A Peptide Modified Hydrogel Therapy for Acute Myocardial Infarction

by

Lewis A. Reis

A thesis submitted in conformity with the requirements for the degree of Doctor o Philosophy The Institute for Biomaterials and Biomedical Engineering University of Toronto

© Copyright by Lewis Reis 2015

A Peptide Modified Hydrogel Therapy for Acute Myocardial Infarction Lewis A. Reis Doctor of Philosophy The Institute of Biomaterials and Biomedical Engineering University of Toronto 2015

Abstract Myocardial infarction (MI) results in the death of cardiomyocytes (CM) followed by scar formation and pathological remodeling of the heart. We postulate that immobilization of the pro-survival angiopoietin-1-derived peptide, QHREDGS, to a chitosan-collagen hydrogel could produce a clinically translatable thermo-responsive hydrogel to attenuate post-MI cardiac remodeling. Conjugation of QHREDGS peptide to chitosan does not interfere with gelation, structure, or mechanical properties of chitosan-collagen hydrogel blends. The storage modulus of 2.5mg/mL 1:1 mass:mass chitosan:collagen was measured to be 54.9±9.1 Pa and loss modulus of 6.1±0.9 Pa. Dose response of the QHREDGS peptide was assessed and it was found that CMs encapsulated in High peptide gel (651±8 nmol peptide/mL-gel) showed improved morphology, viability, and metabolic activity in comparison to the Low peptide (100±30 nmol peptide/mL-gel) and Control (No Peptide) groups. Construct (CMs in hydrogel) functional properties were not significantly different between the groups, however success rate of obtaining a beating construct was improved in the hydrogel with the High amount of QHREDGS

ii

peptide immobilized compared to the Low and Control groups. An adult Lewis rat left anterior descending coronary artery ligation procedure to mimic acute MI model was used to assess in vivo hydrogel performance. QHREDGS-conjugated hydrogel (QHG213H), Control gel, or PBS was injected into 3 locations in the MI zone. By in vivo tracking and chitosan staining, the hydrogel was demonstrated to remain in situ for 2 weeks and cleared in about 3 weeks. By echocardiography and pressure-volume analysis, the QHG213H hydrogel significantly improved cardiac function compared to the controls. Scar thickness and scar area fraction were also significantly improved with QHG213H gel injection compared to the controls. Mechanistically, there were significantly more cardiomyocytes, determined by cardiac troponin-T staining, in the MI zone of the QHG213H hydrogel group. No significant difference in inflammatory response between groups was observed as determined by gene regulation and cytokine analysis of excised heart sections 24 hours after treatment. The interaction of CMs with QHREDGS was found to be mediated by β1-integrins and to increase expression of the pro-survival effector MAPK.

iii

Acknowledgements First, I must thank my supervisor and mentor Dr. Milica Radisic for initially giving me the opportunity to work in her lab, and for the years of support and guidance. I hope the successful completion of this thesis and the project which we have worked on over the last number of years is testament to that. My committee members, Drs. Julie Audet and Molly Shoichet have scrutinized this work many times and given invaluable advice on making it the best it can be, for which I am grateful. The help, advice, and collaborative work from Dr. Ren-Ke Li and his research group, specifically Dr. Jun Wu, have made a large part of this work possible and I must extend my gratitude to them. Dr. Abdul Momen must also be acknowledged for his help in performing much of the early in vivo studies. All the past and present member of the Laboratory for Functional Tissue engineering have given immeasurable support and help over the years and all have contributed to making this project what it is. I must acknowledge Loraine Chiu who’s collaborative efforts undoubtedly shortened my time here by many years, as well as the efforts of Nicole Feric, Carol Laschinger, Jason Miklas, Kent Hyun, and Yan Liang for their contributions to this project and ensuring the work was published. Friends at U of T, Massey College, and those smart enough to not have gone into graduate school, you all know who you are and I thank you for being there throughout. To my family, none of this would have been possible without your love and support. Mom, if it weren’t for you I never would have attempted this challenge nor seen it through to completion. Emily, you came into my life at the beginning of this, changed my life forever, and have stood by me throughout. Thank you and I love you. Finally, I owe a debt of gratitude to all the animals who gave their lives in the name of science and contributed to my work. Know that your sacrifice was not in vain.

iv

Declaration of Co-Authorship The original scientific content of this thesis is comprised of two published, peer reviewed articles in internationally recognized journals. These articles were primarily the work of Lewis Reis. Introduction and literature review are from three published review papers, the transcribed writing being that of Lewis Reis. The contributions of co-authors are stated in the thesis, in conformity with the requirements for the degree of Doctor of Philosophy.

v

Permissions Copyright © 2015 Wolters Kluwer Health. Contents of this thesis have been published in Circulation: Heart Failure: Reis L.A., Chiu L.L.Y., Wu J., Feric N., Momen A., Li R.K., & Radisic M. (2015). Hydrogels with integrin-binding angiopoietin-1-derived peptide, QHREDGS, for treatment of acute myocardial infarction. Circulation: Heart Failure. DOI: 10.1161/CIRCHEARTFAILURE.114.001881. Reuse with permission from Wolters Kluwer Health. A link to the published paper can be found at: http://circheartfailure.ahajournals.org/content/early/2015/01/28/CIRCHEARTFAILURE.114.001 881.abstract Copyright © 2014 John Wiley & Sons, Ltd. Contents of this thesis have been published in the Journal of Tissue Engineering and Regenerative Medicine: Reis L.A., Chiu L.L.Y., Feric N., Fu L., & Radisic M. (2014). Biomaterials in myocardial tissue engineering. Journal of Tissue Engineering & Regenerative Medicine. DOI:10.1002/term.1944. Reuse with permission from John Wiley & Sons, Ltd. A link to the published paper can be found at: http://onlinelibrary.wiley.com/doi/10.1002/term.1944/abstract Copyright © 2012 Frontiers in Bioscience. Contents of this thesis have been published in Frontiers in Bioscience: Chiu LLY, Iyer RK, Reis LA, et al. (2012). Cardiac tissue engineering: current state and perspectives. Frontiers in Bioscience. 17:1533-1550. Reuse with permission from Frontiers in Bioscience. DOI: 10.2741/4002. A link to the published paper can be found at: www.bioscience.org/2012/v17/af/4002/list.htm Copyright © 2012 Elsevier. Contents of this thesis have been published in Acta Biomaterialia. Reis LA et al. (2012). A peptide-modified chitosan-collagen hydrogel for cardiac cell culture and delivery. Acta Biomaterilia. 8(3): 1022-36. Reuse with permission from Elsevier. DOI: 10.1016/j.actbio.2011.11.030. A link to the published paper can be found at: www.sciencedirect.com/science/article/pii/S1742706111005332 Copyright © 2011 Elsevier. Contents of this thesis have been published in Current Opinion in Biotechnology: Iyer RK, Chiu LLY, Reis LA, & Radisic M. (2011). Engineered cardiac tissues. vi

Current Opinion in Biotechnology. 22(5):706-714. Reuse with permission from Elsevier. DOI: 10.1016/j.copbio.2011.04.004. A link to the published paper can be found at: www.sciencedirect.com/science/article/pii/S0958166911000668 Copyright © 2014 Elsevier. Contents of this thesis have been published in Cardiac Regeneration & Repair (Vol. II, Ch. 3). Reis LA et al. (January 2014). Injectable biomaterials for cardiac repair. In Li R.K. & Weisel R.D. (EDs), Cardiac Regeneration & Repair (Vol. II, Ch. 3). Cambridge: Woodhead Publishing. Reuse with permission from Elsevier. DOI: 10.1533/9780857096715.1.49. A link to the book chapter can be found at: http://www.sciencedirect.com/science/article/pii/B9780857096593500037 Copyright © 2011 Springer. Contents of this thesis have been published in Biomaterials for Tissue Engineering Applications: A review of past & future trends: Odedra D, Chiu LLY, Reis LA, et al. (2011). Cardiac Tissue Engineering. In J. A. Burdick & R. L. Mauck (EDs.), Biomaterials for Tissue Engineering Applications: A review of past & future trends (1st ed., p. 562). Vienna: Springer Vienna. Reuse with permission from Springer. DOI: 10.1007/978-37091-0385-2_15. A link to the book chapter can be found at: http://link.springer.com/chapter/10.1007/978-3-7091-0385-2_15/fulltext.html

vii

Abstracts of Published Works Appearing in the Thesis A peptide-modified chitosan-collagen hydrogel for cardiac cell culture and delivery Reis, Lewis A; Chiu, Loraine L Y; Liang, Yan; Hyunh, Kent; Momen, Abdul; & Radisic, Milica. Acta Biomaterialia. (2012). 8(6): 1022-1036. Myocardial infarction (MI) results in the death of cardiomyocytes (CM) followed by scar formation and pathological remodeling of the heart. We propose that chitosan conjugated with the angiopoietin-1 derived peptide, QHREDGS, and mixed with collagen I forms a thermoresponsive hydrogel better suited for the survival and maturation of transplanted cardiomyocytes in vitro compared to collagen and chitosan-collagen hydrogels alone. Conjugation of QHREDGS peptide to chitosan does not interfere with the gelation, structure or mechanical properties of the hydrogel blends. The storage modulus of 2.5 mg ml(-1) 1:1 mass:mass (m:m) chitosan-collagen was measured to be 54.9 ± 9.1 Pa, and the loss modulus 6.1±0.9 Pa. The dose-response of the QHREDGS peptide was assessed and it was found that CMs encapsulated in High-peptide gel (651 ± 8 nmol peptide ml-gel(-1)) showed improved morphology, viability and metabolic activity in comparison to the Low-peptide (100 ± 30 nmol peptide ml-gel(-1)) and Control (No Peptide) groups. Construct (CMs in hydrogel) functional properties were not significantly different between the groups; however, the success rate of obtaining a beating construct was improved in the hydrogel with the High amount of QHREDGS peptide immobilized compared to the Low and Control groups. Subcutaneous injection of hydrogel (Control, Low and High) with CMs in the back of Lewis rats illustrated its ability to localize at the site of injection and retain cells, with CM contractile apparati identified after seven days. The hydrogel was also able to successfully localize at the site of injection in a mouse MI model. Contributions: LAR-concept and design, performed all experiments and data analysis, manuscript writing; LLYC- concept and design, assistance in surgical work, preparation of samples for histology, data analysis; YL- image analysis for vasculature; KH- Pico green DNA analysis; AM-performed surgical work; MR-concept, data interpretation, final approval of manuscript.

viii

Hydrogels with integrin-binding angiopoietin-1-derived peptide, QHREDGS, for treatment of acute myocardial infarction Reis, Lewis A; Chiu, Loraine L Y; Wu, Jun; Feric, Nicole; Laschinger, Carol; Momen, Abdul; Li, Ren-Ke; & Radisic, Milica. Circulation Heart Failure. (2015). DOI: 10.1161/CIRCHEARTFAILURE.114.001881. Background: Hydrogels are being actively investigated for direct delivery of cells or bioactive molecules to the heart post-myocardial infarction (MI) to prevent cardiac functional loss. We postulate that immobilization of the pro-survival angiopoietin-1-derived peptide, QHREDGS, to a chitosan-collagen hydrogel could produce a clinically translatable thermoresponsive hydrogel to attenuate post-MI cardiac remodeling. Methods & Results: In a rat MI model, QHREDGS-conjugated hydrogel (QHG213H), Control gel, or PBS was injected into the peri-infarct/MI zone. By in vivo tracking and chitosan staining, the hydrogel was demonstrated to remain in situ for 2 weeks and was cleared in ~3 weeks. By echocardiography and pressure-volume analysis, the QHG213H hydrogel significantly improved cardiac function compared to the controls. Scar thickness and scar area fraction were also significantly improved with QHG213H gel injection compared to the controls. There were significantly more cardiomyocytes (CMs), determined by cardiac troponin-T staining, in the MI zone of the QHG213H hydrogel group; and hydrogel injection did not induce a significant inflammatory response as assessed by PCR and an inflammatory cytokine assay. The interaction of CMs and cardiac fibroblasts with QHREDGS was found to be mediated by β1integrins. Conclusions: We demonstrated for the first time that the QHG213H hydrogel can be injected in the beating heart where it remains localized for a clinically effective period. Moreover, the QHG213H hydrogel induced significant cardiac functional and morphological improvements post-MI relative to the controls. Contributions: LAR-concept and design, performed all experiments and data analysis, manuscript writing; LLYC- concept and design, assistance in surgical work, preparation of samples for histology, image & data analysis; JW- design, performed surgical work, preparation of samples for histology, data analysis and interpretation, manuscript writing; NF- Attachment assay,

ix

manuscript writing; AM-performed surgical work; RKL & MR-concept, data interpretation, final approval of manuscript.

x

Table of Contents Abstract .................................................................................................................................................... ii Acknowledgements................................................................................................................................. iv Declaration of Co-Authorship ................................................................................................................. v Permissions ............................................................................................................................................. vi Abstracts of Published Works Appearing in the Thesis ....................................................................... viii 1.

Introduction ................................................................................................................................... 1 1.1.

Motivation ................................................................................................................................. 1

1.2.

Hypothesis................................................................................................................................. 2

1.3.

Specific aims ............................................................................................................................. 2

2.

3.

Literature Review.......................................................................................................................... 3 2.1.

Cardiovascular Disease, Outcomes, and Treatment Options .................................................... 3

2.2.

Cell sources, stem cells, & differentiation ................................................................................ 5

2.3.

Design criteria for biomaterials in cardiac tissue engineering .................................................. 7

2.3.1.

Biocompatibility................................................................................................................ 7

2.3.2.

Biodegradability ................................................................................................................ 7

2.3.3.

Mechanical Support .......................................................................................................... 9

2.3.4.

Injectability ..................................................................................................................... 11

2.3.5.

Clinically Relevant Thickness......................................................................................... 11

2.3.6.

Application Time ............................................................................................................ 11

2.4.

Tissue engineered cardiac grafts (in vitro engineering) .......................................................... 12

2.5.

Injection (in vivo engineering) ................................................................................................ 14

2.5.1.

Cell only based therapies ................................................................................................ 15

2.5.2.

Hydrogels to promote endogenous repair ....................................................................... 16

2.5.3.

Hydrogels for the delivery of cells for regeneration ....................................................... 21

2.5.4.

Hydrogels for the artificial maintenance of ventricle geometry and repair .................... 26

2.5.5.

Clinical application ......................................................................................................... 29

2.6.

Chitosan & collagen ................................................................................................................ 30

2.7.

QHREDGS peptide ................................................................................................................. 33

2.8.

Peptide modified chitosan-collagen hydrogel for cardiac tissue engineering ......................... 34 Hydrogel development, in vitro characterization & preliminary in vivo models ........................ 36

xi

3.1.

Materials & Methods .............................................................................................................. 36

3.1.1.

3.1.1.1.

Conjugation .............................................................................................................. 36

3.1.1.2.

Assessing conjugation efficiency ............................................................................. 36

3.1.2.

Chitosan-collagen hydrogel ............................................................................................ 37

3.1.2.1.

Hydrogel formulation ............................................................................................... 37

3.1.2.2.

Scanning electron microscopy .................................................................................. 37

3.1.2.3.

Hydrogel degradation ............................................................................................... 38

3.1.2.4.

Rheological assessment of hydrogels ....................................................................... 38

3.1.3.

In vitro cell culture .......................................................................................................... 38

3.1.3.1.

CM isolation ............................................................................................................. 38

3.1.3.2.

CM media ................................................................................................................. 39

3.1.3.3.

CM encapsulation & culture ..................................................................................... 39

3.1.4.

In vivo studies.................................................................................................................. 39

3.1.4.1.

Subcutaneous injection ............................................................................................. 39

3.1.4.2.

