the development of elastomeric biodegradable
TRANSCRIPT
THE DEVELOPMENT OF ELASTOMERIC BIODEGRADABLE POLYURETHANE
SCAFFOLDS FOR CARDIAC TISSUE ENGINEERING
by
Ian C. Parrag
A thesis submitted in conformity with the requirements for the degree of Doctor of Philosophy
Graduate Department of Chemical Engineering and Applied Chemistry &
The Institute of Biomaterials and Biomedical Engineering
University of Toronto
© Copyright by Ian C. Parrag (2010)
ii
THE DEVELOPMENT OF ELASTOMERIC BIODEGRADABLE POLYURETHANE
SCAFFOLDS FOR CARDIAC TISSUE ENGINEERING
Ian C. Parrag
Doctor of Philosophy, 2010
Department of Chemical Engineering and Applied Chemistry & Institute of Biomaterials and
Biomedical Engineering, University of Toronto
Abstract In this work, a new polyurethane (PU) chain extender was developed to incorporate a
Glycine-Leucine (Gly-Leu) dipeptide, the cleavage site of several matrix metalloproteinases.
PUs were synthesized with either the Gly-Leu-based chain extender (Gly-Leu PU) or a
phenylalanine-based chain extender (Phe PU). Both PUs had high molecular weight averages
(Mw > 125,000 g/mol) and were phase segregated, semi-crystalline polymers (Tm ~ 42°C) with
a low soft segment glass transition temperature (Tg < -50°C). Uniaxial tensile testing of PU
films revealed that the polymers could withstand high ultimate tensile strengths (~ 8-13 MPa)
and were flexible with breaking strains of ~ 870-910% but the two PUs exhibited a significant
difference in mechanical properties.
The Phe and Gly-Leu PUs were electrospun into porous scaffolds for degradation and
cell-based studies. Fibrous Phe and Gly-Leu PU scaffolds were formed with randomly organized
fibers and an average fiber diameter of approximately 3.6 µm. In addition, the Phe PU was
electrospun into scaffolds of varying architecture to investigate how fiber alignment affects the
orientation response of cardiac cells. To achieve this, the Phe PU was electrospun into aligned
and unaligned scaffolds and the physical, thermal, and mechanical properties of the scaffolds
were investigated.
The degradation of the Phe and Gly-Leu PU scaffolds was investigated in the presence of
active MMP-1, active MMP-9, and a buffer solution over 28 days to test MMP-mediated and
passive hydrolysis of the PUs. Mass loss and structural assessment suggested that neither PU
experienced significant hydrolysis to observe degradation over the course of the experiment.
In cell-based studies, Phe and Gly-Leu PU scaffolds successfully supported a high
density of viable and adherent mouse embryonic fibroblasts (MEFs) out to at least 28 days.
Culturing murine embryonic stem cell-derived cardiomyocytes (mESCDCs) alone and with
iii
MEFs on aligned and unaligned Phe PU scaffolds revealed both architectures supported adherent
and functionally contractile cells. Importantly, fiber alignment and coculture with MEFs
improved the organization and differentiation of mESCDCs suggesting these two parameters are
important for developing engineered myocardial constructs using mESCDCs and PU scaffolds.
iv
Acknowledgments
There are numerous people who have contributed to my scientific and personal
experiences during the last several years that have made this work possible. It is my hope that I
have thanked the people who have helped me along the way because I would not have made it
far enough to be writing this now without their help. I would first like to thank my supervisor
Dr. Kimberly Woodhouse for her guidance and support that have been invaluable for my career
development. I am very appreciative of the flexibility she has given me in pursuing my research
and personal interests. Without the opportunity to work with her, I would not have been able to
do and accomplish many things that are important to me. I would also like to thank my
committee members, Dr. Paul Santerre and Dr. Peter Zandstra, for their time and guidance with
this research project along with access to their lab equipment and resources. The members of the
Santerre, Zandstra, and Edwards labs have been very helpful in the training and use of lab
equipment and technical advice. Celine Bauwens, Sylvia Niebruegge, Ting Yin, Kuihua Cai,
and Cheryl Washer have been particularly accommodating in this regard. I would also like to
acknowledge Eric Altman, Frank Gibbs, Dionne White, Gary Skarja, and Tim Burrows for
technical consultations and sample analysis. Funding from the Department of Chemical
Engineering and Applied Chemistry and OGSST was much appreciated. It has been a real
pleasure to work with all of the members of the Woodhouse group both in and out of the lab. I
would especially like to thank Joanna Fromstein, Patrick Blit, Cecilia Alperin-Dalley, Robin
Farmer, Lauren Flynn, Dave Laughren, and Elizabeth Srokowski for all their help with the work
in this thesis. Lastly, I would like to thank my family and friends for all their support and
encouragement that has gotten me through the difficult and enjoyable times that have come along
with this research project. Things never seem that bad when you’ve got good people in your life
and I am very appreciative of every single one of them.
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Table of Contents
Chapter 1: Introduction……………………….………………………………………………..1
1.0. Clinical Problem…...…………………………….………………………………………..1
1.1. Hypothesis………………………...…………………….…………………………………2
1.2. Research Objectives……………………………………………………………………….3
1.3 References…………………………………………………………………………………3
Chapter 2: Literature Review….……………………………………………………………….5
2.0. Introduction……….…..………………..………………………………………………….5
2.1 Heart Tissue..…………………………..…………………………………………………5
2.1.1 Myocardial Cells………………………………………………………………………….6
2.1.2. Extracellular Matrix Organization and Function………………………………………….8
2.1.3. Reparative Response of the Heart to Myocardial Infarctions……………………………..9
2.2. Matrix Metalloproteinases and their Role in Heart Remodeling and Disease…………...10
2.2.1. MMP Expression Following Myocardial Infarctions and in Heart Failure……………...11
2.2.2. Cleavage Sites of ECM Proteins, Peptides and Biomaterials by MMPs………………...12
2.3. Regenerative Approaches to Repair the Heart…………………………………………...14
2.3.1. Inducing Endogenous Mechanisms in Heart Repair……………………………………..14
2.3.2. Cellular Cardiomyoplasty………………………………………………………………..15
2.3.2.1. Fetal and Neonatal Cardiomyocytes..…………………………………………………..17
2.3.2.2. Embryonic Stem Cell-Derived Cardiomyocytes..……………………………………...18
2.3.2.2.1. Differentiation of Murine Embryonic Stem Cells into Cardiomyocytes……………...19
2.3.2.2.2. Large-Scale Production of a Pure Population of Embryonic Stem Cell-Derived
Cardiomyocytes……………………………………………………………………….21
2.3.2.2.3. Transplantation of Human and Murine ESC-Derived Cells into the Heart…………...23
2.3.3. Cardiac Tissue Engineering……………………………………………………………...25
2.3.3.1. Myocardial Tissue Engineering Using Biomaterials with Undefined Structures…..…..25
2.3.3.2. In Situ Cardiac Tissue Engineering..…………………………………………………....27
2.3.3.3. Myocardial Cell Sheets………..………………………………………………………..29
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2.4. Cardiac Tissue Engineering Using Pre-formed Three-Dimensional Scaffolds………….31
2.4.1. Biomaterials for Cardiac Tissue Engineering……………………………………………31
2.4.1.1. Natural Biomaterials..…………………………………………………………………..33
2.4.1.2. Synthetic Biomaterials..………………………………………………………………...33
2.4.1.2.1. Traditional Polymers for Tissue Engineering………………………………………...33
2.4.1.2.2. Elastomeric Biomaterials……………………………………………………………..34
2.4.2. Scaffold Fabrication Techniques………………………………………………………...36
2.4.3. Cells for Cardiac Tissue Engineering……………………………………………………38
2.4.4. Seeding and Cultivation Parameters for Cardiac Tissue Engineering…………………...39
2.5. Biodegradable Segmented Polyurethanes for Tissue Engineering………………………41
2.5.1. Chemistry and Properties of Degradable Polyurethanes………………………………...41
2.5.1.1. Segmented Polyurethane Synthesis…………………………………………………….43
2.5.1.2. Reactant Chemistry for Biodegradable Polyurethanes…………………………………44
2.5.2. Polyurethane Degradation……………………………………………………………….48
2.5.3. Enzyme-Degradable Polyurethanes……………………………………………………..51
2.6. Electrospinning for Tissue Engineering Scaffold Formation……………………………54
2.6.1. Principles and Parameters………………………………………………………………..54
2.6.2. Electrospun Scaffolds for Cardiac Tissue Engineering………………………………….56
2.7. References………………………………………………………………………………..58
Chapter 3: Synthesis and Characterization of Phe and Gly-Leu-containing Segmented
Polyurethanes……………………………………………………………………...84
3.0. Abstract…………………………………………………………………………………..84
3.1. Introduction………………………………………………………………………………85
3.2. Materials and Methods…………………………………………………………………...86
3.2.1. Dipeptide-based Chain Extender Synthesis……………………………………………...86
3.2.2. Gly-Leu-based Chain Extender Purification……………………………………………..88
3.2.3. Chain Extender Characterization………………………………………………………...89
3.2.4. Polyurethane Synthesis and Film Casting………………………………………………..90
3.2.5. Polyurethane Characterization…………………………………………………………...91
3.3. Results and Discussion…………………………………………………………………..92
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3.3.1. Chain Extender Synthesis and Purification………………………………………………92
3.3.1.1. Reaction Systems for Chain Extender Synthesis……………………………………….93
3.3.1.2. Synthesis of Chain Extenders using Gly-Ile or Gly-Leu Dipeptides…………………...96
3.3.1.3. Purification Strategies for the Gly-Leu-based Chain Extender………………………...99
3.3.2. Polyurethane Characterization ………………………………………...………………106
3.3.2.1. Molecular Weight Averages ………………………………………………………….106
3.3.2.2. Thermal Transitions and Phase Segregation…………………………………………..107
3.3.2.3. Chemical Composition………………………………………………………………...107
3.3.2.4. Mechanical Properties…………………………………………………………………109
3.3.2.5. Effect of Amino Acid and Dipeptide-based Chain Extenders on Polyurethane
Properties……………………………………………………………………………...110
3.4. Conclusions……………………………………………………………………………..112
3.5. References………………………………………………………………………………113
Chapter 4: Electrospinning Phe and Gly-Leu Polyurethanes……………………………..116
4.0. Abstract…………………………………………………………………………………116
4.1 Introduction……………………………………………………………………………..117
4.2. Materials and Methods………………………………………………………………….118
4.2.1. Electrospinning Phe and Gly-Leu Polyurethane Scaffolds……………………………..118
4.2.2. Scaffold Characterization……………………………………………………………….120
4.3. Results and Discussion…………………………………………………………………121
4.3.1. Electrospinning Polyurethane Scaffolds………………………………………………..121
4.3.1.1. Effect of PU Concentration on Scaffold Morphology………………………………...123
4.3.1.2. Molecular Weight Averages and Thermal Properties…………………………………129
4.3.1.3. Fiber Size in Electrospun PU Scaffolds for Soft Tissue Engineering………………...129
4.3.2. Aligned and Unaligned Phe PU Scaffolds……………………………………………...130
4.3.2.1. Scaffold Morphology………………………………………………………………….131
4.3.2.2. Molecular Weight Averages and Thermal Properties…………………………………134
4.3.2.3. Mechanical Properties…………………………………………………………………135
4.3.2.4. Electrospun PU Scaffolds for Cardiac Tissue Engineering…………………………...138
4.4. Conclusions……………………………………………………………………………..141
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4.5. References………………………………………………………………………………142
Chapter 5: Polyurethane Degradation by Matrix Metalloproteinases……………………146
5.0. Abstract…………………………………………………………………………………146
5.1. Introduction……………………………………………………………………………..146
5.2. Materials and Methods………………………………………………………………….147
5.2.1. Activation and Activity of MMPs………………………………………………………147
5.2.2. Degradation of Polyurethanes by MMPs……………………………………………….149
5.3. Results and Discussion…………………………………………………………………150
5.3.1. Activation of MMPs……………………………………………………………………150
5.3.2. Activity of MMPs after Incubation with Polyurethanes………………………………..153
5.3.3. Degradation of Polyurethanes by MMPs……………………………………………….155
5.4. Conclusions……………………………………………………………………………..166
5.5. References………………………………………………………………………………166
Chapter 6: Cell Response to Electrospun Polyurethane Scaffolds………………………...170
6.0. Abstract…………………………………………………………………………………170
6.1. Introduction……………………………………………………………………………..171
6.2. Materials and Methods………………………………………………………………….172
6.2.1. Mouse Embryonic Fibroblast Culture and Seeding onto Polyurethane Scaffolds……...172
6.2.2. Characterization of MEFs on Phe and Gly-Leu-containing Polyurethanes…………….172
6.2.3. Culture and Differentiation of Murine Embryonic Stem Cells…………………………174
6.2.4. Monitoring the Differentiation of Cardiomyocytes from mESCs……………………...175
6.2.5. Scaffold Preparation and Cell Seeding…………………………………………………176
6.2.6. Characterization of mESCDCs and MEFs on Aligned and Unaligned Polyurethane
Scaffolds………………………………………………………………………………..176
6.3. Results and Discussion…………………………………………………………………178
6.3.1. Viability of MEFs on Phe and Gly-Leu-containing Polyurethanes…………………….178
6.3.2. Differentiation of mESCs into Cardiomyocytes in Spinner Flasks…………………….181
6.3.3. Effect of Fiber Alignment and Coculture with MEFs on Response of
mESC-derived Cardiomyocytes..………………………………………………………187
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6.3.4. Aligned and Unaligned PU Scaffolds for Cardiac Tissue Engineering………………...200
6.4. Conclusions……………………………………………………………………………..203
6.5. References………………………………………………………………………………203
Chapter 7: Conclusions………………………………………………………………………209
7.0. Conclusions……….……………………………………………………………………209
7.1. Significant Contributions to Literature…………………………………………………214
7.2. Future Work…………………………………………………………………………….214
7.2.1. Polyurethane Design and Synthesis…………………………………………………….214
7.2.2. PU Scaffold Formation and Characterization…………………………………………..214
7.2.3. PU Degradation…………………………………………………………………………215
7.2.4. Cell-based Testing of PU Scaffolds…………………………………………………….215
7.3. References………………………………………………………………………………217
Appendix A: Supplementary Information for Dipeptide-based Chain Extender
Characterization………………………………………………………………220
A.1. C13 NMR Spectra of Reactants, Theoretical Predictions, and Raw Products………….220
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List of Figures
Figure 2.1. The structure of the myocardium…………………………………………….........7
Figure 2.2. The cardiac extracellular matrix…………………………………………………..8
Figure 2.3. Alterations in MMP and TIMP levels in human heart disease…………..............12
Figure 2.4. Illustration of microphase separation in segmented polyurethanes……………...42
Figure 2.5. Standard two-step segmented polyurethane reaction………………………….....44
Figure 2.6. Diisocyanates used to synthesize biodegradable PUs……………………………46
Figure 2.7. Polyols often used in biodegradable PU synthesis………………………………47
Figure 2.8. Model for environmental biodegradation of PUs………………………………..49
Figure 2.9. Schematic of electrospinning apparatus………………………………………….55
Figure 3.1. Chain extender reaction system setups…………………………………………..87
Figure 3.2. Synthesis scheme for Gly-Leu-based diester, diamine chain extender…..............88
Figure 3.3. Synthesis scheme for Gly-Leu PU…………………………………………….....91
Figure 3.4. Mass spectrum of raw Gly-Ile-CDM-PTSA product………………………….....94
Figure 3.5. Mass spectra of raw Gly-Ile-based chain extender products synthesized in different solvent systems.………………………………………………………...95
Figure 3.6. Mass spectrum of crude product from Gly-Leu-CDM-PTSA…………………...97
Figure 3.7. Mass spectra of Gly-Leu-based chain extender using different catalysts and diol linkers…………………………………………………………………….....98
Figure 3.8. HPLC separation of Gly-Leu-based diester product using analytical column and low pH aqueous mobile phase……………………………………………..100
Figure 3.9. HPLC separation of Gly-Leu-based diester product using analytical column and high pH aqueous mobile phase…………………………………………….101
Figure 3.10. Preparative column HPLC purification of chain extender using low and high pH aqueous mobile phases……………………………………………………..102
Figure 3.11. C13 NMR spectra of products collected from preparative column HPLC using the two developed methods of separation………………………………………104
Figure 3.12. FT-IR spectrum of purified Gly-Leu-based chain extender …………………...105
Figure 3.13. The chemical structure of the Phe and Gly-Leu-based chain extenders………..106
Figure 3.14. FT-IR analysis of Phe and Gly-Leu PUs……………………………………….108
Figure 3.15. Representative stress-strain curve for Phe and Gly-Leu PU films……………..109
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Figure 4.1. Illustration of electrospinning apparatus………………………………………..119
Figure 4.2. A comparison of electrospun Phe PU mats formed in the Rabolt laboratory and in our laboratory using conditions established in the Rabolt laboratory…...122
Figure 4.3. Comparison of Phe PU scaffolds formed before and after optimizing electrospinning parameters……………………………………………………..123
Figure 4.4. SEM images of Phe and Gly-Leu PU scaffolds electrospun from different concentrations……………..…………………………………………………...124
Figure 4.5. Fiber diameter distributions of the Phe and Gly-Leu PU scaffolds electrospun from varying concentrations……………………………………………………126
Figure 4.6. Comparison of structural features of the Phe and Gly-Leu PU scaffolds used for degradation and cell-based studies………………………………………….128
Figure 4.7. SEM images of aligned and unaligned Phe PU scaffolds………………………132
Figure 4.8. Characteristics of aligned and unaligned Phe PU scaffolds…………………….133
Figure 4.9. Representative stress-strain curves for aligned and unaligned PU scaffolds stretched in preferred and cross-preferred directions of orientation……………136
Figure 5.1. Activation of MMPs using APMA……………………………………………..151
Figure 5.2. Zymogram of MMP activation solutions……………………………………….152
Figure 5.3. Activity of MMPs after incubation with PU scaffolds…………………………154
Figure 5.4. Mass remaining of PU scaffolds over 28 day degradation study……………….156
Figure 5.5. SEM images of PU scaffolds after 28 day incubation period in various solutions………………………………………………………………………...157
Figure 5.6. Reaction scheme for enzyme activity assay and competitive substrate enzyme activity assay…………………………………………………………..161
Figure 5.7. Inhibition of FS-6 cleavage using the Gly-Leu dipeptide……………………....161
Figure 5.8. Water uptake by Phe and Gly-Leu PU scaffolds……………………………….164
Figure 6.1. Illustration of experimental details for cardiomyocyte production and cell seeding………………………………………………………………………….175
Figure 6.2. AlamarBlue® analysis of MEFs on PU scaffolds over 28 day period..………...179
Figure 6.3. Staining of MEFs on PU scaffolds and TCPS…………………...……………..180
Figure 6.4. Total cell number in spinner flasks during differentiation of mESCs into cardiomyocytes………………………………………………………………....184
Figure 6.5. EB characteristics during differentiation of mESCs into cardiomyocytes……..185
Figure 6.6. Flow cytometry of cells before and after differentiation in spinner flasks……..187
Figure 6.7. AlamarBlue® analysis of cell-seeded PU constructs of varying architecture and TCPS controls………………………………………………………….…..189
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Figure 6.8. Live/Dead® staining of cells on Phe PU scaffolds of varying architecture at day 18+6………………………………………………………………………..191
Figure 6.9. Immunostaining of cells on aligned and unaligned PU scaffolds………………192
Figure 6.10. Immunostaining of cardiac constructs with mESCDCs showing varying levels of differentiation……………………………..…………………………..193
Figure 6.11. Quantifying the alignment of cells on PU scaffolds in coculture constructs…...197
Figure 6.12. Gap junction staining of mESCDCs and MEFs in coculture on aligned and unaligned PU scaffolds…………………………………………………………199
Figure A.1. C13 NMR spectrum of Gly-Leu dipeptide………………………………………220
Figure A.2. C13 NMR spectrum of CDM……………………………………………………221
Figure A.3. Theoretical predictions of Gly-Leu-based diester chain extender using ACD i-Lab software…………………………………………………………….221
Figure A.4. C13 NMR spectrum of raw Gly-Leu-CDM-PTSA…………..………………….222
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List of Tables
Table 2.1. List of cell types considered for cardiac repair…………………………………..16
Table 2.2. Summary of biomaterials and their applications in cardiac tissue engineering.....32
Table 2.3. Summary of electrospinning parameters and effects on fiber morphology……...56
Table 3.1. Molecular weight averages for PUs containing Phe and Gly-Leu-based chain extenders………………………………………………………………………..107
Table 3.2. Thermal properties of the Phe and Gly-Leu PUs as determined by DSC………107
Table 3.3. Summary of mechanical properties of PU films………………………………..110
Table 4.1. GPC and DSC results of Phe and Gly-Leu PU films and scaffolds………….....129
Table 4.2. GPC and DSC results for Phe PU films and electrospun scaffolds of varying architecture……………………………………………………………………...134
Table 4.3. Summary of mechanical properties of aligned and unaligned PU scaffolds stretched in preferred and cross-preferred directions of orientation……………136
Table 4.4. Mechanical properties of films of investigated or potential synthetic biomaterials in cardiac tissue engineering……………………………………...141
Table 6.1. Assessment of cell shape and sarcomere formation of mESCDCs……………..194
Table 6.2. Assessment of mESCDC dimensions……………………………………...…...195
Table 6.3. Average angle of cell axis and orientation index……………………………….198
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List of Abbreviations
3-D three-dimensional
ACN acetonitrile
ANP atrial natriuretic peptide
APMA 4-aminophenylmercuric acetate
BMP bone morphogenic protein
BV blood vessels
CB cardiac body
CDM 1,4-cyclohexane dimethanol
cTnT cardiac isoform of troponin T
Cx-43 connexin-43
DAPI 4',6-diamidino-2-phenylindole
DCM dichloromethane
DMEM Dulbecco’s modified eagle’s medium
Dnp fluorescence-quenching group; 2,4-dinitrophenyl
DSC differential scanning calorimetry
E initial modulus; Young’s modulus; elasticity; stiffness
EB embryoid body
ECM extracellular matrix
EHT engineered heart tissue
ESC embryonic stem cell
ESI electrospray ionization
FACS fluorescent activated cell sorting
FBGC foreign body giant cell
FBS fetal bovine serum
FS-6 fluorogenic substrate for MMPs; Mca-Lys-Pro-Leu-Gly-Leu-Dpa-Ala-Arg-NH2
FT-IR Fourier transform infrared
G418 geneticin; a neomycin analog
G-CSF granulocyte colony stimulating factor
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GFP green fluorescent protein
Gly glycine
Gly-Leu PU segmented polyurethane composed of PCL of molecular weight 1250, LDI, and a
Gly-Leu-based chain extender
GPC gel permeation chromatography
hESC human embryonic stem cell
hESCDC human embryonic stem cell-derived cardiomyocyte
HPLC high performance liquid chromatography
HOCl hypochlorous acid
LDI lysine-based diisocyanate
Leu leucine
LIF leukemia inhibitory factor
Mca fluorescent molecule; (7-methoxycoumarin-4-yl)acetyl
MDM monocyte-derived macrophage
mESC mouse embryonic stem cell
mESCDC mouse embryonic stem cell-derived cardiomyocyte
MHC myosin heavy chain
MHC-neor transgene carrying neomycin resistance gene driven by α-myosin heavy chain
promoter
MI myocardial infarction
MLC-2v myosin light chain-2v
MMP matrix metalloproteinase
NMR nuclear magnetic resonance
ONOO- peroxynitrite
PBS phosphate buffered saline solution
PCL1250 polycaprolactone diol of molecular weight 1250 g/mol
PGA poly(glycolic acid)
PGS poly(glycerol sebacate)
Phe phenylalanine
Phe PU segmented polyurethane composed of PCL of molecular weight 1250, LDI, and a
Phe-based chain extender
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pGK-hygror transgene carrying hygromycin resistance gene driven by phosphoglycerate kinase
promoter
PIPAAm poly (N-isopropylacrylamide)
PLA poly(lactic acid)
PLGA poly(lactic-co-glycolic acid)
PMN neutrophils; polymorphonucleocytes
PTSA p-toluene sulfonic acid
PU polyurethane
SDS sodium dodecyl sulfate
TCPS tissue culture polystyrene
TFA trifluoroacetic acid
Tg glass transition temperature
TGF transforming growth factor
TIPS thermally induced phase separation
Tm melting temperature
1
Chapter 1: Introduction
1.0 Clinical Problem Heart disease is one of the leading causes of disability and death in industrialized nations.
In the most recent study in Canada in 1998 (with updates in 2004), it was found that
cardiovascular diseases affect a quarter of the Canadian population accounting for more than a
third of the deaths and placing an estimated $18 billion burden on the Canadian economy [1, 2].
Topping the list of cardiovascular diseases was coronary heart disease, which leads to ischemic
heart disease, acute myocardial infarctions and congestive heart failure. Similarly, the
prevalence of cardiovascular diseases in the United States in 2005 was 80.7 million, or
approximately 37% of the population. These cases cost the U.S. health care system $448.5
billion and resulted in 869,700 deaths (36.3% of all deaths) [3]. Furthermore, 8.1 million
individuals in the U.S. suffer from the debilitating affects of a myocardial infarction with more
than 920,000 new or recurring cases and 156,800 fatalities in 2004. Interestingly, successes in
treating myocardial infarctions and other cardiac diseases have allowed individuals with
damaged hearts to live longer, but is leading to an increase in the prevalence of congestive heart
failure [2]. In the U.S. alone, 5.3 million people suffer from congestive heart failure with
284,400 deaths in 2004. As a consequence, these studies indicate the huge health care burden of
heart disease and identify the need for effective treatments to combat it.
The heart has a limited capacity to regenerate on its own. Cardiomyocytes that are lost
due to a myocardial infarction (MI), if not fatal, are replaced by the formation of scar tissue, an
adaptive response leading to the loss of contractile function [4]. Subsequent remodeling events
occur in the heart to compensate for this loss of contractile function in an attempt to maintain
cardiac output. Some of these events include changes in cell type, extracellular matrix
composition and organization, ventricular size and architecture, neurohormonal signaling, gene
and protein expression, and paracrine signaling, to name but a few [5]. In the short term, these
remodeling events attempt to maintain cardiac performance but inevitably become destructive to
the heart causing congestive heart failure and ultimately death.
There currently exists several treatment options following a MI and in congestive heart
failure. Pharmacological agents, such as thrombolytic agents, antithrombotics, nitrates, β-
2
blockers, Ca+2 channel blockers, angiotensin converting enzyme inhibitors, statins, and
adrenoceptor antagonists are typically used to increase blood flow, limit the ventricular
remodeling events, and increase cardiac output [6]. Although this therapy may be effective in
temporarily warding off heart failure, it is generally used to manage patients and offers little in
the way of treating the root of the condition. A second form of treatment employs the use of
mechanical devices, such as the left ventricular assist device. This treatment option has
traditionally been used as a bridge-to-transplantation, but has gained wider use as a destination
therapy and as a bridge-to-recovery option [7]. By reducing the workload of the injured heart,
LVAD therapy allows reverse remodeling to occur, whereby the destructive changes occurring in
the heart during remodeling are reversed [8]. This is an exciting new treatment option with some
patients undergoing sufficient recovery for the mechanical device to be removed, but the number
of patients eligible for this therapy remains low. A third treatment option, which remains the
gold standard because the recipient often regains full cardiac function, is heart transplantation.
This option, however, is limited by the lack of suitable donors and has motivated the field of
regenerative medicine to find alternatives that repair, replace, or augment the heart to restore
cardiac functionality.
There are many promising new approaches that are currently being investigated to
regenerate injured myocardial tissue and help fight heart disease. Some of these approaches
include pharmacological strategies, protein and peptide-based methods, gene therapy, cell-based
techniques, and tissue engineering [9]. Cardiac tissue engineering, in particular, offers the
advantage of combining several of these beneficial regenerative techniques along with novel
biomaterials and holds tremendous potential in the treatment of heart disease. As research in
cardiac tissue engineering continues to move forward, so to does the potential of easing the
enormous social and economic burden of this disease.
1.1 Hypothesis The project hypothesis is defined in two parts: 1) glycine-leucine (Gly-Leu) containing
biodegradable segmented polyurethanes can act as temporary scaffolds that support cells; and 2)
fiber alignment within polyurethane scaffolds influences the orientation response of murine
embryonic stem cell-derived cardiomyocytes and mouse embryonic fibroblasts seeded on the
constructs.
3
1.2 Research Objectives 1) Synthesize and characterize a family of biodegradable, segmented polyurethanes using a
Gly-Leu-based diester chain extender, lysine-based diisocyanate, and polycaprolactone
diol
2) Develop and characterize porous, three-dimensional biodegradable polyurethane
scaffolds by electrospinning and investigate the effects electrospinning has on scaffold
and polymer properties
3) Evaluate the in vitro degradation of amino acid and dipeptide-containing polyurethane
scaffolds in the presence of matrix metalloproteinase-1 and matrix metalloproteinase-9
4) Characterize the in vitro cellular response of cells seeded on polyurethane scaffolds
a) Assess the viability of mouse embryonic fibroblasts seeded on Phe and Gly-Leu-
containing polyurethane scaffolds
b) Characterize the response of murine embryonic stem cell-derived cardiomyocytes
and mouse embryonic fibroblasts on aligned and unaligned Phe-containing
polyurethane scaffolds
1.3 References 1. Heart and Stroke Foundation of Canada, Statistics and Background Information -
Incidence of Cardiovascular Diseases. 1998.
2. Heart and Stroke Foundation of Canada, Statistics. 2008.
3. American Heart Association, Heart disease and stroke statistics - 2008 update. American Heart Association, 2008.
4. Kumar, V., R.S. Cotran, and S.L. Robbins, Basic Pathology. 7th ed. 2003, Philadelphia: Saunders. xii, 873.
5. Swynghedauw, B., Molecular mechanisms of myocardial remodeling. Physiological Reviews, 1999. 79(1): p. 215-262.
6. Gelfand, E.V. and C.P. Cannon, Myocardial infarction: contemporary management strategies. Journal Of Internal Medicine, 2007. 262(1): p. 59-77.
7. Deng, M.C., L.B. Edwards, M.I. Hertz, A.W. Rowe, B.M. Keck, R. Kormos, D.C. Naftel, J.K. Kirklin, and D.O. Taylor, Mechanical circulatory support device database of the International Society for Heart and Lung Transplantation: Third Annual Report - 2005. Journal Of Heart And Lung Transplantation, 2005. 24(9): p. 1182-1187.
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8. Burkhoff, D., S. Klotz, and D.M. Mancini, LVAD-Induced reverse remodeling: Basic and clinical implications for myocardial recovery. Journal Of Cardiac Failure, 2006. 12(3): p. 227-239.
9. Puceat, M., Pharmacological approaches to regenerative strategies for the treatment of cardiovascular diseases. Current Opinion In Pharmacology, 2008. 8(2): p. 189-192.
5
Chapter 2: Literature Review
2.0 Introduction The high prevalence and economic burden of myocardial infarctions, congestive heart
failure, and other heart diseases has motivated researchers and clinicians to develop new
strategies to treat patients. Tissue engineering is a field that seeks the development of tissue
constructs to repair, replace, or augment damaged or diseased tissues. This field has already had
some clinical successes that demonstrate how tissue engineering may revolutionize the way
clinicians approach disease management and therapy [1-3]. The first tissue engineered trachea
transplant, for example, was recently performed using a decellularized human donor trachea
combined with the patient’s epithelial and mesenchymal stem cell-derived chondrocytes [1].
This novel procedure not only prevented the need to remove the diseased lung, the conventional
treatment choice, but also eliminated the need for immunosuppression therapy and drastically
improved the quality of life compared to the pre-operation condition and lung-resection option.
The heart is a complex organ composed of many critical components that give rise to its
unique function but also renders it susceptible to various injuries and diseases. Cardiovascular
tissue engineering on a whole explores tissue substitutes for the various components of the heart,
such as blood vessels, heart valves, and cardiac muscle. Advances in cardiovascular tissue
engineering have been made for each of the different components of the heart, but the
development of fully functional cardiac muscle remains one of the most challenging aspects of
this field.
2.1 Heart Tissue The heart is a muscular organ responsible for circulating blood throughout the body. It is
composed of four muscular chambers, the right and left atria, which pump blood to the
ventricles, and the right and left ventricles, which pump blood to the pulmonary and systemic
circuits respectively. Critical to this pumping function is the heart wall. The heart wall is
composed of three distinct layers, the endocardium, myocardium, and epicardium, respectively
located from the lumen of each cardiac chamber out, all surrounded by the pericardium [4].
While each layer plays a critical role for normal cardiac function, the myocardium is the
contractile portion that generates the necessary forces to pump blood to the body and constitutes
6
the bulk of the heart wall. The myocardium consists of multiple interlocking layers of cardiac
muscle tissue and the associated blood vessels, connective tissue, and nerves (Figure 2.1). Cells
within each layer of cardiac muscle tissue are anisotropically organized parallel to each other and
each layer is subsequently oriented at different angles depending on chamber type and location
within each chamber [4]. Due to the high energy requirements associated with contraction, the
myocardium is a highly vascularized structure [4]. The high demand for oxygen within this
muscular layer, however, renders it susceptible to ischemic injury. Disruption of the normal
tissue composition and organization of this portion of the heart wall is observed in many diseases
leading to a loss of contractile function. As a result, regenerative medicine techniques target the
myocardium in an attempt to restore contractility to the tissue. Understanding the cellular
components and tissue organization in the myocardium is therefore a requisite for the design of
engineered myocardial constructs.
2.1.1 Myocardial Cells The myocardium is composed of several cell types including vascular endothelial cells,
vascular smooth muscle cells, fibroblasts, neurons, and cardiomyocytes [5]. Cardiomyocytes are
the contractile cells taking up the bulk of the space in the myocardium. Mature adult ventricular
cardiomyocytes are rod-shaped, typically 10-30 µm in diameter and 80-150 µm in length [5], and
contain a high number of mitochondria and myoglobin to meet the energy requirements of
contraction [4]. Cardiomyocytes are composed primarily of bundles of myofibrils. Myofibrils
consist of a long repeated chain of sarcomeres, the basic contractile unit that give the cells a
striated appearance, composed of actin, myosin, tropomyosin, the troponin complex, and other
associated proteins [6]. In a resting state, the troponin complex and tropomyosin prevent myosin
from interacting with actin filaments. In response to a propagating action potential, the
excitation-contraction coupling mechanism causes an increase in intracellular Ca2+
concentration, the removal of the tropomyosin protein barrier, and, in the presence of ATP,
allows myosin to bind to actin leading to sarcomere shortening [6]. The excitation-contraction
coupling mechanism is made possible by the unique plasma membrane within these cells, the
sarcolemma, along with the transverse tubular system, the sarcoplasmic reticulum, and numerous
protein pumps, ion channels, and regulatory proteins. All working in a coordinated fashion, the
action potential, initiated independently of the nervous system, triggers this complex mechanism
ultimately leading to cardiomyocyte contraction.
7
Individual cardiomyocytes contracting on their own, however, does little in generating
the required forces to pump blood to the body. A coordinated effort is required and as such,
cardiomyocytes are organized into multiple interlocking layers connected to neighboring cells at
intercalated discs unique to cardiac muscle (Figure 2.1) [4]. At intercalated discs, cells are
electromechanically coupled by desmosomes, fascia adherens junctions, and gap junctions [7].
Because of the mechanical, chemical, and electrical connections between cardiomyocytes, the
cardiac tissue acts as a functional syncytium providing synchronous contraction and effective
force production to pump blood from the heart chambers.
Figure 2.1: The structure of the myocardium: a) histology image showing multiple interlocking layers of
cardiomyocytes with arrows indicating intercalated discs. b) Schematic identifying organization of bundles of cardiomyocytes, fibroblasts, blood vessels (BV), and extracellular matrix. Images used with permission from Dr.
Caceci [8].
Cardiac fibroblasts are the most numerous cells in the myocardium, and they play a
pivotal role in regulating tissue organization and function [9]. Fibroblasts are organized adjacent
to groups of myocytes where they interact with other fibroblasts, myocytes, and extracellular
matrix (ECM) macromolecules (Figure 2.1b) [10]. Cardiac fibroblasts are electrically connected
to adjacent fibroblasts and cardiac myocytes via gap junctions that aid in signaling between cells
[10]. The predominant role of cardiac fibroblasts, however, is to regulate the structure and
function of the ECM through deposition of its constituents and secretion of enzymes that degrade
them. The extracellular matrix acts as a mechanical support for tissues and transmits information
from the extracellular environment to regulate cell shape and function. Fibroblasts synthesize
and deposit the majority of the cardiac ECM, especially fibrillar collagen types I and III, elastin,
the proteoglycans laminin and fibronectin, and glycosaminoglycans [5, 11, 12]. ECM
remodeling and turnover is carried out by matrix metalloproteinases secreted by the fibroblasts in
8
both physiological and pathological states. Several growth factors, cytokines, and other
bioactive molecules are also produced by cardiac fibroblasts highlighting their regulatory role in
the heart [9]. In light of their important function in the myocardium, there is an increasing body
of evidence suggesting the critical role of cardiac fibroblasts and other non-myocytes in the
development of engineered cardiac tissue. This will be discussed in further detail in section
2.4.3.
2.1.2 Extracellular Matrix Organization and Function The myocardial extracellular matrix is made up of a fibrillar collagen network, basement
membranes, elastic fibers, proteoglycans, glycosaminoglycans and a variety of bioactive
signaling molecules (Figure 2.2a) [13]. This ECM is subdivided into three groups: the
endomysium, which surrounds individual myocardial cells; the perimysium, which surrounds
groups of myocytes; and the epimysium surrounding the entire muscle [14]. The specific
organization of the ECM layers aid in proper function of the tissue and relay important signaling
cues to the cardiac cells during normal physiology and disease.
Figure 2.2: Cardiac extracellular matrix: a) components and b) organization of endomysium and perimysium. Images used with permission from MacKenna et al. [15] and Goldsmith and Borg [16] respectively.
The endomysium contains the basal lamina, encompassing individual cardiomyocytes,
and fibrillar collagens that form lateral connections between cells (Figure 2.2b). The function of
the endomysium is to support and align myocytes, aid in cell attachment, bring cardiomyocytes
together, and keep blood vessels close to cells for short diffusion distances of nutrients and
9
oxygen [17]. The perimysium is composed of fibrillar collagens, types I and III, in a weave that
connects the basal lamina of the endomysium to the large collagen fibers of the epimysium. The
thick collagen fibers of the epimysium are organized parallel to myofibrils protecting sarcomeres
from overstretch during relaxation [14]. In addition, this parallel organization allows forces to be
transmitted across the tissue layer during contraction to pump blood and aids in tissue elasticity
by pulling back on cardiomyocytes during relaxation. The unique organization of the
endomysium, perimysium, and epimysium, therefore, imparts mechanical integrity to the
myocardium necessary for the dynamic cardiac cycle.
Aside from the structural and functional role of the ECM during contraction and tissue
organization, the ECM also plays a critical role in transmitting signals to cardiac cells during
myocardial development, normal physiology, and in disease. The ECM, for example, provides
micro-structural cues to differentiating cardiomyocytes that regulate sarcomere self-assembly
and guide myofibrillogenesis [18, 19] and influences the rate of maturation of neonatal
cardiomyocytes [20]. Importantly, much of the regulatory information conveyed to cardiac cells
by the ECM is transmitted in the form of physical forces [21]. Cell attachment to the ECM is
primarily carried out by the transmembrane glycoprotein receptors, integrins, present at the cell
surface. Cardiac cells use integrins as mechanotransducers to sense mechanical stimuli within
the tissue leading to intracellular signaling and therefore a cellular response to stresses associated
with normal physiology and in pathological overload [21]. Mechanical forces can help maintain
normal cell shape and an oriented myofibrillar architecture, alter ECM production, gene
expression, cell size, phenotype, and expression and release of paracrine factors, increase
sensitivity to other signaling molecules, and upregulate cell-cell contacts important for the
electrical and mechanical properties of the tissue [15, 22-28]. The response of different cardiac
cells to mechanical forces is very complex and depends on the specific cell type and physical
state of the tissue. Taken together, the ECM is far more than a passive component of the
myocardium but rather an active structural, functional and regulatory component of this tissue.
2.1.3 Reparative Response of the Heart to Myocardial Infarctions Cardiac tissue has a high demand for oxygen due to the high energy consumption
associated with muscle contraction. To ensure that active cardiomyocytes obtain a sufficient
supply of oxygen required to maintain aerobic respiration, the cells are in close proximity to
blood vessels of the coronary arteries. Reduced blood flow in the coronary arteries renders the
10
heart muscle susceptible to ischemic injury. Coronary artery disease refers to degenerative
changes in the coronary circulatory supply resulting in a reduction in the blood flow to the tissue
[29]. Sudden occlusion of the coronary arteries, or a myocardial infarction (MI), occurs in
severe cases of coronary artery disease leading to myocardial necrosis [29].
Several phases characterize the heart’s wound repair process after a myocardial
infarction: cardiomyocyte death, acute inflammation, formation of granulation tissue, ventricular
remodeling, and the formation of organized collagen-rich scar tissue (reparative fibrosis) [30].
In reparative fibrosis, fibroblasts and myofibroblasts, both resident and recruited to the infarct
region, synthesize and deposit ECM proteins including collagen types I, III, and V [28]. While
ECM constituents are being produced, proteases are continuously degrading the ECM to allow
cell migration and remodeling to take place [30]. The end result of the reparative fibrosis is an
adaptive response that maintains the structural integrity of the ventricle, but replaces the injured
myocardial tissue with a dynamic non-contractile scar tissue [31]. Reactive fibrosis, which
occurs in the absence of cell loss around the insulted region, occurs alongside the reparative
fibrosis leading to enhanced myocardial stiffening, arrhythmias, and reduced systolic function
[28, 31]. The hearts response to myocardial insult, therefore, is a reparative one characterized by
a loss of contractile function and myocardial fibrosis. The initial myocardial injury and
secondary effects of the hearts reparative process leads to disruption of the normal cellular and
extracellular composition and organization and the progression to heart failure.
2.2 Matrix Metalloproteinases and their Role in Heart Remodeling and Disease
The matrix metalloproteinases (MMPs) are a family of zinc-binding endoproteinases that
are the driving force behind myocardial matrix remodeling. All MMPs share several functional
features; they degrade ECM components, are activated when zinc is removed from the active
site, need calcium for stability, function at neutral pH, and are inhibited by specific tissue
inhibitors of metalloproteinases (TIMPs) in a 1:1 stoichiometric ratio [32]. MMPs are localized
in the cardiac interstitial space as latent pro-enzymes requiring activation by autoproteolysis, the
serine protease plasmin, oxidized glutathione, or other activated MMPs [33, 34]. Several MMPs
and TIMPs have been identified in the heart and help to maintain normal ECM turnover. In the
developing mouse heart, Nutell et al. [35] found that different levels of MMP-2, MMP-3, MMP-
8, MMP-9, MMP-11, MMP-12, MMP-13, MMP-15, MMP-19, MMP-23, MMP-24, and MMP-
11
28 along with TIMP-1, TIMP-2, TIMP-3, and TIMP-4 are all expressed. Similarly, MMP-1,
MMP-2, MMP-3, MMP-9, MMP-13, and MMP-14 have all been found in human myocardium
[36], and it is likely that others are also expressed. The interplay between these different
proteases and inhibitors creates a balance that contributes to normal myocardial ECM structure
and function. Cardiac pathologies, however, cause an imbalance in these enzymes and play a
role in myocardial collagen accumulation, collagen fibril disruption, myocyte loss, and altered
spatial orientation of cells and intracellular components.
2.2.1 MMP Expression Following Myocardial Infarctions and in Heart Failure
Following a myocardial infarction and in the progression of heart failure, myocardial
fibrosis and remodeling occur due to an imbalance in ECM production, MMP activity, and TIMP
expression [37]. This dysregulation in ECM turnover is a response of cardiac and inflammatory
cells triggered by many different factors including various inflammatory cytokines, growth
factors, and mechanical stresses associated with myocardial injury and pressure overload [37-
39]. Although the exact expression profiles of MMPs and TIMPs depends on the cause, severity,
and stage of heart disease (Figure 2.3), significant increases in MMP-1, MMP-2, MMP-9, MMP-
13, and MMP-14 and reduced levels of TIMP-1, TIMP-3, and TIMP-4 have been observed in
human patients with heart disease [37]. In a study by Webb et al. [40], temporal profiling of
various plasma MMP and TIMP levels was performed on patients following a MI demonstrating
the dynamic changes in MMP and TIMP expression patterns over time. Although in this study
an elevation of MMP-9 levels was linked to left ventricular dilation and adverse myocardial
remodeling months after the initial insult, the exact contribution of the different MMPs and
TIMPs expressed in the progression of heart failure is difficult to determine. Genetic mouse
models, however, have revealed important insight into the role of some of these MMPs and
TIMPs in heart disease. Kim et al. [41] constitutively expressed MMP-1 in the heart and found
compensatory myocyte hypertrophy at 6 months and a loss of cardiac interstitial collagen
concurrent with a marked deterioration of systolic and diastolic function at 12 months. This
study directly demonstrated that disruption of the extracellular matrix in the heart reproduces the
changes observed in the progression of heart failure. Similarly, the targeted deletion of MMP-2
[42] and MMP-9 [43] in knockout mice after an induced myocardial infarction attenuated left
ventricular dilation and ventricular remodeling and improved cardiac function compared to wild
12
type controls. These studies clearly implicate the role of MMP-1, MMP-2, and MMP-9 in
adverse myocardial remodeling in the progression to heart failure following a MI and are being
investigated as potential targets for pharmaceutical intervention [38].
Figure 2.3: Alterations in MMP and TIMP levels in human heart disease. Italic lower case letters depict mRNA
levels, capital letters indicate protein levels, ↑, ↓ and ↔ represent increase, decrease, and no change, respectively. * denotes circulating plasma levels. Image used with permission from Kassiri and Khokha [37].
