RESILIN-LIKE POLYPETIDE-BASED MICROSTRUCTURED HYDROGELS VIA AQUEOUS-BASED LIQUID-LIQUID PHASE SEPARATION FOR
TISSUE ENGINEERING APPLICATIONS
by Hang Kuen Lau
A dissertation submitted to the Faculty of the University of Delaware in partial fulfillment of the requirements for the degree of Doctor of Philosophy in Materials Science and Engineering
Fall 2018
© 2018 Hang Kuen Lau All Rights Reserved
RESILIN-LIKE POLYPETIDE-BASED MICROSTRUCTURED HYDROGELS VIA AQUEOUS-BASED LIQUID-LIQUID PHASE SEPARATION FOR
TISSUE ENGINEERING APPLICATIONS
by Hang Kuen Lau
Approved: __________________________________________________________
Darrin J. Pochan, Ph.D.
Chair of the Department of Materials Science and Engineering
Approved: __________________________________________________________
Babatunde A. Ogunnaike, Ph.D.
Dean of the College of Engineering
Approved: __________________________________________________________
Douglas J. Doren, Ph.D.
Interim Vice Provost for Graduate and Professional Education
I certify that I have read this dissertation and that in my opinion it meets the academic and professional standard required by the University as a dissertation for the degree of Doctor of Philosophy.
Signed: __________________________________________________________
Kristi L. Kiick, Ph.D.
Professor in charge of dissertation
I certify that I have read this dissertation and that in my opinion it meets the academic and professional standard required by the University as a dissertation for the degree of Doctor of Philosophy.
Signed: __________________________________________________________
Xinqiao Jia, Ph.D.
Member of dissertation committee
I certify that I have read this dissertation and that in my opinion it meets the academic and professional standard required by the University as a dissertation for the degree of Doctor of Philosophy.
Signed: __________________________________________________________
Christopher J. Kloxin, Ph.D.
Member of dissertation committee
I certify that I have read this dissertation and that in my opinion it meets the academic and professional standard required by the University as a dissertation for the degree of Doctor of Philosophy.
Signed: __________________________________________________________
Thomas H. Epps, III, Ph.D.
Member of dissertation committee
ACKNOWLEDGMENTS
I would like to thank my advisor, Dr. Kristi Kiick her guidance and support as a mentor, scientist and scholar. I am grateful for her encouraging and guiding me via my time in University of Delaware to improve my research, critical thinking, and writing and presentation skills. This dissertation would not have been possible without her. I also thank my committee members Dr. Thomas Epps from Chemical
Engineering, Dr Xinqiao Jia and Dr. Chris Kloxin from Materials Sciences and Engineering for their support and assistance on my dissertation. I am also thankful for our collaborators Dr. Sapun Parekh from Max Planck Institute for Polymer Research, Dr. Alexandra Paul and Dr. Annika Enejder from Chalmers University of Technology for their contribution on CARS characterization on microstructued hydrogels
micrcomposition. I would also like to thank Dr. Al Crosby, Shruti Rattan, Hongbo Fu and Dylan Barber from the University of Massachusetts Amherst for help and
contribution in micromechanical characterization on microstructued hydrogels. I also thank Dr. Susan Thibeault and Renee King from University of Wisconsin, Madison for their help and contribution on in vivo animal vocal fold injections.
I would like to thank Dr. Linqing Li and Dr. Chris McGann for training and guiding me to start on the resilin-like polypeptides projects. Thanks also to Dr. Becca Scott and Ishnoor Sidhu for help in cell culture experiments. I would also like to thank, from the Delaware Biotechnology Institute Bioimaging Center, Dr. Jeff Kaplan, Dr. Michael Moore and Sylvain Le Marchand for training and help with
AFM. Thanks to Dr. Shi Bai for help with NMR spectroscopy and Dr. PapaNii Asare- Okai for help and training with mass spectrometry.
I would also like to thank all the current and previous Kiick group members Dr. Linqing Li, Dr. Chris McGann, Dr. Nandita Bhadwat, Dr. Eric Levenson, Dr.
Tianzhi Luo, Dr. Yingkai Liang, Dr. Prathamesh Kharkar, Dr. Morgan Urello, Dr.
Chris Koehler, Dr Becca Scott, Bradford Paik, Ishnoor Sidhu, Haocheng Wu, Yu Tian, Michael Haider, Cristobal Garcia, Jingya Qin, Haofu Huang, Luisa Palmese and Ming Fan. Thanks also to all my friends and colleagues Shuang Liu, Shuyu Xu, Ying Hao, Tugba Ozdemir, Anitha Ravikrishnan, Kevin Dicker, Aidan Zerdoum, Eric Fowler. I would also like to thank the current and previous staff in Department of Materials Science and Engineering Charlies Garbini, Christine Williamson, Naima Hall, Robin Buccos, Kathleen Forwood and Judy Allarey for ensuring lab safety and providing administrative supports.
LIST OF TABLES ... ix
LIST OF FIGURES ... x
ABSTRACT ... xvii
Chapter 1 MULTICOMPONENT HYBRID HYDROGELS IN BIOMEDICAL APPLICATIONS ... 1
1.1 Introduction ... 1
1.2 Hydrogel Network Formation ... 3
1.3 Mimics of Natural Proteins ... 8
1.3.1 Elastin ... 9
1.3.2 Resilin ... 10
1.4 Composite Hydrogels ... 13
1.4.1 Nanocrystal-reinforced Matrices ... 14
1.4.2 Particle-reinforced Matrices ... 15
1.4.3 Fiber-reinforced Matrices ... 18
1.5 Hybrid Materials with Engineered Biological Functions ... 19
1.6 Conclusion and Perspectives ... 24
1.7 Dissertation Summary ... 26
2 AQUEOUS LIQUID-LIQUID PHASE SEPRARATION OF RESILIN- LIKE POLYPEPTIDE/POLYETHYLENE GLYCOL FOR FORMATION OF MICROSTUCTURED HYDROGELS ... 28
2.1 Introduction ... 28
2.2 Materials and Methods ... 31
2.2.1 Materials ... 31
2.2.2 RLP Expression and Purification ... 31
2.2.3 Characterization of RLP/PEG Phase Separation ... 32 TABLE OF CONTENTS
2.2.7 Oscillatory Rheology ... 34
2.2.8 AFM Indentation ... 35
2.3 Results and Discussion ... 36
2.3.1 RLP/PEG Phase Separation ... 36
2.3.2 Co-existence Concentration of RLP/PEG ... 41
2.3.3 Microstructured Hydrogels ... 45
2.3.4 Hydrogel Bulk and Micromechanical Properties ... 48
2.4 Conclusions ... 53
3 MICROSTRUCTURED ELASTOMER-POLYETHYLENE GLYCOL HYDROGELS VIA KINETIC CAPTURE OF AQUEOUS LIQUID- LIQUID PHASE SEPRARATION ... 55
3.1 Introduction ... 55
3.2 Materials and Methods ... 58
3.2.1 Materials ... 58
3.2.2 RLP Expression and Purification ... 59
3.2.3 RLP Functionalization and Characterization ... 59
3.2.4 Characterization of RLP-Ac/PEG-4Ac Phase Separation ... 60
3.2.5 Characterization of Equilibrium Concentrations ... 61