Mouse MI model ...................................................................................................... 40

3.1.5.

Construct Characterization .............................................................................................. 40

3.1.5.1.

Gel compaction ......................................................................................................... 40

3.1.5.2.

Live/Dead staining .................................................................................................... 41

3.1.5.3.

Functional testing ..................................................................................................... 41

3.1.5.4.

XTT assay ................................................................................................................. 42

3.1.5.5.

LDH assay ................................................................................................................ 42

3.1.5.6.

PicoGreen DNA Assay ............................................................................................. 43

3.1.5.7.

Histological & immunofluorescent staining ............................................................. 43

3.1.6. 3.2.

Peptide modified chitosan (UP-G113-QHREDGS) ........................................................ 36

Statistical analysis ........................................................................................................... 45

Results & Discussion .............................................................................................................. 45

3.2.1.

Hydrogel composition and characterization.................................................................... 45

3.2.1.1.

Base collagen-chitosan hydrogel .............................................................................. 45

3.2.1.2.

Collagen-chitosan-QHREDGS hydrogel .................................................................. 50

3.2.2.

In vitro studies with cardiomyocytes in collagen-chitosan-QHREDGS hydrogels ........ 52

xii

3.2.2.1.

Distribution, viability, and metabolism of CMs encapsulated in the hydrogels ....... 52

3.2.2.2.

Construct functional & morphological properties .................................................... 57

3.2.3.

3.3. 4.

In vivo studies.................................................................................................................. 59

3.2.3.1.

Subcutaneous injection ............................................................................................. 59

3.2.3.2.

Mouse MI model study ............................................................................................. 64

Summary ................................................................................................................................. 69 Utility of QH-G213-H hydrogel as a treatment for acute MI ..................................................... 70

4.1.

Materials & Methods .............................................................................................................. 70

4.1.1.

Experimental design overview ........................................................................................ 70

4.1.2.

Peptide modified chitosan-collagen hydrogel (QHG213H) ............................................ 71

4.1.3.

Peptide modified polyethylene glycol (PEG-QHREDGS) ............................................. 71

4.1.4.

In vivo studies.................................................................................................................. 72

4.1.4.1.

Assessment of cardiac function (6 Week time point) ............................................... 72

4.1.4.2.

24 hr time point ........................................................................................................ 73

4.1.4.3.

In vivo hydrogel lifespan .......................................................................................... 74

4.1.5.

Immunohistochemistry.................................................................................................... 74

4.1.6.

Quantitative PCR ............................................................................................................ 75

4.1.7.

Western blotting .............................................................................................................. 75

4.1.8.

Image analysis techniques ............................................................................................... 76

4.1.9.

Statistical analysis ........................................................................................................... 76

4.2.

Results ..................................................................................................................................... 76

4.2.1.

In vivo degradation (lifespan).......................................................................................... 76

4.2.2.

Functional Data ............................................................................................................... 77

4.2.3.

Gross morphology and histology .................................................................................... 81

4.2.4.

Cardiomyocyte survival mechanism ............................................................................... 84

4.3.

Discussion ............................................................................................................................... 86

4.4.

Summary ................................................................................................................................. 94

5.

Recommendations for future work ............................................................................................. 96 5.1.

Specific Aim SA-2. i. Show ability to localize in vivo and quantify duration of its in vivo

presence. 96 5.2.

Specific Aim SA-2. iii. Determine mechanism of action........................................................ 96

xiii

6.

Conclusions & Contributions to New Knowledge ...................................................................... 99 6.1.

7.

Publications & Contributions ................................................................................................ 101 References ................................................................................................................................. 105

Appendix A.

Reaction schematics and chemical structures ........................................................... 133

Appendix B.

In vitro construct CM distribution............................................................................. 136

Appendix C.

Utility of QHG213H hydrogel for MI treatment published supplementary data ...... 139

Appendix D.

Unpublished rat MI model supplementary material.................................................. 146

List of Tables Table 3-1: Hydrogel rheological assessment ........................................................................... 48 Table 4-1: Comparative-analysis of reported MI studies......................................................... 94 Table C-1: Genes and primers used in RT-qPCR .................................................................. 139 Table C-2: Description of studies considered for comparative-analysis .............................. 140

List of Figures Figure 2-1: Histological stages of myocardial infarction ........................................................... 4 Figure 2-2: The structures of chitosan and chitin, and their proportions in 85% deacetylated chitosan ......................................................................................................................................... 31 Figure 2-3: Molecular structure of Ang-1 derived peptide QHREDGS .................................. 34 Figure 2-4: QHREDGS peptide modified chitosan (QHG213H) ............................................ 35 Figure 3-1: Hydrogel Characterization .................................................................................... 49 Figure 3-2: Characterization of peptide conjugation and hydrogel morphology ..................... 52 Figure 3-3: Gel encapsulation results in uniform distribution of live cells.............................. 53 Figure 3-4: Metabolic activity and total number of cells encapsulated in 1:1 chitosan:collagen hydrogels ....................................................................................................................................... 56 Figure 3-5: 1:1 Chitosan:collagen hydrogel with immobilized QHREDGS enables cultivation of beating cardiac tissue ................................................................................................................ 58 Figure 3-6: Subcutaneous injection study ................................................................................ 61 Figure 3-7: Immunostaining for different cell populations in subcutaneously injected nodules after 7 days in vivo ........................................................................................................................ 63 xiv

Figure 3-8: Chitosan:collagen gel is suitable for injection into the infarcted heart ................. 65 Figure 3-9: Vascularization and wall thickness in the mouse MI model ................................. 67 Figure 4-1: In vivo experimental design................................................................................... 71 Figure 4-2: In vivo lifespan of Dy-light 800 labelled chitosan-collagen gel............................ 77 Figure 4-3: Cardiac function post MI measured by echocardiography.................................... 79 Figure 4-4: Load dependent cardiac function and LV volumes at 6 weeks post MI ............... 80 Figure 4-5: Load independent pressure-volume analysis at 6 weeks post MI ......................... 81 Figure 4-6: Gross heart morphology at 6 weeks post MI......................................................... 82 Figure 4-7: Vascularization at 6 weeks post MI ...................................................................... 83 Figure 4-8: Apoptosis at 6 weeks post MI ............................................................................... 85 Figure 4-9: Cardiac cell attachment to PEG and PEG-QHREDGS gels ................................. 86 Figure 4-10: Proposed QHREDGS mediated mechanism ....................................................... 92 Figure A-1: Polyelectrolytic complexation of chitosan and collagen .................................... 133 Figure A-2: Modified chitosans used ..................................................................................... 134 Figure A-3: EDC/Sulfo-NHS reaction chemistry for conjugation of biomolecules .............. 134 Figure A-4: QHREDGS conjugation to PEG......................................................................... 135 Figure B-1: In vitro cell leakage and migration ..................................................................... 136 Figure B-2: Quantification of in vitro construct viability ...................................................... 137 Figure B-3: High QHREDGS hydrogel constructs contain elongated cardiomyocytes with visible cross-striations ................................................................................................................. 138 Figure C-1: Quantification of hydrogel lifespan .................................................................... 141 Figure C-2: Histological staining 6 weeks post MI ............................................................... 142 Figure C-3: RT-qPCR of MI and border zone tissues from hearts excised 24 hrs post MI ... 143 Figure C-4: Western blot analysis of border zone tissue from hearts excised 24 hrs post MI144 Figure C-5: Reported cardiac morphological and functional data from studies used in metaanalysis........................................................................................................................................ 145 Figure D-1: Heart morphology & histology 3Wk post MI .................................................... 146 Figure D-2: Further histological staining 3Wks post MI ....................................................... 147 Figure D-3: 3Wk MI quantification measurements ............................................................... 149 Figure D-4: 3Wk MI scar extent assessment ......................................................................... 150 Figure D-5: 6Wk MI quantification measurements (from histology pictures) ...................... 151 xv

Figure D-6: Vascularization within the MI scar zone at 3Wks post infarct........................... 152 Figure D-7: Apoptosis in the MI boundary zone at 3Wk ...................................................... 153 Figure D-8: 3Wk TUNEL/cTnT staining controls ................................................................. 154 Figure D-9: 6Wk TUNEL/cTnT staining raw counts ............................................................ 154 Figure D-10: 6Wk TUNEL/cTnT staining controls ............................................................... 155

xvi

1. Introduction Parts of the text used in this chapter has been published in Circulation Heart Failure (Reis et al. 2015), and is used with permission from Wolters Kluwer Health.

1.1. Motivation Cardiomyocyte (CM) death, myocardial remodeling, and scar tissue formation following myocardial infarction (MI) results in severe cardiac dysfunction and serious health problems. If not treated, and often with treatment, heart failure progresses. Direct transplantation of various cell types and/or bioactive molecules has shown promise, however both strategies have been hampered by low injection site retention and even lower long-term survival of the cells/bioactive molecules (Mayfield et al. 2014; Nelson et al. 2012). To mitigate these problems these treatments are combined with hydrogels to confine the cells/bioactive molecules to the site of injection and promote their long-term survival therein. The acute phase post-MI might be the most appropriate time to utilize a hydrogel-based therapy because therein hydrogel injections can prevent cardiac remodeling, deliver cells to replace the damaged tissue, and/or recruit endogenous stem cells (Menasché 2008). Chitosan and collagen are natural, biodegradable, biocompatible polymers that have been explored for their potential use in the treatment of cardiac dysfunction (Lu et al. 2009; Rask, Dallabrida, et al. 2010; Garbern et al. 2011; Suuronen et al. 2006; Y. Zhang et al. 2008; Wu et al. 2011). Typically, collagen and chitosan, alone or in combination, are cross-linked using exogenous, often toxic, chemical cross-linkers to improve the hydrogel mechanical properties (Lu et al. 2009; Deng et al. 2010; Fujita et al. 2005). We propose that chitosan-collagen composites can gel naturally at physiological temperatures and pH to form mechanically stable hydrogels that are appropriate for in vivo application without the need for exogenous cross-linkers. Furthermore, the collagen-chitosan interaction within the gels resembles the collagen-glycosaminoglycan interaction found in vivo in the extracellular matrix (ECM) (Tan, Krishnaraj, and Desai 2001). Thus, chitosan-collagen may mediate physiological cell-matrix interactions. The functional success of hydrogel-based cardiac and cell therapies can be improved by modifying biomaterials with bioactive molecules because they (cytokines, growth factors, etc.) have the potential to increase transplanted cell survival, reduce resident cell apoptosis, recruit desired regenerative cells, and promote stem/progenitor cell differentiation (Chiu, Iyer, et al. 1

2012; Reis et al. 2014a). One such bioactive molecule is the growth factor angiopoietin-1 (Ang1). In endothelial cells, Ang1 binds the Tie2 receptor (Hayes et al. 1999; I. Kim et al. 2000; Papapetropoulos 2000) but in cells such as neonatal rat cardiomyocytes (NCMs) that lack the Tie2 receptor, Ang1 binds to integrins (Carlson et al. 2001) and activates pro-survival pathways (Dallabrida et al. 2005). We identified the short sequence QHREDGS as the integrin-binding motif of Ang1, and the QHREDGS peptide was found to support CM attachment and survival similar to full-length Ang1 (Rask, Dallabrida, et al. 2010; Rask, Mihic, et al. 2010). It is therefore possible that the QHREDGS peptide could retain/restore cardiac contractile function post-MI by promoting CM survival. Importantly, the QHREDGS peptide is water-soluble, stable, fully-synthetic with a precisely defined composition, and does not require a specific orientation to be functional. We therefore look to incorporate the pro-survival peptide QHREDGS into an optimized chitosan-collagen hydrogel to produce a novel hydrogel-based cardiac regenerative therapy. We chose to immobilize the peptide onto the hydrogel because the effects of QHREDGS were found to be attachment-dependent, and immobilization would ensure sustained localization at the injection site and requires a single low dose.

1.2. Hypothesis Conjugation of Angiopoietin-1 derived peptide QHREDGS to chitosan, injected as a peptidemodified chitosan-collagen gel, will enhance survival of cardiac muscle after myocardial infarction in vivo and attenuate pathological remodeling.

1.3. Specific aims SA-1. Develop a chitosan-collagen hydrogel that supports CM viability and phenotype. Conjugate QHREDGS and demonstrate dose dependent improvement in CM phenotype and function in vitro. SA-2. Assess effect of the developed hydrogel in a rat MI model. i.

Show ability to localize in vivo and quantify duration of its in vivo presence.

ii.

Demonstrate potential to attenuate cardiac remodeling post MI.

iii.

Determine mechanism of action.

2

2. Literature Review Contents of this chapter have been published in (Iyer et al. 2011; Odedra et al. 2011; Chiu, Iyer, et al. 2012; Reis et al. 2012; Reis et al. 2014b; Reis et al. 2015) and are used with permission from respective publishers.

2.1. Cardiovascular Disease, Outcomes, and Treatment Options Cardiovascular disease (CVD) is currently the leading cause of death in the world and it is projected that it will remain as such throughout the next decade (World Health Organization 2011a; World Health Organization 2011b; Mathers and Loncar 2006). In 2008 alone, CVD accounted for 1 in 3 deaths in the US (811,940 of 2,471,984) and of those, half (approximately 1 in 6 American deaths) were attributable to coronary heart disease (Roger et al. 2012). Furthermore it is estimated that each year 785,000 Americans have a new coronary attack, about 470,000 have a recurrent attack and a further 195,000 suffer their first silent myocardial infarction (Roger et al. 2012). While the disorders categorized as CVDs have such divergent causes as atherosclerosis, rheumatic fever, congenital malformations and thrombosis, they all converge to cause damage to the heart muscle. Unfortunately, the damage is irreversible because the heart muscles cells, cardiomyocytes, are thought to be terminally differentiated and nonproliferative (Sutton and Sharpe, 2000; Laflamme and Murry, 2005), which necessarily limits the regenerative potential of the heart. Among CVDs, ischemic heart disease is the most prevalent (World Health Organization 2011b; Roger et al. 2012). It has therefore been the subject of intensive research. The partial or complete blockage of a coronary artery reduces/prevents blood supply to the downstream heart muscle and the affected tissue becomes severely nutrient and oxygen deprived, inducing cardiomyocyte death: an event known as a myocardial infarction (MI). Myocardium undergoes irreversible damage within 20 minutes of MI and a subsequent wave-front of cell death sweeps over the area of ischemia over a three to six hour period finally resulting in the death of up to a billion cells (M. A. Laflamme and Murry 2005). In the weeks that follow, cardiac function is greatly reduced due to the invasion of leukocytes into the infarct area, removal of dead tissue, and deposition of granulation tissue. Occurring in as little as two months the ventricular remodeling results in the formation of tough, rigid, collagenous fibrotic scar tissue at the site of (and surrounding the) infarct, and thinning of the ventricular wall; both leading to increased 3

ventricular wall stress and volume, and decreased ejection fraction and contraction force; see Figure 2-1(Sutton and Sharpe 2000; M. A. Laflamme and Murry 2005). In the end, without intervention and in most cases even with intervention, a myriad of health complications manifest, ultimately leading to irreversible heart failure and premature death.

