulses are generated by an electromagnetic field, displacing the orientation of protons in dipolar molecules and measuring their relaxation times. 130 These parameters are highly tissue‐specific, resulting in excellent soft tissue contrast in optical sections that can be acquired throughout the human body. 128, 130 MRI is therefore regarded as the most powerful modality for pre‐surgical and diagnostic imaging. 144, 145 In vitro high‐resolution MRI has been applied to visualize tissue ultrastructure and to assess biochemical measures such as the glycosaminoglycan content in engineered cartilage. 128 Gründer et al. depicted a correlation between histological polarized light microscopy and MRI images and demonstrated that collagen fiber orientation is visible in MRI. 146 However, resolution of MRI is still limited, requiring the application of contrast enhancing magnetic iron particles to optimize tissue contrast and to resolve cellular structures. 128, 145 In addition, MRI acquisition is time consuming and expensive, 144 making it therefore currently unsuitable for real‐time monitoring experiments. Optical coherence tomography (OCT) employs near‐infrared low‐coherent light sources and measures its optical reflectance to spatially display cross‐sections of tissue microstructures. 130, 147, 148 Resolution of OCT can achieve up to 1 μm and exceeds US imaging. 149 Altered tissue morphologies were visualized and could be assigned to disease stages in correlation to histological and histopathological diagnosis. 149 In clinics, OCT is utilized in cardiology and oncology, as well as to diagnose dermal inflammations and ophthalmic diseases. 129, 130, 150 Due to different refractive indices of polymers and water, pores in polymeric scaffolds can appear darker in OCT images, which then allow the quantification of pore sizes. 151 Penetration depth and resolution of OCT enables the discrimination of different tissue layers such as the epidermal and dermal layer of skin. 129, 147 Monitoring of ECM deposition, remodeling, cell migration, and proliferation was accomplished on tissue‐engineered constructs in vitro by utilizing OCT. 128, 129, 132 Bagnaninchi et al. implemented OCT into a perfusion bioreactor system and detected an increase in ECM formation under shear stress conditions. 152 Currently, penetration depth presents the limiting factor for OCT. 128, 149 In addition, the information gained by OCT does not provide much detail and it is therefore not sufficient to resolve cellular morphologies and characteristics. 129 5. 2 Laser‐Based Strategies to Assess Biomaterials and Tissue Architectures Raman spectroscopy is a non‐destructive method, established in pharmaceutical and polymer industries to identify substances based on their specific molecular vibrational patterns. 153 Laser light is employed for exciting these intrinsic molecular vibrations. Each molecule within the sample triggers specific frequency shifts relative to the incident light that is then detected within the Raman spectrum. 154, 155 Water has a minor Raman scattering signal, which does not interfere with the peaks of organic substances, making the technology advantageous for biological systems. 156, 157 For tissues and viable specimens, near‐infrared (NIR) lasers are utilized for Raman excitation, thereby problems of photo degradation, sample heating or fluorescence emission are nearly excludable and collection of signals from deeper tissue regions is facilitated due to a low level of light absorbance. 156, 157 Raman microspectroscopy, an approach where the laser system is coupled to a microscope, can be conducted on very small sample volumes. 157, 158 Spatially resolved molecular information can be provided by Raman microspectroscopy facilitating the analysis of different tissue structures and the analysis of single cells under physiological conditions 158, 159, 160, 161, 162, 163 ( Figure 4 ). Components of the ECM such as collagens, elastin and proteoglycans exhibit according to their molecular diversity specific Raman fingerprint spectra allowing for their distinction. 162, 164, 165, 166 The enzymatic degradation of collagen fibrils in aortic heart valves has been shown to result in a down‐regulation of collagen‐specific Raman signals. 164 Due to the non‐invasive and contact‐free measurement mode, Raman microspectroscopy is suitable to be implemented into a bioreactor system and was accomplished for time course monitoring of ECM formation under dynamic conditions. 167, 168 As tissue resident cells exhibit specific signals for nucleic acids, which are absent in ECM structures, single cells and even subcellular structures are resolvable by Raman microspectroscopy. 163, 167 In addition, changes of the cell phenotype such as altered protein expression result in a shift of the cell‐specific Raman fingerprint spectrum, empowering the technique as a method to characterize cells after isolation and to monitor the cellular phenotype in 2D and 3D culture systems. 159, 162, 169, 170 The molecular sensitivity of Raman microspectroscopy further facilitates to monitor cell death modalities and to distinguish apoptotic and necrotic cell death in vitro. 160, 161, 171 Acquiring Raman spectra simultaneously with standard fluorescence images allows the identification of defined spectral patterns and assists with the evaluation of the technology with established marker‐based methods. 160, 172 Moreover, comparing molecular vibrational spectra of native and in vitro‐engineered tissues allows the definition of qualitative and quantitative quality criteria for tissue assessment. 173 Scanning whole tissue sections by Raman microspectroscopy and transferring the resultant spatially resolved biochemical information into image contrast has been shown to enable highly sensitive marker‐free depiction of different tissue components and structures, analogously to histological staining 174, 175, 176 ( Figure 5 ). Here, resolutions of up to 1 μm can be obtained; 175 however, for scanning large tissue areas, long acquisition times are required due to the low cross‐section of Raman scattering. 