Science, Faculty of
Chemistry, Department of
DSpace
UBCV
Fu, Linglan
Doctor of Philosophy - PhD
University of British Columbia
Most recently, hydrogels composed of recombinant proteins/peptides have attracted great interest due to their diverse biofunctions, designable amino acids sequence, reactive side chains, excellent biocompatibility and biodegradability. Moreover, conformational switching of folded proteins induces structural and thermodynamic changes in the network, leading to dynamic properties and energy dissipative ability of hydrogels. Despite these advantages, applications of recombinant protein-based hydrogels are often limited by their poor mechanical performance and inefficient network crosslinking, due to steric hindrance of bulky protein domains in the networks. To optimize mechanical behaviors of structural protein-based hydrogels, it is important to understand the design principles of globular protein-based elastomeric networks. This thesis discusses the latest research on protein-based hydrogels, including network designs, crosslinking strategies, and resulting physical and mechanical features of the materials. In addition, classic rubber elasticity theory is introduced to reveal the effects of elastomeric network on physical and mechanical properties of synthetic polymer hydrogels. Based on the theory, a semi-quantitative tool was developed to predict/explain mechanical properties of globular folded polyprotein-based hydrogels, providing guidance for the rational design of globular polyprotein-based hydrogels with desired mechanical properties. Furthermore, we demonstrate for the first time that the macroscopic mechanical energy generated from conformational change of globular proteins closely corresponds to single molecule mechanochemical events by investigating mechanochemical coupling cycles of globular folded proteins at both single-molecule and macroscopic levels. With the guidance of the design principle, we describe synthesis of super tough protein-based hydrogels based on a novel denatured-crosslinking hydrogelation method, which provides potential cartilage-like biomaterials and a general strategy for mechanical enhancement for protein-based hydrogels. We also explored the synthesis of super tough protein-alginate hybrid hydrogels using a double-network approach, further broadening the range of mechanical properties that protein-based biomaterials can reach. Taking advantage of precise mechanical control and dynamic behaviors of mutually exclusive protein-based hydrogels, an application of protein-based hydrogels is reported as extracellular matrices for cell spreading studies, demonstrating a great potential utility of dynamic protein hydrogels in cellular mechano-biology studies.
ENGINEERING GLOBULAR PROTEIN-BASED HYDROGELS WITH TUNABLE MECHANICAL PROPERTIES by Linglan Fu B.Eng., Tianjin University, 2015 A THESIS SUBMITTED IN PARTIAL FULFILLMENT OF THE REQUIREMENTS FOR THE DEGREE OF DOCTOR OF PHILOSOPHY in THE FACULTY OF GRADUATE AND POSTDOCTORAL STUDIES (Chemistry) THE UNIVERSITY OF BRITISH COLUMBIA (Vancouver) March 2021 © Linglan Fu, 2021 ii The following individuals certify that they have read, and recommend to the Faculty of Graduate and Postdoctoral Studies for acceptance, the dissertation entitled: Engineering globular protein-based hydrogels with tunable mechanical properties submitted by Linglan Fu in partial fulfillment of the requirements for the degree of Doctor of Philosophy in Chemistry Examining Committee: Hongbin Li, Chemistry, UBC Supervisor Jayachandran Kizhakkedathu, Pathology, Chemistry, UBC Supervisory Committee Member David Chen, Chemistry, UBC University Examiner Mark Thachuk, Chemistry, UBC University Examiner Additional Supervisory Committee Members: Suzana Straus, Chemistry, UBC Supervisory Committee Member Edward Grant, Chemistry, UBC Supervisory Committee Member iii Abstract Most recently, hydrogels composed of recombinant proteins/peptides have attracted great interest due to their diverse biofunctions, designable amino acids sequence, reactive side chains, excellent biocompatibility and biodegradability. Moreover, conformational switching of folded proteins induces structural and thermodynamic changes in the network, leading to dynamic properties and energy dissipative ability of hydrogels. Despite these advantages, applications of recombinant protein-based hydrogels are often limited by their poor mechanical performance and inefficient network crosslinking, due to steric hindrance of bulky protein domains in the networks. To optimize mechanical behaviors of structural protein-based hydrogels, it is important to understand the design principles of globular protein-based elastomeric networks. This thesis discusses the latest research on protein-based hydrogels, including network designs, crosslinking strategies, and resulting physical and mechanical features of the materials. In addition, classic rubber elasticity theory is introduced to reveal the effects of elastomeric network on physical and mechanical properties of synthetic polymer hydrogels. Based on the theory, a semi-quantitative tool was developed to predict/explain mechanical properties of globular folded polyprotein-based hydrogels, providing guidance for the rational design of globular polyprotein-based hydrogels with desired mechanical properties. Furthermore, we demonstrate for the first time that the macroscopic mechanical energy generated from conformational change of globular proteins closely corresponds to single molecule mechanochemical events by investigating mechanochemical coupling cycles of globular folded proteins at both single-molecule and macroscopic levels. iv With the guidance of the design principle, we describe synthesis of super tough protein-based hydrogels based on a novel denatured-crosslinking hydrogelation method, which provides potential cartilage-like biomaterials and a general strategy for mechanical enhancement for protein-based hydrogels. We also explored the synthesis of super tough protein-alginate hybrid hydrogels using a double-network approach, further broadening the range of mechanical properties that protein-based biomaterials can reach. Taking advantage of precise mechanical control and dynamic behaviors of mutually exclusive protein-based hydrogels, an application of protein-based hydrogels is reported as extracellular matrices for cell spreading studies, demonstrating a great potential utility of dynamic protein hydrogels in cellular mechano-biology studies. v Lay Summary Hydrogels are water-containing soft materials, made of crosslinked polymer networks. This thesis provides a comprehensive understanding of recombinant protein-based hydrogel networks at both molecular and macroscopic levels. A semiquantitative tool is developed to guide the design of globular folded polyprotein-based hydrogels. This allows us to further improve and to precisely control the mechanical performance of protein-based hydrogels for desired applications, such as soft actuators, cartilage-like biomaterials, and artificial substrates for cell biology studies in vitro. vi Preface Portions of this thesis benefited from the intellectual and experimental contributions of others. Chapter 2 has been published. [Linglan Fu] and Hongbin Li. (2020) Toward Quantitative Prediction of the Mechanical Properties of Tandem Modular Elastomeric Protein-Based Hydrogels. Macromolecules. 