First Line of Title

RESILIN-LIKE POLYPETIDE-BASED MICROSTRUCTURED HYDROGELS

VIA AQUEOUS-BASED LIQUID-LIQUID PHASE SEPARATION FOR

TISSUE ENGINEERING APPLICATIONS

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

VIA AQUEOUS-BASED LIQUID-LIQUID PHASE SEPARATION FOR

TISSUE ENGINEERING APPLICATIONS

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

Signed: ______Christopher J. Kloxin, Ph.D. Member of dissertation committee

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 multi-photon and confocal microscopes and also extend thanks to Dr. Chandran Sabanayagam for help with nanoindentation on

iv 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.

v TABLE OF CONTENTS

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 2.2.4 Characterization of Equilibrium Concentrations ...... 33

vi 2.2.5 Fluorescent Labeling of RLP and PEG ...... 33 2.2.6 Hydrogel Formation ...... 34 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

vii 3.3.5 Cell Viability and Growth in 3D LLPS Microstructured Elastomers ...... 89

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

viii LIST OF TABLES

Table 2.1 Concentrations of RLP and PEG in the upper (phase I) and lower (phase II) 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

ix LIST OF FIGURES

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

x 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 coexistence 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

xi 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

xii 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

xiii 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 -1 asymmetric CH3 stretching vibration (2930 cm ); 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 the measurement of swelling ratios...... 89

xiv 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

xv 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

xvi 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

xvii 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-

xviii 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.

xix Chapter 1

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

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

2 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 the precursor solution before inclusion of cells can reduce free-radical induced cell

3 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

4 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

5 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-

6 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

7 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

8 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 surface of a multi-block ELP gel via maleimide-thiol chemistry to promote luminal

9 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

10 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

11 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

12 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 polymeric network then serves to maintain the high elasticity of the hydrogel. In one

13 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.

14 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 filler on the hydrogel.163 An advantage of the synthetic organic particles in the

15 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

16 modulus of nearly 120kPa compared to the modulus of the reinforced single network of approximately 50kPa).170

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

17 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

18 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.

19 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;

20 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

21 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

22 of the implants or scaffolds.35 Recent studies suggest that an active modulated immune response can direct tissue regeneration;204 inflammatory cytokines have an important role in initiation of acute inflammation, cell proliferation, and modulation of tissue healing.35,205 Interleukin-1 (IL-1), granulocytecolony stimulating factor (G- CSF), granulocyte macrophage colony stimulating factor (GM-CSF), CC-chemokine ligand 2 (CCL2), and CCL5 are several of the important factors for tissue healing.25 Hydrogels that deliver GM-CSF topically have been shown to enhance wound healing in patients with second degree burns.206 In addition, chemokines can induce chemotaxis that guides progenitor and stem cell migration and tissue reconstruction. Stromal derived factor 1 (SDF-1), in one such example, was loaded in PEG-heparin hydrogels and showed significant improvement in guiding the migration of early endothelial progenitor cells (eEPCs) compared to gels that did not contain SDF-1;207 the incorporated SDF-1 also reduced scar tissue formation and promoted improved tissue healing.208 Growth factors and tolerance-promoting antigens also have also been shown to enhance tissue regeneration.31,209 For example, regeneration of muscle in a mouse model could be promoted via the use of an RGD-modified alginate hydrogel for co-delivery of vascular endothelial growth factor (VEGF), insulin-like growth factor-1(IGF-1), and myoblasts;210 the VEGF promoted angiogenesis and IGF-

1 promoted myogenesis. Hydrogels able to incorporate and controllably release multiple biomolecules, including cells, cytokines and growth factors, may improve tissue regeneration by the minimization of chronic immune responses and enhancement of tissue growth.

23 1.6 Conclusion and Perspectives Multifunctional hydrogels exhibit improved mechanical and biological properties that can be modulated via chemical and physical methods. The existence of well-developed chemistries for bioconjugation and crosslinking, including an expanding range of click reactions, has enabled the controlled incorporation of a variety of multifunctional groups and the design of specialized crosslinked networks containing composite structures and both synthetic and biological materials. Strategies for increasing crosslinking density (to improve modulus), while at the same time maintaining elasticity, have been of enormous interest and promise. The mechanical properties can be enhanced by judicious design of the matrix polymers (and copolymers) and/or the components in the gel; the combination of synthetic and natural polymers offers interesting opportunities to obtain biomechanically active hydrogels. Materials based on elastin and resilin can provide mechanically active function that mimics the biomechanical properties of the native tissue. However, comprehensive studies on the cellular response and in vivo studies of these synthetic and natural hybrid hydrogels remain limited. The development of composite hydrogels has provided a versatile alternative approach for improving the strength of hydrogels via the use of a stiff second network that reinforces the weak hydrogel network, or via the incorporation of particles in the hydrogel matrix. Hybrid two-phase hydrogels also provide an addition platform for stimuli-induced drug delivery, with the drug stably encapsulated in the second phase until a stimulus is applied. The applications of composite hydrogels as tissue engineering scaffolds has been useful for incorporating drugs into matrices, and modulating the co-delivery drugs or molecules at different release rates, while enhancing the mechanical strength. Additional studies that investigate the ratio of the

24 two phases, and the resulting impact on mechanical properties and release kinetics of cargo from the hybrid hydrogel, are needed to inform the design of materials that can control the release of multiple drugs. In addition, while most composite hybrid hydrogels are produced in two steps (particle fabrication and subsequent encapsulation into hydrogel matrix), strategies that would simplify composite gel production into a single step would find significant value, as it would eliminate the need for additional purification of particles prior to their incorporation into hydrogels for biomedical uses.

Extensive biological studies are needed to evaluate those materials for such use. The ability to encapsulate viable cells in 3D formats is a step toward effective cell delivery and tissue regeneration. The incorporation of bioactive peptides has been widely employed to control cell attachment, proliferation and differentiation within synthetic hydrogels, and cell-mediated degradation of these matrices has improved cell growth and spreading. Appropriate design of multicomponent hydrogels has enabled interesting and mainly untapped opportunities for programming cell behavior to stimulate simultaneous immunotherapeutic treatment and tissue regeneration. While most of the immunomodulating hydrogels studied have been weak physical hydrogels, such as alginate, there is demonstrated and continued need to employ chemically crosslinked and mechanically robust hydrogels for understanding the impact of the matrix on immune response. While it is well known that the mechanical properties of a matrix modulate cell behavior, the impact of the mechanical properties of a matrix on immune system has not yet been studied in detail; further understanding of these processes will inform tissue regeneration. Taken together, the body of work described herein clearly illustrates that the potential of multicompouent hybrid hydrogels for a variety of applications in tissue

25 regeneration and drug delivery. By incorporating and modulating the mechanical functional and bioactive components in the network, the mechanical and biological properties of the hydrogel can be tuned independently without sacrificing one or the other. In the future, hybrid hydrogels are expected to further mimic the microenvironment for cells and tissue reorganization. The mechanically active components should be aimed not only at affecting the bulk mechanical properties, but also should capture the micro-mechanical properties in native tissue. Multicomponent hydrogels with well-organized domains will offer significant opportunities for these materials.

1.7 Dissertation Summary This dissertation reports the investigation of resilin-like polypeptide-based microstructured hybrid hydrogels for tissue engineering applications. Chapter 1 introduces current research focusing on multicomponent hybrid hydrogels and the need to mimic the microenvironment for cells and tissue reorganization. The mechanically active components should affect bulk mechanical properties, and also could capture the heterogeneity in micro-mechanical properties in native tissue.

Chapter 2 presents heterogeneous hybrid hydrogels comprising emerging RLPs, which can be engineered with controlled microstructure and distinct micromechanical properties via the LLPS of aqueous solutions of the RLPs and PEG.

In Chapter 3, photo-triggered crosslinking methodology is employed to capture microstructures in LLPS solutions of a highly elastomeric RLP and PEG, with independent control over microstructure and microscale mechanics by single-step processing methods. Chapter 4 presents further characterization of the micromechanical response of the RLP-PEG mucrostructured hydrogels with small- strain microindentation and large-

26 strain indentation and fracture. Chapter 5 concludes and summarizes this thesis and discusses potential future directions.

27 Chapter 2

AQUEOUS LIQUID-LIQUID PHASE SEPRARATION OF RESILIN-LIKE POLYPEPTIDE/POLYETHYLENE GLYCOL FOR FORMATION OF MICROSTUCTURED HYDROGELS

2.1 Introduction

Heterogeneous hydrogels have been investigated as a novel strategy for enhancing matrix complexity of biomimetic materials for a variety of applications.10,211–213 Microstructures and microdomains induced by macropores, fibers, and/or particles within a hydrogel have been shown to modulate the micromechanical properties and provide desired physical and biological functions. For example, macroporous214,215 and fiber-containing biopolymer-polymer hybrid materials216,217 have been explored to provide stable mechanical properties,163 to enhance diffusion and transport of macromolecules and drugs,168,218 and to promote cell spreading.72,211,219 Multiple approaches to generate 3D heterogeneity in materials have thus emerged in order to capture this complexity and microstructure, with many strategies focusing on photo-patterning as well as the incorporation of microparticles.

While photopatterning has been highly useful for the microfabrication of acrylated polymers and biopolymers,104,220 traditional methods are limited to surface modification and emerging multi-photon techniques, which offer opportunities for 3D patterning of biomolecules221–225 and degradation,226–228 are limited to small-scale patterning. Production of organic polymeric nanoparticles, nanogels, and microgels has also been widely studied for making tough materials229–234 and matrices capable of

28 drug delivery.163,235–238 Current approaches to incorporate those particles into 3D microstructured hybrid hydrogels traditionally employ two steps, with initial fabrication via various methods such as emulsion polymerization162,171–173 or microfluidic techniques,239 followed by particle purification and incorporation into the network. Incorporated particles can be used to improve mechanical or delivery properties or they can be selectively degraded within the hydrogel to provide porous microstructures via hydrolysis,213 enzymatic cleavage,211 or photolytic degradation,226–

228 to further affect cell spreading and delivery profiles. Strategies that simplify the production of microstructured hydrogels into a single step would find significant value for making heterogeneous materials for a variety of applications. Liquid-liquid phase separation (LLPS) has been a well-known phenomenon for more than a century, with the unfavorable interactions of two dissimilar polymer blends, block copolymers, proteins, and polysaccharides causing phase separation of ternary solutions even at a low concentration (< 50 mM).240–242

LLPS provides a method for generating microstructured materials, as phase behavior is readily tuned by factors such as temperature, polymer molecular weight, polymer concentration, ionic strength, pH, and addition of specific salt types.240,243 Indeed, dextran-PEG solutions have been well studied and used for separate fabrication of

PEG174,176 and dextran244 microparticles. Protein-polymer LLPS, in comparison, has been used mainly to understand thermodynamically-driven phenomena including protein crystallization,245,246 the stability of therapeutic monoclonal antibodies,247 and separation/purification technologies.248 Such investigations, however, have not been aimed at utilizing aqueous LLPS for the development of structured protein-based biomaterials.249,250

29 Emerging resilin-like polypeptides (RLPs), based on the insect protein resilin, offer intriguing opportunities in the generation of protein-containing microstructured hydrogels, as they exhibit exceptional rubber-like elasticity characterized by low stiffness, high extensibility, and efficient energy storage.251–253 The mechanical properties of RLP hydrogels are compatible with those needed for mechanically active tissues,216,254–257 and we have previously reported the spreading of hMSCs within RLP-PEG hydrogels, with morphologies distinct from those of hMSCs cultured in

RLP- and PEG-only hydrogels.258 While hybrid RLP polypeptides have been designed to merge the outstanding properties of resilin with those of other proteins such as collagen, elastin, and immunoglobulin-like domains,18,133 little has been reported on RLP-polymer hybrids. In addition, earlier studies that have explored the use of collagen, elastin, and silk in composite materials (fabricated as particles, films,259 fibers,260–262 or hydrogels263) have generally not exerted systematic control of microstructure to alter materials properties. The use of RLP/polymer solutions that show controlled LLPS should thus provide useful opportunities to expand the properties and versatility of protein-polymer matrices. In this study, we exploit the LLPS of aqueous solutions of RLP and PEG to generate heterogeneous hybrid hydrogels with controlled microstructures and microdomains with distinct mechanical properties. Manipulating the phase separation behavior of RLP/PEG solutions, via changes in temperature and protein/polymer concentrations and compositions, was characterized via UV-Vis spectroscopy.

Confocal microscopy studies revealed the presence of microstructures captured during the crosslinking of the hydrogel; the bulk mechanical properties were probed via oscillatory rheological studies and the micromechanical properties via indentation

30 tests. This method provides a simple way of fabricating microstructured hydrogels without extensive manipulations, via aqueous-based LLPS and in situ crosslinking.

2.2 Materials and Methods

2.2.1 Materials Chemically competent cells of E. coli strain M15[pREP4] (for transformation of recombinant plasmids) and Ni-NTA agarose resin (for protein purification) were purchased from Qiagen (Valencia, CA). Amine-terminated, linear (5 kDa and 10 kDa) and 4-arm (10 kDa and 20 kDa) PEG was purchased from Creative PEG Works (Winston Salem, NC). The tri-functional cross-linker tris(hydroxymethyl phosphine) (THP) was purchased from Strem Chemicals (Newburyport, MA). Phosphate-buffered saline (PBS) was purchased from Mediatect (Manassas, VA). Deuterium oxide and NMR solvents were purchased from Cambridge Isotope Laboratories (Tewksbury, MA). All other chemicals were obtained from Sigma-Aldrich (St. Louis, MO) or

Fisher Scientific (Waltham, MA) and were used as received unless otherwise noted.

2.2.2 RLP Expression and Purification RLP protein expression and purification was conducted as previously reported by our laboratories.132,136–138 In brief, a single colony of E coli M15[pREP4] containing the desired RLP construct was inoculated in 100 ml sterile LB media containing 100 μg/mL antibiotics (ampicillin) and grown overnight. 100 mL of overnight culture media was subsequently used to inoculate 750 mL of 2xTYmedia (yeast 10 g/L, NaCl 5 g/L, tryptone 16 g/L) for protein expression. The 750 mL cultures were grown in a shaker at 37 °C until the OD600 reached 0.6-0.8, and then IPTG was added into final concentration of 1 mM to induce the protein expression for

31 4 hrs. Cells were harvested by centrifugation (5000 rpm for 15 mins at 4 °C), and the cell pellets were stored at -20 °C. The frozen cell pellets were lysed by freeze-thaw cycles and the lysed cell pellets were suspended in pH 8.0 native lysis buffer (50 mM

NaH2PO4, 300 mM NaCl, 10 mM imidazole) with 0.45 g lysozyme. Lysed cells were further disrupted via sonication on ice, using a Fisher Scientific model 500 Sonic Dismembrator (10-mm tapered horn) for 20 mins with a 10-second recovery time. The supernatant from centrifugation (10000 rpm for 60 mins) of cell lysate was collected, and pH was adjusted to 8.0, followed by incubation with Ni-NTA resin overnight at 4 °C. The resin was then loaded into a gravitational flow column, washed with native lysis buffer, native wash buffer (50 mM NaH2PO4, 300 mM NaCl, 20 mM imidazole, pH 8.0), and finally eluted with native elution buffer (50 mM NaH2PO4, 300 mM NaCl, 250 mM imidazole, pH 8.0). 100 mL elution fractions were carefully transferred and dialyzed (MWCO 10 kDa) against deionized water (5 L) at 4 °C with at least 6 changes of water before sterile filtration and lyophilization. The protein yield was approximately 30-50 mg per liter of cell culture.