2.2.2 Cleavage Sites of ECM Proteins, Peptides and Biomaterials by MMPs A priori knowledge of the expression profiles and key proteases involved in ECM
remodeling following a myocardial infarction and in heart failure not only identifies key targets
for therapeutic intervention but also allows the development of techniques that exploit the
presence of those enzymes to achieve a particular goal. In tissue engineering, specific sequences
have been incorporated into biomaterial scaffolds that make them susceptible to degradation by
MMPs expressed in various events, such as ECM remodeling, cell migration, angiogenesis, and
wound healing. Biological materials derived from the ECM inherently have these sequences
contained within them and therefore may naturally be degraded by the MMPs. Synthetic
materials, however, may also be developed to incorporate specific MMP-sensitive sequences
conferring unique biological function to these synthetic polymers. When designing novel
enzyme-degradable biomaterials for cardiac tissue engineering, MMP-1, MMP-2, and MMP-9
are rational targets due to their role in heart disease.
After being activated, MMP-1 functions by cleaving various collagens at specific sites in
the native triple helical structure. MMP-1 cleaves collagen type I at Gly775-Ile776 in the α1(I)
13
chain and Gly775-Leu776 in the α2(I) chain, collagen type II at Gly775-Leu776 in the α1(II) chain,
and collagen type III at Gly775-Leu776 in the α1(III) chain [44, 45]. This specific and
characteristic cleavage of collagens at approximately ¾ the length of the collagen fiber from the
N-terminus leads to two fragments that lose stability and unfold to produce single α-chains
called gelatins. Further breakdown of the gelatins and short peptides is not as specific as in the
intact collagen but is traditionally carried out by the gelatinases, MMP-2 and MMP-9, and may
occur at other Gly-Leu and Gly-Ile sites in the polypeptide chains [45]. In addition to MMP-1
breaking down collagens and MMP-2 and MMP-9 degrading gelatins, each of these enzymes are
able to cleave a broad range of substrates, including various collagens, gelatins, elastin,
proteoglycans, regulatory molecules, and other ECM proteins [45-48]. An important result of an
early MMP-1 study was that specificity with this enzyme is largely independent of substrate
conformation and reflects the cleavage site and surrounding amino acid sequences in the native
proteins [49]. As a result, much information has been generated on the specificity requirements
of MMPs by measuring the kinetics of cleaving various short peptide sequences [45]. It was
determined that the recognition of MMPs is based on short peptide sequences up to seven amino
acids in length, three or four amino acids on either side of the scissile bond, and the rate of
cleavage by specific MMPs is determined by the sequence chosen [45]. Several peptides cleaved
by many MMPs do so at Gly-Leu and Gly-Ile sites, mimicking sequences cleaved in native ECM
proteins [45], thus identifying potential sequences and target bonds that may be used in
biomaterial design.
Pioneering work by West and Hubbell [50] incorporated short peptide sequences into
synthetic hydrogel systems making them susceptible to degradation by the cell secreted
proteases, collagenase and plasmin. Hydrogels were developed using the sequences Ala-Pro-
Gly-Leu, with cleavage between the Gly and Leu residues, and Val-Arg-Asn, with cleavage
between the Arg and Asn residues, for degradation by collagenase and plasmin respectively [50].
Subsequent work by Guan and Wagner [51] involved the formation of an elastase sensitive
segmented polyurethane by using the tri-peptide Ala-Ala-Lys in the backbone structure of the
polymer. A critical finding in this study was incorporation of the cleavage site of elastase alone
(Ala-Ala) was enough to confer biological function to the material. Taken together, these
findings suggest that targeting the Gly-Leu and Gly-Ile cleavage sites of MMP-1, MMP-2, and
MMP-9 may confer protease-sensitivity to synthetic biomaterials and may aid in the rational
14
design of biomaterials for cardiac tissue engineering that seek to exploit the presence of these
enzymes.
2.3 Regenerative Approaches to Repair the Heart The need for new therapeutic options to treat an infarcted or failing heart has motivated
researchers to establish techniques to repair, replace, or augment the function of the diseased or
injured tissue. Current techniques being investigated to achieve this fall into a few broad
categories: 1) induction or stimulation of endogenous mechanisms of cardiac repair and
regeneration; 2) the direct transplantation of cells into the damaged tissue; or 3) the use of
biomaterials on their own or in combination with 1 and/or 2 to engineer cardiac tissue either ex
vivo for subsequent transplantation or in situ. Each of these different approaches has potential in
the treatment of heart disease, but it is important to distinguish improvements in physiological
function that occur due to myocardial regeneration and those that occur due to other
mechanisms, such as improved vascularization, reduced scar size, and enhanced cell survival.
Murry et al. [52] recommended that to prove heart regeneration has been achieved, structural,
physiological, and molecular end points must be used to demonstrate the technique has resulted
in newly created cardiomyocytes that are electromechanically connected to host myocardium and
contribute to cardiac function. Any improvement to cardiac function is important for the
treatment of heart disease and warrants extensive investigation, but only the approaches that may
lead to true myocardial regeneration will be discussed here.
2.3.1 Inducing Endogenous Mechanisms in Heart Repair Adult cardiomyocytes have traditionally been considered terminally differentiated cells
that are incapable of proliferating to any significant degree and are therefore unable to regenerate
an injured or diseased heart. Although the inability of the heart to regenerate considerably on its
own appears acceptable, work performed in this field has challenged the view that the heart does
not regenerate at all and stem and progenitor cells that have cardiomyogenic potential have been
identified in the adult heart. Beltrami et al. [53] isolated Lin- c-kitPos cells from the adult rat
heart that were self renewing, clonogenic, and multipotent, giving rise to cardiomyocytes,
smooth muscle cells, and endothelial cells. This same group subsequently identified and isolated
similar c-kitPos cardiac stem cells from human hearts, which mimicked the properties of those
from the rat heart, and demonstrated these cells could form new myocardium in infarcted animal
15
models independent of cell fusion [54]. Oh et al. [55] demonstrated the presence of Sca-1+ (c-
kitNeg) cardiac progenitor cells in the adult murine heart that can differentiate into cells
expressing several cardiac-specific markers in vitro. In response to a myocardial infarction,
endogenous Sca-1+ cells were not mobilized but transplanted Sca-1+ cells homed and
differentiated into cardiac cells in the infarct border zone, half of which fused with host
cardiomyocytes [55]. Matsuura et al. [56] similarly isolated Sca-1+ cells from the adult murine
heart and differentiated these cells in vitro into cardiomyocytes that expressed cardiac
transcription factors and contractile proteins, displayed sarcomeric structures, and contracted
spontaneously. Sca-1+ cardiac progenitor cells have also recently been isolated from adult
human hearts and demonstrate the same potential for deriving cardiomyocytes in vitro as their
mouse equivalents [57]. Martin et al. [58] used an ATP-binding cassette transporter, Abcg2, as a
marker for cardiac side population cells found in the developing and adult murine heart that may
function as a progenitor cell population in developing, maintaining, and repairing the heart. In
addition, Isl1+ cardiac progenitor cells that give rise to cardiomyocytes, smooth muscle cells and
endothelial cells have also been identified in embryonic and postnatal hearts, but it remains
unclear what potential they have in the adult heart [59, 60]. Taken together, several studies have
identified different resident cardiac stem and progenitor cells in an adult heart that may have
potential in regenerating injured myocardium. Evidence has been presented that suggests some
of these cells are activated by injury and inherently contribute to heart regeneration [61], but if it
is occurring, it is not significant on its own to regenerate the tissue. Thus, the next step in
achieving myocardial repair using the body’s endogenous regenerative capacity is determining
how to induce these cells to regenerate significant portions of the injured heart.
2.3.2 Cellular Cardiomyoplasty Cellular cardiomyoplasty, or cell transplantation, is a technique that seeks to promote
cardiac regeneration by introducing cells either directly to the site of injury or to the blood
supply for subsequent homing and integration. This cell-based approach attempts to directly
address the fundamental consequence of a myocardial infarction and a critical component of
heart failure progression; the loss of cardiomyocytes. By introducing cells into the injured heart,
it was hypothesized that the new cells could adapt to the unique myocardial microenvironment
and replace the function of the dead cardiomyocytes. Skeletal myoblasts were the first cell type
to be chosen for transplantation into an infarcted heart and numerous cell types have
16
subsequently been investigated including fetal and neonatal cardiomyocytes, bone marrow-
derived stem cells, endothelial progenitor cells, resident cardiac stem cells, and both mouse and
human embryonic stem cells. Table 2.1 provides a full list of potential cells for myocardial
repair along with some advantages and disadvantages of each for this application. Although
several of these cell types have been shown to improve cardiac function when transplanted in an
infarcted heart and have prompted clinical trials, relatively few of them result in true myocardial
regeneration. Using the recommended guidelines for defining heart regeneration by Murry et al.
Table 2.1: List of cell types considered for cardiac repair. Used with permission from Chen et al. [62].
Cell Source Autologous Easily Obtainable
Highly Expandable
Cardiac Myogenesis
Clinical Trial Safety
Somatic Cells
Fetal Cardiomyocytes No No No Yes No No
Skeletal Myoblasts Yes Yes Depends on age No Yes Yes,
arrhythmias
Smooth Muscle Cells Yes Yes Yes No No No
Fibroblasts Yes Yes Yes No No No
Stem and Progenitor Cells
Mesenchymal Stem Cells Yes No Depends on
age Debated No Yes, fibrosis calcification
Endothelial Progenitor Cells Yes Yes Depends on
age Debated No Yes, calcification
Crude Bone Marrow Cells Yes Yes Depends on
age Debated Yes Yes, calcification
Umbilical Cord Cells No Yes Yes Debated No No
Hematopoietic Stem Cells Yes Yes Yes Debated No Yes
Embryonic Stem Cells No No Yes Yes No Yes, potential
teratomas
17
[52], the ideal cell source for cellular cardiomyoplasty should meet the following criteria: easy to
isolate and expand in vitro; achieve electrical and mechanical integration with host myocardium;
contribute to the structural organization and contractile performance of the heart; and be able to
attain an adult cardiomyocyte phenotype. The literature on cellular cardiomyoplasty is quite
large, so only a brief discussion of the cell types that may lead to true myocardial regeneration
will be presented.
2.3.2.1 Fetal and Neonatal Cardiomyocytes To replace the lost cells associated with a MI, cardiomyocytes are the logical cell choice
for cell-based therapies. Early proof-of-principle work demonstrated transplanted fetal and
neonatal cardiomyocytes can form stable grafts in the myocardium that are electromechanically
connected to host cells via intercalated discs in normal and infarcted hearts [63-65]. Stable
integration of these cells into injured hearts resulted in decreased scar tissue formation, increased
angiogenesis and vascularization, reduced dilatation, and improved ventricular function as
measured using several different techniques [66-71]. Fetal and neonatal cardiomyocytes,
therefore, appear to meet several of the criteria as an ideal cell type for cellular cardiomyoplasty:
electromechanical coupling with host; structural and functional contribution to myocardium; and
the potential of an adult cardiomyocyte phenotype. Importantly, these transplantation studies
provide significant evidence to support the hypothesis that true myocardial regeneration may be
achieved in an infarcted heart and offers continued hope for the development of regenerative
therapeutic options for myocardial repair. Unfortunately, there are a few caveats associated with
the transplantation of fetal and neonatal cardiomyocytes preventing their use in the clinical
setting. First and foremost, human fetal and neonatal cardiomyocytes cannot be used due to
ethical considerations associated with their origin and consequences of harvesting. While
meeting most of the criteria of an ideal cell source for this work, they fall short on being easy to
isolate and expand. Second, only a limited number of transplanted cardiomyocytes engrafted
and survived in the injured heart resulting in the replacement of only a small fraction of the
infarct scar [72, 73]. The low cell engraftment number may be due to cardiomyocyte death
caused by ischemia [73]. These results suggest that a fundamental limitation in cellular
cardiomyoplasty is the delivery, engraftment, and survival of a sufficient number of cells into the
heart to replace the lost cardiac function. Methods to overcome this obstacle may be provided
through the use of biomaterials and will be discussed in greater detail below.
18
2.3.2.2 Embryonic Stem Cell-Derived Cardiomyocytes Embryonic stem cells (ESCs) have the unique advantage over adult stem cells in that they
have the potential of providing a potentially unlimited source of new cardiomyocytes.
Embryonic stem cells are derived from the inner cell mass of the blastocyst stage developing
mammalian embryo [74, 75]. ESCs derived from mice (mESCs) are pluripotent cells capable of
long term undifferentiated proliferation in vitro while retaining the developmental potential of
forming all three embryonic germ layers; endoderm, mesoderm, and ectoderm [76]. Human
ESCs (hESCs) have similar capabilities as mESCs but have the unique advantage of also being
able to give rise to the trophoblast, an extra-embryonic tissue [75, 77]. Being able to give rise to
mesodermal cells, mESCs and hESCs are capable of differentiating into cardiomyocytes with
similar characteristics as those found in vivo and therefore provide a potential source of new
cardiomyocytes for cardiac repair [78, 79]. The potential use of ESCs in myocardial
regeneration, however, requires several criteria be met: 1) a sufficient number of starting ES
cells; 2) efficient and directed differentiation into cardiac progenitor cells or cardiomyocytes; 3)
high production of ESC-derived cardiac progenitors or myocytes; 4) a highly pure population of
desired cells; and 5) resulting phenotype and function similar to adult cardiomyocytes. To be
used in the clinical setting, these criteria must be proven with hESCs. Information and
established techniques for culturing, differentiating, and genetically manipulating murine ESCs,
however, make them a useful model for studying the potential of embryonic stem cell-derived
cardiomyocytes (ESCDCs) in cardiac regeneration. In relation to work conducted in this thesis,
this discussion will mostly focus on mESCs.
To help identify the potential of using ESCs for cellular cardiomyoplasty, several
investigators looked at directly injecting undifferentiated mESCs into the myocardium to test if
the unique microenvironment can drive the differentiation of the cells towards the cardiac
lineage without any adverse affects. Behfar et al. [80] found that transplanted undifferentiated
mESCs differentiated into cardiomyocytes and became functionally integrated into normal and
infarcted myocardium. This work suggests that the host myocardium creates an environment
that can commit undifferentiated mESCs to a specific cardiac lineage and the functionally
integrated cells can lead to an improvement in cardiac function. Similar studies by Min et al.
[81, 82] demonstrated the survival, engraftment, and differentiation of mESCs into mature
cardiac myocytes that attenuated left ventricular hypertrophy, reduced infarct size, improved left
19
ventricular contractility, and increased angiogenesis within the infarcted region. While these
results seem promising for the use of undifferentiated ESCs directly, Nussbaum et al. [83] found
that normal or infarcted hearts do not provide the appropriate cues to guide undifferentiated
mESCs towards a cardiomyocyte fate, but rather leads to teratoma formation and subsequent
rejection in immunocompetent mice. Other groups have similarly found teratoma formation is
the consequence of transplanting undifferentiated mESCs in the heart and other tissues and
represents a major concern for their clinical use [84-86]. As a result of this controversy, a more
restricted ESC-derivative may be more appropriate for use in these studies.
2.3.2.2.1 Differentiation of Murine Embryonic Stem Cells into Cardiomyocytes
Murine ESCs are maintained in an undifferentiated state by coculturing with an
embryonic fibroblast (MEF) feeder layer or by the soluble factor leukemia inhibitory factor (LIF)
and can potentially lead to indefinite self renewal and the generation of a large number of these
cells [87]. Removal of LIF from the culture medium induces differentiation into multiple cell
types and is correlated with a change in expression of the transcription factor Oct-4, a marker of
undifferentiated cells [88]. In vitro differentiation is accomplished via the formation of
aggregated ESCs, called embryoid bodies (EBs), in the absence of LIF leading to the formation
of a number of specialized cells, including cardiomyocytes [89].
Cardiomyocytes derived from mESCs exhibit varying levels of development, which
mimics in vivo differentiation, and appears to be a function of time in culture [89]. In early
beating EBs, mESCDCs may appear as small, round cells with sparse and irregular myofibrils or
more rod-shaped with parallel bundles of myofibrils and A and I bands [78]. As culture time
progresses, cell size increases, ranging greatly from neonatal (diameter ~7-9 µm and length ~20-
45 µm) to adult dimensions (diameter ~10-30 µm and length ~80-150 µm), myofibrils become
densely packed and well organized, and sarcomeres have defined A, I, and Z bands [78]. In
addition, the more developed cells form nascent intercalated discs, with desmosomes, fascia
adherens junctions and gap junctions, and the gap junctions are functional as demonstrated
through dye transfer studies [78]. The cardiac gene expression pattern of mESCDCs follows the
developmental pattern of cardiomyocytes in vivo with the expression of GATA-4 and Nkx2.5
observed prior to ANP, myosin light chain-2v (MLC-2v), α and β-myosin heavy chain, Na+-Ca2+
exchanger, and phospholamban [89]. Sarcomeric protein expression in mESCDCs also follows a
20
progressive developmental pattern observed in vivo [89]. In addition, early mESCDCs express
slow skeletal muscle troponin I, a greater proportion of β-myosin heavy chain, and have a high
sensitivity to calcium similarly seen in embryonic cardiomyocytes [90-92]. Increased culture
time leads to a shift from these fetal isotypes to cardiac troponin I, α-myosin heavy chain, and
decreased sensitivity to Ca2+ more characteristic of mature neonatal and adult cardiomyocytes.
The specialized cardiomyocyte types undergo a shift from pacemaker-like cells early to purkinje-
like cells in the intermediate and atrial and ventricular cells later on [93]. Functionally, the
mESCDCs spontaneously contract, exhibit many features of the excitation-contraction coupling
mechanism found in isolated fetal and neonatal cells, express all major cardiac-specific ion
channels, and may respond to pharmacological agents at later stages of development [89, 93].
Taken together, mESCs can differentiate into cardiomyocytes that initially appear as embryonic-
like cells, but with increased time in culture mature and express more neonatal and adult-like
phenotypes.
Cardiomyocytes spontaneously form in differentiating EBs, but the actual yield of
cardiomyocytes derived from mESCs depends on a number of factors including starting EB size,
culture medium and conditions, ES cell line being used, and time of EB plating [94]. EBs
resemble early post-implantation embryos and the signaling events that drive differentiation into
the different specialized cell lineages loosely mimic those that occur during normal development.
As a result, several extrinsic factors that play a role in cardiomyogenesis in vivo similarly
promote cardiomyocyte formation within EBs. Numerous growth factors and signaling proteins
have been identified that help drive differentiation towards the cardiac lineage in a concentration
and temporal manner including TGF-β1, bone morphogenic protein (BMP)-2, BMP-4, insulin-
like growth factor-1, fibroblast growth factor, hepatocyte growth factor, platelet-derived growth
factor, activin, oxytocin, Wnt/β-catenin inhibition, and erythropoietin [95, 96]. Similarly,
synthetic compounds, such as dimethyl sulfoxide, 5-azacytadine, ascorbic acid, retinoic acid,
opioid, and dynorphin, and free radicals and reactive oxygen species have also been shown to
stimulate cardiomyogenesis [95, 96]. In addition, the physical microenvironment that the
mESCs are placed in may contribute to driving differentiation towards the cardiac lineage.
Features such as matrix composition, topography, 3-D structure, rigidity, and mechanical
stimulation may influence mESCDC yields [97]. Coculture with various cells, such as visceral
endoderm-like cells, may be another method for promoting cardiomyocytes from mESCs [98].
21
Taken together, the identified factors may help drive cells towards the cardiac lineage and
increase the percentage of cardiomyocytes, but have not been able to yield a large and pure
population of ESC-derived cardiac progenitor cells or cardiomyocytes, a requirement for use in
many regenerative applications.
2.3.2.2.2 Large-scale Production of a Pure Population of Embryonic Stem Cell-Derived Cardiomyocytes
One requirement for the successful implementation of ESCs for regenerating the
myocardium is a large and pure population of cardiac progenitors or fully differentiated
cardiomyocytes. As discussed above, any undifferentiated ESCs used in these applications may
lead to undesirable teratoma formation. Similarly, even with the addition of factors to drive cells
towards the cardiac lineage, ESC differentiation results in a mixture of cell types that could have
deleterious consequences when transplanted into the heart.
The recognized need for obtaining a pure population of cells has led to a few novel
techniques for selecting ESCDCs. The first approach involves genetic manipulation of the ESCs
to introduce antibiotic resistance to ESCDCs. Field’s group developed a simple system that
inserted a fusion gene carrying two transcriptional units into mESCs [99]. The fusion gene
contained a phosphoglycerate kinase promoter driving hygromycin resistance gene (pGK-hydror)
to select for mESCs that were stably transfected and a α-myosin heavy chain promoter driving an
aminoglycoside phosphotransferase gene (MHC-neor), which allowed for the selection of
mESCDCs by adding the neomycin analog geneticin (G418) to culture medium. The mESCDCs
selected by this method were >99% pure and expressed markers of highly differentiated cardiac
cells [99]. Kolossov et al. [85] used a similar approach to get a highly purified mESCDC
population (>99%) by having the α-myosin heavy chain promoter drive both a puromycin
resistance gene and a green fluorescent protein (GFP) gene for purification and identification of
the cells. Other groups have also used this genetic selection method for purifying
cardiomyocytes from mESCs and hESCs [100-102]. A second approach to purifying ESCDCs is
to label cell surface markers with fluorescent or magnetic tags and to use fluorescence activated
cell sorting (FACS) or magnetic-activated cell sorting. If appropriate cell surface markers
become available, this approach is advantageous as it avoids any need to genetically modify the
cells. Cell surface markers for cardiac progenitor cells or cardiomyocytes are currently not
known [103], but the molecular signature of ESCDCs is being explored [104] and may lead to
22
potential markers to be used in this purification scheme. Still, the proof of concept for purifying
mESCDCs by FACS has been demonstrated by Muller et al. [105] using a transgenic mESC line
expressing GFP under the cardiac chamber-specific promoter MLC-2v. A Percoll gradient
separation followed by FACS resulted in a >97% pure cardiomyocyte population and
electrophysiological tests identified the cells were preferentially ventricular-like [105]. Hidaka
et al. [106] also used FACS to purify Nkx2.5 positive cardiac progenitor cells from mESCs and
showed the resulting cells differentiated into sinoatrial node, atrial, or ventricular-like cells.
Transgenic GFP expression has been used to identify other cardiac progenitor cells or specialized
cardiomyocyte types including pacemaker, atrial, and ventricular cells [80, 107, 108], suggesting
these cells may also be separated by the FACS approach. Other methods to help purify ESCDCs
are a Percoll gradient separation and manually picking out beating cells, but both methods only
lead to an enriched culture of mESCDCs and these heterogeneous cell populations may inhibit
clinical acceptance [103].
The large number of cells required for use in cell-based regenerative strategies for
myocardial repair has motivated researchers to develop systems for the large-scale production of
ESCDCs. The Zandstra group has been particularly interested in developing and optimizing
bioreactor parameters for the generation of large quantities of ESCDCs. In an early study,
Zandstra et al [109] aggregated MHC-neor/pGK-hygror mESCs in static culture for 4 days and
transferred the EBs to a spinner flask system for subsequent growth and differentiation. On day
9 after initiating differentiation, medium was supplemented with G418 and retinoic acid to select
for and drive differentiation towards cardiomyocytes. A relatively pure mESCDC population
was harvested from the spinner flasks on day 18 with no undifferentiated mESCs and the cells
were spontaneously beating and expressed characteristic markers of mESCDCs [109].
Importantly, this system allowed the generation of ~1.4 x 107 cells in a 250 ml spinner flask
using the CM7/1 mESC line, thus identifying the large-scale production of mESCDCs.
Subsequent work by this group optimized cell and bioreactor conditions to produce even greater
numbers of ESCDCs. Bauwens et al. [110] used a similar approach but encapsulated the EBs in
a hydrogel to prevent aggregation and generated nearly 24 times more ES cell-derived
cardiomyocytes (~3.15 ESCDC per input mESC) than in unencapsulated controls (~0.15 ESCDC
per input ESC) after 9 days of differentiation and 5 days of selection using the D3 mESC line.
These mESCDC yields were increased even further (~3.77 mESCDC/ESC) by perfusion feeding
23
the bioreactor system and culturing under hypoxic conditions (~4% O2 compared to normoxic
levels of ~20%). Recently, Niebruegge et al. [111] directly inoculated bulb-shaped glass spinner
flasks with 2 x 105 CM7/1 mESCs per ml of culture medium for EB formation within the
bioreactor system. Optimization of several parameters including addition of retinoic acid at day
7 instead of day 9, starting selection at day 11 not day 9, and changing 50% of medium every
other day instead of every day resulted in a 4.3 fold increase in number of mESCDCs (~7.6
ESCDC/input mESC for optimized conditions vs. ~1.8 for unoptimized method) with a total of
19 x 107 cardiomyocytes in the 250 ml spinner flask [111]. Similar optimization efforts have
been reported by this group for the large scale generation of human ESCDCs. Factors such as a
homogenous starting EB size, a stirred suspension bioreactor, and hypoxic culture conditions are
more conducive of cardiomyocyte generation and improve hESCDC yields [112, 113].
Ultimately, the ability to derive a large and pure population of cardiomyocytes or cardiac
progenitor cells from ESCs is a critical step in identifying a cell source for regenerating the
myocardium.
2.3.2.2.3 Transplantation of Murine and Human ESC-derived Cells into the Heart
Several studies have been conducted to transplant murine and human ESCDC or cardiac-
committed ESCs for cellular cardiomyoplasty. In a study by Klug et al. [99], a highly pure
population of mESCDCs were transplanted into the heart and formed stable intracardiac grafts
out to at least 7 weeks with a similar frequency of engraftment as fetal murine cardiomyocytes.
Menard et al. [114] transplanted cardiac-committed mESCs into infarcted sheep myocardium and
demonstrated the cells successfully engrafted into the infarct region, differentiated into mature
cardiomyocytes that were electrically connected to host myocardium, improved left ventricular
ejection fraction, and may avoid immune rejection. Kolossov et al. [85] injected purified
mESCDCs either alone or with an equal number of mouse embryonic fibroblasts into an
infarcted heart. It was determined that the mESCDCs had very low engraftment frequency when
injected on their own but was significantly increased when transplanted together with the
fibroblasts. The engrafted mESCDCs formed mature sarcomeric structures, electrically coupled
with host cardiomyocytes, did not develop into teratomas, contributed to ventricular force
contraction, and improved left ventricular function [85]. Thus, using mESCs to form cardiac
24
committed cells or a pure population of cardiomyocytes avoids the concern of teratoma
formation and may contribute to true myocardial regeneration.
While mESCs provide a good model for studying development, disease, and the potential
of ESCs in regenerative medicine, ultimately hESCs must be used if this technology is ever
going to transfer to the clinical setting. Several studies have been conducted using hESCs for
cellular cardiomyoplasty in animal models. Although the transplantation of undifferentiated
hESCs into an infarcted animal myocardium may help drive differentiation towards the
cardiomyogenic lineage without teratoma formation [115], the lessons learned from mESCs
along with the finding that these cells do form teratomas [116, 117] suggests hESCs must be at
least somewhat committed if they are going to gain clinical acceptance. As a result, recent work
has investigated transplanting hESC-derived cardiomyocytes (hESCDCs) or cardiac committed
hESCs into normal and infarcted animal hearts. Evidence has been provided that hESCDCs
survive, proliferate, form mature contractile structures, and electrically couple to host cells
following transplantation into healthy hearts of immunodeficient mice and rats [116, 118, 119].
In infarcted rodent myocardium, cardiac-committed hESCs and hESCDCs similarly appear to
survive and form a mature cardiomyocyte phenotype without any teratoma formation [116, 117,
119-122]. In addition, several of these investigations report an improvement to cardiac function
due to the engraftment of the cells [116, 117, 119, 121].
A few limitations associated with hESCDC transplantation studies include an inefficient
hESC differentiation into cardiomyocytes, a heterogeneous cell population, poor cell survival
after transplantation, and a transient contribution to cardiac function. In an attempt to overcome
some of these limitations, Laflamme et al. [121] treated a high density monolayer of
undifferentiated hESCs with two cytokines to drive differentiation towards the cardiac lineage
and enriched the hESCDCs by a Percoll gradient to get a ~83% pure cardiomyocyte culture (3:1
ratio of generated cardiomyocyte to input hESC). Transplantation of these cells along with a
mixture of prosurvival reagents into infarcted rat hearts resulted in large muscular grafts of
human myocardium in the central infarct region that coupled to host tissue and significantly
improved ventricular structure and contractile function compared to appropriate controls [121].
Despite long-term survival of the transplanted cells, the exact contribution they have on cardiac
function and their long-term benefit remains unclear [119, 123].
25
While ESC appear to be the best candidate for regenerating the myocardium, ethical
considerations associated with harvesting human ESCs has limited wide-spread funding and use
of these cells. Induced pluripotent stem (iPS) cells may offer an alternative source of
cardiomyocytes [124] that may reduce ethical concerns, but this technology is still in a very early
stage with many obstacles to overcome before this option is viable. Regardless of the cell type
or source, cellular cardiomyoplasty on its own is limited by the delivery, engraftment, and
survival of cells in the heart. Interestingly, the prosurvival reagents used by Laflamme et al.
[121] for achieving improved long-term engraftment of hESCDCs included Matrigel, a
gelatinous protein mixture rich in structural extracellular matrix proteins and growth factors. In
this situation, Matrigel acted as a supportive matrix for the cells to adhere to, thus increasing cell
survival. The use of biomaterial scaffolds as a delivery vehicle for cells may overcome the
limitations of cellular cardiomyoplasty and may help to contribute to long-term clinically
relevant cardiac repair.
2.3.3 Cardiac Tissue Engineering Cardiac tissue engineering is a technique that employs the use of biomaterials in
combination with cells and/or various signaling agents towards the development of viable tissue
constructs for regenerating the myocardium. Research in this area has expanded tremendously
over the past several years and continues to grow, but the goal remains the same and to date, a
few promising strategies have emerged. These approaches include using cells along with: 1)
biodegradable synthetic and natural-based scaffolds with pre-formed three-dimensional
structures; 2) biodegradable synthetic and natural-based materials with undefined structures; 3)
injectable biomaterials for in situ myocardial regeneration; and 4) temperature-responsive
biomaterials that act as a substrate for cell sheet formation. The focus of this thesis is on the first
of the four approaches, but a brief discussion of the other three will first be presented.
2.3.3.1 Myocardial Tissue Engineering Using Biomaterials with Undefined Structures
Cardiac cells have an endogenous ability to organize into native-like cardiac tissue when
cultured in an environment that doesn’t provide signaling cues inhibiting them from doing so
[125-127]. A promising approach to engineering myocardial tissue is to exploit this endogenous
capability of the cardiac cells by combining them with biomaterials of undefined structures
allowing tissue organization to be dictated by the cells. The most significant work in the cardiac
26
tissue engineering field using this approach has come from Zimmermann, Eschenhagen and
colleagues. The system they developed uses liquid collagen type I and various species of
cardiomyocytes or cardiac cells for making what they termed engineered heart tissue (EHT). In
early work, Eschenhagen et al. [128] showed spontaneous remodeling of liquid collagen by
embryonic chick cardiomyocytes and the formation of a spontaneously and coherently beating
cardiomyocyte populated matrix. The 3-D structure resembled early myocardial tissue,
prevented the dedifferentiation of primary cardiomyocytes and overgrowth of non-myocytes, and
could be used to measure contractile force generation alone or in the presence of
pharmacological agents. Subsequent work demonstrated this technique could be applied to
mammalian cells through the addition of Matrigel, the EHTs could be formed into various
geometries, can be used for long-term contractile force measurements, can be electrically and
mechanically paced, and used for genetic and pharmacological manipulation [129-131].
Importantly, mechanically stimulating the EHTs led to a differentiated cardiac muscle construct
that had improved organization of parallel, oriented, rod-shaped cardiomyocytes with mature
sarcomeric structures, were electromechanically coupled, secreted ECM proteins to form a
basement membrane surrounding myocytes, and improved contractile function of the tissue [129,
131]. In addition, in vivo assessment of the EHT implanted in rats revealed the transplanted
tissue survived for at least 8 weeks, remained contractile, morphologically integrated into the
host myocardium, and became vascularized and innervated [132, 133].
An initial attempt was made to perform in vivo studies in immunocompetent rats by
developing EHTs with cells and collagen from syngeneic donors, but there was a need to move
to immunosuppression due to immune rejection from the use of Matrigel and/or serum in culture
medium [134]. Recognizing the limitation of this approach for translating to the clinical setting,
methods were investigated to optimize the EHT by looking at different heart cell populations,
defined serum-free medium, and Matrigel-free culturing conditions [135]. A replacement for
Matrigel was achieved by adding insulin and triiodothyronine for 24 hrs on the first day of EHT
culture and a defined serum-free culture was established using a cocktail of reagents and growth
factors. Importantly, it was found that EHTs formed from a native heart cell composition
improved contractile function and response to pharmacological agents and increased the number
of vascular-like structures compared to an enriched cardiomyocyte cell population. In this
system, the non-myocytes may play an important role in remodeling the provisional 3-D matrix
27
formed by the collagen and in secreting new ECM for improved organization and function of the
EHT [134]. Furthermore, the enhanced number of vascular-like structures with the native heart
cell population may also facilitate faster vascularization of the EHT during transplantation into
the heart. Taken together, this study may help resolve some immunological issues associated
with EHTs and highlights the importance of using non-myocyte cells alongside cardiomyocytes
for the formation of myocardial tissue.
In an attempt to move away from neonatal cells, Guo et al. [136] used the EHT
technology to form cardiac tissue with mouse embryonic stem cell-derived cardiomyocytes. In
this study, an enriched population of mESCDCs was achieved by a Percoll separation strategy
and fibroblasts, neurons, and vascular endothelial cells were identified throughout the engineered
tissue with no undifferentiated ESCs. After mechanical stretching, the tissue beat spontaneously
and synchronously, generated contractile force, and responded to pharmacological manipulation.
The degree of cardiomyocyte differentiation and maturity varied, but some cells had highly
organized myofilaments and formed adherens junctions, desmosomes, and gap junctions with
neighboring cells. Implantation of the engineered tissue into nude mice demonstrated survival
and vascularization of the engineered tissue with no teratoma formation [136]. This important
early study identifies the potential of using ESCDCs for cardiac tissue engineering and suggests
the techniques developed for neonatal cardiac cells may be translated to ESC-derived cells.
Work is currently being conducted using the EHT technique along with iPS cell-derived
cardiomyocytes [137]. Still, it has not been determined if implanting EHT into an infarcted heart
improves systolic and diastolic function in a malfunctioning myocardium and is needed before
the potential of this technique is fully understood.
2.3.3.2 In Situ Cardiac Tissue Engineering The appealing concept behind cellular cardiomyoplasty associated with the fairly non-
invasive delivery of cells to the injury site along with the engraftment limitations of injecting
cells on their own have allowed the emergence of in situ cardiac tissue engineering. In this
approach biomaterials are injected either on their own or in combination with cells to form a
supporting matrix in the infarcted heart that may limit the detrimental remodeling events
associated with a MI and promote cardiac regeneration. In addition, when used in combination
with cells, the polymers act as a supportive substrate for cells to adhere to during transplantation
and may help to improve cell engraftment, survival, and distribution of the cells in the heart.
28
Several studies have demonstrated the feasibility and beneficial effect of using injectable
biomaterials in this cardiac tissue engineering approach. Christman et al. [138] demonstrated
that fibrin glue preserves ventricular geometry and improves cardiac function suggesting the
fibrin glue acts as a scaffold that prevents infarct expansion and provides space and/or signaling
for blood vessel ingrowth. Huang et al. [139] found fibrin, collagen I, and Matrigel each on their
own significantly enhance angiogenesis and blood flow in an infarcted myocardium with
collagen I showing a significantly higher number of myofibroblasts in the scaffold. Dai et al.
[140] similarly injected collagen into an infarcted myocardium and reported a thickening of scar
tissue due to the collagen scaffold that improved left ventricular stroke volume and ejection
fraction. Neovascularization and improvement to cardiac function has also been observed when
injecting alginate [141] and self-assembling peptides [142] on their own into injured hearts.
Although injectable biomaterials on their own show promising results, the combination of
these materials with cells leads to a more significant improvement to cardiac function.
Christman et al. [143] injected fibrin glue and skeletal myoblasts into infarcted rat hearts and
showed the fibrin glue significantly improved the survival of myoblasts, increased the
engraftment size, significantly reduced infarct scar size, and improved blood flow compared to
skeletal myoblast transplantation or fibrin alone. Similar beneficial effects on cardiac function
with neovascularization in the infarcted region has been demonstrated with fibrin glue and bone
marrow-derived cells [144]. In addition to fibrin, other biomaterials have also been used as
injectable scaffolds with cells for cardiac tissue engineering. Zhang et al. [145] mixed
ventricular cardiomyocytes with collagen and Matrigel, similar to the EHT technique, and
injected it into an infarcted rat heart before it solidified into a gel. In this system, the
transplanted cardiomyocytes survived, formed condensed tissue, and expressed the gap
junctional protein connexin-43 (Cx-43). Moreover, the transplanted cells and matrix preserved
left ventricular wall thickness and cardiac function better than either the cells or matrix alone, or
sham control. Matrigel has also been used to deliver mESCs into an intramural left ventricular
pouch and into infarcted myocardium with an observed improvement to heart function compared
to appropriate controls [146, 147].
The benefits afforded to injecting biomaterials alone or with cells suggest biomaterials
play an important role in repairing infarcted hearts. The mechanisms behind the beneficial
effects, however, still need to be elucidated. A current limitation with this approach is the types
29
of biomaterials that may be used. Injectable polymers must be liquid or of reasonable viscosity
to be injected into the heart and must form a solid matrix quickly once at the appropriate site.
The development of new temperature-responsive and other injectable biomaterials will certainly
hold an exciting future for in situ cardiac tissue engineering.
2.3.3.3 Myocardial Cell Sheets An emerging technique to myocardial tissue engineering is cell sheet engineering. In this
approach, temperature responsive biomaterials are used as a surface for culturing cardiac cells
into a confluent cell layer without the use of biodegradable biomaterials as an ECM. These
temperature-responsive biomaterials, covalently attached to culture surfaces, are slightly
hydrophobic at 37˚C allowing for cell adhesion, proliferation, establishment of cell-cell contacts,
and secretion of ECM proteins [148]. By lowering the temperature, the biomaterial surfaces
become highly hydrophilic and non-adhesive, allowing the cells to spontaneously detach from
them. In this system, the cell-cell junctional contacts critical for electromechanical coupling of
cardiomyocytes, including gap junctions, desmosomes, and fascia adherens junctions, remain
intact in the cell sheets leading to spontaneously and synchronously beating 2-D functional cell
layers [149, 150]. Furthermore, stacking the individual cell sheets leads to the formation of 3-D
myocardial-like tissue with the formation of appropriate cell-cell connections between layers for
synchronous beating and force production [149-152]. The temperature-responsive biomaterial
used in this work is poly (N-isopropylacrylamide) (PIPAAm) [148], although cell sheets have
also been formed using fibrin-coated dishes [153]. Recently, PIPAAm was grafted onto
microtextured polystyrene substrates that were structurally organized with an array of parallel
grooves [154]. These grooved surfaces led to the formation of cell sheets that had cells highly
aligned in the direction of the grooves and this aligned cell orientation was maintained when
detached from the PIPAAm and further cultured on tissue culture polystyrene dishes. Although
this study was conducted using smooth muscle cells, this technique should work well in the
alignment of cardiomyocytes, an important criterion for developing myocardial tissue. In
addition, cell sheet engineering can be used to form sheets that incorporate multiple cell types,
such as cardiomyocytes, endothelial cells, and fibroblasts, and may contribute important
structural and functional benefit to engineered tissue [155, 156].
The potential of this unique tissue engineering approach has been further demonstrated
through many in vivo studies. Shimizu et al. [157] showed that layers of cardiomyocyte sheets
30
could survive long-term after being implanted into rats with the implanted tissue being well
vascularized, contractile, force producing, and maintaining an elongated and well differentiated
cell structure. When implanted into an infarcted myocardium, cardiomyocyte cell sheets two
layers thick appeared homogenously integrated with host myocardium, were vascularized,
expressed Cx-43, and significantly ameliorated cardiac performance compared to cell sheets
from fibroblasts or untreated controls [158]. Cardiomyocyte cell sheets formed from fibrin-
coated dishes, instead of PIPAAm, showed bidirectional action potential propagation between
host hearts and transplanted tissue demonstrating similar functional integration with infarcted
hearts [159]. Furthermore, forming cell sheets by co-culturing cardiomyocytes with endothelial
cells promotes neovascularization with the transplanted endothelial cells contributing to new
vessel formation [156, 160]. Increasing the endothelial cell to cardiomyocyte ratio enhances
neovascularization and improves cardiac function of infarcted hearts. Although
neovascularization within the cell sheets occurs quickly, especially when they contain
endothelial cells, diffusion limitations leading to hypoxia, insufficient nutrient supply, and
accumulation of waste prevent the transplantation of engineered tissue greater than three cell
layers thick (~80 µm) without getting necrotic cells in the middle of the tissue. To resolve this
issue, Shimizu et al. [161] performed a multistep transplantation of 3 layer thick cell sheets once
every few days allowing each transplant time to become vascularized before adding the next
layer. With this technique, engineered tissue ~ 1 mm in thickness was obtained in vivo without
tissue necrosis and became well vascularized and completely synchronized. This polysurgical
procedure may provide a method for overcoming one of the fundamental limitations with
myocardial tissue engineering - obtaining thick, vascularized engineered tissue in vivo while
preventing necrosis. Similarly, the lack of a suitable cell source for cardiomyocytes limiting
cellular cardiomyoplasty plays an equally large barrier in tissue cardiomyoplasty. As a result,
cell sheet engineering has formed tissue using mesenchymal stem cells [162], autologous skeletal
myoblasts [163], and the coculture of fibroblasts and endothelial progenitor cells [155], all of
which have improved cardiac function when implanted in infarcted myocardium. Despite the
drawback of significant macroscopic strength associated with cell sheet engineering, this
technique holds tremendous potential as a regenerative approach to repairing an infarcted heart.
31
2.4 Cardiac Tissue Engineering Using Pre-Formed Three-Dimensional Scaffolds
Biomaterials are traditionally used in tissue engineering as a temporary extracellular
matrix that provides an environment suitable for tissue development. These biomaterial
scaffolds should have the appropriate physical, chemical, mechanical and degradation properties
to promote cell attachment, proliferation, differentiation, organization, vascularization in vivo,
integration with host tissues, and the replacement of the scaffold by newly formed ECM at a rate
appropriate for the gradual transfer of mechanical load from biomaterial to new tissue. The
biomaterials used in making the scaffolds must be non-toxic and non-immunogenic and the same
should hold true with the materials degradation products. One of the main advantages of using
pre-formed 3-D scaffolds is the ability to tailor the structural and functional properties of the
scaffolds to meet these requirements for site-specific applications. Typically, this approach
utilizes biomaterial scaffolds either in the development of cardiac tissue in vitro for subsequent
implantation or as a transplantation vehicle of cells for in situ myocardial formation. Either way,
there are four main design components that determine the success of this tissue engineering
method; 1) biomaterial choice, 2) scaffold fabrication technique, 3) cell type, and 4) seeding and
cultivation methods. Ultimately, by combining advances made in each of these different
components, an ideal scaffold and tissue formation method may be established leading to
successful regeneration of myocardial tissue.
2.4.1 Biomaterials for Cardiac Tissue Engineering The development of an ideal scaffold for engineering myocardial tissue using pre-formed
3-D scaffolds requires the proper biomaterial be selected. The appropriate starting material may
help to achieve such scaffold requirements as the material and its degradation products being
non-toxic and non-immunogenic, appropriate mechanical properties, suitable degradation rates,
and physiochemical properties that promote cell adhesion, growth, and differentiation. To date,
biomaterials that have been used in cardiac tissue engineering are those that are naturally-
derived, such as collagen, gelatin, fibrin, hyaluronic acid, alginate, and decellularized whole-
ECM, those that are synthetically manufactured, including several polyesters, polylactones, and
polyurethanes, or composites of the two [164]. Table 2.2 highlights several biomaterials and
how they have been used in cardiac tissue engineering. The ideal biomaterial has yet to be
determined or developed, but several of these biomaterials show potential and guidelines for
32
Table 2.2: Summary of biomaterials and their application in cardiac tissue engineering. Used with permission from Chen et al. [62].
Biomaterials Physical State Tissue Engineering Approach
Naturally-derived
Collagen Gel or 3-D porous mesh
Epicardial heart patch & 3-D engineered tissue
Fibrin glue Injectable gel Endoventricular heart patch
Peptide nanofiber Injectable gel Endoventricular heart patch
Collagen – glycosaminoglycans 3-D porous mesh 3-D engineered tissue
Gelatin mesh 3-D porous mesh 3-D engineered tissue
Alginate mesh 3-D porous mesh 3-D engineered tissue
Synthetic Biodegradable
Poly(lactic acid) 3-D porous mesh 3-D engineered tissue
Poly(glycolic acid) and copolymer with poly(lactic acid)
3-D porous mesh 3-D engineered tissue
Polycaprolactone and copolymer with poly(lactic acid)
3-D porous mesh 3-D engineered tissue
Poly(glycerol sebacate) 3-D porous mesh
Epicardial heart patch & 3-D engineered tissue
Segmented Polyurethanes 3-D porous mesh
Epicardial heart patch & 3-D engineered tissue
Synthetic Non-degradable
Poly(ethylene terepthalate) Knitted mesh Left ventricular constraint & cardiovascular grafting
Polypropylene Solid sheet Left ventricular constraint
Poly(tetrafluoroethylene) with or without poly(lactic acid) and or poly(glycolic acid) Solid sheet Treatment of congenital heart disease
& cardiovascular grafting
Poly(N-isopropyl acrylamide) Solid sheet Scaffold-free cell sheet engineering
33
selecting biomaterials for this application are beginning to become more established.
2.4.1.1 Natural Biomaterials Naturally-derived biomaterials offer the advantage of being components of the native
ECM, alginate excepted, and therefore inherently contain signaling motifs that promote cell
attachment, growth, and differentiation and may be degraded by cell-secreted proteases. Several
investigators have shown a high density of spatially uniform cardiac cells can be cultured in pre-
formed collagen and alginate to form constructs that are spontaneously beating and develop force
[165-168]. In addition, in vivo studies have demonstrated the feasibility of transplanting the
constructs onto an infarcted myocardium with cell survival and neovascularization. In one study,
Li et al. [169] seeded fetal rat cardiomyocytes on biodegradable gelatin meshes to form
spontaneously beating grafts that were implanted into cryoinjured hearts. The grafts survived for
at least 5 weeks, were observed to have blood vessel ingrowth, and appeared to integrate with the
host cells. In another study, Leor et al. [170] implanted beating cardiac grafts formed from fetal
cardiac cells grown in alginate scaffolds into the infarcted myocardium of rats. Similar
neovascularization was reported and the transplanted engineered tissue attenuated left ventricular
dilatation and heart failure. While these early studies demonstrate the potential of naturally-
derived materials, constraints of their use include limited control over the mechanical properties
and degradation rates, large scale production, and batch-to-batch variability [171].