53 (12): 4704-4710. This chapter is incorporated in this thesis with permission of American Chemical Society, Copyright 2020. Dr. Li and I conceived the project. Dr. Li and I designed the experiments together. I am responsible for all the experiments and the data analysis. I wrote the manuscript together with Dr. Li. Chapter 3 has been published as [Linglan Fu], Han Wang and Hongbin Li. (2019) Harvesting Mechanical Work From Folding-Based Protein Engines: From Single-Molecule Mechanochemical Cycles to Macroscopic Devices. CCS Chemistry 1(1) 138-147. This chapter is incorporated in this thesis with permission of Chinese Chemical Society, Copyright 2020. Dr. Li conceived the project. Dr. Li and I designed the experiments together. Han Wang performed the optical trapping experiment and related data analysis. I am responsible for the remaining experiments and the data analysis. I wrote the manuscript together with Dr. Li, and Han Wang helped to revise the manuscript. A version of Chapter 4 is in preparation for publication. [Linglan Fu] and Hongbin Li. Converting Muscle-mimetic Biomaterials to Cartilage-like Materials. Dr. Li and I conceived the project. Dr. Li and I designed the experiments together. I am responsible for all the experiments and all the data analysis. I wrote the manuscript together with Dr. Li. vii Chapter 5 is in preparation for publication. [Linglan Fu] and Hongbin Li. Super Tough and Strong Protein-Alginate Double-network Hydrogel. Dr. Li and I conceived the project. I designed the experiments. I am responsible for all the experiments and all the data analysis. I wrote this chapter. Chapter 6 has been published as [Linglan Fu], Amanda Haage, Na Kong, Guy Tanentzapf, and Hongbin Li. (2019) Dynamic protein hydrogels with reversibly tunable stiffness regulate human lung fibroblast spreading reversibly. Chemical Communications. 55 (36): 5235-5238. This chapter is incorporated in this thesis with permission of The Royal Society of Chemistry, Copyright 2019. Dr. Li conceived the project. Dr. Li, Dr. Kong and I designed the experiments together. Dr. Kong constructed the gene of FGR(G-MEP-R)2. Dr. Tanentzapf and Dr. Haage helped with the cell proliferation assay and confocal imaging. I am responsible for the remaining experiments and all the data analysis. I wrote the manuscript together with Dr. Li. viii Table of Contents Abstract ..................................................................................................................................... iii Lay Summary .............................................................................................................................v Preface ....................................................................................................................................... vi Table of Contents ................................................................................................................... viii List of Tables ............................................................................................................................xv List of Figures ......................................................................................................................... xvi List of Symbols .........................................................................................................................xx List of Abbreviations ........................................................................................................... xxiii Acknowledgements ............................................................................................................ xxviii Dedication ............................................................................................................................. xxix Chapter 1: Introduction ........................................................................................................... 1 1.1 Protein- and peptide-based hydrogels ......................................................................... 3 1.1.1 Synthetic protein-based hydrogels .......................................................................... 3 1.1.1.1 Synthetic elastomeric globular protein-based hydrogels ................................ 3 1.1.1.2 Synthetic dynamic protein-based hydrogels ................................................... 6 1.1.1.3 Synthetic protein hydrogels based on protein assemblies ............................... 9 1.1.2 Synthetic polypeptide-based hydrogels ................................................................ 13 1.1.2.1 Elastin-like polypeptide-based (ELP) hydrogels .......................................... 14 1.1.2.2 Hydrogels based on coiled-coil domains. ..................................................... 16 1.1.2.3 Hydrogels based on β-sheet motifs ............................................................... 19 1.1.3 Natural protein-based hydrogels ........................................................................... 22 ix 1.1.3.1 Silk-based hydrogels ..................................................................................... 22 1.1.3.2 Type I Collagen-based hydrogels ................................................................. 24 1.1.3.3 Gelatin-based hydrogels ................................................................................ 25 1.1.3.4 Hydrogels made of other natural proteins ..................................................... 26 1.2 Rubber elasticity theory: design principle of protein-based elastomer mechanics ... 31 1.2.1 Classic rubber elasticity theory ............................................................................. 31 1.2.2 Methods to improve mechanical properties of protein-based hydrogels .............. 36 1.2.2.1 Increase of polymer concentration ................................................................ 37 1.2.2.2 Network Functionalization with cross-linkable groups ................................ 38 1.2.2.3 Increase of the degree of polymerization ...................................................... 41 1.2.2.4 Double networks ........................................................................................... 43 1.2.2.5 Mechanical reinforcement by forming nanostructures ................................. 45 1.2.2.6 Incorporation of nanocomposites .................................................................. 45 1.3 Goals of the thesis ..................................................................................................... 46 Chapter 2: Towards Quantitative Prediction of Mechanical Properties of Tandem Modular Elastomeric Protein-based Hydrogels ................................................................... 50 2.1 Synopsis .................................................................................................................... 50 2.2 Introduction ............................................................................................................... 51 2.3 Materials and methods .............................................................................................. 