2.2.3 Characterization of RLP/PEG Phase Separation RLP and PEG were dissolved into PBS at a 10 wt% (w/v) concentration. The two separated solutions were vortex mixed at designated RLP/PEG mass ratios (0/100,

25/75, 50/50, 75/25 and 100/0) at room temperature. The phase diagrams were determined by heating a series of RLP/PEG solutions from 4 oC to 60 oC at 1 oC intervals and observing when the solutions transitioned from turbid to clear. The samples were incubated at each temperature for 5 minutes and three replicates were performed. UV-Vis spectroscopy was used to characterize the concentration dependence of the phase separation. Samples comprising 10 wt% (w/v) PEG, RLP,

32 and 50/50 RLP/PEG in PBS were diluted sequentially by 0.5 wt% increments. Each solution composition was vortex mixed and the absorbance was measured at room temperature with a 10mm path length quartz cuvette at a wavelength of 600 nm to determine the turbidity of the solution.

2.2.4 Characterization of Equilibrium Concentrations The RLP and PEG were dissolved into d-PBS separately at a 10 wt% (w/v) concentration. The two stock solutions were mixed at designated RLP/PEG mass ratios (0/100, 25/75, 50/50, 75/25 and 100/0) at room temperature and allowed to phase separate overnight into two immiscible layers. Samples were carefully taken from the top and bottom layer to prevent mixing of the two liquids and were dissolved in deuterium oxide (D2O) that contained 0.01 mg/ml DSS as a reference. The concentration of each component was calculated from the proton NMR spectrum acquired (128 scans) with a Bruker AV 600 spectrometer (Bruker Daltonics, Billerica,

MA) under standard quantitative conditions.

2.2.5 Fluorescent Labeling of RLP and PEG RLP and PEG were labeled with Dylight-594 and Dylight-488 respectively. RLP or PEG was dissolved in PBS at a 5 mg/mL concentration. The fluorophore amine-reactive Dylight-594 or Dylight-488 was dissolved in anhydrous DMSO prior to addition to the reaction in a 1:100 molar ratio of NHS to RLP and was reacted for 4 hr at 25 °C (while protected from light). The protein was then dialyzed (MWCO 3.5 kDa) against deionized water (5 L) at 4 °C with at least 3 changes of water to remove the unreacted fluorophore, followed by filtration and lyophilization. The labeling reaction and purification were both protected from light. We did not anticipate the

33 labeling would significantly impact the crosslinking reaction due to the very low amount of the labeling of RLP.

2.2.6 Hydrogel Formation The RLP and PEG were dissolved into PBS separately at 10 wt% (w/v) concentration. The RLP and PEG stock solutions were mixed at RLP/PEG mass ratios (0/100, 25/75, 50/50, 75/25 and 100/0) at room temperature. The 0.6 μl THP crosslinker stock solution (100 mg/ml) was added to the solution of 30 μl 50/50 RLP/PEG precursors (1:3 molar ratio of THP:amine) and vortex mixed. The solutions were incubated at 37 oC in an incubator for 15 min for complete gelation. Hydrogel disks were formed in silicone chambers (45 mm diameter with a thickness of either 0.5 mm or 2 mm) and then were characterized via phase contrast optical microscopy or confocal microscopy for at least three hydrogel replicates. For the confocal microscopy experiments, fluorescently labeled RLP and PEG were used; hydrogels were formed by the same methods.

2.2.7 Oscillatory Rheology The oscillatory rheology experiments were conducted on AR-G2 rheometer (TA Instruments, New Castle, DE) with a 20 mm-diameter stainless steel parallel-plate geometry. The precursor solutions were prepared as described above. The 30 μl precursor solution was deposited on the 25 °C rheometer stage and the geometry was set at a 100 µm gap. Mineral oil was used to seal the geometry and prevented dehydration. The temperature was then quickly raised to 37 °C where it was maintained for the remainder of the experiment. The formation of elastic hydrogels was monitored using a time sweep conducted in the linear viscoelastic regime at 1%

34 strain and an angular frequency of 6 rad/s. This experiment was followed with a frequency sweep from 0.1 rad/s to 100 rad/s conducted at 1% strain. Experiments were repeated on three samples for each cross-linking ratio and the shear modulus was reported as the simple mean. The error is reported as the standard deviation of the samples tested.

2.2.8 AFM Indentation

Force and mechanical measurements were acquired using a Bruker Catalyst AFM. The micromechanical properties of RLP/PEG hydrogels were characterized via indentation (AFM) with a tip equipped with a 25 μm polystyrene (PS) probe, with a spring force constant of 0.06 N/m (NovaScan). The hydrogel was pre-formed between two glass slides (equipped with a 40-μm spacer) using the cross-linking protocol described in the sections above. The thin gels were hydrated in PBS buffer prior to the experiment and the hydrated gel thickness was approximately 40-60 μm. The gels were indented with the AFM probe until the force reached a 1 nN threshold, and the Young’s modulus was evaluated from fitting the force curve data to the Hertzian Spherical model.264–266 The Poisson’s ratio was assumed to be 0.5 for the hydrogel due to the high water content and the incompressibility of the gel samples.267 Over 100 indentations for each of 3 replicates were performed; analysis of the Young’s modulus from the data was conducted with the NanoScope Analysis software. The Young’s modulus data were fit to a Gaussian distribution or double-Gaussian distribution. The

Bruker Catalyst AFM was mounted onto a Zeiss AxioObserver inverted microscope to perform simultaneous bright-field microscopy and force spectroscopy; the light microscope allowed accurate positioning of the AFM probe on individual microdomains and other hydrogel features. Statistical analysis was performed using

35 the ANOVA Tukey-Kramer HSD test with JMP Pro software, and the threshold for statistical significance was set at p<0.05.

2.3 Results and Discussion

2.3.1 RLP/PEG Phase Separation Given the useful properties of RLP-based hydrogels,132,137,138,268,269 the widespread utility of PEGs in biomaterials, and the observation of phase separation in protein-PEG solutions,258 we sought to define the phase separation conditions of select aqueous compositions of RLP/PEG solutions, with the intent of defining conditions to generate RLP-containing microstructured matrices that comprised domains with distinct mechanical behavior. The RLP employed, as reported in our previous work, is a 23 kDa polypeptide containing 12 repeats of the putative consensus sequence (GGRPSDSYGAPGGGN) derived from Drosophila melanogaster and 5 repeats of lysine-rich crosslinking bundles

(GGKGGKGGKGG).132 Various PEGs (linear and 4-arm star, 5 kDa to 20 kDa in molecular weight) were selected owing to their widespread use in hydrogels.

Characterization of 50/50 RLP/PEG solutions at 25 oC via UV-Vis spectrophotometry, as a function of total polymer concentration, offers valuable information to guide the processing of these materials. Mixing a 10 wt% (w/v) solution of RLP and a 20 kDa 4- arm PEG-amine (at a 50/50 mass ratio) in PBS at room temperature results in an opaque solution due to immediate phase separation, with the micro-phase-separated domains growing larger over time, further evolving into two distinct, immiscible layers after incubation overnight at room temperature (Figure 2.1A). To define conditions under which phase separation would permit fabrication of a

36 microstructured hydrogel, the phase boundaries, critical concentrations, and critical temperatures of RLP/PEG solutions were determined. The turbidity of 50/50 (w/v) RLP/PEG solutions of various concentrations at room temperature was analyzed via UV-Vis spectroscopy at an absorbance of 600 nm (Figure 2.1B-C) in order to determine the critical concentration. As shown in Figure 2.1B, neither RLP nor PEG alone exhibit turbidity at any of the concentrations employed (0 wt% to 10 wt% total solids), indicating that they remain in a single phase under these conditions. In contrast, 50/50 solutions of RLP and 4-arm, 10 kDa PEG (PEG-4NH2-10k) became turbid at approximately 8 wt%, whereas an increase the molecular weight of the PEG to 20 kDa (PEG-4NH2-20k) decreases the concentration at which turbidity occurs to 4 wt% (Figure 2.1C). The trends observed for the 4-arm PEG are consistent with those observed for the the linear polymers; solutions of linear PEG-2NH2-5k in PBS become turbid at 12.5 wt% concentration, while those comprising a higher molecular-weight, linear PEG-2NH2-10k become turbid at 11 wt%. The reduction of the critical concentration with increasing PEG molecular weight is consistent with previous literature reports of phase separation in crystallin/PEG270,271 and BSA/PEG aqueous systems.272 Variations in the molecular architectures of PEGs also cause measureable differences in phase separation, with solutions of RLP with the 4-arm PEG-4NH2-10k showing turbidity at lower concentrations (8 wt%) than those comprising linear PEG-

2NH2-10k (11 wt%). The lower phase separation concentration of the multi-arm PEG is likely due to physical restriction of conformational changes that reduces the change of entropy of PEG in the system and thus favors the formation of different phases.273

37 A B 5 C 5 RLP/PEG-4NH2-20k RLP/PEG-2NH2-10k RLP/PEG-4NH2-10k RLP/PEG-4NH2-10k RLP/PEG-2NH2-5k 4 RLP 4 PEG-4NH2-10k Overnight 3 3

2 2

Abs(600) (a.u.) 1 Abs(600) (a.u.) 1 10wt% 50/50 RLP/PEG 0 0

0 2 4 6 8 10 0 2 4 6 8 10 12 14 16 wt% wt%

The phase separation temperature of 10 wt% RLP/PEG solutions (with various PEGs and at various RLP/PEG composition ratios) was determined by incubating the samples at temperatures ranging from 4 oC to 60 oC, and observing the temperature at which the solution transitioned from turbid to clear via UV-Vis spectroscopy. In general, the RLP/PEG solutions phase separated into two distinct liquid phases (LLPS) below a given transition temperature and were miscible above the transition temperature (Figure 2), thus indicating that the RLP/PEG solutions exhibit UCST

(upper critical solution temperature) behavior. The PEGs (PEG-4NH2-20k, PEG-

4NH2-10k, PEG-2NH2-10k and PEG-2NH2-5k), which are generally higher in molecular weight than those used in other phase separation studies,270–272,274 should promote phase separation (e.g., exhibit higher UCSTs), yielding two-phase mixtures at the targeted ambient and physiological temperatures for hydrogel formation. Solutions of 10 wt% RLP or PEG alone maintained optical clarity (indicating a single phase)

38 throughout the entire temperature range (4-60 oC). The behavior of the RLP alone is slightly different than that in previous reports by Dutta, Choudhury and co-workers, in which Rec1-resilin and An16-resilin alone were shown to exhibit reversible UCST and LCST (lower critical solution temperature) transitions in aqueous buffers.135,275 It is also different than the behavior demonstrated by Li et al., in which LCST behavior of the RLP was observed in various aqueous buffers, without the observation of UCST behavior.276 In contrast, for the RLP/PEG solutions, no LCST behavior was observed under the conditions investigated here, although all RLP/PEG solutions showed a detectable UCST, with phase separation occurring below a critical temperature and miscibility above this temperature. As shown in Figure 2.2, the UCST for all of the 10 wt% RLP/PEG samples was observed at RLP/PEG compositions of 25/75 or 30/70 (depending on PEG architecture). While the UCST was approximately 55 oC for solutions of 30/70 RLP/PEG-4HN2-20k; the critical temperature decreased with decreasing molecular weight of the PEG, with transition temperatures of 30 oC for

o 25/75 RLP/PEG-4NH2-10k (and PEG-2NH2-10k) and 15 C for RLP/PEG-2NH2-5k. In contrast to the 25/75 RLP/PEG solutions, the 75/25 RLP/PEG solutions were more miscible across all conditions. For example, the 25/75 RLP/PEG-4HN2-20k exhibited a UCST of only 15 oC, and all remaining 75/25 RLP/PEG compositions comprising lower molecular weight PEGs (PEG-2NH2-5k, PEG-2NH2-10k, PEG-4NH2-10k) were miscible over the entire temperature range of the experiment.

39 70 RLP/PEG -4NH2 (20k)

C) RLP/PEG -4NH2 (10k) o

( 60 RLP/PEG -2NH2 (10k) RLP/PEG -2NH2 (5k) 50

40 Clear 30

10 Turbid Transition TemperatureTransition 0 0 20 40 60 80 100 % RLP

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 (▼)

Phase separation in these systems can be tuned, as expected, by variations in temperature and RLP/PEG ratio. In addition, the critical phase transition temperature

(a UCST for these RLP/PEG solutions) is increased by increasing the molecular weight of the PEG, consistent with results reported for crystalline/PEG,270,271,277,278 lysozyme/PEG,279 and BSA/PEG272 systems. The impact of PEG architecture, however, as assessed for the 10 kDa PEG, does not appear to affect the critical temperature significantly, which is consistent with modeling studies that showed that variations in star-polymers are expected to have a less significant impact on phase

40 separation temperature compared to linear polymers.273,280 The asymmetry of the phase diagram (i.e., the observation of the critical transition temperature at RLP/PEG ratios lower than 50/50) indicates the uneven entropic contribution from the RLP and PEG in the energy of mixing.281 The critical temperatures were higher for the RLP/PEG solutions reported here compared to the reported behavior for other aqueous protein/polymer solutions, in which the transitions often occur below room temperature.270,271 This is likely because of the higher PEG concentration and molecular weight employed in our work relative to those in the previous studies; it is also possible that the protein/polymer compatibility may be less favorable in RLP/PEG solutions compared to other proteins. The higher critical temperatures are advantageous for processing these hydrogels, given that crosslinking of a microstructured material can be conducted at physiological or ambient temperature. Our data taken together demonstrate that the phase separation of these RLP/PEG systems can be easily tuned by temperature, concentration and ratio of RLP/PEG, as well as the PEG molecular weight.