177 Currently, many research groups focus on the combination of Raman spectroscopy with other optical tools, or to identify strategies how to improve the Raman efficiency by utilizing specific enhancement effects. 177, 178, 179 Compared to spontaneous Raman spectroscopy, fluorescence signals have multiple folds higher signal efficiencies. 132, 177 Employing confocal microscopy allows the acquisition of optical sections that can display the spatial distribution and networks of multiple structures in 3D tissue‐engineered constructs. 129, 132 While various fluorescence dyes have been designed for live cell imaging, and genetic modifications can achieve the expression of fluorescence reporter proteins, 129 none of these strategies fulfills the requirement of being completely non‐invasive and label‐free. Imaging endogenous fluorophores employing multiphoton effects represents a completely label‐free approach. 56, 180 Two multiphoton mechanisms can be discriminated – second harmonic generation (SHG) and two‐photon excited fluorescence (TPEF). 56, 180 In both, NIR laser pulses excite the tissue sample, having the advantage of increased axial resolution, penetration depth and low levels of photodamage on the sample. 132, 180 Multiphoton signals are shifted towards shorter wavelengths compared to the incident light. 132, 180 SHG visualizes non‐centrosymmetric molecular structures and results in precise signals at exactly half of the incident laser wavelength. 181 Three‐dimensional organization of fibrillar collagens, microtubuli structures, and myosin in muscle tissues have been successfully monitored using SHG. 181, 182 TPEF, which results in broader emission spectra, was employed to visualize and characterize metabolic states in cell cultures, based on the detection of intrinsic enzymatic co‐factors such as NADPH, porphyrins, and flavins. 183, 184, 185, 186 In addition, elastic fiber networks in native tissues are excitable by TPEF. 56 It has been shown that coupling SHG and TPEF microscopy to a multimodal imaging system can provide qualitatively and quantitatively information on 3D collagen and elastic fiber networks. 56, 76, 187, 188 Multiphoton microscopy can therefore detect damaged ECM structures and represents a possibility to monitor fiber assembly, maturation and remodeling in time lapse. 76, 187, 189, 190 Based on multiphoton signals of cardiac myosin and collagen fibrils, the level of fibrosis in rat hearts after myocardial infarction was quantitatively assessed. 182 Moreover, it was demonstrated that stem cell differentiation and stem cell fate commitment can result in changes of multiphoton signals. 56, 185, 191, 192, 193 Stem cell‐derived cardiomyocytes have been identified by visualizing their mature myosin filaments using SHG. 191, 192 Buschke et al. detected an increased TPEF signal of NADH due to cardiac commitment in embryonic bodies. 193 Figure 4 Principle of Raman microspectroscopy for characterizing cells under physiological conditions. A NIR laser is focused through a microscopic objective and directed onto a single cell or different subcellular regions. Resultant Raman spectra are analyzed using computer‐based multivariate algorithms. Figure 5 Schema of the principle of label‐free identification of different tissue stages based on molecular vibrational signals obtained by Raman spectroscopy: A) healthy tissues, B) early onset of pathological remodeling, and C) end‐stage pathological tissues show different signal patterns, which correlate to the histological findings (Movat's pentachrome staining). Polymeric scaffolds made of synthetic materials, which are widely used for tissue engineering approaches, exhibit broad fluorescence emission in both, SHG and TPEF channels. 129, 180 Although, the integrity and structure of polymeric scaffolds can be assessed by multiphoton imaging, 132, 180 it is challenging to separate strong polymeric autofluorescence signals from those of natural ECM deposits and cellular metabolites. 132 Fluorescence lifetime imaging (FLIM) offers a possibility to improve the resolution for multiple intrinsic fluorophores, exhibiting similar emission spectra. 184, 186 Time‐ and spatially resolved analysis of fluorescence signals generates a new dimension to characterize the architectures of in vitro‐engineered tissues. 186 FLIM represents a time‐resolved analysis of fluorescence signals that are transferred into image contrast. 56, 186 Measured fluorescence time decays enable the distinction of different fluorescence molecules, which can further be impacted when the (bio)chemical environment surrounding a fluorophore is changed due to induced manipulation or disease. 56, 186 FLIM has been demonstrated to be a potential technique to non‐invasively monitor cell fate commitment and stem cell differentiation. 194, 195 Furthermore, FLIM depicts metabolic parameters supporting the assessment of cell viability, which is of great interest when studying cell‐biomaterial interactions. 32 In summary, non‐invasive molecular‐sensitive monitoring techniques are suggested to be more robust and reliable than biochemical assays. 32 These new technologies may complement, or in the near future even potentially replace, histological, histochemical or immunohistological screening, which make it necessary to sacrifice tissues for analyses at a certain time‐point of the manufacturing process. 129, 132 However, to date, validation with classical invasive assays is still required for the interpretation and evaluation of multidimensional datasets obtained by these new techniques. In the future, a device that combines several non‐destructive optical methods could gather multidimensional quality parameters and support continuous monitoring of biological manufacturing processes.