54 2.3.1 Protein engineering ............................................................................................... 54 2.3.2 Hydrogel preparation and dimension measurements ............................................ 58 2.3.3 Tensile tests ........................................................................................................... 59 2.4 Results and discussion .............................................................................................. 60 x 2.4.1 Length-equivalent molecular weight for folded globular domains ....................... 60 2.4.2 Length equivalent molecular weight approach adequately describes the Young’s modulus of (GR)12 hydrogels ............................................................................................ 62 2.4.3 Length equivalent molecular weight approach can be used to predict the change of the mechanical properties of dynamic protein hydrogels ................................................. 65 2.4.4 Mechanical properties of (GR)-based hydrogels .................................................. 67 2.5 Conclusions ............................................................................................................... 69 Chapter 3: Harvesting Mechanical Work from Folding-based Protein Engines: From Single-Molecule Mechanochemical Cycles to Macroscopic Devices .................................. 71 3.1 Synopsis .................................................................................................................... 71 3.2 Introduction ............................................................................................................... 72 3.3 Materials and methods .............................................................................................. 74 3.3.1 Protein engineering ............................................................................................... 74 3.3.2 Optical trapping experiments ................................................................................ 74 3.3.3 Hydrogel preparation and dimension measurements ............................................ 74 3.3.4 Tensile Tests ......................................................................................................... 75 3.3.5 Calculation of the work generated by the folding of proteins at the single molecule level ................................................................................................................... 75 3.4 Results and discussion .............................................................................................. 76 3.4.1 Mechanochemical coupling of GB1 at the single molecule level ......................... 76 3.4.2 Mechanochemical coupling at the macroscopic level .......................................... 80 3.4.3 Bridging the gap between the mechanochemical coupling at single molecule level and macroscopic level ....................................................................................................... 84 xi 3.5 Conclusions ............................................................................................................... 88 Chapter 4: Converting Muscle-mimetic Biomaterials to Cartilage-like Materials .......... 91 4.1 Synopsis .................................................................................................................... 91 4.2 Introduction ............................................................................................................... 92 4.3 Materials and methods .............................................................................................. 95 4.3.1 Protein engineering ............................................................................................... 95 4.3.2 Hydrogel preparation and dimension measurements ............................................ 95 4.3.3 Swelling ratio and water content measurements ................................................... 96 4.3.4 Tensile tests ........................................................................................................... 97 4.3.5 Compressive tests .................................................................................................. 97 4.3.6 Scanning electron microscopy (SEM) imaging .................................................... 98 4.3.7 Characterization of dityrosine cross-links in hydrogel by acid hydrolysis-fluorescence method ......................................................................................................... 98 4.3.8 Cysteine shotgun fluorescence labeling by IAEDANS and fluorescence measurements .................................................................................................................... 99 4.3.9 Single-molecule optical tweezers measurements. ............................................... 100 4.3.10 Hydrogel biocompatibility measurement, LIVE/DEAD viability assay. ....... 100 4.4 Results and discussion ............................................................................................ 101 4.4.1 A denaturing-crosslinking strategy reinforces globular protein-based hydrogel.101 4.4.2 Tough N-DC (FL)8 hydrogels display excellent tensile properties .................... 109 4.4.3 Remarkable compressive performance of N-DC hydrogels ............................... 114 4.5 Conclusions ............................................................................................................. 120 Chapter 5: Super Tough and Strong Protein-Alginate Double-network Hydrogel ....... 123 xii 5.1 Synopsis .................................................................................................................. 123 5.2 Introduction ............................................................................................................. 123 5.3 Materials and methods ............................................................................................ 127 5.3.1 Protein engineering ............................................................................................. 127 5.3.2 Alginate derivatization with phenol moieties ..................................................... 127 5.3.3 Characterization of phenolic density of the modified alginate. .......................... 128 5.3.4 Preparation of alginate hydrogels and double-network hydrogels. .................... 128 5.3.5 Tensile tests ......................................................................................................... 129 5.3.6 Rheological tests ................................................................................................. 130 5.4 Results and discussion ............................................................................................ 130 5.4.1 Physical and chemical gelation of derivatized alginate ...................................... 130 5.4.2 Cooperation of alginate and polyproteins: mechanical improvement of a double-network hydrogel ............................................................................................................ 135 5.4.2.1 Denatured-crosslinking (FL)8 covalently crosslinks with derivatized alginate (D-DC (FL)8/Alg hydrogel). ....................................................................................... 