3.2.6 Hydrogel Formation ... 61
3.2.7 Oscillatory Rheology ... 61
3.2.8 Polymerization Yield ... 62
3.2.9 Confocal Microscopy and Domain Diameter Analysis ... 62
3.2.10 BCARS Imaging and Data Analysis ... 63
3.2.11 AFM Indentation ... 64
3.2.12 Swelling Ratio ...65
3.2.13 Cell Encapsulation and Viability ... 65
3.2.14 Statistical Analysis ... 66
3.3 Results and Discussion ... 66
3.3.1 Design and Synthesis of Photo-crosslinkable RLP-Ac ... 66
3.3.2 Liquid-liquid Phase Separation of RLP-Ac/PEG-Ac ... 68
3.3.3 Characterization of Microstructure and Bulk Mechanics of Microstructured Elastomers ... 73
3.3.4 Distinct Composition and Mechanics of Domains in Micostructured Elastomers ... 81
2.2.5 Fluorescent Labeling of RLP and PEG ... 33
2.2.6 Hydrogel Formation ... 34
3.4 Conclusions ... 91
4 MICROMECHANCICAL PROPERTIES OF MICROSTRUCTURED ELASTOMER-POLYETHYLENE GLYCOL HYDROGELS ... 93
4.1 Introduction ... 93
4.2 Materials and Methods ... 94
4.2.1 Hydrogel Formation ... 94
4.2.2 Oscillatory Rheology ... 95
4.2.3 Microindentation ... 96
4.2.4 Puncture ... 97
4.3 Results and Discussion ... 98
4.3.1 Hydrogel Formation and Characterization ... 98
4.3.2 Oscillatory Shear Rheology ... 100
4.3.3 Microindentation and Resilience ... 102
4.3.4 Large-strain Indentation and Puncture ... 104
4.4 Conclusions ... 111
5 CONCLUSIONS AND FUTURE WORK ... 112
5.1 Conclusions and Perspectives ... 112
5.2 Future Work ... 115
5.2.1 Stabilization of Phase Separation of RLP/PEG ... 115
5.2.2 Injectable Resilin-like Polypeptide Hybrid Hydrogels for Vocal Folds ... 118
REFERENCES ... 123
Appendix A COPYRIGHT PERMISSION ... 159
3.3.5 Cell Viability and Growth in 3D LLPS Microstructured Elastomers ... 89
phases of RLP/PEG solutions of various compositions... 43 Table 3.1 Fitting Parameters used on Raman-like spectra. Seed values with
lower/upper bounds. ... 64 Table 3.2 RLP-Ac functionalization ... 68
LIST OF TABLES
Table 2.1 Concentrations of RLP and PEG in the upper (phase I) and lower (phase II)
Figure 1.1 PNIPAAm - coiled-coil peptide - PNIPAAm thermally responsive self- assembled hydrogel. (a) The hydrogel is crosslinked by the coiled-coil structure formed by the polypeptide and by PNIPAAm after its
collapse and aggregation above its LCST. (b) Schematic of PNIPAAm - coiled-coil peptide - PNIPAAm and peptide sequence.82
Reproduced with permission from ref 108. Copyright (2013) WILEY- VCH Verlag GmbH & Co. KGaA. ... 6 Figure 1.2 The PAMPS and PAAm network of the double network hydrogel under
tensile test. The highly crosslinked PAMPS network fractured while loosely crosslinked PAAm network still holding the gel stucture during extension.95 Reproduced with permission from ref 121.
Copyright (2010) The Royal Society of Chemistry. ... 8 Figure 1.3 Resilin-like polypeptide hydrogels demonstrate useful mechanical
properties and biological functions.132 Reproduced with permission from ref 158. Copyright (2012) The Royal Society of Chemistry. ... 12 Figure 1.4 Microgel-reinforced double network PAAm hydrogel that exhibited
excellent extension (a) and torsion (b). Microgel before tensile deformation (c) and after deformation (d).169 Reproduced with permission from ref 195. Copyright (2013) American Chemical
Society. ... 17 Figure 1.5 Important materials design considerations for tissue engineering,
including cell adhesion peptide, protease sensitive peptide for cell- mediated matrix degradation, and presence of signaling molecules. ... 20 Figure 2.1 Solution of PEG and RLP undergo phase separation to yield two
aqueous phases in PBS buffer. A) Liquid-liquid phase separation of 10 wt% 50/50 RLP/PEG in PBS in room temperature and the mixture partitioned into two immiscible aqueous phases. B) UV-vis
absorbance of RLP, and 50/50 RLP/PEG-4NH2-10k C) UV-Vis absorbance of 50/50 RLP/PEG solutions with different molecular weight 4-arm and linear PEGs; total solution concentrations ranged from 15 wt% to 0 wt% (w/v). ... 38
LIST OF FIGURES
Figure 2.2 Phase diagram of 10 wt% RLP/PEG solutions. Phase transition of RLP/PEG with different PEG molecular weight and architecture, the line indicates the phase separation transition from a two-phase to a homogeneous solution. PEG-2NH2-5k (■), PEG-2NH2-10k (●), 4- arm PEG-4NH2-10k (▲) and 4-arm PEG-4NH2-20k (▼)... 40 Figure 2.3 Co-existence concentrations of the RLP/PEG solution characterized via
1H NMR. (A) The 10 wt% 50/50 RLP/PEG formed two immiscible layers upon overnight incubation at 25 oC, with a clear phase I and yellowish phase II observed. (B) RLP and PEG concentrations
characterized by 1H NMR. The x indicates the initial concentration of the mixtures that phase separated into PEG-rich (red) and RLP-rich (blue) phases. The dashed line indicates the tie lines connecting pairs of the PEG-rich and RLP-rich phases. The black line is rendered for visual clarity. ... 43 Figure 2.4 Schematic of the crosslinking reaction and illustration of the phase-
separated hydrogels. The chemical crosslinker THP reacts with primary amines from both RLP lysine residues and PEG amine end groups. The reactions allow both RLP-rich and PEG-rich domains to be crosslinked to provide a stable network and capture the
microstructure during phase separation. ... 46 Figure 2.5 Morphology of the RLP-PEG hydrogels. (A) RLP and PEG
concentration characterized by NMR. The x indicates the initial concentration of the mixtures that phase separated into PEG-rich (red) and RLP-rich (blue) phases. (B) Particle size distribution curves for the hydrogels formed with composition (i) and (ii) from the co- existence curve. The inset shows confocal images of hydrogels produced from solutions of Dylight-594 labeled RLP (red) and
Dylight-488 labeled PEG (green). Scale bar = 50 μm. ... 47 Figure 2.6 Rheology of 10wt% 50/50 RLP/PEG hydrogel. Time sweep of the
10wt% PEG, 10wt% RLP and 10wt% 50/50 RLP/PEG hydrogel crosslink with THP indicated fast gelation time and storage modulus G’ above loss modulus G” of the hydrogel indicated stable gels. The storage modulus reaches its highest value within~15 min, with the storage modulus being significantly greater than the loss modulus, indicating the formation of a solid-like gel. The subsequent reduction of the G’ at times greater than 15 min may be due to the coarsening and rearrangement of the micro-domains, resulting in a reduced interfacial area, but nevertheless, the gels attain a steady modulus after 60 minutes. ... 49
Figure 2.7 Micromechanical differences in hydrogel domains as assessed via AFM.