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10. 1002/adhm. 201400781
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Advanced healthcare materials
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Fabricated elastin
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The mechanical stability, elasticity, inherent bioactivity, and self-assembly properties of elastin make it a highly attractive candidate for the fabrication of versatile biomaterials. The ability to engineer specific peptide sequences derived from elastin allows for precise control of these physicochemical and organizational characteristics, and further broadens the diversity of elastin-based applications. Elastin and elastin-like peptides can also be modified or blended with other natural or synthetic moieties, including peptides, proteins, polysaccharides and polymers, to augment existing capabilities or confer additional architectural and biofunctional features to compositionally pure materials. Elastin and elastin-based composites have been subjected to diverse fabrication processes, including heating, electrospinning, wet spinning, solvent casting, freeze-drying, and cross-linking, for the manufacture of particles, fibers, gels, tubes, sheets and films. The resulting materials can be tailored to possess specific strength, elasticity, morphology, topography, porosity, wettability, surface charge and bioactivity. This extraordinary tunability of elastin-based constructs enables their use in a range of biomedical and tissue engineering applications such as targeted drug delivery, cell encapsulation, vascular repair, nerve regeneration, wound healing, and dermal, cartilage, bone and dental replacement.
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No full text available
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10. 1002/adhm. 201500005
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Advanced healthcare materials
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Photocrosslinkable Gelatin Hydrogel for Epidermal Tissue Engineering
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Natural hydrogels are promising scaffolds to engineer epidermis. Currently, natural hydrogels used to support epidermal regeneration are mainly collagen- or gelatin-based, which mimic the natural dermal extracellular matrix (ECM) but often suffer from insufficient and uncontrollable mechanical and degradation properties. In this study, a photocrosslinkable gelatin (i. e. , gelatin methacrylamide (GelMA)) with tunable mechanical, degradation and biological properties is used to engineer the epidermis for skin tissue engineering applications. The results reveal that the mechanical and degradation properties of the developed hydrogels can be readily modified by varying the hydrogel concentration, with elastic and compressive moduli tuned from a few kPa to a few hundred kPa and the degradation times varied from a few days to several months. Additionally, hydrogels of all concentrations displayed excellent cell viability (>90%) with increasing cell adhesion and proliferation with increase in hydrogel concentrations. Furthermore, the hydrogels are found to support keratinocyte growth, differentiation and stratification into a reconstructed multi-layered epidermis with adequate barrier functions. The robust and tuneable properties of GelMA hydrogels have suggested that the keratinocyte laden hydrogels can be used as epidermal substitutes, wound dressings or substrates to construct various in vitro skin models.
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No full text available
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10. 1002/adhm. 201500168
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Advanced healthcare materials
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3D Printing of Scaffolds for Tissue Regeneration Applications
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The current need for organ and tissue replacement, repair and regeneration for patients is continually growing such that supply is not meeting the high demand primarily due to a paucity of donors as well as biocompatibility issues that lead to immune rejection of the transplant. In an effort to overcome these drawbacks, scientists working in the field of tissue engineering and regenerative medicine have investigated the use of scaffolds as an alternative to transplantation. These scaffolds are designed to mimic the extracellular matrix (ECM) by providing structural support as well as promoting attachment, proliferation, and differentiation with the ultimate goal of yielding functional tissues or organs. Initial attempts at developing scaffolds were problematic and subsequently inspired a growing interest in 3D printing as a mode for generating scaffolds. Utilizing three-dimensional printing (3DP) technologies, ECM-like scaffolds can be produced with a high degree of complexity and precision, where fine details can be included at a micron level. In this review, we discuss the criteria for printing viable and functional scaffolds, scaffolding materials, and 3DP technologies used to print scaffolds for tissue engineering. A hybrid approach, employing both natural and synthetic materials, as well as multiple printing processes may be the key to yielding an ECM-like scaffold with high mechanical strength, porosity, interconnectivity, biocompatibility, biodegradability, and high processability. Creating such biofunctional scaffolds could potentially help to meet the demand by patients for tissues and organs without having to wait or rely on donors for transplantation.