135 5.4.2.2 Renaturation of the denatured-crosslinking (FL)8 enhances double-network density (N-DC (FL)8/Alg). .......................................................................................... 136 5.4.2.3 N-DC (FL)8/Alg in the presence of calcium ions further improved toughness of the double-network hydrogels (I-N-DC (FL)8/Alg). .............................................. 139 5.5 Conclusions ............................................................................................................. 142 Chapter 6: Dynamic Protein Hydrogels with Reversibly Tunable Stiffness Regulate Human Lung Fibroblasts Spreading Reversibly ............................................................... 144 6.1 Synopsis .................................................................................................................. 144 xiii 6.2 Introduction ............................................................................................................. 145 6.3 Materials and methods ............................................................................................ 149 6.3.1 Protein engineering ............................................................................................. 149 6.3.2 Hydrogel preparation for cell spreading and tensile tests ................................... 149 6.3.3 Characterization of dityrosine cross-links in hydrogel by acid hydrolysis-fluorescence method ....................................................................................................... 150 6.3.4 Cell spreading and redox switching .................................................................... 151 6.3.5 Tensile tests ......................................................................................................... 153 6.3.6 Scanning electron microscopy (SEM) imaging .................................................. 153 6.4 Results and discussion ............................................................................................ 153 6.4.1 Dynamic protein hydrogel based on FGR(G-MEP-R)2 shows reversible redox-responsive microstructural, physical and mechanical properties .................................... 153 6.4.2 The MEP hydrogels support cell proliferation .................................................... 157 6.4.3 HLF cells dynamically change morphology in response to a cyclic mechanical change of the MEP hydrogel ........................................................................................... 158 6.5 Conclusions ............................................................................................................. 164 Chapter 7: Conclusions and Outlook .................................................................................. 166 7.1 Conclusions ............................................................................................................. 166 7.2 Outlook ................................................................................................................... 169 7.2.1 Exploring smart protein-based hydrogels with actuation functions ................... 169 7.2.2 Designing dynamic protein-based hydrogels supporting cell mechano-biology studies. ............................................................................................................................ 170 7.2.3 Designing recombinant protein-based hydrogels as 3D ECMs .......................... 171 xiv 7.2.4 Engineering protein-based biomaterials for bioprinting ..................................... 173 References ...............................................................................................................................175 Appendix: Amino acid sequences of the proteins studied ..................................................199 xv List of Tables Table 1-1 Summary of the advantages and disadvantages of natural protein-based hydrogels in Section1.1.3. .................................................................................................................................. 30 Table 2-1. Average results applying length equivalent molecular weight on (GR)12, (GB1)8, (FL)8 hydrogels. ............................................................................................................................ 64 Table 2-2 Prediction of MEP-based dynamic hydrogel (Eoxidized/Ereduced) based on Me. ............... 67 Table 4-1 Mechanical properties of the state-of-the-art protein hydrogels and polymer hydrogels...................................................................................................................................................... 118 Table 4-2 Enhancement of mechanical properties via the DC hydrogelation method. .............. 120 Table 5-1 Concentration of alginate, EDC, NHS and tyramine for the synthesis of alginate-tyramine conjugates .................................................................................................................... 128 Table 5-2 Young’s modulus (E) and breaking strain of (FL)8/Alg double network hydrogels. . 142 xvi List of Figures Figure 1-1 Tandem modular elastomeric polyproteins for hydrogel engineering. ......................... 4 Figure 1-2 Tensile properties of (FL)8 hydrogels. .......................................................................... 6 Figure 1-3 A dynamic hydrogel network based on the conformational change of calmodulin (CaM). ............................................................................................................................................. 8 Figure 1-4 Representative protein-protein interaction strategies that have been utilized for hydrogel engineering. ..................................................................................................................... 9 Figure 1-5 4-arm star-like tetrameric recombinant protein-based hydrogel based on specific protein-protein interaction. ........................................................................................................... 11 Figure 1-6 ELP-PEG-based rubber like hydrogels. ...................................................................... 15 Figure 1-7 Hydrogels based on coil-coiled domains. ................................................................... 19 Figure 1-8 Packed fibrillar network structures made of β-sheet motifs. ...................................... 21 Figure 1-9 Adhesion and cohesion of mussel byssal plaque. ....................................................... 28 Figure 1-10 The rubber elasticity theory based on affine deformation and statistical thermodynamics. ........................................................................................................................... 33 Figure 1-11 Schematics of hydrogel mechanical enhancement approaches. ................................ 38 Figure 1-12 Schematics of ELP-CCP-MCP tough hydrogels. ...................................................... 41 Figure 2-1 Schematics of the length-equivalent molecular weight approach. .............................. 54 Figure 2-2 Engineering the gene of desired polyprotein through multiple steps. ......................... 57 Figure 2-3 Monodispersed molecular weight of GB1-Resilin (GR)n tandem modular polyproteins. .................................................................................................................................. 