Optical microscopy images of the RLP-PEG thin gel with the AFM probe located at the RLP-rich domain (A) and PEG-rich matrix (B) scale bar = 50 μm. (C) Micromechanical properties characterized via indentation. The distribution of Young’s moduli from indentation of 10 wt% RLP (red dashed line), 10 wt% PEG (black dashed line), and 10 wt% (w/v) 50/50 RLP-PEG crosslinked with 1:3 THP:amine molar ratio. The phase-separated RLP-rich domains and PEG-rich matrix were visualized via optical microscopy and indented separately; the data were fit to a Gaussian distribution for the PEG-rich matrix (black solid line) in the RLP-PEG and a double-Gaussian distribution for the RLP-rich domains (red solid line). (D) The box plot indicates the statistical distribution of the data, and the asterisk indicates
statistically significant differences between the mean values of the marked samples and all other samples (p < 0.05). ... 51 Figure 2.8 Fitting of AFM micromechanical data for RLP-rich domains. The
Young’s moduli distribution of the RLP-rich domain (black) and fitting of double Guassian (blue) with peak 1 (red) at 5.3 ± 2.3 kPa and peak 2 (green) at 15.6 ± 8.1 kPa. ... 52 Figure 3.1 Acrylamide functionalization of RLPs. A) Schematic of RLP
functionalization. Lysine residues along the polypeptide chain were reacted with an acrylic acid N-hydroxysuccinimide ester through simple amide bond coupling reactions; B) NMR spectrum of RLP-Ac showing the 3 vinylic peaks which increase in intensity with an increase in the NHS-Ac:Lysine ratio; and C) Various degrees of RLP- Ac functionalization achieved with various NHS-Ac:Lysine molar ratios from 0.2 to 4. ... 68 Figure 3.2 Phase separation of 50/50 RLP-XAc/PEG-4Ac in PBS buffer. A) UV-
Vis transmittance of 50/50 RLP-XAc/PEG-4Ac solutions where X was varied between 2 and 10 as a function of total polymer wt%.
RLP-Ac/PEG-4Ac solutions with increasing RLP-Ac/PEG-4Ac ratios with B) RLP-2Ac and C) RLP-6Ac as a function of increasing total polymer wt%. ... 70
Figure 3.3 Phase diagram of RLP-XAc/PEG-4Ac in PBS buffer. A) Coexistence curve for 50/50 RLP-6Ac/PEG-4Ac as determined by 1H NMR. The x indicates the initial concentration of the mixtures before phase
separation. The diamond data represent the phase separation concentrations from UV-Vis data. Final concentrations after phase separation in the PEG-rich and RLP-rich domains are shown as circles and triangles, respectively. The dashed lines connect pairs of the PEG- rich and RLP-rich phases. The black line is rendered for visual clarity only. B) Comparison of concentrations in PEG- and RLP-rich
domains of 10wt% 50/50 RLP-XAc/PEG-4Ac for X=4 and 6. ... 73 Figure 3.4 Temporally controlled microstructured hydrogels. A) Schematic of
hydrogel formation and microstructure development. B) Time sweep of 10 wt% 50/50 RLP-6Ac/PEG-4Ac with UV irradiation at 0, 5 and 10 min after vortex mixing. C) Modulation of hydrogel mechanical properties with variations in the time of irradiation. Data shown are oscillatory rheology time sweeps of 10 wt% 50/50 RLP-6Ac/PEG- 4Ac. All samples were monitored for 10 minutes, but the various samples were irradiated with UV-light for different durations starting at time 0 (e.g., irradiation for 30 sec, 1 min, 2 min, 4 min and 10 min), illustrating the control over mechanical properties that is afforded by these methods. D) Storage moduli comparison for RLP-2Ac and RLP- 6Ac with UV irradiation at 0, 5 and 10 min after mixing. ... 75 Figure 3.5 Phase contrast images of photo-crosslinked RLP and PEG hydrogels. A)
10 wt% PEG-4Ac, B) 10 wt% RLP-2Ac and C) 10 wt% RLP-6Ac hydrogels crosslinked immediately after mixing with UV irradiation for 4 min. The lack of contrast observed in these experiments indicates the absence of microstructure in pure PEG and RLP hydrogels (A-C). (D- F) 10 wt% 50/50 RLP-6Ac/PEG-4Ac and (G- I)10 wt% 50/50 RLP-2Ac/PEG-4Ac hydrogels UV irradiation at (D, G) 0, (E, H) 5 and (F, I) 10 min after mixing. Samples in panels D through I were also crosslinked with UV irradiation for 4 min. ... 77 Figure 3.6 Evolution of domain diameters in microstructured hydrogels. A)
Autofluorescence images of photo-crosslinked 10 wt% 50/50 RLP- 6Ac/PEG-4Ac and 10 wt% 50/50 RLP-2Ac/PEG-4Ac hydrogels;
microscale RLP-rich domains grow in diameter when precursors were incubated at room temperature for 0, 5 and 10 min prior to photo- crosslinking. B) Average particle diameters of the RLP-rich domains over time for RLP-2Ac and RLP-6Ac solutions with PEG-4Ac. C) Domain diameter distribution of the RLP-rich domains, with different times of incubation prior to photocrosslinking of RLP-PEG hydrogels. 78
Figure 3.7 Oscillatory rheological characterization of 10 wt% 50/50 RLP-Ac/PEG- 4Ac hydrogels. The comparison of storage moduli of microstructured hydrogels and equilibrium PEG-rich and RLP-rich phases for RLP- 2Ac and RLP-6Ac. ... 81 Figure 3.8 A) BCARS spectra of 10wt% RLP-6Ac and PEG-4Ac in PBS. B)
BCARS images of 10 wt% 50/50 RLP-6Ac/PEG-4Ac at the asymmetric CH3 stretching vibration (2930 cm-1); highest intensity correlates with RLP-rich domains (scale bar: 100 μm). ... 84 Figure 3.9 BCARS spectra of 10 wt% 50/50 RLP-6Ac/PEG-4Ac hydrogels.
Fingerprint RL spectra, normalized to ~1660 cm-1, from B) PEG-rich and C) RLP-rich areas in hydrogels at different crosslinking times. ... 84 Figure 3.10 BCARS images for 10 wt% 50/50 RLP-6Ac/PEG-4Ac hydrogels. A)
Ratio images (integrated intensities at 1468 cm-1/1660 cm-1)
representing the [PEG] relative to [RLP] within the hydrogels when photo-crosslinked at 0, 5 and 10 min. Yellow to red represents a high value of [PEG]/[RLP] and blue indicates high [RLP]/[PEG] ratios (i.e., low values of [PEG]/[RLP]) (scale bar: 10 µm). B) . Peak area ratios of the vibrations 1468cm-1 and 1660 cm-1 plotted versus crosslinking time (0, 5 and 10 min) and after 240 and 1440 minutes (obtained by peak fitting). ... 85 Figure 3.11 Micromechanical characterization of hydrogel domains via AFM
indentation. A) The distribution of Young’s moduli from indentation of RLP-rich domains and PEG-rich matrix for 10 wt% 50-50 RLP- 6Ac/PEG-4Ac hydrogels crosslinked 0, 5 and 10 min after mixing.
The phase-separated RLP-rich domains and PEG-rich matrix were visualized via optical microscopy and indented separately. Optical microscopy images of the RLP-PEG thin hydrogel with the AFM probe located at the RLP-rich domain and PEG-rich matrix. B) The box plot shows the statistical distribution of the data for the phase- separated domains crosslinked at 0, 5, and 10 minutes post-mixing, as well as the mechanical properties of the individual phases
photocrosslinked after bulk phase separation (Equil). The asterisk indicates statistically significant differences between the mean values of the marked samples and all other samples (p < 0.01). ... 88 Figure 3.12 Swelling ratios for the microstructured hydrogels and also for
hydrogels formed from the individual PEG-rich and RLP-rich phases after bulk phase separation. The bulk phase-separated solutions were isolated and crosslinked into hydrogels in separate samples prior to
Figure 3.13 Cytocompability and cell localization in microstructured hydrogels.
Confocal z-stack maximum intensity projections images for 3D cultures of encapsulated A-B) hMSC in 10wt% RLP-MMP-RGD- 2Ac/PEG-4Ac hydrogels and C-D) 10wt% RLP-MMP-RGD- 6Ac/PEG-4Ac hydrogels at A, C) day 1 and B, D) day 7. Colors indicate live cells (calcein, green), dead cells (ethidium homodimer, red), and autofluorescence of RLP-rich domains (white). ... 91 Figure 4.1. Evolution of domain diameters in hydrated RLP-PEG hydrogels.