2.4.1.2 Synthetic Biomaterials In contrast to the naturally-derived materials, synthetic biomaterials offer more control
over their physical, chemical, mechanical, and degradation properties. Importantly, synthetic
biomaterials may be combined with ECM proteins or are being developed to incorporate specific
peptide sequences that confer unique biological functionality to them. As a result, the benefits
that were once afforded to natural biomaterials can be exploited and incorporated into these bio-
mimetic polymers leading to synthetic polymers that meet several of the requirements of an ideal
biomaterial for cardiac tissue engineering. Although this has yet to be achieved, progress in this
field is proceeding rapidly towards reaching this goal.
2.4.1.2.1 Traditional Polymers for Tissue Engineering Traditional synthetic scaffolds used in tissue engineering include the polyesters poly
(lactic acid) (PLA), poly (glycolic acid) (PGA), and the copolymer poly (lactic-co-glycolic acid)
34
(PLGA) and have been applied to cardiac tissue engineering. Some successes have been made
using these scaffolds for generating engineered cardiac tissue in vitro. Bursac et al. [172], for
example, seeded primary neonatal rat ventricular cells onto porous PGA scaffolds and cultured
the constructs in a bioreactor for 1 week. The samples formed a peripheral tissue-like region 50-
70 µm thick containing differentiated cardiomyocytes organized into multiple layers in a 3-D
configuration. In this study, the engineered tissue had homogenous electrical properties and
continuous impulse propagation and was deemed a good model for electrophysiological studies.
A follow up study found that coating the PGA scaffolds with laminin improved the molecular,
structural, and electrophysiological properties of the constructs that closer mimicked native
ventricular tissue [173]. The use of these polyesters in developing in vitro models of myocardial
tissue warrants their investigation, but it has been recognized that these polymers are
inappropriate for implanting in the body for cardiac repair. Acidic degradation products from
these polyesters affect cell viability and may lead to an intense inflammatory response and an
adverse tissue reaction after implantation [174]. Upon degradation, the mechanical integrity of
these polyester scaffolds is lost rapidly and may not provide enough time for the engineered
tissue to buildup the appropriate structure to support itself [175]. In addition, these polymers are
stiff and possess inappropriate mechanical properties for cardiac tissue engineering. As a result
of the downfalls associated with PGA, PLA, and PLGA, other synthetic biomaterials have been
developed that are more promising for this application.
2.4.1.2.2 Elastomeric Biomaterials The dynamic cardiac environment requires the use of biomaterials that are soft and
flexible and can withstand the repetitive forces associated with the cardiac cycle without losing
mechanical integrity. Furthermore, the critical role mechanical forces play in the development of
normal cardiac tissue requires the polymers have the appropriate mechanical properties to
promote mechanotransduction between cells and the external cardiac or cardiac-mimicked
environment [21]. An important recent study by Engler et al. [176] suggests the elasticity of the
surface used to culture cardiomyocytes plays a critical role in developing force, in organizing the
cytoskeleton, and in sustaining a contractile phenotype. It was determined that matrices that
mimic the elastic modulus of the myocardial microenvironment (initial modulus (E) ~ 10-15 kPa
for embryonic quail myocardium) are optimal for cardiomyocytes to produce transmittable force,
promote a striated myofibrillar structure, and sustain rhythmic contractions long-term.
35
Cardiomyocytes cultured on hard matrices that mimic post-infarcted scar tissue (E ~ 35-70 kPa)
overstrain themselves, lack a striated appearance, and stop beating after a few days in culture.
Cells on soft matrices (E ~ 1 kPa) remain contractile for several days in culture but produce little
work. This study clearly identifies the influence elastic modulus has on cardiomyocyte structure
and function and suggests that understanding the mechanical properties of the native
myocardium may help to establish the appropriate parameters for developing biomaterials to be
used in the heart. Towards this goal, the ultimate tensile strength (σ), under static loading, of
human cardiac muscle taken from the middle layer of left ventricular myocardium is
approximately 115 mN/mm2 (115 kPa) in the direction parallel to contraction and about 38 kPa
in the transverse direction [177]. The ultimate percentage elongation (ε) of the myocardial tissue
is 66% in the parallel direction and 86% in the transverse direction [177]. This tissue has a
Young’s modulus of ~20-500 kPa [177, 178]. In addition, a normal resting heart undergoes 75
beats/min [4] with a stress of at least 60 kPa generated by the myocardium for ~0.3 s/beat [179].
Taken together, mechanical properties of biomaterials should mimic those of the native
myocardium and are important design criteria in biomaterial development.
The importance of soft, flexible and elastic scaffolds suggests elastomeric biomaterials
may be appropriate for use in the heart. Several biodegradable polyesters,
polyhydroxyalkanoates, segmented polyurethanes, and composite materials have been developed
with elastic mechanical properties that are promising for cardiac tissue engineering applications
[51, 178, 180-188]. Chen et al. [178] recently synthesized poly(glycerol sebacate) (PGS) at
different temperatures yielding biocompatible, degradable, elastomeric materials that mimic the
mechanical properties of myocardial tissue with an elastic modulus ranging from 0.056 – 1.2
MPa. Similarly, Englemayr et al. [189] used PGS to create accordion-like honeycomb scaffolds
that closely mimicked the mechanical properties of the right ventricular myocardium of adult rats
and successfully cultured cardiac cells on these scaffolds towards the formation of anisotropic
cardiac-like tissue. Park et al. [190] developed an elastomeric composite scaffold made from the
FDA-approved materials poly(lactide-co-caprolactone), PLGA, and collagen type I and seeded it
with neonatal cardiac cells. Although the mechanical properties of the scaffold were not
determined, the construct showed improved cellularity, expression of cardiac markers and
contractile function compared to controls of collagen I and PLGA alone. Importantly, the
authors attributed the improved contractile function to the elastic mechanical properties of the
36
composite scaffold which promoted mechanotransduction within the construct and the formation
of electromechanical coupling of cells [190]. Scaffold stiffness associated with the PLGA
scaffold prevented construct contraction despite cell attachment and expression of cardiac
markers. McDevitt et al. [191] used segmented polyurethane (PU) films composed of
polycaprolactone, a lysine-based diisocyanate, and a phenylalanine-based chain extender to
culture neonatal rat cardiomyocytes and demonstrated the cells retain a differentiated phenotype
and contractile function. This polymer has an initial modulus that is much greater than native
myocardium (~50 MPa in ambient conditions), but it was suggested that the elasticity may have
contributed to the cells ability to preserve a normal morphology and function. More recently,
this same polyurethane has successfully been used as films and 3-D scaffolds to culture primary
cardiac cells and embryonic stem cell-derived cardiomyocytes with positive results [192-194].
Wagner’s group similarly developed a series of elastomeric segmented polyurethanes using
triblock soft segments of polycaprolactone-polyethylene oxide-polycaprolactone, 1,4-
butanediisocyanate, and either putrescine or peptide chain extenders that show great promise in
cardiac tissue engineering [51, 195, 196]. This group demonstrated these polyurethanes have the
appropriate mechanical and biocompatibility properties for use in the cardiac environment and
improved cardiac function when implanted on infarcted rat hearts [197, 198]. One of the
advantages of using segmented polyurethanes for this work is the flexibility in chemistry used to
synthesize these polymers allows the ability to tune the mechanical properties and degradation
rates and incorporate chemical moieties to form bio-mimetic materials, discussed in greater
detail in section 2.5. Taken together, elastic mechanical properties are a requisite for
biomaterials that will be used in the heart. These materials promote mechanotransduction within
engineered constructs, help prevent detachment from surfaces, aid in transmitting forces during
contraction, and may contribute to elastic recoil during relaxation. As a result, elastomeric
biomaterials are going to play a major role in identifying ideal biomaterial scaffolds for cardiac
tissue engineering.
2.4.2 Scaffold Fabrication Techniques The proper biomaterial choice can contribute to the success of a biomaterial scaffold for
engineering myocardial tissue, but the scaffold fabrication technique also plays a major role in its
outcome. Scaffolds provide a delivery vehicle and framework for cells to attach to, migrate
along and organize into functional homogenous tissue. Some of the parameters which must be
37
considered when forming scaffolds for tissue engineering include: a high surface area-to-volume
ratio to achieve a high density of electromechanically coupled cells and efficient transport of
oxygen, nutrients, and waste; pore sizes that allow cell migration; surface chemistries and
features that promote cell adhesion, growth, and differentiation; and the appropriate architecture
to organize the tissue [171, 199]. While a scaffold that meets all these requirements may be
obtained by decellularizing a cadaveric heart [200], this is not an option for synthetic
biomaterials.
Several polymer processing techniques can be used to form synthetic polymeric scaffolds
that meet many of the requirements for tissue engineering applications, including thermally
induced phase separation, molecular self assembly, laser ablation, electrospinning, solvent
casting and particulate leaching, polymer extrusion, gas foaming, phase inversion, rapid
prototyping, and others [171, 201]. Not all of these, however, can be easily used to form
scaffolds with the appropriate architecture that promotes anisotropic organization of
cardiomyocytes. The anisotropic organization of cardiomyocytes is critical for the structure and
function of the native myocardium and is increasingly being recognized as an important design
consideration for engineered myocardial tissue. Recent studies have identified different methods
of producing scaffolds with architectures that promote anisotropic cardiac tissue formation.
Bursac et al. [202] melted sucrose and extruded it to form a fibrous template with which to form
anisotropic fibrous and foamy scaffolds from PLGA through a solvent casting and particulate
leaching method. Neonatal cardiac cells cultured on the scaffolds aligned along the polymeric
fibers, were interconnected, and supported macroscopically continuous, anisotropic impulse
propagation as determined by optical mapping. Englemayr et al. [189] used a laser excimer
ablation technique to create accordion-like honeycomb scaffolds from PGS. These scaffolds
were physically and mechanically anisotropic, closely mimicked the mechanical properties of the
right ventricular myocardium of adult rats, and promoted the anisotropic alignment of cardiac
cells that beat in synchrony when electrically stimulated. Guan et al. [203] applied a thermal
gradient during the thermally induced phase separation (TIPS) process to form physically and
mechanically anisotropic scaffolds from an elastase-sensitive segmented polyurethane. The
oriented scaffolds supported muscle-derived stem cell growth and were completely degraded in 8
weeks after subcutaneous implantation in rats. Work in the same group similarly formed
physically and mechanically anisotropic scaffolds by electrospinning a different biodegradable
38
PU onto a rotating collection mandrel [204]. Using this technique, Rockwood et al. [193]
cultured neonatal rat ventricular cells on aligned and unaligned electrospun PU scaffolds. Both
scaffold architectures supported the culture of the cardiac cells, but the aligned scaffold
promoted cardiomyocyte alignment and significantly decreased expression of atrial natriuretic
peptide (ANP), indicative of a more mature ventricular phenotype. Others have similarly used
electrospinning to form anisotropic scaffolds for culturing cardiac cells [205] and this approach
appears to have much potential in cardiac tissue engineering, discussed in greater detail in
section 2.6. The combination of the appropriate elastomeric biomaterial and scaffold fabrication
technique for an anisotropic organization may ultimately lead to ideal tissue engineering
scaffolds that have the appropriate physical, chemical, mechanical, and degradation properties
required to promote the regeneration of an infarcted heart.
2.4.3 Cells for Cardiac Tissue Engineering As discussed in the section on cellular cardiomyoplasty above (section 2.3.2), the
uncertainty in a suitable source of de novo cardiomyocytes is a limitation of cardiac tissue
engineering. The various cell types used in cellular cardiomyoplasty that have demonstrated
potential in cardiac repair may hold equal or greater potential when combined with biomaterial
scaffolds due to better cell survival, increased cell engraftment area, and associated benefits from
the biomaterials themselves. Still, most studies in the cardiac tissue engineering field have been
conducted using primary cardiac cells as a proof-of-principle in establishing the potential of
various techniques and parameters needed for developing engineered myocardial tissue.
Embryonic stem cells and induced-pluripotent stem cells appear to be the best source of bona
fide cardiomyocytes and are beginning to be used in combination with the different myocardial
tissue engineering strategies.
While the search for a source of cardiomyocytes continues, an increasing body of
evidence suggests the other cardiac cells that are critical in developing and maintaining normal
cardiac structure and function in vivo also play an important role in engineering cardiac tissue.
Naito et al. [135] demonstrated a native heart cell population was better than an enriched
cardiomyocyte population in the formation of EHT. Implanted cell sheets formed from the
coculture of endothelial cells and cardiomyocytes enhanced neovascularization, with endothelial
cells contributing to formation of new vessels, and may reduce periods of ischemia of
transplanted tissue [156, 160]. In regards to using preformed 3-D constructs, coculturing
39
cardiomyocytes and cardiac fibroblasts promoted synchronous contractions and a more
organized and mature cardiomyocyte phenotype, potentially due to fibroblast secretion of growth
factors, compared to constructs made by cardiomyocytes alone [206]. Cardiac fibroblasts have
been shown to promote remodeling of a preformed collagen matrix and aid in the formation of
compact cardiac-like tissue when cocultured with cardiomyocytes towards engineering cardiac
organoid chambers [207]. In this study, a compact tissue allowed the cardiac organoid chambers
to exhibit several physiological characteristics of cardiac pump function, including
electromechanical coupling, translation of wall tension into pressure, generation of positive
stroke work, a functional Frank–Starling mechanism, and a positive inotropic response to
calcium, thus leading to a tissue engineered model for studying ventricular function, injury and
repair. In addition, recent work in the same group demonstrated that coculturing neonatal
cardiomyocytes with cardiac fibroblasts, but not foreskin fibroblasts, augments the
electromechanical function of engineered myocardial tissue through a combination of improved
cellular structure and alignment and enhanced electrical conduction, potentially through direct
fibroblast-myocyte gap junction formation [208]. The pre-treatment of synthetic elastomeric
scaffolds with cardiac fibroblasts has also been shown to create a supportive environment for
cardiomyocyte attachment, differentiation, and contractile function [209]. This increasing body
of knowledge clearly identifies the importance of non-myocytes in engineering myocardial tissue
and is an important design criterion for future studies in this field.
2.4.4 Seeding and Cultivation Parameters for Cardiac Tissue Engineering The last main component for developing engineered tissue using pre-formed 3-D
scaffolds, which has equal significance for the other approaches to cardiac tissue engineering, is
the seeding and cultivation parameters required to develop constructs with morphological,
functional, and mechanical properties similar to native tissue. These techniques are more
directed at the development of functional cardiac tissue in vitro for subsequent implantation or as
models for studying development, disease, and regeneration, but some may also apply to
transplantation of cells for in situ tissue formation. In both of these approaches, biomaterial
scaffolds are used as transport vehicles to deliver cells into the body to help regenerate tissues.
High scaffold porosity is a tissue engineering requirement to achieve the transportation of a high
density of cells but if the seeding technique isn’t conducive of promoting cell infiltration, then
this won’t be realized. Towards this goal, it was determined that a high cell seeding density
40
(>1x105 cells/mm3) in combination with medium perfusion promotes a high density of uniformly
distributed cells in the scaffolds [168, 210]. Other dynamic seeding methods, including
centrifugation, pulling a vacuum, or applying pressure to push cells in the scaffolds, may also
work in achieving this goal [165, 194].
Once the cells have been seeded, if they are not to be implanted directly, bioreactors and
other cultivation conditions may aid in formation of the tissue. Medium perfusion (0.5-1.5
ml/min) during cultivation promotes the distribution of cardiac cells, oxygen, and nutrients
throughout scaffolds resulting in thick functional tissue ~1.5-2 mm thick [168, 210-213].
Perfusion of culture medium provides convective and diffusive transport of oxygen and nutrients
to all parts of the construct allowing for aerobic metabolism to take place in the scaffolds away
from the peripheral surface layer, >200 µm [211-213]. The ability for cells to undergo aerobic
metabolism throughout the scaffold promotes spatially uniform cell distributions, cell viability,
expression of cardiac markers, and cardiomyocyte function [211-213]. Cultivation of tissue
engineered cardiac constructs under cyclic mechanical strains has also been shown to improve
tissue engineered constructs [129, 131, 134, 214]. As mentioned above, external mechanical
stimuli trigger intracellular signaling pathways and are critical for the development of cardiac
tissue. Mechanical stimulation induces normal myofibril formation and organization,
electromechanical coupling, cell hypertrophy, tissue compaction and orientation, and improves
contractile function of engineered tissue [129, 131, 134, 214]. Interestingly, in a study
performed with cultured fibroblasts, it was found that anisotropic tissue constructs may
overcome contact guidance cues and remodel in response to mechanical signals [215], and it
would be interesting to see if this is true of cardiac constructs as well. Electrical pacing of
cardiac constructs has also been shown to improve tissue formation and function [216]. In
contrast to mechanical signals, though, physical guidance cues from surface topography
dominate in directing cell organization and elongation compared to electrical field stimulation
[217].
While these different construct cultivation methods are important for the development of
engineered tissue, one caveat of their use for in vitro formation and subsequent transplantation is
the increased susceptibility to ischemia due to thicker tissues and greater cell maturity. The issue
of quick vascularization of implanted tissue remains a great challenge in the tissue engineering
field and must be resolved if approaches other than polysurgery are going to be used in the
41
clinical setting. Advances in ischemic preconditioning and oxygen microcarriers hold potential
in helping to resolve this issue [218, 219], but timely vascularization will still be required for
successful transplantation of the engineered tissue.
2.5 Biodegradable Segmented Polyurethanes for Tissue Engineering Segmented polyurethanes represent an important class of synthetic polymers for tissue
engineering applications. A significant advantage of PUs in comparison to other biomaterials is
the flexibility in chemistries used in the synthesis process allows the development of polymers
with diverse physical, chemical, mechanical, and degradation properties. A priori knowledge of
the particular properties needed for site specific applications enables researchers to tailor the PU
properties to meet these requirements. Moreover, the ability to incorporate specific peptide
sequences into the backbone structure of the polymer confers unique biological functionality to
these synthetic materials. Thus, the flexible PU chemistry provides the opportunity of building
synthetic polymers that give both researchers and cells some control over the properties and
performance of these materials. The formation of bio-mimetic synthetic polymers with the
ability to adjust the material properties as new design criteria are identified could have important
implications in the development of ideal biomaterial scaffolds for tissue engineering. As such,
recent years have seen a large expansion in the number of new biodegradable polyurethanes
available for regenerative medicine and this section will focus mostly on these.
2.5.1 Chemistry and Properties of Degradable Polyurethanes Polyurethanes are a heterogeneous family of polymers used in a variety of biomedical
applications due to their diverse material properties and good biocompatibility [220]. Segmented
polyurethanes are thermoplastic block copolymers of the (AB)n type consisting of alternating
sections of hard segments, composed of a diisocyanate and a low molecular weight diol or
diamine chain extender, and soft segments, generally composed of either polyethers, polyesters,
polycarbonates, or polyalkyldiols [220]. Segmented polyurethanes are characterized by a
thermodynamically driven microphase separation, resulting from incompatibilities between hard
and soft segments [221]. The degree of phase segregation is based on the morphology of the
different segments, which may appear as isolated or interconnected hard segments in a
continuous soft segment matrix (Figure 2.4) [222]. The actual morphology, however, depends
on the relative amounts of hard and soft segments and is affected by various factors. Hard
42
segment regularity, rigidity, inter-segment interactions, and the presence or absence of bulky side
groups all have a major effect on hard segment interconnectivity and phase segregation [223].
The segregated morphology is also contingent on hard and soft segment length, polarity,
crystallinity, overall composition, and mechanical and thermal history [224-226].
Figure 2.4: Illustration of microphase separation in segmented polyurethanes.
For segmented polyurethanes, microphase separation impacts mechanical and
degradation properties, two important factors that influence the performance of biomaterial
scaffolds. Increased phase segregation generally results in improved mechanical properties due
to hard segments acting as physical crosslink sites and soft segments adding some flexibility to
the polymer [227]. The elastic nature of segmented polyurethanes is a consequence of the
thermodynamic incompatibility that serves as the driving force for phase separation [221]. When
the polymers are stretched, an energetically unfavorable phase mixing occurs which is restored
when released. The design of elastic polyurethanes therefore requires the appropriate chemistry
to promote phase segregation [228, 229]. Phase separation, however, may also impact the
degradation rates of the polymer. Increased phase segregation, for example, has been shown to
protect susceptible functional groups present in hard blocks from hydrolytic agents [230, 231].
Similarly, increased phase separation promotes uninterrupted packing of polymer molecules
leading to semicrystalline domain formation. Water and other hydrolytic agents are restricted
43
from reaching labile bonds within the semicrystalline domains and reduces the rate of
degradation compared to amorphous domains [232]. As a result, high phase segregation may
lead to increased PU stability and relatively slow degradation rates. Biodegradable elastomeric
PUs designed for tissue engineering applications should therefore promote phase segregation for
achieving elastic mechanical properties while still allowing access to labile bonds for
degradation to occur at the appropriate rate. In an aqueous tissue environment where hard
segment surface enrichment occurs, the introduction of degradable sequences into the chain
extender chemistry is a logical approach in achieving the balance between elastic mechanical
properties and reasonable degradation rates. Ultimately, the reactant chemistries must be chosen
to obtain the appropriate properties for the site specific applications.
2.5.1.1 Segmented Polyurethane Synthesis Segmented polyurethane synthesis is a step growth polymerization occurring in a two
step process (Figure 2.5). In the first step, the diisocyanate is reacted with the soft segment
polyol to form the prepolymer. Here, the characteristic urethane linkages are formed through the
reaction between the isocyanate groups and the hydroxyl-terminated end groups of the polyol. In
the second step, the low molecular weight chain extender is used to link the prepolymer
segments yielding a high molecular weight polymer. Terminal isocyanate groups from the
prepolymer react with terminal amines of diamine chain extenders to form urea groups, the end
product being a polyurethaneurea. Alternatively, if a diol chain extender is used, additional
urethane functional groups are formed leading to a polyurethane. Throughout this report,
polyurethaneureas will often be referred to simply as polyurethanes, but it is important to note
the distinction in chemistry as it has an important influence on polymer properties and the
mechanisms for observed behavior. Urea groups, for example, have an additional N-H group
that may act as a donor in hydrogen bond formation. Hydrogen bonding is extensively observed
in PUs and plays a role in inter- and intra-molecular interactions, cohesiveness of the polymer,
and may be a driving force for phase segregation [220]. Consequently, polyurethaneureas have
improved physical and mechanical properties compared to equivalent diol extended
polyurethanes [220].
44
Figure 2.5: Standard two-step segmented polyurethane reaction. Step 1: Prepolymer formation.
Step 2: Chain extension.
During polyurethane synthesis, several side reactions may occur leading to branching,
crosslinking, or changes to the stoichiometric balance of reactants [220]. The presence of water
causes isocyanate groups to form unstable carbamic acids, which subsequently decompose to
amines with the liberation of CO2 gas. These newly formed amines may then quickly react with
isocyanates, thus changing reactant stoichiometry and leading to lower molecular weight
polymers. Isocyanates may also react with carboxylic acids, amides, urethanes, and ureas.
Undesirable branching and crosslinking may occur at elevated temperatures between isocyanates
and urethanes and/or ureas of the growing polymer chains. To minimize these side reactions,
PUs are synthesized under anhydrous conditions at temperatures below 100ºC.
2.5.1.2 Reactant Chemistry for Biodegradable Polyurethanes The flexibility in chemistry that gives PUs their diverse properties is a consequence of
having several choices in diisocyanates, soft segment polyols, and chain extenders for PU
synthesis. The diisocyanates and polyols used in making biodegradable PUs are typically those
that are commercially available, while the chain extenders appear more diverse and unique. The
structural composition of the reactants as well as the synthesis and processing conditions dictate
polymer morphology, phase segregation, and chain interactions and affect overall polymer
45
properties. In general, a few points can be made regarding the structure of reactants, the effects
they have on polymer properties, and the choice of chemistries for obtaining an appropriate
polymer. Aliphatic reactants have the most molecular flexibility, then cycloaliphatic groups, and
then aromatic groups, which are more regular and rigid. While some chain mobility is required
to achieve a regular, ordered structure, too much flexibility prevents the formation of crystalline
structures, inter-chain interactions, and domain cohesion [233]. As a result, polyurethanes
composed of aromatic groups in the backbone structure generally have enhanced mechanical
properties due to improved phase segregation and domain cohesion, while aliphatic PUs are
generally softer and weaker [220, 234, 235]. Similar to high molecular flexibility, asymmetric
reactants and bulky side chains prevent the alignment of polymer segments thus reducing
interchain interactions, increasing phase mixing, and leading to reduced mechanical properties
[235, 236]. As mentioned above, additional hydrogen-bond forming groups in
polyurethaneureas equates to improved phase segregation and mechanical properties compared
to diol chain-extended polyurethanes. As a consequence, having some knowledge about how
reactant chemistry will affect polymer properties allows the selection of complimentary
components to obtain PUs with appropriate properties for specific applications. For example, if
an aliphatic diisocyanate that has bulky side chains is to be used, then a symmetric diamine chain
extender that isn’t too flexible may help to achieve a moderately phase segregated PU with soft,
flexible mechanical properties.
The commonly used, commercially available diisocyanates for synthesizing
biodegradable polyurethanes are shown in Figure 2.6. Most of the diisocyanates are aliphatic or
cycloaliphatic, with the exception of 4,4’-diphenylmethane diisocyanate (MDI). Our laboratory
has been particularly interested in synthesizing biodegradable PUs using 2,6 diisocyanto methyl
caproate, a lysine-based diisocyanate (LDI). LDI is an aliphatic molecule that has an asymmetric
methyl ester side chain and polyurethanes created using this diisocyanate reflect its aliphatic
asymmetric nature. Segmented PUs containing LDI had lower molecular weights and no hard
segment glass transition temperature, indicative of reduced hard segment cohesion, when
compared to an equivalent PU composed of the symmetric aliphatic 1,6-diisocyanatohexane
(HDI) [238, 239]. Predictably, Caracciolo et al. [236] recently demonstrated that the mechanical
properties of PUs formed with LDI were weaker than those made with HDI and was contributed
to reduced phase segregation due to LDI asymmetry. Despite the reduced physical and
46
Figure 2.6: Diisocyanates used to synthesize biodegradable PUs. Image used with permission from Guelcher [237].
mechanical properties of LDI-based PUs, the major appeal of it comes from the non-toxic
degradation products that are associated with its liberation in the body. This reagent is derived
from the amino acid L-lysine and if hydrolysis of the urethane linkages generate LDI in vivo, this
diisocyanate will react with water to form L-lysine methyl ester, a non-toxic product [240].
Biodegradable PUs formed using LDI found no toxic or tumorigenic responses to these materials
upon implantation into animal models [241-243]. In addition, the methyl ester side chain of LDI
provides a functional group that may be modified to attach biological agents to the surfaces of
the PUs. Ernsting et al. [244] utilized the side chain of LDI in attaching the natural anti-oxidant
vitamin E to surface modifying macromolecules as a model system for bringing bioactive
molecules to the surfaces of PUs. This technique was developed in an attempt to promote PU
stability, but has potential in conjugating biological agents relevant to tissue engineering, such as
the cell adhesion peptide RGD or various growth factors.
The polyols generally used as the soft segment of biodegradable PUs are polyethers,
polyesters, polyalkyl diols, or a combination of these. Figure 2.7 shows the chemical structure of
some common soft segments for biodegradable PUs. Polyols are responsible for the flexibility
and elongation limit of the PUs and given that they make up roughly 50-75% of the material, the
choice of soft segment often dictates the properties of the overall PU. Degradable PUs made
from polylactide, for example, are generally stiff, have a high strength, and are not very
47
Figure 2.7: Polyols often used in biodegradable PU synthesis. Image used with permission from Guelcher [237].
extensible, similar to polylactide homopolymers [245, 246]. For elastomeric biodegradable PUs,
polycaprolactone (PCL) soft segments lead to the formation of highly extensible, soft, elastic
polymers [239, 245, 246] and are particularly attractive for soft tissue applications. Skarja and
Woodhouse [239, 247] formed PUs with PCL and PEO soft segments of varying molecular
weight and found the mechanical and degradation properties of the PUs were directly related to
soft segment composition and molecular weight. All the PCL-based PUs were elastomeric, but
an increase in crystallinity, hydrophobicity, initial modulus, ultimate tensile stress and strain, and
resistance to degradation was observed with an increase in starting PCL molecular weight. PEO-
based PUs, on the other hand, were amorphous, soft, and tacky polymers that were much more
hydrophilic and had faster degradation rates than the PCL-based PUs [239, 247]. To take
advantage of good elastomeric mechanical properties of PCL and degradation characteristics of
PEO, several groups have developed a series of biodegradable segmented PUs with triblock soft
segments of PCL-PEO-PCL [51, 196, 248, 249]. A wide range of elastomeric PUs with different
mechanical properties and degradation characteristics were observed in these triblock soft
segment-based PUs with similar trends in PU properties with changes in PCL and PEO
molecular weights as observed by Skarja and Woodhouse. Thus, PUs with varying elastomeric
mechanical properties and degradation characteristics may be formed by adjusting the chemistry
and molecular weight of soft segment polyols.
A wide range of difunctional reactants can be used as the chain extender in synthesizing
biodegradable PUs, but the majority of these are diols and diamines. While commercially
48
available diisocyanates and polyols are typically used, several groups have developed novel
chain extenders as a method of introducing degradable chemical moieties into the hard segment
of these polymers [51, 196, 238, 239, 247, 250-252]. Using this approach allows the choice of
soft segment polyols for achieving specific physical and mechanical properties while still
promoting overall polymer degradation. Surface enrichment of the hard segment in aqueous
tissue environments also promotes accessibility of labile bonds to various hydrolytic agents, thus
enabling polymer degradation. In addition, amino acids and short peptides can easily be
incorporated as chain extenders either directly, by adding a lysine residue on the end of desired
peptide sequences, or using linker molecules to create diamine reactants [51, 238, 251-253]. As
a result, PUs can be developed that are susceptible to cell-secreted proteases or exhibit other
unique biological functionality. The ability to create synthetic bio-mimetic PUs with very
diverse physical, chemical, mechanical, and degradation properties provides much promise for
segmented PUs in tissue engineering applications.
2.5.2 Polyurethane Degradation The diverse properties that may be achieved by segmented polyurethanes make them
attractive candidates for use in several biomedical applications. Traditionally, these synthetic
polymers have gained use as long-term implants in pacemaker lead insulation, cardiac assist
devices, breast implants, and others [220]. In these applications, polymer stability was critical to
the success of the device. The observation that these PUs were degrading in vivo leading to
device failure prompted many studies in trying to understand the mechanisms behind this PU
degradation. While most of this work was directed towards the development of more bio-stable
PUs, this knowledge is equally important in tissue engineering applications where these
mechanisms may be exploited in achieving specific PU degradation characteristics.
There are several mechanisms in which polymers degrade including by thermal,
radiation, mechanical, and chemical means [254]. All of these mechanisms may have some
effect on polyurethanes, for example during production, processing, and sterilization, but
mechanical and chemical degradation will have the most significant effects on PU degradation in
vivo. “Environmental biodegradation” is a term given by Santerre et al. [255] for describing the
multifaceted mechanical and chemical mechanisms that work synergistically in degrading PUs in
the body (Figure 2.8). Environmental biodegradation encompasses the degradation resulting
from environmental stress cracking, the aqueous tissue environment, hydrolytic enzymes,
49
oxidative agents, calcification, mechanical stresses, metallic ions, salts, acids, and cells and is a
function of PU chemistry, morphology, and processing [255]. The complex interplay of all the
different factors makes a complete mechanistic understanding of PU degradation difficult, but
insight can be gained by understanding the contribution of the individual components on their
own.
Figure 2.8: Model for environmental biodegradation of PUs. Image used with permission from Santerre et al. [255].
Biodegradable PUs used in the mechanically demanding cardiac environment are
exposed to repetitive stresses that may influence the degradation of the material. A stress of at
least 60 kPa is generated in the myocardium per heartbeat [179] and this is repeated
approximately 75 times per min in a normal resting heart. The cyclic mechanical stresses
experienced during the cardiac cycle can physically cleave bonds within polymer chains,
resulting in polyurethane fatigue [220]. Similarly, strains placed on the PU can alter the
structural morphology and phase segregation of the polymer [255]. This may expose labile
bonds to various hydrolytic agents that may have otherwise been protected, thus promoting
degradation. Strain induced crystallization may also occur in the heart through the organization
of polymer chains under stress [220]. Although crystallization may reduce PU degradation, it
may also alter the mechanical properties making it stiffer and potentially unable to withstand
subsequent stresses leading to structural failure. Environmental stress cracking through the
formation and propagation of cracks in the PU may also occur in the mechanically active heart.
Environmental stress cracking and PU degradation result from the combination of mechanical
stresses from the biological environment or residual effect of processing and several other factors
50
such as oxidative agents from inflammatory cells, the presence of foreign body giant cells
(FBGC) and monocyte-derived macrophages (MDM), polyether soft segments, PU morphology,
and various proteins including α2-macroglobulin [256, 257]. Polyether-based PUs and polymers
with inappropriate mechanical properties, such those that are not elastomeric, are too hard or too
soft, or show significant hysteresis, may have increased susceptibility to mechanical-based
degradation in the heart and may lead to inappropriate degradation rates for specific applications.
Chemical degradation of polyurethanes is mediated by oxidation and hydrolysis. One of
the most important processes associated with chemical degradation of polyurethanes in vivo is
inflammation. Inflammatory cells, such as neutrophils (polymorphonucleocytes, PMNs) and
MDMs, are recruited to the site of biomaterial implant. During the initiation of inflammation,
PMNs are the first cell type to arrive, but shortly after MDMs and FBGCs, formed by the fusion
of macrophages, take over and are the main cell types involved in biomaterial degradation and
chronic inflammation [258]. Inflammatory cells contain several reactive oxygen species that are
released during FBGC formation and chronic inflammation, including superoxide, hydrogen
peroxide, hypochlorous acid (HOCl), and peroxynitrite anion (ONOO-) [29, 259, 260].
Sutherland et al. [257] investigated the effects of neutrophils, HOCl, and ONOO- on
polyetherurethanes. It was found that all three oxidative factors induced polymer degradation,
with HOCl reacting with urethane-aliphatic ester linkages and ONOO- targeting aliphatic ether
groups. Other studies have similarly shown these reactive oxygen species released from
inflammatory cells contribute to the degradation of synthetic polymers [261].
Hydrolysis is another mechanism of chemical degradation involved in polyurethane
degradation in vivo and can be either passive or enzyme-mediated. In general, polyurethane
physiochemical properties, including type of chemical bonds, chemical composition,
hydrophilicity, crystallinity, phase segregation, and porosity, all affect the rate of hydrolysis
[262]. Local changes in pH may arise from hydrolysis through the formation of new functional
groups during chain cleavage and this may also influence degradation [263]. Importantly, the
release of several hydrolytic enzymes by MDMs and FBGCs during inflammation led to in vitro
studies investigating the effects of these cells and enzymes on the stability of polyurethanes
[264]. In an early study, Santerre et al. [265] identified a polyester urea-urethane susceptibility
to cholesterol esterase, an enzyme released by MDMs. Follow-up studies suggested cholesterol
esterase hydrolyzes urethane linkages in the soft segments of phase segregated polyurethanes
51
[266, 267] and that the relative amount of hard segment correlates with the degree of
biodegradation [230]. Specifically, an increased hard segment content was indirectly
proportional to hydrolytic degradation in three polyether-urea urethanes even though the hard
segments contained hydrolysable groups [230]. Similarly, the level of enzymatic hydrolysis is
inversely proportional to hydrogen bonding between hydrolysable bonds in the hard and soft
segments due to domain cohesion and the inability of enzymes to reach the labile bonds [268].
Other MDM-released esterases have also been shown to play a role in PU degradation adding to
the important role of MDMs in understanding the mechanisms behind environmental
biodegradation of PUs [255]. For segmented PUs used as biomaterial scaffolds for cardiac
tissue engineering, a combination of cyclic stresses, inflammatory cells, oxidative agents,
inflammatory and injury-related enzymes, as well as other factors will likely contribute to the
environmental biodegradation of these synthetic polyurethanes.
The knowledge gained in elucidating environmental biodegradation of PUs allows the
rational design of novel biomaterials for biomedical applications. MDM-induced degradation of
PUs, for example, has been exploited by Santerre and colleagues in the development of novel
antimicrobial-containing PUs for medical devices [255, 269]. The antimicrobial agents are
incorporated into the backbone structure of the PUs through labile bonds that have been shown
to be susceptible to MDM-secreted enzymes and are released as a direct consequence of polymer
degradation through inflammatory cells recruited during device implantation and infection. As a
consequence, the drug release profiles are dictated by the tissue in a time-course that responds to
the tissue environment and directly corresponds with tissue healing [255, 269]. By incorporating
specific chemical moieties into the backbone structure of PUs, a similar tissue-responsive
degradation mechanism may be exploited by cell-secreted proteases that are not traditionally
associated with PU degradation and is the subject of the next section.
2.5.3 Enzyme-Degradable Polyurethanes The mechanisms behind the biodegradation of PUs in vivo suggest potential pathways
through which to promote degradation for tissue engineering applications. While these
mechanisms will play a role in the resorption of PU scaffolds, PU degradation can be further
promoted by exploiting the presence of cell-secreted proteases that are not traditionally involved
in PU degradation but are present in the microenvironment in which biomaterial scaffolds may
be beneficially employed. This has led to the development of novel enzyme-sensitive
52
elastomeric segmented polyurethanes by incorporating amino acids and peptides into the
polymer backbone and has much potential as bio-mimetic biomaterials in soft tissue engineering.
Skarja and Woodhouse developed a novel family of enzyme-sensitive biodegradable
polyurethane elastomers using an L-phenylalanine-based diester chain extender [238, 239]. In
this work, 1,4-cyclohexane dimethanol was used to link phenylalanine (Phe) residues to form a
diamine, diester chain extender. It was conceived that this approach would produce non-toxic
degradation products that could be readily metabolized in vivo and the choice of amino acid or
peptide could be tailored for hydrolysis by specific enzymes. Phenylalanine was incorporated
into the polyurethane backbone to introduce enzymatic susceptibility to the polymer by
chymotrypsin-like serine proteinases found in the body, such as cathepsin G and chymase [247].
Segmented PUs synthesized with the Phe-based chain extender showed enhanced degradation in
the presence of chymotrypsin, and to a lesser extent trypsin, and could achieve a range of
physical, chemical, mechanical, and degradation properties by changing the soft segment
chemistry [239, 247]. One polymer from this family, synthesized from PCL of molecular weight
1250, LDI, and the Phe-based chain extender (Phe PU), has been successfully used in cardiac
tissue applications. Films of this PU supported a differentiated cardiac phenotype and contractile
function using neonatal rat cardiomyocytes [191]. Similarly, mESCDCs maintained a normal
cardiomyocyte morphology, expressed the cardiac specific markers myosin heavy chain, desmin,
and α-sarcomeric actinin, and were contractile on ECM-coated Phe PU films [192]. Recently,
porous 3-D scaffolds formed by thermally induced phase separation and electrospinning were
used with mESCDCs and neonatal rat cardiac cells with similar positive results [193, 194].
Extensions of this work, described in this thesis, may provide further evidence of the potential of
the Phe PU and a newly developed dipeptide-containing PU for cardiac tissue engineering.
Biomaterial scaffolds that seek to mimic the biological structure and function of the
native ECM should also degrade and reorganize through the same mechanisms. Synthetic
hydrogels were developed to target polymer degradation by ECM degrading cell secreted
proteases [50]. Guan and Wagner [51] extended this work in the formation of elastomeric
segmented polyurethanes that were susceptible to degradation by the ECM degrading enzyme
elastase. In this work, the tri-peptide Ala-Ala-Lys (AAK) was used as a chain extender to
facilitate enzyme-mediated degradation of hard segments. Elastase is known to cleave peptides
between alanine residues and the lysine residue was added to generate a diamine structure [51].
53
PUs were synthesized using the peptide chain extender, 1,4-butanediisocyanate, and either PCL
or PCL-PEO-PCL triblock soft segments of varying block molecular weights. The resulting PUs
were flexible elastomers with a range of mechanical properties, breaking strains of 670-890%
and tensile strengths of 15-28 MPa, and showed enhanced degradation in the presence of elastase
through incorporation of the Ala-Ala cleavage site [51]. In subsequent work by this group, an
elastase-sensitive PU was formed into anisotropic and randomly oriented 3-D scaffolds by
thermally induced phase separation [203]. Scaffold architecture and physical and mechanical
properties were a function of the fabrication conditions. The oriented scaffolds exhibited
anisotropic mechanical properties with a much higher breaking strain and ultimate tensile
strength in the longitudinal direction compared to the transverse direction or random scaffolds
and better supported muscle-derived stem cells than random scaffolds [203]. Of particular note
were the degradation characteristics of this elastase-sensitive PU scaffold. In vitro degradation
demonstrated the PU scaffold had similar passive hydrolysis as films of the same material after 8
weeks (~12% mass loss), but showed significantly enhanced mass loss in the presence of elastase
(~42% for scaffold vs. 27% for film), likely due to increased surface area and accessibility of
elastase to labile Ala-Ala peptide bonds [203]. Significantly, subcutaneous implantation of the
elastase-sensitive PU scaffold showed almost no signs of the scaffold after 8 weeks, whereas a
comparable PU scaffold with the elastase insensitive chain extender putrescine was still present,
albeit with considerable degradation [203]. Although the exact mechanism behind this in vivo
degradation work is difficult, it highlights a few critical aspects of PU degradation. The
significant difference in degradation seen from the 8 week in vitro study compared to that done
in vivo suggests the in vitro models with enzymes alone, while providing useful mechanistic
information, cannot accurately replicate the complex in vivo environment. A higher elastase
concentration in vivo than was used in the in vitro studies may have contributed to enhanced
degradation, but more likely reflects the multifaceted mechanisms of environmental
biodegradation discussed above. Similarly, the AAK peptide sequence promotes specific
elastase-mediated PU degradation but may also confer susceptibility to a variety of other
enzymes or PU biodegradation mechanisms irrespective of the specific peptide sequence [203].
Thus, the incorporation of peptide sequences into the polymer backbone structure may have a
primary effect of promoting specific enzyme-mediated PU degradation and a secondary effect of
conferring susceptibility to the other mechanisms of environmental biodegradation of PUs. The
54
Wagner lab has shown the putrescine extended PUs have favorable properties for cardiac tissue
engineering applications [197, 198] and it will be exciting to see how the enzyme-sensitive PUs
function in the heart.
2.6 Electrospinning for Tissue Engineering Scaffold Formation Electrospinning is a technique that may be used to form non-woven 3-D scaffolds from
various natural and synthetic biomaterials for tissue engineering applications [270]. This
relatively simple, cost effective, and versatile method allows the processing of polymers into
fibers with diameters on the nanometer and micrometer scale, often with control of the specific
range. The resulting scaffolds have a topography and porosity that mimics the native ECM and
has been shown to influence cell behavior [270, 271]. The electrospinning process may also be
used to form anisotropic scaffolds with fibers oriented in one direction [272], an important
design criterion for engineering anisotropic tissues such as that found in the myocardium.
Electrospinning therefore is a promising technique in forming anisotropic 3-D scaffolds from
biodegradable elastomeric biomaterials for cardiac tissue engineering.
2.6.1 Principles and Parameters Electrospinning uses electrostatic forces to process polymers into fibers. Figure 2.9 is a
schematic that illustrates the electrospinning technique and apparatus setup. A polymer solution
is fed to the end of a syringe needle via a mechanical syringe pump forming a polymer droplet at
the needle tip. Under an applied electric field, an electrostatic attraction between the oppositely
charged polymer solution and collection plate combined with an electrostatic repulsion from
similar charges within the liquid cause the droplet to move from a rounded meniscus to a cone-
shaped arrangement, called the Taylor cone [270]. A polymer jet is formed from the Taylor cone
when the electrostatic charge of the polymer solution overcomes the surface tension of the
droplet. As the ejected polymer jet approaches the target, the solvent evaporates and continuous
electrical forces cause stretching on the entangled polymer chains resulting in a decreased jet
diameter and fiber formation [273]. A single polymer fiber is typically pulled from the needle
and deposited on the collection plate, but polymer jet splaying may cause the polymer fiber to
split into two or more fibers leading to the deposition of several polymer fibers simultaneously
[273]. By altering several fabrication parameters, scaffolds can be formed with varying
porosities, fiber diameters, pore sizes, fiber morphologies, thicknesses, and architectures [270,
55
272, 274]. If electrospinning parameters are not appropriate for fabricating continuous fibers,
such as inappropriate polymer molecular weight, concentration, or viscosity, polymer droplets
are deposited instead in a process called electrospraying [275]. Both electrospinning and
electrospraying have been used in tissue engineering applications [276].
Figure 2.9: Schematic of electrospinning apparatus. Image used with permission from Kenawy et al. [277].
Many different electrospinning processing parameters can influence scaffold properties
including solvent type, polymer concentration, flow rate, polymer molecular weight, dielectric
constant, voltage, syringe needle tip design, distance to collector, ambient conditions, presence
of additives, and collector composition, geometry and motion [270, 272, 274]. Table 2.3
summarizes the effect that several of these parameters have on fiber morphology. Adjusting the
various parameters can lead to polymer morphologies such as droplets, beads-on-a-string, and
round, flat, porous, or fused fibers [270, 272, 274]. As well, the fibers may have diameters
ranging from a few hundred nanometers to several microns and may be randomly oriented or
aligned to various degrees. The diverse polymer morphologies that may be formed by altering
the electrospinning parameters allows the formation of scaffolds that exhibit a variety of
physical, mechanical, and degradation properties that will influence the performance of the
scaffold in the presence of cells and in vivo.