59 Figure 2-4 Schematics of an Instron-5500R tensometer. .............................................................. 60 xvii Figure 2-5 Physical dimensions and mechanical properties native and denatured (GR)12 hydrogel rings............................................................................................................................................... 63 Figure 2-6 Using length equivalent molecular weight approach to predict the change of the mechanical properties of MEP-based dynamic hydrogels. ........................................................... 66 Figure 2-7 Molecular weight of globular polyproteins in crosslinked network contributes to mechanical properties. .................................................................................................................. 68 Figure 2-8 Native and denatured (GR)12 hydrogels at protein concentration ranging from 10% to 20%. .............................................................................................................................................. 70 Figure 3-1 Force-distance curves of (GB1)4 show that GB1 can refold against a stretching force........................................................................................................................................................ 77 Figure 3-2 Mechanochemical cycles lead to the harvesting of mechanical work generated by the folding of proteins. ........................................................................................................................ 78 Figure 3-3 Mechanochemical cycles at the macroscopic level lead to the harnessing of mechanical work generated by protein folding in protein hydrogels. .......................................... 83 Figure 3-4 Dynamic response of GRG5RG4R-based hydrogels during mechano-chemical coupling. ........................................................................................................................................ 84 Figure 3-5 Mechanical work generated by the folding of GB1 domains at the single molecule level. .............................................................................................................................................. 85 Figure 3-6 The output of the mechanical work can be tuned by the bias load in mechanochemical cycles. ............................................................................................................................................ 86 Figure 3-7 Dynamic response of the 20% GRG5RG4R-based hydrogel at zero load when the buffer is switched from PBS to 7 M GdHCl. ................................................................................ 87 Figure 4-1 Representative force-distance curves of FL domain at a pulling speed of 50 nm/s. ... 94 xviii Figure 4-2 Schematics of the NC-(FL)8 hydrogels. .................................................................... 101 Figure 4-3 Physical entanglements enhanced the stiffness of the (FL)8 hydrogels. ................... 103 Figure 4-4 Fluorescence spectra of (FL)8 hydrogels. .................................................................. 104 Figure 4-5 DC (FL)8 hydrogels can be cycled between its N-DC and D-DC states reversibly. . 105 Figure 4-6 Scanning electron microscopy images of the N-DC (A) and N-NC (B) (FL)8 hydrogels. .................................................................................................................................... 106 Figure 4-7 Schematics of the chain entangled network structure of the N-DC (FL)8 hydrogel. 107 Figure 4-8 Typical tensile stress-strain curves of 20% N-DC (FM23C)8 hydrogels. ................. 108 Figure 4-9 Mechanical properties of N-DC and D-DC (FL)8 hydrogels in tensile testing. ........ 109 Figure 4-10 Mechanical properties of N-DC (FL)8 and (FL)16 hydrogels at different protein concentrations. ............................................................................................................................ 110 Figure 4-11 The N-DC (FL)8 hydrogels can dissipate a large amount of energy. ...................... 112 Figure 4-12 N-DC (FL)8 hydrogels displayed superb compressive mechanical properties. ...... 113 Figure 4-13 Compression tests and stress-strain curves of a N-DC (FL)8 hydrogel compressed to failure. ......................................................................................................................................... 115 Figure 4-14 Consecutive compression-unloading curves of a N-DC (FL)8 hydrogel at a frequency of 0.08 Hz (A) and 0.67 Hz (B). ................................................................................................. 116 Figure 4-15 Cell viability tests on N-DC hydrogels. .................................................................. 119 Figure 5-1 Schematic design of the Alginate-(FL)8 double-network hydrogel. ......................... 125 Figure 5-2 UV-Vis spectra are used to determine phenolic density of the modified alginate. ... 131 Figure 5-3 Physical and chemical gelation of alginate with phenol moieties. ............................ 132 Figure 5-4 Mechanical properties of D-DC (FL)8/Alg double-network hydrogels. ................... 135 xix Figure 5-5 Mechanical performance and swelling behavior of (FL)8/Alg double network hydrogels. .................................................................................................................................... 138 Figure 5-6 Representative stress-strain curves of (FL)8/Alg gels. .............................................. 140 Figure 6-1 Dynamic protein hydrogels constructed from MEP. ................................................. 148 Figure 6-2 Swelling behavior and microstructure of oxidized and reduced FGR(G-MEP-R)2 hydrogels. .................................................................................................................................... 154 Figure 6-3 Reversible mechanical properties of FGR(G-MEP-R)2 hydrogels governed by redox stimuli. ........................................................................................................................................ 156 Figure 6-4 HLF cells can proliferate on MEP hydrogels. ........................................................... 158 Figure 6-5 Control experiments were carried out with HLF cells spreading on 7% FRF4RF4R hydrogels. .................................................................................................................................... 160 Figure 6-6 Control HLF culturing experiments on FRF4RF4R hydrogels with different Young’s modulus. ...................................................................................................................................... 161 Figure 6-7 Morphology of human lung fibroblasts can dynamically respond to the change of the hydrogel stiffness. ....................................................................................................................... 