Autofluorescence images of hydrated RLP-PEG hydrogels with
precursors were A) immediate (0 min) photo-crosslinked or B) incubated at room temperature for 5 min prior to photo-crosslinking. C) Surface area to volume distribution with different RLP-rich domains of RLP-PEG hydrogels... 99 Figure 4.2. Schematic of mechanical and micromechanical measurement for
microstructured hydrogels: A) oscillatory shear rheology, B) small- strain microindentation, and C) large-strain puncture mechanics ... 100 Figure 4.3. Oscillatory shear rheology of hydrogels A) Frequency sweep of
oscillatory rheology with representative average shear moduli of preformed RLP-PEG microstructured hydrogels compared with RLP- rich, PEG-rich hydrogels. The storage modulus G (solid) and loss modulus G (open) are both presented. B) Comparison of the storage modulus G, loss modulus G and C) tan(δ) of the hydrogels at
0.01Hz. ... 102 Figure 4.4 Microindentation of RLP-PEG microstructured hydrogels A) Stress-
strain loading curves during small-strain microindentation for RLP- PEG_25um and RLP-PEG_10um hydrogels, presented up to a strain of ~20%. B) Summary of elastic modulus measured from small-strain microindentation and oscillatory shear rheology at a comparable applied strain rate. C) Representative loading and unloading stress- strain cycle for RLP-PEG hydrogels from microindentation
experiments at different applied strains. D) Summary of averaged
resilience values for RLP-PEG hydrogels. ... 104 Figure 4.5. Force-displacement curves for puncture experiments with an indenter of
radius R=15 m for RLP-PEG hydrogels. ... 106 Figure 4.6 Confocal images during large strain indentation. A) Image of RLP-PEG
hydrogel was forced out along the side of the indenter in 3D view. B) Top x-y view of the indenter approached, indented and deformed the RLP-PEG hydrogel in 2D plane in each steps indentation. ... 107
Figure 4.7 Comparison of RLP-PEG, RLP-rich and PEG-rich hydrogels on puncture. A) Critical load for puncture (Pc) and B) effective elastic modulus (kʹE) with dependent indenter radius R. The data is further compared to the composite model with the upper and lower limits shown in red. ... 109 Figure 4.8 Impact of indenter radius on puncture mechanics of RLP-PEG hydrogels.
A)Critical load for puncture (Pc) and B) critical nominal stress at puncture (σc) normalized by shear modulus (G’) dependence on
indenter ratios R. ... 110 Figure 5.1 A) Phase diagram of RLP/PEG. B) Phase separation kinetics of 10wt%
50/50 RLP/PEG and 15wt% 80/20 RLP/PEG. ... 116 Figure 5.2 Stability of 10 mg/ml peptides incorporated in 15wt% 80/20 RLP/PEG
with x-y view images on top and z-view images in the bottom ... 117 Figure 5.3 Schematic of hydrogel formation. Physically crosslinked RLP/HA gel
and chemically crosslinked RLP-Ac/HA-SH gel crosslinking
reactions. ... 120 Figure 5.4 Oscillatory rheology of physical and chemically crosslinked hydrogels.
A) Time sweep data for RLP/HA physically crosslinked hydrogels, which formed immediately at 37oC. B) Frequency sweep
measurements of RLP/HA physical shows the formation of a physical gel at 37oC. C) Time sweep data for RLP-Ac/HA-SH chemically crosslinked hydrogels showing a gel point of approximately 40 min.
D) Frequency sweep data for RLP-Ac/HA-SH gel indicating that
stable hydrogels with solid-like properties are formed. ... 121 Figure 5.5 Rabbit vocal folds days 5 and 21 after injection with RLP hydrogels.
Hematoxylin and eosin stain. All demonstrate intact epithelium without inflammation and normal muscularis. A) RLP/HA gel at day 5. B) RLP/HA gel at day 21. C) RLP-Ac/HA-SH gel at day 5. D) RLP-Ac/HA-SH gel at day 21. E) RLP solution at day 5. F) RLP
solution at day 21. Scale bars 100 μm. ... 122
ABSTRACT
Hydrogels provide mechanical support and a hydrated environment that offer good cytocompatibility and controlled release of molecules, and myriad hydrogels thus have been studied for biomedical applications. Recent research has increasingly focused on multicomponent hydrogels that better capture the multifunctional and microstructural nature of native biological environments.
Multiple approaches to generate microstructured hydrogels have emerged in order to control microscale properties for applications ranging from mechanical reinforcement to regenerative medicine. In this thesis, we introduce new
heterogeneous hybrid hydrogels comprising emerging resilin-like polypeptides (RLPs) and poly(ethylene glycol) (PEG). Phase diagrams of the RLP/PEG system were
generated in order to define solution parameters that would yield micron-scale domains in the hydrogels. The hydrogels can be engineered with controlled microstructure and distinct micromechanical properties via the liquid-liquid phase separation (LLPS) of aqueous solutions of the RLPs and PEG. The microstructure in the hydrogels was captured by crosslinking a phase-separated RLP and PEG solution via a Mannich-type reaction with the crosslinker tris(hydroxymethyl phosphine) (THP). The production of RLP-rich domains and PEG-rich matrix was confirmed via confocal microscopy. The hydrogel mechanical properties were assessed via
oscillatory rheology and atomic force microscopy (AFM), with the hydrogels
exhibiting a moderate bulk shear storage modulus (ca. 600 Pa) and micromechanical properties of the domains (Young’s modulus ca. 13 kPa) that were distinct from those
of the matrix (ca. 6 kPa). These results demonstrate that tuning the parameters of the aqueous-aqueous phase-separated RLP/PEG solutions provides a simple,
straightforward methodology for fabricating microstructured protein-containing hydrogels, without extensive material processing or purification.
Despite the range of such microstructured materials described, few methods permit independent control over microstructure and microscale mechanics by precisely controlled, single-step processing methods. We further reported a photo- triggered crosslinking methodology that traps microstructures in LLPS solutions of RLP and PEG. RLP-rich domains of various diameters could be trapped in a PEG continuous phase, with the kinetics of domain maturation dependent on the degree of acrylation. The chemical composition of both hydrogel phases over time was assessed via in situ hyperspectral coherent Raman microscopy, with equilibrium concentrations consistent with the compositions derived from NMR-measured coexistence curves.
Atomic force microscopy revealed that the local mechanical properties of the two phases evolved over time, even as the bulk modulus of the material was constant, showing that our strategy permits control of mechanical properties on micrometer length scales, of relevance in generating mechanically robust materials for a range of applications. The successful encapsulation, localization, and survival of stem cells (hMSCs) was demonstrated and suggests the potential application of phase-separated RLP/PEG hydrogels in regenerative medicine applications.
Furthermore, micromechanical properties of RLP-PEG microstructured hydrogels were characterized via oscillatory shear rheology, small-strain
microindentation, and large-strain indentation and fracture. Oscillatory shear rheology and small-strain microindentation measured the small-strain elastic response of RLP-
PEG hydrogels. The elastic moduli calculated from rheology were comparable with the elastic moduli obtained from microindentation. Repeated cyclic loading and unloading microindentation revealed high resilience values (>85%) for RLP-PEG hydrogels even up to 80% strain. Large-strain puncture under a confocal microscope enabled the visualization of the microstructured hydrogel under indentation and deformation of RLP-rich domains. Puncture experiments also characterized the mechanical response and effective elastic moduli of the RLP-PEG, RLP-rich and PEG-rich hydrogels. The impact of spherical indenter sizes on puncture mechanics were also evaluated and extracted a fracture energy and maximum stress of the microstructured RLP-PEG hydrogels. Microstructured RLP-PEG maintain excellent mechanical properties and biocompatibility suggesting their potential in tissue engineering applications.