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No full text available
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10. 1002/adhm. 201500197
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Advanced Healthcare Materials
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Cell Invasion in Collagen Scaffold Architectures Characterized by Percolation Theory
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No abstract available
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Collagen scaffolds are biological templates for the regeneration of damaged or diseased tissues. Since cell invasion into these porous scaffolds is vital for healthy tissue regeneration, the characteristics that influence cellular response must be accounted for in scaffold design. For instance, characterization of mean pore size has revealed that cell migration is highly influenced by scaffold structure. [ 1, 2 ] However, in order to invade at all, cells require a pathway of connected pores: a characteristic that is not described by mean pore size. The extent to which such pathways are present is termed the interconnectivity, and this term may be used to include the number, size, and shape of these pathways. Interconnectivity also has acknowledged importance for nutrient supply and the formation of vascularized tissue deep within the scaffold. [ 3, 4 ] However, there has so far been very little emphasis on its characterization. This gap in understanding between structure and function is a considerable limitation for efficient scaffold design. The difficulty with scaffolds fabricated from natural materials is that they are intrinsically variable; therefore their structural features can be difficult to characterize. Such scaffolds are frequently fabricated by a freeze-drying technique, in which the pore structure is defined by the solidification of ice crystals from an aqueous slurry of polymer fibers. Whereas a thorough understanding of the relationship between freeze-drying conditions and resulting pore size has been developed, [ 5 ] no such understanding exists for interconnectivity. Although the pore space will be predominantly interconnected, due to the characteristic interlocking of the ice crystals, [ 6 ] the difficulty is in assessing the extent of this interconnectivity, especially with relevance to cell invasion. To some extent, interconnectivity can be estimated by a visual assessment from microscopy, and in these cases, greater interconnectivity has been linked to improved cell distribution. [ 7, 8 ] For more rigorous structural characterization, 3D tomographic representations of the scaffolds of interest are often used, such as those from X-ray microcomputed tomography (Micro-CT). One approach is to measure the fraction of pore space accessible from the scaffold exterior. [ 9, 10 ] However, problems exist in scale-up of these measured values from Micro-CT to results that are meaningful at the scale of a bulk sample. [ 11 ] Percolation theory, which deals with the mathematical treatment of transport properties in porous solids, is a recognized solution to the problem of Micro-CT scalability. [ 12 ] It has not yet, however, been implemented for the study of cell accessibility in tissue engineering scaffolds. Whereas existing Micro-CT characterization methods focus on thorough parameterization of individual pores and fenestrations, [ 13, 14 ] with this approach there is little emphasis placed on the spatial distribution of the fenestrations, and whether or not they provide continuous pathways suitable for cell invasion. The ideal method for interconnectivity assessment should include scale-independent characterization of the transport pathways through the structure, and the tools necessary for this approach may be found in percolation theory. In this communication, we demonstrate the use of percolation theory to investigate scaffold interconnectivity in terms of a characteristic feature size for cell transport. We introduce a scale-invariant parameter, termed the “percolation diameter, ” to describe the characteristics of the transport pathways encountered by an invading object. In combination with measurement of pore size, we demonstrate the relevance of the percolation diameter for predicting the extent of cell invasion. So named after the work of Saxton, [ 15 ] the percolation diameter is the size of the largest spherical object able to travel through an infinitely large scaffold. This may be considered a critical value in terms of interconnecting pathways: scaffolds with a certain percolation diameter will impede the transport of any object larger than this diameter. The methodology for its calculation, illustrated in Figure 1, is based on successive measurements of L and d, where d is the diameter of the largest sphere able to travel a linear distance L through the pore space. Using a scaling relationship from percolation theory, these measurements may then be extrapolated to find the value of d as L approaches infinity: this value d c is termed the percolation diameter. Figure 1 Pictorial representation of the methodology used for calculation of pore size and percolation diameter. Pore size was calculated by ellipse fit to z-slices sampled from the Micro-CT dataset. Percolation diameter was calculated by measuring the maximum accessible z -distance, L, to invading objects of varying diameter, d, and extrapolating to infinite scaffold sizes. Full details are given in the Experimental Section. To demonstrate the power of this approach, a series of freeze-dried collagen scaffolds with measurable differences in structure was required, for correlation to observed biological response. Several of the variables in the freeze-drying process were investigated, to assess their potential for producing differences in scaffold interconnectivity. One of the most promising was the choice of suspension medium; a variable that has previously been shown to produce dramatic differences in scaffold architecture. [ 16 ] We chose to compare the structures obtained from two common variants of suspension medium: 0. 