162 xx List of Symbols b bond length of a freely jointed chain d0 diameter DL the contracted length of the protein polymer chain upon protein folding E Young’s modulus f retraction force F0 the constant force G shear modulus G' storage modulus G'' loss modulus h0 height kB Boltzmann constant l change in length ℒ the Langevin function Laa the average length of an amino acid residue Ld the length of the folded protein from its N- and C-termini Lec effective length Maa the average Mw of an amino acid residue Mc molecular weight between cross-links mdry mass of dried gel xxi Me length-equivalent molecular weight for folded globular domains mi mass of freshly prepared gel ms mass of equilibrium gel Mw molecular weight n the number of active network strands per unit volume of the network N the number of bonds of a freely jointed chain Naa the number of amino acid residue in the disordered polypeptide chain Nc the number of crosslinking points per chain p persistence length Pf the folding probability of the protein R gas constant r swelling ratio "#$% the end-to-end distance of a freely jointed chain S entropy T absolute temperature U internal energy V volume w water content x the distance between the native and unfolded states of protein xxii ΔG0 the free energy difference between the unfolded and native states of the protein at zero force ΔLc the contour length increment upon protein unfolding/folding Δxf the distance between the unfolded and the transition state Δxu the distance between the native and transition state  the end-to-end distance of a worm-like chain λ extension ratio ρ density of polymer σ retractive stress xxiii List of Abbreviations 2D two-dimensional 3D three-dimensional aa amino acid AC10A the triblock protein with a middle random coil domain (C10) flanked by two associative leucine zipper domains (A) AFM atomic force microscope Alg alginate AlgH alginate with high tyramine modification density AlgL alginate with low tyramine modification density AlgM alginate with medium tyramine modification density BSA bovine serum albumin CaM calmodulin CarHc C-terminal adenosylcobalamin binding domain CCP covalently crosslinkable protein CnaB2 immunoglobulin-like collagen adhesion domain 2 Cph1 Phytochrome-like protein DAPI 4',6-diamidino-2-phenylindole DC denatured crosslinking D-DC denatured crosslinking gel stored in unfolded state DMEM Dulbecco's Modified Eagle Medium DMSO dimethyl sulfoxide xxiv DN double network DNA deoxyribonucleic acid D-NC native crosslinking gel stored in unfolded state DOPA 3,4-dihydroxyphenyl- L-alanine Dronpa145N A Dronpa mutant with Lysine145 on β-strand 7 changed to Asparagine dsDNA double-stranded DNA DTT dithiothreitol E. coli Escherichia coli ECM extracellular matrix EDC N-(3-dimethylaminopropyl)-N’-ethylcarbodiimide hydrochloride EDTA ethylenediaminetetraacetic acid EdU 5-ethynyl-2’-deoxyuridine ELP elastin-like protein FL ferredoxin-like protein FLM23C a cysteine variant of FL G GB1 G block L-guluronic acid GB1 the B1 binding domain of protein G from Streptococcus GdHCl guanidine hydrochloride GL5 a loop insertion mutant of GB1 in which five residues (GGGLG) were inserted into the second loop of GB1 xxv GL5CC bi-Cysteine mutant of GL5 GO Graphene oxide GSH Glutathione HA hyaluronic acid HDPE high density polyethylene HLF human lung fibroblasts I27 the 27th Ig domain of human titin IAEDANS 5-((2-[(iodoacetyl)amino]ethyl)amino)naphthalene-1-sulfonic acid Ig Immunoglobulin I-N-DC denaturing-crosslinking-renatured with calcium cations IPTG isopropyl-1-β-D-thiogalactoside LB Luria-Bertani broth LOV2 Light-oxygen-voltage-sensing domain M block D-mannuronic acid MCP metal crosslinkable protein MEP mutually exclusive protein MES morpholinoethanesulfonic acid buffer Mfp Mussel foot proteins NC native crosslinking N-DC denatured crosslinking gel stored in folded state NHS N-hydroxysuccinimide ester N-NC native crosslinking gel stored in folded state xxvi OD optical density PAAm polyacrylamide PAMPS poly (2-acrylamido, 2-methyl, 1-propanesulfonic acid) PBS phosphate buffered saline PDB protein data bank PEG polyethylene glycol PEGDMA poly(ethylene glycol) dimethacrylate pHEMA 2-hydroxyethyl methacrylate PhoCl photocleavable protein PNIPAM Poly(N-isopropylacrylamide) PVA polyvinyl alcohol R resilin RNA ribonucleic acid Ru(II) ruthenium(II) [Ru(bpy)3]2+ ruthenium (II) tris-bipyridyl dication Sc SpyCatcher SDS-PAGE sodium dodecyl sulfate polyacrylamide gel SEM scanning electron microscopy sGFP split green fluorescent protein St SpyTag THPC Tetrakis(hydroxymethyl)phosphonium chloride TIP1 tax interactive protein-1 TNFn3 the third fibronectin type 3 domain of human tenascin xxvii Tris tris(hydroxymethyl)aminomethane buffer ULD ubiquitin-like domain unAlg unmodified alginate UVR8 ultraviolet-B receptor UV-Vis Ultraviolet-visible wt. weight xxviii Acknowledgements I offer my enduring gratitude to my supervisor Dr. Hongbin Li for his tireless support, skillful guidance, and stoic patience. He gave me constructive suggestions and helped as my thesis developed. He challenged my knowledge and encouraged me to become independent. In himself, Dr. Li has modelled what it is to be an academic professional and a creative researcher. I have been inspired and learned in many ways during my PhD study in his lab. I thank Dr. Suzana Straus, Dr. Jayachandran Kizhakkedathu and Dr. Edward Grant for enlarging my vision of science and providing guidance as my supervision committee. I especially thank Dr. Straus and Dr. Kizhakkedathu for critical evaluation of this manuscript. I would like to thank past and present members of the Li group. I owe particular thanks to Dr. Na Kong, who helped and trained me to initiate the first research project. I had many discussions with Dr. Shanshan Lyu, Dr. Chengzhi He, Dr. Ruidi Wang and Dr. Wenxu Sun about scientific research and technical details. Dr. Yi Cao, Dr. Jie Fang, Dr. Xiaoye Gao, Yabin Guo, Dr. Yongliang Wang, Han Wang, Tianyu Duan, Jiayu Li, Adam Xiao, Haoquan Lin, Jiacheng, Zuo, Qingyuan Bian and Jess Fong have provided valuable suggestions, help and encouragement. I also would like to thank Dr. Elena Polishchuk and Jessie Chen for their countless help and technical support. I thank Dr. Margo Lillie and Dr. Robert Shadwick for their generous help and assistance with Instron tests. I thank Dr. Amanda Haage and Martin Arostegui for sharing their skills and knowledge in mammalian cell culture and imaging. Thanks to all my friends for being supportive over the years. To Jason Hedges, thank you for all the scientific support and making my life happy and wonderful every day. Finally, special thanks are owed to my parents, Shichao and Xiaoling, for the unconditional love and support. xxix Dedication To my grandma, Wang Qingfang. 