MULTICOMPONENT HYBRID HYDROGELS IN BIOMEDICAL APPLICATIONS
1.1 Introduction
Three-dimensional (3D) hydrogel networks provide mechanical support and hydrophilic properties that are advantageous for myriad applications ranging from those in consumer to biomedical products. The highly porous structure allows for fast diffusion of small molecules,1 and hydrogels thus have been used in separation and purification,2 biosensor,3,4,5 and tissue regeneration.6–8 Hydrogels provide a hydrated environment for cells, which improves their suitability for tissue engineering
applications.8–10 For tissue engineering purposes, hydrogels not only need to provide a physical support for cell growth, but also need to maintain a mechanically active and biochemically appropriate environment that provide cell-matrix interactions to direct cell proliferation and differentiation. Given the variety of properties necessary for optimizing material activity in the biological environment, multicomponent hybrid hydrogels have been of significant research interest.
The formation of a multicomponent hybrid network can be achieved via either chemical or physical means. Many biologically active proteins or peptides can simply be reacted with synthetic polymers via radical polymerization or other conjugation strategies, including click protocols,11–13 yielding multiple opportunities to easily produce multicomponent hydrogels. In particular, highly specific click reactions provide a simple way to produce macromolecules or hydrogel networks with a
Chapter 1
controllable network structure and patternable design. The non-toxic and mild
chemistries enable cell encapsulation and provide opportunities for hydrogel formation in vivo. In addition, the use of physical networks, including those formed from self- assembling peptides and proteins, has expanded the versatility of these physical approaches for producing self-assembling hydrogels.14–16
Both synthetic and natural polymers have been utilized for fabricating
scaffolds. For biological application, the materials must be inherently biocompatible, biodegradable, and cell adhesive. Additionally, they must have a porous, mechanically stable, and three-dimensional structure with facile manufacture. Synthetic materials provide a wide range of molecular structures and chemical capability,7,17 while
biomimetic materials, and in particular structural proteins such as collagen and elastin, provide mechanical characteristics unique to native tissue.18,19 Hybrid polymeric scaffolds combining natural and synthetic polymers have thus gathered significant and continued interest for their potential to mimic the extracellular matrix (ECM). In addition, to further improve the mechanical robustness of the hydrogel network,
composite hybrid hydrogels provide an additional mechanical reinforcement.20–22 Drug delivery can also be enhanced when a second phase, such as drug-loaded nanoparticles and microparticles, is incorporated in the hydrogel matrix.23,24
For most of the biochemically inert polymers, the lack of interaction between cells and hydrogels can limit the utility of the materials for directing cellular behavior, and accordingly, the purposeful design and production of multicomponent hydrogels to fulfill different biological function has grown.6,10,19,25–27 In addition to providing cell adhesion and cell-mediated degradation, incorporation of biofunctional biomolecules, including growth factors28–31 and signaling molecules17,32,33 can also facilitate cell
proliferation and differentiation. Controlled delivery of biomolecules to modulate immune response,34–36 with co-delivery of therapeutics and DNA, can further expand the functions of hydrogels beyond tissue regeneration to cancer and gene therapies.37–
39 The applications of these tunable hydrogels in biomedical engineering are numerous, owing to the ease by which functions can be altered by simple
incorporation of the components that are required for particular applications. This review focuses on the recent development and applications of multicomponent hybrid hydrogels.
1.2 Hydrogel Network Formation
Stable hydrogel networks are essential to provide structural support and can be formed by chemical and physical crosslinking; given the wide selection of
crosslinking methods available, multiple components can be randomly or selectively incorporated into the hydrogel networks. Chemically crosslinked hydrogel networks, employing covalent bonds, generally provide a stronger and more stable network, although chemical degradation or other strategies are then necessary for elimination of the hydrogels from a biological environment. Covalently crosslinked hydrogels can be formed via various reactions, including free radical polymerization,40–42 click
chemistry,12,43–45 and thiol-ene chemistry.46–48
The advantage of radical polymerization is that multiple, vinyl-functionalized components can react and form multicomponent hybrid hydrogels, such as
PEGDMA/GelMA49 and PEGDA/Hep-MA50 in a one-pot reaction. Incorporating bioactive components (e.g., gelatin and heparin) in the matrix imparts desired bioactivity while maintaining necessary mechanical strength. Pre-polymerization of
damage during in situ cell encapsulation,51,52 and there are multiple types of
photoinitiators (such as Igracure 295953 and lithium arylphosphinate (LAP)54,55) that maintain high cell viability, and conditions can be employed to make free radical polymerization useful for forming hybrid hydrogels in vivo.56
Click chemistry has been widely used in conjugation due to its fast, highly specific, and efficient reaction, which allows selective modification and incorporation of biologically active molecules (such as cell adhesion and enzymatically degradable peptides) in specific sites even in the presence of various functional groups and under physiological conditions.11,57 Hydrogels utilizing click chemistry have a well defined network structure and can show significantly improved mechanical properties.58 The most commonly used click reactions include alkyne-azide,59,60 Diels-Alder, 45,61 and thiol-ene reactions.44,59,62
Physical hydrogels, in contrast, are formed by secondary interactions, including hydrogen bonding, ionic interactions, and hydrophobic interactions.63 Cooperative physical interactions can be used to form stable hydrogels via crystallization, self-assembly, and thermally induced crosslinking. Although secondary interactions can provide stable hydrogels, the strength of the physical network can be altered by pH, temperature or organic solvent.64,65,66 Specific ligand- receptor binding events and self-assembling peptides also can be employed to form physical hydrogels, permitting the elimination of any potential toxic crosslinker or initiator. Although physical gels may suffer from weak mechanical properties and dissociation from the bulk material, physical crosslinks formed via multiple methods have been shown to be valuable in the production of multicomponent hydrogels.67–69
Spontaneous self-assembly, generally driven from cooperative physical interactions,70 has also been widely used in the formation of physical networks. A large range of biomacromolecules, including peptides and proteins can form network structures via formation of coiled-coil, triple helix and β-sheet structures; canonical examples include collagen-based71–74 and silk-based75–77 hydrogels. Peptide sequences that form self-assembled structures have thus been incorporated into hybrid hydrogels.
For example, the peptide sequence (AKAAAKA)2 has been conjugated to Pluronic®
polymers to form a self-assembled peptide/polymer hybrid hydrogel78,79 that showed a compressive modulus similar to that of native elastin and was capable of supporting cell adhesion.
Thermally responsive polymers, such as poly(N-isopropylacrylamide) (PNIPAAm), have also been employed in self-assembly and the formation of
injectable materials for biomedical and drug delivery applications.8 Many peptides and proteins conjugated to PNIPAAm exhibit materials with dual self-assembly and
thermally responsive properties.80–83 Hydrogels have been produced via the
interactions of coiled-coil domains of PNIPAAm- coiled-coil polypeptide –PNIPAAm triblock polymers. Below the LCST of the PNIPAAm, the hydrogel is only
crosslinked by the coiled-coil interactions of the polypeptide (Figure 1.1), and thus exhibits shear-thinning behavior, which is useful for injection. With an increase of temperature to above 37oC (e.g., upon injection in vivo), the thermally responsive PNIPAAm segments collapse and aggregate, resulting in a stiff hydrogel with a modulus up to 60kPa.82
Figure 1.1 PNIPAAm - coiled-coil peptide - PNIPAAm thermally responsive self- assembled hydrogel. (a) The hydrogel is crosslinked by the coiled-coil structure formed by the polypeptide and by PNIPAAm after its collapse and aggregation above its LCST. (b) Schematic of PNIPAAm - coiled- coil peptide - PNIPAAm and peptide sequence.82 Reproduced with permission from ref 82. Copyright (2013) WILEY-VCH Verlag GmbH &
Co. KGaA.