56
Table 2.3: Summary of electrospinning parameters and effects on fiber morphology. Used with permission from
Murugan and Ramakrishna [272]. Processing Parameter Effect on Fiber Morphology Concentration/viscosity • Low concentrations/viscosities yield beads or beads-on-a-string
morphology; increasing concentration/viscosity reduces beaded defects • Fiber diameters increase with increasing concentration/viscosity
Conductivity/solution charge density
• Increasing conductivity aids in production of uniform bead-free fibers • Higher conductivities yield smaller fibers with a few exceptions
Polymer molecular weight • Increasing molecular weight reduces number of beads and droplets Dipole moment and dielectric
constant • Successful spinning can be achieved in solvents with high dielectric
constants Flow rate • Lower flow rates yield fibers with smaller diameters
• High flow rates produce fibers that are not dry upon reaching collection plate
Field Strength/voltage • At high voltages, beading is observed • No direct correlation with field strength and fiber diameter
Distance between needle tip and collector
• A minimum distance is required to obtain dry fibers • At distances too close or too far, beading is observed
Needle tip design • Using a coaxial, 2-capillary spinneret, hollow fibers are produced • Multiple needle tips are used to increase throughput
Collector composition and geometry
• Smooth fibers result from metal collectors; porous fiber structures are obtained using porous collectors
• Aligned fibers are obtained by using a conductive frame, rotating drum, or wheel-like bobbin collector
Ambient parameters • Increasing temperature reduces viscosity and leads to smaller fiber diameters
• Increasing humidity causes circular pores to appear on fibers
2.6.2 Electrospun Scaffolds for Cardiac Tissue Engineering The electrospinning technology is a promising approach to forming 3D polymeric
scaffolds that resemble the size scale and organization of the natural ECM. Natural materials
[278-280], synthetic polymers [277, 281], and combinations of the two [282] have been
electrospun into scaffolds for tissue engineering applications. Importantly, thermoplastic
segmented PUs may be processed by conventional solvent-based methods and can be used with
the electrospinning technique [237]. As a result, several different biodegradable PUs have been
made into scaffolds using this process [276, 282-287]. Rockwood et al. [286], for example,
successfully electrospun the enzyme-susceptible, elastomeric Phe PU to form scaffolds with a
range of fiber sizes from several hundred nanometers to tens of microns. In this work, the
electrospinning process did not affect the molecular weight averages, thermal properties, or
chemical composition of the polymer, but enhanced degradation compared to films due to
increased surface area. In the Wagner lab, Stankus et al. [282] used a biodegradable, elastomeric
PU mixed with type I collagen to electrospin elastic matrices. The resulting mats had randomly
57
oriented fibers ranging from 100 to 900 nm in diameter and supported smooth muscle cells when
seeded on the scaffolds. Interestingly, investigations into the mechanical properties of the
electrospun PU matrices suggested this processing technique did not compromise the materials
mechanical properties and had a tensile strength and distensibility similar to films formed with
the same material [282]. These results indicate that several of the favorable PU properties are
maintained when processed into scaffolds via electrospinning.
Although there are many benefits to forming tissue engineered scaffolds by
electrospinning, it has been recognized that one limitation with this is the fiber mats have
relatively small pore sizes that inhibit cellular infiltration into the scaffold [270]. Nanofibrous
meshes in particular, have small pore sizes and these scaffolds behave more as a 2-D sheet on
which cells can attach, grow, and migrate along rather than a 3-D scaffold where cells can grow
and migrate into [288]. To address this issue, a few studies have been conducted to improve the
ability of attaining cells within a 3-D electrospun scaffold. Nam et al. [289] combined
electrospinning with salt leaching to create thick scaffolds that had deliberate 100-200 µm gaps
that promoted cellular infiltration. They demonstrated that after 3 weeks of culture, cells
infiltrated as far as 4 mm into the scaffold with a cell density of ~70% within the delaminations.
In another approach, Pham et al. [290] observed that cells attached to fibers of different size, but
cell spreading was enhanced by nanofibers whereas cellular infiltration was promoted by
microfibers, which had pore sizes ranging from 20 – 45 µm. To exploit the benefits of each fiber
size, bilayered constructs composed of a layer of nanofibers followed by a layer of microfibers
were utilized in this study with positive results on cell infiltration and spreading. In a novel
study by Stankus et al. [276], smooth muscle cells were electrosprayed concurrently with
electrospinning a biodegradable PU to achieve a high cellular density and infiltration within the
elastic matrix during scaffold formation. It was determined that electrospraying the cells did not
affect viability or proliferation either alone or while electrospinning the polyurethane matrix.
This cellular integration technique followed by culturing in a perfusion bioreactor resulted in
significantly higher smooth muscle cell numbers compared to static cultures and a uniform
distribution and an elongated cell morphology [276]. In addition, by altering the electrospinning
parameters, scaffolds with aligned fibers were formed and cells electrosprayed with these
scaffolds were oriented parallel to fiber orientation, which was independent of perfusion flow
direction [276]. The physical cues provided by fibrous electrospun scaffolds may play an
58
important role in the formation of engineered tissue and the techniques developed for promoting
cellular infiltration may contribute to the success of these scaffolds.
In cardiac tissue engineering, cellular alignment is critical to the structural organization
and functionality of the engineered tissue. Physical cues provided by biomaterial scaffolds may
help in the proper organization of cardiac cells. The ability to form aligned scaffolds from
electrospinning is an important consideration in choosing a scaffold fabrication technique for use
in the heart. Different collection plate designs in the electrospinning apparatus or post-scaffold
processing may lead to electrospun scaffolds with aligned fibers [205, 272]. Using a rotating
mandrel collection plate, Courtney et al. [204] electrospun an elastomeric biodegradable PU into
scaffolds that exhibited a direct relationship between mandrel rotational velocity and fiber
alignment. The aligned scaffolds were both physically and mechanically anisotropic with low
breaking strains in the preferred direction of fiber orientation and high breaking strains in the
cross-preferred direction [204]. The random scaffolds were physically and mechanically
isotropic. Using a similar approach to producing aligned and unaligned biodegradable PU
scaffolds, Rockwood et al. [193] showed the physical cues provided by the electrospun PU fibers
influenced the organization and phenotypic expression of cardiac cells. Specifically, cardiac
cells seeded on the aligned PU scaffold organized along the fibers leading to highly oriented
cells organized parallel to each other that were electrically connected and had a more mature
ventricular phenotype compared to the same cells cultured on unaligned PU scaffolds or tissue
culture polystyrene. Zong et al. [205] fabricated random electrospun scaffolds and post-
processed them by heating and uniaxial stretching to achieve aligned scaffolds. Cardiac cells on
the aligned PLA scaffolds were highly aligned, developed mature sarcomeric structures and
intercalated discs, and were contractile. Several other studies have successfully demonstrated the
efficacy of electrospun scaffolds for culturing cardiac cells and highlight its potential in cardiac
tissue engineering [194, 291, 292].
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Chapter 3: Synthesis and Characterization of Phe and Gly-Leu-containing Segmented Polyurethanes § Large sections of this chapter have been accepted for publication [1]: Parrag IC, and KA
Woodhouse. Development of Biodegradable Polyurethane Scaffolds using Amino Acid and
Dipeptide-based Chain Extenders for Soft Tissue Engineering. Journal of Biomaterials Science –
Polymer Edition, in press.
† Sections of this chapter also contributed to the publication [2]: Rockwood DN, RE Akins, IC
Parrag, KA Woodhouse, and JF Rabolt. Culture on Electrospun Polyurethane Scaffolds
Decreases Atrial Natriuretic Peptide Expression by Cardiomyocytes In Vitro. Biomaterials,
2008, 29(36): p. 4783-4791.
3.0 Abstract Biodegradable segmented polyurethanes (PUs) are promising biomaterials for soft tissue
engineering applications. The inherent flexibility of PU chemistry allows the incorporation of
specific chemical moieties into the back bone structure conferring unique biological function to
these synthetic polymers. Previous work has shown that chain extenders incorporating an amino
acid or short peptide sequence exhibit enhanced degradation in the presence of specific enzymes.
Building on this, the cleavage site of several matrix metalloproteinases, a Gly-Leu dipeptide, was
introduced into a chain extender through the reaction with 1,4-cyclohexane dimethanol. The
Gly-Leu-based diester chain extender was purified by high performance liquid chromatography
and successful synthesis and purification was confirmed by several techniques. PUs were
synthesized with polycaprolactone diol of molecular weight 1250 g/mol, a lysine-based
diisocyanate, and either the Gly-Leu-based diester chain extender (Gly-Leu PU) or a
phenylalanine-based diester chain extender (Phe PU). Both PUs had high molecular weight
averages (Mw > 125,000 g/mol) and were phase segregated, semi-crystalline polymers (Tm ~
42°C) with a low soft segment glass transition temperature (Tg < -50°C). Uniaxial tensile testing
of PU films revealed that the polymers could withstand high ultimate tensile strengths (~ 8-13
MPa) and were flexible with breaking strains of ~ 870-910%. Importantly, the Gly-Leu PU had
85
a significantly higher initial modulus, yield stress and ultimate stress compared to the Phe PU,
suggesting the Gly-Leu-based chain extender allowed for better hard segment packing and
hydrogen bonding leading to enhanced mechanical properties. The newly developed Gly-Leu
PU had several properties that were promising for soft tissue engineering applications and
warranted further investigation into scaffold fabrication and biomaterial testing.
3.1 Introduction The goal of tissue engineering scaffolds is to mimic the native ECM of tissues by
providing a temporary material that plays both a structural and functional role in guiding the
development of tissue constructs for regenerative medicine [3]. The appropriate biomaterial for
these applications must meet several requirements; the material and its degradation products
must be non-toxic and non-immunogenic, the mechanical properties should be appropriate for
the tissue under investigation, degradation should take place at a rate that allows newly secreted
ECM to replace the biomaterial, and physiochemical properties should promote cell adhesion,
growth, and differentiation. In mechanically active soft tissues, such as the heart, mechanical
forces play a critical role in the development and maintenance of the structural organization and
function of the tissue [4]. Tissue engineering scaffolds must therefore have the appropriate
mechanical properties to promote mechanotransduction within the tissue and degradation
characteristics that allow the gradual transfer of mechanical load from biomaterial to newly
formed ECM. Elastomeric segmented polyurethanes are important biomaterials for use in the
heart where soft, flexible and elastic mechanical properties are requisite. The inherent flexibility
in PU chemistry also allows the incorporation of amino acid and peptide sequences into the
backbone structure of the polymer making them susceptible to cell-secreted proteases [5, 6].
This may lead to degradation characteristics that are responsive to the tissue environment and are
therefore more appropriate for cardiac and other soft tissue engineering applications than
synthetic polymers that degrade predominantly by passive hydrolysis.
Building on previous work in our lab of enzyme-sensitive segmented polyurethanes, the
goal of this work was to develop a family of polyurethanes that incorporate either a Gly-Leu or
Gly-Ile dipeptide into the backbone structure of the polymer. The peptide bonds between
glycine and leucine and glycine and isoleucine residues are the cleavage sites of several MMPs
that have been found to play a critical role in heart disease [7-10]. Incorporation of the Gly-Leu
86
or Gly-Ile dipeptide was suspected to confer enzyme susceptibility to the PUs, thereby improving
the degradation characteristics by making them more responsive to ECM remodeling proteases.
This chapter describes the successful synthesis and characterization of a Gly-Leu-
containing polyurethane (Gly-Leu PU) that may hold potential as a new biomaterial in cardiac
tissue engineering. Several synthesis and purification methods were initially investigated for
incorporating the Gly-Leu and Gly-Ile dipeptides into an amine terminated chain extender.
Ultimately, the Gly-Leu dipeptide was reacted with a cycloaliphatic diol linker to form a
symmetric diamine diester chain extender. A Gly-Leu PU was subsequently synthesized and
structure-property comparisons were made using a phenylalanine-containing polyurethane (Phe
PU) previously developed in our laboratory. The two PUs only differed structurally by the
presence of the amino acid phenylalanine in the chain extender instead of the Gly-Leu dipeptide.
Consequently, the PUs exhibited very similar physical, chemical, and thermal properties with
only subtle differences in mechanical behavior due to structural differences between the two
polymers.
3.2 Materials and Methods All materials were purchased from Sigma-Aldrich Canada (Oakville, ON, Canada) unless
otherwise stated.
3.2.1 Dipeptide-based Chain Extender Synthesis Initial work in developing a dipeptide-based chain extender investigated different
synthesis and purification methods. The different synthesis parameters that were tested include
the solvent, catalyst, dipeptide, and linker molecule. Two equipment setups were used to
investigate the synthesis methods and are illustrated in Figure 3.1. The first system (system 1)
utilized a round bottom reaction vessel, a Dean-Stark water trap, water condenser, stir bar, drying
tube, heating mantle, and magnetic stir plate. This setup was used to test either toluene alone or
toluene and dimethyl formamide (DMF) as solvents, the Gly-Leu and Gly-Ile dipeptides, the
catalysts p-toluene sulfonic acid (PTSA) and methane sulfonic acid (MSA), and the linkers 1,4-
cyclohexane dimethanol (CDM) and 1,6-hexane diol. Reaction contents were added to the
reaction vessel at a 2:1 molar ratio of dipeptide to diol linker in an excess of acid catalyst. The
reactions were carried out under reflux conditions for periods between 2 and 48 h. Any water
that formed during the reaction was collected in the trap to drive the reaction in favor of the
87
diester product. After the reaction, crude products were collected directly or by allowing the
solvents to evaporate in a fume hood. Collected products were dried in a fume hood at room
temperature for 24 h, then in a vacuum oven at 60ºC for 48 h, and were stored in a desiccator
until analysis.
In the second system (system 2), an addition funnel, round bottom reaction vessel, tygon
tubing, sparge, spin bar, 3-neck round bottom flask, and drying tube were employed. Using this
setup, a hydrogen chloride gas was generated by slowly dripping concentrated sulfuric acid into
a reaction vessel containing sodium chloride. The hydrogen chloride gas acted as an acid
catalyst and was sparged into the 3-necked reaction vessel containing the reactants with the 2:1
molar ratio of dipeptide to diol linker. This system was used to test the hydrogen chloride gas
catalyst and tert-amyl alcohol solvent with the Gly-Ile dipeptide and CDM linker. After
allowing the reaction to occur for ~4 h, the product was recovered by vacuum drying at 60 ºC
using a rotary evaporator for 6-12 h. The product was further dried in a vacuum oven at 60ºC for
48 h and was stored in a desiccator until analysis.
Figure 3.1: Chain extender reaction system setups. a) System 1 was used to test the solvents toluene and
toluene/DMF, the catalysts PTSA and MSA, the dipeptides Gly-Leu and Gly-Ile, and the CDM and hexane diol linkers. b) System 2 was used for hydrogen chloride gas catalyzed synthesis using t-amyl alcohol as a solvent.
§Ultimately, a Gly-Leu-based chain extender was synthesized with system 1 using the
method by Huang et al. [11] for amino acid-based diester products. A Fischer esterification
reaction was used to react the Gly-Leu dipeptide and CDM at a 2:1 molar ratio (Figure 3.2). The
reaction was carried out in excess p-toluene sulfonic acid catalyst in toluene under refluxing
§ This section was published in [1].
88
conditions and was driven to completion using a water trap. The reaction was stopped when no
more water appeared to evolve (approximately 3 h). The crude chain extender, which appeared
as a clear viscous product, was collected and residual solvent was allowed to evaporate in a fume
hood for 24 h and subsequently in a vacuum oven at 60˚C for 48 h. The resulting product was
ground into a fine powder and stored in a desiccator until purification.
Figure 3.2: Synthesis scheme for Gly-Leu-based diester, diamine chain extender.
3.2.2 Gly-Leu-based Chain Extender Purification§ The raw chain extender product was purified by high performance liquid
chromatography (HPLC) using Gemini C-18 HPLC analytical and preparative columns
(Phenomenex, Torrance, CA, USA). Methods for separating the desired diester product were
initially tested with the analytical column using a series of mobile phase steps. Two methods
were developed on the analytical scale that were scaled up and optimized using the preparative
column. The first method developed used a low pH aqueous mobile phase made up of 0.1%
trifluoroacetic acid (TFA) in ultrapure deionized water (pH ~2.1) and acetonitrile (ACN) as the
organic phase. The series of steps used to isolate and collect peaks on the preparative column
corresponding to the Gly-Leu-based diester chain extender product included: a hold at 15%
organic phase (85% aqueous phase) for 4 min; gradient step from 15-40% ACN in 10 min;
§ Part of this section was published in [1].
89
gradient step from 40-60% organic phase in 1 min; hold at 60% organic phase for 3 min;
gradient step from 60-15% ACN in 1 min; and re-equilibration hold at 15% organic phase for 7
min prior to injecting the next sample. The second method used a higher pH aqueous mobile
phase composed of 10 mM ammonium bicarbonate in ultrapure deionized water (pH 8.0) and the
same ACN organic phase. The higher pH profile on the preparative column had a hold at 32%
organic phase for 6 min; gradient step from 32-50% ACN in 6 min; 50% ACN hold for 2 min;
gradient step from 50-100% organic phase in 1 min; hold at 100% for 4 min; a gradient step from
100-32% ACN in 1 min; and a re-equilibration hold at 32% organic phase for 9 min. In both
methods, the raw chain extender was dissolved in aqueous phase at 75 mg/ml and 1 ml of
solution was injected into the column running at a flow rate of 18 ml/min using a Waters HPLC
system (Waters Chromatography, Mississauga, ON). Solvents were kept in an ice bath to reduce
hydrolysis of the desired product. Peaks were monitored using a UV detector (Waters
Chromatography) at a wavelength of 214 nm. Two prominent peaks were collected between
approximately 14.5-16.0 min with the low pH method and between approximately 12.5-15.5 min
with the high pH method. The purified chain extender was recovered following lyophilization
and the dry product was stored in a desiccator until use. After initial analysis of purified
products, the low pH method was chosen for subsequent use.
3.2.3 Chain Extender Characterization§ Raw products from the different synthesis and purification methods were characterized
by electrospray ionization (ESI) mass spectrometry. The ESI mass spectrometry was performed
at the Proteomic and Mass Spectrometry Centre at the University of Toronto (Toronto, ON,
Canada) using an API QSTAR XL mass spectrometer (Applied Biosystems Inc., Foster City,
CA) running in positive mode within the charge/mass ratio detection range of 150-550 amu.
Purified chain extenders were further characterized by C13 nuclear magnetic resonance
(NMR) spectroscopy and Fourier transform infrared spectroscopy (FT-IR). Samples for C13
NMR analysis were dissolved in deuterated methanol at approximately 4% w/v and analyzed
using a Varian Mercury 300 spectrometer (Varian Inc, Palo Alto, CA) for a total of 1000 scans.
Reactant C13 NMR spectra and theoretical Gly-Leu-based chain extender predictions using ACD
i-Lab software (Advanced Chemistry Development Inc, Toronto, ON) were used to help identify
§ Section was published in [1].
90
and assign spectral peaks. FT-IR samples were prepared as KBr crystals. Approximately 1 mg
of sample was added to ~300 mg of KBr powder, the mixture was homogenized, and crystals
were formed by high pressure using a hydraulic press (Carver Inc, Wabash, IN) at 15,000
pounds. Samples were analyzed using a Spectrum 1000 FT-IR spectrometer (PerkinElmer,
Waltham, MA) from 4000 - 400 cm-1 at a resolution of 4 cm-1 for 64 scans using PerkinElmer
Spectrum software.
3.2.4 Polyurethane Synthesis and Film Casting§† Two polyurethanes were synthesized for use in this study. The first PU contains a
phenylalanine (Phe)-based diester chain extender, polycaprolactone diol of molecular weight
1250 (PCL1250), and 2,6 diisocyanto methyl caproate (LDI; Kyowa Hakko Kogyo Co. Ltd,
Japan) and is referred to as the Phe PU. The Phe-based chain extender was synthesized by
Toronto Research Chemicals Ltd (North York, ON) according to procedures developed by
Skarja and Woodhouse [12]. The Phe PU was synthesized as previously described [13] using a
similar procedure as described next. The second PU contained the Gly-Leu-based chain extender
described here and was synthesized using PCL1250 and LDI. This polyurethane was made using
a 2:1:1 molar ratio of LDI, PCL1250, and the Gly-Leu-based chain extender respectively (Figure
3.3). LDI, distilled prior to synthesis, was reacted with PCL1250 in anhydrous DMF at 85ºC for
2.5 h under nitrogen in the presence of 0.1% stannous octoate catalyst to form the prepolymer.
Triethylamine was added to the chain extender in DMF at twice the molar concentration of chain
extender to neutralize it and this solution was added to the reaction vessel after allowing the
temperature to drop below 60˚C. Polymer chain extension took place at room temperature for
approximately 18 h and the reaction contents were subsequently precipitated in aqueous KCl
(~0.05 M). The polymer was incubated in deionized water at 37ºC for 48 h to remove residual
salt and soluble reactants, dried under vacuum at room temperature for at least 48 h, and stored
in a desiccator until use.
The final PU products were dissolved in chloroform at 3% w/v and gravity filtered using
Whatman 40 filter paper. Approximately 20 ml of the polymer solutions were cast in 5 cm x 5
cm Teflon dishes, the dishes were covered, and the solvent was allowed to evaporate at room
§ Section was published in [1] and † the Phe PU synthesized and characterized here was used in [2].
91
temperature for 72 h. The resulting thin films were dried under vacuum at room temperature for
24 h and stored in a desiccator until use.
Figure 3.3: Synthesis scheme for Gly-Leu PU.
3.2.5 Polyurethane Characterization§† Polyurethane molecular weight averages were determined by gel permeation
chromatography (GPC) using a Waters GPC system (Waters Chromatography). One sample
from three separate synthesis trials was dissolved in a mobile phase consisting of 0.05 M lithium
bromide in DMF at 0.2% w/v and 200 μl of the resulting solution was injected into a column
running at 1 ml/min and 80°C. A calibration curve generated from polystyrene standards
(Varian, Sunnyvale, CA) ranging from 2,980 to 706,000 g/mol was used along with Waters
Empower 2 software (Waters Chromatography) to produce retention time data, number and
weight average molecular weights, and polydispersity.
Thermal properties of the PUs were characterized by differential scanning calorimetry
(DSC) using a Thermal Analyst 2100 thermal analyzer (DuPont Canada, Mississauga, ON)
§ Section was published in [1] and † the Phe PU synthesized and characterized here was used in [2].
92
performed at the Brockhouse Institute for Materials Research (McMaster University, Hamilton,
ON). Scans were conducted at a rate of 15˚C/min from -100 to 200˚C. One sample from three
separate synthesis trials was subjected to one scan to erase the thermal history, and were
subsequently quenched to room temperature and exposed to a second scan. Results from both
scans were used to identify the glass transition temperature (Tg) and melting temperature (Tm)
of the polymer. Crystallinity of the PUs was calculated assuming an enthalpy of fusion of 32.4
cal/g for 100% crystalline PCL [14] and the theoretical mass fraction of PCL1250 expected from
a 2:1:1 molar ratio of either LDI:PCL1250:Phe chain extender or LDI:PCL1250:Gly-Leu chain
extender.
FT-IR analysis was carried out by dissolving the PU samples in chloroform at 5% w/v
and placing the solution directly onto NaCl plates. The solvent was evaporated at room
temperature under vacuum for 30 min and stored in a desiccator until analysis. Spectra were
generated from 256 scans using parameters described above for the chain extender.
Uniaxial tensile properties of the PU films were determined according to ASTM D1708
using an Instron 4301 testing machine (Instron, Norwood, MA). Samples were cut from films
that were 70-90 µm thick using an ASTM D1708 die. The PUs were preconditioned at
approximately 24˚C and 55% relative humidity for 48 h. Testing was carried out using a 50 N
load cell running at a crosshead speed of 100 mm/min. Samples were stretched to break and
stress-strain curves were generated from force-displacement data. Stress and strain at yield,
stress and strain at break, and initial modulus were inferred from the stress-strain graphs.
Statistical comparisons were made using a two-tailed independent t-test using the SPSS Statistics
17.0 statistical software package (SPSS Inc, Chicago, IL).
3.3 Results and Discussion
3.3.1 Chain Extender Synthesis and Purification The goal of this work was to develop PUs that could be degraded by ECM remodeling
proteases expressed in an injured heart. Previous work demonstrated that enzyme-responsive
synthetic polymers can be developed by incorporating amino acids or short peptides into the hard
segment of biodegradable PUs [5, 6]. Through this approach, the choice of soft segment could
be employed to achieve diverse polymer properties while still promoting PU degradation.
Methods to incorporate amino acids and short peptides into the hard segment of PUs have
93
included: 1) the direct use of amino acids and dipeptides, producing asymmetrical chain
extenders with urea and amide functional groups being formed during PU synthesis [15]; 2) by
adding a lysine residue on the end of the desired peptide sequences to form an asymmetric
diamine chain extender [5]; or 3) using a diol linker to create symmetric diamine reactants [16].
In regards to the last approach, several aliphatic and cycloaliphatic diols, including ethylene
glycol, 1,3-propane diol, 1,4-butane diol, 1,6-hexane diol, dodecane diol, and 1,4-cyclohexane
dimethanol, have been used as linkers to attach the amino acids phenylalanine, alanine, glycine,
and lysine and fabricate amine-terminated diester products [16-21]. Continuing with work that
was done in our laboratory using diol linkers, this approach was utilized to produce a dipeptide-
based amine terminated diester chain extender.
3.3.1.1 Reaction Systems for Chain Extender Synthesis Initial work in forming dipeptide-based diester products involved the Gly-Ile dipeptide
with system 1 (Figure 2.1) using a similar approach as Skarja and Woodhouse [16] for
synthesizing the Phe-based diester chain extender. In this approach, the Gly-Ile dipeptide was
reacted with CDM at a 2:1 molar ratio in the presence of excess p-toluene sulfonic acid catalyst
under refluxing conditions. The acid catalyst limited the peptide bond-forming side reaction
between two or more dipeptide molecules by protonating the amine functional groups. The raw
product obtained from this (denoted Gly-Ile-CDM-PTSA) was analyzed by ESI mass
spectroscopy and is shown in Figure 3.4. The unreacted dipeptide was observed at a charge to
mass ratio of 189 amu, the partially reacted monoester product at 315 amu, and the desired
diester product at 243 and 485 amu corresponding the diprotonated and mono-protonated forms
respectively. The major peaks seen in the mass spectrum of the Gly-Ile-CDM-PTSA
corresponded to these molecular weights suggesting side reactions and the formation of
undesirable side products were not significantly detected. The monoester and diester products
were formed to some extent indicating the esterification reaction took place. The desired diester
product, however, had a lower relative peak intensity (~30%) in comparison to the monoester
product (~42%) and unreacted dipeptide peaks (~28%) suggesting the reaction did not go to
completion. As a result, reaction time was varied from ~2 - 48 hrs to achieve better diester
product yields, but this had little affect on the outcome and the monoester product and unreacted
dipeptide were consistently higher than the diester product. It was hypothesized that one
explanation for this observation was due to a mixing issue within the reaction vessel. Once the
94
reaction started, a viscous product formed on the bottom that limited the stir bar from effectively
mixing the vessel contents. Ineffective reactant mixing may prevent the reaction from
progressing to completion and may lead to the higher monoester product observed.
Figure 3.4: Mass spectrum of raw Gly-Ile-CDM-PTSA product. Relevant peaks are highlighted. A higher proportion of the monoester product was observed over the diester product suggesting the reaction did not go to
completion. In an attempt to obtain a higher Gly-Ile-based diester product, two other solvent systems
were investigated. These solvents were explored to test if reaction heterogeneity and reactant
mixing influenced successful esterification. The first attempt used system 1 and the same
reactants with a 95% toluene/5% DMF co-solvent instead of toluene alone. Adding DMF to the
reaction vessel led to a more homogenous solution and better mixing in this system. The second
attempt used tert-amyl alcohol as a solvent in system 2. This solvent was tested in combination
with the HCl gas catalyst and led to a homogenous reaction mixture. Raw products from the two
alternative synthesis trials were analyzed by ESI mass spectroscopy and are shown in Figure 3.5.
The diester product in both of these systems was very low in comparison to the monoester
products and side reactions. As a result, the toluene-based solvent in system 1 appeared to be the
95
Figure 3.5: Mass spectra of raw Gly-Ile-based chain extender products synthesized in different solvent systems. a)
Raw Gly-Ile-CDM-PTSA product in toluene/DMF co-solvent and b) raw Gly-Ile-CDM-HCl product in t-amyl alcohol. Very little diester product was observed in both systems.
96
best synthesis method tested for producing the desired Gly-Ile-based diester product. Similarly,
the PTSA catalyzed system produced better results than the HCl catalyzed system and therefore
the HCl gas catalyzed system was not further explored.
3.3.1.2 Synthesis of Chain Extenders using Gly-Ile or Gly-Leu Dipeptides Several MMPs cleave ECM proteins at Gly-Leu and Gly-Ile sites in the polypeptide
chains [10]. Incorporating either the Gly-Leu or Gly-Ile dipeptides into the backbone structure
of PUs may introduce labile bonds susceptible to cleavage by the MMPs. Initial experiments in
producing a dipeptide-based diester chain extender suggested some success when using CDM,
PTSA and toluene as synthesis parameters. The inability to drive the desired diester product to
completion with the Gly-Ile dipeptide, however, prompted synthesis strategies with the Gly-Leu
dipeptide. The Gly-Leu dipeptide was reacted with CDM at a 2:1 molar ratio with PTSA in
toluene under reflux for ~3 hr. Crude chain extender products were characterized by mass
spectrometry to identify the success of the synthesis. A representative spectrum for the raw Gly-
Leu-CDM-PTSA product is shown in Figure 3.6. Using this dipeptide, the reaction appeared to
proceed in favor of the diester product. A higher relative peak intensity was observed for the
diester product at 243 and 485 amu (~80%) compared to the peaks of the monoester product at
315 amu (~18%) or unreacted dipeptide at 189 amu (~2%). This contrasted results of the Gly-Ile
dipeptide, which favored the partially reacted monoester product. The difference in reaction
success may be a result of the chemical structure of the two dipeptides. Although very similar in
structure, the isoleucine residue contains a methyl group on the side chain in close proximity to
the carboxylic acid group that is needed for the esterification reaction to occur. This creates
greater steric hindrance and may limit access of the fairly rigid CDM molecule to the reactive
carboxylic acid end. The leucine residue has the methyl side-chain group farther from the
carboxylic acid and may allow greater access to this reactive functional group. Interestingly,
once the reaction started with both dipeptides, a viscous product formed on the bottom of the
reaction vessel limiting mixing. A moderate scale reaction setup (100-150 ml reaction vessel)
was used because this scale system allowed the stir bar to continue spinning even after the
formation of this product, thereby promoting mixing of reaction contents. Approximately 10-20
g of raw product could be produced on this scale. Scaling up production from this system,
however, inhibited mixing of the viscous product and may be a limiting factor in producing
large-scale batches of the Gly-Leu-based diester chain extender. Regardless, these results
97
suggested a dipeptide-based chain extender could be synthesized with a reaction system that
favored the diester product. Purification strategies were subsequently investigated before PUs
were synthesized with this new chain extender.
Figure 3.6: Mass spectrum of crude product from Gly-Leu-CDM-PTSA. The diester product dominates the
spectrum at 243 and 485 amu.
Other systems that were investigated for the Gly-Leu-based chain extender include a
methane sulfonic acid catalyst and 1,6-hexane diol linker. MSA is a liquid catalyst and its use
was suspected to improve mixing within the reaction vessel if reaction scale-up was necessary.
The aliphatic diol linker was tested as an alternative to CDM as an option for influencing PU
properties. Results from these reactions are shown in Figure 3.7 and both show similar success
in forming a Gly-Leu-based diester product. Although the hexane diol linker showed positive
esterification results in this system, CDM was chosen for subsequent work due to the increased
rigidity associated with this molecule. This was anticipated to reduce hard segment mobility,
promote phase separation, and improve mechanical properties in comparison to the aliphatic diol,
especially in light of using the aliphatic, asymmetric side chain-containing diisocyanate LDI.
More importantly, PUs created with dipeptide-based chain extenders employing CDM as a linker
allowed a direct comparison to the Phe-containing PUs to test the effects chain extender
chemistry had on PU properties. The hexane diol linker and MSA were not investigated further,
98
Figure 3.7: Mass spectra of Gly-Leu-based chain extender using different catalysts and diol linkers. a) Raw Gly-Leu-CDM-MSA and b) raw Gly-Leu-HD-PTSA. High relative peak intensities corresponding to diester product
were observed. Note: diester peaks with hexane diol linker are 230 and 459 amu and monoester peak is at 289 amu.
99
but offer alternative options for future work in scaling up and synthesizing alternative Gly-Leu-
based chain extenders.
3.3.1.3 Purification Strategies for the Gly-Leu-based Chain Extender§ After identifying that a large portion of the Gly-Leu-CDM-PTSA raw product was the
desired diester product, purification of the chain extender was investigated. Solubility tests on
the reactants and raw products were conducted using several common solvents in an attempt to
establish a similar purification protocol to the phenylalanine-based chain extender. No
significant differences were observed between the reactants and raw product so an alternative
approach was undertaken. High performance liquid chromatography is commonly used to
separate peptides and was investigated for purifying the dipeptide-based diester chain extender.
An analytical Gemini C-18 column was used to establish a method for purifying the chain
extender and was subsequently scaled up using a preparative column for large scale purification.
The Gemini C-18 column is a silica-based column containing covalently bonded hydrophobic C-
18 ligands (n-octadecyl alkyl chains). An advantage of using the Gemini column is the working
pH can range from 1-12, thus allowing the flexibility of developing and optimizing a separation
scheme for purifying the diester product in either a protonated or neutral state.
Two methods for purifying the Gly-Leu-based diester product were established on an
analytical scale. The first method involved an aqueous mobile phase consisting of ultrapure
deionized water with 0.1% TFA and an organic phase of ACN. The aqueous mobile phase
maintained a low pH (~2.1) to keep the analytes protonated and allowed separation to be based
on charge and relative hydrophobicity. A solvent flow rate of 1 ml/min was used along with a
UV detector at a wavelength of 214 nm. Using a gradient elution profile of 0-50% ACN in 30
min and 50-100% ACN in 5 min, peaks associated with the desired diester product were isolated
and collected. Figure 3.8 shows a representative chromatogram with collected peaks highlighted
and the resulting mass spectroscopy analysis. The mass spectrum indicated a pure Gly-Leu-
based diester chain extender. No peaks corresponding to the monoester product, unreacted
dipeptide, or other contaminants were observed suggesting successful product purification.
§ Parts of this section were published in [1].
100
Figure 3.8: HPLC separation of Gly-Leu-based diester product using analytical column and low pH aqueous mobile phase. a) Chromatogram of run going from 0-50% ACN in 30 min and 50-100% ACN in 5 min with collected
peaks highlighted. b) Mass spectrum of peaks identified in a). A pure diester product was identified. The elution peaks in the TFA-based profile described above corresponding to the diester
product were relatively close to other peaks and there was a concern about peak broadening and
successful separation when scaling up to a prep column. As a result, a second method was also
developed on the analytical scale. A higher pH aqueous mobile phase was used in this system
composed of 10 mM ammonium bicarbonate (pH~8.0). One advantage to using this method was
the retrieved product would be neutralized, thus eliminating a potential step in the purification
scheme. Using liquid chromatography-mass spectroscopy in tandem, peaks corresponding to the
desired product were identified. Subsequent method development attempted to separate the
desired peaks by using a series of mobile phase gradient and hold steps. A typical chromatogram
from this method is shown in figure 3.9 along with the corresponding mass spectrum of collected
peaks. The high relative proportion of diester product and relatively few other peaks suggested
this method to be promising for purifying the chain extender with the prep column.
101
Figure 3.9: HPLC separation of Gly-Leu-based diester product using analytical column and high pH aqueous
mobile phase. a) Chromatogram from an elution profile of gradient-hold steps with collected peaks highlighted. b) Mass spectrum of peaks identified in a). Both the diester and monoester products are observed using this method.
On the analytical scale, mass spectrometry results suggested the two purification schemes
developed may be used to separate out the Gly-Leu-based diester chain extender. These two
methods were then scaled up to a prep column system to obtain enough of the purified product
for further characterization and use in PU synthesis. Adjustments were made to the elution
profiles using a combination of gradient and hold steps to isolate desired peaks. Representative
chromatograms from the two methods and corresponding mass spectra are shown in Figure 3.10.
The spectra from both methods appeared similar with relatively high levels of the diester product
along with the presence of some monoester product and few other peaks. Using peak intensities
from the spectra, the relative percent of monoester and diester products in the low pH method
was approximately 10% monoester and 90% diester product indicating a relatively pure
population of the desired diester chain extender. The high pH method had approximately 30%
monoester and 65% diester product. The higher monoester product with the high pH method
may have resulted from increased hydrolysis of the diester product during separation despite
keeping the mobile phases on ice during the HPLC runs. Although the hydroxyl-terminated end
of the monoester product may still react with the isocyanates during PU synthesis, this product
was not desirable due to its asymmetric structure, different end reactivities, and reduced potential
for hydrogen bonding. As a result, the low pH method that produced a higher relative percent of
diester product was preferred for chain extender purification.
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Figure 3.10: Preparative column HPLC purification of chain extender using low and high pH aqueous mobile
phases. a) Chromatogram of low pH method with collected peaks identified. b) Mass spectrum of low pH method. c) Chromatogram of high pH method with collected peaks. d) Mass spectrum of high pH method. The low pH
method had a higher relative intensity of desired diester product suggesting this was preferred over high pH method.
Using the prep column system, enough product was recovered from the two HPLC
purification methods to further characterize the chain extender and provide further evidence for
the preferred purification scheme. C13 NMR was conducted to help characterize the products
and the corresponding C13 NMR spectra are shown in Figure 3.11. Using the low pH method,
the spectrum revealed the presence of 10 prominent non-identical carbon peaks in the purified
product. This corresponds well with the number of non-identical carbons found in the theoretical
103
structure of the Gly-Leu-based diester chain extender and provided further proof of successful
purification using this method. Analysis of the reactants and theoretical predictions using ACD
i-Lab software (Appendix A) were used to assign the chemical shifts in the spectrum to the
corresponding carbons in the chemical structure of the chain extender. The unshielded carbonyl
carbon from the ester functional group was assigned downfield at the highest chemical shift of
172 ppm in relation to the unaffected carbonyl carbon of the amide functional group found at 167
ppm. A slight change in chemical shift was noted for the methylene carbon from CDM, which
shifted downfield following esterification, but the unreacted form of this carbon present in the
monoester product was not observed. This may be due to the relatively low percent of
monoester product compared to diester product as well as the low signal to noise ratio with
obtained spectra. Importantly, the aromatic carbons normally found in the spectrum of the
unpurified product (Appendix A) between 120-140 ppm are absent, indicating separation from
the acid catalyst p-toluene sulfonic acid. The large unassigned peaks centered at approximately
49 ppm correspond to the deuterated methanol solvent.
The high pH method was similarly able to separate the product from the PTSA catalyst,
but this spectrum had several additional non-identical carbons in the collected product. Four
carbonyl peaks were identified in the region from 160-220 ppm. Although only a speculation,
the additional carbonyl carbons may be carboxyl functional groups from the Gly and Leu amino
acids that would be formed through hydrolysis of the amide and ester linkages in the diester and
monoester products. The Gly and Leu amino acids have molar masses of 75 and 131 g/mol
respectively and would not show up in the mass spectroscopy range investigated. These results
suggested that the low pH method was able to achieve a relatively pure population of the desired
Gly-Leu-based diester product whereas the high pH method was not. The low pH method was
therefore chosen for purifying the chain extender for subsequent studies. Approximately 30 mg
of product was recovered for each 30 min run when injecting 75 mg of raw product giving an
approximate 70% yield. Despite the limitation of being time consuming, this system produced
enough of the purified chain extender for synthesizing the Gly-Leu PU.
104
Figure 3.11: C13 NMR spectra of products collected from preparative column HPLC using the two developed
methods of separation. a) C13 NMR spectrum of low pH HPLC method. This method led to the same number of chemical shifts as non-identical carbons in the theoretical structure of chain extender and was the preferred method
of separation. The chemical structure of chain extender along with corresponding carbon peak assignments is provided. * Corresponds to the large unassigned solvent peaks at ~49 ppm. b) C13 NMR spectrum of high pH
HPLC method. This method led to more than 10 non-identical carbons suggesting this method was not preferred for the purification of the chain extender.
105
Once the preferred method of separation was chosen, FT-IR analysis was conducted on
the purified products to identify functional groups within the chain extender. The FT-IR
spectrum of the purified chain extender is shown in Figure 3.12. The formation of an ester peak
at 1736 cm-1 confirms successful esterification between the carboxyl group of the Gly-Leu
dipeptide and the hydroxyl groups of the 1,4-cyclohexane dimethanol. Amide I and amide II
functional groups are indicated by the peaks at 1675 cm-1 and 1560 cm-1 respectively indicating
the peptide bond between the glycine and leucine residues was not disrupted during chain
extender synthesis or purification. A protonated amine group is indicated by the peak at 2960
cm-1, suggesting the need for chain extender neutralization prior to PU synthesis. This was
performed by adding triethylamine at twice the molar concentration of chain extender in DMF
before adding the mixture to PU synthesis reaction.
Figure 3.12: FT-IR spectrum of purified Gly-Leu-based chain extender. Key functional groups are identified.
106
3.3.2 Polyurethane Characterization§† Polyurethanes were synthesized by the standard two-step process using PCL of molecular
weight 1250, LDI, and either the Gly-Leu-based chain extender developed here or the previously
established phenylalanine-based chain extender (Gly-Leu PU and Phe PU respectively). The two
polyurethanes only differ by the substitution of a Gly-Leu dipeptide in place for the amino acid
phenylalanine around the CDM linker in the chain extender (Figure 3.13). The similarity in PU
chemistries provided the opportunity to directly investigate the structure-property relationship of
incorporating the Gly-Leu dipeptide into the backbone polymer chain. The molecular weight
averages, thermal properties, chemical composition, and mechanical properties of the two PUs
are discussed here.
Figure 3.13: The chemical structure of the Phe and Gly-Leu-based chain extenders.
3.3.2.1 Molecular Weight Averages Molecular weight averages were determined by GPC. As shown in Table 3.1, the Gly-
Leu PU had high molecular weight averages (Mw > 125,000 g/mol) and a relatively low
polydispersity (PD < 1.55). These molecular weight averages were comparable to those obtained
for the Phe PU, consistent with previously published data on the Phe PU [13]. The similarities in
chain extender molecular weights (484 g/mol and 439 g/mol for the Gly-Leu and Phe-based
chain extenders respectively) suggest the Gly-Leu-based chain extender may react with the
prepolymer to a similar extent as the Phe-based chain extender. Importantly, the presence of any § Section was published in [1] and † characterization of the Phe PU contributed to [2].
107
Gly-Leu-based monoester product that may not have been separated during purification or
formed from hydrolysis of the diester product during purification did not greatly impact the
ability to obtain a high molecular weight PU. The molecular weight averages obtained for both
polyurethanes are well within the range to impart elastomeric mechanical behavior to polymer
films, Mn > 25,000 g/mol [23]. Table 3.1: Molecular weight averages for PUs containing Phe and Gly-Leu-based chain extenders. Mean ±
standard deviation. n=3 Polymer Mn (g/mol) Mw (g/mol) PD (Mw/Mn)
Gly-Leu PU Films 82700 ± 6000 126600 ± 9400 1.531 ± 0.159 Phe PU Films 76100 ± 4300 128600 ± 1800 1.696 ± 0.099
3.3.2.2 Thermal Transitions and Phase Segregation Thermal properties of the PUs were characterized by DSC. These results suggest the
Gly-Leu PU is a thermoplastic polymer with similar thermal properties to the Phe PU (Table
3.2). A low soft segment glass transition temperature (Tg) of approximately -51˚C was
observed, indicating a soft, flexible polymer at room and body temperature. Importantly, this
low Tg suggests the Gly-Leu-based chain extender promoted good phase segregation for the
formation of a biphasic morphology with hard and soft segment domains. As a result, the
polymer was semi-crystalline with a soft segment melting temperature of approximately 42˚C
and a soft segment crystallinity of ~28%. No hard segment glass transition temperature or
melting temperature was observed out to the tested temperature of 200ºC. Although it can not be
ruled out that hard segment thermal transitions may occur above 200 ºC, the bulky methyl ester
side chain of LDI as well as the pendant side chains on the leucine residue within the chain
extender may limit hard segment packing thus preventing hard segment crystallinity. The Phe
PU also had no hard segment transitions, which may also be contributed to the bulky side chains
within the hard segment [24]. Table 3.2: Thermal properties of the Phe and Gly-Leu PUs as determined by DSC. Mean ± standard deviation. n=3
Polymer Soft Segment Polymer Tg (˚C) Tm (˚C) Crystallinity (%) Crystallinity (%)
Gly-Leu PU Films -51.5 ± 1.1 42.0 ± 1.5 28.4 ± 2.6 16.5 ± 1.5 Phe PU Films -50.6 ± 1.2 43.0 ± 0.7 27.4 ± 1.6 16.0 ± 0.9
3.3.2.3 Chemical Composition FT-IR analysis confirmed the presence of the appropriate functional groups within the
Gly-Leu PU (Figure 3.14). The broad absorption band centered at 3370 cm-1 indicated N-H
stretching from the urethane, urea, and amide groups. A similar band was observed in the Phe
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PU centered at 3373 cm-1 corresponding to the urethane and urea functional groups. The slight
difference in N-H stretching peaks between the Phe and Gly-Leu PU may result from a higher
degree of hydrogen bonding within the Gly-Leu PU, which causes a shift to a lower wavenumber
[25]. A broad ester band centered around 1732 cm-1 was observed for the Gly-Leu PU along
with a shoulder band at 1736 cm-1 corresponding to a hydrogen bonded and free functional
groups respectively. The ester peaks are predominantly contributed to the PCL soft segment,
however, some contribution may also be made from the ester linkages within the chain extender.
Similar bands were seen with the Phe PU. Although the ester functional group dominated the
carbonyl stretching region from 1650-1800 cm-1, a shoulder band centered at 1654 cm-1 and 1648
cm-1 was observed for the Gly-Leu and Phe PUs respectively. These bands correspond to the
amide I carbonyl stretching within the amide, urea, and urethane groups in the Gly-Leu PU and
the urea and urethane groups in the Phe PU. Similarly, an amide II band caused by N-H bending
and C-N stretching was observed at approximately 1549 cm-1 for both PUs. No unreacted
isocyanate groups were found for either PU due to the absence of a peak at approximately 2267
cm-1.
Figure 3.14: FT-IR analysis of Phe and Gly-Leu PUs with functional groups highlighted.