05 m acetic acid and 0. 001 m hydrochloric acid (HCl). Acetic acid is a good solvent for collagen and therefore interacts strongly with the suspended fibrils. [ 17 ] This leads to the formation of discrete pore walls, by collagen rejection at the ice crystallization interface. Conversely, the use of HCl at low concentrations is known to change the morphology of the solidifying ice. [ 18 ] We hypothesized that these effects may influence the interconnectivity of the resulting scaffolds. Initial microstructural examination under scanning electron microscopy (SEM) revealed clear differences in pore structure according to the choice of acidic suspension medium. The SEM images in Figure 2 reveal the contrast between the smooth, planar pore walls obtained with acetic acid, and the much more fibrous structure obtained with HCl. This effect can be attributed to the limited interaction between HCl and the collagen fibers, resulting in fiber trapping within the solidifying ice. [ 18 ] Micro-CT analysis revealed that on the change from acetic acid to HCl, the percolation diameter decreased from 72 ± 5 to 32 ± 2 μm. The range of scaffold architectures could be extended even further by the control of additional variables, such as the suspension cooling rate: a recognized factor in determining both pore size and morphology. [ 19, 20 ] By simultaneously adjusting such variables, it was found to be possible to produce a range of scaffolds with percolation diameters between 32 and 100 μm. The pore size range of these scaffolds, 52–70 μm, was comparatively small. Importantly, statistical analysis revealed that significant differences in percolation diameter were achievable in scaffolds with no significant difference in pore size (see the Supporting Information for full data). The initial part of this study therefore showed, for the first time, that pore size and interconnectivity can be independently controlled in freeze-dried collagen scaffolds. Figure 2 Scanning electron micrographs (left, scale bar 50 μm) and Micro-CT visualization of accessible pore space to an invading object of diameter d (right), for scaffolds fabricated with a) 0. 05 m acetic acid and b) 0. 001 m HCl. The emboldened/red-highlighted Micro-CT volumes (1 mm 3 ) represent the case where the invading object is larger than the percolation diameter, d c. The aim of the second part of the study was to assess the invasion potential of connective tissue cells in response to changes in percolation diameter. By controlling the freeze-drying conditions as discussed above, five key scaffolds were produced, with constant pore size but varying percolation diameter. Primary human fibroblasts were chosen for the cell invasion tests, as these were considered the most relevant for the intended end application of connective tissue regeneration. Three days after fibroblast seeding onto the scaffold surfaces, the scaffolds were harvested and fluorescently stained to reveal the actin cytoskeleton of the fibroblasts. Cross-sections were then taken in order to test the hypothesis that percolation diameter would have a measurable influence on cell invasion. Representative cross-sections of each scaffold after three days of culture may be seen in Figure 3. It can be seen from these images that the extent of cell invasion varies widely between the different architectures. Most notably, there appeared to be limited invasion in the two scaffolds of lowest percolation diameter. To confirm this result, we measured the intensity of actin fluorescence, I, as a function of distance from the scaffold surface, Z, as illustrated in Figure 3 b for the two scaffolds previously displayed in Figure 2. A shallower decline in fluorescent intensity corresponds to a more even distribution of actin across the scaffold cross-section, and therefore more efficient cell invasion. Measured intensity values were normalized to total measured intensity, such that the integral of each intensity profile was kept constant. It can be seen from Figure 3 b that percolation diameter affected not only the maximum invasion distance achieved by the cells but also the proportion of cells that remained close to the scaffold surface. To compare this invasion efficiency quantitatively across all tested scaffolds, we used these intensity plots to measure the median cell position for each scaffold. This was plotted as a function of percolation diameter, as shown in Figure 3 c. This plot highlights an interesting relationship between percolation diameter and cell invasion efficiency. For scaffolds with percolation diameter greater than 40 μm, cell invasion efficiency remains roughly constant, with only a gradual possible increase in median cell position with percolation diameter. However, median cell position shows a sharp decrease for the scaffolds with percolation diameter below 40 μm. It seems, therefore, that there may be a critical interconnectivity threshold for persistent, directed cell invasion. Figure 3 Cell invasion results after three days culture, shown a) in cross-section for scaffolds of successively increasing percolation diameters (scale bar 1 mm), b) as fluorescent intensity profiles (for the scaffolds shown in Figure 2, SEM scale bar 50 μm as before), and c) as a plot of median cell position against percolation diameter. By isolating the effect of interconnectivity from that of pore size, we have therefore demonstrated that characterization of only one of these parameters is not sufficient for reliable prediction of cell invasion. This result has profound implications for the design of scaffolds for soft tissue engineering. It implies that