The versatility of polymer synthesis and modification enables the production of synthetic polymers in different molecular structures, including star and branched polymers and multiple networks. The widely employed tetra-functionalized PEG has been useful for forming hydrogel networks;25,84–89 tetra-PEG hydrogels have become popular owing to their simple, robust, and versatile chemistries.90 The networks have demonstrated improvements in extension and strength compared with conventional hydrogels,91 and more recent reports have shown that there are negligible local defects so that the networks produced from the tetra-PEGs act as a nearly ideal elastic
network.92 In another example, a reducible micelle hydrogel has been formed, using a multi-arm PEG-containing copolymer, for drug delivery applications. The 8-arm PCL-
PEO copolymer was linked by a disulfide core and exhibited a micellar structure;93 the micelles then further crosslinked to form hydrogels. Micelle size could be reduced in the presence of a reducing agent, which cleaved the di-sulfide core linkage and
reduced the sizes of the multi-arm polymer by half (to yield a 4-arm architecture). The mechanical strength of the 8-arm hydrogel was nearly 10-fold that of a control
hydrogel formed with a crosslinked linear copolymer, and the modulus of the 8-arm micellar hydrogel was decreased 58% when the multi-arm polymer was reduced to the 4-arm polymer.
In addition to these variations in polymer architecture, hybrid networks formed with two different polymers have been shown to exhibit excellent mechanical
properties. Interpenetrating polymer networks (IPNs), for example, are among the earliest multicomponent, hybrid polymer networks; the concept of IPNs was
introduced in the 1960s and remains an active research area.94 Double networks are one unique type of IPN system that contains two types of polymers with asymmetric network structure95 (Figure 1.2) and has provided significant improvement in the strength of hydrogels compared to that of single networks.96–99 A double poly(2- acrylamido-2-methylpropanesulfonic acid) (PAMPS)/PAAm network hydrogel, formed via a two-step polymerization, has improved the compressive strength of the hydrogel over 20 times relative to PAMPS and PAAm single network hydrogels while retaining highly elastomeric behavior.99 Other groups have combined biopolymers such as gelatin and bacterial cellulose (BC) to form double network hydrogels with high mechanical strength (up to 5MPa in compression),98 or PVA/PAAm materials for load-bearing cartilage substitution.100
Figure 1.2 The PAMPS and PAAm network of the double network hydrogel under tensile test. The highly crosslinked PAMPS network fractured while loosely crosslinked PAAm network still holding the gel stucture during extension.95 Reproduced with permission from ref 95. Copyright (2010) The Royal Society of Chemistry.
1.3 Mimics of Natural Proteins
Natural hydrogels, including proteins and polysaccharides, have been used in biological applications and tissue engineering due to their biocompatibility,
biodegradability and biological functions.17 Natural polymers, such as alginate,101 chitosan,102,103 gelatin73,104,105 and elastin106,107 are able to form physical hydrogels, but often have poor mechanical properties.9 However, modification of natural polymers is often more difficult, with fewer chemical options compared to those available with synthetic polymers, and the purification of natural polymers often suffers from batch- to-batch variability. In addition, natural polymers extracted from animals or bacteria raise concerns about immunogenic reactions.90 A recent review includes details regarding polysaccharide-based hydrogels for tissue engineering applications;108 we
include here descriptions of protein-based hydrogels based on recombinant polypeptides109 for tissue engineering applications.
1.3.1 Elastin
Elastin is one of the most important structural proteins in mammals, providing the elastomeric behavior of most tissues, including tendons and blood vessels.110 The canonical amino acid sequence that gives rise to the mechanical properties of elastin is the flexible VPGXG repeat, where X can be any natural amino acid except proline.
Recombinant methods have enabled the development of an enormous variety of biosynthetic elastin-like-polypeptides (ELPs).19,106,107,111–116
The inverse transition behavior of elastin, in which ELP forms coacervates above a critical transition temperature, has been widely studied as a function of pH, salt concentration, and temperature.117 The transition temperature can be tuned by variations in the amino acid sequence, where the addition of hydrophobic residues reduces the transition temperature.118 ELP nanoparticles have been produced to encapsulate and release bone morphogenetic proteins (BMP) for potential protein and drug delivery applications.119 With the advantages of ELPs, they have been
incorporated into multicomponent materials (both chemically and physically
crosslinked) to enhance both the mechanical and biological functions.120 Multi-block elastin polypeptides containing the hydrophobic IPAVG end block for physical crosslinking have shown high extension and tensile strength.120
To further improve the biological properties of ELPs, various cell adhesion peptide and degradation domains have been added to the ELP sequences to improve cell adhesion, spreading and migration.121 An RGD peptide was incorporated on the
endothelialization in vascular grafts;106 the surface-specific conjugation enhanced the adhesion and proliferation of both endothelial cells and mesenchymal stem cells.
Other groups have taken advantage of the reversible, thermally responsive behavior of ELPs to form low-concentration, injectable hydrogels that can be crosslinked via disulfide bonding of cysteine residues in vivo.122 It has been possible to predict and tune the inverse transition temperature of a wide range of ELPs via sequence design.68,114,115,123–125
In addition to hydrogel matrix materials, ELPs also can form nanoparticles and nanofibers. Silk-elastin milti-block polypeptides can self-assemble into nanoparticles with the silk block in the core.77 Nanoparticles have also been formed from the elastin- mimetic hybrid copolymer PAA-VPGVG;126 in this particular case, the nanoparticles were formed by collective hydrogen binding and hydrophobic interactions, rather than by coacervation of the elastin-like domains, and are of interest in drug delivery
applications. ELP electrospun fibers, crosslinked with glutaraldehyde in a vapor- initiated process and then rehydrated in NaCl buffer,127 have provided opportunities for the use of hydrogel fibers to guide cell direction and to mimic the orientations of cells in native tissue. The opportunities for employing ELPs in biomedical fields continue to expand, not only as a result of the mechanical properties that are
comparable to those of native elastin, but also due to the responsive behavior of ELPs in which makes them highly versatile for drug delivery applications.
1.3.2 Resilin
Resilin is another structural protein, found in insects, where it is located primarily in active ligament and tendons.128 The excellent resilience and energy storage allows resilin to recover from repetitive high-strain cyclic loading with
essentially no hysteresis, even under high frequency conditions, which has an important role in insect flight and jumping129 and in sound production.130 Repetitive constructs of the consensus sequence of resilin from D. melanogaster
(GGRPSDSYGAPGGGN) have been produced from the first exon of the Drosophila CG15920 gene via recombinant methods, and the polypeptide showed excellent mechanical properties comparable to those of native resilin.131 The unique resilience of crosslinked RLP and hybrid RLP hydrogels has motivated their use in applications requiring highly elastomeric and biomechanical functions, such as vocal fold
therapeutics,132 artificial muscles,133 and cardiovascular applications.134 The RLPs show pH- and temperature-responsive behavior related to that of ELPs, although in addition to the inverse transition temperature, select RLPs can show dual phase transitions with both upper and lower solution critical temperatures.135
To improve the biological functionality of the RLP, our group has produced multiple constructs that incorporate cell adhesion domains (RGD), enzymatic
degradation domains (MMP-sensitive), and heparin-binding domains (HBD) to yield a multi-biofunctional material (Figure 3).132,136–139 RLP-based hydrogels can be
crosslinked by the reaction of amines in the RLP sequence (Lys) with the small- molecule crosslinker tris(hydroxymethyl phosphine) propionic acid (THPP) or tris(hydroxymethyl phosphine) (THP). Hydrogels formed by these methods exhibited excellent mechanical properties characteristic of resilin, while improving cell adhesion and cell-mediated degradation. In studies from other groups, the bone morphogenetic protein-2 (BMP-2) peptide has been incorporated into RLP films derived from A.
gambiae; the resulting surfaces promoted osteogenic differentiation of mesenchymal stem cells.140
Figure 1.3 Resilin-like polypeptide hydrogels demonstrate useful mechanical
properties and biological functions.132 Reproduced with permission from ref 132. Copyright (2012) The Royal Society of Chemistry.