109
3.3.2.4 Mechanical Properties The uniaxial tensile properties of Phe and Gly-Leu PU films were determined according
to ASTM D1708. Stress-strain curves were generated from force-displacement data (Figure
3.15). Initial modulus was determined from the slope of the linear portion of the curve (elastic
region of curve) from ~1-10% strain. The yield stress and strain were taken from the point after
the initial modulus where an increase in strain led to a decrease in stress. Stress and strain at
break were inferred from the curves at the point of material failure. The mechanical properties
for the two PUs are summarized in Table 3.3. The stress-strain curves for the two PUs were
typical of semi-crystalline elastomers. Both PU film types were able to withstand high ultimate
stresses from 8-13 MPa and were highly extensible, being able to withstand deformations greater
than 850%. Importantly, a comparison of means using a two-tailed independent t-test indicated a
significant difference between the two PUs for all the tensile properties determined (p<0.05),
with the exception of strain at break. The Gly-Leu PU had a significantly higher initial modulus,
stress at yield, and stress at break, while the Phe PU had a significantly higher strain at yield.
Several factors can influence the tensile properties of the PUs, including molecular weight, soft
segment length, crystallinity, and phase segregation [13], however, GPC and DSC results
indicated the Phe and Gly-Leu PU were similar in all these regards. Differences in chain
Figure 3.15: Representative stress-strain curves for Phe and Gly-Leu PU films.
110
extender structure may be one factor causing the observed differences in mechanical properties.
Although DSC scans revealed hard segment cohesiveness was not sufficient to see any thermal
transitions, the bulkier aromatic side chain in the Phe-based chain extender may reduce mobility
and limit hard segment packing to a larger extent than the side chain in the Gly-Leu-based chain
extender. As a result, the hard segment domains within the Gly-Leu PU may be more organized
to promote hydrogen bonding. The Gly-Leu PU had greater hydrogen bonding within the N-H
stretching region as suggested by FT-IR and further supports this notion. As well, amide groups
within the Gly-Leu-based chain extender add additional hydrogen bond forming groups and may
also play a role in the enhanced tensile properties of the Gly-Leu PU.
Table 3.3: Summary of mechanical properties of PU films. Mean ± standard deviation. n=9. a Statistical difference between means at p<0.05 using two-tailed independent t-test. *not in accordance with ASTM standard
Polymer Initial
Modulus (MPa)
Stress at Yield (MPa)
Strain at Yield (%)
Stress at Break (MPa)
Strain at Break (%)
Gly-Leu PU 50.9 ± 7.5*a 4.6 ± 0.3a 19.4 ± 2.2a 12.9 ± 1.2a 908.2 ± 64.5 Phe PU 44.1 ± 4.8*a 3.5 ± 0.2a 22.5 ± 0.6a 8.0 ± 0.9a 871.9 ± 80.1
3.3.2.5 Effect of Amino Acid and Dipeptide-based Chain Extenders on Polyurethane Properties
The development of a Gly-Leu-based chain extender and its incorporation into a
segmented PU were successfully demonstrated. PUs incorporating the amino acid and dipeptide
had very similar molecular weight averages, thermal transitions, crystallinity, degree of phase
segregation, and chemical functional groups. The comparable polymer properties of the two PUs
are reflected by the structural similarities between the two chain extenders. As seen in Figure
3.13, both are symmetric diester molecules of similar molecular weight whose reactive amine-
terminated ends promote uniform polymerization and hydrogen bond forming urea functional
groups. In addition, they both contain the rigid cycloaliphatic linker CDM and have bulky
hydrophobic pendant side chains. The ester linkages in both PUs offer additional hydrogen
bonding and introduce potential sites for hydrolytic cleavage. These ester groups have been
shown to be cleaved by chymotrypsin in the Phe PU [26] and may promote chain cleavage in
vivo by other biological agents, such as the esterases released by inflammatory cells [27], thus
enhancing overall PU degradation. Subtle differences in mechanical properties, however, were
observed between the two PUs. This may be accounted for by the presence of the amide
functional groups in the Gly-Leu PU, which adds addition H-bonding to the polymer and
111
enhances hard segment cohesion. In general, though, the Gly-Leu PU and Phe PU exhibited very
similar polymer properties based on the analytical techniques employed here.
In synthesizing the PUs, PCL1250 was the only soft segment investigated. Previous
work demonstrated a family of PUs using LDI and the Phe-based chain extender as the hard
segment could be formed that exhibit a wide range of physiochemical, thermal, mechanical, and
degradation characteristics by altering the soft segment chemistry and soft segment molecular
weights [6, 28]. Increasing PCL molecular weight to 2000 g/mol resulted in elastomers that had
an increase in crystallinity, hydrophobicity, initial modulus, ultimate tensile stress and strain, and
resistance to degradation. Decreasing PCL molecular weight to 530 had the opposite effect on
PU properties. Polyethylene oxide (PEO)-based PUs were amorphous, soft, and tacky, were
much more hydrophilic, and had faster degradation rates than the PCL-based PUs [6, 28]. These
trends have been verified by several other groups using triblock soft segments of PCL-PEO-PCL
and different hard segments [5, 29-31]. This work suggests that polymer properties using the
Gly-Leu-based chain extender may also be adjusted by altering the soft segment chemistry and
molecular weights as required for particular applications. Ultimately, a family of PUs with
diverse characteristics may be achieved that contain a Gly-Leu dipeptide.
The method of incorporating the Gly-Leu dipeptide into the chain extender was chosen
based on previous work in our laboratory. A few other methods have been demonstrated to
introduce peptides into the backbone structure of segmented PUs. A direct comparison to other
amino acid and peptide containing PUs is difficult due to the differences in soft segments and
diisocyanates used in these studies but a few points may be made about these other techniques.
In the study by Lipatova et al. [15], amino acid and dipeptides were used directly as chain
extenders in the production of segmented PUs. Isocyanates can react with several different
functional groups and in this situation, reacted with the amine end to form urea groups and the
carboxylic acid end to make amide linkages. Although details of the polymer synthesis were not
provided in this report, the PUs were all of lower molecular weight, ~10,000 – 20,000 g/mol
[15], which is in the range where molecular weight averages have a large impact on overall
polymer properties [23]. Typical elastomeric PUs have number average molecular weights
ranging from 20,000 – 100,000 g/mol [23], and this approach may not be desirable for achieving
high molecular weight polymers that exhibit good elastic mechanical properties. Chain
extenders with different end groups may lead to lower molecular weights and high
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polydispersities due to different end group reactivities and irregular polymerizations [32, 33].
An important finding made by Lipatova et al. [15], however, was that increased phase separation
was observed when moving from a Phe amino acid chain extender to a Phe-Phe dipeptide and
was attributed to the addition of the peptide linkage. In comparing the Phe and Gly-Leu PUs, a
similar degree of phase segregation was identified between the two PUs, but the presence of the
amide bonds in the Gly-Leu PU may account for the enhanced mechanical properties.
A second method for introducing peptides into the chain extender involves the addition of
a lysine residue at the C-terminus of a specific peptide sequence to create a diamine structure.
Guan and Wagner [5] used this approach with an Ala-Ala-Lys sequence to make PUs that
display elastase sensitivity. The PU made with PCL of molecular weight 2000, 1,4-
butanediisocyanate, and the AAK chain extender (AAK PU) is structurally the closest polymer
of this family to the Phe and Gly-Leu PUs. Although the AAK PU had lower molecular weight
averages than the Phe and Gly-Leu PUs, other properties were fairly similar. The AAK PU was
a semicrystalline phase segregated elastomer that had a comparable soft segment Tg, -54˚C, and
breaking strain, 830%, to the Phe and Gly-Leu PUs, but did have a higher ultimate tensile
strength, 28 MPa. The difference in ultimate tensile strength between the AAK PU and the Phe
and Gly-Leu PUs may be attributed to the higher PCL molecular weight. PUs made using
PCL2000, LDI, and the Phe-based chain extender showed a similar high ultimate tensile
strength, ~30 MPa [28]. It is important to note, though, that the AAK chain extender, which has
relatively small side chains, was used with a symmetric, linear diisocyanate, which may have
allowed for good phase segregation and mechanical properties. The symmetric diol linker
method used with the Gly-Leu PU may have helped to achieve adequate PU properties in
conjunction with the asymmetric, bulky side chain containing LDI. Based on the knowledge of
structure-property relationships, one might speculate that a PU made from a Gly-Leu-Lys chain
extender, PCL1250, and LDI may not have resulted in the favorable properties observed with the
Gly-Leu-based chain extender developed here.
3.4 Conclusions Enzyme-susceptible segmented polyurethanes may be made through amino acid and
peptide-containing chain extenders. In an attempt to create MMP-sensitive PUs, a synthesis and
purification method was developed to successfully fabricate an amine terminated, Gly-Leu-based
diester chain extender. PUs incorporating this chain extender had high molecular weight
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averages and were phase segregated polymers. The PUs were soft, flexible elastomers that
exhibited a high breaking strain and high ultimate tensile strength. The Gly-Leu PU had very
similar physical, thermal, and chemical properties compared to the Phe PU. Chain extender
chemical structure reflected the subtle differences in mechanical properties observed between the
two PUs. Previous success with the Phe PU as biomaterial scaffolds and the comparable
properties of the Gly-Leu PU suggest this synthetic polymer may also hold promise as a
biomaterial in soft tissue engineering applications.
3.5 References 1. Parrag, I.C. and K.A. Woodhouse, Development of Biodegradable Polyurethane
Scaffolds Using Amino Acid and Dipeptide-based Chain Extenders for Soft Tissue Engineering. Journal of Biomaterials Science-Polymer Edition, In Press.
2. Rockwood, D.N., R.E. Akins, I.C. Parrag, K.A. Woodhouse, and J.F. Rabolt, Culture on electrospun polyurethane scaffolds decreases atrial natriuretic peptide expression by cardiomyocytes in vitro. Biomaterials, 2008. 29(36): p. 4783-4791.
3. Atala, A., Engineering tissues, organs and cells. Journal Of Tissue Engineering And Regenerative Medicine, 2007. 1(2): p. 83-96.
4. Parker, K.K. and D.E. Ingber, Extracellular matrix, mechanotransduction and structural hierarchies in heart tissue engineering. Philosophical Transactions of the Royal Society B-Biological Sciences, 2007. 362(1484): p. 1267-1279.
5. Guan, J.J. and W.R. Wagner, Synthesis, characterization and cytocompatibility of polyurethaneurea elastomers with designed elastase sensitivity. Biomacromolecules, 2005. 6(5): p. 2833-2842.
6. Skarja, G.A. and K.A. Woodhouse, In vitro degradation and erosion of degradable, segmented polyurethanes containing an amino acid-based chain extender. J Biomater Sci Polym Ed, 2001. 12(8): p. 851-73.
7. Ducharme, A., S. Frantz, M. Aikawa, E. Rabkin, M. Lindsey, L.E. Rohde, F.J. Schoen, R.A. Kelly, Z. Werb, P. Libby, and R.T. Lee, Targeted deletion of matrix metalloproteinase-9 attenuates left ventricular enlargement and collagen accumulation after experimental myocardial infarction. J Clin Invest, 2000. 106(1): p. 55-62.
8. Hayashidani, S., H. Tsutsui, M. Ikeuchi, T. Shiomi, H. Matsusaka, T. Kubota, K. Imanaka-Yoshida, T. Itoh, and A. Takeshita, Targeted deletion of MMP-2 attenuates early LV rupture and late remodeling after experimental myocardial infarction. American Journal Of Physiology-Heart And Circulatory Physiology, 2003. 285(3): p. H1229-H1235.
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9. Kim, H.E., S.S. Dalal, E. Young, M.J. Legato, M.L. Weisfeldt, and J. D'Armiento, Disruption of the myocardial extracellular matrix leads to cardiac dysfunction. Journal Of Clinical Investigation, 2000. 106(7): p. 857-866.
10. Woessner, J. and H. Nagase, Specificity requirements of the MMPs, in Matrix Metalloproteinases and TIMPs. 2000, Oxford University Press: New York. p. 98-108.
11. Huang, S., D. Bansleben, and J. Knox, Biodegradable Polymers: Chymotrypsin Degradation of a Low Molecular Weight Poly (ester-Urea) Containing Phenylalanine. J. Appl. Polym. Sci., 1979. 23: p. 429-437.
12. Skarja, G.A. and K.A. Woodhouse, Synthesis and characterization of degradable polyurethane elastomers containing an amino acid-based chain extender. J. Biomater. Sci. Polym. Ed., 1998. 9(3): p. 271-95.
13. Skarja, G.A. and K.A. Woodhouse, Structure-property relationships of degradable polyurethane elastomers containing an amino acid-based chain extender. J. Appl. Polym. Sci., 2000. 75: p. 1522-1534.
14. Crescenz.V, G. Manzini, Calzolar.G, and C. Borri, Thermodynamics of Fusion of Poly-beta-propiolactone and Poly-epsilon-caprolactone- Comparative Analysis of Melting of Aliphatic Polylactone and Polyester Chains. Eur. Polym. J., 1972. 8(3): p. 449-&.
15. Lipatova, T., G. Pkhakadze, D. Vasil'chenko, V. Vorona, and V. Shilov, Structural peculiarities of block copolyurethanes with peptide links as rigid block extenders. Biomaterials, 1983. 4: p. 201-204.
16. Skarja, G.A. and K.A. Woodhouse, Synthesis and characterization of degradable polyurethane elastomers containing an amino acid-based chain extender. J Biomater Sci Polym Ed, 1998. 9(3): p. 271-95.
17. Arabuli, N., G. Tsitlanadze, L. Edilashvili, D. Kharadze, T. Goguadze, V. Beridze, Z. Gomurashvili, and R. Katsarava, Heterochain polymers based on natural amino acids. Synthesis and enzymatic hydrolysis of regular poly(ester amide)s based on bis(L-phenylalanine) a,w-alkylene diesters and adipic acid. Macromol. Chem. Phys., 1994. 195: p. 2279-2289.
18. Huang, S., D. Bansleben, and J. Knox, Biodegradable Polymers: Chymotrypsin Degradation of a Low Molecular Weight Poly (ester-Urea) Containing Phenylalanine. J Appl Polym Sci, 1979. 23: p. 429-437.
19. Parades, N., A. Rodriguez-Galan, and J. Puiggali, Synthesis and Characterization of a Family of Biodegradable Poly(ester amide)s Derived from Glycine. J Polym Sci A: Polym Chem, 1998. 36: p. 1271-1282.
20. Parades, N., A. Rodriguez-Galan, J. Puiggali, and C. Peraire, Studies on the Biodegradation and Biocompatibility of a New Poly(ester amide) Derived from L-Alanine. J Appl Polym Sci, 1998. 69: p. 1537-1549.
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21. Skarja, G.A., The development and characterization of degradable, segmented polyurethanes containing amino acid-based chain extenders, in Department of Chemical Engineering and Applied Chemistry. 2001, University of Toronto: Toronto.
22. Sutherland, K., J.R. Mahoney, 2nd, A.J. Coury, and J.W. Eaton, Degradation of biomaterials by phagocyte-derived oxidants. J Clin Invest, 1993. 92(5): p. 2360-7.
23. Lamba, N.M.K., S.L. Cooper, M.D. Lelah, and K.A. Woodhouse, Polyurethanes in biomedical applications. 1998, Boca Raton: CRC Press. 277.
24. Skarja, G.A. and K.A. Woodhouse, In vitro degradation and erosion of degradable, segmented polyurethanes containing an amino acid-based chain extender. J. Biomater. Sci. Polym. Ed., 2001. 12(8): p. 851-73.
25. Socrates, G., Infrared characteristic group frequencies. 2nd ed. 1994, Chichester; New York: Wiley. viii, 249 p.
26. Elliott, S.L., J.D. Fromstein, J.P. Santerre, and K.A. Woodhouse, Identification of biodegradable products formed by L-phenylalanine based segmented polyurethanes. Journal of Biomaterial Science Polymer Edition, 2002. 13: p. 691-711.
27. Santerre, J.P., K. Woodhouse, G. Laroche, and R.S. Labow, Understanding the biodegradation of polyurethanes: From classical implants to tissue engineering materials. Biomaterials, 2005. 26(35): p. 7457-7470.
28. Skarja, G.A. and K.A. Woodhouse, Structure-property relationships of degradable polyurethane elastomers containing an amino acid-based chain extender. J Appl Polym Sci, 2000. 75: p. 1522-1534.
29. Abraham, G.A., A. Marcos-Fernandez, and J. San Roman, Bioresorbable poly(ester-ether urethane)s from L-lysine diisocyanate and triblock copolymers with different hydrophilic character. Journal Of Biomedical Materials Research Part A, 2006. 76A(4): p. 729-736.
30. Cohn, D., T. Stern, M.F. Gonzalez, and J. Epstein, Biodegradable poly(ethylene oxide)/poly(epsilon-caprolactone) multiblock copolymers. Journal Of Biomedical Materials Research, 2002. 59(2): p. 273-281.
31. Guan, J.J., M.S. Sacks, E.J. Beckman, and W.R. Wagner, Biodegradable poly(ether ester urethane)urea elastomers based on poly(ether ester) triblock copolymers and putrescine: synthesis, characterization and cytocompatibility. Biomaterials, 2004. 25(1): p. 85-96.
32. Shi, F.Y., L.F. Wang, E. Tashev, and K.W. Leong, Synthesis And Characterization Of Hydrolytically Labile Poly(Phosphoester Urethanes). Acs Symposium Series, 1991. 469: p. 141-154.
33. Wirpsza, Z. and T.J. Kemp, Polyurethanes: chemistry, technology, and applications. Ellis Horwood series in polymer science and technology. 1993, Chichester; New York: E. Horwood. xv, 517 p.
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Chapter 4: Electrospinning Phe and Gly-Leu Polyurethanes § Sections of this chapter have been accepted for publication [1]: Parrag IC, Woodhouse KA.
Development of Biodegradable Polyurethane Scaffolds using Amino Acid and Dipeptide-based
Chain Extenders for Soft Tissue Engineering. Journal of Biomaterials Science – Polymer Edition,
in press.
† Sections of this chapter also contributed to the publication [2]: Rockwood DN, Akins RE,
Parrag IC, Woodhouse KA, Rabolt JF. Culture on Electrospun Polyurethane Scaffolds
Decreases Atrial Natriuretic Peptide Expression by Cardiomyocytes In Vitro. Biomaterials,
2008, 29(36): p. 4783-4791.
4.0 Abstract Electrospinning was used to form porous three-dimensional PU scaffolds. The
concentration of the Phe and Gly-Leu PUs in the electrospinning system was systematically
adjusted to form scaffolds with different structural features. Consistent with the literature, the
PUs went through concentration-dependent structural transitions with the formation of beads, to
beads-on-a-string, to good fibers lacking beaded features with an increase in concentration.
Average fiber diameter showed a direct correlation to electrospinning concentration. The Phe
and Gly-Leu PUs formed from a 14% and 10% w/v concentration respectively both had
randomly organized fibers, an average fiber diameter of approximately 3.6 µm, similar molecular
weight averages and thermal properties, and were chosen for subsequent degradation and cell-
based studies. In addition, the Phe PU was electrospun into scaffolds of varying architecture to
investigate how fiber alignment influences the orientation response of cardiac cells. The Phe PU
was electrospun into physically aligned and unaligned scaffolds. Uniaxial tensile testing of the
scaffolds was conducted in the preferred and cross-preferred direction of fiber orientation. The
unaligned scaffold was soft, flexible, and highly extensible in both directions while the aligned
scaffold was much stiffer, not very extensible, and had a high ultimate tensile stress when
stretched in the direction of fiber orientation. The molecular weight averages, thermal
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properties, and average fiber diameter were similar between the Phe PU substrates of different
architecture and were used for subsequent cell-based studies.
4.1 Introduction Tissue engineering scaffolds act as a temporary ECM for the attachment, organization,
and delivery of a high density of cells to an injured or diseased tissue. Some of the parameters
which must be considered when forming scaffolds for tissue engineering include: a high surface
area-to-volume ratio to achieve a high density of appropriately coupled cells and efficient
transport of oxygen, nutrients, and waste; pore sizes that allow cell migration; surface
chemistries and features that promote and guide cell adhesion, growth, and differentiation; and a
suitable architecture to organize the engineered tissue [3]. In addition, the combination of
starting biomaterial and polymer processing method will dictate if the scaffold has the proper
mechanical properties and degradation rates for the tissue under investigation.
Elastomeric segmented polyurethanes are thermoplastic polymers and may be processed
by conventional solvent-based methods. Two of the more promising techniques used in making
PU scaffolds for cardiac and other soft tissue engineering applications are electrospinning and
thermally induced phase separation (TIPS). Fujimoto et al. [4, 5] demonstrated PU scaffolds
formed by TIPS have the appropriate physical, mechanical, and biocompatibility properties for
use as a patch in normal and infarcted hearts. In another study using TIPS, Guan et al. [6] used
an elastase-sensitive PU to form porous anisotropic scaffolds that may be used to guide
anisotropic tissue formation. These scaffolds completely degraded in vivo in 8 weeks by passive
hydrolysis, enzyme-mediated cleavage, and other mechanisms and may have appropriate
degradation rates that allow the transfer of structural and mechanical functionality to the
engineered tissue [6]. Previous work in our laboratory by Fromstein et al. [7] described the
formation of Phe PU scaffolds by both electrospinning and TIPS and investigated how the
different scaffold morphologies influenced murine embryonic stem cell-derived cardiomyocytes
(mESCDCs). Cardiac cells on both scaffold types were contractile and expressed the proteins
sarcomeric myosin heavy chain and Cx-43 but exhibited markedly different morphologies. In
contrast to the round cells found on TIPS scaffolds, mESCDCs were more elongated on the
electrospun PU fibers and were characteristic of more differentiated and mature mESC-derived
cells [7]. As a result, while both TIPS and electrospinning may be used to form anisotropic PU
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scaffolds that may help to control cell organization, electrospinning may be preferred in
promoting a differentiated and mature cardiomyocyte phenotype.
Electrospinning is a relatively simple, cost effective, and versatile method for processing
polymers into fibers with diameters on the nanometer or micrometer scale [8]. The non-woven,
3-D scaffolds that are formed from electrospinning have a topography and porosity that mimics
the native ECM and has been shown to positively influence cell behavior [8, 9]. This polymer
processing technique provides control over fiber diameters, structural features, and fiber
orientation and has become widely used for tissue engineering [8, 9].
Previous experience in our lab with scaffold fabrication techniques suggested
electrospinning as a very promising approach to scaffold formation in the development of a
cardiac patch [7]. Consequently, this chapter describes the successful employment of
electrospinning in developing porous 3-D PU scaffolds. After investigating several
electrospinning parameters, scaffolds with good fiber formation were developed with the Phe
and Gly-Leu PUs. Characterizing the morphology and fiber diameter allowed the selection of
Phe and Gly-Leu PU scaffolds that shared similar structural features and were deemed
appropriate for subsequent degradation and cell-based studies. The Phe PU was also used to
form scaffolds with aligned and unaligned fibrous structures. The scaffolds developed here were
used for cell-based studies with mESCDCs and MEFs, discussed in chapter 6, and primary
cardiac cells, which contributed to the publication by Rockwood et al. [2].
4.2 Materials and Methods All materials were purchased from Sigma-Aldrich Canada (Oakville, ON, Canada) unless
otherwise stated.
4.2.1 Electrospinning Phe and Gly-Leu Polyurethane Scaffolds§† Initial tests were conducted using the Phe PU to identify the appropriate conditions for
forming porous, 3-D fibrous scaffolds. A custom built electrospinner (Spark Engineering, Glen
Allen, VA) with a rotating and translating cylindrical collection plate was used for this work
(Figure 4.1). Previous experience electrospinning the Phe PU by our collaborators in the Rabolt
laboratory at the University of Delaware provided starting conditions for this work [7, 11].
§ The Phe and Gly-Leu PU scaffolds electrospun from 14% and 10% w/v were used in [1] and † the conditions established here for producing aligned and unaligned Phe PU scaffolds contributed to the scaffolds used in [2].
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Differences in electrospinner design required the optimization of these starting parameters. A
few of the electrospinning parameters that were investigated include concentration, solvent type,
flow rate, distance from needle tip to collection mandrel, solution temperature, and needle gauge.
Ultimately, conditions appropriate for obtaining a porous scaffold were identified. The two PUs
were subsequently dissolved in dichloromethane (DCM) at concentrations from 2-16% w/v. The
polymer solutions were loaded into a syringe (Becton Dickenson, Franklin Lakes, NJ) with a 1”
22 gauge blunt end needle (Kontes Glass Company, Vineland, NJ) and placed into a syringe
pump (KD Scientific, Holliston, MA) running at 3 ml/h. The grounded collection mandrel was
stationary and was positioned 30 cm from the tip of the needle. Reynolds Wrap non-stick
aluminum foil was placed over the collection area to aid in removing the polyurethane from the
mandrel. Using a high voltage power supply (Gamma High Voltage Research Inc, Ormond
Beach, FL), a 12 kV voltage was applied to the needle tip generating an electric field between
the needle and grounded collection plate. The process was carried out under ambient conditions
(approximately 25˚C and 60% relative humidity). The resulting mats were placed in a vacuum
chamber overnight to remove residual solvent and stored in a desiccator until use. Scaffolds
used in degradation and cell based studies were made from 14% and 10% w/v concentrations in
DCM for the Phe and Gly-Leu PUs respectively. Electrospinning was stopped when scaffold
thickness reached ~90 µm as determined by measuring with a micrometer.
Figure 4.1: Illustration of electrospinning apparatus. Image used with permission from Kenawy et al. [12].
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Aligned and unaligned scaffolds were made using the Phe PU. The Phe PU was
dissolved in DCM at 14% w/v and was electrospun using the conditions described above.
Randomly oriented fibers were obtained by using a collection mandrel rotating and translating at
5 cm/s. To achieve fiber alignment, the mandrel rotational velocity was increased to 270 cm/s
and the translation velocity was decreased to 3 cm/s. The electrospinning process was carried
out under ambient conditions until the deposited polymer reached approximately 100 µm in
thickness (~8 h) when measured with a micrometer. The resultant polymer mats were placed in a
vacuum chamber overnight to remove residual solvent and stored in a desiccator until use.
4.2.2 Scaffold Characterization§† Scaffold architecture and fiber characteristics were examined using scanning electron
microscopy (SEM). Discs (5 mm in diameter) were punched out of the electrospun mats using
biopsy punches (Fray Products Corp., Buffalo, NY) away from the edges of the scaffolds. These
discs were then mounted onto metal stubs and sputter coated with a 5 nm layer of platinum.
Samples were analyzed using an S-2500 scanning electron microscope (Hitachi, Japan) with an
accelerating voltage of 10 kV and images were taken using the Quartz PCI software (Quartz
Imaging Corporation, Vancouver, BC).
Fiber diameters were measured manually by selecting 100 individual fibers randomly
away from points of fusion from at least 3 different images in the Phe and Gly-Leu PU scaffolds
of different concentrations using Image Pro Express image analysis software (Media
Cybernetics, Bethesda, MD). The image analysis software was similarly used to manually
measure fiber diameters and fiber axis angles from at least 150 randomly selected fibers from 6
different images with the aligned and unaligned Phe PU scaffolds. For fiber angle measurements
taken on fibers whose angles were changing, random areas of the images were zoomed in on and
the major angle of fiber axis within this focused region was measured. Fiber diameter averages,
distributions, and alignment were quantified from the fiber measurements. Statistical
comparisons of fiber diameters were made using either a two-tailed independent t-test or one-
way analysis of variance (ANOVA) with Bonferroni or Dunnett post-hoc analysis using the
SPSS Statistics 17.0 statistical software package (SPSS Inc, Chicago, IL). Comparison of fiber
§ Characterization of the Phe and Gly-Leu PU scaffolds was published in [1] and † part of the characterization of the aligned and unaligned Phe PU scaffolds was published in [2].
121
angles was performed with a two-tailed independent t-test using the absolute value of fiber
angles.
Molecular weight averages and thermal properties of the electrospun scaffolds were
determined by GPC and DSC analysis respectively as described in chapter 3. The mechanical
properties of the aligned and unaligned electrospun scaffolds were determined by uniaxial tensile
testing in the preferred and cross-preferred direction of fiber orientation. All parameters during
testing were conducted according to ASTM D1708 with the exception of sample dimensions.
Rectangular samples 3 cm long and 1 cm wide were used instead of the dumbbell-shaped
specimens to conserve material. Scaffolds were preconditioned at ~24˚C and ~55% relative
humidity for 48 h. Testing was carried out using an Instron 4301 testing machine with a
crosshead speed of 13 mm/min and a 10 N load cell. Samples were placed between grips 2 cm
apart, were stretched to break, and stress-strain curves were generated from force-displacement
data. All sample failure occurred away from the grips. Stress and strain at break and initial
modulus were inferred from the stress-strain graphs. Statistical comparisons were made with a
one-way ANOVA with Bonferroni post hoc analysis.
4.3 Results and Discussion
4.3.1 Electrospinning Polyurethane Scaffolds Electrospinning is a polymer processing technique used in tissue engineering to create
fibrous scaffolds while controlling fiber size and organization as a means to mimic the native
extracellular matrix. In collaboration with Dr. Rabolt’s lab at the University of Delaware, the
electrospinning method was used to develop scaffolds with the Phe PU. In the study by
Rockwood el al. [11], the Phe PU was electrospun using 18% and 20% w/v concentrations in
DCM, a 23 gauge needle, a positive 10 kV applied voltage, and a 25 cm distance from needle tip
to collection plate. The resulting scaffolds generally lacked the beaded formations that occur at
lower polymer concentrations, but exhibited extensive fiber fusion and melting that masked the
porosity and fibrous nature of the scaffolds [11]. In subsequent work, the Phe PU was
electrospun using the same conditions but with a 15% w/v concentration and 20 cm distance to
collection plate [7]. Under these conditions, the scaffolds exhibited more uniform fiber
structures that were missing the melted morphology observed previously. As a result, these new
conditions were initially investigated to electrospin the Phe PU with the custom built
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electrospinner in our laboratory. Representative SEM images of the scaffolds formed from the
same conditions using the electrospinner in the Rabolt lab and the one in our lab are shown in
Figure 4.2. A markedly different scaffold morphology was obtained when using the two
different systems. Both electrospinners produced fibers without any beads but the fibers formed
with our system appeared much smaller and had a more melted appearance than the mat
fabricated in the Rabolt lab. These results indicated that the Phe PU could be successfully
electrospun using our apparatus but suggested that differences between the two setups required
some optimization of parameters to reduce this melted scaffold appearance.
Figure 4.2: A comparison of electrospun Phe PU mats formed in a) the Rabolt laboratory and b) our laboratory
under the same conditions. A distinct difference in scaffold morphology resulted from the different electrospinners.
To optimize the Phe PU scaffolds and better understand our electrospinning system,
several parameters were adjusted including concentration, solvent type, flow rate, distance from
needle tip to collection mandrel, solution temperature, and needle gauge. Although a systematic
investigation of all these parameters was not conducted, many of the conditions tested led to
ubiquitous fiber fusion and a melted scaffold morphology (Figure 4.3a). Fiber fusion may result
from insufficient solvent evaporation or could be due to polymer chain rearrangement, as occurs
with PU chains in an attempt to reduce interfacial free energy when placed in different
environments [13]. Ultimately, conditions were identified for achieving a highly porous scaffold
with a fibrous morphology (Figure 4.3b). The Phe PU was electrospun from a 14% w/v
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concentration in DCM, a 22 gauge needle, a 3 ml/hr flow rate, a +12 kV applied voltage, and a
30 cm distance from needle tip to collection plate under ambient conditions (~25˚C and ~60%
relative humidity). Although concentration was varied, these other conditions were held constant
in forming subsequent PU scaffolds.
Figure 4.3: Comparison of Phe PU scaffolds formed before and after optimizing electrospinning parameters. a)
unoptimized parameters with scaffolds displaying ubiquitous fiber fusion and a film-like morphology. b) optimized scaffold parameters leading to good fiber formation with reduced fiber fusion.
4.3.1.1 Effect of PU Concentration on Scaffold Morphology§ To help identify appropriate scaffolds for degradation and cell-based studies, the Phe and
Gly-Leu PUs were electrospun from concentrations of 2-16% w/v in DCM using the previously
established conditions. The resulting morphology was characterized by SEM and is shown in
Figure 4.4. Both polymers displayed similar trends with an increase in concentration but the
transitions occurred at slightly different concentrations. At low concentrations with insignificant
chain entanglements, electrospraying occurred. This was observed at concentrations less than
6% w/v for the Phe PU and less than 4% for the Gly-Leu PU by the presence of spherical
structures and the absence of polymer fibers. As the concentration was increased from the
electrospraying concentrations (8% and 6% for Phe and Gly-Leu PU respectively), a beads-on-a-
string morphology was observed with spherical structures found along polymer fibers. Further § Part of this section was published in [1].
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Figure 4.4: SEM images of Phe (top) and Gly-Leu (bottom) PU scaffolds electrospun from different concentrations. The diameters of 100 individual fibers from 3 different images were measured. Average fiber diameter ± standard
deviation is given for defect-free fibers. A significant difference in average fiber diameter was observed between all concentrations for Phe PU (ANOVA, p<0.05) with the exception of the 14-16% comparison. Similarly, a significant
difference in average fiber diameter was observed for all concentrations quantified for the Gly-Leu PU (p<0.05).
125
increases in concentration led to fibers lacking the beaded formations. This trend of morphology
transitions from electrospraying, to beads-on-a-string, to good fiber formation with increasing
concentration is consistent with other electrospun polymers in the literature [14]. More
specifically, McKee et al. [15] correlated electrospinning morphology to the concentration
regime of the polymer solution for polyester copolymers: lower polymer concentrations in the
semi-dilute unentangled regime caused the polymer to be electrosprayed; concentrations just
above the transition to the semi-dilute entangled regime led to a beads-on-a string morphology;
good fiber formation lacking the beaded defects occurred at concentrations 2-2.5 times the start
of the semi-dilute entangled regime. Rockwood et al. [11] confirmed the concentrations for
defect-free fibers formed from the Phe PU had specific viscosities that were greater than 2-2.5
times the start of the semi-dilute entangled regime. The transitions in polymer morphology with
the Phe and Gly-Leu PUs observed here likely correspond to these different polymer solution
regimes. The difference in concentrations at which these transitions occur for the two PUs may
be the result of observed differences in solubility. The Gly-Leu PU was completely dissolved in
DCM at all the concentrations tested but the Phe PU was not completely soluble in DCM, where
it appeared partially in suspension. This solubility influences surface tension and viscosity of the
polymer solutions and will affect the electrospinning process and resulting fiber morphology
[14].
Fiber diameters were measured in the electrospun mats corresponding to the
concentrations that led to defect-free fibers. Average fiber diameter was obtained after
measuring 100 different fibers and is quantified in Figure 4.4. A trend of increasing average
fiber diameter was observed with increasing concentration and is consistent with the literature on
electrospun polymers [14, 16]. A significant difference in average fiber diameter was observed
between all concentrations investigated for the Phe PU (ANOVA, p<0.05) with the exception of
the 14-16% comparison (p>0.05). Similarly, a significant difference in average fiber diameter
was observed between all concentrations quantified for the Gly-Leu PU (p<0.05). Therefore,
altering PU concentration has a significant impact on average fiber diameters that are formed.
Characterizing the fiber diameters further identifies a range of microfibers formed under each
polymer concentration tested. The distribution of fiber diameters within each concentration are
shown in Figure 4.5. The range of fibers obtained may be the result of polymer splaying, or fiber
splitting, observed while electrospinning the two PUs at different concentrations. The splaying
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Figure 4.5: Fiber diameter distributions of the Phe and Gly-Leu PU scaffolds electrospun from varying
concentrations. a) Phe PU electrospun in concentrations from 12-16% w/v in DCM and b) Gly-Leu PU electrospun in concentrations from 10-14% w/v in DCM. A range of microfibers are observed under all conditions.
127
phenomenon is caused by high repulsive forces on the polymer chains due to an increase in
surface charge density as the polymer jet travels through the electric field [17]. Deitzel et al.
[16] identified that polymer splaying occurred at higher concentrations in the electrospinning
range with a polyethylene oxide/water system and led to a bimodal fiber diameter distribution.
Although a bimodal distribution was not seen here, polymer splaying likely contributed to the
range of fiber diameters obtained with the two PUs. Fibers at all concentrations fused to other
fibers at points of intersection, which may be due to incomplete solvent evaporation or polymer
chain rearrangement. The different electrospun PU scaffolds had a porous, 3-D structure and
were considered promising for future experiments.
Fiber diameter has been shown to affect the physical, mechanical, and degradation
properties of biomaterial scaffolds [18] and also plays a role in influencing cell behavior [19,
20]. Therefore, the Phe and Gly-Leu PU concentrations that yielded scaffolds with the most
similar physical morphology and average fiber diameters were identified to reduce the variables
that may influence scaffold degradation and cell-response to the PUs. A comparison of fiber
diameter means was conducted by a one-way ANOVA with Dunnett post hoc analysis using the
different Gly-Leu PU concentrations as a control to compare to the Phe PU concentrations. The
Gly-Leu PU scaffolds electrospun at 12 and 14% w/v concentrations had average fiber diameters
that were significantly different than average fiber diameters of all the Phe PU scaffold
concentrations measured (p<0.05). The 10% w/v Gly-Leu PU scaffold, however, was not
significantly different than the Phe PU scaffolds electrospun at 12 and 14% w/v (p>0.05).
Importantly, this Gly-Leu PU scaffold had approximately the same average fiber diameter as the
Phe PU formed at 14% w/v (3.6 ± 1.2 µm and 3.6 ± 1.4 µm for the Gly-Leu and Phe PUs
respectively). In addition both of these scaffolds had a comparable randomly organized
morphology (Figure 4.6a) and a similar fiber diameter distribution (Figure 4.6b). The structural
similarities between the Gly-Leu and Phe PUs electrospun at 10 and 14% respectively provided
justification for their use in subsequent studies.
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Figure 4.6: Comparison of structural features of the Phe and Gly-Leu PU scaffolds used for degradation and cell-based studies. a) SEM images of Phe and Gly-Leu PU scaffolds formed from 14% and 10% w/v concentrations
respectively. Data represents average fiber diameter ± standard deviation. No statistical difference in average fiber diameter was observed between the two scaffolds. b) Fiber diameter distribution of Phe and Gly-Leu PU scaffolds.
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4.3.1.2 Molecular Weight Averages and Thermal Properties Degradation characteristics and polymer performance are influenced by the molecular
weight, crystalline structure, and thermal transitions of biomaterials. The electrospun PU
scaffolds were analyzed by GPC and DSC to help identify how the electrospinning process
affected these polymer properties. Table 4.1 summarizes the GPC and DSC results of the pre-
processed PU films and post-processed PU scaffolds. No major deviations in molecular weight
averages or thermal properties were noted for the electrospun scaffolds when compared to the
PU films, consistent with previous findings with the Phe PU [11]. The high molecular weight
averages observed with the PU scaffolds suggested the shear stresses associated with this
processing technique did not lead to degradation of the polymer. Similarly, no changes in the
thermal transitions indicated that the PUs remain as soft, flexible polymers at physiological
temperature and were phase segregated, semi-crystalline materials that may exhibit good
elastomeric mechanical properties. A slight decrease in overall percent crystallinity was noted
for the electrospun Gly-Leu PU scaffolds. This may be a result of the polymer chains not having
enough time to form ordered 3-D crystal structures before solidifying during solvent evaporation
[16, 18]. The slight change in crystallinity (~3%), however, was not suspected to have a major
impact on scaffold performance. These results suggested that the PUs retained several essential
polymer properties of the base material. Table 4.1: GPC and DSC results of Phe and Gly-Leu PU films and scaffolds. Data is given as mean ± standard
deviation. n=3.
Gel Permeation Chromatography Differential Scanning Calorimetry
PU Scaffold Type/Morphology
Mn (kDa)
Mw (kDa) Polydispersity Tg (ºC) Tm
(ºC) Crystallinity
(%)
Phe PU Films 76.1 ± 4.3
128.6 ± 1.8 1.696 ± 0.099 -50.6 ±
1.2 43.0 ±
0.7 16.0 ± 0.9
Phe PU Scaffolds 79.5 ± 7.0
135.8 ± 6.3 1.709 ± 0.171 -49.5 ±
1.4 42.5 ±
1.9 15.8 ± 2.5
Gly-Leu PU Films 82.7 ± 6.0
126.6 ± 9.4 1.531 ± 0.159 -51.5 ±
1.1 42.0 ±
1.5 16.5 ± 1.5
Gly-Leu PU Scaffolds
82.3 ± 1.9
125.4 ± 4.6 1.523 ± 0.066 -52.4 ±
0.3 40.3 ±
0.1 13.6 ± 0.8
4.3.1.3 Fiber Size in Electrospun PU Scaffolds for Soft Tissue Engineering Fibrous scaffolds with micron-sized fibers have been shown to be promising in tissue
engineering applications. Pham et al. [20] observed that cells attached to electrospun microfibers
and the larger pore sizes within these scaffolds were more conducive to cellular infiltration than
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those with nanofibers. Henry et al. [21] formed electrospun PU scaffolds with fiber diameters
ranging from 3-10 µm and found a higher number and improved morphology of fibroblasts
cultured on the porous scaffolds compared to films of the same material. Fromstein et al. [7]
successfully cultured mESCDCs onto electrospun Phe PU scaffolds that had fiber diameters
ranging from 2-10 µm. The mESCDCs on these scaffolds were adherent and had a sarcomeric
phenotype that appeared more mature than cells on TIPS scaffolds. Increased surface area
associated with the Phe PU microfiber scaffolds also led to significantly higher enzyme-mediated
degradation than films of the same material [11]. These results suggest that the Phe and Gly-Leu
PU microfiber-based scaffolds formed here with fibers ranging from 1-7 µm may also be
successfully employed in degradation and cell-based studies for which they were developed.
The focus on producing nanofibers as a means of mimicking the size of fibril-forming
proteins found in the native ECM has had a positive influence on cell attachment, morphology,
and proliferation [8, 9]. While micron-sized fibrous scaffolds were formed from the two PUs for
this work, electrospinning parameters can be adjusted that may allow the formation of
nanofibrous scaffolds using these polymers. Changing the solvent is one method that has been
shown to successfully move from microfiber to nanofiber-based scaffolds. Bolgen et al. [18], for
example, found that electrospinning pure PCL in chloroform produced microfibrous scaffolds
but switching to a co-solvent of chloroform and DMF caused a shift to nanofiber-based
scaffolds. Adjusting the relative percent of DMF in this co-solvent system allowed the formation
of fibers with a variety of diameters on the hundreds of nanometer scale range [18]. Stankus et
al. [22] electrospun a biodegradable PU elastomer in hexafluoroisopropanol at 5% w/v and
achieved good fiber formation with diameters from 100-900 nm. Cha et al. [23] reported a
tetrahydrofuran-DMF co-solvent system for electrospinning PU scaffolds with a range of fibers,
several of which were on the submicron scale. Others have similarly reported electrospinning
submicron fibers with biodegradable PUs by various solvent systems and electrospinning
parameters [24]. Future work should be conducted to explore different electrospinning
parameters for achieving submicron-sized fibers using the Phe and Gly-Leu PUs.
4.3.2 Aligned and Unaligned Phe PU Scaffolds The emerging evidence that fiber orientation can be used to influence cell alignment and
aid in the development of anisotropic tissues suggests aligned scaffolds may be preferred over
random scaffolds in the anisotropic myocardium [25-27]. Physical cues provided by underlying
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substrates have a strong influence over cellular organization in cardiac tissue engineering [28]
and this organization is critical to the functionality of the tissue [29]. Elastic biodegradable
biomaterials with an aligned orientation may yield a material with the appropriate mechanical
properties and organizational cues for the development of myocardial constructs. To further
investigate the potential of electrospun PU scaffolds for cardiac tissue engineering, the Phe PU
was formed into scaffolds with aligned and unaligned orientations. The physical, thermal, and
mechanical characteristics of the scaffolds were determined to identify how fiber alignment
affects these biomaterial scaffold properties.
4.3.2.1 Scaffold Morphology† The Phe PU was electrospun at a 14% w/v concentration in DCM onto a rotating and
translating grounded mandrel collection plate using the conditions described above. By
adjusting the speed of the electrospinning collection mandrel, scaffolds were formed into two
different morphologies with a general trend of either aligned or unaligned polymer fibers. The
unaligned scaffold was formed at a low rotational and translational mandrel velocity (~5 cm/s for
both motions). Fiber alignment was achieved by increasing the rotational mandrel velocity to
270 cm/s and decreasing the translational velocity to 3 cm/s. Representative SEM images of the
scaffolds formed under these conditions are shown in Figure 4.7. Qualitatively, electrospinning
the PU at a low mandrel rotational speed appeared to result in polymer fibers being randomly
oriented on the collection plate. The fibers collected at a high rotational speed showed a general
alignment in the direction of mandrel rotation. This was achieved at a somewhat lower mandrel
rotational velocity than has been reported in the literature with other electrospun polymers (~12-
18 m/s) [25, 30, 31]. Image analysis software was used in conjunction with the SEM images to
quantify the angle of fiber axis and is presented in Figure 4.8a. The observation of randomly
oriented fibers at a low rotational speed and predominantly aligned fibers at a high speed was
confirmed by this analysis. A fairly uniform distribution of fiber angles was observed with the
unaligned scaffold whereas the aligned scaffold had a more normal distribution with the majority
of fibers within 20º of the reference line. The average absolute value of fiber angle was 16.7 ±
19.9º for the aligned scaffold and 37.2 ± 27.9º for the unaligned scaffold with a significant
difference between the two (t-test, p<0.05). In this analysis, a value of 0º theoretically represents
† The work in this section contributed to [2].
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perfectly aligned fibers while a value of 45º theoretically represents perfectly random fibers,
supporting the trend of fiber alignment at high mandrel rotational speeds and random fibers at
low rotational speeds.
Figure 4.7: SEM images of aligned and unaligned Phe PU scaffolds. a) Phe PU electrospun at a high mandrel
rotational speed and b) at a low mandrel rotational speed. Average absolute fiber angle and average fiber diameter ± standard deviation were calculated from measuring fiber angle and diameter from 150 individual fibers from 6
different images. * Indicates statistical difference in average values using a two-tailed independent t-test (p<0.05).