2.3.2 Co-existence Concentration of RLP/PEG Development of the co-existence (binodal) curve and tie lines connecting the two equilibrium phases, as shown in Figure 2.3, indicates the set of initial solution conditions that will yield a distinct pair of liquid phases. After the RLP/PEG-4NH2- 20k solutions completely phase separated into two immiscible layers, a clear phase I

(PEG-rich) and yellowish phase II (RLP-rich) were observed (Figure 2.3A). Co- existence curves for the RLP/PEG solutions were determined via 1H NMR spectroscopic evaluation of the phases; 10 wt% RLP/PEG solutions in deuterated PBS were incubated at room temperature overnight. Quantitative 1H NMR was used to

41 determine the RLP and PEG concentrations in both phase I and phase II (Table 2.1); the characteristic protons from the RLP and PEG are both clearly observed. For the

50/50 RLP/PEG-4NH2-20k solution, the PEG-rich phase I contained 63.9 ± 2.8 mg/ml PEG and 22.9 ± 7.6 mg/ml RLP, and the RLP-rich phase II contained 7.6 ± 4.2 mg/ml PEG and 175.9 ± 23.2 mg/ml RLP. When the RLP/PEG ratio was increased from 50/50, the difference in the concentrations of RLP and PEG in each of the final phases was correspondingly reduced (i.e., the RLP concentration in the upper PEG-rich layer increased, while it decreased in the lower RLP-rich phase), which will permit tuning of the concentrations of the phases based on initial conditions.272,277 The PEG concentrations observed in the equilibrium phases are in the range of those reported in the literature for dextran/PEG solutions, in which higher PEG concentrations (> 50 mg/ml PEG) were observed in phase I and lower PEG concentrations (< 20 mg/ml PEG) were observed in phase II.281,282

42 A B 100 Inital conc. 80 PEG-rich phase RLP-rich phase

60 Phase I 40 PEG (mg/ml)

Phase II 20

0 0 50 100 150 200 250 RLP (mg/ml)

Table 2.1 Concentrations of RLP and PEG in the upper (phase I) and lower (phase II) phases of RLP/PEG solutions of various compositions.

Phase I Phase II RLP PEG RLP PEG (mg/ml) (mg/ml) (mg/ml) (mg/ml) 25/75 6.5 83.4 230.0 19.2 50/50 14.5 65.4 170.7 12.3 50/50 25.4 65.8 155.8 6.0 50/50 28.8 60.6 201.3 4.5 75/25 46.9 32.9 96.5 14.5

43 The co-existence curve for RLP/PEG solutions of different initial concentrations was constructed from the equilibrium concentrations of the individual separated phases; the tie line permits the choice of any given starting concentration that will yield RLP-rich and PEG-rich phases with concentrations indicated by the binodal curve; variations in the starting composition can be used to control the volume of the resulting phases. The volume fraction of each phase (φI and φII) can be 퐵퐶 퐴퐶 determined from the tie line according to the lever rule,244 휙퐼 = and 휙퐼퐼 = , 퐴퐵 퐴퐵 where A refers to the composition in Phase I, B refers to the composition in Phase II, and C is the initial composition on the phase diagram. The production of solutions with concentrations on the tie line permits manipulation of the volume fraction and composition of the phases, and also (see below) the domain sizes of the microstructures in RLP-PEG hydrogels. The volume fraction of RLP-rich phase II was determined from the co-existence curve to be 0.087, 0.235 and 0.552 in the 25/75, 50/50 and 75/25 RLP/PEG solutions, respectively. The RLP and PEG concentrations in the co-existing phases were lower in protein and PEG concentrations compared to those reported for BSA/PEG272 and crystallin/PEG,270,271 due to the higher molecular weight of PEG employed in our studies, but provide a large range of compositions for producing microstructured hydrogels. The use of this information to design the composition and structures of RLP/PEG hydrogels is consistent with the reported use of the co-existence curve of Dextran (Dex)-PEG to select the size of Dex-rich particles fabricated from a Dex-PEG aqueous phase-separated system;244 however, the crosslinking of the bulk materials directly after phase separation was not studied. Taken together, these data illustrate the manner in which the aqueous LLPS of RLP/PEG can be tuned by temperature, initial concentration, and ratio of RLP/PEG.

44 The equilibrium concentrations are highly tunable with the starting compositions of the RLP/PEG; by selecting the initial concentrations of the RLP and PEG, the final concentration and volume fraction of each phase can be tuned, and thus the composition of each domain when crosslinked in the hydrogels.

2.3.3 Microstructured Hydrogels Microstructured heterogeneous hydrogels were chemically crosslinked under conditions in which aqueous LLPS was observed (Figure 2.4); PEG-4NH2-20k was selected in these experiments due to its greater extent of phase separation compared to that of the solutions containing a lower molecular weight PEG. The fast, one-step, Mannich-type reaction of a small-molecule crosslinker, tris(hydroxymethyl phosphine) (THP), with primary amines from both RLP lysine residues and PEG amine end groups, was used to crosslink and capture the microstructure during phase separation. The reaction allowed the crosslinking between RLP and PEG, and within the RLP-rich and PEG-rich domains independently, to provide a stable network and to permit capture of micron-size domains in the resulting hydrogels. The resulting materials are distinct from other studies employing LLPS to produce microstructured matrices in that elastomeric microdomains of various mechanical strength can be produced, rather than the microporous scaffolds that have been the subject of most other reports.234,283,284

45 Lys NH2 NH2 Lys Lys Mixed Mixed RT PBS Lys NH2 NH2 THP Incubated (crosslinker) 37oC Hydrogel RLP PEG-4NH2 RLP/PEG 1:1 NH :OH in PBS 2

The microstructure of hydrogels with various RLP/PEG ratios was characterized via confocal microscopy. Fluorescently labeled RLP (Dylight-594) and PEG (Dylight-488) were used to form the RLP-PEG hydrogels by mixing the two components with THP at room temperature, then allowing gelation at 37 oC for 15 min; green fluorescence thus indicates the PEG-rich phases and red indicates the RLP- rich phase. The hydrogel morphology and the microstructure can be controlled by the RLP/PEG ratio and compositions, coupled with the use of rapid crosslinking, yielding hydrogels with domain compositions and sizes (5-50 μm) that can be tuned depending upon the volume fraction of the separated phases (Figure 2.5). For example, two initial compositions on the same tie line (Figure 2.5A; (i), 50 mg/ml RLP and 58 mg/ml PEG and (ii), 80 mg/ml RLP and 50 mg/ml PEG) were selected to form hydrogels. As indicated in Figure 2.5B, the size of the RLP-rich domains (red) was smaller in composition (i) (21.1 ± 11.7 μm) compared to those in composition (ii) (93.5 ± 45.6 μm), which is consistent with the fact that the volume fraction of the

RLP-rich phase of composition (i) ( 휙퐼퐼 = 0.16 ) was lower than that of composition

46 (ii) (휙퐼퐼 = 0.31). The broad distribution of the domain sizes in the gels may result from an unstable interfacial energy between the phases; stabilizing this interface could reduce the distribution of microdomain sizes and will be investigated in future studies. Nevertheless, these data illustrate that the RLP/PEG solutions show well-behaved phase separation, that the solutions can be crosslinked into hydrogels, and that the size and composition of the domains can be controlled via choice of initial solution compositions.

A 100 B 50 i ii Inital conc. 80 PEG-rich phase RLP-rich phase 40 i 60 ii 30

40 20 i Volume %

PEG (mg/ml) ii

20 10

0 0 0 50 100 150 200 250 0 20 40 60 80 100 120 140 160 180 200 RLP (mg/ml) Particle Sizes (um)

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 2.3.4 Hydrogel Bulk and Micromechanical Properties The bulk and micromechanical properties of the phase-separated composite hydrogels were also investigated. Oscillatory rheology was used to characterize the gelation time and storage modulus of the hydrogels (Figure 2.6). The fast gelation time was indicated by the storage modulus crossover with the loss modulus in under 5 minutes; the gels attain a steady modulus (620 ± 150 Pa) within an hour (see Supporting Information). These data confirm that hydrogels with tunable domain sizes can be formed via the THP crosslinking chemistry and that the shear modulus of the hydrogels is relevant for applications in soft tissue engineering (0.1-10 kPa).285 That the shear storage modulus of the RLP-PEG gel was lower than that of 10 wt% PEG- only (7590 ± 1730 Pa) and 10 wt% RLP-only hydrogels (3300 ± 1200 Pa), may be a result of heterogeneous crosslinking arising from the phase separation of RLP and PEG, resulting in dense domains and a loose matrix that can be easily deformed by shear forces. Manipulation of chemical protocols for crosslinking, which are under investigation, may be used to address these properties and further evaluate the mechanical properties of the coexisting RLP-rich and PEG-rich phases with the ideal mixing rule.

48 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 microdomains, resulting in a reduced interfacial area, but nevertheless, the gels attain a steady modulus after 60 minutes.

The micromechanical properties of the hydrogel domains were characterized via atomic force microscopy (AFM) indentation. An AFM tip with a 25 μm spherical probe was employed to determine that there are distinct regional mechanical differences within the RLP-PEG hydrogels (Figure 2.7).286–288 A large-sphere probe was required to provide a greater contact surface area and stronger signal for force measurement on these soft materials; however, the resolution was thus reduced in scanning mode. Solutions comprising 10 wt% (w/v) RLP, 10 wt% (w/v) PEG, or 10 wt% (w/v) RLP and 20 kDa 4-arm PEG (at a 50/50 ratio) were crosslinked between

49 two glass slides with THP at a 1:3 THP:amine molar ratio according to the methods described above. To minimize optical scattering within the hydrogels, thin hydrogels (40-60 μm) were formed in order to observe the domains via phase contrast optical microscopy and to identify target locations for indentation. The Young’s modulus determined from each indentation is presented in Figure 2.7, showing the distribution of values for 10 wt% RLP, 10 wt% PEG, and also for the RLP-rich domains (Figure 2.7A) and the PEG-rich matrix (Figure 2.7B) of 10 wt% 50/50 RLP-PEG microstructured hydrogels. The 10wt% RLP and 10wt% PEG gels show Gaussian distributions with means of 17.0 ± 3.6 kPa and 7.8 ± 5.4 kPa respectively. The compressive Young’s moduli determined via AFM were higher compared to those determined from the shear rheology data, likely due to the fact that thin gels were required for the AFM experiments in order to observe the domains and the matrix clearly. Substrate effects might thus be significant, yielding a surface elastic modulus that tends to be higher compared to the bulk modulus.289 Nevertheless, the data clearly show that the RLP and PEG thin hydrogels have distinct mechanical properties that are similar to those of the bulk matrices and that can be probed by microscale mechanical methods (Figure 2.7C-D). The data for the RLP-PEG hydrogels indicate the presence of a PEG-rich matrix with a Gaussian distribution of moduli values

(mean of 3.1 ± 0.7 kPa), as well as RLP-rich domains that exhibited a double- Gaussian distribution of moduli with maxima at 5.3 ± 2.3 kPa and 15.6 ± 8.1 kPa (representing 58% and 42% of the overall distribution respectively) (Figure 2.8). The large degree of dispersity of the Young’s moduli for the RLP-rich domains likely results from contributions from the PEG-rich continuous phase, which is present at the surface and complicates the measurements of the RLP-rich discontinuous phase.267,290

50 A C 120 D 40 PEG RLP PEG-rich RLP-rich 100 Matrix Domain * 30 80 * 20 60

B 40 10

20 Young's Modulus (kPa) 0 Normalized Probability (%) Probability Normalized 0 0 10 20 30 PEG-rich RLP-rich PEG RLP Young's Modulus (kPa) Matrix Domain

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.

Interestingly, microstructured hydrogels have not, to our knowledge, been reported from other phase-separating protein/PEG systems,250 which is likely a result of the focus of those studies on the purification and crystallization of globular proteins. In addition, microstructure like that observed in the RLP-PEG hydrogels has not been reported in other protein-polymer based hybrid hydrogels (e.g., gelatin- PEG49,105 or ELP-PEG291) perhaps as a result of differences in the physicochemical properties of the ELP compared to RLP as well as the use of lower molecular weight PEGs in those studies. The hydrophobic character of PEG supports its favorable interactions with unfolded protein at higher temperatures, reducing the sizes of the

ELP hydrophobic aggregate (~1μm in sizes) and improving the transparency of the ELP-PEG hydrogels.292 The size of the microdomains in the RLP-PEG systems are comparable to those observed in dextran-PEG aqueous LLPS, which exhibited diameters in the range from 2.5-20 μm.244,293 The simple methods reported here offer

52 the advantages of fully aqueous processing conditions, as well as the ability to generate microstructured hydrogels in a single step; there is no need in the RLP-PEG system for additional processing. Furthermore, microstructured hydrogels that contain elastomeric microdomains offer opportunities to systematically study the role of domain properties and interfaces in the behavior of soft materials. In addition, the length scales of the domain sizes in the RLP-PEG hydrogels are comparable to microstructures employed for tissue regeneration,219,294,295 thus offering opportunities to correlate the impact of microdomains on cell spreading and proliferation.

2.4 Conclusions The aqueous LLPS of RLP/PEG solutions provides a simple, one step fabrication method to form microstructured hydrogels that does not require toxic organic solvents nor extensive purification or extra washing steps. The aqueous-based LLPS of RLP/PEG can be tuned by temperature, concentration, and the ratio of RLP to PEG in initial solutions. The microstructure of hydrogels can be controlled via judicious selection of the initial concentrations and compositions of the RLP and PEG, which are guided by the phase diagram for RLP/PEG solutions. By carefully selecting the composition (e.g., near the phase boundary) and volume fractions of each of the phases, these approaches have the potential to create hydrogels with various additional heterogeneities, including biocontinuous structures. The selection of the small molecule crosslinker THP that can crosslink both RLP and PEG-amines provided stable crosslinking in both RLP-rich and PEG-rich phases. Because the microstructure of the hydrogel can be captured by the fast crosslinking kinetics and relatively slow phase separation kinetics, it should be possible to control the microstructures by varying the relative kinetics of phase separation and crosslinking, to provide hydrogels

53 with different micromechanical environments with benefit for controlling mechanical properties and cellular behavior.

54 Chapter 3

MICROSTRUCTURED ELASTOMER-POLYETHYLENE GLYCOL HYDROGELS VIA KINETIC CAPTURE OF AQUEOUS LIQUID-LIQUID PHASE SEPRARATION

3.1 Introduction

The heterogeneity and biophysical properties of native extracellular matrix (ECM) are essential in regulating cell behavior and guiding tissue regeneration. Microstructured hydrogels that mimic the complexity of ECM thus have emerged as useful materials for controlling material mechanical properties, diffusion of macro- and biomolecules, and mammalian cell behavior.170,175,233,294,296 A variety of evidence suggests that heterogeneity in hydrogels can promote cell growth and organization in three dimensions (3D); indeed, macroporous scaffolds and hydrogels containing degradable microparticles have been shown to promote osteogenic differentiation in vitro and bone tissue formation in vivo.211,297,298 Furthermore, the mechanical properties of hydrogel materials, including stiffness, elasticity, and viscoelastic properties, impact the differentiation of mesenchymal stem cells as well as tissue regeneration.285,299–302 Hydrogel geometries and surface curvature also influence cell morphology and migration, thus providing additional materials handles for regulating gene expression and cell functions,303,304 independently of bulk mechanical properties.