Other recombinant constructs have combined the properties of multiple structural proteins into a hybrid resilin-elastin-collagen (REC) polypeptide.18 This polypeptide self-assembles into fibrous structures via the interactions of collagen, yield materials with a Young’s modulus between 0.1 and 3 MPa, consistent with those observed for native resilins and elastins. In a related example, the well-characterized GB1 domain was combined with random-coil resilin-like domains to produce
multiblock mimics of the passive elastic muscle protein titin.133 The material showed high resilience at low strain and was durable at high strain, consistent with the observed properties of muscle.
We have also explored hybrid RLP materials produced with synthetic polymers as matrices for cardiovascular tissue engineering.134 The RLP was synthesized via biosynthetic methods and contained the RGD integrin-binding
domain, MMP degradation domain, and heparin-binding domains of the sequences described above. Four-arm vinyl sulfone-terminated PEG was reacted with the cysteine-containing RLP via Michael-type addition. The resulting hybrid hydrogel maintained the mechanically active and biologically active domains, and supported the spreading of AoAFs during in vivo culture to a significantly greater extent than RLP- only hydrogels. Incorporating RLP and PEG together provides the mechanically durable and resilient hydrogel, with improved cell interactions, that may be useful in the engineering of mechanically active tissues.
1.4 Composite Hydrogels
Conventional hydrogels often exhibit weak mechanical strength and poor deformation (e.g., gels from gelatin and agarose),92 and increasing crosslinking density has been a common method for improving mechanical properties both natural and synthetic polymeric hydrogels.7 However, high crosslinking density results in restriction of the chains which yields stiff materials with limited extensibility and reduced water content in the swelled state,63 as well as compromised permeability and slow molecular diffusion.141 Composite hydrogels have thus been investigated as a strategy for improving the mechanical strength of hydrogel-based materials.142 These strategies employ traditional composite approaches, in which a filler is either
physically entrapped or chemically crosslinked within the hydrogel matrix to produce materials with increased mechanical strength. Mechanically stiff fillers, such as nanoclays, in the composite networks serve as reinforcement and as a multi-point crosslinker to improve the mechanical strength of the composite hydrogel, obviating the requirement for a high network density.143 The reorientation of the filler and
example, nanocomposite hydrogels utilized exfoliated nanoclay to reinforce a
PNIPAAm hydrogel; these materials showed both excellent mechanical strength (up to 1000 kPa) and high elasticity (up to 1000% strain-to-break).144–146 Composite
hydrogels have since been produced to incorporate a broader scope of inorganic species including SiNPs,147–149 metal nanoparticles,142,150 hydroxyapatite,22,29,151 carbon nanotubes (CNTs),152 and graphene oxide (GO) sheets153as reinforcement.
Although the strength and modulus of these organic-inorganic systems is significantly improved with the addition of the inorganic matrix, leaching of the inorganic species is a concern. In recent decades, the development of organic nanocrystals, organic particles, and electrospun polymer fibers have provided alternatives that avoid the need for the inorganic filler.
1.4.1 Nanocrystal-reinforced Matrices
Polysaccharide nanocrystals, formed primarily by crystal-forming cellulose and chitin, have been utilized to replace inorganic filler in nanoparticle-reinforced hydrogels.21 The rod-like nanocrystals, also referred to as nanowhiskers, can be
extracted from natural materials; cellulose nanocrystals are often extracted from cotton or ramie, and chitin nanocrystals are extracted from shrimp or crab.154,155 These
nanocrystals have the advantage of being biocompatible and biodegradable, as well as having mechanical strength and moduli that are comparable to those of inorganic fillers (over 100GPa).154 Different groups have incorporated cellulose nanocrystals (CNC) or chitins as reinforcement fillers for PAAm,156,157 PVA,158 chitosan,159 and CMC/HEC160 hydrogels. The mechanical properties of the composite hydrogels generally increase with increased nanocrystal content.
CNCs have also been used, in electrospinning of PEO, to reinforce the resulting nanofibers;161 the composite nanofibers showed an increased modulus (38 MPa) compared to that of PEO fiber (15 MPa), and these properties depended on the CNC content. CNC-reinforced, injectable hydrogel comprising a carboxymethyl cellulose and dextran matrix have also been produced;21 chemically crosslinked, CNC- reinforced hydrogels showed a higher modulus compared to physically blended CNC hydrogels. The development of such polysaccharide nanocrystal composites has provided biocompatible and biodegradable fillers, which has enabled the use of nanocrystal composite hydrogels in tissue engineering. However, the sizes of the nanocrystals are limited in scope due to their extraction from naturally occurring materials, thus the options for engineering properties by altering filler dimensions is also limited.
1.4.2 Particle-reinforced Matrices
In addition to nanocrystal-containing composite hydrogels, synthetic organic nanoparticles and microparticles also have been incorporated into hydrogels for mechanical reinforcement. For example, the uniform dispersion of monodisperse cationic polystyrene (c-PS) nanoparticles into a PAAm hydrogel improved the
compression strength to 40MPa compared to the original 70 kPa modulus of a PAAm- only hydrogel.162 The improvement in mechanical properties was attributed to the uniform dispersion of monodisperse c-PS that were pre-fabricated by emulsion polymerization. Another group incorporated the thermoresponsive PNIPAAm microgels into the PAAm matrix and evaluated the mechanical properties below and above the LCST of the PNIPAAm that led to understanding the effect of soft and hard
163
composite hydrogel is that they can be used not only reinforce the mechanical properties, but can also serve as a vehicle for drug and/or protein delivery. The incorporation of block copolymer micelles (BCMs) in PAAm hydrogels via free radical polymerization resulted in hydrogels that sustain significant elongation (up to 480%),164 and that could also be loaded with hydrophobic drugs (via loading of the hydrophobic core of the BCMs during micelle formation) to permit drug delivery upon mechanical deformation of the hydrogel. Other organic nanoparticles, including hyperbranched polymers,165 polymeric nanoparticles,162,166 micelles,164 and/or nanogels,150,167,168 have also been used in the production of composite hydrogels for controllable drug delivery. For example, hyperbranched polyester (HPE) hydrogels enabled the entrapment of the hydrophobic drug dexamethasone acetate within the HPE hydrophobic cavities without causing drug aggregation, and showed longer sustained release compared to drug encapsulated in a PEG hydrogel.165 The drug- loaded nanoparticle composite hydrogel was able to achieve sustained release and a high drug concentration for local delivery,172 and drug delivery could also be triggered with stimuli such as temperature or mechanical deformation.168
Composite hydrogels are not limited to those formed with nanoparticles;
microgel hydrogels have also been shown to improve strength and torsion resistance.
Poly(2-acrylamido-2-methylpropanesulfonic sodium) (PNaAMPS) microgel-
reinforced the PAAm double-network hydrogel films have shown high tensile strength (up to 2.6MPa with a strain up to approximately 10%; Figure 1.4).169 Pre-formed microgels were incorporated into a PAAm hydrogel to form two-phase composite materials. The additional PAAm double network resulted in even greater mechanical enhancement compared to microgel-reinforced single-network hydrogels (e.g., a
modulus of nearly 120kPa compared to the modulus of the reinforced single network of approximately 50kPa).170
Figure 1.4 Microgel-reinforced double network PAAm hydrogel that exhibited excellent extension (a) and torsion (b). Microgel before tensile deformation (c) and after deformation (d).169 Reproduced with
permission from ref 169. Copyright (2013) American Chemical Society.