Figure 2-1: Histological stages of myocardial infarction (Reproduced with permission from Nature Publishing Group, Laflamme and Murry 2005)

Even with treatment of the underlying cause of heart disease, approximately 50% of patients will experience heart failure within five years of an acute MI (Pantilat and Steimle 2004; Go et al. 2013). At present, the gold standard treatment for those who have reached end-stage heart failure is a heart transplant; but the insufficiency of donors combined with the need for patientdonor matched organs severely limits the number of patients that can be treated (Zammaretti and Jaconi 2004; Menasché 2008). In 2010, there were 2,333 heart transplants performed in the US (Roger et al. 2012). In 2011, greater than 3,000 people were on the waiting list for a heart transplant and thus living with end-stage heart failure (Roger et al. 2012). Over the last few years left ventricular assist devices (LVADs) have shown great improvement in safety and efficacy. As such, LVADs have been increasingly used as a bridge to transplantation, giving those awaiting a transplant a credible life-saving therapy option by supporting the failing heart, with the consequence of reducing transplant waitlist deaths (Wilson et al. 2009; Kirklin et al. 2013). Additionally, after months to a year on an LVAD, many patients experience improvement in the global contractile properties of the heart (e. g. ejection fraction) allowing their removal from the device. More recently, some VADs have been approved as a destination therapy, providing individuals ineligible for a heart transplant with a long-term therapy option (Wilson et al. 2009; Kirklin et al. 2013). However, there remains a subset of 4

patients that do not experience a significant and/or long-lasting improvement on an LVAD. Moreover, there is a great need to develop therapies that prevent end-stage heart failure. Consequently, the field of cardiac tissue engineering emerged as a means of developing alternative sources of cardiac tissue and methods for replacing tissue damaged by CVD. Biomaterials have featured prominently in cardiac regenerative therapy and can be divided among two main strategies for inducing functional repair to the heart muscle post-injury: (i) the production of functional cardiac patches in vitro that can be implanted onto the damaged area and thereby directly replace the non-viable portion of a damaged heart, and (ii) injectable biomaterials (hydrogels) to help prevent cardiac remodeling, deliver cells to replace the damaged tissue, and recruit endogenous cell types in an attempt to restore functionality (Menasché 2008). In considering possible delivery of cells to the heart for tissue repair, the first issue that arises is that of cell source.

2.2. Cell sources, stem cells, & differentiation One of the biggest issues with cardiac repair is that adult cardiac cells do not proliferate, so damaged tissue cannot repair itself. One cannot take a heart tissue biopsy, expand the cells to a sufficient number, and re-introduce them back into the patient. This has led to the need to find a cell source, ideally autologous, capable of differentiating into functional cardiomyocytes for use in cardiac tissue engineering. Many cell types have been assessed or used as potential cardiomyocyte replacements post infarction (Q. Z. Chen et al. 2008). The use of various adult cell types including skeletal myoblasts and mesenchymal and hematopoietic stem cells have all been assessed and although advantageous as they can come from autologous sources, results have been varied, as will be discussed in coming sections. Neonatal cardiomyocytes are a common cell source for they retain some proliferative capacity and have shown ability to couple with adult myocytes, however their clinical use in humans is not feasible (Reinecke et al. 1999; Reffelmann and Kloner 2003). They do however remain a good model cell source in developing methods for regeneration as they provide a model of what can be achieved (Reffelmann and Kloner 2003). Embryonic and induced pluripotent stem cells (ESC and iPS cells, respectfully) are at the moment the most promising cell sources for true myocardial repair. Studies from a number of groups have shown that it is possible to generate CMs from mouse and human ESCs, and human iPS cells (Kattman, Huber, and Keller 2006; Yang et al. 2008; M. Zhang et al. 2001). 5

Flk1+ cells derived from ESCs and cultivated as embryonic bodies (EBs) were identified as multipotent cardiovascular progenitors, and it was further demonstrated that Flk1 expression can be used to enrich for cardiac specific progenitors (Kouskoff et al. 2005). When isolated from the differentiated EBs, on day four of culture, and cultured as a monolayer, these progenitors generate cultures highly enriched for contracting CMs. As these progenitors differentiate they progress through the developmental stages thought to be involved in the establishment of the cardiovascular lineages in vivo and the presence of specific cytokines is required. A combination of activin-A and bone morphogenetic protein-4 over the first four days of EB differentiation induces the initial stages of specification and subsequent exposure to dickkopf homolog 1 (DKK1) and vascular endothelial growth factor (VEGF) significantly enhances the differentiation of Flk1+ progenitors into CMs (Kattman, Huber, and Keller 2006; Yang et al. 2008). The method reported shows the potential of the indentified progenitor to successfully differentiate into the three cell types found in healthy adult tissue and thus is of great interest as a cell source. There does remain some concern in that ESCs are associated with teratoma formation due to their proliferative capacity and as they are not an autologous source there is immune response issues, as well as the ethical issues of using ESCs (Vunjak-Novakovic et al. 2010). A method for producing iPS cells from mouse, and subsequently human, somatic cells was originally reported by Yamanaka and colleagues, and is significant as iPS cells present a much needed autologous cell source for transplantation into the heart and differentiation into CMs and vasculature (Takahashi and Yamanaka 2006; Takahashi et al. 2007). Problems with the reported method, however, are low success rate at producing iPS cells (1/10000) and the use of a viral vector to deliver the four factors needed to induce pluripotency. The latter is disconcerting as it entails modification of the cells genome and thus raises questions of long-term effects. A new method recently reported by Warren et al. (2010) uses mRNA to deliver the same transcription factors as Yamanaka but with much higher efficiency and success rate at producing iPS cells very close in phenotype to true ESCs. Furthermore, addition of further mRNA with differentiation factors showed success at directing iPS cell fate (Warren et al. 2010). Advantages of the newly reported method are that cell reprogramming to a pluripotent state and subsequent differentiation can be achieved without compromising genetic integrity (Warren et al. 2010). As advances in iPS cell production continues they remain a clear favorite in terms of a bench mark 6

cell source for transplantation, however the concern of teratoma formation upon transplantation, as with ESCs, remains.

2.3. Design criteria for biomaterials in cardiac tissue engineering The complexity of the native heart environment and pathophysiology of post-MI remodeling challenges the development of strategies for the treatment and design of biomaterials. An important aspect of biomaterial design is the consideration of the objective of the biomaterial application. Design criteria will be different whether the biomaterial is a vehicle for cell delivery, a functionalized material aimed at artificially retaining normal ventricular geometry, or a scaffold to generate tissue patches. However, there are some common design criteria that must first be addressed: (i) biocompatibility, (ii) biodegradability, (iii) mechanical support, (iv) injectability (in the case of hydrogels), (v) clinically relevant thickness (in the case cardiac patches), and finally (vi) envisioned application time post-infarct (Leor, Landa, and Cohen 2006; Q. Z. Chen et al. 2008; Vunjak-Novakovic et al. 2010; Bouten et al. 2011). 2.3.1. Biocompatibility

Biocompatibility is generally defined as the “ability of a material to perform with an appropriate host response in a specific application” (Williams 1987). In the context of cardiac tissue engineering this encompasses the need for the material to function without initiating a significant foreign body response in vivo, while retaining the ability to support both cardiomyocyte survival in vitro and in vivo readily without cytotoxicity, as well as the contractile function of the myocardium (Leor, Landa, and Cohen 2006; Q. Z. Chen et al. 2008; VunjakNovakovic et al. 2010). This does not preclude the activation of the host inflammatory and immune response but rather focuses on mitigating and controlling the type of response in order to prevent further injury to the heart and not impede its function. Specifically, a biocompatible material should be resistant to blood clotting and bacterial colonization, and if immunogenic should not recruit cell types that can exacerbate the remodeling process. For example, activation of the host immune response such that there is preferential recruitment of reparative M2 macrophages over cytotoxic M1 macrophages is generally considered to be a beneficial trait for biomaterials. 2.3.2. Biodegradability

7

Biodegradability refers to the mechanism through which an implanted material breaks down and the inherent life-span of the material. A thorough discussion of the various mechanisms and subsequent definitions are covered in detail elsewhere (Treiser et al. 2010). However, in brief, there are three modes to consider: (i) bioerosion is degradation of a material through hydrolytic mechanisms (covers both surface and bulk erosion); (ii) bioresorption is degradation through cellular activity; and (iii) biodegradation is degradation through enzymatic activity. In the context of cardiac tissue engineering, a biomaterial is considered to be biodegradable if degradation occurs through disintegration, a hydrolytic mechanism or by enzymatic activity that the biomaterial will encounter in vivo; and that the degradation products similarly conform to the requirements of both biocompatibility and biodegradability (Leor, Landa, and Cohen 2006; Q. Z. Chen et al. 2008; Vunjak-Novakovic et al. 2010; Bouten et al. 2011). While biocompatibility and biodegradability are distinct concepts, they are often considered in tandem during biomaterial design as there is little use in designing a biocompatible material that degrades into toxic components. Many physiological extracellular matrix (ECM)-based biomaterials readily fit these dual criteria (e. g. fibronectin and collagen) as they inherently contain the correct molecular composition required for cell attachment and survival, and they are readily degraded in vitro and in vivo within days to weeks by enzymes secreted by cells into biocompatible and biodegradable degradation products (Q. Z. Chen et al. 2008). Notably, cells can turnover these ECM biomaterials and replace them with their own ECM components, thereby remodeling their environment as necessary (Z. Li and Guan 2011). However, sourcing these materials can be complicated by the fact that they may retain many of their surface antigens and may elicit an immune response if used in xeno-transplantation. Despite this limitation there are sources of the ECM-biomaterials that have been approved for human use, including fibrin that can be isolated from a patient’s own blood (Odedra et al. 2011). Conversely, synthetic materials have been developed such as polyethylene glycol (PEG), poly (glycerol-sebacate) (PGS), and poly (tetrafluoroethylene) (PTFE) with defined chemical compositions and designed to have no foreign body response. Modification of the chemical composition can permit the selection of degradation rates in the range of a few weeks to years. However there is the concern as to whether the degradation products are truly being removed from the body or rather accumulate, the long-term effects of which are unknown (Q. Z. Chen et 8

al. 2008; Z. Li and Guan 2011). An additional limitation is that these synthetic biomaterials often do not support cell adhesion and survival, and therefore need to be modified with appropriate bioactive molecules (Vunjak-Novakovic et al. 2010). Another important consideration in terms of biodegradability is the issue of how quickly the material should be removed in order to properly execute its desired function when applied for cardiac ventricular repair. Biomaterials designed for cell delivery, recruitment, and survival (anti-apoptotic, pro-angiogenic) should survive in vivo at least one week, based on the fact that most cell death occurs within the first few days post-MI; and should be fully degraded in 6-8 weeks (in animal models) recognizing that pathological remodeling is complete by approximately 6 weeks after MI (Patten et al. 1998; Krzemiński et al. 2008; Z. Li and Guan 2011). Thus, the biomaterial should remain long enough to have the desired effect but no longer than necessary as it may become a hindrance to repair. For example, improved cardiac function post treatment was demonstrated with fibrin glue used to transplant skeletal myoblasts into ischemic myocardium wherein it degraded in 7-10 days (Christman, Vardanian, et al. 2004). Scaffold degradation is likely to require a similar timeframe. Most scaffolds are designed to be quickly replaced by new ECM secreted from the seeded cells, with the engineered tissues that are implanted containing very little of the original scaffold material. For scaffolds that are present at the time of implantation, in vivo degradation should not exceed weeks to months and should be quickly replaced by functional tissues. Zimmermann et al. have spent considerable time developing functional cardiac tissue constructs in vitro demonstrating that their collagen-based scaffold can be remodeled and replaced by maturing cardiomyocytes resulting in a scaffold-free transplantable engineered heart tissue construct (Eschenhagen et al. 1997; W.-H. Zimmermann 2001). In the case of biomaterials designed to provide support to the failing ventricle, they should have relatively slow degradation rates on the order of months to years and while controversy persists as to the timeframe required, it may be desirable for such materials to remain for the very long-term (see Section 2.5.4 for examples and further discussion) (Nelson et al. 2011). It is important to note however that there is still much debate as to the mechanical properties and degradation requirements that are necessary for certain outcomes due to the large disparity between studies (Nelson et al. 2011). These topics therefore continue to be active areas of investigation. 2.3.3. Mechanical Support

9

The third criteria of mechanical support requires some forethought as to the envisioned application with consideration of whether the biomaterial can withstand the mechanical demands that will be placed upon it after ventricular application and also whether it is likely to interfere with the normal mechanical functioning of the surrounding tissue. The in vivo model system for the biomaterial should be considered with respect to species-specific mechanical demands. This is because the mechanical forces placed upon a biomaterial by the human heart will vastly differ from the forces exerted by a small rodent heart in an animal application. Specifically, the human myocardium ranges in stiffness from 20kPa (end of diastole) to 500kPa (end of systole), whereas rat myocardium ranges from 0. 1 to 140kPa (Q. Z. Chen et al. 2008; Vunjak-Novakovic et al. 2010; Bouten et al. 2011; Venugopal et al. 2012). A material envisioned to artificially thicken the ventricle wall and maintain ventricular geometry during remodeling should have a stiffness in the high end of the range characteristic for the native ventricle; whereas a material designed to be injected, to act as a temporary matrix for transplanted cells and/or to recruit endogenous cells can have a low-end stiffness so long as it is sufficiently stiff to withstand the contraction/dilation of the heart. Biomaterials intended for in vitro applications can have a very low stiffness, as long as the cells seeded into it (hydrogels or scaffold) are able to remodel it into a material that as a final product is mechanically similar to the native myocardium. This is especially evident in the case of cardiac tissue grafts for repair of full thickness defects, wherein mismatching the mechanical properties can result in the grave consequences of inducing undue strain on the injured heart if the graft is too stiff or graft failure due to the stresses experienced in vivo if the graft is insufficiently stiff (Ozawa et al. 2002). Furthermore, issues such as burst pressure and suture retention must be considered as the graft has no time to integrate with the host tissue since it experiences full cardiac load immediately upon implantation. Results from tissue engineering of arterial grafts provide useful benchmarks for these properties that are achievable using current tissue engineering methods e.g. burst pressure of over 3000mmHg and suture retention strength of over 160g (L’Heureux et al. 2006; Dahl et al. 2011). Motivated by these challenges, Lang et al have been developing a surgical glue to improve cardiac grafts and achieved success in repairing a full thickness defect (N. Lang et al. 2014). In general, naturally-derived biomaterials have weak mechanical properties, with moduli in the tens of Pa to tens of kPa range (Q. Z. Chen et al. 2008). Moreover, there is a batch-to-batch and source-to-source variability in the physical properties of these biomaterials. As a 10

consequence, biomaterials made purely of naturally-derived components are limited as to their mechanical support applications. Synthetic materials, on the other hand, are more consistent in their composition between batches and have mechanical properties such as stiffness, elasticity, and porosity that can be precisely controlled (Q. Z. Chen et al. 2008). 2.3.4. Injectability

A hydrogel that can pass through a fine gauge needle (~27G) is described as injectable as it is possible to safely administer it into the heart in a minimally invasive manner. Injectability can be achieved by two approaches wherein gelation of the hydrogel (by temperature, chemical, lightinduced cross linking (Yeo et al. 2007; Habib et al. 2011), etc.) is: (i) initiated but not completed prior to the hydrogel passing through the needle, and (ii) initiated after delivery to the desired site. Importantly, the polymerization time should be in the order of minutes to tens of minutes to ensure the hydrogel is delivered and successfully localized at the site of injection and not completely washed out (Vunjak-Novakovic et al. 2010). This is because for polymerization that requires tens of minutes to hours to complete there is enough time for the biomaterial to be subjected to the contraction of the heart and to be carried away in the blood stream, rather than gel properly in the ventricle wall. 2.3.5. Clinically Relevant Thickness

Biomaterials used for tissue engineering strategies where cells are cultured with the biomaterial in vitro then implanted in vivo have their own distinct requirement in that they must be capable of supporting the cultivation of tissues of clinically relevant thickness: up to ~10 mm for full thickness cardiac grafts, whereas tissue patches can be thinner (Chiu and Radisic 2010). The limits of oxygen diffusion within a metabolically active tissue of high cell density (e.g. 108cells/cm3) restricts tissue thickness to approximately 200µm, thus scaffolds often require that a primitive vascular network or a channel array for culture medium perfusion be incorporated into the design to allow sufficient nutrient exchange to the centre of the tissues during in vitro cultivation (Radisic et al. 2006). This is a major issue in producing in vitro cardiac patches that will be discussed in greater detail in Section 2.4. 2.3.6. Application Time