These studies together indicate the importance of engineering hydrogel microstructures with independently tunable microscale structure and micromechanical properties for controlling cell behavior. Accordingly, multiple production strategies

55 have been pursued, such as photopatterning,223,224,226,295,305 selective degradation,226–228 and the incorporation of microparticles169,170 synthesized via emulsion polymerization162,171–173 and/or microfluidic239 technologies. However, these methods suffer from either low throughput, as in the case of photopatterning, or multiple processing steps, as in particle fabrication. Liquid-liquid phase separation (LLPS), a well-known phenomenon based on the unfavorable interactions of two dissimilar liquids (e.g., solutions of polymers or biopolymers), results in phase separation of solutions even at a low concentration,306– 308 and aqueous two-phase solutions (ATPS) have thus been widely employed for the purification of proteins, polysaccharides, and DNA.281,309,310 In addition to these purification applications, LLPS also provides a highly versatile, and as yet underexplored, platform for fabricating microstructured materials. Phase behavior can be modified by factors such as temperature, polymer molecular weight, polymer concentration, ionic strength, pH, and addition of specific salts.308,311,312 The timescales over which LLPS occurs range from seconds to minutes near the critical point,313 and crosslinking of the solutions within this time frame can enable the generation of microstructured hydrogel materials.314 In such approaches, the use of solutions with comparable kinetics of phase separation and crosslinking can yield, from ATPS, materials with microstructures of various length scales.174,214,244 Our group recently reported such methodology for the single-step fabrication of microstructured, elastomeric hydrogels with distinct micromechanical properties, via the aqueous LLPS of resilin-like polypeptides (RLPs) and poly(ethylene glycol) (PEG) with stable crosslinking in both RLP-rich and PEG- rich phases.315 The rubber-like elasticity of RLPs, including low stiffness (0.6-2 MPa),

56 high extensibility (~ 300%), and efficient energy storage (> 90% resilience),251–253 provides distinct mechanical properties compatible with biomaterial applications for mechanically active tissues.216,254–257 Despite our demonstration of the one-step generation of microstructured elastomers, however, independent manipulation of the microstructure and mechanical properties of the previously reported RLP-PEG hydrogels was not possible, given the interdependence of the relative kinetics of phase separation and chemical crosslinking during hydrogel formation. The crosslink density

(affecting mechanical properties) and crosslinking kinetics (affecting microstructure) thus could not be independently and easily tuned. Given the demonstrated importance of microstructure in guiding cell behavior, we thus sought to develop alternative approaches that would facilitate the formation of RLP-based elastomeric hydrogels with independently tunable domain diameters, domain compositions, and bulk and domain mechanics. Accordingly, we developed a photo-crosslinkable RLP-PEG LLPS system to enable on-demand control of microstructure in hydrogels across a range of compositions of target mechanical properties. Although photo-crosslinking reactions have been adapted widely in the formation of microstructured materials, they have been employed largely in the production of microporous hydrogels, mesoporous organohydrogels, microparticles and copolymerized materials.316–319 The development of methods that permit selective initiation of crosslinking of both phases in an ATPS thus offer substantial new opportunities to systematically control the encapsulation of microgels within a continuous matrix, providing needed flexibility for tuning (micro)mechanical properties as well as a vast parameter space for customization.211,219,298

57 We report here the facile production of photo-crosslinkable RLPs and demonstrate the feasibility of producing a photo-crosslinked ATPS with RLP- Acrylamide (RLP-Ac) and PEG-Acrylate (PEG-Ac) phases. The rapid photo- crosslinking enabled the capture of microstructures with select domain properties – diameter, composition, and mechanics – as shown by multiple microscopic and spectroscopic methods. Confocal microscopy permitted facile characterization of the average domains sizes in a hydrogel after various periods of incubation prior to crosslinking, and broadband coherent anti-Stokes Raman scattering microspectroscopy (BCARS) was used to analyze the compositions of the hydrogel domains and continuous phases throughout the demixing process. The mechanical properties of the various hydrogels were analyzed via oscillatory rheology and compared with the rule of mixtures, and AFM indentation was used to confirm the variations in the micromechanical properties of the microstructured materials. Finally, our studies also demonstrated that mesenchymal stem cells can be localized in select regions of the cell-compatible scaffolds.

3.2 Materials and Methods

3.2.1 Materials Chemically competent cells of E. coli strain M15[pREP4] (for transformation of recombinant plasmids) and Ni-NTA agarose resin (for protein purification) were purchased from Qiagen (Valencia, CA). Acrylate-terminated 4-arm (20 kDa) PEG was purchased from Jenkem. Phosphate-buffered saline (PBS) was purchased from Mediatect (Manassas, VA). Deuterium oxide and NMR solvents were purchased from Cambridge Isotope Laboratories (Tewksbury, MA). All other chemicals were obtained

58 from Sigma-Aldrich (St. Louis, MO) or Fisher Scientific (Waltham, MA) and were used as received unless otherwise noted.

3.2.2 RLP Expression and Purification RLP protein expression and purification was conducted as previously reported by our laboratories.132,136–138 The protein yield was approximately 30-50 mg per liter of cell culture.

3.2.3 RLP Functionalization and Characterization The RLP proteins were functionalized with acrylamide groups via modification of regularly positioned lysine residues on the polypeptide chain. First, the RLP proteins were dissolved in PBS (10mg/ml). The acrylic acid N-hydroxysuccinimide ester (NHS-Ac) was dissolved in dimethyl sulfoxide (DMSO) in 50mg/ml separately and drop-wise added into the RLP solution. The ratio of NHS-Ac to lysine was varied depending upon the desired functionality of the conjugate. The reaction was stirred at room temperature for approximately 4h. This reaction solution was diluted 8 times with DI water to prevent precipitation and dialyzed (Snakeskin, 3.5kDa, Thermo Scientific) against DI water at 4 oC (in a cold room) to remove byproducts and DMSO. The purified RLP-Ac was filtered and lyophilized and stored at -20 oC prior to experiment. The functionality of the RLP-Ac was characterized via 1H NMR spectrometry.

The purified RLP-Ac (~2mg) was dissolved in (600μl) D2O (Cambridge Isotope Laboratories, Tewksbury, MA) and analyzed using an AVIII 600MHz NMR spectrometer (Bruker Daltonics, Billerica, MA). The protons from the eight phenylalanine residues per RLP molecule were used as an internal reference for the

59 quantification of acrylamide group functionality. The integration of the aromatic

1 protons of phenylalanine ( H NMR (600 MHz, D2O, δ): 7.15–7.40 (m, 5H)) was compared to the integration of the vinylic protons of the acrylamide that resulted from

1 the reaction of the acrylamide and lysine amine groups ( H NMR (600 MHz, D2O, δ): 5.65–6.30 (d, 3H)).

3.2.4 Characterization of RLP-Ac/PEG-4Ac Phase Separation

RLP-Ac and PEG-4Ac were dissolved in phosphate-buffered saline (PBS) at various concentrations from 10wt% to 20wt%. The two solutions (at 50/50 RLP- Ac/PEG-4Ac mass ratios) were vortex mixed at room temperature. The phase diagrams were determined by heating a series of RLP/PEG solutions from 4 oC to 60 oC at 1 oC intervals and observing when the solutions transitioned from turbid to clear; the transition point was identified as the point of 50% transmittance. UV-Vis spectroscopy was also used to characterize the concentration dependence of the phase separation at room temperature. Samples comprising 15wt% (w/v) 50/50 RLP- Ac/PEG-4Ac in PBS were diluted sequentially by 0.5 wt% increments. Each solution composition was vortex mixed and the absorbance was measured at room temperature with a 10mm path length quartz cuvette at a wavelength of 600 nm to determine the turbidity of the solution. Although phase separation was observed in all solutions comprising the PEG-4Ac, it should be noted that variations in PEG-4Ac viscosity were apparent in different lots of PEG-4Ac ordered from the same manufacturer; these viscosity differences affect phase separation kinetics, so all comparisons here are made between separate samples, but those employing PEG-4Ac of similar viscosity.

60 3.2.5 Characterization of Equilibrium Concentrations The RLP-Ac and PEG-4Ac were dissolved into d-PBS separately at a 10, 15, 20 wt% (w/v) (w/v) concentration. The two stock solutions were mixed at 50/50 RLP- Ac/PEG-4Ac mass ratios at room temperature and allowed to phase separate overnight into two immiscible layers. Samples were carefully taken from the top and bottom layer to prevent mixing of the two liquids and were dissolved in deuterium oxide

(D2O) that contained 0.01 mg/ml DSS as a reference. The concentration of each component was calculated from the proton NMR spectrum acquired (128 scans) with a Bruker AVIII 600 MHz NMR spectrometer (Bruker Daltonics, Billerica, MA) under standard quantitative conditions.

3.2.6 Hydrogel Formation The RLP-Ac and PEG-4Ac were dissolved into PBS separately at 10 wt% (w/v) concentration. A stock solution of the photo-initiator LAP was prepared in PBS at a concentration of 13.4 mg/ml. The RLP-Ac and PEG-4Ac stock solutions were mixed at 50/50 RLP-Ac/PEG-4Ac mass ratios at room temperature. 5μl of LAP solution was added to 100μl RLP-Ac/PEG-4Ac mixture and vertex mixed the samples to obtain 2.2mM LAP in the precursor mixture. An UVP Blak-Ray B-100AP high Intensity UV lamp (UVP, Upland, CA) with 365nm wavelength ~ 5mW/cm2 intensity was irradiated for 4 min on the hydrogels in 0, 5 and 10 min after vertex mixing. The UV intensity was confirmed with a radiometer.

3.2.7 Oscillatory Rheology The oscillatory rheology experiments were conducted on an AR-G2 rheometer (TA Instruments, New Castle, DE) with an attached UV Light Guide accessory and UV lamp source (OmniCure S2000 (Excelitas)), with an 8 mm-diameter stainless steel

61 parallel-plate geometry. The precursor solutions were prepared as described above. The 10 μl hydrogel precursor solution was deposited on the quartz rheometer stage and the geometry was set at a 200 µm gap. Mineral oil was used to seal the geometry and prevented dehydration of the hydrogel. The 365nm UV with 5 mW/cm2 intensity was applied in 0, 5 or 10 min to induce UV crosslinking. The mechanical properties of the hydrogels were measured in the linear viscoelastic regime where the modulus is independent of the level of applied stress or strain. The gelation of hydrogels was monitored using a time sweep conducted in the linear viscoelastic regime at 1% strain and an angular frequency of 6 rad/s. This experiment was followed with a frequency sweep from 0.1 rad/s to 100 rad/s conducted at 1% strain and amplitude sweep from 0.1% to 1000% strain. Experiments were repeated on three samples for each condition and the shear modulus reported as the simple mean. The error is reported as the standard deviation of the samples tested.

3.2.8 Polymerization Yield The precursor solutions were prepared in deionized water and a 20 μl hydrogel precursor solution was crosslinked with UV irradiation for either 2 min or 4 min. The hydrogels was then incubated with 400 μl of deionized water for 2 hrs. The water was then removed from the hydrogels and lyophilized both supernate and hydrogels separately.

3.2.9 Confocal Microscopy and Domain Diameter Analysis Hydrogel disks were formed in silicone chambers with 5 mm diameter and 0.5 mm thickness, by the same methods above, for confocal microscopy analysis of domain sizes. The confocal Z-stack images were acquired with a Zeiss 780

62 multiphoton microscope (Carl Zeiss, Inc., Thornwood, NY). A Chameleon Vision II Multiphoton laser with a 755nm wavelength was used to excite the autoflorescence of the RLP, and the NDD detection system was used for imaging the multiphoton florescence. The domain diameters were analyzed with Volocity software.

3.2.10 BCARS Imaging and Data Analysis The custom-built BCARS setup has been described in detail by Billecke et al.320. Briefly, a commercial laser source (Leukos-CARS, Leukos) delivered 1 ns pulses of a 1064 nm laser with 32 kHz repetition rate which was split to serve as pump/probe beam and to pump a photonic crystal fiber generating a super-continuum (>100 µW nm-1, 1100-1600 nm) employed as the Stokes beam. Both beams were overlapped in the focal plane of an inverted microscope (Eclipse Ti-U, Nikon) equipped with a xyz piezo stage (Nano-PDQ 375 HS, Mad City Labs) using a 100X, 0.85 NA (Olympus) objective. The spectral signal was collected in forward direction via an air objective (M-10x, 0.25 NA, Newport), where the laser lines were removed with a notch (NF03-532/1064E-25, Semrock) and a shortpass filter (FES1000, Thorlabs) before being dispersed by a spectrometer (Shamrock 303i, 300 lines mm-1, 1000 nm blaze, Andor), and finally collected on a deep-depletion CCD (Newton, DU920P-BR-DD, Andor). Several stitched images (4 tiles, 250 nm steps, 101x101 pixels per tile) with 500 ms pixel dwell time were collected per sample with a spectral range between 600 and 3400 cm-1.

Raw spectra were treated with a modified Kramers-Kroenig algorithm321,322 and an error phase correction via an interactive noise-maintained approach323 made model-free using a second-order Savitky-Golay filter (404 cm-1 window size) to remove non-resonant contributions and to generate a Raman-like spectrum from the

63 imaginary component of the third-order susceptibility. Data analysis as well as the generation of images for specific Raman frequencies was done in IgorPro 6.34

(Wavemetrics). Images of the PEG- and RLP-rich phases were generated by integrating over 10 wavenumbers around 1468 cm-1 and 1660 cm-1 and then dividing the two images. These images were then used to identify 10 regions within each phase, where spectra were extracted with a 3x3 pxel bin. Then the region between 1350 cm-1 and 1700 cm-1 was fitted in Matlab (Mathworks) with a sum of Lorentzian functions as summarized in Table 3.1 and a ratio of the respective peak areas were formed to follow the composition of the two phases with crosslinking time.

Table 3.1 Fitting parameters used on Raman-like spectra. Seed values with lower/upper bounds.

Peak position [cm-1] FWHM [cm-1] Vibration 324 1384 (1382-1386) 12 (10-15) CH3 325 1412 (1408-1416) 12 (10-15) CH2 wagging 324 1440 (1435-1445) 12 (10-15) CH2 deformation 326 1468 (1460-1470) 12 (10-15) CH2 bending 1607 (1600-1615) 12 (10-15) Ring modes phenylalanine 324 1666 (1650-1670) 12 (10-40) Amide I 325

3.2.11 AFM Indentation

64 to the experiment; the hydrated hydrogel thickness was approximately 40-60 μm. The hydrogels were indented with the AFM probe until the force reached a 1 nN threshold, and the Young’s modulus was evaluated from fitting the force curve data to the

Hertzian Spherical model.264–266 The Poisson’s ratio was assumed to be 0.5 for the hydrogel due to the high water content and the incompressibility of the hydrogel samples.267 Over 100 indentations for each of 3 replicates were performed; analysis of the Young’s modulus from the data was conducted with the NanoScope Analysis software. The Bruker Catalyst AFM was mounted onto a Zeiss AxioObserver inverted microscope to perform simultaneous bright-field microscopy and force spectroscopy; the light microscope allowed accurate positioning of the AFM probe on individual microdomains and other hydrogel features.

3.2.12 Swelling Ratio The precursor solutions were prepared as described above. Hydrogels were then incubated with PBS overnight. The weight of swollen hydrogels (Ws) were recorded, and the hydrogels were then washed with deionized water 3 times and lyophilized to obtain the dry weight (Wd). The swelling ratios were calculated as

Ws/Wd.