1 Chapter 1: Introduction Soft materials have gained great interest for researchers in the 21st century with the rapid development of material science and biomedical engineering. Soft materials generally refer to condensed matters that easily undergo deformations or structural alterations, when subjected to thermal fluctuations or external stresses at room temperature. Soft materials include gels, liquid crystals, colloidal assemblies, polymers, granular materials and other ordered superstructures.1-3 Additionally, soft materials are essential components of biological systems, such as muscles, tendons, brain tissue and fat.4-7 As one of the most typical soft materials, hydrogels are crosslinked three-dimensional polymer networks which are highly water-absorbent yet water-insoluble. Formation of hydrogels requires hydrophilic and crosslinkable groups on the polymer chains. The crosslinking can be generated by chemical covalent bonding or physical crosslinking (such as entanglements, hydrogen bonding, metal-ligand coordination, host-guest complexation and ionic interactions). Hydrogels have been widely studied and utilized in a variety of applications in multidisciplinary fields due to their exceptional properties, such as tunable mechanical properties, adhesive surface, self-healing effects, and sensitive responses to environmental stimuli. Owing to these unique features, diverse hydrogel-based novel technologies and devices have been developed, including drug delivery,8-9 tissue engineering materials,10-11 soft robots,12-13 artificial muscles and joints,14-16 smart adhesive materials,17-18 wound dressing and skincare patches,19-21 biosensors, and electrochemical devices.22-25 The study of hydrogels can be traced back to the pioneering work of Wichterle and Lim in 1960, who described the synthesis of crosslinked pHEMA (2-hydroxyethyl methacrylate) networks for biological use.26 This work opened up a novel methodology toward engineering soft 2 materials with high water content. Numerous designs of hydrogel networks have been reported to meet increasing biomedical needs. Most recently, hydrogels built from natural polymers, especially recombinant proteins/peptides as building blocks, have become attractive due to their biofunctions, designable amino acids sequence, monodispersed molecular weight, and excellent biocompatibility.27 Additionally, conformational switching of structural proteins can induce thermodynamic energy change in the network, leading to the energy dissipative ability of hydrogels. Despite these advantages, applications of recombinant protein-based hydrogels are often limited by their poor mechanical properties and low crosslinking density, considering steric hindrance of bulky protein domains in the networks. Comparing folded polyproteins with coiled macromolecular structure, synthetic polymer chains or low molecular weight gelators, which are more flexible and anisotropic, generally result in a higher degree of crosslinking and contain more physical entanglements in the polymer-based hydrogel network. To optimize the mechanical strength of structural protein-based hydrogels, it is important to understand the design principles of such crosslinking networks. This chapter focuses on introducing the latest research progress on protein-based hydrogels, including network designs, crosslinking strategies, especially their physical and mechanical features. Classic rubber elasticity theory is also presented, as it reveals the effect of elastomeric network structures on physical and mechanical properties of hydrogels. Based on the theory, I discuss the advantages and drawbacks of prevailing network design strategies for hydrogel mechanical reinforcement, as guidance for future design of protein-based hydrogels. 3 1.1 Protein- and peptide-based hydrogels Hydrogels have been constructed using diverse polymers. Typically, hydrogels can be divided into three categories based upon their building blocks: artificial synthetic polymers, natural polymers, and synthetic-natural hybrids. Owing to sustainable sources, biocompatibility and biodegradation, natural polymer hydrogels have been extensively studied, such as protein-based hydrogels, polysaccharides hydrogels, and DNA hydrogels. For biomedical purposes, one of the most promising biomolecules for gel fabrication is noted as proteins/peptides, since soft tissues in human bodies are largely composed of numerous proteins. Proteins/peptides, such as naturally derived collagen, gelatin, silk fibroin, as well as synthetic peptides and globular proteins, have been employed to develop a vast array of biofunctional materials to mimic natural soft tissue and environmental responsiveness. While biologically relevant performance is surely important, mechanical behaviors of protein-based hydrogels are also essential for biomedical applications. These features are mainly governed by hydrogel building blocks and network architectures. Herein I present representative network design by categories of protein- /peptide-based building blocks. 1.1.1 Synthetic protein-based hydrogels 1.1.1.1 Synthetic elastomeric globular protein-based hydrogels Globular proteins are the most abundant proteins in nature, displaying an enormous variety of biofunctions. Unlike synthetic polymers or intrinsically disordered polypeptides, globular proteins possess compact three-dimensional structures, mostly comprised of folded α-helices and β-sheets, which are generally stabilized by noncovalent interactions. The folded domains can undergo dynamic unfolding transitions under mechanical stress. The dynamic unfolding transition engenders significant changes in free energy, effective chain length and flexibility of proteins. The 4 use of globular proteins as building blocks provides elasticity, toughness and large energy dissipation to the hydrogels. Additionally, benefiting from modern molecular biology and protein engineering, these techniques allow for exquisite control over protein sequence, structure, and function at the molecular level. Protein engineering allows for the rational design of hydrogel building blocks by combining different globular domains. Figure 1-1 Tandem modular elastomeric polyproteins for hydrogel engineering. A) Schematic Ru(II)-mediated photo-crosslinked network of the GB1-Resilin based hydrogel. (Adapted from Ref28) B) Young’s modulus of different constructs of GR-based hydrogels.29 C) Representative tensile stress-strain curves of GRG5RG4R hydrogels.28 Copyright 2010, American Chemical Society. Lv et al. first engineered a tandem modular polyprotein hydrogel by combining the artificial elastomeric protein GB1 domains (PDB: 1PGA) and unstructured resilin to mimic the passive elasticity of muscle.29 The muscle’s passive elastic behavior is regulated by the I-band section of the giant muscle protein titin, which contains a vast array of immunoglobulin-like (Ig) domains and certain unstructured sequences. GB1 is a small globular protein derived from the B1 IgG binding domain of protein G from Streptococcus30. GB1 displays comparable mechanical 5 stability to the Ig domains found in titin. To construct a hydrogel network, the tandem modular polyprotein GB1-Resilin in solution was crosslinked through a Ru(II)-mediated photochemical reaction where tyrosine residues in close proximity form a dityrosine that served as crosslinking sites (Figure 1-1A). Mechanical characterizations and single molecule techniques were employed to study the rubber-elasticity behavior of the hydrogel under stretching loads. During stretching, mechanical unfolding of some GB1 domains occurred in the hydrogel network which dissipated energy, thereby preventing fracture from overstretching. GB1 domains refolded quickly and spontaneously upon withdrawal of the stretching stimuli. These characteristics of the GB1-resilin network closely mimicked the I-band part of titin in muscle