An additional consideration that can influence biomaterial design for cardiac regeneration therapy is the time post-infarction at which the biomaterial is to be applied since new and old infarcts present their own unique challenges. The rapid cell death that results from nutrient and 11

oxygen deprivation downstream of a coronary artery blockage suggests that a cell injectionbased strategy should be most effective if applied shortly after an MI. The application of a biomaterial modified with both cells and growth factors within hours or days after an MI may promote directed wound repair such that the scar tissue formed would be minimized, the contractile function maintained in the border zone, the ischemic area reduced, and consequently pathological remodeling attenuated. Recent work has shown that early application of a collagen based hydrogel 3 hours after MI significantly improved pathological remodeling when measured 4 weeks after treatment in comparison to application of gel 7 or 14 days after infarct (Blackburn et al. 2014). Injection 7 days after MI did show significantly improved cardiac function in comparison to injection 14 days after, which showed no improvement over PBS injection, however not to the same level injection 3 hours after infarct imparted, thus indicating the earlier the application the better the outcomes (Blackburn et al. 2014). Notably, incorrectly timed administration of the therapy could potentially exacerbate the problem. While early delivery of cells might conceptually be more effective by initiating early re-vascularization and contributing to the protection of the spared myocardium, it may at the same time expose the delivered cells to a very hostile environment because of the significant immune response, the presence of cell death-associated cytokines and the by-products of dead and dying cells in the infarcted area soon after insult, which can compromise the viability of the cells. The correct time to deliver cells has been investigated in the clinic with both the TIME and SWISS-AMI trials showing that injection of cells from 3 days to 4 weeks after infarct did not improve LV cardiac functional or pathological outcomes (Traverse 2011; Traverse 2012; Sürder et al. 2013). A scaffold-based contractile tissue engineering strategy while applicable in the acute phase may have a more significant effect if implanted after scar formation. Similarly, larger areas of damaged cardiac muscle evident in chronic cases might benefit from a scaffold-based regenerative therapy approach. Therefore, choosing the right time point post-MI for an intervention is a challenge and no clear consensus has been reached.

2.4. Tissue engineered cardiac grafts (in vitro engineering) In vitro tissue engineering approaches involve pre-organizing cells seeded into scaffolds or on their own into a functional tissue (i.e. one capable of contracting and propagating electrical signals). To be clinically relevant any engineered cardiac tissue must have properties (functional and morphological) similar to that of native heart tissue and remain viable after implantation 12

(Wolfram-Hubertus Zimmermann, Melnychenko, and Eschenhagen 2004). Challenges in the development of such in vitro tissues is that oxygen diffusion limits tissue sizes to 100-200 µm, thus vascularization is required, and generating tissues that can electromechanically couple with the native myocardium (and generate significant contractile force) (Vunjak-Novakovic et al. 2010). A number of different groups have tried to tackle these problems using different approaches ranging from using decellularized tissue as a scaffold, using mechanical/electrical stimulation of cells seeded in hydrogel scaffolds, to layering individual cell sheets. In an effort to improve the properties of engineered heart tissue (EHT) Eschenhagen et al. (1997) seeded a mix of collagen I, extracellular matrix proteins (Matrigel), and neonatal rat cardiomyocytes into lattice or circular molds and saw spontaneous remodeling of the liquid reconstitution mixture and a development of spontaneously and synchronously contracting solid EHTs after 5-7 days (Eschenhagen et al. 1997). Subsequent culture of EHTs with cyclic mechanical strain improved morphological, functional, and mechanical properties of the EHT, and by stacking a number of these constructs together formed spontaneously contracting tissue 14 mm thick that could be implanted at the site of infarct in a rat MI model. One month later it was observed that the constructs showed un-delayed electrical coupling with the native tissue and improved diastolic and systolic function of rats who received the patches in comparison to sham-operated rats (W. H. Zimmermann et al. 2006). Another approach to mechanical stimulation for achieving functional improvement of EHT was noted by Radisic et al. (2004). It was demonstrated that neonatal cardiomyocytes seeded on collagen sponges with Matrigel and exposed to physiologically relevant electric field stimulation during culture induced the formation of mature myocardium with elongated, viable cells aligned in parallel. Ultra-structural organization in stimulated cardiac constructs was remarkably similar to that present in the native myocardium, with hallmarks of electromechanical cell coupling including gap junctions, intercalated disks and sarcomeres all markedly more frequent than in non-stimulated constructs (Radisic et al. 2004). A pioneering study done by Ott et al. (2008) looked to tackle the larger problem of wholeorgan transplant, while at the same time illustrating an approach for answering the questions of oxygen transport and tissue thickness in engineered heart tissues. By decellularizing whole, adult, cadaveric Fisher rat hearts by coronary artery perfusion of detergents they were able to obtain whole heart scaffolds. The process preserved the underlying extracellular matrix (ECM), 13

and produced an acellular scaffold with perfusable vascular architecture, competent acellular valves and intact chamber geometry. Reseeding the scaffolds with neonatal rat CMs or endothelial cells and culturing by perfusion in a bioreactor mimicking cardiac physiology they observed macroscopic contractions by day four. At day eight they were able to show a small, but significant pump function comparable to about 2% of the adult heart when the engineered hearts were exposed to physiological loads and electrical stimulation (Ott et al. 2008). A method for developing functional EHT was also accomplished using no scaffold at all, but stacking sheets of neonatal rat cardiomyocytes cultured on the surface of Poly(Nisopropylacrylamide)-grafted polystyrene dishes. The treated dishes allow for cell attachment at 37°C, but culturing at 20°C changes the properties of the surface causing cell layers to detach, at which point individual layers can be stacked, and these stacked sheets looked like homogeneous heart-like tissue (Miyagawa et al. 2005). When these scaffold free cell stacks were transplanted onto infarcted rat hearts they became attached to the infarcted myocardium, showed angiogenesis, expressed connexin-43 (a gap junction protein associated with functionally coupled cells), appeared as a homogeneous tissue in the myocardium, and significantly improved cardiac performance (determined by echocardiography) in comparison to those that did not receive the treatment (Miyagawa et al. 2005). Furthermore, connexin-43 staining and transmission electron microscopy (TEM) showed the existence of gap junctions and intercalated disks (associated with coordinated cardiac muscle contraction) between implanted and host tissue. One limitation of this technique is that it is still limited to the hundreds of µm in thickness unless methods for oxygen transport or vascularization are introduced. Many of the issues with engineering tissue in vitro are the same as those for direct in vivo application (injection methods), including appropriate biomaterial (scaffold) choice, cell source, and survival/integration once transplanted, thus the two methods are not mutually exclusive. Injection of cells/hydrogels into heart tissue shortly after MI has the potential to minimize the formation of scar tissue and attenuate the pathological remodeling process whereas tissue engineered cardiac patches may be more useful in later stages of remodeling such as the complete replacement of non-contractile areas. In vitro tissue engineering may also provide living patches for repair of congenital malformations that cannot be addressed through injectable methods (Soonpaa and Field 1998).

2.5. Injection (in vivo engineering) 14

The first type of injection that must be discussed is that of direct cell transplantation. Based on the successes and limitations from such injections the use of biomaterials was developed to aid in the successful transplantation of cells. In general, injectable hydrogels for cardiac regeneration and repair can be divided into three categories according to their state of development and current application. The first group of hydrogels is designed to promote endogenous repair through maintaining cell survival, recruiting endogenous cells or inducing neovascularization. The second group acts a temporary matrix for cell transplantation and exogenous repair. The final group consists of injectable hydrogels that act as bulking material to support the failing left ventricle, in turn maintaining or restoring normal geometry of the heart, and improving cardiac function. It should be noted that there is a significant overlap between the categories, and most hydrogels fit into more than one category. 2.5.1. Cell only based therapies

The first type used in clinical studies for cell based cardiac repair was skeletal myoblasts and results were varied. Skeletal myoblasts were initially chosen due to their relative ease of isolation and culture (expansion) from small muscle biopsies, pre-clinical efficacy, non-immunogenicity, and ability to cope well with ischemic environments (Menasché et al. 2003). Injection showed reduced infarct size, survival of transplanted cells, and slight improvement of cardiac function, however it was found that most cells did not trans-differentiate into cardiomyocytes (as hoped, becoming skeletal muscle) and therefore lack the proteins needed to electromechanically couple with the native myocardium and it was reported that the grafts did not beat in synchrony with the surrounding tissue (Reinecke, Poppa, and Murry 2002; Reinecke et al. 2000). Even so, Phase I trials were deemed successful and Phase II trials are currently underway. Neonatal cardiomyocytes retain the ability to proliferate and have received much attention as a potential cell source. Injection of CMs isolated from neonatal rats and transplanted to the site of infarct in adults have shown attenuation of pathological remodeling leading to improved function, and even integration of cells into the host myocardium through gap junctions and intercalated disks (Reinecke et al. 1999; Muller-Ehmsen 2002; R. R. K. Li et al. 1996). Further studies, however, show that direct injection of these cells results in survival of only ~50% of cells immediately after injection, loss of up to ~90% of cells due to extrusion, and only ~10% survival of the remaining cells after as little as one week (Müller-Ehmsen et al. 2002; M. Zhang et al. 2001). 15

Mesenchymal and hematopoietic stem cells have also been looked at as potential cell sources for direct injection and have had some promising results. They are advantageous for the two cell lines have multi-lineage differentiation potential, and in vitro studies showed promise in the two cell type’s ability to become cardiomyocytes. Direct injection of these cell types into MI areas showed incorporation into host myocardium, reduced infarct size, and even showed functional improvements, however it has been noted that fusion events are rare and no significant muscle mass has been generated in vitro (Bittira et al. 2002; Makino et al. 1999; Shake et al. 2002; Orlic, Kajstura, Chimenti, Jakoniuk, et al. 2001; Murry et al. 2004). Perhaps the most interesting observation about cell transplantation is that regardless of cell type, developmental stage, or origin most studies showed some improvement in cardiac function post injection; yet all seem to have the same issues with incorporation into the native tissue, survival, and retention (W. H. Zimmermann et al. 2006). Meta-analysis of clinical trial results demonstrated a significant, albeit low 3%, increase in left ventricular ejection fraction (LVEF) as well as a significant reduction in infarct size (-5.6%) and end systolic volume (-7.4ml) in patients treated by intracoronary cell injection after acute MI (Lipinski et al. 2007). Dose-response between injected cell volume and LVEF change was reported (Lipinski et al. 2007). Although these studies are encouraging, modest improvements motivate investigation of new cell sources and methods that increase survival and retention of injected cells. It has also been suggested that the improvements seen with cell transplantation may be due to release of cardio-protective and angiogenic cytokines from transplanted cells that modulate remodeling rather than integration of the cells into the host myocardium (Menasché 2008). As such, groups have also looked at identifying specific cytokines and directly injecting them in place of cells. While the delivery of cells to the injury site is a major aspect of cardiac regeneration therapy and the cell types used are many and varied, the topic is discussed in further detail elsewhere (Shiba et al., 2009; Hilfiker et al., 2011; Martinez and Kofidis, 2011; Liau et al., 2012). 2.5.2. Hydrogels to promote endogenous repair

The group of injectable hydrogels that promote endogenous repair works based on either the bioactivity of the base biomaterial or the incorporation of bioactive molecules that are antiapoptotic or cardioprotective (i.e. QHREDGS, Tβ4, insulin-like growth factor-1), angiogenic (i.e. Tβ4, bFGF, VEGF, PDGF, hepatocyte growth factor, pleiotrophin plasmid), or chemotactic 16

(i.e. stromal cell-derived factor-1). The incorporation of bioactive molecules into injectable hydrogels is important for the successful delivery of the biomolecules and in turn cardiac regeneration, since biomolecules in the soluble form are rapidly diffused from the injection site and proteolytically degraded in the in vivo environment. The existence of different cardiac progenitor cell types has been reported in native heart tissue, and it was demonstrated that these progenitors can be induced to migrate to areas of infarct with the addition of specific peptides or cytokines (Beltrami et al. 2003; Oh et al. 2003; Martin et al. 2004; Cai et al. 2003). One such peptide is thymosin β-4 (Tβ4) which has been shown to promote migration and survival of CMs, endothelial cells, smooth muscles cells, and vascular progenitors from cardiac explants in vitro and has garnered attention for clinical trials as a drug for patients with acute MI (Smart et al. 2007; Bock-Marquette et al. 2004; Crockford 2007). A temperature-responsive hydrogel made of chitosan and glycerol phosphate was used to encapsulate bFGF, which was released through hydrogel degradation (H. Wang et al. 2010). After 4 weeks of injection in a rat MI model, the bFGF-hydrogel improved cardiac function (i.e. decreased LV end-diastolic diameter and LV end-systolic diameter, as well as increased LV ejection fraction and LV fractional shortening), as compared to PBS control, bFGF in PBS, and chitosan hydrogel alone (H. Wang et al. 2010). This may be due to the increase in arteriole density and in turn decreased infarct size in the bFGF-hydrogel group compared to all other groups. The chitosan hydrogel alone also increased arteriole density and decreased infarct size compared to PBS control and bFGF in PBS (H. Wang et al. 2010). This shows that the presence of the injectable hydrogel was necessary for mechanical support and sustaining the release of encapsulated bFGF, while the incorporation of bioactive molecules such as bFGF further improved endogenous cardiac repair through promoting arteriogenesis. By contrast, treatment with soluble bFGF showed no improvement compared to PBS control due to the short half life of the growth factor and its high diffusibility from the injection site (H. Wang et al. 2010). Other groups also investigated the incorporation and release of bFGF from various hydrogel systems, including encapsulation into biodegradable gelatin hydrogel microspheres (Iwakura et al. 2003; Yamamoto et al. 2001; Y. Liu et al. 2006), UV crosslinkable chitosan hydrogels (Fujita et al. 2005), and synthetic random copolymer poly(N-isopropylacrylamide-co-propylacrylic acid-cobutyl acrylate) (Garbern et al. 2011). These studies all showed increased vascular density with 17

the injection of bFGF hydrogels, and in turn improved cardiac function after myocardium infarction in rats, rabbits, or dogs. Conjugation of VEGF to a temperature-sensitive, aliphatic polyester hydrogel, poly (δvalerolactone)-block-poly (ethylene glycol)-block-poly (δ-valerolactone) (PVL-b-PEG-b-PVL), showed that the hydrogel provided localized and sustained VEGF activity, which increased blood vessel density (Wu et al. 2011). While the hydrogel alone maintained thicker ventricular wall and improved cardiac function (i.e. fractional shortening, ventricular volumes, preload recruitable stroke work, end-systolic elastance) after MI in rats, the conjugation of VEGF to the hydrogel further enhanced cardiac repair (Wu et al. 2011). While alginate, a bio-inert natural material, has been shown to prevent cardiac remodeling and dysfunction in the rat MI model, it has also been modified with peptides and growth factors to create a bioactive injectable hydrogel for endogenous cardiac repair (Landa et al. 2008). Alginate that was modified with RGD conjugation led to increased angiogenic response compared to nonmodified alginate when injected into the infarct area in rats 5 weeks post-MI (Yu, Gu, et al. 2009). In addition, alginate hydrogel was found to be an effective injectable delivery system for cardiac repair post-MI, since it could sequentially deliver VEGF and PDGF to induce formation of mature blood vessels and in turn improve cardiac function (Hao et al. 2007). In this alginate system, PDGF was released more slowly than VEGF likely due to the difference in the affinities of VEGF and PDGF to alginate. As such, VEGF induced angiogenesis, while PDGF was present in the later time to mature the newly formed capillaries. This sequential delivery approach improved vascularization and cardiac repair compared to the delivery of VEGF or PDGF individually (Hao et al. 2007). Injectable affinity-binding alginate hydrogel microbeads were also used to sequentially release cytoprotective IGF-1 and pro-angiogenic HGF to increase angiogenesis and formation of mature blood vessels, as well as to prevent infarct expansion and scar fibrosis after 4 weeks in a rat acute MI model (Ruvinov, Leor, and Cohen 2011). The injection of these hydrogel microbeads also reduced cell apoptosis, induced cell cycle re-entry of CMs and enhanced presence of GATA-4 positive cells, indicating an improved endogenous cardiac regeneration (Ruvinov, Leor, and Cohen 2011).