3.2.13 Cell Encapsulation and Viability The hMSCs (Lonza, MD) were encapsulated in 10wt% 50/50 RLP-MMP-

RGD-2Ac/PEG-4Ac hydrogels at a cell density of 1 x 106 cells/ml (hMSC). The precursors were vortex mixed and then pipette mixed with the cells before being deposited onto silicone chambers that were 5 mm in diameter and 1 mm in thickness, which were then placed under a 365nm UV lamp (UVP, Upland, CA) and irradiated at

65 5mW/cm2 for 2 min. The hydrogels were then placed into the cell growth media from the MSCGM bullet kit for (hMSC, Lonza, MD) and placed in a cell culture incubator at 37 oC, 5% CO2. Live/Dead stain (Life Technologies) was utilized to determine the viability of the hMSCs. The hydrogels were washed in PBS and placed in PBS containing 2 mM Calcein AM and 4 mM ethidium homodimer-1 for 20 min at 37 oC, 5% CO2. Cells were then washed twice with PBS and imaged while still alive. The cell–gels were then imaged via laser scanning confocal microscopy on a Zeiss LSM

710 microscope (Carl Zeiss, Inc., Thornwood, NY). The excitation (Ex) and emission (Em) waveleght of Calcein AM (Ex 488 nm/Em 490-552 nm), ethidium homodimer-1 (Ex 561 nm/Em 560-735 nm). Z-stack images were acquired from hydrogels and representative images in maximum intensity projections are reported.

3.2.14 Statistical Analysis Statistical analysis was performed using the ANOVA Tukey-Kramer HSD test with JMP Pro software, and the threshold for statistical significance was set at p<0.01.

3.3 Results and Discussion

3.3.1 Design and Synthesis of Photo-crosslinkable RLP-Ac The RLP employed in these studies is a 23 kDa polypeptide containing 12 repeats of the putative consensus sequence (GGRPSDSYGAPGGGN) derived from Drosophila melanogaster and 5 repeats of lysine-rich bundles (GGKGGKGGKGG) that can be used for crosslinking or RLP functionalization.132 The RLPs were expressed following procedures extensively employed in the Kiick laboratories132,136– 138 and were functionalized with acrylamide groups to facilitate the desired on-demand photo-crosslinking of microscale domains. Chemical modification of RLP with N-

66 acryloxysuccinimide (NHS-Ac) via reaction of lysine residues yielded RLP-Ac (Figure 3.1A) via the protocols detailed in the experimental section; The degree of modification was confirmed via 1H NMR; the appearance of the 3 vinylic peaks at δ 5.65-6.30 ppm327 indicated the successful functionalization of the RLP-Ac, and comparison of the area of these peaks to that of the aromatic protons from phenylalanine (δ 7.15-7.40 ppm) permitted determination of the degree of acrylation of the RLP (Figure 3.1B). The degree of acrylamide functionalization can be easily varied via modulation of reaction stoichiometry, where NHS-Ac:Lys molar ratios ranging from 0.2 to 2 yield RLP with 2 to 10 acrylamide groups (RLP-2Ac to RLP- 10Ac, Figure 3.1C and Table 3.2) per chain. Further increases in the NHS-Ac:Lysine ratio (up to a ratio of 4) did not result in any additional increase in the number of acrylamide reacted per RLP. Although there are 15 lysine residues present in each RLP chain, these are distributed in short (GGKGGKGGKGG) ‘domains’ at regular intervals in the RLP sequence; the close proximity of the lysines in these short domains possibly results in steric hindrance that prohibits complete coupling to all lysinses. Morevoer, the competing hydrolysis reaction in aqueous conditions328 is almost certainly the origin of the plateau in the degree of functionalization of the RLP. Nevertheless, the NHS-mediated acrylation yielded a sufficiently wide range of acrylation (2Ac to 10Ac) to test the impact of acrylation on the phase separation and mechanical properties of the crosslinked RLPs.

67 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.

Table 3.2 RLP-Ac functionalization

# Ac per RLP % Functionalization RLP-2Ac 2.2 +/-0.5 14.9 +/-3.6 RLP-4Ac 4.0 +/-0.5 26.7 +/-3.1 RLP-6Ac 6.7 +/-0.4 44.7 +/-2.9 RLP-10Ac 9.9 +/-0.5 66.0 +/-3.0

3.3.2 Liquid-liquid Phase Separation of RLP-Ac/PEG-Ac LLPS has been observed in a variety of biopolymer solutions, such as those including gelatin and dextran. In addition, elastin-like polypeptide (ELP) solutions undergo LLPS with an increase in temperature, which is driven by a preference for

68 homotypic self-interactions over protein–solvent interactions,329 and RLP solutions of various compositions have also been demonstrated to show similar behavior. In our studies reported here, the phase separation of ternary solutions of RLP-Ac and commercially available PEG-4Ac was confirmed via UV-Vis spectrophotometric analysis (Figure 3.2). PEG-induced phase separation of proteins and polysaccharides is widely reported in the literature (e.g., crystallin/PEG270,271 and BSA/PEG aqueous systems272) and has been a mainstay in the purification of protein-based biopharmaceuticals.330 In these previous studies, as well as in our previous investigation of RLP/PEG solutions,315 the use of PEGs of higher molecular mass drives phase separation more owing to a less favorable entropy of mixing in comparison to PEGs at similar concentrations but with lower molecular masses. However, when the PEG molecular mass is too large, the volumetric swelling is significant, thereby adding substantial, and in some cases unwanted, external stresses to the material.50 Therefore, 20kDa PEG-4Ac was selected as the second component in the RLP-Ac LLPS system in order to balance the ability to drive phase separation while avoiding excessive swelling. Moreover, 20kDa PEG-4Ac gels are mechanically appropriate for soft tissue applications50,105,291 (due to the multi-arm PEG molecular architecture and its moderate molecular weight). Turbidity (OD600) observed in 50/50 RLP-Ac/PEG-4Ac solutions upon phase separation was used as an indicator of LLPS and was measured at 25 oC as a function of total polymer concentration from 0 to 15wt% (Figure 3.2A).

Consistent with our previous investigations in which turbidity was present for

315 RLP-NH2/PEG-NH2 solutions employing 4-arm, 20kDa PEGs, turbidity was also observed in RLP-Ac/PEG-4Ac solutions for all of the various functionalized RLP-Ac,

69 regardless of the number of acrylamide groups per RLP chain. The observed trends in the LLPS behavior of these RLP-Ac/PEG-4Ac were similar to those observed for the

315 RLP-NH2/PEG-4NH2 solutions, i.e., with phase separation occurring at high polymer concentrations and miscibility at lower concentrations. The phase separation concentration for a given solution was defined as that at the total concentration at which the transmittance was reduced to 50%.331,332 Differences in the number of acrylamide groups (X) reacted per RLP chain lead to slight variation in the phase separation concentration (between 6.5-9.5 wt% for X between 2 and 10) as illustrated in Figure 3.2A, which did not show any clear trend. The phase separation concentration was also determined for solutions of RLP-XAc (where X=2 or 6) with PEG-4Ac with different mixing ratios (Figure 3.2B-C); the phase separation wt % of polymer increased from 7.7% for 25/75 RLP-XAc/PEG-4Ac to 11.3% for 75/25 RLP- XAc/PEG-4Ac.

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 LLPS ultimately yields two immiscible, coexisting phases at equilibrium. An understanding of the compositions of the two phases as a function of the initial concentrations of RLP and PEG permits determination of the binodal curves and tie lines, which then affords opportunities to predict phase compositions and relative phase volumes as a function of initial solution compositions. The coexistence curve for RLP-6Ac/PEG-4Ac solutions of different initial concentrations (10, 15, and 20wt%) was constructed from the equilibrium concentrations of the individual separated bulk phases from 50/50 RLP-6Ac/PEG-4Ac solutions determined via 1H NMR as previously reported.315 The coexistence curve of RLP-6Ac/PEG-4Ac is shown in Figure 3.3A; the coexistence concentrations of RLP and PEG indicated the stronger partitioning of RLP into phase-separated domains (with at least 9.5-fold and up to 50-fold greater RLP concentration than that in the PEG-rich phase) compared to the concentration of PEG in the PEG-rich phase (just 2.5 – 5 times the concentration of PEG as that in the RLP-rich phase) (Figure 3.3A). We also tested if the number of acrylamides on RLP chains affects phase separation equlibria with 10 wt% RLP- 2Ac/PEG-4Ac and found statistically identical concentrations in each phase (Figure 3.3B). The similarity in the coexistence concentrations for the two RLP-XAc shows that differences in the number of acrylamide groups of the RLP does not significantly affect the coexistence concentrations of RLP and PEG in the RLP-rich and PEG-rich phases. The volume fraction of the of PEG-rich (top phase I, ϕI) and RLP-rich (bottom phase II, ϕII) phases can be determined from the tie line according to the lever rule,244 퐵퐶 휙퐼 = (1) 퐴퐵 퐴퐶 휙퐼퐼 = (2) 퐴퐵

71 where A refers to the composition in the PEG-rich phase, B refers to the composition in the RLP-rich phase, and C is the initial composition on the phase diagram. The volume fraction of the RLP-rich phase (ϕII) was determined, from the coexistence curve for the 10wt% 50/50 RLP-XAc/PEG-4Ac, to be 0.22 ± 0.05 for RLP-2Ac and 0.16 ± 0.04 for RLP-6Ac; the volume fraction of the PEG-rich phase (ϕI) was 0.89 ± 0.09 and 0.86 ± 0.07 for RLP-2Ac and RLP-6Ac, respectively. Both RLP-2Ac and RLP-6Ac show a lower volume fraction for the RLP-rich phase versus the PEG-rich phase, which will result in an RLP-rich dispersed and PEG-rich continuous phase. Statistical analysis from the ANOVA Tukey-Kramer HSD test illustrates that the volume fractions of the RLP-rich phase (p-value 0.69) and PEG-rich phase (p-value 0.98) were statistically similar regardless of the number of acrylamide groups linked to the RLP.

72 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.

3.3.3 Characterization of Microstructure and Bulk Mechanics of Microstructured Elastomers The possibility of employing photo-crosslinking of the phase-separated RLP- Ac/PEG-4Ac solutions to permit temporal control over the microstructure of resulting hydrogels was evaluated (Figure 3.4A). Photochemical methods were employed to crosslink solutions of functionalized RLP-XAc with PEG-4Ac via incorporation of the biocompatible photo-initiator lithium phenyl-2,4,6-trimethylbenzoylphosphinate

73 In situ oscillatory shear rheology with a UV accessory was used to characterize photo-crosslinking gelation kinetics upon application of UV light (365nm, 5mW/cm2, for a duration of 10 min) and to measure the shear storage moduli of the crosslinked phase-separated RLP-PEG hydrogels at time points 0, 5 or 10 min after vortex mixing. Time-sweep rheology data for a 10wt% 50/50 RLP-6Ac/PEG-4Ac hydrogel show in Figure 3.4B that the storage modulus (G’) increased and stabilized rapidly with gelation, which occurred in less then 30 sec upon application of UV light, although irradiation for longer than 30 sec resulted in corresponding increases in final storage modulus; a plateau in the storage modulus was observed within 4 min. The data illustrate that the final shear storage modulus can be modulated over the range of 1 kPa to 17 kPa by varying the duration of UV irradiation from 30 sec to 10 min (Figure 3.4C), regardless of at what point the UV irradiation is applied after mixing. In addition, there was no significant change in the gelation time or final plateau storage modulus (Figure 3.4D, G’ values of 16.5 ± 4.0 kPa, 17.1 ± 3.2 kPa, and 17.1 ± 5.4 kPa at 0, 5 and 10 min) as a function of the point at which the UV irradiation was applied, suggesting that the extent of phase separation does not alter the bulk mechanical properties of these materials although the microstructures are trapped out of equilibrium. Furthermore, the storage modulus of RLP-2Ac/PEG-4Ac hydrogels was similarly insensitive to the variation in times of crosslinking, with a G’ of approximately 7.7 ± 1.7kPa at for hydrogels crosslinked at various timepoints after mixing (Figure 3.4D). The overall lower moduli in the RLP-2Ac/PEG-4Ac relative to the RLP-6Ac/PEG-4Ac hydrogels is almost certainly a result of the lower acrylamide functionality and consequently lower crosslinking density of the RLP-2Ac compared to RLP-6Ac.

74 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.

Multiphoton microscopy was employed to image microstructures that developed in these hydrogels; visualization of the domains was enabled by the autofluorescence of the RLP phase with 755 nm excitation. Phase contrast optical microscopy confirmed that both pure RLP and pure PEG hydrogels lacked any microstructure after crosslinking (Figure 3.5). Mixing of the 50/50 RLP-XAc/PEG- 4Ac solutions in PBS immediately results in an opaque solution for both RLP-2Ac and RLP-6Ac; the evolution of domains captured by UV crosslinking at 0, 5 and 10 mins

75 are shown in Figure 3.6A. In both solutions, the diameters (and distribution of diameters) of the RLP domains increased with time. The 10 wt% 50/50 RLP- 6Ac/PEG-4Ac hydrogels developed RLP-rich domains with an initial diameter of 10.2 ± 3.4µm (volume fraction of 0.06) after immediate UV crosslinking; the domains increase in diameter to 14.5 ± 6.0 µm (volume fraction of 0.11) when photo- crosslinked 10 min after mixing. In contrast, the evolution of the diameters of the RLP domains in 10 wt % 50/50 RLP-2Ac/PEG-4Ac solutions occurred more rapidly, with diameters of 12.8 ± 5.5µm (volume fraction of 0.08) upon immediate crosslinking and 67.1 ± 20.3µm (volume fraction of 0.12) when photo-crosslinked at 10 min after mixing (Figure 3.6B). The particle size distribution data in Figure 3.6C also delineate the more rapid growth of RLP-rich domains over time for the RLP-2Ac versus RLP- 6Ac-containing solutions.

76 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.

The phase separation kinetics of both systems was distinct while the equilibrium coexistence concentrations were almost identical.335 The evolution of the coalescence of the domains is balanced by gravitational, flotational, and frictional forces,336 and thus the rate of the phase separation is a function of the relative density,

78 interfacial tension, and viscosity of both phases, which do not influence equilibria.