The electrospun polymers of different architecture were further characterized by
measuring the fiber diameters within each scaffold. As seen in Figure 4.8b, the aligned and
unaligned PU scaffolds had a very similar fiber diameter distribution. A wide fiber diameter
range was observed for both scaffolds, from approximately 600 nm to 7 µm (Figure 4.8),
indicating the presence of both nano and microfibers. Interestingly, electrospinning the Phe PU
under the same conditions but using a stationary collection plate led to the formation of only
micro-fibers without any fibers on the nanometer scale. Adding motion to the mandrel collection
plate led to the formation of nanofibers, which may have been caused by stretching and thinning
of the fibers upon deposition onto the moving mandrel. Rockwood et al. [11] found no
difference in fiber diameters when electrospinning the Phe PU on a stationary or rotating
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Figure 4.8: Characteristics of aligned and unaligned Phe PU scaffolds. a) Quantification of fiber angle. The
aligned scaffold had the majority of fibers within 20º of a reference angle while the unaligned scaffold had a more uniform distribution. b) Fiber diameter size distribution. A similar fiber diameter distribution was observed for the
two scaffold architectures. Fiber diameters and angles were measured from 150 individual fibers from 6 images.
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mandrel and this discrepancy may reflect the different electrospinner setups, which was
demonstrated above to affect scaffold features. The aligned and unaligned scaffolds formed
using the dynamic collection mandrel had an average fiber diameter of 2.9 ± 1.7 µm and 2.7 ±
1.5 µm respectively with no statistical difference between the two (t-test, p>0.05). A high
mandrel rotational speed used to align fibers has been reported in the literature to decrease fiber
diameters due to additional stretching during deposition on the rotating mandrel [31], but this
was not observed here. The lower mandrel rotational velocity (2.7 vs.13 m/s) that induced fiber
alignment may not have been sufficient to see this phenomenon. The formation of aligned and
unaligned scaffolds with a similar average fiber diameter prevented the need to optimize
electrospinning conditions and allowed further investigations into how fiber alignment affected
scaffold properties.
4.3.2.2 Molecular Weight Averages and Thermal Properties The molecular weight averages determined by GPC and thermal properties given by DSC
for the Phe PU films and electrospun scaffolds of different architecture are given in Table 4.2.
The electrospinning process did not appreciably affect molecular weight averages or thermal
properties of the PU scaffolds, even when running at high mandrel rotational speeds. The PU
scaffolds had high molecular weight averages (Mw>130,000 Da) and a low polydispersity (~1.7)
consistent with the polymer films. Similarly, the thermal properties for all PU morphologies
were comparable with a soft segment glass transition temperature of approximately -49ºC, a soft
segment melting temperature of approximately 43ºC, and an overall percent crystallinity of 16%.
Increased crystal structure formation that may be observed at high mandrel rotational speeds due Table 4.2: GPC and DSC results for Phe PU films and electrospun scaffolds of varying architecture. No
appreciable differences in molecular weight averages or thermal properties were observed. Data is given as mean ± standard deviation. n=3.
Gel Permeation Chromatography Differential Scanning Calorimetry
PU Scaffold Type
Mn (kDa)
Mw (kDa) Polydispersity Tg (ºC) Tm
(ºC) Crystallinity
(%)
Phe PU Films 76.1 ± 4.3
128.6 ± 1.8 1.696 ± 0.099 -50.6 ±
1.2 43.0 ±
0.7 16.0 ± 0.9
Unaligned Phe PU Scaffolds
79.5 ± 7.0
135.8 ± 6.3 1.709 ± 0.171 -49.5 ±
1.4 42.5 ±
1.9 15.8 ± 2.5
Aligned Phe PU Scaffolds
75.4 ± 4.2
131.6 ± 6.2 1.746 ± 0.128 -48.3 ±
0.7 43.7 ±
2.2 16.7 ± 1.0
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to increased stretching and molecular chain orientation was not observed here [25, 32]. The high
molecular weight averages and thermal transitions of soft, flexible semi-crystalline polymers
suggest the potential of good PU performance at physiological temperatures in further studies.
4.3.2.3 Mechanical Properties† The tensile properties of the electrospun PU scaffolds were investigated by uniaxial
tensile testing in both the preferred and cross-preferred direction of fiber orientation. The
unaligned scaffold was soft, flexible, and highly extensible in both directions while the aligned
scaffold exhibited these characteristics only in the cross-preferred direction. When stretched in
the direction of fiber orientation, the aligned scaffold was much stiffer and not as extensible.
Typical stress-strain curves for the scaffolds are shown in Figure 4.9. Initial modulus, ultimate
tensile stress, and ultimate tensile strain were inferred from these curves and are summarized in
Table 4.3. Fiber orientation had a significant impact on the mechanical properties of the
scaffolds. The unaligned scaffold exhibited similar mechanical behavior in both directions with
a relatively high ultimate tensile strain, ~300%, but a low ultimate tensile stress, ~1 MPa, and
initial modulus, ~8 and ~4 MPa when stretched in the preferred and cross-preferred directions
respectively. In contrast, the aligned scaffold exhibited a significant difference in all tensile
properties when stretched in the preferred and cross-preferred directions of fiber orientation
(ANOVA, p<0.05). When stretched in the preferred direction, the scaffold was stiff and
incapable of withstanding high deformations. Samples stretched in this direction had a relatively
high initial modulus, ~24 MPa, and ultimate tensile stress, ~4 MPa, but exhibited a relatively low
ultimate elongation, ~60%. When stretched in the cross-preferred direction, the aligned scaffold
was more similar to the unaligned scaffold with a high ultimate tensile strain and a low ultimate
stress and modulus. In addition, a statistically significant difference was observed for all
mechanical properties when comparing Phe PU films to the different scaffold architectures and
stretch directions. The PU films had a significantly higher initial modulus, ultimate stress, and
strain at failure compared to the electrospun scaffolds (ANOVA, p<0.05) suggesting the
electrospinning process had a major impact on resulting mechanical properties of the scaffolds.
† The work in this section contributed to [2].
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Figure 4.9: Representative stress-strain curves for aligned and unaligned PU scaffolds stretched in preferred and
cross-preferred directions of orientation.
Table 4.3: Summary of mechanical properties of aligned and unaligned Phe PU scaffolds stretched in preferred and cross-preferred directions of orientation. Data is given as mean ± standard deviation. N=8 unless otherwise stated. One-way ANOVA with Bonferroni post hoc analysis showed statistical difference (p<0.05) compared to a) aligned scaffold stretched in preferred direction, b) unaligned scaffold stretched in preferred
direction and c) Phe PU films.
Scaffold Architecture Initial Modulus(MPa)
Ultimate Tensile Stress (MPa)
Strain at Failure (%)
Cast Phe PU Films (N=9) 44.09 ± 4.77a,b 8.01 ± 0.94a,b 872 ± 80a,b
Aligned Stretched in Preferred Direction 24.42 ± 5.04b,c 3.94 ± 1.04b,c 61 ± 11b,c
Aligned Stretched in Cross-Preferred Direction 0.60 ± 0.11a,b,c 0.29 ± 0.02a,c 263 ± 25a,c
Unaligned Stretched in Preferred Direction 8.36 ± 1.86a,c 1.07 ± 0.11a,c 298 ± 70a,c
Unaligned Stretched in Cross-Preferred Direction
(N=5) 4.12 ± 0.59a,c 0.94 ± 0.04a,c 345 ± 29a,c
The observed mechanical behavior of the PU scaffolds is typical of electrospun polymers
with aligned and unaligned morphologies. Courtney et al. [25] electrospun a biodegradable PU
elastomer using various mandrel rotational speeds and found a correlation between the degree of
physical alignment of polymer fibers and mechanical properties of scaffolds. A mandrel
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rotational velocity of 3 m/s began to lead to the alignment of fibers and further increases in
rotational speed induced higher degrees of physical anisotropy. As fiber alignment became
higher, scaffolds became increasingly stiff in the preferred direction of fiber orientation and had
the opposite affect in the cross preferred direction. Other groups similarly demonstrated aligned
scaffolds exhibit high strengths and low elongations when stretched in the preferred direction of
orientation and have much higher strains and lower strengths when stretched in the perpendicular
direction [30, 31]. Importantly, a study by Johnson et al. [30] used pure electrospun PCL to
investigated the microstructure of scaffolds under strain to identify the response of polymer
fibers to mechanical stretch. It was determined that electrospun fibers initially not aligned in the
direction of stretch rearrange and undergo a strain-induced orientation of fibers parallel to
applied force [30]. In this situation, applied loads are carried by weak inter-fiber interactions and
structural reorganization events. As more fibers become aligned with increasing strain, the
applied stress is efficiently transferred to unstrained areas of the scaffold where fibers have yet to
be aligned [30]. This leads to scaffolds that have low strengths and high elongations due to the
weak load-bearing features and high strains needed to align the fibers. In contrast, the aligned
scaffold stretched in the preferred direction of orientation has little capacity for fiber
reorganization in the direction of applied force. The applied load is therefore carried directly by
the strong fibers leading to high strengths and relatively low elongations [30].
What is less known about the mechanical properties of aligned and unaligned PU
scaffolds is the microstructural organization of hard and soft segments within the electrospun
fibers and their contribution to the mechanical properties. GPC and DSC results suggested the
electrospinning process did not affect molecular weight averages or thermal properties of the PU,
so the polymer retained a high molecular weight, phase segregated, and semi-crystalline
structure. These characteristics have a major impact on the mechanical properties of
polyurethane films [13] but it is unclear what their contribution is to the mechanical properties of
the electrospun fibrous scaffolds. High strain rates and polymer fiber stretching experienced
during the electrospinning process can lead to highly oriented molecular chains within the
polymer fibers and a correspondingly stiff polymer behavior [32, 33]. Further mechanical
testing of individual PU fibers and determination of the molecular structure within the fibers is
required to elucidate this microstructure-property relationship. Understanding this behavior may
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be an important consideration for determining if the mechanical properties of these scaffolds are
suitable for cardiac applications.
4.3.2.4 Electrospun PU Scaffolds for Cardiac Tissue Engineering Aligned and unaligned scaffolds were formed from the Phe PU for subsequent cell-based
studies in the development of a cardiac patch. Electrospun scaffolds have previously been
shown to be promising in cardiac tissue engineering applications. Shin et al. [34] electrospun
PCL into thin fibrous scaffolds 10 µm in thickness that had a randomly oriented fiber
organization and fiber diameters ranging from 100 nm to 5 µm with an average diameter of 250
µm. Neonatal cardiac cells were successfully cultured on the scaffolds for 14 days and were
adherent, contracted spontaneously and in synchrony, expressed sarcomeric proteins in a striated
fashion, and had diffuse gap junction formation [34]. Fromstein et al. [7] cultured mESC-
derived cardiomyocytes on unaligned electrospun Phe PU scaffolds that had microfibers ranging
from 2-10 µm. The cardiomyocytes were adherent and contractile on the electrospun meshes
and had a striated sarcomeric phenotype that appeared more mature than cells on TIPS scaffolds
[7]. Recognizing the importance of physical cues for guiding cell organization, Zong et al. [27]
fabricated random electrospun PLA scaffolds and post-processed them by heating and uniaxial
stretching to achieve aligned scaffolds. Cardiac cells on the aligned scaffolds were highly
aligned, developed mature sarcomeric structures and intercalated discs, and were excitable and
contractile. Continuing with the concept of using anisotropic scaffolds to organize cardiac cells,
the aligned and unaligned electrospun Phe PU scaffolds developed and characterized in this
chapter contributed to the work by Rockwood et al. [2]. In this study, we demonstrated that the
physical cues provided by the electrospun PU fibers influenced the organization and phenotypic
expression of neonatal cardiac cells. Specifically, aligned PU scaffolds led to highly oriented
cells and a more mature ventricular phenotype compared to the same cells cultured on unaligned
PU scaffolds or tissue culture polystyrene. These studies clearly indicate the potential of the
aligned and unaligned electrospun PU scaffolds for cardiac tissue engineering and support their
use in subsequent cell-based studies with mESCDCs and MEFs.
Contractile constructs have been successfully formed by culturing cardiac cells on
electrospun scaffolds using a few different polymers, including the rigid PLA. Scaffolds that are
soft, flexible and elastic, however, are required to sustain long-term contractile function [35] and
will be critical for the success of these biomaterials in the heart. The dynamic cardiac
139
environment exposes scaffolds to repetitive stresses and the scaffolds should be able to withstand
these forces without losing mechanical integrity. Moreover, mechanical forces play a critical
role in the formation of normal cardiac structure and function. Tissue engineering scaffolds
should therefore have the appropriate mechanical properties to promote mechanotransduction
between cells and the external cardiac or cardiac-mimicked environment [36]. Fujimoto et al. [5]
provided evidence that biodegradable PU elastomers may be formed into scaffolds with the
appropriate physical, mechanical, and biocompatibility properties for use as a patch in improving
the function of infarcted hearts. Similarities in the mechanical properties between the
biodegradable PU elastomer used in that in vivo study and the Phe PU used in this work suggests
that the electrospun Phe PU scaffolds may also have success in the heart [5, 37, 38].
The exact mechanical properties required for cardiac tissue engineering scaffolds have
not been explicitly defined, but recent work in the field is moving towards mimicking the
mechanical properties of the native myocardium [39-41]. Engler et al. [35] identified that
cardiomyocytes develop force, organize a striated sarcomeric structure, and sustain a contractile
phenotype long-term when cultured on substrates that have an elastic modulus similar to that of
myocardial tissue. Cardiomyocytes on hard matrices do not form striated sarcomeres or remain
contractile for long periods whereas these cells cultured on very soft substrates remain
contractile but do little work [35]. Human myocardium has a Young’s modulus of ~20-500 kPa,
a tensile strength of ~3-115 kPa, and elongations from ~60-90% [41, 42]. The aligned and
unaligned electrospun PU scaffolds formed here have mechanical properties that generally
exceed those of the native myocardium, with the exception of properties of the aligned scaffold
stretched in the cross-preferred direction. The higher ultimate percent elongations (~60-300%)
may be beneficial to PU scaffold performance in the heart, whereas the high ultimate stresses
(~1-4 MPa) and initial modulus (~4-24 MPa) may be more problematic. In particular, scaffold
stiffness is an important determinant of cell organization and function [35] and the higher
stiffness of these scaffolds may limit these properties in engineered tissue. To date, this has not
been observed in in vitro studies with cardiac cells on either the electrospun Phe PU scaffolds or
Phe PU films, which have a higher modulus than the scaffolds [2, 7, 43, 44]. It is important to
point out that the tensile testing of these scaffolds was conducted under ambient conditions and
the mechanical properties may be altered in the physiological environment where water may
disrupt the hydrogen bonding within these materials. In addition, the modulus that cells
140
experience may be quite different than the modulus of overall scaffold where mechanical
properties were a function of numerous individual fiber strengths, inter-fiber interactions, and
fiber alignment events.
The initial modulus and ultimate tensile strength of several synthetic biomaterials that
have been used or hold potential in cardiac tissue engineering are shown in Table 4.4 and are
ranked based on their similarity to the modulus of native myocardial tissue. It has been
suggested that biomaterials stiffer than heart tissue may be used as a heart patch to reduce heart
wall stress and attenuate cardiac deterioration following a myocardial infarction [45]. Diastolic
dysfunction, however, will occur if the biomaterial is too stiff [45]. Until the mechanical
properties of scaffolds for cardiac tissue engineering become strictly defined, comparing the
modulus of biomaterials to native heart tissue is one approach to develop biomimetic matrices
for use in the heart. As seen in Table 4.4, the Phe PU is significantly closer to native
myocardium than some biomaterials, such as the traditional tissue engineering synthetic
polymers PGA and PLA, but is not as close as other materials, such as PGS. The electrospun
Phe PU scaffolds had reduced mechanical properties that were closer to the native myocardium
than the films, but there still may be room for improvement. Since unaligned scaffolds have a
lower modulus than those with fibers aligned in one direction, a bi-layered scaffold with a thin
layer of aligned fibers over unaligned fibers may provide physical cues to cells while decreasing
scaffold stiffness. This may not address the issue of the modulus the cells experience, but may
have implications as a cardiac patch. Alternatively, one of the major advantages of using
polyurethanes in tissue engineering is the ability to tailor the properties for site specific
applications. Therefore, the soft segment of the Phe PU may be changed to lower the modulus
and more closely mimic the mechanical properties of the heart. Synthesizing the Phe PU with a
lower PCL diol molecular weight (e.g. 530 Da) or introducing PEO of different molecular
weights by using tri-block PCL-PEO-PCL soft segments may help to achieve this goal [37, 38].
Further cell-based and in vivo studies with the Phe PU scaffolds are required to test whether
these modifications are necessary.
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Table 4.4: Initial modulus and ultimate tensile strength for films of investigated or potential synthetic biomaterials in cardiac tissue engineering. Polymers are ranked from top to bottom based on similarities in modulus to human
myocardium [37, 38, 45, 46]. Polymer Young’s Modulus Ultimate Tensile Strength
Human Myocardium 0.02-0.5 MPa 3-115 kPa
Poly(glycerol sebacate) 0.04-1.2 MPa 0.2-0.5 MPa
Poly(1,8-octandiol-co-citric acid) 1-16 MPa 6.7 MPa
Poly(1,3-trimethylene carbonate) 5-6 MPa 2-12 MPa
Poly(ether ester urethane)urea 5-75 MPa 8-20 MPa
Phenylalanine-based Polyurethanes 12-30 7-80 MPa
Poly(1,3-trimethylene carbonate-co-
lactic acid) (50:50) 16 MPa 10 MPa
Polycaprolactone 320 MPa 32 MPa
Poly(p-dioxanone) 600 MPa 12 MPa
Polylactic Acid 1-5 GPa 30-80 MPa
Polyglycolic Acid 7-10 GPa 70 MPa
4.4 Conclusions The ability to form porous 3-D scaffolds with the appropriate physical and mechanical
properties is critical to the success of tissue engineering scaffolds. Electrospinning is an
important scaffold fabrication technique that allows the formation of fibrous polyurethane
scaffolds that are promising for soft tissue applications. Phe and Gly-Leu PU microfiber
scaffolds were formed via electrospinning with randomly organized structures and a similar
average fiber diameter. The similarities in structural, physical, and thermal properties warranted
their use in subsequent degradation and cell based studies. Phe PU scaffolds were processed into
different fibrous architectures by electrospinning the polymer at varying mandrel rotational
speeds. The aligned and unaligned PU scaffolds were fabricated for subsequent coculture studies
with mESCDCs and MEFs. The different scaffolds developed and characterized in this chapter
have promising properties for cardiac and other soft tissue engineering applications and warrant
their further investigation in degradation and cell-based studies.
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4.5 References 1. Parrag, I.C. and K.A. Woodhouse, Development of Biodegradable Polyurethane
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Chapter 5: Polyurethane Degradation by Matrix Metalloproteinases
5.0 Abstract The degradation of Phe and Gly-Leu PU scaffolds was investigated to test MMP-
mediated and passive hydrolysis of the PUs. Inactive proMMP-1 was activated by incubation
with (4-aminophenyl)mercuric acetate for 24 h. A fluorogenic substrate assay and zymography
were used to confirm the presence of an active form of both MMP-1 and MMP-9. Incubating the
MMPs with PU scaffolds showed MMP-1 remained stably active in free solution for at least 24 h
whereas MMP-9 activity in free solution dropped to less than 50% in 6 h. Degradation
experiments conducted with the Phe and Gly-Leu PUs in the presence of active MMP-1, active
MMP-9, or a buffer solution were carried out over a 28 day period. Mass loss and structural
assessment suggested that neither PU experienced significant hydrolysis to observe degradation
over the course of the experiment. The Gly-Leu PU was specifically designed to confer
enhanced degradation in the presence of the MMPs. The lack of MMP-mediated Gly-Leu PU
degradation may be due to enzyme concentration, time frame of study, enzyme adsorption onto
PU surfaces, length of peptide sequence in PU, and accessibility of enzymes to labile bonds.
5.1 Introduction Degradation rates and characteristics are an important consideration for the success of
tissue engineering scaffolds. One criterion for biomaterial scaffolds is that the temporary matrix
must degrade at a rate that is appropriate for tissue regeneration. In the case of heart tissue,
degradation should simultaneously occur as ECM proteins are synthesized, allowing the gradual
transfer of mechanical load from the biomaterial to the newly secreted ECM [1, 2]. This enables
the cells and ECM to respond to the dynamic cardiac environment by promoting proper tissue
organization and remodeling that is critical to heart function [3]. Moreover, the scaffolds should
be present long enough to guide integration with the host myocardium but should not last so long
as to interfere with electromechanical coupling of cells [4]. Similarly, degradation should occur
in a time frame that prevents fibrous tissue encapsulation that may inhibit proper integration with
the host.
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Normal ECM degradation and tissue remodeling is carried out by MMPs. MMPs play a
critical role in the wound healing events that occur in the heart following a MI and continue to
participate in ECM degradation during detrimental ventricular remodeling in the progression to
heart failure [5-7]. Specifically, MMP-1, MMP-2, and MMP-9 have been shown to have major
roles in the heart after injury. Disrupting the normal expression of these proteases attenuates
ventricular remodeling and improves cardiac performance following an infarction [8-10]. A
priori knowledge of the critical MMP involvement in an injured heart and their sites of cleavage
in ECM proteins allow the design of biomimetic synthetic polymers that may exploit the
presence of these cell-secreted proteases to degrade the material. An MMP-sensitive synthetic
biomaterial may degrade at a rate that is more appropriate for promoting ECM production, tissue
remodeling, and vascularization as required by the host tissue than other synthetic polymers that
degrade predominantly by passive hydrolysis. This may be one method of establishing
degradation rates that are appropriate for cardiac tissue engineering and would be an appealing
property of candidate biomaterials for use in the heart.
The cleavage site of several MMPs was built into the backbone structure of PUs through
the formation of a Gly-Leu-based diester chain extender. The Gly-Leu PU and chymotrypsin-
sensitive Phe PU were synthesized, characterized, and formed into electrospun scaffolds as
discussed in previous chapters. In this chapter, the degradation of the PU scaffolds by passive
hydrolysis and enzymatic mechanisms using the proteases MMP-1 and MMP-9 is described.
The degradation experiments were conducted to provide a better understanding of how altering
PU hard segment chemistry and incorporating a Gly-Leu dipeptide into the chain extender
affected PU properties. The ability to obtain activated MMPs was initially confirmed and the
activity of these enzymes was tested. Subsequently, a PU scaffold degradation study was carried
out and changes in mass and physical structure were assessed.
5.2 Materials and Methods All materials were purchased from Sigma-Aldrich Canada (Oakville, ON, Canada) unless
otherwise stated.
5.2.1 Activation and Activity of MMPs The MMPs used in this study were secreted as pro-enzymes requiring activation prior to
obtaining protease activity. Therefore, initial studies were conducted to explore the activation of
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the MMPs. The organomercurial agent (4-aminophenyl)mercuric acetate (APMA) was used in
the activation studies. ProMMP-1 (Calbiochem, San Diego, CA) or proMMP-9 at a
concentration of 40 μg/ml was added to a mixture containing 1 mM of APMA in a buffer
solution (5 mM Tris, 10 mM CaCl2, 150 mM NaCl, 0.02% (w/v) NaN3, pH 7.5). Activating
enzyme solutions were incubated at 37˚C and aliquots were taken at 15 min, 30 min, 1 hr, 2 hr, 4
hr, and 24 hr.
To quantify protease activity and identify MMP activation, a fluorogenic substrate assay
was employed. The fluorogenic substrate FS-6 (Mca-Lys-Pro-Leu-Gly-Leu-Dpa-Ala-Arg-NH2
where Mca is (7-methoxycoumarin-4-yl)acetyl and Dnp is 2,4-dinitrophenyl; Calbiochem, San
Diego, CA) is a peptide-based substrate that initially has the fluorescent group Mca quenched by
Dnp due to proximity within the peptide chain. FS-6 has a high substrate specificity for all
MMPs and upon cleavage of the Gly-Leu peptide bond, Mca becomes free of Dnp and
fluorescence can be measured to quantify MMP activity [11]. Aliquots taken from the activation
solution at different time points and MMP-9 purchased in its active form (67 kDa active product;
Calbiochem, San Diego, CA) were diluted with buffer solution to a final enzyme concentration
of 750 ng/ml. Enzyme solutions were incubated with FS-6 at a final FS-6 concentration of 5
μM. Fluorescence was measured with a microplate reader at an excitation/emission wavelength
of 324/400 nm after incubating the activated enzyme solution with FS-6 for 1 hr. Fluorescence
values were normalized to buffer solution alone with FS-6.
In addition to measuring activation by a fluorogenic substrate assay, zymography was
used to confirm the presence of active MMPs. A protocol described by Hawkes et al. [12] was
used with a few modifications. A 10% zymogram gel was made by mixing stock solutions of
30% acrylamide/bisacrylamide, 1.5 M Tris-HCl (pH 8.8), 10% w/v gelatin, 10% w/v sodium
dodecyl sulfate (SDS), ultrapure deionized H2O, 10% w/v ammonium persulfate, and N,N,
N’,N’-tetramethylethylenediamine. Samples were prepared by mixing at a 1:1 ratio with
Laemmeli sample buffer (0.125 M Tris-HCl of pH 6.8, 4% w/v SDS, 20% v/v glycerol, and
0.04% w/v bromophenol blue) at final dilution of 240 ng per 10 μl solution with the exception of
active MMP-9 (40 ng per 10 μl solution). Electrophoresis was carried out at 180 V for
approximately 45 min. The gel was incubated with zymogram renaturing buffer (Invitrogen,
Carlsbad, CA) for 30 min, then with zymogram developing buffer (Invitrogen) for ~4 h, and
finally with SimplyBlue SafeStain (Invitrogen) for ~1 h all at room temperature with ultrapure
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deionized water rinse steps in between. A standards lane of pre-stained calibrated molecular
weight proteins (Bio-Rad Laboratories, Mississauga, ON) was used as a reference to help
identify the molecular weights of the MMP species.
An activity study was conducted to identify the duration the MMPs remain active in free
solution. Active MMP-9 and proMMP-1 activated by APMA for 24 h were diluted in buffer
(200 ng/ml) and were incubated with PU scaffolds for 6, 12, and 24 h. Enzyme solution was
removed at each time point and protease activity was measured using the fluorogenic substrate
assay. Percent activity was calculated by comparing to initial fluorescent measurements prior to
incubation with scaffolds.
5.2.2 Degradation of Polyurethanes by MMPs Phe and Gly-Leu PU scaffold degradation was carried out in buffer, MMP-1, or MMP-9
solutions to investigate the passive and enzyme-mediated hydrolysis of the scaffolds in vitro.
Electrospun PU scaffolds approximately 80-100 µm thick were cut into discs ~1.5 cm in
diameter. Scaffolds were weighed and were incubated in buffer or active MMP solutions (200
ng/ml). Solutions were changed every 24 h. After 7, 14, 21, or 28 days in the different
solutions, samples were rinsed three times in triton X-100 solution (1% v/v in water) and three
times in ultrapure deionized water. Scaffolds were padded dry with a Kimwipe, placed in a
vacuum chamber at room temperature for 48 hrs, and reweighed. Differences between initial and
final mass were used to calculate percent mass remaining at various time points. Changes in
scaffold structure and surface features were determined by SEM as described in chapter 4.
Preliminary water uptake measurements and an enzyme inhibition assay were performed
to better understand the mechanisms that may influence PU degradation. PU scaffolds were
weighed and then incubated in ultrapure deionized water for 4 days to promote water absorption
into the matrices. Samples were padded dry with a Kimwipe and were reweighed. The
difference in mass before and after incubation in water was used to calculate percent water
uptake. Enzyme inhibition was performed by measuring MMP activity using the fluorogenic
substrate assay in the presence of various concentrations of the Gly-Leu dipeptide. Fluorescence
levels in the presence of Gly-Leu dipeptide were compared to those without the dipeptide to test
if enzyme binding to the Gly-Leu dipeptide inhibited cleavage of the fluorogenic peptide
substrate.
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5.3 Results and Discussion
5.3.1 Activation of MMPs The degradation characteristics of amino acid and dipeptide-containing PU scaffolds were
investigated to understand the effects of incorporating the Gly-Leu dipeptide into the backbone
polymer structure. MMPs are secreted as latent pro-enzymes requiring activation before being
able to cleave ECM proteins, peptides, or biomimetic polymers. MMP-9 (gelatinase B, 92 kDa
gelatinase) is secreted as a 92 kDa protein while MMP-1 (interstitial collagenase, collagenase-1)
is secreted as a major 52 kDa pro-enzyme and a minor 57 kDa glycosylated species, all requiring
cleavage of peptides > 10 kDa in size for activation [13-15]. MMP-1 and MMP-9 can be
activated in vitro by proteases, such as trypsin and plasmin, or organomercurial agents such as
APMA [13, 14, 16]. Incubation of MMP-1 with APMA results in initial collagenolytic activity
prior to a drop in molecular weight, followed by the rapid conversion to a 44 kDa intermediate
and eventually to a more stable 42 kDa active enzyme [16]. APMA-induced activation of MMP-
1 occurs through a cysteine switch mechanism when APMA interacts with a cysteine residue in
the pro-peptide domain of the MMP [16, 17]. This interaction interferes with the zinc-binding
coordination bond leading to conformational changes that disrupts the pro-domain and renders
the active site accessible for proteolytic activity. The free active site and corresponding protease
activity leads to cleavage of the pro-peptide domain autocatalytically and a reduced molecular
weight active protein [16, 17]. MMP-9 has also been shown to be activated by incubation with
APMA through a similar cysteine switch mechanism [14, 17]. The 92 kDa pro-enzyme is
converted to a 83 kDa form exhibiting proteolytic activity through cleavage of the pro-peptide
domain at the N-terminus of the protein [14, 17]. Full activation of MMP-9 is achieved through
longer incubation periods (~24 hr) with APMA leading to the formation of a 67 kDa active
species by cleaving a peptide at the C-terminus of the protein chain [14, 17]. Several other
active forms of both MMP-1 and MMP-9 have been reported in the literature and are a function
of the agents used in activation [18].
Human fibroblast-derived MMPs were supplied in solution containing a mix of the pro-
enzyme and its activated equivalent. Degradation experiments required catalytic activity of the
MMPs and therefore conversion of the inactive zymogens to their active form was explored. The
proMMPs were incubated with APMA for different periods and proteolytic activity was
monitored using a fluorogenic substrate assay. As indicated by the results in Figure 5.1, MMP-1
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demonstrated protease activity after incubation with APMA at all time points while MMP-9
showed little protease activity after incubation with APMA. MMP-1 displayed a high level of
fluorescence, 5337 ± 451 units, after only 15 min of incubation with APMA. This fluorescence
decreased slightly at 30 min and slowly increased over the next few time points with the highest
level observed at 24 hr, 5619 ± 413 units. In contrast, low levels of fluorescence were observed
at all time points with MMP-9 incubated with APMA. Given the proven ability of APMA to
activate proMMP-9 [14, 17, 18], the problem with activating MMP-9 was not likely due to the
method of activation but rather reflected a problem with the MMP-9 enzyme. As a result, MMP-
9 was supplied in its active 67 kDa form (Calbiochem) and this enzyme showed high protease
activity after incubation with the fluorogenic substrate (Figure 5.1). Therefore, active forms of
both MMP-1 and MMP-9 were obtained that exhibited protease activity for use in subsequent
studies.
Figure 5.1: Activation of MMPs using APMA. Fluorescence was normalized to buffer background with FS-6.
MMP-1 showed protease activity at all time points after incubation with APMA while MMP-9 showed little activity. Active MMP-9 (Calbiochem) demonstrated protease activity. Error bars represent ± standard deviation. n=3, N=3.
The results from the fluorogenic substrate assay were supported by zymography of the
different MMP solutions. Zymography involves the electrophoretic separation of proteins
through a polyacrylamide gel containing gelatin under denaturing and non-reducing conditions.
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Following separation, the resolved proteins are renatured and proteolytic activity is detected
through degradation of gelatin. A 10% polyacrylamide gel was used and the gel was stained
using a Coomassie blue staining solution after the enzymes were renatured. The resulting gel is
shown in Figure 5.2 with clear bands indicating proteolytic activity against the non-degraded
blue gelatin background. Several high molecular weight bands were observed with the
proMMP-9 and MMP-9 solution incubated with APMA indicating a mixed protein solution.
MMP-9 forms complexes with several proteins after secretion from cells and these bands may
reflect a complexed form of MMP-9 [17]. No difference in band patterns was observed between
the proMMP-9 and APMA-incubated MMP-9 solution suggesting the absence of an active form
of this enzyme. In contrast, MMP-9 supplied in an active form exhibited additional bands
not present in the other MMP-9 lanes. Gelatin degradation was observed at all points within the
Figure 5.2: Zymogram of MMP activation solutions. No differences in proMMP-9 and MMP-9 incubated with
APMA were observed. Active MMP-9 resulted in two additional lower molecular weight bands, one corresponding to the 67 kDa active species. Incubation of MMP-1 with APMA led to a loss of the 52 kDa pro-enzyme band with
the 44 kDa intermediate and 42 kDa active species observed in both lanes.
153
lane above a certain cutoff indicating either a mixed protein solution similar to proMMP-9 or the
presence of active intermediate products that were not completely denatured during
electrophoresis. In addition, two lower molecular weight bands with distinct gelatinase activity
were observed. The enzyme was supplied in the 67 kDa form and the bands likely corresponded
to this active species and a further processed form. An active 55 kDa MMP-9 species has been
reported [18] and could account for this lower molecular weight species. MMP-1 was supplied
as a mixture of both the inactive zymogen and activated species. As a result, the proMMP-1 lane
displayed three bands corresponding to the 52 kDa pro-enzyme, the 44 kDa intermediate species,
and the 42 kDa active form. After activation by APMA for 24 h, the 52 kDa pro-enzyme was no
longer present while the 44 kDa intermediate and 42 kDa active species remained. This is
consistent with the activation mechanism of APMA reported in the literature [16-18]. These
results support the fluorogenic substrate assay for MMP activation and subsequent experiments
were performed using MMP-1 activated by APMA for 24 h and the active MMP-9 directly.
5.3.2 Activity of MMPs after Incubation with Polyurethanes After successfully obtaining active forms of MMP-1 and MMP-9, an activity study was
performed to test the stability of these proteolytic enzymes and determine when solutions should
be changed for the degradation experiment. This activity study was carried out using a similar
procedure to the one described by Skarja and Woodhouse [19] for trypsin and chymotrypsin.
Active enzymes were incubated with and without the PU scaffolds and aliquots were taken at
various time points. Results from the fluorogenic substrate assay with the two enzymes are
shown in Figure 5.3. Aliquots taken from MMP-1 incubated with either PU scaffold showed
high levels of protease activity for at least 24 h. In contrast, MMP-9 activity in solution dropped
to less than 50% its initial level after 6 h for the two PUs. As a control, the enzyme solutions
were incubated in the multi-well plates without a PU scaffold and showed a similar trend; MMP-
1 in solution remained active for at least 24 h while MMP-9 in solution dropped below 50% its
initial activity in 6 h. The exact mechanisms behind the observed MMP-1 and MMP-9 activity is
not clear and may be due to differences in enzyme stability, adsorption onto the PU scaffolds and
multi-well plates, and/or intermolecular cleavage. Further studies investigating the adsorption of
154
Figure 5.3: Activity of MMPs after incubation with PU scaffolds. a) MMP-1 activity remains at initial level for at
least 24 h. b) MMP-9 activity drops to less than 50% its initial level after 6 h. Error bars represent ± standard deviation (n=3, N=3).
155
the enzymes onto the PU scaffolds are required to help elucidate these mechanisms. For the
degradation study, fresh enzyme solutions were added to PU scaffolds each day to replenish any
lost activity.
5.3.3 Degradation of Polyurethanes by MMPs In vitro degradation experiments were conducted to help identify if specific and enhanced
degradation by MMP-1 and MMP-9 is achieved by incorporating the Gly-Leu dipeptide into the
polyurethane structure. Active enzymes were incubated with the PU scaffolds and mass loss was
assessed after 7, 14, 21, and 28 days in the different solutions. Figure 5.4 shows the mass of PU
scaffolds remaining at different time points in the degradation study. The Phe PU scaffolds
exhibited no changes in mass after incubation with the MMP-1, MMP-9, or buffer solutions over
the 28 day period. Similarly, the Gly-Leu PU scaffolds also had no changes in mass after
incubation with any of the solutions during the degradation study. These results suggested that
neither PU type was susceptible to either MMP-mediated degradation or passive hydrolysis over
the time period investigated. SEM images of the PU scaffolds after 28 days in the different
solutions are shown in Figure 5.5. No major change in surface or scaffold morphology for either
PU type was observed confirming the mass loss data. There was no evidence of broken fibers or
rough fiber surfaces that would indicate susceptibility to the different solutions. The slightly
fused fiber morphology that makes the Phe PU scaffolds look somewhat different from one
another is due to slight variations within the starting PU scaffold structure and was not the result
of degradation.
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Figure 5.4: Mass remaining of PU scaffolds over 28 day degradation study. a) Gly-Leu PU scaffolds and b) Phe PU
scaffolds after incubation with MMP-1, MMP-9, and buffer solutions over a 28 day period. No appreciable difference in mass was observed over the 28 day period. Error bars represent ± standard deviation (n=3, N=1).
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Figure 5.5: SEM images of PU scaffolds after 28 day incubation period in various solutions. A) Gly-Leu PU in
buffer. B) Gly-Leu PU in MMP-1. C) Gly-Leu PU in MMP-9. D) Phe PU in buffer. E) Phe PU in MMP-1. F) Phe PU in MMP-9. No changes in surface or scaffold morphology were observed.
The results from the initial degradation study of the PU scaffolds in MMP and buffer
solutions suggested that little or no observable degradation occurred over the 28 day time period.
The Phe PU has been shown to be relatively resistant to passive hydrolysis while being more
susceptible to enzyme-mediated degradation by chymotrypsin and to a lesser extent by trypsin
[19]. Phe PU films exhibited ~1%, 3.5%, and 4.5% mass loss after 28 days in the presence of
buffer, trypsin, and chymotrypsin solutions respectively and enzymatic degradation occurred
through a surface-mediated mechanism [19]. Electrospun Phe PU scaffolds were shown to have
higher mass loss (~10%) over 28 days in chymotrypsin due to increased surface area [20].
Increased fiber surface roughness and breaking of PU fibers were observed with increasing
chymotrypsin exposure and is consistent with the surface-mediated degradation mechanism [20].
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In the current work, the Phe and Gly-Leu PU scaffolds did not exhibit changes in mass or surface
features that would be expected if the scaffolds were susceptible to passive hydrolysis or
enzymatic degradation. Although previous work with PCL-based PU films suggested little
passive hydrolysis of the PU scaffolds would occur, the increased surface area of the electrospun
mats may be anticipated to cause a higher mass loss in the buffer solution (>1%) compared to
films [19, 20]. Nonetheless, Bolgen et al. [21] found that electrospun scaffolds made from pure
PCL were more hydrophobic than PCL films, thereby reducing the degradation rates. A similar
phenomenon may be observed here with the PCL-based PUs. It was not anticipated that the Phe
PU would be susceptible to degradation by the MMPs. The lack of visible degradation with this
polymer provides further support that the Phe-based chain extender confers specific and
enhanced degradation in the presence of certain enzymes.
More importantly though is the lack of observable MMP-mediated degradation with the
Gly-Leu PU. The Gly-Leu dipeptide was incorporated into the chain extender as a means of
introducing the MMP cleavage site into the polymer backbone. Surface enrichment of hard
segments as is commonly observed with PUs was anticipated to present the labile bonds to the
MMPs and lead to PU degradation. Unfortunately, in the degradation study conducted here, the
Gly-Leu PU did not exhibit enhanced degradation in the presence of MMP-1 or MMP-9.
There are several potential explanations for the absence of any enzyme-mediated Gly-
Leu PU degradation observed in this study. One may be related to the concentration and time
period investigated. The degradation experiment was conducted by incubating the PU scaffolds
with active MMPs at a concentration of 200 ng/ml, which is similar to the concentration found in
the plasma of patients with congestive heart failure [22]. This concentration was chosen in an
attempt to carry the degradation study out under physiologically-relevant conditions, but MMP
concentrations found in the plasma may not necessarily correlate to the actual enzyme level
scaffolds will be exposed to when implanted in the heart. MMP expression following a MI may
become elevated in the presence of biomaterial scaffolds due to the body’s inflammatory
response and the increased number of MMP-secreting cells recruited to implant site [6, 23].
Moreover, dose-dependent PU degradation has been demonstrated with polycarbonate-based
PUs using cholesterol esterase whereby some PUs were stable at low enzyme concentrations but
exhibited significant degradation at higher concentrations [24]. The MMP concentration used in
this study may have been too low to observe PU degradation in the time period investigated.
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Increasing the enzyme levels and length of the study may lead to substantial and observable PU
degradation.
Enzyme adsorption and activity may also have been a factor in explaining the absence of
PU degradation. Santerre et al. [25] provided a mechanistic model for degradation of PUs by
hydrolytic enzymes: 1) PU surface contact with aqueous environment and structural
rearrangement of polymer chains; 2) enzyme adsorption/desorption with the surface; 3) active
adsorbed enzymes react with susceptible bonds; 4) degradation products are released from main
polymer chain into solution exposing new polymer surfaces; 5) water, electrolytes, and enzymes
establish interactions with newly exposed surface; and 6) process is propagated over time to
degrade the polymer. Therefore, enzymes must adsorb onto the PUs and must remain active in
the surface-bound state for polymer chain cleavage to occur. Enzyme adsorption onto the PU
scaffolds was difficult to determine from the activity study conducted here and may have been a
limiting step in the degradation study.
A third explanation for the lack of Gly-Leu PU degradation could be that introducing the
Gly-Leu dipeptide alone without any flanking amino acid sequences into the PU structure was
not sufficient for recognition by the MMPs for binding and substrate cleavage. It was
hypothesized that this minimalistic approach of incorporating the Gly-Leu cleavage site may
enhance degradation rates in PUs, but it was not known conclusively whether the dipeptide alone
would confer enzyme-susceptibility to the biomaterial. While most enzymes have very specific
substrates in which they bind to and act upon, all the MMPs have broad substrate specificities
and are able to cleave several ECM and non-ECM proteins [17, 26]. An early MMP-1 study
suggested that specificity with this enzyme was largely independent of substrate conformation
and reflected the amino acids surrounding the cleavage site [27].
Much work was subsequently conducted on short peptide sequences to identify substrate
specificities of the MMPs. Octapeptides are typically used in these studies, as this is
approximately the peptide length that can fit into the active site of the enzymes [26]. One group
systematically altered the peptide sequence Gly-Pro-Gln-Gly-Ile-Ala-Gly-Gln, known to be
cleaved by several MMPs between the Gly and Ile residues, to investigate the effects amino acid
residues surrounding the cleavage site had on substrate specificity [17]. Hydrolysis rates were
dependent on both the specific amino acid and position within the peptide sequence for different
MMPs [17]. In addition, successively removing amino acids from the C and N-terminus of this
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reference peptide led to rapid decreases in hydrolysis rates once the peptide was smaller than six
amino acids long [17]. While this would suggest that a hexapeptide is the minimum length for
high rates of hydrolysis, it does not necessarily indicate that peptide cleavage by MMPs will not
occur with shorter peptide lengths.
Hubbell and West [28] incorporated the peptide Ala-Pro-Gly-Leu into synthetic
hydrogels that rendered the biomaterial susceptible to collagenase-mediated degradation.
Hydrogel discs 5 mm in diameter and 2 mm thick were completely degraded after 5 and 7 days
of incubation with 2 mg/ml and 0.2 mg/ml collagenase solutions, respectively. Recently, Guan
et al. [29] used the same sequence with a lysine residue on the end (Ala-Pro-Gly-Leu-Lys) as a
diamine chain extender in the development of segmented polyurethanes that were also
susceptible to collagenase. Although the details of this work have not been fully published, PU
films showed approximately 20% and 30% mass loss after 4 and 8 weeks in the presence of
collagenase, respectively. The exact peptide requirements for achieving MMP susceptibility
with biomaterials are not known. However, the literature would suggest that peptide length and
sequence will affect the rates of hydrolysis. It is possible that additional amino acids flanking
the Gly-Leu cleavage site are required for degradation by the MMPs. The lack of observed
degradation may also be due to low rates of hydrolysis with the dipeptide alone. Higher enzyme
concentrations as well as longer time periods may be necessary to detect any Gly-Leu PU
degradation by the MMPs.
In an attempt to further elucidate the mechanisms behind PU degradation, a preliminary
competitive substrate assay was carried out. This was conducted to test if the Gly-Leu dipeptide
alone is a substrate for MMPs and therefore if it can act as a site for potential cleavage by the
proteases. If the Gly-Leu dipeptide is a binding substrate for the MMPs, then introducing it into
the fluorogenic substrate assay may lead to a competitive reaction and lower fluorescence
(Figure 5.6). Therefore, active MMPs were incubated with FS-6 and different concentrations of
the Gly-Leu dipeptide were introduced into the solution. Fluorescence levels were compared to
those with the MMPs incubated with FS-6 alone to get a relative percent of FS-6 cleavage
(Figure 5.7). No change in FS-6 cleavage was observed when the Gly-Leu dipeptide was
introduced at 100x the molar concentration of FS-6 but a drop of approximately 10% occurred at
1000x molar concentration. Although repeated trials are necessary to confirm these results, this
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Figure 5.6: Reaction scheme for enzyme activity assay and competitive substrate enzyme activity assay.
Figure 5.7: Inhibition of FS-6 cleavage using the Gly-Leu dipeptide. No difference in FS-6 cleavage was observed with 100x molar concentration of dipeptide but a drop of ~10% occurred at 1000x. Error bars represent ± standard
deviation. n=3, N=1.
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preliminary study suggested that FS-6 is a much better substrate than the Gly-Leu dipeptide. The
unchanged FS-6 cleavage rates in the presence of 100x excess of Gly-Leu dipeptide may reflect
a much higher binding efficiency to the length and sequence of the FS-6 peptide substrate
compared to the dipeptide alone. This would be consistent with the trend of higher rates of
hydrolysis with peptides longer than 6 amino acids in length and dramatically reduced rates with
shorter peptides [17]. The drop in FS-6 cleavage at 1000x molar excess of dipeptide, however,
may suggest that while having much lower binding kinetics than FS-6, the Gly-Leu dipeptide
may still act as a substrate for the MMPs. This would support the notion that the Gly-Leu
dipeptide is susceptible to MMPs, but hydrolysis rates were too slow to be observed in the study
conducted here.