18

Hydrogels have also been used to enhance gene therapy. The human VEGF plasmid was injected in a rat MI model with an amphiphilic, thermo-responsive polymer synthesized by alternatively cross-linking poloxamer and di-(ethylene glycol) divinyl ether (Kwon et al. 2009). The sustained human VEGF expression within the infarct area increased capillary density and formation of larger blood vessels. In a separate study, the pleiotrophin plasmid was shown to induce neovasculature formation in the ischemic rat myocardium without angioma formation, but the injection of the plasmid in saline resulted in low transfection efficiency due to limited exposure of cells to the plasmid (Christman et al. 2005). The incorporation of pleiotrophin plasmid in fibrin glue created a gene-activated matrix that increased neovasculature formation compared to injection of the plasmid in saline. To deliver stromal cell-derived factor-1 (SDF-1) as a chemokine to attract stem cells for cardiac regeneration, a modified chemokine S4V was designed to retain chemotactic bioactivity while being resistant to cleavage by matrix-metalloproteinase-2 and exopeptidase (Segers et al. 2007). S4V was then tethered to self-assembling peptides that formed nanofibers suitable for local delivery. The intra-myocardial delivery of S4V-nanofibers post-MI in rats led to the recruitment of c-kit positive stem cells and increased capillary density, as well as improved cardiac function (i.e. increase in ejection fraction) (Segers et al. 2007). Hydrogels derived from the native cardiac extracellular matrix (ECM) retain the complexity of the extracellular environment normally found in vivo and were demonstrated to improve maturation of cardiomyocytes derived from human embryonic stem cells (DeQuach et al. 2010), thus they constitute a promising approach for myocardial repair (Duan et al. 2011). ECM components have been isolated from the healthy myocardium and used as injectable hydrogels for endogenous cardiac repair. Cardiac ECM-derived hydrogels are prepared by decellularizing heart tissues with sodium dodecyl sulfate, followed by lyophilization and milling into a fine powder that is then solubilized by enzymatic digestion using pepsin and HCl and neutralized using NaOH (Singelyn et al. 2009; Seif-Naraghi et al. 2010; Singelyn et al. 2012). Myocardial matrix yielded from the decellularization of porcine myocardial tissue gelled at 37°C and showed collagen and glycosaminoglycan content (Singelyn et al. 2009). In addition, the matrix induced the homing of endothelial and smooth muscle cells, as shown by cell migration towards the matrix in vitro, as well as cell infiltration into the gel and increased arteriole formation in vivo 11 days post-injection within the rat myocardium. While the solubilized form of the decellularized 19

porcine pericardium was a stronger chemo-attractant for endothelial and smooth muscle cells, both decellularized porcine and human pericardium were shown to retain the native protein and glycosaminoglycan and induce neovascularization when injected into the left ventricular free wall of rats (Seif-Naraghi et al. 2010). Moreover, a low number of c-kit+ cells were found within the injection site, indicating an increase in endogenous CMs with the injection of the material (Seif-Naraghi et al. 2010; Singelyn et al. 2012). As such, the use of decellularized cardiac matrix as injectable hydrogels can promote endogenous homing and in turn cardiac-specific tissue formation, as mediated through the presence of cardiac-specific cues. In addition, cardiac function was maintained without arrhythmias in the rat MI model, when treated with the ECM hydrogel (Singelyn et al. 2012). Besides demonstrating the use of decellularized ECM in small animal models, it has also been shown that hydrogels made from decellularized porcine ventricular ECM can be delivered via percutaneous, transendocardial catheter injection in both healthy and infarcted porcine myocardium (Singelyn et al. 2012). The injection was successful, with no catheter clogging. While there was some leakage of the injectate into the ventricle, it was confirmed that a hydrogel was formed within the porcine myocardium without any hydrogel found in other organs, since any leaked matrix would be rapidly diluted in the blood and prevented from gelation. This study shows the feasibility of delivering in situ gelling materials via transendocardial injection in a large animal model and motivates the translation of injectable materials for MI treatment in humans (Singelyn et al. 2012). Keeping with ECM derived materials, gels derived from the small intestinal submucosa extracellular matrix were found to be a suitable injectable material for cardiac repair, due to the presence of bFGF (Okada et al. 2010). The gel induced angiogenesis and reduced infarct size compared to saline control in the murine MI model (Okada et al. 2010). In another study, the injection of small intestine ECM emulsion into the myocardium recruited c-kit positive cells, myofibroblasts and macrophages after rat MI (Zhao et al. 2010). It also increased VEGF levels, enhanced angiogenesis and improved cardiac function (i.e. fractional shortening, ejection fraction, stroke volume) compared to saline control, likely mediated by the cell recruitment (Zhao et al. 2010). Similarly, fibrinogen, the precursor to fibrin, has also been used in biomaterial strategies. Rufaihah et al have demonstrated the applicability of a photocrosslinkable semi-synthetic PEGfibrinogen hydrogel loaded with VEGF in rat models of MI. They showed that a PEG-fibrinogen 20

hydrogel was able to store and release VEGF in a sustained and controlled fashion, as well as significantly improve arteriogenesis and cardiac performance after infarct (Rufaihah et al. 2013). With tunable mechanical properties they have also shown the ability of the PEG-fibrinogen hydrogel on its own to mechanically support the failing LV and to improve cell transplantation, giving the developed hydrogel multiple courses of action as well as promise as a clinically relevant biomaterial (Habib et al. 2011; Rufaihah et al. 2013). G-CSF has also been studied in depth but has seen varied results in experiments with various animal species (Orlic, Kajstura, Chimenti, Limana, et al. 2001; Iwanaga et al. 2004; Deten et al. 2005). While it undoubtedly promotes migration of progenitors to the site of infarct the population is most likely non-cardiac lineage. It has been suggested, and to some extent shown, by a number of groups that G-CSF may not induce cardiac regeneration but mitigate the damage done during MI and influence cardiac repair (Nygren et al. 2004; Minatoguchi et al. 2004). One group showed in mice that G-CSF treatment reduced final infarct size and resident cell apoptosis, and increased vascular density and ventricular function in comparison to mice who did not receive the treatment (Harada et al. 2005). Despite some conflicting studies and varied results GCSF has already entered clinical trials being administered to patients with recent MI, and has shown some promising results (H.-J. Kang et al. 2004; H.-J. Kang et al. 2007; J. Kang et al. 2014). 2.5.3. Hydrogels for the delivery of cells for regeneration

Early studies in cell transplantation for cardiac repair lacked control over cell retention, survival, and function after injection of the cells into the heart. Later studies used natural and synthetic injectable hydrogels that polymerized in situ as a cell delivery vehicle to overcome such limitations. Some hydrogels used in cell delivery for cardiac regeneration include fibrin, chitosan, collagen, Matrigel™, and various synthetic polymers. Fibrin was one of the earliest biomaterials studied for use in the heart and for cell transplantation. It exhibits several advantages: it is naturally occurring and could be patientspecific; it is biocompatible, biodegradable, and pro-angiogenic. Furthermore, fibrin glue has ideal properties for cell delivery in cardiac repair due to its binding domains for growth factors and receptors. A three-dimensional fibrin hydrogel is formed upon injection with a dual-barreled syringe, which contains fibrinogen and fibrinolysis inhibitor aprotinin in the first barrel and thrombin and CaCl2 in the second barrel (Christman, Fok, et al. 2004). The cross-linking 21

mechanism of the fibrin hydrogel is similar to how clotting occurs in vivo. Skeletal myoblasts that were injected with fibrin glue into the myocardium post-MI were found in both the border zone and within the infarct, while myoblasts were found only in the border zone when injected in bovine serum albumin (Christman, Fok, et al. 2004). Fibrin gel provided a temporary scaffold for transplanted cells and contained RGD motifs that bind to cell receptors, thus entrapping the cells upon injection and improving their survival (Christman, Vardanian, et al. 2004). The surviving myoblasts at five weeks post-injection into the myocardium were located in clumps around arterioles (Christman, Vardanian, et al. 2004). The injection of fibrin glue with or without skeletal myoblasts led to smaller infarcts compared to injection of bovine serum albumin or myoblasts in bovine serum albumin (Christman, Vardanian, et al. 2004). There was also increased arteriole density in the infarct area when skeletal myoblasts were delivered in the fibrin glue compared to delivery in bovine serum albumin (Christman, Vardanian, et al. 2004). Although fibrin seemed to significantly decrease infarct size and increase arteriole formation without the presence of transplanted cells, it could be seen that transplantation of cells without an appropriate injectable hydrogel was not capable of appreciable cardiac repair (Christman, Fok, et al. 2004; Christman, Vardanian, et al. 2004). In a separate study, rat adipose-derived stem cells isolated from subcutaneous adipose tissues were injected in fibrin glue into the rat left ventricular wall post-MI (X. Zhang et al. 2010). At 4 weeks after injection, the co-injection of cells and hydrogel showed greater graft size, better cardiac function (decreased LV end-diastolic diameter and LV end-systolic diameter, as well as increased fractional shortening and ejection fraction), and increased arteriole density compared to injection of cells alone (X. Zhang et al. 2010). In addition, another study was conducted to deliver human bone marrow-derived mesenchymal stem cells in fibrin glue in a nude rat MI model using percutaneous injection catheters (Martens et al. 2009). Fibrin glue increased cell retention and survival. More importantly, this study determined the viscosity limits for the delivery of the cell-hydrogel suspension using several commercially available catheters. By controlling the composition of fibrinogen and thrombin, the fibrin hydrogel could be tailored to be compatible with the catheters. This also shows the potential of using catheter delivery for cell transplantation therapies involving other cell types and hydrogels (Martens et al. 2009). While fibrin-based biomaterials show promise and have the additional benefit of fibrin being FDA approved for human use, these biomaterials may only provide short-term cardiac functional 22

benefits that decrease with time (Yu, Christman, et al. 2009; Nelson et al. 2011; Segers and Lee 2011). Another natural hydrogel used for cell transplantation is a temperature responsive hydrogel made of chitosan and glycerol phosphate similar to that described by Wang et al above (Lu et al. 2009). This hydrogel gels in approximately 10 to 15 minutes at physiological temperatures. Embryonic stem cells were injected into the infarcted heart wall of rats in the chitosan hydrogel, showing better cell retention at 24 hours and increased graft size, improved cardiac function, thicker ventricular wall and greater micro-vessel density at 4 weeks as compared to injection of the cells in phosphate buffered saline (PBS) (Lu et al. 2009). Histological staining of excised heart sections showed a significant presence of chitosan at 24 hours after injection, while sparse positive chitosan staining was found at 4 weeks. No trace of the hydrogel was found at 6 weeks post-injection, showing complete degradation of the hydrogel in vivo (Lu et al. 2009). Chitosan has also been explored for the delivery of adipose-derived mesenchymal stem cells into the infarcted heart (Z. Liu et al. 2012). The chitosan hydrogel improved MI microenvironment by recruiting SDF-1, which is a key chemokine for homing of stem cells, and scavenging reactive oxygen species generated by ischemia that would otherwise impair adhesion molecules of the delivered stem cells (Z. Liu et al. 2012). Pre-treatment of animals with gelatin hydrogel microspheres incorporating bFGF was used to improve transplantation of fetal rat CMs in rat MI models (Sakakibara et al. 2002). It has been shown that bFGF can enhance viability of CMs through neovascularization and also directly affect CMs by stimulating DNA synthesis, myocyte proliferation and differentiation during development. In the experimental groups with bFGF microspheres alone and with bFGF microspheres followed by CM transplantation, there was neovascularization in the scar tissue after 1 week. Importantly, more transplanted cells survived in the scar area with the pretreatment with bFGF microspheres compared to no pretreatment, in which transplanted cells only survived in the peri-infarct regions (Sakakibara et al. 2002). In another study, human embryonic stem cell (hESC)-derived CMs were delivered to the heart within Matrigel™ using a cocktail of pro-survival factors (M. a Laflamme et al. 2007). The prosurvival cocktail resulted in a 4-fold improvement in myocardial graft size relative to hESC-CMs injected with Matrigel™ alone after 1 week. After 4 weeks, increased LV wall thickening and reduced ventricular dilation, improved global function (fractional shortening and ejection 23

fraction), and attenuated changes in LV chamber dimensions were observed in hearts injected with hESC-CMs along with the pro-survival cocktail relative to pro-survival cocktail-only, buffer-only, or non-cardiac cell controls (M. a Laflamme et al. 2007). An example of a synthetic hydrogel aimed at improving cell transplantation therapies to reduce LV remodeling post-MI is an injectable α-cyclodextrin/poly(ethylene glycol)–bpolycaprolactone-(dodecanedioic acid)-polycaprolactone–poly(ethylene glycol) (MPEG–PCL– MPEG) hydrogel (T. Wang, Jiang, et al. 2009). Culturing bone marrow-derived stem cells (BMSCs) in the hydrogel showed it was non-toxic and allowed for the maintenance of cell morphology in vitro. In a rabbit LV MI model co-injection of α-cyclodextrin solution with BMSCs and MPEG–PCL–MPEG solution in the infarct area 7 days post-MI showed immediate gelation and localization (T. Wang, Jiang, et al. 2009). Histological analysis of excised heart cross sections 4 weeks after injection showed the hydrogel was absorbed (degraded), and that the hydrogel significantly increased BMSC retention and vessel density around the infarct compared to injecting BMSCs alone. Functional echocardiography studies demonstrated an increased LV ejection fraction (~77% increase) and attenuated left ventricular dilatation with hydrogel-BMSC co-injection compared to PBS and BMSC only controls (T. Wang, Jiang, et al. 2009). A synthetic injectable hydrogel made of oligo[poly(ethylene glycol) fumarate] (OPF) was developed to deliver mouse embryonic stem cells into the LV wall of a rat MI model (H. Wang et al. 2012). The hydrogel improved cell retention compared to delivery in PBS, and was biodegradable with complete degradation by 6 weeks after injection in the infarcted heart. The hydrogel alone significantly reduced the infarct size and improved cardiac function at 4 weeks after injection (H. Wang et al. 2012). However, the combination of cells and hydrogel further increased revascularization, improved cardiac function, decreased infarct size and fibrotic area, as compared to hydrogel alone and cells injected in PBS. Importantly, the transplanted embryonic stem cells expressed cardiovascular markers, such as cardiac troponin-T, von Willebrand factor and α-smooth muscle actin. As such, the OPF hydrogel not only supported cell transplantation by improving cell retention and survival, but the stem cells also underwent differentiation into the cardiac linage, thus showing potential of this delivery method for heart regeneration (H. Wang et al. 2012). Synthetic hydrogels have also been modified by incorporation of bioactive peptides and growth factors to improve their use for cell delivery. A synthetic PEG-based hydrogel modified 24