Given the similarities in the compositions of the phases for the RLP-2Ac and the

RLP-6Ac, the slower rate of domain growth in the RLP-6Ac/PEG-4Ac solutions is thus suggested to arise from a lower interfacial tension for the RLP-6Ac, likely as a result of increased hydrophobic interactions and miscibility between the RLP-6Ac and the PEG-4Ac (relative to RLP-2Ac). Analysis of the normalized particle size distribution data in Figure 3.6C suggests a sharp asymmetric peak for the RLP-6Ac/

PEG-4Ac samples, consistent with an Ostwald ripening model,337 while the RLP-2Ac/

PEG-4Ac samples are more consistent with the Smoluchowski model, with a size distribution independent of the volume fraction338 (Figure 3.6C), suggesting quantitative differences in the RLP-2Ac and RLP-6Ac phase separation kinetics that can be exploited to generate microscale domains of various sizes in these phase- separated hydrogels. Interestingly, as clearly indicated by the data in Figure 3.4D, the storage moduli of these materials were insensitive to the size of the microstructures that evolved over 0, 5 and 10 min prior to crosslinking (Figure 3.6A and 3.6B), which is particularly notable for the RLP-2Ac solutions. To first order, the modulus of composite materials depends on the volume fraction and the mechanical properties of each phase according to the rule of mixtures,339 with the upper bound storage modulus (GU) and lower bound storage modulus (GL) related by

퐼 퐼 퐼퐼 퐼퐼 퐺푈 = 퐺 휙 + 퐺 휙 (3) 1 휙퐼 휙퐼퐼 = 퐼 + 퐼퐼 (4) 퐺퐿 퐺 퐺

79 where 퐺 is the storage modulus and 휙 is the volume fraction in phase I (top, PEG-rich) and phase II (bottom, RLP-rich). The similarity in the moduli of these materials at different crosslinking times is consistent with the observed relatively small changes in the volume fractions of the phases over time; indeed, the fraction of the PEG-rich phases remains at ca. 0.90 over the 10-minute time period of the experiments. The shear storage modulus for each phase was determined by photo- crosslinking of each of the individual macroscale phases on the rheometer after overnight incubation of RLP-XAc/PEG-4Ac solutions. As shown in Figure 3.7, the shear storage moduli for the PEG-rich phase (퐺퐼) from the RLP-2Ac/PEG-4Ac and RLP-6Ac/PEG-4Ac solutions were 10.4 ± 0.9 kPa and 14.3 ± 0.4 kPa, respectively; while the storage moduli for the RLP-rich phase (퐺퐼퐼) from the RLP-2Ac/PEG-4Ac and RLP-6Ac/PEG-4Ac solutions were 3.1 ± 0.3 kPa and 30.5 ± 5.1 kPa, respectively. The higher storage modulus in RLP-6Ac was expected owing to the higher crosslinking density afforded by the higher degree of acrylation for RLP-6Ac, as noted above. By applying the rule of mixtures (Equation 3 and 4) to these phase-separated hydrogels, the calculated G’ values for RLP-6Ac (15 - 17 kPa) and RLP-2Ac (6.4 - 9.9 kPa) matched well with the measured values (RLP-6Ac 16.8 ± 3.8 kPa and RLP-2Ac 7.7 ± 1.7kPa). The difference between the storage moduli in the RLP-rich and PEG- rich phases suggests that the micro-mechanical properties will be heterogeneous in the microstructured hydrogels, which could be used to selectively and locally promote cell proliferationa and/or differentiation within the domains or matrix.301,302,340

80 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.

3.3.4 Distinct Composition and Mechanics of Domains in Micostructured Elastomers Characterization of microstructured hydrogels has been previously reported via cryogenic electron microscopy (cryo-EM),341–343 small angle neutron scattering (SANS),319,341,344 and coherent anti-Stokes Raman scattering (CARS) microscopy;291,345,346 Given the simplicity and non-invasiveness of the technique, broadband coherent anti-Stokes Raman scattering microspectroscopy (BCARS) was employed to assess the relative compositions of the domains and continuous phases of

RLP/PEG solutions as a function of demixing time during the phase separation process for the 10 wt% 50/50 RLP-6Ac/PEG-4Ac hydrogels; representative results (images and spectra) are shown in Figure 3.8. Raman-like (RL) spectra were produced from the BCARS dataset collected from hydrated hydrogels, for which the

81 intensities then are linearly related to the concentration of the polymer probed. In order to quantify the RLP and PEG content in different phases of the RLP-PEG mixtures in situ in a microstructured hydrogel, pure RLP and PEG spectra (Figure 3.8A) were first acquired to determine suitable Raman bands for selective imaging of these molecules. The RL spectra of PEG showed good correspondence with conventional Raman spectra in the fingerprint and CH-stretching regions.326,347 A comparison of the features of the RLP and PEG spectra indicated that a spectral ratio

of I1468 cm-1 (δCH2)/I1660 cm-1 (Amide I) accurately captures the [PEG]/[RLP] ratio. This characteristic Raman peak ratio was used to analyze the composition of the RLP- and PEG-rich regions of the RLP-6Ac/PEG-4Ac hydrogels as a function of incubation time. Figure 3.8B shows the intensity distribution of the vibration at 2930 cm-1, representing the density of CH3 bonds, which is reflective of the regions with large

[RLP] due to its relatively large number of CH3 groups. These results show similar domain diameters as observed with multiphoton microscopy (Figure 3.6D); the corresponding RL fingerprint spectra of the RLP-rich phase and PEG-rich phase within the hydrogels (normalized to the maximum value in the amide I region (~1660 cm-1)) are shown in Figure 3.9A-B and show clear spectral differences in the RLP and PEG rich domains. Figure 3.10A shows a map of the RLP-rich phase via plotting of

I1468 cm-1/I1660 cm-1 in the 10wt% 50/50 RLP-6Ac/PEG-4Ac hydrogels crosslinked at 0, 5, and 10 min after mixing. These BCARS images illustrate the multiphase structure of these mixtures (consistent with Figure 3.8B) wherein RLP-rich domains (that still contain PEG) grow larger with time within a PEG-rich continous phase that becomes enriched with PEG during the phase separation process.

82 -1 The 1468cm (δCH2) peak was observed to increase in the PEG-rich phase and decrease in the RLP-rich phase over time, indicating that the PEG polymer was partitioning out of the RLP-rich droplets and into the PEG-rich matrix. Plotting the peak ratio I1468 cm-1/I1660 cm-1, allowed for quantification of the phase separation with mixing time (Figure 3.10B), confirming previous observations (from domain size ripening) that the PEG-rich phase appears to achieve equilibrium after 10 minutes. The low and relatively consistent spectral ratio observed in the RLP-rich phase is likely a result of the high RLP concentration in the RLP-rich phase, which may obscure changes of the I1468 cm-1/I1660 cm-1 ratio. Both the spectral data and the ratiometric analysis are consistent with the 1H NMR analysis (Figure 3.3A) and suggest that at equilibrium there is almost no PEG in the RLP-rich domains while the PEG-rich continuous phase contains a significant amount of RLP. The relatively constant [RLP]/[PEG] ratio in the RLP-rich phase over the first 240 minutes suggests that the RLP-rich domains first grow in size (coalesce) and then later exclude the PEG. The heterogeneous PEG-rich phase, in contrast, appears to always contain a certain amount of RLP. In concert with photocrosslinking, BCARS microspectroscopy provides an important capability for facile observation of the spatial organization and chemical compositions of phases during the course of phase separation.

83 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 3.10 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).

The micromechanical properties of the hydrogel domains were characterized via atomic force microscopy (AFM) indentation. Thin hydrogels (40-60 μm) were formed to minimize optical scattering within the hydrogels to observe the domains with phase contrast microscopy and to identify target locations for indentation. An AFM tip with a 1 μm spherical probe was employed to determine the mechanical different in distinct regions within the RLP-PEG hydrogels286–288 and reduce the

85 substrate effects that would occur by probing a thin hydrogel with a large probe.267,348 The Young’s modulus determined from each indentation is presented in Figure 3.11A, showing the distribution of values for the RLP-rich domains and the PEG-rich matrix for 10 wt% 50/50 RLP-6Ac/PEG-4Ac microstructured hydrogels crosslinked at 0, 5 and 10 min. The hydrogels photo-crosslinked in 0 min show similar Young’s moduli for both the domains and matrix, with means of 1.5 ± 0.3 kPa for the RLP-rich domains and 1.4 ± 0.3 kPa for the PEG-rich matrix (Figure 3.11B; p = 0.28).

Likewise, the Young’s moduli of the hydrogels photo-crosslinked at 5 min show similar values. In contrast, hydrogels photo-crosslinked at 10 min showed RLP-rich domains (2.2 ± 0.5 kPa) that are mechanically distinct from the PEG-rich matrix (1.7 ± 0.9 kPa; p < 0.01), with differences in modulus that are qualitatively consistent both with the differences observed in bulk rheology measurements of the individual phases (Figure 3.7) and with the difference in composition indicated by the BCARS data at later timepoints (Figure 3.9). The significantly lower Young’s modulus measured from AFM indentation versus that predicted from the bulk oscillatory rheology measurement is consistent with previous studies,349 and likely arises from the confinement of water in a hydrogel during bulk rheological measurements; this confinement increases the resistance to deformation and consequently the modulus. In contrast, the lack of confinement, coupled with the small force applied during the AFM indentation, minimizes resistance to deformation.349 Although evaluation of the local water content and swelling within the domains and matrices was not possible to measure in the RLP-PEG hydrogels, the swelling ratio for the bulk microstructured hydrogels as well as that of the crosslinked PEG-rich phase and crosslinked RLP-rich phase (each separately, isolated after bulk phase separation and then crosslinked) in

86 PBS is shown in Figure 3.12. The RLP-rich phase exhibited a lower swelling ratio compared to that of the PEG-rich phase, as might be expected based on the known swelling of PEG-based hydrogels;350,351 the higher water content of the PEG phase is expected to lower the mechanical properties of the PEG-rich phase relative to that of the RLP-rich phase. Indeed, AFM measurements of the bulk RLP-rich and PEG-rich bulk phases crosslinked after overnight phase separation, yield Young’s moduli (2.7 ± 0.2 kPa and 1.1 ± 0.4 kPa, respectively, (indicated by the data marked Equil in Figure

3.11B) that are perfectly consistent with those measured from the domains and matrix in the microstructured RLP/PEG hydrogels. These data thus illustrate the power of the combined BCARS and micromechanical measurements; the evolution of the phases and their compositions over time can be visualized and related to the micromechanical properties within the phase-separated hydrogels.

87 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 the measurement of swelling ratios.

3.3.5 Cell Viability and Growth in 3D LLPS Microstructured Elastomers Previous reports have suggested the promise of hierarchically structured hydrogels for directing cell-matrix interactions (spreading, migration and adhesion),72,219,258 as well as promoting mesenchymal stem cells toward osteogenic differentiation.211,297,298 Thus, the cytocompatibility of the RLP/PEG hydrogels was evaluated via 3D encapsulation of human mesenchymal stem cells (hMSC). The hydrogel precursors were first vortex mixed alone and then gently pipette mixed with the cells to minimize cell death induced by the high shear stress of vortex mixing. The rapid photocrosslinking of the hydrogels permitted cell encapsulation with good cell viability and cell distribution. The 10wt% 50/50 RLP-MMP-RGD-2Ac/PEG-4Ac was used in the cell encapsulation studies owing to the wider range of domain diameters accessible; the RLP-2Ac contained a matrix metalloproteinase (MMP)-senstive peptide for cell-mediated degradation and an RGD cell-adhensive ligand to facilitate

89 integrin-mediated cell attachment. Cell viability was evaluated via staining with calcein and ethidium homodimer and confocal imaging at day 1 and day 7 as illustrated in Figure 3.13. The hMSCs maintained an exceptionally high cell viability of 95% (Figure 3.13A and 3.13B) within the microstructured hydrogels over a period of 7 days, with spreading clearly observed by day 7. Interestingly, the cells showed elongated morphologies only around the RLP-rich domains (Figure 3.13B). Similar cell viability was observed for hMSCs encapsulated in RLP-MMP-RGD-6Ac/PEG-

4Ac hydrogels, as shown in Figure 3.13C-D, as well as for human microvascular endothelial cells in both RLP-6Ac and RLP-2Ac systems (data not shown). The high viability and organization of hMSCs around the RLP-rich domains demonstrates the promise of employing these organized RLP/PEG hydrogels to localize cells via methods that could be extended to multiple cell types and co-cultures.

90 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).

3.4 Conclusions Photo-crosslinking methods were exploited as a facile method to capture morphologically, chemically, and mechanically distinct phases in microstructured hydrogels during LLPS of RLP-Ac/PEG-4Ac solutions. Evaluation of the LLPS for RLPs with various degrees of acrylamide functionalization established that equilibrium phase diagrams were not significantly affected by the degree of functionalization. Photo-triggered crosslinking of RLP-Ac/PEG-4Ac during the phase separation permitted the production of RLP/PEG hydrogels with RLP-rich domains with various diameters; a higher degree of RLP acrylation reduced the rate of domain growth, presumably by increased miscibility mediated by hydrophobic interactions

91 between the RLP-6Ac and PEG-4Ac. Controllable photocrosslinking permits the modulation of the modulus via UV exposure time, allowing for on-demand and independent tuning of microstructure and mechanical properties. Significant differences in the compositions (and thus mechanical properties) of the developing domains and continuous phase were indicated to occur only after 10 minutes of phase separation, as indicated by BCARS microscopy and AFM; interestingly, the microstructured matrices exhibit bulk mechanical properties that correspond to the rule of mixtures theory and do not vary over time. Furthermore, the materials demonstrated spatial localization of multiple cell types, at high viabilities, around RLP-rich domains. Overall, the LLPS of RLP-Ac/PEG-4Ac, when captured via photo-crosslinking, permits independent tuning of the microstructure and micromechanical properties that can be used to design complex materials for biomedical and other applications. The high cell viability and capability to guide cell organization within the microstructured hydrogels indicates their potential use in regenerative medicine applications.

92 Chapter 4

MICROMECHANCICAL PROPERTIES OF MICROSTRUCTURED ELASTOMER-POLYETHYLENE GLYCOL HYDROGELS

4.1 Introduction

The mechanical properties of hydrogel materials, including stiffness, elasticity, and viscoelastic properties, has been well known to influence cell morphology, migration, differentiation for tissue regeneration.285,299–302 Microstructured hydrogels that mimic the heterogeneity and biophysical properties of native extracellular matrix

(ECM) have been developed as useful materials for controlling local mechanical properties, diffusion of macro- and biomolecules, and mammalian cell behavior.170,175,233,294,296 Heterogeneity in hydrogels has been shown to promote cell growth and organization in three dimensions (3D),211,297,298 and further regulate gene expression and cell functions.303,304 However, local mechanical properties of those complex biomaterials and microstructure materials have not been well studied.

Resilin-like polypeptides (RLPs) exhibit unique rubber-like elasticity properties,251–253 which provide distinct mechanical properties compatible with biomaterial applications for mechanically active tissues.216,254–257 Our recent study of the micromechanical characterization of RLP hydrogels highlighted the effects on local mechanical properties at various polypeptide concentrations and demonstrated the unique elastomeric features, with maintenance of excellent resilience and toughness, above the overlap polymer concentration (c*) for resilin-based hydrogels, confirming their used for applications in regenerative medicine. Our group has also recently reported the single-

93 step fabrication of microstructured, elastomeric hydrogels with distinct micromechanical properties, via the aqueous liquid-liquid phase separation of resilin- like polypeptides (RLPs) and poly(ethylene glycol) (PEG), with stable crosslinking in both RLP-rich and PEG-rich phases,315 and independently tunable microscale structure and micromechanical properties.

To study the influence of micromechanical properties with microstructured hydrogels, we employed oscillatory shear rheology, small-strain microindentation, and large-strain indentation and fracture ranging from different length scales. Oscillatory shear rheology and small-strain microindentation measure the small-strain elastic response of hydrogels. Large-strain indentation and puncture provide insight of mechanical response at large deformations as well as the resistance for fracture initiation.352 Microindentation and puncture studies require minimum sample volume, exhibit ease of implementation, and have been used on various other synthetic polymer gels and biological tissues.352–354 We were able to observe the indentation and deformation of the microstructured hydrogels with confocal microscopy, and, in addition, we characterized the mechanical response, indenter dependence and extracted a fracture energy of the microstructure hydrogels.