optimization of pore size is an incomplete approach to the enhancement and control of cell invasion, and that an understanding of interconnectivity is required to ensure efficient cell movement into a structure. In this study, percolation diameters above 40 μm were required to ensure scaffold accessibility to connective tissue cells. It is interesting to note that in microfluidic channels, both mouse fibroblasts and human mesenchymal stromal cells show a step change in migration rate at channel widths of 40 μm. [ 21 ] Similarly, obstruction of tumor-derived (HT1080) cell migration through two-photon polymerized structures has been observed when the polymer walls making up the structure are more closely spaced than 50 μm. [ 22 ] In each case, a change in cell morphology was observed as the size of the obstructions approached the approximate length scale of a cell (10–30 μm). Importantly, whereas these studies used synthetic polymers to allow microfabrication of scaffolds with discrete feature sizes, we have implemented the percolation diameter methodology to demonstrate that similar structure–function correlations can be drawn in natural polymer scaffolds. Percolation diameter is readily assessed using Micro-CT analysis, and thanks to its scale-invariance from percolation theory, it is relevant for the study of cell invasion in bulk samples. This novel approach will enable future investigation of the relative importance of pore size and interconnectivity, as well as the cell-type dependence of the observed critical interconnectivity threshold for cell invasion. Analysis of the largest object that may traverse the pore space has also been previously investigated using 3D confocal microscopy of collagen networks, although not in a scale-independent manner. [ 23 ] This does, however, signify that such a methodology may be implemented using datasets from a range of imaging techniques. It should be noted that for this study, percolation diameter was measured from scaffolds in the dry state. Further investigation into the effect of scaffold hydration will be an interesting extension, crucial for understanding the response of cells in culture. In addition, complementary techniques such as permeability measurement are currently being investigated in our labs, to enable comparison of fluid flow characteristics with cell invasion potential. We suggest that the percolation diameter approach could provide the means to quantify the necessary interconnectivity threshold required for soft tissue regeneration. Although interconnect sizes above 100 μm are recognized as necessary for bone tissue regeneration, [ 24 ] this is not necessarily a transferable principle to the very different application of soft tissue engineering, which often requires pore structures on a smaller scale. [ 25 ] In summary, we have introduced a novel method for the interconnectivity characterization of collagen scaffolds, and using this method we have demonstrated the independence of interconnectivity and pore size. Using percolation diameter as a parameter for the description of scaffold interconnectivity, we have demonstrated its relevance in determining the extent of fibroblast invasion. This methodology has the potential to enable quantification of the necessary scaffold characteristics for cell invasion and soft tissue regeneration, and to help develop a more thorough understanding of the relationship between scaffold architecture and biological function. By optimization of pore size and percolation diameter, in combination with other properties such as contact guidance by collagen alignment, it is hoped that this approach will open up new strategies for inducing controllable, directed cell invasion. Experimental Section Scaffold Fabrication : Insoluble fibrillar type I collagen from bovine Achilles tendon (Sigma-Aldrich, UK) was hydrated overnight at 1% (w/v) in either 0. 05 m acetic acid (Alfa-Aesar, UK) or 0. 001 m hydrochloric acid (Sigma-Aldrich). After homogenization and centrifugation for air bubble removal, the resulting collagen suspension was poured into stainless steel molds, filling height < 1 cm. The molds were placed into a VirTis AdVantage benchtop freeze-drier (Biopharma Process Systems, UK), which was either precooled to −35 °C or ramped to −35 °C from room temperature at 1. 2 °C min −1. An additional scaffold was created in a silicone mold, filling height 2 cm, by quenching the acetic acid suspension to −20 °C. After complete freezing, a pressure of 80 mTorr and a temperature of 0 °C were maintained for ice sublimation. The resulting scaffolds were chemically cross-linked using 1-ethyl-3-(3-dimethylaminopropyl) carbodiimide hydrochloride (EDC, Sigma-Aldrich) and N -hydroxysuccinimide (NHS, Sigma-Aldrich), with 95% ethanol as solvent. EDC and NHS were used in the molar ratio 5:2:1 relative to the collagen carboxylic acid groups (EDC:NHS:COOH). Scaffolds were immersed in the cross-linking solution for two hours, before thorough washing with distilled water (5 × 5 min), and drying using the same freeze-drying cycle as before. Scaffold Image Acquisition and Pore Size Measurement : For qualitative SEM analysis, scaffolds were sectioned in the plane containing the freezing direction (the x – z plane) using a scalpel, and sputter-coated with gold/platinum. A JEOL JSM-820 SEM was used for image acquisition, in secondary electron mode at 10 kV. For quantitative Micro-CT analysis a Skyscan 1072 system (Bruker, BE) was used to image scaffold samples cut with a 5 mm biopsy punch. Projection images were taken at 25 kV and 138 μA, with 0. 23° rotation steps and 7. 5 s image acquisition time, averaged over four frames. Magnification was set at 75x, pixel size 3. 