with matrix metalloproteinase (MMP) cleavable peptide domains was developed for the codelivery of Tβ4 and vascular cells (Kraehenbuehl et al. 2009). Tβ4 was physically entrapped in the hydrogel, which gels in approximately 30 minutes at physiological conditions. Due to the MMP cleavable peptide domains, the hydrogel was prone to cell-mediated proteolytic degradation and the encapsulated Tβ4 could be released due to the MMP secretion of encapsulated human umbilical vein endothelial cells in vitro. At the same time, the presence of Tβ4 led to improved adhesion, survival, migration and organization of the cells (Kraehenbuehl et al. 2009). In vivo, the MMP-responsive PEG hydrogel was used to co-deliver Tβ4 and hESCderived vascular cells into the infarcted myocardium of rats (Kraehenbuehl et al. 2011). By 6 weeks after MI, there was improved cardiac function (i.e. decreased end-systolic volume and increased ejection fraction) in the group with vascular cells injected in Tβ4-encapsulated hydrogel, as compared to treatment with PBS control, Tβ4-free hydrogel and Tβ4-encapsulated hydrogel. This improvement was due to the better organization of native CMs and endothelial cells in the combined Tβ4 and cell delivery, as supported by the presence of Tβ4 and the secretion of human vascular cytokines (i.e. VEGF, EGF, HGF) and pro-survival factors (i.e. survivin) by the transplanted vascular cells (Kraehenbuehl et al. 2011). They hypothesized that the gel provided temporary support and pro-survival factors as it substituted for the degrading endogenous matrix, and hESC-derived vascular cells contributed to formation of capillary-like vessels, stabilization of host vessels, secretion of paracrine factors, or induction of paracrine factor secretion from native rat cells (Kraehenbuehl et al. 2011). The injection of cell-free PEGhydrogel with Tβ4 improved cardiac function compared to PBS control, showing that the injection of the bioactive hydrogel alone could also be explored for endogenous cardiac repair (Kraehenbuehl et al. 2011). Injection of a synthetic self-assembling hydrogel RAD16-II (initially developed by Zhang et al. 1993) formed nanofiber microenvironments in the host myocardium and promoted recruitment of endogenous ECs and vascular smooth muscle cells, which appeared to form functional vascular structures, necessary for improving blood flow and improving the general environment for CM survival. Furthermore, co-injection of the self-assembling peptide with CMs showed survival (albeit low) of the transplanted cells, but significant increase in recruitment of endogenous progenitors (M. E. Davis et al. 2005). The addition of insulin-like growth factor-1 (IGF-1, a CM growth and differentiation factor) proved to be anti-apoptotic and 25

increased cell growth in transplanted (injected) CMs. It showed sustained, controlled, and targeted release of IGF-1 from the nanofibers for 28 days. Improved systolic function was observed when CMs were injected with the new functionalized hydrogel in a rat MI study (M. E. Davis et al. 2006). 2.5.4. Hydrogels for the artificial maintenance of ventricle geometry and repair

The third approach for preservation of cardiac function is the injection of a bulking material into the ventricular wall to prevent progressive adverse remodeling due to high wall stresses that develops after MI. Theoretical modeling suggests that hydrogels change the geometry and dilation mechanics of the ventricle in such a way as to reduce the elevated local wall stresses, and further that the mechanical support provided by the hydrogels improves the ejection fraction and the stroke volume/end-diastolic volume relationship (Z. Li and Guan 2011; Wall et al. 2006). To this end Fujimoto et al developed an N-isopropylacrylamide (NIPAAm), acrylic acid (AAc), and hydroxyethyl methacrylate-poly(trimethylene carbonate) (HEMAPTMC) based synthetic hydrogel (Fujimoto et al. 2009). It was found that a feed ratio of 86/4/10 poly(NIPAAm-co-AAcco-HEMAPTMC) monomers formed a hydrogel at 37°C, which gradually became soluble over a 5 month period, and furthermore that no degradation product cytotoxicity was observed in vitro (Fujimoto et al. 2009). The hydrogel or PBS was injected into the MI zone two weeks after MI in a rat model, and analysis was done 8 weeks after injection. In the PBS group, LV cavity area increased and contractility decreased at 8wk post-MI, while in the hydrogel group both parameters were preserved during this period, and furthermore the hydrogel was still present at 8 weeks (Fujimoto et al. 2009). Echocardiography was used to measure heart function and it was reported that the hydrogel maintained fractional area change (contractility) at the same level measured just prior to injection, and was ~55% better than the PBS group at 8 weeks post injection (10 weeks post MI). Tissue in-growth was observed in the hydrogel injected area and a thicker LV wall (~100% thicker) and higher capillary densities were found for the hydrogel versus PBS group (Fujimoto et al. 2009). Thermo-responsive materials that transform from solution to gel near body temperature are frequently employed for tissue engineering because their transitional properties make them easy to deliver and manipulate. T. Wang et al developed a series of hydrogels by introducing hydrophobic PCL (polycaprolactam)-grafted polysaccharide chains into the thermo-responsive poly(N-isopropylacrylamide) (PNIPAAm) network (T. Wang, Wu, et al. 2009). The resulting 26

hydrogel (Dex-PCL-HEMA/PNIPAAm) transitions from solution to gel within 30s at 37°C and is reversible within the same time frame. To test the reparative effect of this hydrogel to damaged cardiac function, 200µL was injected four days after MI in a rabbit model. One month after the treatment, histological staining of heart sections showed a significant increase in scar thickness and reduction in infarct size, compared with the control group (T. Wang, Wu, et al. 2009). Moreover, LV end systolic pressure was increased and end diastolic pressure reduced in comparison with the MI control group. It was also observed that hydrogel implantation partially reduced the dilatation of the LV, as shown by a significant increase in ejection fraction and a significant decrease in LV end diastolic and systolic diameters (T. Wang, Wu, et al. 2009). These results suggest that this hydrogel could serve as an injectable biomaterial that prevents LV remodeling and dilation for the treatment of MI. Similarly, Wu et al developed a temperature-sensitive, aliphatic polyester hydrogel (HG) conjugated with VEGF and looked at its effects on cardiac recovery after MI (Wu et al. 2011). It was shown that the hydrogel gelled in ~10 minutes at physiological conditions, and that in vitro it was stable for up to 5 weeks, but completely degraded by 6 weeks when injected subcutaneously. Injection around the infarct site was done seven days after coronary artery ligation and rats were monitored up to five weeks post-injection (six weeks post-MI). Compared with outcomes in the PBS group, ventricular volumes, preload recruitable stroke work, and endsystolic elastance were all significantly preserved with HG-VEGF, and furthermore the VEGF conjugated hydrogel led to improved blood vessel density in the infarct area (Wu et al. 2011). Functionally, ejection fraction was ~60% higher, fractional shortening ~37% better with the HGVEGF group over the PBS injected group, and while the infarct thinned and dilated after PBS injection it was ~30% smaller and 50% thicker in hearts treated with HG-VEGF (all measured at 5 weeks post injection) (Wu et al. 2011). They did note that most of these effects could be seen with injection of hydrogel (no VEGF), or hydrogel with VEGF mixed in (not conjugated). However, none of the effects were as significant as those observed with their VEGF conjugated polyester hydrogel leading to the conclusion that conjugation provided sustained, localized VEGF function promoting regeneration (Wu et al. 2011). Furthermore, an important relationship on how hydrogel mechanical properties correlate their therapeutic outcomes has been brought up and studied by Ifkovits et al (Ifkovits et al. 2010). They compared two injectable methacrylated hyaluronic acid (MeHA) formulations that exhibit 27

similar degradation and tissue distribution upon injection but have differential moduli, using infarct only as a control group. The MeHA or control hydrogels were injected 30 min after the infarction and the result showed both treatments significantly increased the wall thickness in the apex (>200% increase) and basal (>40% increase) infarct regions compared with the control infarct (Ifkovits et al. 2010). Specifically, the treatment with the higher modulus (MeHA high) group had significantly smaller infarct area (approximately 16% smaller) compared with the control infarct group. Moreover, the normalized end-diastolic and end-systolic volumes were reduced (though not significantly) for the MeHA High group (1.7 and 1.9, respectively) compared to control (2.1 and 2.5, respectively) (Ifkovits et al. 2010). Functionally, real time 3D echocardiography was used to assess the LV dimension and cardiac function for each animal before and immediately after infarction, suggesting that MeHA group also tended to have a better cardiac output and higher ejection fraction (again however, not significant) than the lowmodulus (MeHA Low) and control infarct groups (Ifkovits et al. 2010). Some care must be taken in interpreting the results of bulking material strategies as recent work suggest passive wall support on its own does not prevent LV remodeling and preservation of cardiac function. Rane et al looked to decouple biomaterial effects from mechanical effects of such treatments by injecting bio-inert, non-degradable PEG hydrogel into a rat MI model nine days after infarction (Rane et al. 2011). While infarct wall thickness was significantly increased in the injected group verses saline injected control animals, there was no difference in cardiac function (ejection fraction, end diastolic and systolic volumes) between groups. Confirmation of the decoupling between material bioactivity and mechanical effects was done by comparing cellular response between groups and showing no difference, thus leading to the conclusion that benefits reported from other studies are likely due to differences in inflammatory and cellular response to the material and not due to the bulking effect of the material itself (Rane et al. 2011). For an injected hydrogel to act as a long term mechanical support, it would be assumed its presence long-term (slow or no degradation) is required. To investigate this, a non-degradable, in situ polymerizable PEG hydrogel was assessed for its effect on short and long term cardiac repair after MI. Chosen for its degradation properties (i.e. lack thereof) as well as its very high mechanical stiffness and inertness, gel components were injected into the infarcted region and polymerized in situ following permanent ligation of the left anterior descending artery in male Wistar rats (Dobner et al. 2009). Benefits were seen after 4 weeks with increased wall thickness 28

(100%) and modestly attenuated dilation (43%) compared to a saline treated control group. However, morphological improvements were only seen during the first 4 weeks as by 3 months the hydrogel injection group had regressed to a cardiac geometry and function similar to the control group (Dobner et al. 2009). The results suggest that while maintaining ventricle geometry using stiff biomaterials is beneficial in the short term, successful long term application will most probably require substantial further research toward developing a material with the necessary characteristics and a greater understanding of the degree of stress relief needed to halt pathological remodeling (Dobner et al. 2009). As is evident, complex synthetic polymer hydrogels are generally the choice for ventricle geometry preservation strategies due the high tunability of mechanical properties to match the needs of the failing LV. This does not exclude groups from using natural polymers as well, however. Dai et al investigated whether commercially available collagen injection could thicken the infarcted LV wall and improve function by preventing systolic bulging in Fischer rats (Dai et al. 2005). 100μL of collagen or saline was injected into the scar area of rats with week old MIs, and then cardiac morphology and function was monitored a further six weeks. The results showed significantly increased scar thickness in the collagen injected group (719±26μm) compared with the saline-treated group (440±34μm), and also that injected collagen remained present in the scar at six weeks. Furthermore, stoke volume was significantly larger in the collagen-treated group (163±8μL) than in the saline-treated group (129±6μL), and LV ejection fraction was greater in the collagen-treated group (48.4±1.8%) than in the saline-treated group (40.7±1.0%) (Dai et al. 2005). 2.5.5. Clinical application

The true implication of any of these injectable biomaterial strategies can only really be assessed in a clinical setting. If the hydrogels are inert like alginate, they could be regulated as devices, leading to a shorter time to reach the clinic. Two groups have developed alginate based hydrogels currently in Phase II clinical trials (NCT01226563, & NCT01311791). The calciumcross-linked alginate hydrogel developed by Leor et al. was first tested in a swine anterior MI model (Leor et al. 2009). Animals were monitored up to 60 days (56 days post-injection) and control animals (intracoronary saline injection) showed an increase in LV diastolic area by 44%, LV systolic area by 45%, and LV mass by 35%, whereas intracoronary hydrogel injection (2mL) prevented and even significantly reversed LV enlargement by greater than 100% compared to 29

control for all measures. Furthermore scar thickness was 53% greater with 2mL hydrogel injection compared to the control, and further proved that their developed biomaterial was feasible, safe, and effective (Leor et al. 2009). Dubbed IK-5001, the hydrogel has been approved for Phase II clinical trials and they are currently recruiting patients to test the safety and effectiveness of the device. Intracoronary injection of 4mL of IK-5001 into the blocked artery (after successful stent placement) will be done with the goal of prevention of ventricular remodeling and congestive heart failure when administered after a recent acute MI (NCT01226563). By contrast, the hydrogel developed by Lee et al is designed for direct injection into the infarcted wall for restoration of LV geometry. Called Algisyl-LVR™, it is another calciumcross-linked alginate biomaterial that gels in 3-4 minutes and achieves a material strength of 35kPa (Lee et al. 2012). An initial pilot study in 6 patients suffering from dilated cardiomyopathy demonstrated sustained improvements in LV size and function that were accompanied by statistically significant improvements in clinical status and quality of life 3 months postinjection, with no implant related complications (Lee et al. 2012). The approved Phase II clinical trial will evaluate the concept that direct mid LV intramyocardial injections of the biomaterial into the free wall of the failing LV of patients with dilated cardiomyopathy will reduce LV size, restore LV shape, lower LV wall stress and improve global LV function (NCT01311791). The goal is to show significantly improved peak maximum oxygen uptake 6 months after treatment between Algisyl-LVR™ treated patients versus those receiving medical management.

2.6. Chitosan & collagen Chitosan is a naturally occurring linear polysaccharide, composed of N-acetyl-D-glucosamine and D-glucosamine monomers linked by β(1-4) glycosidic bonds, as seen in Figure 2-2. It is derived from the deacetylation of chitin, the most abundant organic molecule and main component of the exoskeleton of crustaceans, molluscs, and insects. It is biocompatible, and its degradation products are nontoxic and non-immunogenic (Prasitsilp et al. 2000; Muzzarelli 1993). Its biodegradation products are lower molecular weight chitosans, chito-oligomers, and the monomers (N-acetyl-D-glucosamine, itself an anti-inflammatory drug, and D-glucosamine). Additionally, the degradation of chitosan through enzymatic hydrolysis by lysozyme can be controlled by modifying the degree of deacetylation (I. Y. Kim et al. 2008).