The exact interpretation of the competitive substrate enzyme activity results, however,
remains difficult. This competitive substrate enzyme inhibition study may be better validated by
comparing the peptide Ala-Pro-Gly-Leu, which is susceptible to collagenase within PUs and
other synthetic polymers [28, 29]. Moreover, it is important to consider the length of the
dipeptide as a substrate as this may influence MMP binding to the molecule. Inhibition of
MMPs has been observed with a short sequence consisting of Pro-Leu-Gly-NHOH [18], but it is
unclear if a dipeptide alone is long enough to see similar behavior. It has been suggested that the
active site of MMPs can accommodate peptide sequences up to 8 amino acids in length [26].
Shorter peptide sequences may lead to reduced rates of hydrolysis because these peptides are not
stabilized in the active site of the enzymes as well as the longer sequences. When the Gly-Leu
dipeptide is contained within a polymer chain, the MMPs may have a larger substrate to bind to
allowing a more stabilized complex to form between the enzyme and potential cleavage site. A
better model for this system may be to replace the Gly-Leu dipeptide with the Gly-Leu-based
chain extender, thus providing a longer chain for the MMPs to bind to. Similarly, using the Gly-
Leu-based chain extender as a linker to make a short polymer chain may provide a model
polymer segment to test both competitive reaction enzyme inhibition and cleavage of polymer
chains at the Gly-Leu peptide bond. Guan and Wagner [30] verified the sensitivity of the Ala-
Ala-Lys chain extender to elastase by synthesizing a methyl-PEG flanked peptide conjugate
(mPEG-Ala-Ala-Lys-mPEG) and identifying the molecular weight of products formed after
exposure to the enzyme. This may better indicate if the Gly-Leu dipeptide is sufficient to get
chain cleavage by the proteases or if additional amino acids are required.
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A fourth possible explanation for the Gly-Leu PU scaffold stability may be due to
polymer hydrophobicity and hard segment packing within the PU. Enzyme-mediated PU
degradation is surface limited [19] and polymer hydrophobicity and hard segment packing may
limit accessibility of the MMPs to potentially susceptible cleavage sites. Work by Santerre and
colleagues found that increased hard segment content was indirectly proportional to hydrolytic
degradation of polyurethanes even though the hard segments contained hydrolysable groups [31].
In addition, the level of enzymatic hydrolysis was inversely proportional to hydrogen bonding
between hydrolysable bonds in the hard and soft segments due to domain cohesion and the
inability of enzymes to reach the labile bonds [32]. A similar phenomenon may account for the
stability of the Gly-Leu PU to enzyme-mediated hydrolysis. Characterization of the Gly-Leu PU
indicated a phase segregated polymer with enhanced mechanical properties compared to Phe PU
and increased hydrogen bonding within the hard segment. These results suggested the Gly-Leu
PU had some hard segment interactions but the microphase was amorphous with no observed
hard segment thermal transitions due to the bulky side chains within the LDI and chain extender.
Hard segment surface enrichment can also occur with segmented PUs in aqueous environments,
suggesting that the labile peptide bonds may have been at the surface where enzymes would have
access. Water uptake by the PU scaffolds (Figure 5.8) revealed that both PU scaffold types
absorbed some water but the Gly-Leu PU scaffolds absorbed more water than the Phe PU
scaffolds. This trend suggests that the Gly-Leu PU was not as hydrophobic as the Phe PU and
that MMPs may indeed have had access to polymer fiber surfaces.
Interestingly, cleavage of polymer chains by the MMPs requires that the susceptible
bonds not only be accessible but also presented in a conformation that is suitable for binding to
the active site of the enzymes [17]. Santerre et al. [25] demonstrated that similar levels of
adsorbed cholesterol esterase to different PU surfaces did not necessarily correlate to the same
amount of polymer degradation. If the Gly-Leu dipeptide alone is sufficient to be cleaved by
MMPs at a rate that would have been observed in this study, then polymer chain organization
and hard segment interactions may have played a role in preventing hydrolysis due to the
conformation of the labile bonds.
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Figure 5.8: Water uptake by Phe and Gly-Leu PU scaffolds. A trend of higher water uptake was observed with the Gly-Leu PU compared to the Phe PU suggesting the Gly-Leu PU was not as hydrophobic as the Phe PU. n=9, N=1.
The cleavage site of MMPs was incorporated into the PUs as a means of developing a
MMP-sensitive PU. It remains unclear what contribution the Gly-Leu-based chain extender had
on polymer degradation. Subtle differences in mechanical properties and polymer processing
were previously identified between the Phe and Gly-Leu PUs, but the majority of the polymer
properties were very comparable. It was anticipated that one of the main differences in polymer
properties between the two PUs would come from their degradation characteristics. This was not
observed in the degradation study conducted here. Little or no evidence of degradation for either
PU scaffold occurred over the 4 week time period from passive hydrolysis or MMP-mediated
cleavage. This would suggest that introducing the Gly-Leu dipeptide into the backbone structure
of the PUs did not confer specific and enhanced degradation to the PU scaffolds.
To better understand the degradation properties of the PU scaffolds, in vitro and in vivo
cell-based degradation studies are required. The PU degradation system using C14-labeled hard
segments extensively utilized by Santerre and colleagues [33-35] may be a useful model for
elucidating the mechanisms of Phe and Gly-Leu PU degradation by inflammatory cells,
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including PMNs and MDMs, or other MMP-secreting cells found in the heart such as cardiac
fibroblasts. MMP-mediated PU degradation in this in vitro model could be determined by
measuring MMP activity and corresponding PU degradation and then inhibiting MMP activity
and again quantifying PU degradation. In vivo degradation testing should also be performed in
animal models to determine the residence times for the Phe and Gly-Leu PU scaffolds in the
cardiac environment. The multifaceted mechanical and chemical mechanisms that work
synergistically in degrading PUs in vivo make it difficult to fully appreciate the performance of
the PU scaffolds by in vitro studies alone and are required if these materials will ever move to
the clinical setting. Factors such as the aqueous tissue environment, hydrolytic enzymes,
oxidative agents, calcification, mechanical stresses, environmental stress cracking, metallic ions,
salts, acids, and cells will all contribute to the environmental biodegradation of PU scaffolds
[36]. Guan et al. [37] demonstrated an elastase-sensitive PU scaffold exhibited ~42% mass loss
over an 8 week in vitro enzyme-based degradation study but was completely degraded after the
same time period when implanted subcutaneously in rats. The significant difference in
degradation rates between the in vitro and in vivo studies highlights the multifaceted mechanisms
of environmental biodegradation that may be difficult to predict using enzyme-based in vitro
studies alone. The work presented here further identifies the need to move to in vitro and in vivo
cell-based studies to obtain a better appreciation of the degradation properties of the Phe and
Gly-Leu PU scaffolds.
If after additional cell-based testing the degradation rates for the Phe and Gly-Leu PU
scaffolds are not appropriate for use in the heart, several modifications can be made to the PU
chemistry to alter these properties. Previous work has demonstrated that the degradation rates of
the Phe-containing family of PUs by both passive hydrolysis and enzymatic means may be
significantly increased by lowering the PCL soft segment molecular weight or by using PEO of
various molecular weights [19]. A similar approach may be used to increase degradation rates
with the Gly-Leu PU. However, this may not address MMP-mediated degradation of the
scaffolds. If the Gly-Leu dipeptide alone is not sufficient to confer MMP susceptibility to the
PUs, then numerous other peptide sequences may be used to tune the degradation properties of
the PUs [17, 26]. Altering the length and peptide sequence can dramatically alter rates of peptide
hydrolysis and could lead to synthetic polymers with a large repertoire of degradation
characteristics for specific applications [17, 28]. The method of incorporating larger peptides
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into the PUs will be important because of the higher molecular weights associated with the
additional amino acid residues. If the diol linker method used in developing the Phe and Gly-
Leu-based diester chain extenders were used with larger peptide sequences, the ratio of soft
segment to hard segment would consequently decrease. As a result, several of the favorable
properties conferred to the PUs by soft segment choice may be lost due to the large hard segment
domains. The method used by Wagner’s group of adding a lysine residue at the C-terminal of a
desired peptide sequence may work better with larger peptide sequences without allowing the
chain extender molecular weight to get too big [29, 30]. Alternatively, it may be interesting to
characterize PU properties of hexa or octapeptide-based diester chain extender. Further testing
of the Phe and Gly-Leu PU scaffolds will help identify whether modifications to the chemistry
are necessary to achieve appropriate degradation properties for use in the heart.
5.4 Conclusions The degradation characteristics of PU scaffolds are critical to their success in tissue
engineering applications. MMP activation and activity studies were conducted to confirm active
enzymes could be obtained for degradation trials. A subsequent MMP and buffer-based
degradation experiment showed little passive hydrolysis or enzyme-mediated chain cleavage.
These results suggested that incorporating the Gly-Leu dipeptide into the backbone structure was
not sufficient to see enhanced degradation in the presence of MMP-1 and MMP-9. Several
factors may have played a role in the stability of the Gly-Leu PU including enzyme
concentration, time period investigated, MMP adsorption onto the PUs, length of peptide
sequence incorporated in the chain extender, and accessibility of enzymes to potentially labile
bonds. It remains unclear what contribution the Gly-Leu dipeptide has on the degradation rates
and characteristics of the PU scaffolds. Future work using in vitro and in vivo cell-based studies
will help to better characterize the degradation behavior of the PU scaffolds.
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27. Gross, J., E. Harper, E.D. Harris, P.A. McCroskery, J.H. Highberger, C. Corbett, and A.H. Kang, Animal collagenases: specificity of action, and structures of the substrate cleavage site. Biochem Biophys Res Commun, 1974. 61(2): p. 605-12.
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28. West, J.L. and J.A. Hubbell, Polymeric biomaterials with degradation sites for proteases involved in cell migration. Macromolecules, 1999. 32(1): p. 241-244.
29. Guan, J.J. and W.R. Wagner. Development of collagenase and plasmin sensitive elastomeric scaffolds for soft tissue engineering. in 8th World Biomaterials Congress. 2008. Amsterdam.
30. Guan, J.J. and W.R. Wagner, Synthesis, characterization and cytocompatibility of polyurethaneurea elastomers with designed elastase sensitivity. Biomacromolecules, 2005. 6(5): p. 2833-2842.
31. Santerre, J.P. and R.S. Labow, The effect of hard segment size on the hydrolytic stability of polyether-urea-urethanes when exposed to cholesterol esterase. J Biomed Mater Res, 1997. 36(2): p. 223-32.
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33. Labow, R.S., D.J. Erfle, and J.P. Santerre, Neutrophil-mediated degradation of segmented polyurethanes. Biomaterials, 1995. 16(1): p. 51-9.
34. Labow, R.S., E. Meek, and J.P. Santerre, Model systems to assess the destructive potential of human neutrophils and monocyte-derived macrophages during the acute and chronic phases of inflammation. Journal Of Biomedical Materials Research, 2001. 54(2): p. 189-197.
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Chapter 6: Cell Response to Electrospun Polyurethane Scaffolds § Sections of this chapter have been accepted for publication [1]: Parrag IC, and KA
Woodhouse. Development of Biodegradable Polyurethane Scaffolds using Amino Acid and
Dipeptide-based Chain Extenders for Soft Tissue Engineering. Journal of Biomaterials Science –
Polymer Edition, in press.
6.0 Abstract Cell-based studies were conducted to investigate the cellular response to PU scaffolds
and identify the potential of these biomaterials in soft tissue engineering applications. The Phe
and Gly-Leu PU scaffolds were seeded with a high density of mouse embryonic fibroblasts
(MEFs). AlamarBlue analysis, Live/Dead staining, and immunostaining of the cell-seeded
constructs indicated that both PUs could support a high density of viable cells out to at least 28
days. Cells were adherent and spread out with no regular organization on the randomly oriented
substrates. For cardiac applications, the Phe PU was electrospun into scaffolds with aligned and
unaligned architectures and two culture conditions were investigated: 1) murine embryonic stem
cell-derived cardiomyocytes (mESCDCs) seeded alone onto the scaffolds or 2) mESCDCs
seeded onto the PUs pre-seeded with MEFs. In both culture conditions, viable mESCDCs
attached to the PU scaffolds and were functionally contractile out to at least 28 days post seeding
the mESCDCs. Importantly, the aligned scaffolds led to the anisotropic organization of rod-
shaped cells, improved sarcomere organization, and increased mESCDC aspect ratio (length to
diameter ratio) when compared to cells on the unaligned scaffolds. In addition, pre-seeding the
scaffolds with MEFs improved sarcomere formation, increased cell alignment and aspect ratio,
and led to a mESCDC morphology that was more extended on the PU scaffolds than the
mESCDCs cultured alone. These results suggest that both fiber alignment and pre-treatment of
scaffolds with fibroblasts improved the differentiation and organization of mESCDCs. The
results of this work are very promising for cardiac tissue engineering and further characterization
of the PU constructs will help to understand the potential of the PU scaffolds for use in the heart.
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6.1 Introduction Tissue engineering scaffolds that seek to mimic the native ECM should provide a
temporary polymer matrix that plays a structural and functional role in guiding tissue
development. Studying the cellular response to biomaterial scaffolds can help demonstrate if the
scaffolds meet several of the properties required in developing functional constructs for
regenerating injured or diseased tissues. Some of these requirements include the biomaterial and
degradation products being non-cytotoxic, appropriate physiochemical properties to promote cell
adhesion, growth, and differentiation, and a suitable architecture for proper tissue organization
and cell-cell coupling [2]. In vitro cell-based studies will help to understand if scaffolds elicit a
favorable response from the cells and may help in better defining criteria needed for specific
tissue engineering applications.
Previous work has demonstrated that the Phe PU has favorable properties to support cells
for cardiac tissue engineering. Phe PU films were successfully used to culture neonatal rat
cardiomyocytes and murine embryonic stem cell-derived cardiomyocytes (mESCDCs) while
maintaining normal phenotypic and contractile properties [3, 4]. Using thermally induced phase
separation (TIPS) and electrospinning, 3-D Phe PU scaffolds supported contractile mESCDCs
with cells on the electrospun fibrous scaffolds having a striated sarcomeric phenotype that
appeared more mature than cells on TIPS scaffolds [5]. Recently, the aligned and unaligned
electrospun Phe PU scaffolds developed for this work were used to culture primary neonatal
cardiac cells [6]. The aligned scaffolds provided physical cues for the anisotropic organization
of cardiac cells resulting in a more similar organization to native cardiac tissue. In addition, the
aligned cardiac constructs were associated with a decrease in atrial natriuretic peptide compared
to unaligned constructs suggesting a more mature, ventricular-like cardiac phenotype [6].
To further investigate the use of PUs for cardiac and other soft tissue engineering
applications, cell-based studies were conducted with electrospun PU scaffolds. Phe and Gly-Leu
PU scaffolds were seeded with mouse embryonic fibroblasts (MEFs) to test if the Gly-Leu PU
could support a high density of adherent cells and whether the biomaterial or any degradation
products released during the culture period were cytotoxic. Results from these experiments
suggest that both the Phe and Gly-Leu PU scaffolds could support a high density of adherent and
morphologically extended cells out to at least 28 days. As this work focuses on cardiac
applications, embryonic stem cells were differentiated into cardiomyocytes and seeded onto
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aligned and unaligned Phe PU scaffolds on their own or in coculture with MEFs. The
mESCDCs attached to the electrospun scaffolds, formed striated sarcomeric structures, expressed
the gap junctional protein Cx-43, and were functionally contractile on the scaffolds for at least 28
days. In addition, cells were organized in arrangements that were dictated by the underlying
fibrous scaffold substrate and helped to form constructs that were physically anisotropic,
resembling the organization in the native myocardium. Future work will help to characterize and
optimize the cardiac constructs but the results from this study support the use of PU scaffolds for
cardiac tissue engineering applications.
6.2 Materials and Methods All materials were purchased from Sigma-Aldrich Canada (Oakville, ON, Canada) unless
otherwise stated.
6.2.1 Mouse Embryonic Fibroblast Culture and Seeding onto Polyurethane Scaffolds§
Mouse embryonic fibroblasts (Calbiochem, San Diego, CA) were grown on tissue culture
polystyrene (TCPS) dishes in Dulbecco’s Modified Eagle’s Medium (DMEM) supplemented
with 10% fetal bovine serum (FBS), 1% non-essential amino acids, 1% penicillin/streptomycin,
1% L-glutamine, and 0.1 mM β-mercaptoethanol. Medium was changed every other day and
cells were split approximately every third day at a 1:3 ratio. The polymers were synthesized and
scaffolds were prepared as previously described in chapters 3 and 4 of this thesis. Electrospun
Phe and Gly-Leu PU scaffolds were punched into 5 mm discs, UV sterilized for 20 min per side,
and equilibrated in phosphate buffered saline solution (PBS) for at least 2 h. Scaffolds were
passively coated with fibronectin at 0.1 mg/ml in PBS at 37ºC overnight. Cells were statically
seeded on one side of the scaffold at 100,000 cells/scaffold and were cultured in 48-well non-
tissue culture-treated polystyrene plates. Constructs were characterized over a 28 day culture
period with ~75% of the medium changed everyday until analysis.
6.2.2 Characterization of MEFs on Phe and Gly-Leu-containing Polyurethanes§
Viability, density, and morphology of the cell-seeded Phe and Gly-Leu PU constructs
were determined during 28 days of culture by alamarBlue® analysis, Live/Dead® staining, and
§ Sections were published in [1].
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immunostaining. AlamarBlue® analysis was conducted on the cell-seeded constructs after 7, 14,
21, and 28 days in culture. The alamarBlue® dye (Invitrogen, Carlsbad, CA) was diluted with
fresh culture medium at 10% v/v prior to analysis. The constructs were incubated with the
solution for 4.5 h and were analyzed at 570 nm and 600 nm with a spectra thermo microplate
reader (SLT Labinstruments, Salzburg, Austria). Percent reduction of the alamarBlue® was
calculated according to the manufacturer’s instructions to give relative cell viability of MEFs on
the PU scaffolds. At least 5 samples (n>5) were analyzed at each time point and the experiment
was repeated three times (N=3). Statistical comparisons were made using a two-tailed
independent t-test or one-way analysis of variance (ANOVA) with Bonferroni post hoc analysis
using the SPSS Statistics 17.0 statistical software package (SPSS Inc, Chicago, IL).
A Live/Dead® assay (Invitrogen, Carlsbad, CA) was used to visualize live and dead cells
on the scaffolds after 28 days. Three samples in three separate trials (n=3, N=3) were stained
with calcein AM and ethidium homodimer-1 at concentrations of 2 μM and 20 μM in PBS
respectively for 20 min at 37˚C. Cells were examined with a confocal microscope (Carl Zeiss
Canada, Toronto, ON) at the Advanced Optical Microscopy Facility (AOMF; Princess Margaret
Hospital, Toronto, ON) equipped with Zeiss LSM software using excitation/emission spectrum
of 494/517 and 528/617 for calcein and ethidium homodimer-1 (in the presence of DNA)
respectively. The PU scaffolds were visualized in the confocal images using a 360 nm
wavelength due to auto-fluorescence of the polymers.
Cell density and morphology were determined by immunostaining MEFs with a focal
adhesion kit (Chemicon® International, Temecula, California). The actin cytoskeleton, cell
nuclei, and focal adhesions were stained using TRITC-conjugated phalloidin, 4',6-diamidino-2-
phenylindole (DAPI), and an anti-vinculin monoclonal antibody respectively. After 28 days in
culture, samples were removed from culture medium, rinsed in PBS, and fixed with 4%
paraformaldehyde (Electron Microscopy Sciences, Ft. Washington, PA) in PBS for 20 min at
room temperature. Fixed constructs were washed 3 times in blocking solution (Hank’s Balanced
Salt Solution containing 2% FBS and 4% bovine serum albumin), permeabilized with an
Intraprep® permeabilization reagent (Beckman Coulter, Mississauga, ON) for 10 min, and
incubated in blocking solution for 30 min. Samples were labeled with anti-vinculin (1:250
dilution in blocking solution) for 1 h at room temperature, followed by labeling with TRITC-
conjugated phalloidin (1:500 dilution) and Alexa 488 anti-mouse secondary antibody (10 μg/ml;
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Molecular Probes) for 1 h, and subsequently with DAPI (0.5 μg/ml) for 10 min. The constructs
were washed 3X in blocking buffer between each step. Three labeled samples from three
separate trials (n=3, N=3) were imaged using a two-photon confocal microscope with Zeiss LSM
software.
6.2.3 Culture and Differentiation of Murine Embryonic Stem Cells The D3 murine embryonic stem cell line was previously transfected via electroporation
with a vector carrying both a phosphoglycerate kinase promoter driving a hygromycin resistance
gene and an α-cardiac myosin heavy chain promoter in front of the neomycin resistance gene
(MHC-neor/pGK-hygror) to select for mESC-derived cardiomyocytes [7]. Undifferentiated
ESCs were cultured as described for the CM 7/1 cell line [3] on 0.2% v/v gelatin coated tissue
culture polystyrene dishes in DMEM supplemented with 15% ESC-screened FBS (Hyclone,
Logan, UT), 1% L-glutamine, 1% non-essential amino acids, 1% sodium pyruvate, 1%
penicillin/streptomycin, 0.1 mM β-mercaptoethanol, and 1000 units/ml leukemia inhibitory
factor (LIF). Cells were maintained in culture for at least two passages after thawing before
initiating differentiation.
The large-scale differentiation of mESCs and selection of desired cardiomyocytes was
carried out using a similar system to that described by Zandstra et al. [7]. Figure 6.1 illustrates
the experimental details for this study. The mESCs were grown on gelatin coated TCPS and
subsequent growth and differentiation took place in 250 ml glass bulb spinner flasks (CELLspin
250; Integra Biosciences, Switzerland). Differentiation was initiated (day 0) by suspending
mESCs in 125 ml of culture medium in the absence of LIF (differentiation medium; no LIF, no
sodium pyruvate) at 100,000 cells/ml. Cells were placed in the spinner flasks running at an
impeller speed of 60 rpm in an incubator (37˚C, 5% CO2). After 24 h, cells aggregated to form
embryoid bodies (EBs) and 125 ml of fresh differentiation medium was added to the spinner
flask after an additional 24 h (250 ml total volume). Every subsequent day, the cells were
settled, and 50% of the media was exchanged for fresh differentiation media. Starting on day 9,
the differentiation medium was supplemented with the antibiotic G418 (400 μg/ml) and all-trans
retinoic acid (10-9 M) to select for cardiomyocytes (selection medium); 50% media was exchange
with fresh selection medium everyday until the cells were harvested from the spinner flask on
day 18.
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Figure 6.1: Illustration of experimental details for cardiomyocyte production and cell seeding. Large scale cell
growth, differentiation, and selection was performed in spinner flasks. Cardiac bodies were dissociated into individual mESCDCs and were seeded onto the PU scaffolds on their own or pre-seeded with MEFs.
6.2.4 Monitoring the Differentiation of Cardiomyocytes from mESCs Cell growth kinetics were monitored throughout the 18 day differentiation/selection
period by removing a 1 ml sample from the spinner flask to count the number of EBs/ml and by
fully dissociating a 1 ml sample of EBs to determine cell number. EBs were dissociated into
single cells up to day 9 by incubating with 0.25% trypsin-EDTA (Gibco, Invitrogen Corporation,
Carlsbad, CA) for 5 min followed by the addition of culture medium and physical disruption by
pipetting up and down. Starting day 9 and on, EBs were dissociated by incubating the cells with
collagenase type IV (1mg/ml) supplemented with 2% FBS at 37°C for 20 min followed by
addition of DNAse (1 mg/ml). After an additional 10 min, fresh medium was added and cells
were centrifuged at 900 rpm. Medium was aspirated and 0.25% trypsin-EDTA solution was
added for 5 min. Cells were physically disrupted by pipetting after neutralizing the enzymes
with fresh medium and were counted using trypan blue dye.
The successful differentiation of the embryonic stem cells into cardiomyocytes was
investigated by flow cytometry. Cells dissociated after removal from the spinner flask were
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placed in a Falcon tube with Hank’s Balanced Salt Solution containing 2% FBS at approximately
1 million cells per tube. Ethidium monoazide (1 µg/ml) was added to the samples and were
placed on ice under bright light for 30 min. An Intraprep permeabilization kit was then used to
fix and permeabilize the cells according to kit instructions. Samples were subsequently labeled
with either an anti-cardiac isoform of troponin-T (clone 13-11; 2 µg/ml; Lab Vision, Fremont,
CA) or an anti-Oct 4 (2.5 µg/ml; BD Biosciences, San Jose, CA) primary antibody and a rabbit
anti-mouse IgG FITC-conjugated secondary antibody (1:100 dilution; Invitrogen Corporation,
Carlsbad, CA). Rinses with Hank’s Balanced Salt Solution with FBS were performed between
steps. Samples were analyzed at the Faculty of Medicine Flow Cytometry Facility (University of
Toronto, Toronto, ON) using a BD FACS Calibur analyzer with an Argon-488 nm blue laser
excitation. FITC was detected with a 530/30 nm band pass filter and EMA was detected using a
670 nm long pass filter.
6.2.5 Scaffold Preparation and Cell Seeding MEFs were cultured on TCPS dishes and aligned and unaligned Phe PU scaffolds were
prepared as described above. One day prior to the end of mESCDC selection in the spinner flask
(i.e. day 17), MEFs were seeded statically onto one side of the aligned and unaligned scaffolds at
a density of approximately 12,500 cells per scaffold. Cell seeded constructs were placed in an
incubator (37ºC, 5% CO2) in differentiation medium overnight. On day 18, aggregated
cardiomyocytes (cardiac bodies, CBs) were removed from the spinner flask, allowed to settle,
and excess medium was removed. Cardiac bodies were dissociated to obtain single cells by
incubation with collagenase type IV-DNAse and trypsin-EDTA steps as outlined above.
Following dissociation, the cells were filtered with a 40 µm cell filter (Becton Dickenson,
Mississauga, ON) and approximately 2 million ESCDCs were seeded on the same side of the
scaffolds as the MEFs. MEFs or ESCDCs alone on aligned and unaligned scaffolds were also
prepared. MEFs or ESCDCs alone or in coculture on TCPS served as controls. Constructs were
cultured in differentiation medium with 75% of the medium exchanged everyday until analysis.
6.2.6 Characterization of mESCDCs and MEFs on Aligned and Unaligned Polyurethane Scaffolds
AlamarBlue® analysis and Live/Dead® staining were used as described above to confirm
the success of the seeding protocol. Cell morphology, organization, and protein expression were
characterized by immunostaining the cell-seeded PU constructs. The actin cytoskeleton, cell
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nuclei, and either a sarcomeric structural protein or gap junctional protein were stained using
TRITC-conjugated phalloidin (Chemicon), DAPI (Chemicon), and either an anti-α-actinin or
anti-connexin-43 (Cx-43) monoclonal antibody respectively. The protocol for fixing and
immunostaining MEFs on Phe and Gly-Leu PU scaffolds described above was used here with the
aligned and unaligned constructs. Samples were labeled with either anti-α-actinin or anti-Cx-43
at a 1:1500 dilution for both primary antibodies and the same dilutions for TRITC-conjugated
phalloidin, Alexa 488 anti-mouse or anti-rabbit secondary antibodies, and DAPI as previously
outlined. Labeled samples were imaged using a two-photon confocal microscope with Zeiss
LSM software. Prior to obtaining images, polymer auto-fluorescence was used to help orient the
constructs so the major axis of fiber alignment was parallel with the top to bottom plane of each
image. Image Pro Express image analysis software (Media Cybernetics, Bethesda, MD) was
used to quantify the angle of cell axis and give qualitative measurements of cell length and
diameter, aspect ratio (length/diameter), and cell area (length*diameter). Cell angles and
dimensions were measured manually using the image analysis software from at least 200 cells
from 6 to 9 different images. Image Pro Analyzer 6.3 software (Media Cybernetics, Bethesda,
MD) was used to determine the orientation index and standard cell area as a second method of
quantifying cell orientation and area respectively. The orientation index was determined by
performing a Fast Fourier Transform (FFT) on 6 to 9 different immunostained images as
described by Nichol et al. [8]. Subtracting the minor to major axis ratio of the thresholded FFT
image from 1 resulted in the orientation index. A value of 0 corresponded to a completely
random orientation whereas a value of 1 corresponded to a perfectly aligned orientation. The
standard cell area was calculated by quantifying the area of f-actin expression and dividing by
the number of cell nuclei quantified in the same image. Statistical comparisons were made using
either a two-tailed independent t-test or a one-way ANOVA with Bonferroni post hoc analysis.
Beating constructs were captured by video microscopy using an Olympus microscope
equipped with a Sony EXwaveHAD colorwide digital camera. Image J software was used to
record cell contraction and produce the videos. Samples were assessed after 6 and 14 days in
culture.
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6.3 Results and Discussion
6.3.1 Viability of MEFs on Phe and Gly-Leu-containing Polyurethanes§ In cell-based approaches to tissue engineering, scaffolds must be able to support the
attachment of cells and neither the material itself nor its degradation products should be
cytotoxic. To assess whether the Gly-Leu PU could meet these fundamental criteria, MEFs were
seeded onto the electrospun PU scaffolds at a high seeding density and viability and morphology
was evaluated during a 28 day culture period. MEFs were chosen as the cells for this work for a
few reasons. First, fibroblasts are one of the main cell types involved in tissue remodeling
through the secretion of matrix metalloproteinases. Culturing MEFs on the PU scaffolds may
result in the expression and activation of MMPs and may promote the degradation of the Gly-
Leu PU scaffolds. Assessing the viability of the cells over time may help to identify if any
cytotoxic degradation products are released from the polymers. In addition, this work will help
to establish culture conditions that will directly translate to future cell-based degradation studies.
Second, increasing evidence suggests a critical role of cardiac fibroblasts in the development of
myocardial tissue engineered constructs [9-12]. Our lab has been particularly interested in
developing myocardial constructs using mESCDCs. It has been identified that coculturing these
cells with MEFs improves mESCDC adhesion to material surfaces and aids in their functional
properties [13]. Towards the goal of a cardiac patch using PU scaffolds described later in this
chapter, MEFs were used in a coculture system with the mESCDCs. Demonstrating the ability
of the electrospun PU scaffolds to support MEFs provided a starting point for pre-seeding the
aligned and unaligned scaffolds with MEFs for the coculture studies.
After preparing fibronectin-coated Phe and Gly-Leu PU scaffolds, MEFs were statically
seeded on the matrices at a high seeding density. Viability on the cell-seeded constructs was
determined by alamarBlue® analysis yielding a relative cell number on the PU scaffolds (Figure
6.2). AlamarBlue® is reduced during cellular respiration and, assuming all the MEFs undergo a
similar rate of respiration, a higher percent reduction will correlate to a higher number of viable
cells on the scaffolds. Both PUs had a similar trend in cell viability over the 28 day period.
Initially, both PUs showed an increase in percent reduction from day 7 to day 14 identifying
some cell growth during this period, although this change was only significant for the Phe PU
(ANOVA, p<0.05). From day 14 to day 21, the Phe PU had a slight increase in percent § Large portions of this section were published in [1].
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reduction while the Gly-Leu PU exhibited a slight decrease. Day 21 to day 28 showed a
significant decrease in alamarBlue® reduction for both PUs (ANOVA, p<0.05) potentially
indicating the occurrence of cell death. Interestingly, despite similar trends over the 28 day
culture period, the Phe PU had a significantly higher percent reduction compared to the Gly-Leu
PU for all time points investigated (t-test, p<0.05) suggesting the Phe PU was able to support a
higher number of viable cells. TCPS controls were conducted alongside the PU scaffolds but the
relative percent reduction with the control was always 100% due to increased culture area
associated with the 48-well TCPS plates. A TCPS control group containing a similar number of
cells as observed on the PUs may help to determine if the drop in relative cell number at day 28
was a function of the culture substrates.
Figure 6.2: AlamarBlue® analysis of MEFs on unaligned Phe and Gly-Leu PU scaffolds over 28 day period. A two-
tailed independent t-test indentified a statistical difference in means between PU types for each time point. * represents a statistical difference in means using a one-way ANOVA with Bonferroni post hoc analysis (p<0.05)
when looking at the time progression of each PU. Error bars represent ± standard deviation. n>5, N=3.
To further investigate the cell seeded constructs, cells were visualized by Live/Dead®
staining (Figure 6.3a) and immunostaining (Figure 6.3b). The results from the Live/Dead®
staining of the cell seeded constructs identified a high density of viable cells on both scaffold
types and TCPS controls after 28 days post seeding. Despite the alamarBlue® results suggesting
some cell death from day 21 to day 28 on the PU scaffolds, there was little evidence of any dead
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cells on the scaffolds at 28 days. Although a decrease in percent reduction may indicate cell
death, it could also be a consequence of cell metabolism slowing down. In addition, dead cells
may be washed off the scaffolds during medium changes or washes performed during assays and
this may account for their absence in the obtained images.
Figure 6.3: Staining of MEFs on unaligned Phe and Gly-Leu PU scaffolds and TCPS by a) Live/Dead staining
(green = live cells, red = dead cells, blue = PU scaffold) and b) Immunostaining (green = vinculin, red = f-actin, blue = cell nuclei). n=3, N=3.
Immunostaining the actin cytoskeleton, cell nuclei, and focal adhesions confirmed the
Live/Dead® results of a high density of cells on the PU scaffolds. Cells appeared attached and
extended on all substrates with no apparent orientation. A fairly confluent layer of cells was
identified for the culture surfaces with cells in close contact or on top of each other in many
spots. This may be a result of cell overgrowth if MEFs are not contact inhibited. However, this
phenomenon more likely reflects the high cell density used for seeding as a similar result was
observed when immunostaining the cells after 7 days in culture. Interestingly, the Phe PU
scaffolds had a confluent layer of cells on both sides of the scaffolds even though cells were
seeded on only one side. In contrast, the Gly-Leu PU scaffolds had a confluent layer on the cell-
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seeded side but only high density patches on the non-seeded side. This observation correlates
with the alamarBlue® results that suggested a higher relative number of viable cells seen on the
Phe PU scaffolds compared to the Gly-Leu PUs. The non-cell-seeded side of the Phe PU may
have been in closer contact with the bottom of the culture wells during and after seeding and may
have allowed for better cell attachment and migration than the Gly-Leu PU scaffolds. Seeding
both sides of the PU scaffolds may help to resolve this difference, as it may not accurately reflect
the Phe PUs ability to support a greater number of cells.
A reduced cell number after 28 days in culture compared to earlier time points may have
occurred, but the results from this cell-based study suggest both PU scaffolds are capable of
supporting a high density of MEFs out to at least 28 days. The ability of the electrospun PU
scaffolds to support the attachment and viability of cells was not surprising. Electrospun fibrous
scaffolds with random architectures and micron-sized fibers have been shown to promote cell
attachment [14, 15]. Moreover, mESCDCs seeded on electrospun Phe PU scaffolds attached to
the material and maintained a phenotype typical of this cell type [5]. Previous cytotoxicity
testing of the Phe-based family of PUs in the presence of keratinocytes suggested the PUs were
not cytotoxic but that PU degradation was associated with a decrease in cell viability [16]. The
chemical reactants used in synthesizing the Phe and Gly-Leu PUs were chosen because they are
non-toxic [17, 18]. A systematic investigation of chymotrypsin-mediated cleavage sites in the
Phe PU similarly suggested that non-toxic products are released from the material during
degradation [19]. The large similarities between the Phe and Gly-Leu PU, including polymer
chemistry and properties, scaffold architecture, and fiber diameters, would suggest a similar
ability of the Gly-Leu PU to support the attachment and viability of cells and this is supported by
the studies conducted here.
6.3.2 Differentiation of mESCs into Cardiomyocytes in Spinner Flasks The source of cells used in cardiac tissue engineering is a critical component to
generating functional, force generating tissue that can be used to regenerate infarcted or diseased
myocardium. Embryonic stem cells represent a source of a potentially unlimited number of de
novo cardiomyocytes and are of great interest in approaches to cardiac repair [7]. ESCDCs have
the ability to form electromechanical coupling with host myocardium that is required in
establishing true regeneration of the cardiac tissue [13, 20, 21]. One of the criteria required for
the successful employment of ESCs in the formation of engineered myocardial tissue is a large
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and pure population of cardiac progenitors or fully differentiated cardiomyocytes. The Zandstra
laboratory has been particularly interested in achieving this goal by developing and optimizing
bioreactor parameters for the generation of large quantities of ESCDCs [7, 22, 23]. The systems
developed by this group utilize genetically modified mESCs (MHC-neor/pGK-hygror) that allow
for the selection of cardiomyocytes to obtain a pure population of desired cells. In one study,
Zandstra et al [7] generated ~14 x 106 mESCDCs in a 250 ml spinner flask that were
spontaneously beating and expressed characteristic markers of this cell type. The
cardiomyocytes were produced by aggregating mESCs in static culture for 4 days followed by
the additional growth and differentiation in a spinner flask system. On day 9 after initiating
differentiation, medium was supplemented with G418 and retinoic acid to select for and drive
differentiation towards cardiomyocytes. On day 18 the relatively pure mESCDC population was
harvested from the spinner flasks. In collaboration with the Zandstra group, a similar system to
generate mESCDCs was previously performed in our lab towards the development of a cardiac
patch using the Phe PU [3, 5] and was used again for the work described here.
Several methods have been used to form EBs including suspension culture in bacterial-
grade dishes, culture in methylcellulose semisolid media, the hanging drop method, aggregation
in a 96-well round bottom plate or conical tube, and methods for the scalable production of EBs
[24]. The large scale production of mESCDCs for tissue engineering applications requires the
scalable production of these cells in spinner flasks and bioreactor systems. Towards this goal,
EBs have been formed by suspension culture with bacterial-grade dishes for 4 days before
transfer to spinner flasks or the static suspension culture for 1 day followed by encapsulation of
the EBs in size-specific agarose hydrogel capsules [7, 22]. This was performed to prevent
undesirable EB agglomeration using spinner flasks with paddle-type impellers. Other systems
have been developed that allow the direct formation of EBs within the spinner flasks or
bioreactors thereby eliminating this 2-step process [24]. Schroeder et al. [25] compared the
paddle-type and glass bulb-shaped impeller spinner flasks and found that the bulb-shaped
impellers reproducibly formed homogenous EBs by directly inoculating mESCs whereas the
paddle impellers caused EB agglomeration. Adjusting impeller speed affected the shear stress on
cells and influenced EB size, formation, and subsequent cardiomyocyte differentiation [25].
A culture period of 18 days in spinner flasks was used to generate mESCDCs for
subsequent investigations with the Phe PU scaffolds. The D3 MHC-neor/pGK-hygror mESC line
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was grown on gelatin-coated TCPS dishes in culture medium containing LIF and 12.5 million
cells were inoculated in 250 ml glass bulb-shaped impeller spinner flasks in the absence of LIF
(day 0). Cell aggregation occurred over the first 24 h to form embryoid bodies and subsequent
cell growth and differentiation took place within EBs. Starting on day 9, G418 and retinoic acid
were added to the culture medium to select for and drive differentiation towards cardiomyocytes
respectively. Cells were harvested from the spinner flasks on day 18 for the cell-based studies.
Cell growth and differentiation were monitored during the 18 day period by counting cells and
EBs and flow cytometry for expression of cell markers.
Cell number in the spinner flasks was measured by dissociating EBs and counting the
cells at various time points during the 18 day differentiation and selection period. Results are
shown in Figure 6.4. High cell expansion within the EB structures was observed during the first
6 days of culture reaching a maximum number of approximately 1 x 109 total cells on day 6.
From day 6 on, a steady drop in total cell number occurred out to day 18 when the cells were
harvested from the spinner flasks. The drop in cell number from day 9 on was expected due to
the heterogeneous cell population generated during EB differentiation and the start of antibiotic
selection of cardiomyocytes. The drop in cell number from day 6 to day 9, however, occurred
prior to the addition of G418 and was not an observation previously recorded in the large-scale
production of mESCDCs in spinner flasks [7, 23]. Diffusion limitations as EB size increased
most likely affected cell viability. Alternatively, insufficient nutrient exchange and
accumulation of toxic waste products due to the high cell population could account for the drop
in cell numbers. Monitoring EB size and glucose/lactate levels may help identify the cause of
cell death prior to selection and may indicate potential methods for optimizing this protocol.
Despite cell death prior to selection, a large number of cells were harvested from the spinner
flasks after 18 days. The mESCDC output to mESC input ratio using similar systems for
mESCDC production have recorded values from ~1.5-3.5 [22, 23]. Although some differences
exist between those studies and the one conducted here, the actual viable cardiomyocyte yields
likely correspond to these numbers more closely than the total cell number may suggest (~12
mESCDC/input mESC).
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Figure 6.4: Total cell number in spinner flasks during differentiation of mESCs into cardiomyocytes. Cell growth
was observed for 6 days followed by a drop in cell number for additional 12 days in suspension. A drop in cell number was observed prior to the start of selection indicating room for optimization of culture parameters. Error
bars represent ± standard deviation. n=4.
In addition to total cell number, EB number and cells/EB were monitored during the first
9 days of culture in the spinner flasks. As seen in Figure 6.5a, a high and variable EB number
was observed on day 2 followed by a lower and more constant EB number for the remainder of
the time period investigated. The large deviation on day 2 indicated that the EB formation in the
spinner flasks was not as reproducible as previous studies have shown [25]. In addition,
although shear stresses in the glass bulb-shaped impeller spinner flasks limit EB agglomeration
[25], some smaller EBs may still have aggregated during the first 4 days in suspension. This is
consistent with the expression of E-cadherin and EB agglomeration seen in other spinner flask
and bioreactor systems [26]. The number of cells per EB during the first 9 days in the spinner
flask is shown in Figure 6.5b. A similar trend as total cell number was observed with a steady
increase in number of cells per EB from inoculation to day 6 followed by a decrease from day 6
to day 9. This was expected given the trend of total cell number and the relatively constant EB
number. The maximum number of cells per EB on day 6 was ~9,000 ± 1,800 and may suggest a
maximum size before diffusion limitations to the center of the EBs became a limiting factor to
cell viability. Niebruegge et al. [23] recorded a higher number of cells per EB with the CM7/1
cell line, ~25,000 cells per EB, before cell number began to drop. This could reflect differences
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in cell line or that it was not EB size but rather total number of cells and culture medium
limitations that decreased cell viability. Interestingly, the trials that had fewer EB numbers on
day 2 and a correspondingly higher number of cells per EB produced a higher number of cells on
day 18 and had a higher frequency of beating EBs and total cardiomyocytes. Starting EB size
has been identified as an important determinant of differentiation towards specific lineages and
higher mesoderm and cardiac induction has been associated with larger EB sizes [27].
Controlling the starting EB size has therefore emerged as a way to increase cardiomyocyte yields
from ESCs [28]. The results here suggest that the larger EB sizes similarly produced a higher
number of cardiomyocytes. Further characterization of ESC differentiation in the spinner flasks
should be conducted to better understand the production of cardiomyocytes using the D3 mESC
line and the glass bulb-shaped impeller spinner flasks. Specifically, characterizing EB size,
glucose/lactate levels, frequency of beating EBs, and better quantifying viable cardiomyocytes
will help determine the cause of cell death prior to selection and may identify potential
parameters in the process that need to be optimized.
Figure 6.5: EB characteristics during differentiation of mESCs into cardiomyocytes. a) EB number and b) cells/EB
for first 9 days of culture in spinner flasks. Error bars represent ± standard deviation. n=4.
The cell populations that were used to inoculate the spinner flask and were harvested
after 18 days of culture were identified by flow cytometry. Oct 4 is a transcription factor
expressed by undifferentiated mESCs that is down regulated during differentiation and is used as
a marker of pluripotent mESCs [29]. Oct 4 has been commonly used to identify undifferentiated
mESCs in cell populations for the scalable production of cardiomyocytes [7, 22] and was used to
identify mESCs in the cell populations before and after differentiation in spinner flasks. Cardiac
cells were identified by expression of the cardiac isoform of troponin T (cTnT). Troponin T is a
component of the troponin complex involved in regulating actin-myosin interactions during
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cardiomyocyte contraction and has been used as a marker to identify mESCDCs [30, 31]. Flow
cytometry analysis of the cell populations at day 0 and day 18 is shown in Figure 6.6. Ethidium
monoazide was used to indicate cell viability and the data was gated to remove dead cells. The
cell population used to inoculate the spinner flasks on day 0 expressed high levels of Oct 4
(>80%) indicating a high percentage of undifferentiated ESCs. Cardiac cells were not identified
in the starting cell population. Following 9 days of cell growth and differentiation and an
additional 9 days of selection (day 18), EBs were contracting spontaneously (referred to as
cardiac bodies, CBs) indicating the presence of cardiomyocytes. Cardiac-specific troponin T
staining of cells at day 18 identified that the viable cells in the CBs were composed of a
relatively pure population of cardiomyocytes (~97% cTnT positive). In addition, no Oct 4
positive cells were observed demonstrating the absence of any undifferentiated ESCs. The
selection of mESCDCs by antibiotics using genetically modified mESCs has been shown to
result in a pure population of cardiac cells [20, 32] and is currently the best method of purifying
cardiomyocytes from a heterogeneous cell population [33]. The potential of teratoma formation
associated with the transplantation of undifferentiated ESCs in the heart is a major concern of
using ESCs for cardiac repair [13, 34-36]. Ensuring no undifferentiated ESCs are present in cell
populations used in developing engineered myocardial constructs is therefore of utmost
importance. Although repeated trials are needed to confirm the results, a relatively pure
population of mESCDCs with no evidence of undifferentiated mESCs was obtained for
subsequent experiments using PU scaffolds.
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Figure 6.6: Flow cytometry of cells before and after differentiation in spinner flasks. Cells were analyzed at a) day 0
and b) day 18. A high Oct 4 expression was observed on day 0 indicating a high proportion of undifferentiated ESCs. On day 18, a high population of cardiac troponin T (cTnT) positive cells with no Oct 4 positive cells was
observed consistent with a pure population of cardiomyocytes and no undifferentiated ESCs. n=1.
6.3.3 Effect of Fiber Alignment and Coculture with MEFs on Response of mESC-derived Cardiomyocytes
Biological, chemical and mechanical signals provided by tissue engineering scaffolds can
be used to direct cell behavior. Fiber alignment within biomaterial scaffolds has been shown to
influence cell orientation, growth, and other cellular processes [37] and may provide the
appropriate physical cues required for the formation of anisotropic cardiac tissue. Previous work
towards the development of a cardiac patch has demonstrated the Phe PU has suitable properties
for culturing cardiac cells [3-6]. To further investigate the use of PU scaffolds in cardiac tissue
engineering, the Phe PU was electrospun into scaffolds with aligned and unaligned architectures
as described and characterized in chapter 4. The scaffolds of different architecture were seeded
with mESCDCs alone and with MEFs to investigate the influence that fiber alignment has on cell
attachment, organization, contractile function, and protein expression.