Nanoparticles and microparticles can be fabricated via various methods, including emulsion polymerization,162,171–173 self-assembly77,117,119 and phase separation.174–176 In one example, 8-arm PEG has been used to form PEG
microspheres via phase-separation in aqueous media.174–176 The PEG microspheres could be crosslinked via the reaction of amines with vinyl sulfone or with acrylate, and the sizes of the microspheres were controllable in different media, with improved cell viability in a microsphere-based scaffold.174 Compared to microspheres formed via emulsion polymerization, these microspheres do not require extensive solvent exchange or washing to remove organic solvent, although the reaction conditions needed to be precisely controlled to prevent bulk gel gelation. Improved control over
the reaction kinetics and changes in particle sizes over time will enable better control of the microspheres and properties of the resulting matrices.
1.4.3 Fiber-reinforced Matrices
The native ECM comprises a complicated and often anisotropic structure, with a combination of fibers and network polymers, such as collagen fibers aligned in tissue.27 Thus, the use of fibrous structures in designed materials has been employed to better mimic native ECM and guide cell direction; electrospinning has been a widely used and simple method to produce controlled nanoscale fibers.177 The applied high- voltage electrostatic force draws a polymer fiber from polymer solutions,178 and the resulting fibers can collected into isotropic or aligned fibrous mats. The activities of cardiomyocytes cultured on random and aligned electrospun biodegradable
polyurethane fiber mats were different, with greater multi-cellular organization on the aligned fiber mats.179 Materials comprising PLGA/gelatin electrospun nanofibrous have also been produced to mimic cardiac tissue;180 after electrospinning, the
hydrophilic gelatin could be rehydrated to yield fiber-like hydrogels. Cardiomyocytes cultured on the PLGA/gelatin nanofiber showed enhanced attachment and spreading.
Thermoresponsive multiblock poly(PEG/PPG/PCL urethane) hydrogel nanofibers have also been produced for temperature-mediated BSA release from fibers,181 and encapsulated proteins, such as nerve growth factor (NGF)182 and lysozyme,183 maintained their bioactivity after release from PCL-based electrospun fibers.
Nanofibers are also commonly employed fillers used to enhance the
mechanical properties of hydrogels. Fibers produced from several biocompatible and biodegradable polymers – including PCL, PLLA and chitosan – have been studied in different hydrogel systems. Chitosan nanofibers (CNF) incorporated in a PAAm
hydrogel improved the mechanical properties of the CNF/PAAm hydrogel compared with those of chitosan/PAAm hydrogels, showing a 2.5-fold higher compressive stress to 50.2 kPa (at 95% strain) than the chitosan/PAAm hydrogels.184 In another example, biodegradable PCL was electrospun with gelatin to forma PCL-gelatin core-shell fiber,20 which was mixed with gelatin and crosslinked to form a composite hydrogel.
The fibrous composite hydrogel showed an improvement in modulus to 20.3kPa from 3.2kPa (for a gelatin-only hydrogel). In addition, the fibrous structure of the PCL- gelatin alone served to direct cell orientation in a 2D aligned electrospun fiber mat,179 similar to other studies described above. The fibrous composite hydrogel provides a hydrated local environment and 3D support for cells, which is an advantage over traditional fiber mat scaffolds. The construction of aligned fiber hydrogel constructs for cell culture applications remains an active research area owing to its potential in various therapies, including the cardiovascular area.
1.5 Hybrid Materials with Engineered Biological Functions
Although the strategies described above have provided alternatives for achieving mechanically robust networks, a lack of cell-matrix interaction often leads to the failure of the biomaterials in in vitro and in vivo studies.185,186,187 Various cell- matrix interactions, including cell adhesion and matrix degradation are required for cell growth and migration,25 and hybrid hydrogels can be employed to capture these properties in a chemically and mechanically versatile substrate.
Figure 1.5 Important materials design considerations for tissue engineering, including cell adhesion peptide, protease sensitive peptide for cell-mediated matrix degradation, and presence of signaling molecules.
An inherent limitation of synthetic materials in biological applications is the lack of cell-matrix interactions, which limits cell attachment, remodeling, and migration in a scaffold. Incorporating ECM molecules and cell adhesive peptides (such as those from fibronectin and laminin) in the matrix materials has been widely shown to provide significant enhancement in cellular interactions with various scaffolds.26,27,90,187–189 The integrin-mediated cell adhesion facilitated by these
macromolecules provides for cell attachment, spreading, actin organization, and focal adhesion.187 The Arg-Gly-Asp tripeptide (RGD) has been the most commonly
employed cell adhesive peptide in hybrid hydrogel systems because of its effective cell adhesion through most integrins.188 Besides the RGD peptide, sequences derived from laminin (LN) (such as IKVAV, YIGSR) and fibronectin (FN) (such as
KQAGDV, REDV) also have been used to induce cell adhesion on hydrogel matrices.90 Cell adhesion peptides that have been employed in hydrogel matrices;
these sequences, and others, have shown value for stabilizing cells in matrices, as well as facilitating cell migration and maintaining cell functions.190–194
Besides cell adhesion, controllable degradation of the matrix material is also important for cell growth and tissue regeneration. The designed scaffold has to degrade at a rate comparable with cell growth and deposition of ECM molecules.
Perhaps the most commonly used degradation mechanism for synthetic hydrogels is hydrolytic degradation of ester linkages or polyester segments in polymers.90 Despite the widespread and simple application of these hydrolytic strategies, however,
hydrolytic degradation rates are difficult to control in vivo and are not controlled by cell growth.109,195,196 Therefore, cell-mediated degradation strategies have been employed to optimize scaffold degradation with ECM deposition.25,54,191,192,197,198
Matrix metalloproteinase (MMP)-sensitive peptides are a class of enzyme- sensitive peptides derived from native ECM proteins, such as collagen or elastin, that promote cell-mediated matrix degradation;90 The use of these sequences offers substantive flexibility in controlling matrix degradation, as the substitution of amino acids in a MMP-sensitive peptide modifies degradation kinetics.186 The degradation rates of the materials can extend over a wide range of time scales by simple variations of the amino acids in the sequences, which can provide sufficient control for
achieving degradation times that match the needs of a given application. In one example, the morphology of hMSCs encapsulated in MMP-sensitive peptide
crosslinked PEG hydrogel depends on the concentration of MMP-sensitive peptide in the hydrogel; variations in the peptide concentration in the hydrogel also permitted the control of hMSC differentiation in different culture media.54
In addition to the use of MMP-sensitive peptides for cell-mediated matrix degradation, hydrogels with controlled degradation rates have also been widely employed in drug delivery. The incorporation of a human neutrophil elastase (HNE)- sensitive peptide in a PEG hydrogel via thiol-ene chemistry87,199 was employed to trigger the release of a model protein upon triggered degradation of the HNE-sensitive sequence,199 indicating the potential for cell-mediated degradation in drug delivery applications.200,201 Controllable matrix degradation is also important in 3D cell culture.
Relevant examples include the use of a substrate, carboxybetaine methacrylate (CBMA), for reaction with a disulfide containing crosslinker via radical
polymerization to form a hydrogel in the presence of cells. During cell culture, this hydrogel rapidly degrades owing to the reaction of the disulfide-containing crosslinker with the cysteine-containing media, permitting recovery of the encapsulated cells.52 Recent exploitation, in our laboratories, of retro Michael-type addition has also been employed to control hydrogel degradation. In these cases, degradation of select thioether succinimide bonds has been employed to degrade PEG/heparin hydrogels and release heparin at glutathione (GSH) concentrations consistent with intracellular concentrations.202 The degradation mechanism can also be employed for GSH- triggered release of model proteins from PEG-only hydrogels, providing an opportunity for targeted protein delivery over timescales unique from those of disulfide- or hydrolytic-mediated mechanisms.203 A recent review provides a
comprehensive description of hydrogel degradation in cellular microenvironments via hydrolytic, enzymatic, thiol-exchange, and photolytic mechanisms.195
The recognition of materials by macrophages, which release chemokines to recruit immune cells, and subsequent chronic immune responses often lead to rejection