74 μm. Projections were processed into 3D datasets using the Skyscan reconstruction software NRecon, before binarization using the Trainable Segmentation plugin within the ImageJ software distribution FIJI. Image noise was reduced using individual z-slice despeckle, followed by a 2 × 2 × 2 median filter in 3D. Z-slices were sampled from the dataset at 50 μm spacing and mean pore size over 20 slices was calculated using FIJI: after removal of outliers larger than 2 pixels, a watershed algorithm was applied to the dataset, to allow ellipse fit to each pore. Pore size refers to the mean diameter of the circle of equivalent area to these best-fit ellipses. Percolation Calculations : The median-filtered scaffold dataset was imported into the Skyscan analysis software CTAn. A cubic region of interest (ROI) was defined such that only one face of the cube, an x – y face, was accessible to invasion. Face dimensions were set at 1 mm × 1 mm (for visualization as in Figure 2 ) or at 2 mm × 2 mm (for numerical analysis). The CTAn function “ROI Shrinkwrap” was then used to identify the volume accessible to a virtual object. The diameter of this object, d, was controlled, and the corresponding length of the accessible pore volume in the z -direction, L, was measured. These measurements were then plotted according to the following relationship from percolation theory:[ 15 ] where v is a percolation constant with value 0. 88 for 3D systems. [ 26 ] Values of d were plotted as a function of to allow calculation of the intercept: the percolation diameter, d c. Cell Culture : Human periodontal ligament fibroblasts (Lonza, CH) were cultured in high glucose Dulbecco's Modified Eagle Medium (LifeTechnologies, CH) with 5% fetal bovine serum and 1% penicillin/streptomycin. Trypsin-EDTA was used to detach the subconfluent fibroblasts, which were seeded at passage five. Scaffold samples approximately 10 mm × 10 mm × 2 mm were sterilized in 70% ethanol, before washing twice in phosphate buffered saline (PBS, LifeTechnologies) and subsequent prewetting in medium. Excess medium was aspirated from the scaffolds before seeding in triplicate onto the 10 mm × 10 mm face, at a concentration of 64 000 cells in 50 μL medium per scaffold. Extra medium was added after one hour at room temperature. Culture conditions were maintained at 37 °C and 5% CO 2 for three days, with one medium change. At day three, medium was removed and the scaffolds were washed in PBS, before fixing with 10% formalin (Sigma-Aldrich). Staining and Microscopy : Once washed in PBS, scaffolds were immersed in 0. 1% Triton X-100/PBS (Sigma-Aldrich) for 10 min and then washed in PBS before cytoskeletal actin staining with Alexa Fluor 488 Phalloidin (MolecularProbes, CH) at 2. 5 μL/200 μL in 1% bovine serum albumin/PBS (BSA, Sigma-Aldrich). Scaffolds were then embedded in 15% gelatin/PBS (BioGel, CH), and the solidified gelatin blocks were fixed with 10% formalin. A Leica VT1000 S Vibratome was used to section these blocks at a thickness of 200 μm, to reveal the scaffold cross-section. A Yokogawa CV1000 Cell Voyager confocal microscope was used to record the maximum fluorescent intensity over 11 z -slices, spacing 20 μm, for each scaffold cross-section. For each scaffold condition, two biological replicates were chosen for analysis, deliberately selected such that local collagen wall orientation was kept constant between scaffold conditions. Three sections were taken from each of these replicates, giving a total of six images for study per scaffold condition. Fluorescent intensity profiles were averaged over a width of 4 mm (300 pixels) and background intensity values (measured from an empty area of the image) were subtracted. Measured intensity values I were normalized to the total summed intensity over the profile, ∑ I, and the median cell position was calculated by finding the distance at which the cumulative intensity equaled half the total summed intensity. Statistics : For pore size and percolation diameter, the mean of three measurements was calculated, along with standard error of the mean. For median cell position, the mean and standard error of six measurements was calculated. Statistical significance was tested using one-way ANOVA and Games-Howell analysis was used for pairwise comparisons (significance level p < 0. 05).
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10. 1002/adhm. 201500236
| 2,015
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Advanced healthcare materials
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ECM-Based Biohybrid Materials for Engineering Compliant, Matrix-Dense Tissues
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An ideal tissue engineering scaffold should not only promote, but take an active role in, constructive remodeling and formation of site appropriate tissue. ECM-derived proteins provide unmatched cellular recognition, and therefore influence cellular response towards predicted remodeling behaviors. Materials built with only these proteins, however, can degrade rapidly or begin too weak to substitute for compliant, matrix-dense tissues. The focus of this review is on biohybrid materials that incorporate polymer components with ECM-derived proteins, to produce a substrate with desired mechanical and degradation properties, as well as actively guide tissue remodeling. Materials are described through four fabrication methods: (1) polymer and ECM-protein fibers woven together, (2) polymer and ECM proteins combined in a bilayer, (3) cell-built ECM on polymer scaffold, and (4) ECM proteins and polymers combined in a single hydrogel. Scaffolds from each fabrication method can achieve characteristics suitable for different types of tissue. In vivo testing has shown progressive remodeling in injury models, and suggests ECM-based biohybrid materials promote a prohealing immune response over single component alternatives. The prohealing immune response is associated with lasting success and long term host maintenance of the implant.