Embryonic stem cells are an important source of cardiomyocytes for regenerating injured
myocardium. Investigating the development of myocardial constructs using murine ESCDCs
provides an important model for testing scaffold properties for cardiac tissue engineering and can
give important insight to translate to human ESCDCs for clinically relevant cardiac repair. In
addition to cardiomyocytes, increasing evidence suggests a critical role of cardiac fibroblasts in
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the development of tissue engineered cardiac constructs [9-12, 38]. Similarly, coculturing
mESCDCs on top of MEFs improved mESCDC adhesion, helped maintain an elongated
phenotype for several weeks, and aided in the electrical properties of the cells [13]. Cellular
transplantation of mESCDCs with an equal number of MEFs improved engraftment of the cells
and significantly improved cardiac performance in infarcted hearts [13]. This work suggested
that the coculture of mESCDCs and MEFs may improve the attachment, elongation, and function
of mESCDC on the Phe PU scaffolds.
Pretreatment of elastomeric scaffolds with fibroblasts has been shown to improve cell
density, tissue compaction, and functional properties of myocardial constructs by providing a
supportive environment for cell attachment, differentiation, and contractile function [11, 38]. It
was hypothesized that seeding MEFs one day before seeding mESCDCs may lead to a higher
density of cardiomyocytes on the Phe PU scaffolds by improving cell attachment and could have
functional benefits on the engineered tissue compared to seeding mESCDCs alone. After the 18
day differentiation and selection period, mESCDCs were harvested from the spinner flasks and
were seeded onto the elastomeric matrices either on their own or after the scaffolds had been pre-
treated with MEFs. Preliminary work looked at seeding partially dissociated cardiac bodies, but
most cells remained as CBs and did not appear to be interacting with the scaffolds. In order to
test how fiber alignment influenced cell behavior, it was important that the cells interact with the
PU scaffolds and therefore the CBs were fully dissociated into individual cells prior to seeding.
The constructs were subsequently analyzed by several techniques to characterize the cellular
response to the different scaffold architectures.
The cell-seeded constructs were assessed by alamarBlue® analysis to identify if the
different architectures influenced the number of cells on the two scaffold types following cell
seeding. Figure 6.7 shows the percent reduction of alamarBlue® by MEFs, mESCDCs, and
coculture of the two cell types on aligned and unaligned Phe PU scaffolds and TCPS controls at
different time points post seeding. A similar percent reduction of alamarBlue® was observed at
each time point for all cell groups on the aligned and unaligned scaffolds and suggested that
architecture did not affect cell attachment to the PU substrates (ANOVA, p>0.05). In contrast,
cells seeded on TCPS had a significantly higher percent reduction compared to either PU
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Figure 6.7: AlamarBlue analysis of cell-seeded PU constructs of varying architecture and TCPS controls. a)
coculture, b) MEFs alone, and c) mESCDCs alone. No statistical difference was observed in alamarBlue reduction between the aligned and unaligned PU scaffolds for any time point or any cell type (ANOVA, p>0.05). Cells
cultured on TCPS were significantly different than both aligned and unaligned scaffolds for all cells and all time points investigated (p<0.05). n>3, N=3 (except mESCDCs alone, N=2).
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scaffolds for all cells and time points (p<0.05). Although it is unclear the reason for this,
possible explanations may include: a higher number of seeded cells adhered to the TCPS than
those seeded on PU scaffolds; the increased culture area with TCPS wells allowed for
proliferation of cells; or the cells on TCPS expressed a phenotype that was more metabolically
active than cells on PU. A direct comparison of alamarBlue® reduction between different cell
groups was difficult due to differences in metabolic activity for each cell type. A higher percent
reduction, however, was observed for coculture group compared to MEFs alone suggesting that
some mESCDCs adhered to the PU scaffolds.
To visualize viable and non-viable cells on the PU scaffolds, cells were prepared by
Live/Dead® staining. The heterogeneous nature of mESC differentiation in EBs leads to
formation of non-cardiac myocytes and selecting for the cardiomyocytes leads to cell death in the
non-myocyte population. No attempt was made to separate out the dead cells prior to mESCDC
seeding and it was of interest to determine if any dead cells remained attached to the scaffolds or
if they were washed away during media changes. Cytokines released from dead cells may affect
cell behavior and can elicit a strong inflammatory response in vivo [39] that could adversely
affect function and integration with the host if these constructs are implanted in animal models.
Figure 6.8 shows images from the Live/Dead® staining of MEFs and coculture cell groups on the
aligned and unaligned scaffolds. In these images, green corresponds to viable cells, red spheres
to dead cells, and red and blue fibers to the PU scaffold due to auto-fluorescence of the polymer
at the investigated wavelengths. All of the images show a fairly high number of viable cells on
the PU scaffolds. MEFs on the scaffolds had a fairly uniform cell distribution but cell-cell
contact was limited. There was no sign of dead MEFs on the scaffolds. The coculture group
similarly had viable cells over most of the scaffolds but the culture surface had several large high
density patches of viable cells that were in close contact. A few small round red spots were
observed on the coculture constructs indicating the presence of some dead cells. Although some
dead cells were identified, the frequency was not high suggesting the majority of the dead cells
were washed away following seeding or during medium changes. As a result, there was no
motivation to change the seeding protocol to separate viable and dead cells by Percoll gradient
prior to seeding. Visualizing the cells also demonstrated preliminary differences in cell
organization on the aligned and unaligned scaffolds. Cells elongated along the PU fibers and
appeared aligned with the orientation of fibers in the aligned scaffold whereas they were more
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randomly oriented on the unaligned PU. This was consistent with cells on other electrospun
polymers [37] and suggested that the underlying fibrous substrates were providing physical cues
for cellular organization. The results from the Live/Dead® staining confirmed the ability to
successfully seed a high number of viable cells on the PU substrates and that the cells appeared
to be interacting with PU fibers. This was an important step in developing a system to
investigate the influence fiber alignment had on the cardiac cells.
Figure 6.8: Live/Dead staining of cells on Phe PU scaffolds of varying architecture at day 18+6. Green = viable cells, Blue = PU scaffold, Red = dead cells (round spots) and PU scaffold (fibers). High density patches of viable
cells were observed for coculture constructs with the presence of some dead cells. MEFs on PU matrices were viable and had a more scattered cell distribution with no dead cells. n=3, N=3.
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Cell morphology, organization, and protein expression were characterized by
immunostaining the cell-seeded aligned and unaligned PU scaffolds. Constructs were stained for
the actin cytoskeleton (f-actin), cell nuclei, and sarcomeric structure (α-actinin) 6 days after
mESCDC seeding (Figure 6.9). In the coculture images, both cell types are positively stained for
the actin cytoskeleton (red) and cell nuclei (blue), but only the mESCDCs will express the
sarcomeric structure (green). The initial results suggested by Live/Dead® regarding cell
attachment and distribution on the substrates were confirmed by immunohistochemistry.
Adherent and extended cells were observed for the cell groups on all substrates indicating the
surfaces promoted cell attachment and interaction with the two PU architectures. The MEFs
were uniformly distributed on the scaffolds but had limited cell-cell contact. The mESCDCs
seeded alone and in coculture with the MEFs were observed as single cells, in small cell clusters,
or more frequently in high-density patches around the scaffolds. Cells within these patches were
in close contact with the PU substrates and other cells suggesting the potential of cell-matrix and
cell-cell contacts that are critical to the functionality of myocardial tissue. The mESCDCs
Figure 6.9: Immunostaining of cells on aligned and unaligned PU scaffolds. Red = cytoskeleton (f-actin), Green =
sarcomere (α-actinin), Blue = cell nuclei. n=3, N=3.
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cultured alone and with MEFs on both scaffold architectures exhibited different degrees of
differentiation as indicated by cell shape and the appearance of a striated sarcomeric structure.
Early stages of mESCDC differentiation within EBs are typically characterized by small round
cells with no evidence of sarcomere development [40, 41]. With increased time, mESCDCs
become elongated, rod-shaped cells with highly developed and organized sarcomeric structures,
Figure 6.10: Immunostaining of cardiac constructs with mESCDCs showing varying levels of differentiation. a)
coculture on unaligned PU, b) coculture on aligned scaffolds, c) mESCDCs alone on unaligned PU, and d) mESCDCs alone on aligned scaffolds. Red = cytoskeleton (f-actin), Green = sarcomere (α-actinin), Blue = cell
nuclei, Yellow arrows = round cardiomyocytes with poorly defined sarcomeric structures, White arrows = elongated cardiomyocytes with well defined and organized sarcomeres.
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consistent with cardiomyocyte development in vivo [40, 41]. The mESCDCs on PU scaffolds
varied from immature cells that were round or triangular shaped with α-actinin staining in the
cytosol and no evidence of sarcomere assembly, to more mature cells that were rod-shaped with
highly defined and organized sarcomeric structures (Figure 6.10). Importantly, ranking cell
shape and the appearance of striated structures within the immunostained images by blinded
observers identified a trend that both fiber alignment and coculture with fibroblasts improved the
differentiated state of the mESCDCs (Table 6.1). A higher percentage of rod-shaped mESCDCs
were found on the aligned scaffolds compared to the unaligned scaffolds. In addition,
coculturing with MEFs reduced the percentage of circular cells and increased the striated cell
appearance compared to mESCDCs cultured alone.
Table 6.1: Assessment of cell shape and sarcomere formation for mESCDCs cultured alone and in coculture with MEFs on PU scaffolds. Three blinded observers ranked 60-200 cells from 6-9 different images for cell shape and
appearance of striated structure within each culture group. Values are given as mean ± standard deviation.
Cell Shape (%) Striated Appearance (%)
Cylindrical Triangular Circular Aligned mESCDCs alone 67.0 ± 28.2 10.7 ± 12.5 22.3 ± 15.6 47.3 ± 0.8
Unaligned mESCDCs alone 26.3 ± 28.8 41.3 ± 20.7 32.5 ± 8.2 48.6 ± 6.6 Aligned Coculture 70.7 ± 25.6 15.7 ± 18.7 13.5 ± 6.9 73.0 ± 0.7
Unaligned Coculture 36.7 ± 29.2 32.0 ± 23.9 31.3 ± 5.2 71.5 ± 12.9
The trend that fiber alignment and coculture with MEFs improved the differentiation of
mESCDCs was further supported by measuring cell dimensions. Qualitative cell dimensions of
mESCDCs alone and in coculture on the PU scaffolds were measured using image analysis
software and the α-actinin expressing immunostained samples (Table 6.2). Typical mESCDCs
range in size from neonatal (diameter ~7-9 µm and length ~20-45 µm) to adult dimensions
(diameter ~10-30 µm and length ~80-150 µm) [41] and was similarly observed with the
mESCDCs on PU scaffolds. The mESCDCs alone ranged from approximately 30-85 µm and
20-60 µm in length, 7-21 µm and 8-21 µm in diameter, and had an average aspect ratio
(length/diameter ratio) of 4.5 ± 1.4 and 2.8 ± 0.8 and area (length*diameter) of 540 ± 201 µm2
and 604 ± 176 µm2 on aligned and unaligned scaffolds respectively. In coculture, the mESCDCs
were approximately 40-100 µm and 30-75 µm in length, 9-18 µm and 9-24 µm in diameter, and
had an average aspect ratio of 4.7 ± 1.4 and 3.3 ± 0.9 and area (length*diameter) of 717 ± 237
µm2 and 818 ± 198 µm2 for the aligned and unaligned scaffolds respectively. The standard
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average cell area resulted in a similar area as that calculated from length and diameter
measurements. Importantly, the aligned architecture led to a more elongated cell morphology
with a higher length and aspect ratio, more similar to adult cardiomyocytes (aspect ratio of 5.3
for adult and 2.9 for neonatal cardiomyocytes [41]), than the unaligned architecture for both
culture groups. In addition, coculturing the mESCDCs with MEFs increased cell length on both
architectures and led to a higher aspect ratio on the unaligned scaffold compared to mESCDCs
alone.
Table 6.2: Assessment of mESCDC dimensions. Cell dimensions were measured from at least 200 cells from 6 to 9 different immunostained images. Values given as mean ± standard deviation.
Architecture Length (range,
µm)
Average Length (µm)
Diameter (range,
µm)
Average Diameter
(µm)
Aspect Ratio (L/D)
Area (L*D, µm2)
Standard Area (µm2)
Aligned mESCDCs
alone ~30-85 49 ± 11 ~7-21 11 ± 2 4.5 ± 1.4 540 ±
201 583 ± 112
Unaligned mESCDCs
alone ~20-60 41 ± 9 ~8-21 15 ± 3 2.8 ± 0.8 604 ±
176 706 ± 254
Aligned Coculture ~30-100 58 ± 14 ~8-18 12 ± 2 4.7 ± 1.4 717 ±
237 724 ± 185
Unaligned Coculture ~30-75 52 ± 10 ~9-24 16 ± 3 3.3 ± 0.9 818 ±
198 702 ± 188
Assessments of cell shape, sarcomere formation and cell dimensions suggested that both
fiber alignment and coculturing with MEFs improved mESCDC differentiation and maturity
compared to cells on unaligned scaffolds or mESCDCs cultured alone. This is consistent with
previous work in the literature using primary neonatal cardiac cells. We recently demonstrated
that culturing neonatal cardiac cells on the aligned PU scaffolds led to a more mature ventricular-
like phenotype compared to the same cells on unaligned PU scaffolds or TCPS controls [6]. In
addition, recent studies have shown that pre-seeding cardiac fibroblasts alone or with endothelial
cells improved cardiomyocyte elongation, viability, compaction, and functional properties
compared to simultaneous seeding of cells or enriched cardiomyocyte cultures [11, 38]. Thus,
scaffold architecture and culturing with non-myocytes affects both primary cardiomyocyte and
mESCDC phenotypic maturity. Future work should be conducted to assess the mESCDC-based
constructs by ultrastructural analysis using transmission electron microscopy to confirm the
results obtained here.
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Initial results from Live/Dead® staining suggested that cells on aligned scaffolds were
elongated and organized parallel to PU fibers whereas cells on unaligned scaffolds showed no
regular organization. This trend was much more pronounced by immunostaining the cytoskeletal
and sarcomeric structures and further indicated that fiber alignment had a major influence on the
organization of cells. Using auto-fluorescing PU fibers as a reference, all cell types on the
aligned scaffolds had the major cell axis oriented parallel to the PU fibers while no general trend
of cellular organization appeared to occur with cells on the unaligned scaffolds. Measuring the
angle of cell axis for mESCDCs alone (Figure 6.11a) and when cocultured with MEFs (Figure
6.11b) on the two PU architectures confirmed the qualitative assessment regarding cellular
organization. In both culture groups, the majority of the cells on the aligned PU were within 20˚
of the reference angle while cells on the unaligned scaffold were found at nearly all orientations.
Using the absolute value of the angle of cell axis, the average cell angle was calculated for the
mESCDCs alone and coculture on PU scaffolds and is shown in Table 6.3. A significant
difference in average angle of cell axis was observed between mESCDCs on aligned and
unaligned scaffolds both on their own and in coculture (ANOVA, p<0.05) suggesting that fiber
architecture significantly influences cell organization. This was similarly confirmed by
determining the orientation index (Table 6.3), where cells on the aligned scaffolds were
significantly higher, and therefore more aligned, than cells on unaligned scaffolds (ANOVA,
p<0.05). This observation was also made with neonatal primary cardiac cells cultured on the
aligned and unaligned Phe PU scaffolds [6]. The parallel anisotropic organization of mESCDCs
observed here with the aligned constructs is more similar to the organization of cardiomyocytes
found in native cardiac tissue than on the unaligned scaffolds [42]. Anisotropic cardiomyocyte
organization is critical to myocardial function and is increasingly being identified as an
important criterion for generating engineered cardiac tissue [6, 43-45]. Interestingly, a
significant difference in absolute cell axis angle was also observed between mESCDCs alone and
in coculture on unaligned PU scaffolds (p<0.05), with the coculture group showing more
alignment than the mESCDCs alone. This may suggest that in the absence of physical alignment
cues from the underlying scaffold, the presence of MEFs promoted some degree of alignment of
the mESCDCs. Nichol et al. [8] found that coculturing primary neonatal cardiomyocytes with
cardiac fibroblasts induces cardiomyocyte alignment via a MMP-2 mediated mechanism and
limits apoptosis compared to enriched cardiomyocyte cultures. The improved cardiomyocyte
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Figure 6.11: Quantifying the alignment of a) mESCDCs alone and b) coculture of mESCDCs and MEFs on PU
scaffolds. The majority of cells on aligned scaffolds were within 20º of reference angle consistent with the alignment of the underlying PU fibers. Cells on the unaligned scaffolds exhibited no general organization. Angles
were measured from at least 200 cells taken from 6 to 9 different images.
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elongation, compaction, electrical connectivity and contractile function that has been observed
with pretreatment of scaffolds with non-myocytes [11, 38] may also play a role in promoting
mESCDC alignment.
Table 6.3: Average angle of cell axis and orientation index (mean ± SD). Cell angles were measured from at least 200 cells from 6 to 9 different images. Orientation index was determined from 6 to 9 different images. a Significant
difference for same culture group on different architecture. b Significant difference for different culture group on same scaffold architecture (ANOVA, p<0.05).
Architecture & Cell Type Average Angle of Cell Axis Orientation Index
Aligned mESCDCs 9.4 ± 10.1a 0.63 ± 0.04a
Unaligned mESCDCs 37.5 ± 24.7a,b 0.26 ± 0.12a
Aligned Coculture 9.0 ± 7.4a 0.65 ± 0.06a
Unaligned Coculture 29.1 ± 22.5a,b 0.30 ± 0.11a
The expression of the gap junctional protein connexin-43 was identified by
immunostaining the cell-seeded constructs. Figure 6.12 presents immunostained images of Cx-
43 (green), the cytoskeleton (f-actin; red), and nuclei (blue) for the coculture cell group on the
PU matrices 6 days post seeding. The results from these images indicated that the cells on both
PU architectures were expressing gap junctional proteins. The expression pattern appeared
random and may be observed both intracellularly, with newly synthesized Cx-43, and also
around the cell periphery. The gap junction expression was not limited to lateral ends of cells as
observed with mature intercalated discs but rather dispersed around the cell and was similar to
previous work with neonatal cardiac cells on the PU scaffolds [6]. The expression of Cx-43 is
critical to electrical connectivity of cardiac cells and these results suggested that the cells showed
some electrical coupling. This was further supported by synchronous beating of the high density
patches of mESCDCs as observed by light microscopy. Higher Cx-43 expression and end-to-end
patterning may be observed at later time points and should be characterized for periods longer
than one week on the scaffolds. The expression of the gap junctional proteins may have
important implications for functionality of engineered tissue and may further elucidate the
differences between aligned and unaligned cardiac constructs.
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Figure 6.12: Gap junction staining of mESCDCs and MEFs in coculture on a) aligned and b) unaligned PU
scaffolds. Cells on both scaffold architectures expressed gap junctional proteins indicating cell-cell contacts and potential electrical connectivity. n=3, N=1.
The mESCDCs cultured alone or with MEFs were contracting on the PU scaffolds
indicating a contractile phenotype was retained on the PU substrates. Moreover, the mESCDCs
were contracting synchronously and with enough force to observe PU scaffold movement at the
biomaterial edges. The contracting constructs were captured by video microscopy indicating an
important functional component of the engineered cardiac tissue. This was consistent with
previous findings that suggested the Phe PU may have appropriate elastic mechanical properties
to allow cell contraction while being attached to PU surfaces [3-6]. PU movement associated
with cell contraction was initially observed on day 4 post cardiomyocyte seeding and continued
out to at least day 28. Engler et al. [46] determined that cardiomyocytes cultured on surfaces
with an elastic modulus that mimics native cardiac tissue (E ~ 10-15 kPa) is optimal for staining
long-term rhythmic contractions whereas cells on hard matrices (E ~ 35-70 kPa) stop beating
after a few days in culture. Although the aligned and unaligned PU scaffolds had an initial
modulus much greater than 35-70 kPa and were therefore stiffer surfaces, the cells remained
contractile for at least 4 weeks. The mechanical properties of the scaffolds were not measured in
a hydrated state nor were individual PU fibers measured, so it is not known what modulus the
cells experience. The ability of the PU scaffolds to maintain cellular contraction for several
weeks while being attached to the PU substrate provides additional evidence of the potential of
the Phe PU for cardiac applications. In addition, scaffold movement became higher over time
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suggesting increased strength of cell contraction and cell maturity with increasing time in
culture. No observable differences were made between aligned and unaligned scaffolds, but the
anisotropic organization of cells on the aligned scaffold would suggest cells were pulling on the
PU in one direction. There was no apparent difference in contraction between mESCDCs
cultured alone or with MEFs. The presence of high density patches of cardiac cells made it
difficult to quantify construct “beating” because a higher degree of PU movement may not
necessarily correlate to higher strength of contraction but rather a higher number of cells at the
edge of scaffolds contributing to the movement between samples and groups. Achieving a
confluent layer of cells on scaffolds would allow quantification of beating and may help to better
identify how fiber alignment affected contractile properties of the mESCDCs and the
contribution coculturing mESCDCs with MEFs had on functional properties.
6.3.4 Aligned and Unaligned PU Scaffolds for Cardiac Tissue Engineering Topographical cues provided by surface substrates have been known to influence cell
behavior through alterations in cell orientation, migration, proliferation, secretion of ECM, and
cytoskeletal arrangements [47]. A wide variety of structural features, including grooves, ridges,
steps, pores, wells, nodes, and adsorbed protein fibers, have been tested on different substrates
with several cell types in an attempt to control cellular behavior [47]. In the context of tissue
engineering scaffolds, 3-D structural features will similarly affect the cellular response to the
scaffolds. Much work has therefore been conducted in fabricating scaffolds that mimic the
native ECM of specific tissues in order to induce a favorable response of cells. Electrospinning
has emerged as an important technique in tissue engineering due to the ability to form polymer
fibers on the size scale of native ECM proteins [37, 48]. Cells attach and stretch along
electrospun fibers suggesting that the physical cues provided by the fibers could be used to direct
cell orientation and promote the alignment of cardiac cells [37, 48]. Anisotropic organization is
critical to the structure and function of native myocardium and is an important criterion of
cardiac tissue engineering. Electrospun scaffolds with fiber alignment is one technique for
achieving anisotropic organization of cardiac cells and is promising in developing myocardial
tissue constructs.
The results presented here provided evidence of the potential of developing a cardiac
patch using PU scaffolds and mESCDCs. A higher density of mESCDCs was cultured on the
PU scaffolds than previous work in our lab and a distinct cellular organization was achieved by
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aligning PU fibers in the underlying substrate. This system identified that fiber alignment and
coculture with fibroblasts improved the differentiation, maturity, and organization of the
mESCDCs on the PU constructs but more work is necessary to optimize and characterize the
constructs to fully understand the influence these factors have on cell behavior and engineered
tissue function. There are several optimization and characterization methods that should be
carried out in the future. First, the seeding protocol should be adjusted to achieve a high density
of cardiomyocytes uniformly distributed on the electrospun substrates as opposed to having high
density patches. This would promote electromechanical coupling of cells critical to tissue
function and cell-cell signaling that may influence tissue formation. A Percoll gradient could be
used to isolate only the viable cardiomyocytes thereby limiting any adverse signaling from the
presence of dead cells during seeding. Perfusion or other dynamic seeding method has been
shown to improve the distribution of cardiac cells in polymeric matrices [49, 50] and may be
another method for obtaining a high density of uniformly distributed cells.
Second, constructs should be characterized by ultrastructural analysis to verify cell
dimensions and mESCDC differentiation and maturity. Cell size, sarcomere organization and
intercalated disc formation identified by transmission electron microscopy may be used as
markers of mESCDC differentiation and maturity [41]. Immunostaining suggested the
mESCDCs exhibited a higher level of differentiation and organization on aligned scaffolds and
in coculture with MEFs, but this analysis should be confirmed by other methods.
Third, cell maturation should be tested using quantitative polymerase chain reaction to
look at differences in gene expression of atrial natriuretic factor, the α- and β-myosin heavy
chain isoforms, and skeletal and cardiac troponin I. Early mESCDCs express slow skeletal
muscle troponin I and a greater proportion of β-myosin heavy chain [51-53]. More mature
mESCDCs shift from these fetal protein isoforms to cardiac troponin I and α-myosin heavy chain
characteristic of more mature neonatal and adult cardiomyocytes [51-53]. In addition, it was
found that primary neonatal cardiac cells cultured on aligned PU scaffolds had a significant
decrease in ANF expression compared to cardiac cells on unaligned PU or TCPS [6]. A lower
expression of ANF is associated with maturation of ventricular myocardium suggesting the
anisotropic organization of cardiac cells led to more mature ventricular-like engineered tissue
than isotropic cellular organization. Alignment of mESCDCs may similarly lead to more mature
and highly differentiated cells compared to cells with no general organization.
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Fourth, electrical signal propagation should be determined by optical mapping to
investigate electrical coupling of cells and synchronous contractions on the PUs. Different
cellular organizations may lead to significant differences in how electrical signals are propagated
through the constructs. In addition, optical mapping will identify if the high density patches of
mESCDCs are electrically connected and if the presence of MEFs influences electrical
connectivity of the cells. Other functional components of the constructs, such as excitation
threshold and maximum capture rate may provide additional evidence for differences in
construct functionality.
Lastly, the cell population needs to be characterized to identify the number of cells on the
scaffolds and ratio of mESCDCs and MEFs in the coculture cell group. Emerging evidence
suggests that the ratio of cardiomyocytes to fibroblasts and other cardiac cells is an important
consideration for developing cardiac constructs [54]. Understanding the number of cells on the
Phe PUs and ratio of each cell type will help to determine the success of the seeding protocol and
suggest changes to this procedure to optimize cell ratios that have shown to improve functional
characteristics of engineered tissue.
The work described here suggests the electrospun Phe PU scaffolds have potential in the
development of a cardiac patch. More importantly, this system may be used as a model for
testing biomaterial scaffold properties and better defining which of these properties are important
and required for cardiac tissue engineering. Numerous PU scaffold properties may be adjusted
to systematically investigate their influence on cell behavior and tissue function. In addition to
defining scaffold properties for cardiac tissue engineering, systematically changing properties of
aligned and unaligned PU scaffolds could help investigate if thermoplastic polymeric scaffolds
can be used to drive the differentiation of ESCs to the cardiac lineage. Kraehenbuehl et al. [55]
systematically varied the modulus, cell adhesion ligands, and other parameters of 3-D cell-
responsive hydrogels to differentiate mESCs into cardiomyocytes. A similar system may be
used with the PUs and mESCs to directly differentiate and select cells on the biomaterials.
Similarly, the Phe PU matrices could be seeded with cardiac progenitor cells and allow
differentiation to occur directly on the scaffolds. Kattman et al. [30] identified a population of
cardiac progenitor cells that have the potential to differentiate into cardiomyocytes, smooth
muscle cells and endothelial cells. Directly using these progenitors is an alternative to seeding
several cell types separately while still achieving a mixed population of myocardial cells that are
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important for cardiac tissue function. Ultimately, if further studies continue to identify PU
scaffolds may be used in cardiac tissue engineering applications, clinically relevant cardiac
constructs may be developed by seeding PUs with human ESCDCs or induced pluripotent stem
cell-derived cardiomyocytes. This will be an important step in the field of cardiac tissue
engineering and allow the development of tissue constructs for regenerating the myocardium.
6.4 Conclusions Polyurethanes are promising biomaterials for soft tissue engineering applications. Phe
and Gly-Leu PUs were formed into scaffolds that have several similar physical, chemical, and
thermal properties. A high density of MEFs was seeded onto the PU substrates and several
analytical techniques were used to characterize the PU constructs. The results from this study
indicated that the scaffolds could support a high density of viable cells out to at least 28 days and
suggested that the Phe and Gly-Leu PUs meet some fundamental requirements of tissue
engineering scaffolds. For cardiac applications, the Phe PU was electrospun into scaffolds with
varying architecture. Embryonic stem cells were differentiated in a spinner flask system to
generate a large and pure population of cardiomyocytes. The mESCDCs were seeded onto the
aligned and unaligned Phe PUs alone or onto the PUs pre-seeded with MEFs. Viable cells
attached to both architectures and the substrates supported a contractile phenotype. Importantly,
the aligned scaffolds led to the anisotropic organization of rod-shaped cells, improved sarcomere
organization, and increased mESCDC aspect ratio when compared to cells on the unaligned
scaffolds. In addition, pre-seeding the scaffolds with MEFs improved sarcomere formation,
increased cell alignment and aspect ratio, and led to a mESCDC morphology that was more
extended on the PU scaffolds than the mESCDCs cultured alone. These results suggest that both
fiber alignment and pre-treatment of scaffolds with fibroblasts improves the differentiation and
organization of mESCDCs and are important parameters for developing engineered myocardial
tissue constructs using ESC-derived cardiac cells and PU scaffolds.
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Chapter 7: Conclusions and Future Work
7.0 Conclusions The current options for treating myocardial infarctions and congestive heart failure,
including medical therapy, ventricular assist devices, and heart transplantation, do not offer a
long term solution to an increasing number of patients with heart disease. Cellular
cardiomyoplasty has emerged as a method of introducing cells into the heart to improve cardiac
function but to date, clinical trials using skeletal myoblasts and bone marrow-derived cells have
produced mixed results [1-12]. Cellular cardiomyoplasty has been limited by the delivery,
engraftment, and survival of a sufficient number of cells into the heart [1, 13, 14]. This may be
improved through the use of the appropriate delivery vehicle, such as biomaterial scaffolds.
Cardiac tissue engineering employs the use of biomaterials in combination with cells to
develop viable tissue constructs to improve or regenerate injured or diseased myocardium.
Significant achievements have been made in the last several years towards developing cardiac
tissue in vitro but scaffold requirements for this application have not been well defined.
Biodegradable segmented polyurethanes (PUs) are excellent biomaterials for research in the
cardiac tissue engineering field due to its flexible chemistry. This provides an opportunity to
develop new biomaterials that incorporate specific chemical moieties that confer unique
biological functionality to the synthetic polymers.
The work conducted in this thesis involved the development and testing of new
biodegradable PU scaffolds to better identify important scaffold properties for in vitro cardiac
tissue formation. This was broken up into two closely related parts. The first involved the
synthesis, characterization, processing, and in vitro testing of a segmented PU that incorporates
the Gly-Leu cleavage site of several matrix metalloproteinases (Gly-Leu PU). Most synthetic
polymers degrade predominantly by passive hydrolysis, so the development of an MMP-
sensitive segmented PU may help to tailor the degradation properties to the cardiac environment.
The results presented in Chapter 3 demonstrated the successful synthesis and characterization of
the Gly-Leu PU using a previously developed chymotrypsin-sensitive phenylalanine-based PU
(Phe PU) as a comparison. A Gly-Leu-based diester chain extender was developed as a
minimalistic means of incorporating the Gly-Leu cleavage site. This allowed flexibility in the
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choice of the soft segment which in turn may be used to obtain diverse polymer properties with
this family of PUs. This also allowed a direct comparison to the Phe PUs to investigate how
changing the chain extender chemistry from the phenylalanine amino acid to the Gly-Leu
dipeptide affected polymer properties. Following successful synthesis, purification and
characterization of the Gly-Leu-based diester chain extender, the Gly-Leu PU was synthesized
using a soft segment composed of polycaprolactone diol of molecular weight 1250. This soft
segment was chosen because the Phe PUs incorporating this have been very promising in cardiac
applications [15-18]. As may have been predicted by the chemistry, the Gly-Leu PU had several
properties that were very similar to the Phe PU. Both PUs had high molecular weight averages,
were phase segregated, semi-crystalline polymers, and were soft, flexible elastomers with high
breaking stresses and strains. Interestingly, despite similarities in stress-strain curves, the Gly-
Leu PU had a significantly higher initial modulus, yield stress and ultimate stress compared to
the Phe PU.
Once the Gly-Leu PU had been synthesized and characterized, it was processed into
porous, 3-D scaffolds for further characterization and testing. Electrospinning was used to
process the PUs because it forms fibrous matrices that meet several criteria of biomaterial
scaffolds including; a high surface area-to-volume ratio, structural features for promoting cell
adhesion, growth and differentiation, and an architecture to help organize cells. In Chapter 4, it
was demonstrated that both the Phe and Gly-Leu PUs could be formed into scaffolds with a
variety of structural features, including beads, beads-on-a-string, and defect-free fibers of
varying diameters, by adjusting the polymer solution concentration during electrospinning. The
structural features obtained were similar to those found with other synthetic polymers in the
electrospinning literature [19]. The Phe and Gly-Leu PU scaffolds fabricated from a 14% and
10% w/v concentration respectively were chosen for subsequent studies due to their similarities
in structural features. Both scaffolds had a randomly organized fiber structure, an average fiber
diameter of approximately 3.6 µm, and similar fiber diameter distributions. In addition, the
electrospinning process did not affect the molecular weight averages or thermal properties of
either PU suggesting the electrospun PUs retained several important properties of the base
material.
The remaining experiments for this part of the thesis were conducted to test the
performance of the electrospun Phe and Gly-Leu PU scaffolds for tissue engineering
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applications. This was carried out by characterizing the enzyme-mediated and passive
hydrolysis of the two PU scaffolds and by assessing the viability of cells cultured on the two PUs
in vitro. In the degradation study in Chapter 5, changes in mass and structural features of the PU
scaffolds following incubation in either MMP-1, MMP-9, or buffer solutions were investigated.
Neither the Phe nor Gly-Leu PU scaffolds exhibited any detectable passive hydrolysis or MMP-
mediated chain cleavage over the 28 day experiment. This result does not support the original
hypothesis and provides evidence against the rational for developing the Gly-Leu PU. Several
mechanisms may account for the Gly-Leu PU stability but the minimalistic approach of
incorporating the Gly-Leu dipeptide alone without any flanking amino acid sequences along with
the phase segregated nature of the PUs is suspected to have the most influence. A segmented PU
was recently developed that incorporates a Pro-Ala-Gly-Leu-Lys sequence in the chain extender
chemistry and exhibits enhanced degradation in the presence of collagenase [20]. This supports
the need for a longer peptide sequence to achieve MMP-susceptibility.
Despite the degradation results, the cell-based studies with the Gly-Leu PU were more
promising. The experiments in Chapter 6 using the Gly-Leu PU were conducted to test whether
the PU scaffolds could support the attachment of cells and whether the material or its
degradation products were cytotoxic. These fundamental criteria of biomaterial scaffolds were
assessed by seeding mouse embryonic fibroblasts (MEFs) onto the PUs at a high seeding density
and evaluating cell viability and morphology during a 28 day culture period. MEFs were
specifically chosen as the cell type for these experiments because fibroblasts are involved in
tissue remodeling through the secretion of matrix metalloproteinases. As well, fibroblasts play
an important role in the development of engineered myocardial tissue constructs [21-24].
Therefore, establishing conditions to culture MEFs on the PU scaffolds provided a starting point
for subsequent coculture studies with murine embryonic stem cell-derived cardiomyocytes
(mESCDCs). Characterizing the MEFs on the Phe and Gly-Leu PU scaffolds demonstrated that
both PUs could support a high density of viable cells out to at least 28 days. Cells were adherent
and spread out with no regular organization on the randomly oriented substrates. These
encouraging results suggested the scaffolds meet some fundamental requirements for tissue
engineering applications. Further work is required to fully characterize the Gly-Leu PUs,
particularly testing the degradation properties using in vitro cell-based and in vivo animal
models. This may help to identify whether the Gly-Leu PUs may be used for cardiac
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applications or if further modifications to the chemistry are needed to tailor the PUs to this
unique environment.
While the first part of this thesis was motivated by the degradation characteristics of the
PUs, the second part focused on physical properties of elastomeric PUs and their influence on
cardiac cells. The anisotropic organization of cardiac cells is critical to the structure and function
of the native myocardium and achieving this is an important goal in engineering myocardial
tissue constructs. The electrospinning process allows control over parameters such as the size
and organization of the fibrous scaffold structure. Therefore, the conditions for achieving PU
scaffolds with varying architecture were explored in Chapter 4 as a means of investigating how
fiber alignment influences the organizational response of mESCDCs and MEFs. Consistent with
previous reports using other polymers [25, 26], the Phe PU was electrospun into scaffolds with
aligned and unaligned architectures by adjusting the rotational speed of the collection mandrel in
the electrospinning system. A high mandrel rotational speed was used to achieve fiber alignment
whereas a low rotational speed was used to obtain randomly organized fibers. Electrospinning
the PU at different mandrel rotational speeds did not influence the molecular weight averages or
thermal properties of the resulting scaffolds but significantly affected their mechanical
properties. Specifically, the unaligned PU scaffold exhibited similar mechanical behavior when
stretched in the two perpendicular directions with a relatively low initial modulus and ultimate
tensile stress and high ultimate elongation. The aligned PU scaffold was mechanically
anisotropic with a relatively high initial modulus and ultimate tensile stress and low elongation
when stretched in the preferred direction of fiber orientation.
Most studies in the cardiac tissue engineering field have been conducted using primary
cardiac cells as a proof-of-principle in establishing the potential of various techniques and
parameters needed for developing engineered myocardial tissue. Work performed with our
collaborators at the University of Delaware demonstrated that culturing primary cardiac cells on
the aligned Phe PU scaffold significantly influenced the phenotype of the cells when compared
to either the unaligned PU scaffold or a tissue culture polystyrene control [18]. Specifically, we
found that the aligned scaffolds provided physical cues for the anisotropic organization of
cardiac cells that resembled native cardiac tissue. In addition, the aligned cardiac constructs
were associated with a decrease in atrial natriuretic peptide compared to unaligned constructs
suggesting a more mature, ventricular-like cardiac phenotype [18].
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Embryonic stem cell-derived cardiomyocytes are currently the best source of de novo
cardiomyocytes and are therefore the best candidates for achieving true regeneration of
myocardial tissue. To be used in cardiac tissue engineering, scaffolds that promote ESCDC
differentiation and functionality must be identified. Relatively few studies have been conducted
using ESCDCs on pre-formed scaffolds and even less work has looked at the coculture of
ESCDCs with other cells found in the myocardium on 3-D scaffolds. Thus, the cell experiments
carried out in Chapter 6 using the aligned and unaligned electrospun scaffolds were designed to
test the response of mESCDCs alone and in coculture with MEFs on the two scaffold
architectures. Murine ESCs were differentiated in a spinner flask system to generate a large and
pure population of cardiomyocytes. The mESCDCs were seeded onto the aligned and unaligned
Phe PUs on their own or pre-seeded with MEFs. In both culture conditions, viable mESCDCs
attached to the PU scaffolds and were functionally contractile. Importantly, the aligned scaffolds
led to the anisotropic organization of rod-shaped cells, improved sarcomere organization, and
increased mESCDC aspect ratio when compared to cells on the unaligned scaffolds. In addition,
pre-seeding the scaffolds with MEFs improved the differentiated phenotype of the mESCDC as
observed by an increase in sarcomere formation, elongation, alignment, and aspect ratio of
mESCDCs compared to this cell type alone. These results build on previous work in our
laboratory that showed unaligned electrospun PU scaffolds supported contractile mESCDCs with
a striated sarcomeric phenotype and appeared more differentiated than cells on scaffolds formed
by thermally induced phase separation [16]. The current work showed that both fiber alignment
and coculturing the mESCDCs with MEFs further improved the organization and maturity of the
differentiated cells. This suggests that an aligned scaffold architecture and coculture with
fibroblasts are important considerations for the formation of cardiac constructs using PU
scaffolds and mESCDCs. Future work should further characterize the cell-seeded PU constructs
to determine if the organization and mature differentiated phenotype associated with the aligned
cocultured constructs has an influence on the functional behavior of the engineered tissue. If so,
this aligned coculture system may be used as a platform to better define critical scaffold
properties for cardiac tissue engineering by altering the mechanical, degradation, and structural
characteristics of the PU scaffolds and testing how each affects the functional properties of
cardiac constructs.
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7.1 Significant Contributions to Literature • Designed, synthesized, characterized, processed, and tested a new family of elastomeric
biodegradable PUs using a Gly-Leu-based diester chain extender. This new family of
PUs had several properties that were promising for soft tissue engineering applications.
• Developed a system for coculturing mESCDCs with MEFs and demonstrated that both
scaffold architecture and coculture with fibroblasts improved cell orientation,
morphology, and differentiated phenotype. A higher cell density, organization and
differentiated phenotype were achieved with the mESCDCs than previous work in our
laboratory and in the literature.
7.2 Future Work A discussion of future work including rationale and potential methods for conducting the
experiments was presented towards the end of each experimental chapter where the
corresponding work applies. These points are briefly highlighted again here.
7.2.1 Polyurethane Design and Synthesis • Adjust the soft segment chemistry of the Gly-Leu PU to PCL of different molecular
weights or tri-block soft segments of PCL-PEO-PCL of varying molecular weight to
expand the available polymer properties with the Gly-Leu-containing family of PUs. Try
to obtain a PU with mechanical properties that more closely mimics the mechanical
properties of the native myocardium.
• Design peptide-based chain extenders with varying sequences to produce an array of
enzyme-sensitive synthetic elastomers with different degradation rates. Scaffolds formed
by blending the different PUs or co-electrospinning the polymers may help in developing
scaffolds with degradation rates that will allow the gradual transfer of mechanical load
from the PU to the newly secreted ECM. This may help in forming biomimetic materials
with appropriate degradation properties for mechanically active soft tissues.
7.2.2 PU Scaffold Formation and Characterization • Systematically adjust electrospinning parameters to achieve nanofibers with Phe and Gly-
Leu PUs and characterize scaffold properties to further contribute to knowledge of how
fiber size influences scaffold properties and performance.
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• Test the mechanical properties of individual fibers to understand what cells will
experience. Compare the mechanical properties of individual fibers of different sizes and
PUs synthesized with different soft segments. Investigate the response of cardiac cells to
PU fibers of different size and elastic modulus.
• Determine the mechanical properties of Gly-Leu PU scaffolds and perform all testing in a
hydrated state.
• Characterize the dynamic mechanical properties of PU scaffolds along with mechanical
stimulation of cell-seeded constructs.
• Implant the PU scaffolds in a normal and infarcted heart to determine if scaffolds have
appropriate biocompatibility properties for this application.
7.2.3 PU Degradation • Test MMP adsorption on the PU scaffolds
• Carry enzyme-based degradation experiments out to longer time points and with higher
concentration of MMPs.
• Validate the competitive substrate enzyme inhibition assay with the peptide Ala-Pro-Gly-
Leu.
• Synthesize a short polymer chain using the Gly-Leu-based chain extender and look at
cleavage of Gly-Leu dipeptide by MMPs.
• Conduct cell-based PU degradation studies with C14-labelled PUs and MEFs or
inflammatory cells. Test the susceptibility of the Gly-Leu PU to MMP-mediated
degradation by measuring activity of MMPs and inhibiting MMPs while characterizing
the release of degradation products.
• Perform in vivo degradation studies to identify residence times and foreign body reaction
to Phe and Gly-Leu PUs in an infarcted rat heart.
7.2.4 Cell-based Testing of PU Scaffolds • Optimize mESCDC seeding to obtain a high and uniform density of cells on the aligned
and unaligned PUs. This may be achieved by removing dead cells and using a perfusion
seeding technique.
• Use ultrastructural analysis to characterize cardiomyocyte size, sarcomere organization,
and intercalated disc formation.
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• Determine gene expression by quantitative polymerase chain reaction to identify protein
isoforms and markers for cell maturation. Some of the genes to test include atrial
natriuretic factor, the α- and β-myosin heavy chain isoforms, and skeletal and cardiac
troponin I. Test the expression of the different markers with mESCDCs alone and in
coculture with MEFs to identify what influence the presence of fibroblasts had on
cardiomyocyte gene expression.
• Investigate functional properties of aligned and unaligned cardiac constructs. Optical
mapping should be used to look at action potential propagation and electrical connectivity
throughout the constructs. Contractile properties may be determined by measuring
excitation threshold and maximum capture rate.
• Characterize the ratio of mESCDCs and MEFs in coculture on PU matrices and optimize
this ratio as needed.
• Systematically adjust properties of aligned PU scaffolds and test functional behavior of
cardiac constructs. This may help to better define requirements of biomaterial scaffolds
for cardiac tissue engineering.
• Implant PU scaffolds seeded with mESCDCs and MEFs into infarcted mouse hearts as a
model system for investigating coupling with host myocardium, improvements to cardiac
function, and potential regeneration of myocardial tissue.
• Differentiate mESCs on aligned and unaligned PU scaffolds with varying mechanical
properties and fiber diameters and investigate the ability of the scaffolds to drive
differentiation towards the cardiac lineage.
• Culture cardiac progenitor cells on PUs and allow differentiation to occur as a method for
achieving different myocardial cells in one seeding procedure. Identify if the scaffolds
direct differentiation towards a particular cell type and alter scaffold properties to test if
this may lead to different ratio of cell types.
• Test the ability to form cardiac constructs using PU scaffolds and human ESCDCs or
iPSC-derived cardiomyocytes.
• Transplant skeletal myoblasts or bone marrow-derived cells on PU scaffolds into
infarcted animal myocardium to determine if the PU scaffolds improve engraftment and
cardiac function to a greater degree than either cells or scaffolds alone.
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Appendix A – Supplementary Information for Dipeptide-based Chain Extender Characterization
A.1 C13 NMR Spectra of Reactants, Theoretical Predictions, and Raw Products
C13 NMR on the reactants, theoretical Gly-Leu-based chain extender predictions using
ACD i-Lab software, and raw products were used to aid in assigning chemical shifts in the
purified chain extender and assess successful purification. Below are the spectra used as
references to assign peaks to the corresponding chemical structure of the chain extender.
Figure A.1: C13 NMR spectrum of Gly-Leu dipeptide
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Figure A.2: C13 NMR spectrum of CDM
Figure A.3: Theoretical predictions of Gly-Leu-based diester chain extender using ACD i-Lab software.
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Figure A.4: C13 NMR spectrum of raw Gly-Leu-CDM-PTSA. The prominent peaks found around 120-140 ppm
correspond to the aromatic carbons of p-toluene sulfonic acid.