4.2 Materials and Methods

4.2.1 Hydrogel Formation

Resilin-like polypeptides were expressed and purified as previously described with a final product yield of approximately 20-30 mg/L after dialysis and lyophilization. The RLP-Ac were functionalized with acrylamide groups via modification of regularly positioned lysine residues on the polypeptide chain as in

94 Chapter 3. The RLP-2Ac and PEG-4Ac were dissolved into PBS separately at 10 wt

% (w/v) concentration. The RLP/PEG precursor solution containing 2.2 mM photoinitiator LAP was prepared as reported elsewhere.355 The precursor solution was vortex mixed for 30 sec and was either immediately (0 min) crosslinked, or allowed to phase separate for 5 min, prior to photo-crosslinking with UV light. A UV lamp

(Thorlabs, Newton, NJ) with 365 nm wavelength ≈5 mW cm−2 intensity was used to irradiate the samples for 5 min; the UV intensity was confirmed with a radiometer. The hydrogels were formed in circular silicone molds with 8 mm diameter and 2 mm thickness for rheology and small-strain microindentation characterization and puncture mechanics. Hydrogels were incubated in PBS for overnight hydration before testing.

4.2.2 Oscillatory Rheology The oscillatory rheology experiments were conducted on an AR 2000 rheometer using 8 mm diameter stainless steel parallel-plate geometry. The swollen hydrogels were prepared as described above. The preformed hydrogels were mounted on the rheometer plates at room temperature (20ºC), and the geometry was set at a gap at which the normal force equalled 0.02 N to prevent slippage. Frequency sweep experiments, from 0.01-10 Hz, were performed at a fixed strain amplitude of 1%, which is in the linear viscoelastic regime of the hydrogels as determined by a strain sweep at a fixed frequency of 1 Hz. Shear storage moduli (G′) and loss moduli (G″) were recorded over the entire frequency range with ten data points per decade.

Experiments were repeated at least on three individual, separate samples for each type of hydrogel.

95 4.2.3 Microindentation Hydrogels were prepared as described above. A piezo-controlled linear actuator was used to bring a cylindrical probe of radius, a = 1000 m, in contact at a constant displacement rate of 10 m/s with thin circular disks of hydrogels (thickness, h~2mm), is well known that soft polymer gels often have strain rate-dependent mechanical properties.356–359 To facilitate comparison of low-strain elastic moduli between small-strain microindentation and rheology, a small indentation rate was chosen in order to yield quasi-static loading conditions during the experiments. The average strain rate for microindentation was computed from the indentation rate and the radius of the probe used, (average strain rate = (10 μm/sec)/(1000 μm)) and was calculated as ~0.01Hz. Upon loading and unloading, the relative displacement, , and resulting force, P, was measured with a custom-designed load cell. Average stresses were calculated from the force normalized by the cross-sectional area of the probe 푃 (휎 = ) 휋푎2 and average strains were calculated by normalizing the displacement by the radius of the probe (휖 = 훿/푎) for hydrogels (h~2000μm). Repeated loading and unloading cycles were performed for all hydrogels to measure the reversible energy storage capability or resilience at different peak loads (or applied strains). Resilience is defined as the ability of a material to undergo significant deformations with minimum energy loss, and it is measured by dividing the area under the unloading curves by the area under the loading curves. The experiments were repeated at least 3 times on three different samples at each peak load. In order to quantify an effective linear elastic moduli, Eindentation for the three hydrogel concentrations, a linear approximation was employed up to ~10% strain for the force- displacement curves obtained from the loading regime of the stress-strain cycles. The

96 effective elastic modulus, Eindentation, was then determined from the slopes of the linear fits at low strain using the following equation:360 (1 − 휐2) 푎 푎 3 −1 퐸 = { 1 + 1.33 ( ) + 1.33 ( ) } 푖푛푑푒푛푡푎푡푖표푛 2퐶푎 ℎ ℎ where the compliance, C, (퐶 = 휕훿/휕푃), is the inverse of the slope of the linear force- displacement curve (during loading); ʋ is the Poisson’s ratio (which was approximated as 0.5 assuming incompressibility under time scales of loading361 ); and h is the thickness of the hydrogel sample. The average of the slopes obtained from the loading curves was used to quantify C. Note that equation (1) includes corrections for finite values of a/h, which deviate from the classical constraints of Hertzian contact mechanics.360,362

4.2.4 Puncture For puncture experiments, the experimental set-up with the piezo-controlled linear actuator was utilized to perform large-strain indentation to the point of failure with hollow, needles with a tip radius, R, equal to 5-35 μm. Upon indentation, the force increases as the needle deforms the hydrogel until fracture occurs at a peak puncture force, Pc. Fakhouri et al have proposed semi-empirical relationships to analyze the force-displacement response during loading until the point of fracture in terms of the material's elastic modulus, E, and the needle tip geometry.352 The deformation behavior of a material is modeled by Hertzian contact mechanics under small strains, and a Neo-Hookean model under large strains to develop the ‘puncture equation’352

푃 = 푘″퐸푅훿 + 푘′퐸훿2 where P is the resistive force on the indenter,  is the indentation depth, R is the indenter tip radius, kʹ is an empirical constant, and kʺ is a constant depending upon

97 the indenter tip geometry, obtained from Hertzian contact mechanics for flat and spherically tipped indenters.363

4.3 Results and Discussion

4.3.1 Hydrogel Formation and Characterization The photo-crosslinking of the phase-separated RLP-2Ac/PEG-4Ac solutions that permit temporal control over the microstructure of resulting hydrogels was evaluated in Charter 3.355 Photochemical methods were employed to crosslink solutions of functionalized RLP-2Ac with PEG-4Ac via incorporation of the biocompatible photo-initiator lithium phenyl-2,4,6-trimethylbenzoylphosphinate

(LAP) and irradiation with 365nm light;333,334 crosslinking should therefore occur both throughout the individual phases as well as across the phase-separated interface in the hydrogel and easily allow domains of targeted diameters to be captured during the course of phase separation. The RLP-PEG hydrogels, crosslinked immediately (0 min) or after 5 or 10 min after mixing, were hydrated overnight in PBS, and the domain sizes of the hydrogels were characterized as described in Chapter 3. The

RLP-PEG hydrogels with RLP-rich domains of ~10 µm resulted with cross-linking upon immediate crosslinking, or ~25 µm when cross-linked after 5 min and ~60 µm after 10 min incubation (after mixing).355 The surface area is 4πr2 and volume is

4/3πr3; therefore, surface area to volume ratio is 4πr2/(4/3πr3) = 3/r. The surface area to volume ratio of the RLP-rich domains in different RLP-PEG hydrogels is shown in

Figure 4.1C. In nanocomposite systems, higher surface area to volume ratio provides a higher interface density, which is important to mechanical properties.364 The mechanical and/or micromechanical properties of the microstructured RLP-PEG hydrogels were compared with oscillatory shear rheology, small-strain

98 microindentation, along with large-strain indentation and puncture at length scales comparable with the domain sizes (Figure 4.2).

10 um 25 um 60 um 25

10 % Distribution

0 0.0 0.5 1.0 1.5 2.0 2.5 3.0 3.5 4.0 Surface area:Volume (1/m)

Figure 4.1. Evolution of domain diameters in hydrated RLP-PEG hydrogels. Autofluorescence images of hydrated RLP-PEG hydrogels with precursor 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

4.3.2 Oscillatory Shear Rheology Oscillatory shear rheology was employed to characterize the mechanical properties of preformed, equilibrium swollen RLP-PEG hydrogels under shear, and frequency sweep experiments were conducted to obtain the shear storage moduli, G′

(averaged for n=3). The average shear storage moduli (G′) and loss moduli (G″) are plotted as a function of frequency over 0.01-10 Hz in Figure 4.3A. The storage modulus of RLP-2Ac/PEG-4Ac hydrogels was insensitive to the variation in times of crosslinking, with a G’ of approximately 7.7 ± 1.7kPa for hydrogels crosslinked at various timepoints after mixing (Chapter 3)355. Upon overnight hydration in PBS, the hydrogels swelled up to 2 times (ws/wo = 2.1 ± 0.3) their initial weight and the modulus decrease slightly to ~ 5 kPa, similarly insensitive to the domain sizes. The

100 domain sizes or extent of phase separation does not alter the bulk mechanical properties although the microstructures are trapped out of equilibrium. The microstructure hydrogels were further compared with equilibrium RLP- rich, PEG-rich phases and PEG only hydrogels. All hydrogels exhibited frequency- independent deformation response at low frequency suggesting their highly elastic nature (Figure 4.3A). The G’ and G” values at 0.01 Hz are shown in Figure 4.3B, where 0.01 Hz was comparable to average strain rate of microindentation and puncture. Interestingly, the RLP-PEG microstructure hydrogels (5.2 ± 1.4 kPa for RLP-PEG_25um and 4.7 ± 0.7 kPa for RLP-PEG_10um) had the highest G values whereas the RLP-rich (2.4 ± 0.7 kPa) and PEG-rich (2.1 ± 0.7 kPa) ones had lower values. The higher shear modulus of the microstructure RLP-PEG maybe due to the micro-scale RLP-rich domains resistance to the shear force. The 10wt% PEG only hydrogels with storage modulus G’ in 2.5 ± 1.0 kPa, however, had the highest loss modulus G” in 1.1 ± 0.6 kPa. To compare the loss tangent, 퐺″ tan(훿) = 퐺′ was also determined from oscillatory shear rheology data to provide a relative measure of the viscous to elastic properties of materials. The tan (δ) <1 for all the hydrogels indicated an elastic solids behavior (Figure 4.3C). PEG only hydrogels had the highest tan (δ) value (0.47 ± 0.29) and PEG-rich phase hydrogels also had a higher tan (δ) value (0.21 ± 0.07). The RLP-PEG and RLP-rich hydrogels had smaller tan () value which indicated that RLP based materials had low energy dissipation. The results were consistent with the RLP only hydrogels previously reported (tan () ~0.02) and consistent with reported behavior for a range of elastomers which exhibit lower energy storage with increasing loss tangent.365–367

101 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.

4.3.3 Microindentation and Resilience Small-strain microindentation was employed to determine the linear elastic modulus of hydrated RLP-PEG microstructured hydrogels. The stress-strain curves for the RLP-PEG hydrogels are relatively linear at small strains (Figure 4.4A), where the slope of the initial linear regime is a measure of elastic modulus. The elastic moduli

Eindentation, calculated from small-strain microindentation measurements were 17.4 ± 0.4 kPa, 20.0 ± 0.6 kPa and 18.3 ± 0.4 kPa for RLP-PEG_10um, RLP-PEG_25um and RLP-PEG_60um respectively. To compare the elastic moduli from small-strain microindentation and oscillatory shear rheology, elastic storage moduli, Erheology= 3G, were calculated from the shear storage moduli (G) at 0.01 Hz, which is comparable to the average strain rate of microindentation. The elastic moduli calculated from shear rheology were 14.1 ± 2.1 kPa, 15.6 ± 4.2 kPa and 15.4 ± 0.9 kPa for RLP-PEG_10um, RLP-PEG_25um and RLP-PEG_60um respectively (Figure 4.4B). The elastic

102 modulus measurements from shear rheology are comparable with moduli measurements obtained from microindentation.

Microindentation was utilized to measure multiple loading and unloading cycles and determine the resilience of hydrated RLP-PEG microstructured hydrogels.

The stress-strain cycles with different applied force for RLP-PEG hydrogels are shown in Figure 4.4C. Loading and unloading curves for RLP-PEG overlapped and minimal hysteresis was observed even up to ~80% strains. The resilience was determined by comparing the areas under the loading and unloading curves for each cycle. High resilience of >85% for both RLP-PEG hydrogels (domain sizes 25um and 10um) were observed for the range of 12-80% strains (Figure 4.4D). Interestingly, RLP-PEG hydrogels showed lower resilience at low strain (~12% strain) with maximum resilience around 30-40% strain (~94% resilience). Although the resilience was slightly reduced to 91% at high strain, the RLP-PEG hydrogels had resilience comparable with RLP-based hydrogels.368–370

103 Figure 4.4 Microindentation of RLP-PEG microstructured hydrogels A) Stress-strain loading curves during small-strain microindentation for RLP-PEG_25um 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.

4.3.4 Large-strain Indentation and Puncture

Large-strain indentation and puncture experiments were performed on RLP- PEG hydrogels using a spherical indenter to provide insight into inherent material properties associated with fracture initiation, or crack nucleation. The puncture

104 experiments characterized the effective stiffness of hydrogels at large strains to the point of failure, which is not captured by the low-strain microindentation. The critical force for puncture (Pc) is the maximum force associated with the peak in the loading curve, that corresponds to the needle penetrating through the surface of the gels, and the critical depth (dc) corresponds to the critical force. Figure 4.5 shows the force- displacement curves for RLP-PEG hydrogels, where the force increased with compression displacement that corresponding related to the hydrogels is forced out along the side of the indenter at high strain, and then critical puncture occurred at the sharp drop in force. Pc can be related to the critical nucleation stress or critical nucleation energy for fracture formation, depending on the size scale of the indenting probe radius.352

By coupling confocal microscopy and autofluorescence of RLP, the RLP-PEG hydrogel deformation under high strain indentation can be visualized as shown in

Figure 4.6. Three dimension (3D) images were constructed with confocal Z-stack images and performed in between a step indentation of 50 µm each. It is clear that the rhodamine -coated indenter with radius of 15 µm (red) probed both RLP-rich domains (green) and PEG-rich matrix (black) with large deformation. Under the x-y view in two dimensions (2D), the indenter approaching and deforming the RLP-rich domains was clearly visible (Figure 4.6B). RLP-rich domains were spherical in shape and distributed evenly in the RLP-PEG hydrogel before the indenter approached.

While the indenter approached and deformed the RLP-rich domains, it generated a deformation zone around the indenter as indicated by the white dashed line in Figure

4.6B. The deformation zone grew larger as the indenter reached deeper into the hydrogels and deformed RLP-rich domains in a radial pattern.

105 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) 3D view of the indenter forced out along the side of RLP-PEG hydrogel. B) Top x-y view of the indenter approaching, indenting and deforming the RLP-PEG hydrogel in 2D plane for each indentation step.

At small depth, indenters follow the force–displacement behavior predicted by for small-strain contact;31 while at large depth, the force responds to a second order transition dependence on depth for materials compression described by a Neo- Hookean constitutive relationship.29 Force-displacement curves can be fitted as a function of the indentation depth () in the puncture equation.352 The effective elastic modulus (kʹE) is a measure of the large-strain response of materials and has been shown to correlate well with elastic moduli measurements from other mechanical

107 testing techniques, such as shear rheology, for various polymer gels and elastomeric materials.352 Three different indenter probe sizes (R = 5, 13, and 20 µm) were used to compare the puncture mechanics of RLP-PEG, RLP-rich and PEG-rich hydrogels.