30

Figure 2-2: The structures of chitosan and chitin, and their proportions in 85% deacetylated chitosan

The presence of amino groups on the polysaccharide backbone of chitosan allows for conjugation with various bioactive molecules (Figure 2-2) (Shi et al. 2006; I. Y. Kim et al. 2008). Furthermore, the amino groups make chitosan poly-cationic, meaning it can adhere more easily to negatively-charged tissue (Chenite et al. 2000), and can be used for the controlled release of negatively-charged molecules (Fujita et al. 2004). However, chitosan suffers from mechanical weakness and instability, as well as an inability to maintain predefined shapes (Puppi et al. 2010). Chitosan is not a novel material but thanks to the diversity of potential uses it has been well studied and utilized in many biomedical applications including wound repair, tissue engineering, and drug delivery (as discussed above and reviewed extensively in Q Li et al. 1992; Majeti and Kumar 2000; IY Kim et al. 2008). Collagen I is a poly-anionic natural protein polymer and is a major constituent of the cardiac extracellular matrix. Currently, collagen based materials are one of the most common ECM materials used for culturing cells in vitro, and form the backbone of many hydrogels used in vivo (as described earlier) as it is the most characterized and abundant component of the ECM (Johnson, Lin, and Christman 2011). Native fibrous collagen can be processed into an injectable solubilized form, which can self-assemble under specified conditions into a hydrogel with nanoscale fibrous structure (Johnson, Lin, and Christman 2011). Collagen molecules assemble to form microfibrils, which, in turn, form fibrils to create collagen fibers. Collagen can be reconstituted to form hydrated gels that are similar to loose connective tissue in vivo, with cell behavior in three dimensional hydrated collagen gels being more typical of in vivo behavior than those grown on two-dimensional plastic culture surfaces. The material properties—material and gelation kinetics, stiffness, nanoscale fibrous architecture—and cell-matrix interactions of 31

collagen gels have been well characterized and can be manipulated by changes in temperature, ion concentration, phosphate content, and pH to specify the properties of the hydrogel. It is desirable as a biomaterial since it is biodegradable, has attachment sites for many cell types, and has a mild inflammatory response when implanted (Pulapura and Kohn 1992). Limitations in its use for cardiac tissue engineering, however, are its mechanical properties and suitability for culturing cardiomyocytes, as well as its rapid biodegradation (Puppi et al. 2010). Numerous methods for improving the mechanical properties of collagen gels have been explored including addition of chemical cross-linkers, but the effect of the by-products of these chemical reactions can face scrutiny from regulatory and clinical bodies (Nicodemus and Bryant 2008). In contrast, in the natural ECM the structural properties of collagen are known to be modified by glucosaminoglycans. The role they (glucosaminoglycans) play in modifying collagen fibril formation is by intertwining the microfibrils, forming thicker fibrils, and thus improving structural properties and increasing pore size. The structural characteristics of chitosan make it very similar to glucosaminoglycans, thus it was proposed and shown that addition of poly-cationic chitosan improves the mechanical stability and compressive strength of the highly poly-anionic collagen based gels through ionic interactions between the two components (Tan, Krishnaraj, and Desai 2001). Furthermore, the addition of chitosan to collagen gels has been shown to reduce the rate of degradation by collagenase (H. Chen et al. 2008). The interactions between collagen and chitosan in polyelectrolyte complexed hydrogels is illustrated in Figure A-1 (Appendix A). Collagen-chitosan hydrogel mixes have shown positive results by other groups as well. The work of Lu et al suggested a thermo-responsive chitosan based hydrogel increased vasculogenesis in rat infarcts when co-injected with ESCs in comparison to phosphate buffered saline (PBS) control (Lu et al. 2009). Based on this study, chitosan was added to an already established collagen hydrogel from another group, forming a chemically cross-linked collagenchitosan hydrogel aimed at improving recruitment, differentiation and survival of vascular progenitors (Suuronen et al. 2006; Deng et al. 2010). While Lu et al were focused primarily on CM regeneration (they were successful in improving ESC survival and differentiation), the work by Deng et al showed that addition of chitosan stabilized the collagen hydrogel, was better suited for the maturation of endothelial cells (in vitro), and promoted greater vascular growth and recruited more endothelial/angiogenic cells than collagen alone (Lu et al. 2009; Deng et al. 32

2010). Addition of the oligosaccharide sialyl LewisX (sLeX, which binds to the adhesion molecule L-selectin) to the hydrogel improved mobilization and recruitment of endogenous and transplanted circulating progenitor cells. Furthermore it enhanced neovascularization, and improved tissue perfusion in a rat hind-limb ischemia model (Suuronen et al. 2009).

2.7. QHREDGS peptide The angiopoietins are a family of growth factors that are known to bind to the receptor tyrosine kinase Tie2 acting as both agonists and context-dependent antagonists (S. Davis et al. 1996; Maisonpierre 1997). In endothelial cells, angiopoietin 1 (Ang1) was found to act as a survival factor by activating the phosphoinositide-3-kinase (PI3K)/Akt anti-apoptotic pathway (Hayes et al. 1999; I. Kim et al. 2000; Papapetropoulos 2000). It was later found that fibroblasts lacking the Tie2 receptor could still bind to Ang1 and that the resultant adhesion could be inhibited by blocking integrin receptors (Carlson et al. 2001). This led to the discovery that the αV integrin subunit, and more specifically the integrin receptor αVβ3, was responsible for angiopoietin-mediated attachment (independent of Tie2), which induced endothelial cell adhesion, migration, and activated characteristic integrin signal transduction pathways (Camenisch et al. 2002). It has also been reported that while neonatal rat cardiomyocytes (NCM) do not express the Tie2 receptor, they could still adhere to Ang1; adhesion to Ang1 conferred a significant survival advantage to NCMs; Akt and mitogen-activated protein kinase (MAPK) pathways were involved in the Ang1-mediated survival of CMs; and finally adhesion to Ang1 was integrin-dependent (Dallabrida et al. 2005). Further study identified the short sequence QHREDGS, conserved among mouse, rat, human, and other species, as the integrin-binding motif in the fibrinogen-like domain of Ang1, and it was demonstrated that the peptide QHREDGS was capable of supporting CM attachment and survival similar to full-length Ang1 (Figure 2-3) (Rask, Mihic, et al. 2010; Rask, Dallabrida, et al. 2010). While chitosan on its own does not promote cell attachment previous work has shown the ability of the peptide QHREDGS to promote CM attachment and growth, as well as survival and maturation when covalently attached to a photo-crosslinkable Az-chitosan hydrogel (Rask 2009; Rask, Mihic, et al. 2010; Rask, Dallabrida, et al. 2010). Although a well known integrin ligand, RGD, also promotes attachment of CMs, it prevents caspase-3 activation, it decreases the force of contraction of papillary muscle, and leads to pathological hypertrophy of CMs (Boateng et al. 2005;

33

Balasubramanian and Kuppuswamy 2003). The work previously described with Az-chitosan also showed the superiority of QHREDGS to an RGD based peptide, specifically RGDS.

Figure 2-3: Molecular structure of Ang-1 derived peptide QHREDGS

2.8. Peptide modified chitosan-collagen hydrogel for cardiac tissue engineering As it is known chitosan can interact with collagen to improve the mechanical stiffness of collagen based gels it is proposed that a mixture of the two components can form a hydrogel suitable for the culture of CMs in vitro as well as be mechanically stable enough for injection into the heart for treatment of acute MI. The QHREDGS peptide sequence has been demonstrated to promote the attachment and survival of CMs, suggesting it might have the therapeutic potential to help restore cardiac contractile function post MI. Moreover, since the beneficial effects of the QHREDGS peptide were found to be attachment dependent, we postulated that immobilizing the peptide onto a hydrogel would promote its pro-survival activity in addition to ensuring its sustained localization at the site of injection. Importantly, as a peptide, QHREDGS has an advantage over the protein-based therapies in that it is water-soluble, very stable, fully-synthetic with a precisely defined composition, it can be produced in a cost-effective and facile manner, and does not require a specific orientation to be functional. As discussed earlier, the presence of amino groups on the backbone of chitosan allows for its conjugation with the carboxyls of bioactive molecules. Conjugation can be achieved through many routes; however one of the most common methods for stable cross-linking of molecules is 34

the use of carbodiimide compounds as they provide the most popular and versatile method for labeling or cross-linking to carboxylic acids. The most readily available and commonly used water soluble carbodiimide is 1-ethyl-3-(-3-dimethylaminopropyl) carbodiimide hydrochloride (EDC) (Hayworth 2014). EDC reacts with carboxylic acid groups to form an active, unstable Oacylisourea intermediate that is easily displaced by nucleophilic attack from primary amino groups in the reaction mixture. The primary amine forms an amide bond with the original carboxyl group, and a soluble urea derivative is released as the EDC by-product. Failure of the intermediate to quickly react with an amine results in hydrolysis of the intermediate, regeneration of the carboxyls, and the release of an N-unsubstituted urea (Hayworth 2014). The efficiency of the reaction can be greatly improved with the addition of sulfo-N-hydroxysuccinimide (SulfoNHS). Sulfo-NHS replaces EDC during the reaction to form a Sulfo-NHS ester that is considerably more stable than the O-acylisourea intermediate, yet is still highly reactive with primary amines at physiologic pH. The Sulfo-NHS is regenerated upon successful conjugation of the carboxyl and amine, and it along with other reaction byproducts can subsequently be removed and the products purified through methods such as dialysis or chromatography. The modified chitosan-QHREDGS product is illustrated below (Figure 2-4), with further modifications used seen in Figure A-2 (Appendix A). A full schematic of the reaction and involved components can be seen in Figure A-3 (Appendix A).

Figure 2-4: QHREDGS peptide modified chitosan (QHG213H)

35

3. Hydrogel development, in vitro characterization & preliminary in vivo models The original work described in this chapter was published in (Reis et al. 2012), and is used with permission from Elsevier.

3.1. Materials & Methods All animal experimental procedures were approved by the Animal Care Committee of the Toronto General Research Institute and the University of Toronto Committee on Animal Care, according to the Guide for the Care and Use of Laboratory Animals. 3.1.1. Peptide modified chitosan (UP-G113-QHREDGS) 3.1.1.1.

Conjugation

QHREDGS peptide was conjugated to chitosan using EDC chemistry in a manner similar to that previously described for conjugation to Az-chitosan (Rask, Dallabrida, et al. 2010). To assess the dose response of the peptide two levels were used, a Low and a High concentration of the peptide in proportion to the amount of chitosan present. Briefly, chitosan (UP-G113, Novamatrix) was dissolved at 20 mg/mL in 0.9% normal saline and peptide at 10 mg/mL in phosphate buffered saline (PBS, Lonza). These were then mixed with EDC and Nhydroxysulfosuccinimide (S-NHS) dissolved in PBS to obtain final reaction solutions of 5 mg/mL chitosan and 0.5 mg/mL (Low) or 3 mg/mL (High) peptide with the ratio of [EDC]/[peptide] and [S-NHS]/[EDC] kept constant in the two reaction mixtures at 0.8 and 2.75, respectively. The reaction solution (1.5 mL) was then left on a vortex mixer (VWR) at 650 RPM for 3 hours. The solutions were diluted 4X with PBS and dialysed (using a dialysis membrane, Spectra/POR MWCO 3500, Spectrum Labs) against two 2 L changes of distilled water for ~24 hours. Sterilization of the dialyzed reaction mixture was done by passing the solution through a 0.2 µm syringe filter (Progene) and then the material was recovered through lyophilisation for approximately 48 hours. The recovered material was stored at -20°C until use. 3.1.1.2.

Assessing conjugation efficiency

36

Fluorescently labelled peptide, Fitc-Ahx-QHREDGS was used to assess the conjugation efficiency and true final concentration of peptide attached to chitosan post dialysis. Fitc-labelled QHREDGS (Biomatik) was substituted for regular peptide in the protocol above (3.1.1.1) and all steps were protected from light. As the molecular weight cut-off of the dialysis membrane is at most a tenth that of chitosan it can be safely assumed that all of the chitosan is retained and recovered, and the peptide present is only that which was successfully attached to the chitosan. Standards of the FITC-Ahx-QHREDGS in PBS was made ranging from 0.0005 to 0.01 mg/mL. For both the standards and the reaction solutions recovered post dialysis, the pH was adjusted to 7, as fluorescence is greatly affected by pH (Graber et al. 1986). The samples (Low & High) were diluted a further 1:10 and 1:100 and then run, with the standards, through a fluorometer (Spectra Max Geminin EM, Molecular Devices) at an excitation wavelength of 490 nm and emission of 520 nm, with all samples run in triplicate. The true final concentrations of peptide and conjugation efficiency were calculated by comparing the fluorescence of the samples to the standards, correcting for the dilution factor and for the volume recovered post dialysis. 3.1.2. Chitosan-collagen hydrogel 3.1.2.1.

Hydrogel formulation

Chitosan-collagen hydrogels were made through modification of the standard protocol for collagen gelation from BD Biosciences. Briefly, pure or peptide-modified chitosan was dissolved in sterile 0.9% normal saline and mixed with 10X PBS, collagen (BD Biosciences), and 1N NaOH on ice, in that order. The final solution consisted of 2.5 mg/mL chitosan, 2.5 mg/mL collagen, 10% of the final (solution) volume of 10X PBS (V10XPBS=0.1*VTotal), and 2.3% the volume of collagen added of 1N NaOH (VNaOH=0.023*VCollagen). The final hydrogel solution was mixed thoroughly and kept on ice until needed. An appropriate amount of hydrogel solution was pipetted onto a tissue culture plate bottom, or mixed with cells first, and allowed to gel for 30 minutes in a humidified 5% CO2 37°C incubator. To make 0:1 or 3:1 mass:mass (m:m) chitosan:collagen hydrogels the final solution was the same as above, but with 0 mg/mL or 7.5 mg/mL chitosan, respectively. 3.1.2.2.

Scanning electron microscopy

Samples were imaged using environmental scanning electron microscopy (Hitachi S-3400 N). Hydrogel samples (N=2/group) were allowed to gel for 30 min and individually placed into the 37

specimen chamber. A filter paper was used to gently remove the excess water from the sample. The chamber was closed and the temperature of the chamber was decreased to -20°C. The samples were imaged under variable pressure mode at 70 Pa and 15 kV. 3.1.2.3.

Hydrogel degradation

To assess the in vitro degradation of the hydrogel 2.0 mL Eppendorf tubes were weighed after having 5 holes punched in the caps with an 18G1/2 needle. 500 μL of 2.5 mg/mL 1:1 m:m chitosan-collagen hydrogel was then added to each tube and allowed to gel as previously described. Culture media (1 mL) was added to each tube and incubated for different amounts of time, with media changes (1 mL removed and replaced) every 48 hours. Gels were left for 1, 72, 120, or 240 hours (N=3/time point) when 1 mL of media was removed and the tubes and gel frozen at -80°C. All groups were lyophilized together for 72 hours, and then weighed again. Degradation was assessed by comparison of the lyophilized material weight to that of the 1 hour group. 3.1.2.4.

Rheological assessment of hydrogels

To assess differences in the mechanical stability of hydrogels with varying ratios of chitosan present, as well as the effect of conjugating QHREDGS peptide to chitosan, a rheological assessment measuring the loss and storage moduli of varying hydrogel compositions was performed. Rheology was performed using a TA Instruments AR1000 rheometer with a 6 cm acrylic cone and plate geometry and calibrated as per the manufacturer’s instructions prior to seeding approximately 1 mL of appropriate hydrogel (0:1, 1:1, 3:1, or 1:1 with conjugated QHREDGS peptide). Temperature was controlled using an integrated Pelletier plate and equilibrated to 4°C and prior to hydrogel seeding. To determine the Linear Visco-elastic Range (LVR), and show gel transition, a strain sweep was done at 4°C from 0-5% strain and frequency of 1Hz, and after allowing the hydrogel to equilibrate at 37°C for 30 minutes. Second, from the data collected from the strain sweeps a frequency sweep was performed on samples from all formulations at 37°C. The frequency sweeps were performed by seeding hydrogel onto the rheometer at 4°C, raising the temperature of the plate to 37°C and gelling for 30 minutes. A total of N=3 samples per group per analysis (strain or frequency sweep) were performed. 3.1.3. In vitro cell culture 3.1.3.1.

CM isolation

38

Cardiomyocytes used for in vitro work were isolated from neonatal (1-2 day old) SpragueDawley rat hearts as previously described (Radisic et al. 2003; Naito et al. 2006; Rask, Dallabrida, et al. 2010). Briefly, rats were euthanized; hearts were removed, quartered and then digested overnight with trypsin followed by five rounds of collagenase digests. Cells were then pre-plated for 1hr to enrich for CMs, and pre-plate supernatant was collected. CM number in the pre-plate supernatant was determined using a haemocytometer, and then cells were used for experiments. Quantification of the cell population before and after pre-plating to enhance the CM fraction has been performed previously, and it is reported after pre-plating the cell fractions are 63±2% CM, 33±3% cardiac fibroblasts, 3-4% smooth muscle cells, and 2-3% endothelial cells (Naito et al. 2006). A second pre-plate can be performed to improve the CM fraction to ~80% if desired (Iyer, Chui, and Radisic 2009). 3.1.3.2.