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No full text available
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10. 1002/adhm. 201500304
| 2,015
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Advanced healthcare materials
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Temporally Tunable, Enzymatically-responsive Delivery of Pro-angiogenic Peptides from Poly(ethylene glycol) Hydrogels
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Pro-angiogenic drugs hold great potential to promote reperfusion of ischemic tissues and in tissue engineering applications, but efficacy is limited by poor targeting and short half-lives. Methods to control release duration or provide enzymatically-responsive drug delivery have independently improved drug efficacy. However, no material has been developed to temporally control the rate of enzymatically-responsive drug release. To address this void, hydrogels were developed to provide sustained, tunable release of Qk, a pro-angiogenic peptide mimic of vascular endothelial growth factor, via tissue-specific enzymatic activity. After confirmation that sustained delivery of Qk is necessary for pro-angiogenic effects, a variety of previously-identified matrix metalloproteinase (MMP)-degradable linkers were used to tether Qk to hydrogels. Of these, three (IPES↓LRAG, GPQG↓IWGQ, and VPLS↓LYSG) showed MMP-responsive peptide release. These linkers provided tunable Qk release kinetics, with rates ranging from 1. 64 to 19. 9 × 10 −3 hours −1 in vitro and 4. 82 to 8. 94 × 10 −3 hours −1 in vivo. While Qk was confirmed to be bioactive as released, hydrogels releasing Qk failed to induce significant vascularization in vivo after one week, likely due to non-enzymatically degradable hydrogels employed. While Qk was the focus of this study, the approach could easily be adapted to control the delivery of a variety of therapeutic molecules.
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No full text available
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10. 1002/adhm. 201500411
| 2,016
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Advanced healthcare materials
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Recombinant Resilin-based Bioelastomers for Regenerative Medicine Applications
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The superior elasticity, excellent resilience at high-frequency, and hydrophilic capacity of natural resilin have motivated investigations of recombinant resilin-based biomaterials as a new class of bio-elastomers in the engineering of mechanically active tissues. Accordingly, we report here the comprehensive characterization of modular resilin-like polypeptide (RLP) hydrogels and introduce their suitability as a novel biomaterial for in vivo applications. Oscillatory rheology confirmed that a full suite of the RLPs can be rapidly crosslinked upon addition of the tris(hydroxymethyl phosphine) (THP) cross-linker, achieving similar in situ shear storage moduli (20k±3. 5Pa) across various material compositions. Uniaxial stress relaxation tensile testing of hydrated RLP hydrogels under cyclic loading and unloading showed negligible stress reduction and hysteresis, superior reversible extensibility, and high resilience with Young’s moduli of 30±7. 4kPa. RLP hydrogels containing MMP-sensitive domains are susceptible to enzymatic degradation by MMP-1. Cell culture studies revealed that RLP-based hydrogels supported the attachment and spreading (2D) of human mesenchymal stem cells (hMSCs) and did not activate cultured macrophages. Subcutaneous transplantation of RLP hydrogels in a rat model, which to our knowledge is the first such reported in vivo analysis of RLP-based hydrogels, illustrated that these materials do not elicit a significant inflammatory response, suggesting their potential as materials for tissue engineering applications with targets of mechanically demanding tissues such as vocal fold and cardiovascular tissues.
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No full text available
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10. 1002/adhm. 201500517
| 2,016
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Advanced healthcare materials
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Textile Technologies and Tissue Engineering: A Path Towards Organ Weaving
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Textile technologies have recently attracted great attention as potential biofabrication tools for engineering tissue constructs. Using current textile technologies, fibrous structures can be designed and engineered to attain the required properties that are demanded by different tissue engineering applications. Several key parameters such as physiochemical characteristics of fibers, pore size and mechanical properties of the fabrics play important role in the effective use of textile technologies in tissue engineering. This review summarizes the current advances in the manufacturing of biofunctional fibers. Different textile methods such as knitting, weaving, and braiding are discussed and their current applications in tissue engineering are highlighted.
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No full text available
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10. 1002/adhm. 201500531
| 2,015
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Advanced healthcare materials
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Hierarchical Nanofibrous Microspheres with Controlled Growth Factor Delivery for Bone Regeneration
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The integration of controlled growth factor delivery and biomimetic architecture into a microsphere is a challenging but attractive strategy for developing new injectable biomaterials for tissue engineering. In this work, we developed a unique hierarchical nanosphere-encapsulated-in-microsphere scaffolding system for bone tissue regeneration. First, we synthesized heparin-conjugated gelatin (HG) that provides binding domains for bone morphogenetic protein 2 (BMP2) to stabilize this growth factor, protect it from denaturation and proteolytic degradation, and subsequently prolong its sustained release. Next, we developed a unique approach that includes a water-in-oil-in-oil (W/O/O) double emulsion process and a thermally induced phase separation to encapsulate BMP2-binding HG nanospheres into nanofibrous microspheres. The nanofibrous microsphere was self-assembled from synthetic nanofibers, and had superior surface area, high porosity, low density, and was an excellent carrier to support cell adhesion and tissue in-growth. BMP2 in the hierarchical microsphere was released in a multiple-controlled manner (by the binding with heparin and encapsulation of the nanosphere and microsphere) and retained its high bioactivity. An in vivo calvarial defect model confirmed that this unique hierarchical microsphere was an excellent osteoinductive scaffold for enhanced bone regeneration. By choosing different growth factors, this hierarchical microsphere system can be easily applied to other types of tissue regeneration. Our work expands the ability to develop new injectable biomaterials for advanced regenerative therapies.
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No full text available
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