The indentation depth was independent of the hydrogels, and showed a linear relationship with the indenter radius (dc ~ 100 R). The Pc and k’E of RLP-PEG hydrogel was in between the RLP-rich and PEG-rich hydrogels (Figure 4.7), which indicated the high strain puncture to be able to probe both the RLP-rich and PEG-rich phase as observed in confocal images (Figure 4.6). The k’E of RLP-PEG with small indenter radius was comparable to the PEG-rich hydrogels, while for the large indenter radius (R = 20 µm), k’E approached that of the RLP-rich hydrogels.

According to the rule of mixing, composites properties depend on the volume fraction and the mechanical properties of each phase,339 with the upper bound modulus (EU) and lower bound modulus (EL) related by

퐼 퐼 퐼퐼 퐼퐼 퐸푈 = 퐸 휙 + 퐸 휙

1 휙퐼 휙퐼퐼 = 퐼 + 퐼퐼 퐸퐿 퐸 퐸 where 퐸 is the modulus and 휙 is the volume fraction in phase I (PEG-rich) and phase II (RLP-rich). However, the k’E did not follow the rule of mixtures in RLP-PEG phase-separated hydrogels, the calculated k’E values for RLP-PEG were higher than the measured values, especially for the 20 µm probe radii. The larger indenter radius was able to indent deeper into the hydrogels and create a larger volume of deformation zone and to sample a greater number of RLP-rich domains compared to the smaller indenter, which may account for this deviation. Further systematic study of confocal imaging during the puncture experiment will be able to clearly visualize the deformation zone with different indenter radius in the future.

108 푃푐 Fracture initiation energies (o) of the materials can be calculated by Γ = , and 표 푅 shows a dependence of indenter radius. The fracture energies were 0.89 ± 0.21 kJ/m2, 0.23 ± 0.05 kJ/m2 and 0.97 ± 0.05 kJ/m2 respectively for RLP-PEG, RLP-rich and

2 PEG-rich hydrogels for a 5 m indenter; and their o increased to 2.00 ± 0.37 kJ/m , 0.75 ± 0.15 kJ/m2 and 2.96 ± 0.58 kJ/m2 respectively. Nevertheless, the calculated fracture energies were within the order of magnitude.

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 The impact of the spherical tip indenter size on puncture was further evaluated with a range of tip radius R from 1 to 35 μm. The critical force for puncture (Pc) was linearly proportional to R for small indenter sizes and exponentially related to R for large indenter sizes (Figure 4.8A). Normalizing the critical load by the cross-sectional

2 area of the indenter defines a critical nominal stress at puncture (σc=Pc/R ); where σc decreases with small indenter R and eventually plateaus for large indenter R. (Figure 4.8B). The puncture behavior was consistent with the transition between energy-limited and stress-limited failure RT for spherical tip indenters reported by

352 Fakhouri et al.. The RT for RLP-PEG hydrogels were around R ~ 20 µm. For the energy-limited regime, the fracture energy (Γo) of the materials can be obtained from the slopes of the Pc vs R plot. The fracture energy of RLP-PEG is 1.31 ± 0.11 kJ/m2 for the energy-limited fracture, and the maximum stress at fracture was 1.62 ± 0.12 kPa for the stress-limited fracture with large indenters.

110 4.4 Conclusions The micromechanical properties of RLP-PEG microstructured hydrogels were evaluated with oscillatory shear rheology, small-strain microindentation, and large- strain indentation and fracture ranging from different length scales. Oscillatory shear rheology and small-strain microindentation measured the small-strain elastic response of 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% high 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 were used for extracting fracture energy and maximum stress of the microstructured

RLP-PEG hydrogels.

111 Chapter 5

CONCLUSIONS AND FUTURE WORK

5.1 Conclusions and Perspectives Multicomponent hybrid hydrogels have been widely used for a variety of applications in tissue regeneration and drug delivery. By incorporating and modulating the mechanically functional and bioactive components in the network, the mechanical and biological properties of the hydrogel can be tuned independently without sacrificing one or the other. Hybrid hydrogels are expected to further mimic the microenvironment for cells and tissue reorganization. The mechanically active components should be aimed not only at affecting the bulk mechanical properties, but also should capture the micro-mechanical properties in native tissue. Multicomponent hydrogels with well-organized domains provided significant opportunities for hydrogel materials for tissue engineering applications.

Emerging resilin-like polypeptides (RLPs), based on the insect protein resilin, offer intriguing opportunities in the generation of protein-containing microstructured hydrogels, as they exhibit exceptional rubber-like elasticity characterized by low stiffness, high extensibility, and efficient energy storage. Mechanical properties of

RLP hydrogels are compatible for mechanically active tissues, such as vocal fold and cardiovascular tissues. Liquid-liquid phase separation (LLPS) has been a well-known phenomenon for more than a century, with the unfavorable interactions of two

112 dissimilar polymer blends, block copolymers, proteins, and polysaccharides causing phase separation of ternary solutions. LLPS provides a method for generating microstructured materials, as phase behavior is readily tuned by factors such as temperature, polymer molecular weight, polymer concentration. Aqueous LLPS of protein-polymer solutions has been ideal for the development a simple one-step process for controlled structured protein-based biomaterials.

Heterogeneous hybrid hydrogels comprising RLPs can be engineered with controlled microstructure and distinct micromechanical properties via the LLPS of aqueous solutions of the RLPs and PEG, as shown in Chapter 2. The aqueous-based LLPS of RLP/PEG can be tuned by temperature, concentration, and the ratio of RLP to PEG in initial solutions. Phase diagrams of the RLP/PEG system were generated to define solution parameters that would yield micron-scale domains in the hydrogels.

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), which provided stable crosslinking in both RLP-rich and PEG-rich phases. The fast crosslinking kinetics and relatively slow phase separation kinetics was important for capturing microstructure of the hydrogel during LLPS. Hydrogels with a microstructure comprising 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 and micromechanical properties of the domains that were distinct from those of the matrix.

113 Despite the microstructured materials described, few methods permit independent control over microstructure and microscale mechanics by single-step processing methods. In Chapter 3, photo-triggered crosslinking methodology was shown to trap microstructures in LLPS solutions of a highly elastomeric RLP and PEG. Evaluation of the LLPS for RLPs with various degrees of acrylamide functionalization established that equilibrium phase diagrams were not significantly altered with acrylation. RLP-rich domains of various diameters can be trapped in a PEG continuous phase via photo-triggered crosslinking, with the kinetics of domain maturation dependent on the degree of acrylation, as a result of miscibility mediated by hydrophobic interactions between the RLP-Ac and PEG-Ac. 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.

Successful encapsulation, localization, and survival of hMSCs cells was demonstrated and suggests the potential application of phase-separated RLP/PEG hydrogels in regenerative medicine applications.

Microscale mechanical responses of the RLP-PEG microstructured hydrogels were characterized in Chapter 4. Oscillatory shear rheology, small-strain microindentation, and large-strain indentation and fracture assessed the bulk and micromechanical properties of RLP-PEG hydrogels. Oscillatory shear rheology and

114 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 reveal a high resilience value (>85%) for RLP-PEG hydrogels even up to 80% high strain. Large-strain puncture monitored with confocal microscopy enabled visualization of the microstructured hydrogel under indentation and deformation of RLP-rich domains. The puncture experiments were also employed to capture 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 evaluated and a fracture energy and maximum stress of the microstructured RLP-PEG hydrogels was extracted. The microstructured RLP-PEG hydrogels maintained excellent resilience and toughness, which further supports their potential use for biomaterials-based tissue engineering.

5.2 Future work

5.2.1 Stabilization of Phase-Separated RLP/PEG Solutions and Hydrogels Aqueous based liquid-liquid phase separation of RLP/PEG will ultimately yield two immiscible layers. The process of macro phase separation is dependent on the evolution of the coalescence of the domains, which is balanced by gravitational, flotational, and frictional forces.336 The rate of the phase separation can be correlated with physical properties of the RLP/PEG system as function of the relative density, interfacial tension, and viscosity of both phases,371–373 and the time scales of LLPS can range from seconds to minutes near the critical point.313 In order to capture microstructure during phase separation, the relative kinetics of phase

115 separation has to be slower than chemical crosslinking during hydrogel formation.

Selecting the initial composition from the phase diagram in Figure 5.1A permits tuning of the phase separated RLP-rich and PEG-rich concentrations, as well as the microstructure.272,277 However, the coalescence rate increases with increasing volume fractions. 374 As a result, higher RLP compositions in RLP/PEG (such as 15wt%

80/20 RLP/PEG) shows rapid macro phase separation in less than minutes (Figure

5.1B), which thus do not yield evenly dispersed PEG-rich domains in an RLP-rich matrix, nor bicontinuous structured hydrogels.

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.

As mentioned above, methods to reduce the interfacial energy of the phases could be employed to try to increase the homogeneity of the domain sizes.375–379

Incorporating surfactant or nanoparticles has been shown to reduce interfacial tension at the water-water interface.380–383 We thus tried to design a RLP-PEG peptide conjugate to stabilize RLP/PEG phase separation system. Three peptide

116 conjugates PEG-RLP, PEG-4FRLP and PEG-4WRLP were designed and synthesized.

The purified peptides with 10 mg/ml were incorporated in 15wt% 80/20 RLP/PEG and crosslinked immediately (Figure 5.2). Without peptides, the RLP/PEG completely phase separated into two layers. Incorporating with 10 mg/ml peptides showed a top layer with RLP-rich domains and PEG-rich matrix, while bottom layers shown an inverse RLP-matrix and PEG-rich domains. However, the hydrogels were not evenly dispersed, indicating the need to optimize the concentration of the peptides and RLP or PEG diblock length in the conjugates.

PEG-RLP = PEG1000-GGRPSDSFGAPGGGN

PEG-4FRL = PPEG1000-FFFFGGRPSDSFGAPGGGN

PEG-4WRLP = PEG1000-WWWWGGRPSDSFGAPGGGN

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 5.2.2 Injectable Resilin-like Polypeptide Hybrid Hydrogels for Vocal Folds Vocal fold tissue has unique mechanical properties that serve important functions in voice production, such as vibration at frequencies of 100-1000 Hz and reversible recovery after transient stretch to high strain.384–387 Materials for vocal fold treatment therefore not only should have minimum foreign body and inflammatory reactions, but also should match the biomechanical and viscoelastic properties of the vocal fold tissue. Resilin-like polypeptides (RLPs) offer advantages in engineering new elastomeric and highly mechanically active biomaterials that can be tailored for vocal fold tissue repair.128,252,253,388,389 Native ECM-based biomaterials including HA and collagen have also been employed in vocal fold injection; the HA-based materials are likely the most promising materials in the field.390 The possibility of introducing elastomeric RLP to HA-based biomaterials would thus offer exciting opportunities to further improve the mechanical match of the biomaterial, maintain mechanical integrity during vibration under high frequencies, and exhibit dynamic mechanical response for vocal fold tissue engineering.

RLP/HA-based materials are employed in these studies to capture both the high-frequency robustness of the RLP and anti-inflammatory properties of HA. Three compositions of materials for in vivo injection into rabbit vocal fold included a physical hydrogel (RLP/HA gel), a chemically crosslinked hydrogel (RLP-Ac/HA-SH gel) and an RLP solution to evaluate the biocompatibility of the RLP-based materials in vivo. Schematics of the physically crosslinked RLP/HA gel and the chemically crosslinked RLP-Ac/HA-SH gel are shown in Figure 5.3. The physical RLP/HA gel was formed via electrostatic interactions between the positively charged lysine residues of the RLP and the negatively charged HA backbone. The chemically crosslinked RLP-Ac/HA-SH gel was crosslinked via Michael-type addition chemistry,

118 which is a highly efficient reaction with no by-products, which has been employed widely for in situ reaction in biological systems.391 Mechanical properties of these hydrogels in vitro, prior to their use in vivo, were evaluated via oscillatory rheology to confirm that the solutions would form hydrogels as expected. RLP/HA gel solutions were expected to form viscoelastic networks as a result of electrostatic interactions between the lysine-rich RLP and the negatively charged HA. RLP/HA immediately formed a hydrogel at 37oC with a storage modulus of 632 ± 35 Pa (Figure 5.4A-B). The chemically crosslinked RLP-

Ac/HA-SH gel formed a stiffer hydrogel via Michael-type addition reaction between the RLP and the HA; given the functionalization of the biopolymers, no additional crosslinker was added and no side products were produced. The functionalization of the biopolymers reduces their net charge, and thus the immediate formation of a physical network was not observed; the gel point was observed 40 min after mixing at

37oC and the hydrogel reached a stable modulus of 1570 ± 500 Pa (Figure 5.4C-D), similar to those reported for human vocal fold tissue (shear modulus G’ = 400-2000

Pa).392,393 In vivo, rabbit vocal folds injected with RLP-based biomaterials retained normal viscoelastic properties days 5 and 21 post injection. These findings demonstrate that, RLP/HA and RLP-Ac/HA-SH gels replicate mechanical properties of vocal fold mucosa in vitro and in vivo. Rabbit vocal fold tissues were histologically normal after injection of RLP solution, demonstrating biocompatibility of RLP in vocal fold lamina propria (Figure 5.6). RLP/HA and RLP-Ac/HA-SH hydrogel injections into lamina propria resulted in no inflammation at day 5 but mild inflammatory changes at day 21. Significant inflammatory reaction to intramuscular

119 RLP-Ac/HA-SH gel at day 21 in one rabbit in our study emphasizes the need to control depth of injection in vocal fold procedures. It is noteworthy that epithelium was normal and lamina propria was only mildly inflamed, and that this rabbit, like the others in this study, did not suffer adverse health effects after injection. Given only mild, non-fibrotic histopathological changes induced by lamina propria injections of RLP/HA and RLP-Ac/HA-SH gels by day 21 and intact vocal fold viscoelasticity, both hydrogels appear suitable as candidates for vocal fold tissue engineering.

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.

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158 Appendix A

Chapter 1 Reproduced in its partially from:

Title: Opportunities for Multicomponent Hybrid Hydrogels in Biomedical Applications

Author: H. K. Lau, K. L. Kiick

159 Figure 1.1

Figure 1.2

160 Figure 1.3

Figure 1.4

Chapter 2 Reproduced in its entirety from:

Title: Aqueous Liquid–Liquid Phase Separation of Resilin-Like Polypeptide/Polyethylene Glycol Solutions for the Formation of Microstructured Hydrogels

Author: H. K. Lau, L. Li, A. K. Jurusik, C. R. Sabanayagam, K. L. Kiick

Publication: ACS Biomaterials Science & Engineering Publisher: American Chemical Society Date: May 1, 2017

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Chapter 3 Reproduced in its entirety from: Title: Microstructured Elastomer‐PEG Hydrogels via Kinetic Capture of Aqueous

Liquid–Liquid Phase Separation

Author: H. K. Lau, A. Paul, I. Sidhu, L. Li, C. R. Sabanayagam, S. H. Parekh, K. L. Kiick

Publication: Advanced Science Publisher: Wiley-VCH Verlag GmbH & Co. KGaA Date: March 12, 2018 https://onlinelibrary.wiley.com/doi/abs/10.1002/advs.201701010

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