CM media

The cardiomyocyte culture media was comprised of 10% (v/v) FBS, 1% Hepes, 100 U/mL penicillin-streptomycin, 0.02 U/mL Insulin, 5 μg/mL vitamin C, and the remainder Dulbeco’s Modified Eagle medium (10% F.I.V. media). Cells were always re-suspended prior to encapsulation in a serum rich media, specifically 30% F.I.V (30% FBS). 3.1.3.3.

CM encapsulation & culture

Unless otherwise specified each sample prepared for in vitro experiments consisted of 5x105 CMs re-suspended in 2.5 μL 30% F.I.V. media and encapsulated in 12.5 μL of appropriate gel prepared as described in 3.1.2.1. The gel, media, and CM mix was pipetted onto the surface of a 24-well tissue culture plate and put in an incubator for 30 minutes to allow for gelation, after which samples were supplemented with 500 μL warm 10% F.I.V. media and cultured for 120 hours (five days). Media was aspirated, or collected, and 100% changed every 48hrs (two days). 3.1.4. In vivo studies 3.1.4.1.

Subcutaneous injection

CMs from neonatal Lewis rat pups were collected in the same manner as described in 3.1.3.1, spun down and 2 million CMs were re-suspended in 10μL 30% F.I.V. media first, followed by the addition of 100μL appropriate gel (Control (No Peptide), Low, and High) in Eppendorf tubes. The cell/gel suspension was then randomly injected subcutaneously (100 μL) with a 23G1/2 needle into the back of adult female Lewis rats (four injections per rat). The area where samples 39

were injected had been shaved and rough injection sites were marked with permanent marker in order to locate the injected material later. Samples (N=6/group) were left for seven days and then the animals were sacrificed, the skin surrounding the area where injections were made was removed and the material (hydrogel with cells) recovered and fixed with 10% formalin at 4°C for approximately 12hrs. Recovered samples were then transferred to PBS and sent to the Pathology Research Program (PRP) histology lab at University Health Network for paraffin-embedding and sectioning. 3.1.4.2.

Mouse MI model

To assess the potential of the developed hydrogel to be injected into the infracted heart, a mouse MI model was used. Adult male C57 Black-6 mice were subjected to a left ventricular anterior descending coronary artery ligation procedure (LAD procedure) to mimic severe MI with the help of Dr. A. Momen from the University Health Network. The mouse MI model has become important for studying myocardial regeneration following interventions such as cell based therapies, and studies have shown the LAD procedure gives rise to MIs of reproducible size and severity as well as an immediate surgical survival of 60%, and 2 week survival of ~83% (or 50% overall survival) (Kumar et al. 2005). Animals were divided into three groups: Sham, MI Only, and MI+Control gel. Briefly, animals were initially anaesthetized with 5% isoflurane and maintained with 2-2.5% isoflurane during surgery. Animals were intubated and ventilated using a Minivent ventilator (Harvard Apparatus, March-Hugstetten, Germany) at ~200 breaths/min. A left thoracotomy was performed, and the left coronary artery was ligated using a 7-0 silk suture (Ethicon) passed with a reverse cutting needle. Sham mice did not receive the LAD ligation and the chest was closed immediately. MI+Control gel mice had 50µL of 2.5mg/mL 1:1 m:m chitosan:collagen Control (No Peptide) gel injected with a 27G1/2 needle into the area immediately below the ligation suture. The hydrogel was warmed to ~37°C for ~10min prior to injection to allow gelling to initiate. The animals were removed from the ventilator and allowed to recover below a heat lamp for approximately 30 minutes. (N=3 animals for MI, N=3 for MI+Control Gel, N=1 for Sham). 3.1.5. Construct Characterization 3.1.5.1.

Gel compaction

40

To assess the chitosan-collagen ratio that was most appropriate for formation of beating cardiac constructs, CMs were encapsulated in the same manner as described in 3.1.3.3, with groups consisting of CMs in 0.5:1, 1:1, or 3:1 m:m chitosan-collagen hydrogel, with the concentration of collagen kept at 2.5 mg/mL. The cell/hydrogel mix was pipetted onto 12mm circular glass cover-slips located in the wells of a 12-well tissue culture plate. This was done to obtain uniform initial construct sizes and prevent cell/gel attachment to the bottom of the plates, which would confound results. After 144 hours (6 days) of incubation (with media changes every 48 hours) the media was removed and pictures of the hydrogel constructs taken (N=4/group). Adobe Photoshop CS3 was then used to measure the area of each hydrogel sample, and mean values compared between groups. 3.1.5.2.

Live/Dead staining

Live/dead staining was performed according to the manufacturer’s protocol (Invitrogen) using 5-carboxyfluorescein diacetate acetoxymethyl ester (CFDA, green for live cells) and propidium iodide (PI, red for dead cells) on N=2 constructs/group. CFDA is a non-fluorescent compound that can penetrate intact cell membranes wherein it is cleaved by non-specific esterases to release polar fluorescent fluorescein which is retained inside viable cells, whereas PI is a nucleic acid probe that is unable to cross intact cell membranes but penetrates the membranes of dead cells. Quantification was performed using ImageJ as previously described (Iyer, Chiu, and Radisic 2009). To show live cell distribution confocal images (z-stacks) were taken at the edge and center of cell/gel constructs. 3.1.5.3.

Functional testing

Two measurements were taken to assess the electrical function of encapsulated cardiomyocytes in chitosan-collagen hydrogels (No peptide, Low, and High peptide) cultured for five days as described in 3.1.3.3. Excitation threshold (ET) is the minimum voltage required to pace most of the cells in a construct and the maximum capture rate (MCR) is the maximum frequency at which the encapsulated CMs could be induced to beat simultaneously (Radisic et al. 2004). N=11 constructs/group were cultured as described above and then ET/MCR were measured using bi-phasic electrical pulses as previously reported (Rask, Mihic, et al. 2010). Briefly, media was removed, a pair of parallel carbon electrodes were placed surrounding the gels, and the system was immersed in warm Tyrode’s solution (Sigma, pH≈7.3). All 41

measurements were done in an environmental chamber at 37°C and ET was measured by increasing the voltage until ~90% of cells in the field of view were seen to be beating synchronously with square biphasic pulses of 2 ms and frequency of 1 Hz. MCR was measured by setting the voltage at 12 V increasing the frequency until most of the cells were no longer beating synchronously with the driving signal. Success rate was measured as the percentage of constructs for which ET/MCR measurements were able to be taken (i.e. constructs with significant beating portions) to the total number of constructs seeded for that group. Chi-square and Fisher’s Exact tests were performed to analyze the association between groups. 3.1.5.4.

XTT assay

CMs were cultured for five days as per 3.1.3.3 at which time media was removed and replaced with fresh 10% F.I.V. media and XTT assay solution as per the manufacturer’s protocol (XTT cell viability assay kit, Biotium, Inc.). Samples were incubated for a further three hours with the mix and then three 100 μL samples of mix were taken from each hydrogel sample and put in corresponding wells of a 96 well plate. The absorbance of wells was read using an absorbance meter (Apollo LB911, Berthold Technologies) at a wavelength of 450 nm and a reference of 620 nm. The triplicate readings per hydrogel were averaged to get a value for the sample, and then all sample values were averaged to get a value for the group. A total of N=6 per group were tested. In addition, N=2 per group of no-cell control gels were used. The mean absorbance reading of the no-cell controls was subtracted from all test groups. A standard curve was also made to show that XTT is indeed a good indicator of cell metabolism and indirectly cell number by encapsulating varying numbers of CMs in the same volumes of serum rich media and No-QHREDGS Control hydrogel and incubating for 1 or 120 hours. Relative absorbance is reported, with mean absorbance of samples relative to that of constructs with no CMs. Higher relative absorbance indicates a higher level of viability/metabolic activity. 3.1.5.5.

LDH assay

The LDH cytotoxicity assay (Cayman Chemical Company) was used to assess cell death at varying time points. It was performed as per the manufacturer’s protocol on media samples taken every 24 hours from samples prepared as per 3.1.3.3. Briefly, 50 μL media samples (in triplicate) were put into the wells of a 96 well plate and an equivalent volume of LDH reagent was added to 42

each well. The plate was placed on an orbital shaker for 30 minutes at 600 rpm protected from light and then the absorbance of each well read using an absorbance meter (Apollo LB911, Berthold Technologies) at a wavelength of 492 nm and reference of 620 nm. A total of N=6 gels/group plus 2 gels/group of no-cell controls were assessed. The absorbance of the no-cell controls was subtracted from the readings of CM samples to get a more accurate value. A cell standard was made in the same manner as for the XTT assay, and the LDH assay performed on media samples from the standard (N=4/group) as well as the LDH standard. 3.1.5.6.

PicoGreen DNA Assay

Constructs cultured as per 3.1.3.3 for four days were subsequently immersed in 1 mL of lysis buffer (0.2% Triton X-100, 200 mM Tris-HCl, 20 mM EDTA, pH 7.5) and transferred to 2 mL Eppendorf tubes. Several autoclaved 1 mm diameter silica beads were added to each vial, and the constructs were homogenized using a Mini-Beadbeater-16 (Biospec), via six cycles of ten seconds each. 500 μL of each homogenate was then transferred to separate wells of a 24-well plate, and the Picogreen assay was performed as directed by Invitrogen. Briefly, 500 μL of Picogreen reagent was added to each well, and incubated in the dark for 3-5 minutes, after which fluorescence was read using a fluorescence micro-plate reader (Spectra Max Gemini EM, Molecular Devices; excitation 480 nm, emission 520 nm). A separate assay, which was done on gels lysed immediately after encapsulation with a known quantity of cells, served as a control to determine the cell number in gels at the end of the experiment. 3.1.5.7.

Histological & immunofluorescent staining

Samples for immunohistochemical and immunofluorescent staining were fixed at 4°C with 10% neutral buffered formalin (Sigma, HT501129) for approximately 12 hrs, and then transferred to PBS. For paraffin embedding and sectioning, as well as hematoxylin and eosin (H&E), Mason’s trichrome, smooth muscle actin (SMA), Factor VIII (F8), and CD31 staining, the fixed samples were sent to the Pathology Research Program (PRP) at the University Health Network. Chitosan staining was performed using Cibacron Brilliant Red-3BA (CBR-3BA, Sigma Aldrich) and Weigert’s Iron Hematoxylin as described (Rossomacha, Hoemanni, and Shive 2004). Paraffin sections were also stained for cardiac troponin T (cTnT, mouse monoclonal antibody, Thermo Scientific MS-295-P). Deparaffinised slides were blocked using Normal Horse Serum (Gibco 16050-122) for 40 minutes and then primary antibody (cTnT) was 43

applied at a dilution factor of 1:200 for at least 12 hrs at 4°C. Secondary antibody solution (Alexafluor 488 donkey anti-mouse IgG (H+L), Invitrogen A21202) was applied at a dilution of 1:200 with DAPI (Sigma) nuclear stain at 1:200 dilution for ~40 min at room temperature. The slides were mounted using Fluoromount aqueous mounting medium (Sigma-Aldrich, F4680), covered with cover-slips (VWR, 22x50 mm) and imaged. Sections stained for CD3 and vimentin were deparaffinised and rehydrated with subsequent baths (3 of each) of 100% xylene, 100% EtOH, 95% EtOH, 75% EtOH, and distilled water for 3 minutes per change. Antigen retrieval was performed by microwaving slides immersed in TRISEDTA buffer for 5 min, and blocked using Dako Serum Free Protein Block (Dako X0909) for 30 min at room temperature. Primary antibody (poly rabbit anti-human CD3 diluted 1:200, Dako A0452 or mouse monoclonal anti-vimentin diluted 1:400, Sigma V6630) was applied to sections for 2 hrs at room temperature followed by secondary antibodies (goat anti-rabbit FITC IgG (H+L), Jackson ImmunoResearch 111-095-144 or Alexafluor 488 donkey anti-mouse IgG (H+L), Invitrogen A21202) at 1:400 dilution for 1 hr at room temperature. DAPI (Sigma) nuclear stain was applied at 1:1000 dilution for 10 minutes and finally slides were mounted using Dako Fluorescence Mounting Medium (Dako, S3023) and imaged. Quantification of area covered by positively stained cells in immunostained sections was performed using Adobe Photoshop CS3 and ImageJ updated from a previously reported method (Chiu and Radisic 2010). Up to five distinct images per section per sample were gathered. The colour profiles were saved such that the same colour selection profile could be applied automatically to every image to select for positive staining (removing subjective human staining selection). Unselected image areas were deleted and the image was converted to black and white and saved as a JPEG. Adjusted images were then opened in ImageJ, thresholded to remove potential background, and then analyzed to count particles. The resultant output was Area Fraction of black particles (positive stain) in relation to the whole image which gives a relative measure of positive immunostaining. The algorithm was tested on images from every group to ensure consistency, colour profile refined as necessary, and then applied to all images. The five images per section were averaged to obtain a % Positive Area measure for the sample, and then sample values averaged to obtain Mean±SD for each group. The number of F8+ vessels and their diameters were determined using ImageJ according to a previously published method (Chiu and Radisic 2010). Five randomly selected 300x300pixels 44

(131x131mm2) areas were cropped and magnified from each original F8 image. The numbers of F8+ vessels were manually counted from the cropped images and normalized to the area of the image to determine the vessel density. The diameters of the vessels were evaluated by drawing a line across the short axis of the vessel and then measuring the length of the line in ImageJ. 3.1.6. Statistical analysis

Statistical analysis was performed using SPSS Statistics 17.0 and GraphPad Prism 5.0. Differences between experimental groups were analyzed by using one-way ANOVA with posthoc Tukey tests or two-way ANOVA with Bonferroni post-tests. Categorical data was compared using Chi-square or Fisher’s Exact tests. PG11)which is not a practical measure for our application. Furthermore, during the elastic storage modulus (G1) never equilibrated during these experiments to give a meaningful measure of time to gelation. Mechanical properties of collagen-chitosan hydrogels were studied previously (L. Wang and Stegemann 2011; Tan, Krishnaraj, and Desai 2001). The initial increase in storage modulus and subsequent decrease with further addition of chitosan is likely due to the marked changes in the gel microstructure (Figure 3-1 D&E). The resultant structure of collagen only gels is a highly interconnected fibrous network with fibers of fairly uniform size. Addition of chitosan does not seem to affect fiber formation, however they interfere with the resultant network by sequestering fibers together and thus forming thicker “struts”, yet retaining a high level of interconnectivity in the structure (i.e. 1:1 gels). By bringing together many collagen fibers along a chitosan backbone, formation of these thicker struts results in an increase in mechanical integrity. Increasing the fraction of chitosan (i.e. 3:1 chitosan:collagen gels) causes a further increase in collagen fiber sequestration (thus thicker struts), however the degree of connectivity between resultant fibers is greatly diminished (see Appendix A, Figure A-1). Furthermore, Tan et al (2001) found previously that in the case of a human hematopoietic cell line, changes in cell growth were determined by chitosan:collagen ratio only, and that increasing or decreasing the total protein concentration (while keeping the ratio constant) had no effect on cell proliferation (Tan, Krishnaraj, and Desai 2001). Protein based gels are commonly tested at strain and frequency values on the order of 1% and 1 Hz respectively (L. Wang and Stegemann 2011; Ikeda and Foegeding 2003; Whittingstall 2003). As these values fall in the LVR for all chitosan:collagen formulations tested here they were chosen as a basis to compare the loss and storage moduli between gel formulations. The values are summarized in Table 3-1, below, and as expected confirm that a 1:1 collagen:chitosan gel formulation provides the most mechanically stable gel in terms of storage modulus, but also shows the greatest difference (separation) between loss and storage moduli (P