tissue at a molecular level, while the macroscopic artificial hydrogels showed high resilience at low strain, fast hysteresis recovery and desirable elasticity with Young’s modulus reaching ~65 kPa (Figure 1-1B and C). Inspired by this pioneering study, many strong and extensible hydrogels have been successfully developed using globular proteins as energy dissipative building blocks. The Popa group has demonstrated that photo-crosslinked BSA-based hydrogels can be used to make gels with tunable and controllable elasticity.31 Additionally, Da Silva and co-workers explored the rheological properties and small-angle scattering response of hydrogels obtained from mechanically robust globular polyproteins, I27 and protein L.32 Moreover, guided by single-molecule force spectroscopy at a molecular level, Wu et al. demonstrated that it was possible to semi-quantitatively design synthetic protein hydrogels with tailored mechanical performance by considering the mechanical properties of the cross-linkers and the load-bearing modules.33 Similarly, Fang et al. reported the engineering of a highly elastic and tough hydrogel based on extremely mechanically labile ferredoxin-like polyprotein domains (FL)8. The (FL)8 hydrogels were demonstrated with a Young’s modulus of ~16 kPa, breaking strain of 450%, and toughness 6 of 57 kJ/m3 (Figure 1-2).34 Similarly, the Ru(II)-mediated photo-crosslinking strategy was applied to constitute the gel networks. FL domain exhibited very low mechanical stability and only 5 pN of force was adequate to unfold the domain, resulting in the finding that a large portion of FL domains in the hydrogels were unfolded by the force generated by swelling. Furthermore, they suggested that possibly hydrophobic collapse and aggregation of unfolded FL introduced extra physical entanglements into chemical crosslinked networks leading to the marked stretchability, energy dissipation and toughness of the hydrogel. Figure 1-2 Tensile properties of (FL)8 hydrogels. A) Representative stress-strain curve of (FL)8 hydrogels loading to failure, showing the excellent extensibility. B) The loop area between the stretching-relaxation curves of (FL)8 hydrogels indicates massive energy dissipation. Additionally, plastic deformation occurs at high stretching stain. C) and D) Kinetics of the hysteresis shows that the hysteresis can be mostly recovered within the first 2 secs.34 Copyright 2013, Macmillan Publishers Limited. 1.1.1.2 Synthetic dynamic protein-based hydrogels In a classic hydrogel network, the functionalization of hydrogels is normally realized by decorating, inserting or introducing functional groups to network chains that are mainly static 7 systems and only undergo simple degradation. As the field of biomedical engineering has developed, one of the most attractive functionalities of hydrogels is the use of artificial extracellular matrices (ECM), which support and direct cellular responsive behaviors by providing biocompatible and dynamic local microenvironments. Through cell-ECM interactions, these external mechanical signals are transduced to the cell’s interior, triggering biochemical signaling cascades that determine a wide variety of cellular behaviors, ranging from cell adhesion, cell spreading to cell differentiation.35-37 However, use of relatively inert biomaterials, which only provide a simple biomimetic environment for cell culture, limits studies related to cell interactions induced by mechanical cues from local matrices. Moreover, conventional hydrogels with dynamic properties generally involve irreversible or uni-directional degradation or disconnection of networks through artificial triggers, such as pH, temperature, light, and additional chemical agents.38-40 Fortunately, utilization of dynamical globular proteins with conformational switching as building blocks fulfills the requirements and opens up a novel avenue for dynamic hydrogel design. Conformational change of proteins in the networks can not only reversibly tune certain properties of the materials, but also maintains the stability of the overall system, while degradation and disruption of polymeric primary networks typically lead to gel-sol phase transition. An early example from the Murphy group selected the well-characterized protein calmodulin (CaM) as a dynamic unit which undergoes a hinge-like conformational transition controlled by ligand binding within PEG-CaM-PEG gel networks, shown in Figure 1-3.41-42 CaM exhibits an extended dumbbell shape in the absence of calcium ions, while it undergoes a rapid transition from the extended conformation (5.2 nm) to a collapsed conformation (1.4 nm) in the presence of calcium. Consequently, the addition of calcium to the hydrogel drove contraction of the crosslinking networks leading to shrinkage of the hydrogel. The authors developed a systematic 8 approach to achieve reversible variation in network density, showing dynamically tunable responses at a macroscopic level. Figure 1-3 A dynamic hydrogel network based on the conformational change of calmodulin (CaM). A) Conformational switching of CaM between the extended and the collapsed state by adding or removing ligands. B) Schematic network structures of the dynamic PEG-CaM-PEG gel hydrogel.42 Copyright 2008, WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim. In principle, the greater the difference in size between two conformational states of a protein unit, the greater the responsiveness of the hydrogel. This was demonstrated by Kong and coworkers, who reported an entirely protein-based hydrogel with a relatively large change in Young’s modulus.43 This dynamic protein-based hydrogel was fabricated by using mutually exclusive proteins (MEP, GL5CC/I27) as building blocks. MEP is a specially designed, redox-responsive protein-folding switch. In the oxidized state, the host domain GL5CC is folded; in the reduced state, the GL5CC unfolds and the effective chain length of the MEP changes from 2.6 to 26 nm, leading to a change in the polymer chain length between two crosslinking points. This dynamic hydrogel could be reversibly tuned between a stiff and soft state via a redox reaction, and the Young’s modulus of the hydrogels could be reversibly cycled between ~10 kPa and 40 kPa. 9 Additionally, the dynamic change of the MEP hydrogel in terms of Young’s modulus, swelling ratio as well as resilience are proportionally dependent on the protein concentrations. Building upon these successful hydrogel developments, it is possible to engineer dynamic protein-based hydrogel matrices mimicking soft tissues for cell-ECM interaction studies in vitro. Figure 1-4 Representative protein-protein interaction strategies that have been utilized for hydrogel engineering. A) Coiled coil binding (PDB: 2XU6). B) SpyTag-SpyCatcher chemistry by forming an isopeptide bond (PDB: 4MLI). C) Split florescent protein binding (PDB: 2B3P). D) Docking and dimerization domain and anchoring domain self-assembling (PDB: 2H9R). 1.1.1.3 Synthetic protein hydrogels based on protein assemblies Protein-protein interactions and protein-ligand binding primarily exist in all living systems and play essential roles in cell signaling, membrane transportation and immune response. Taking advantage of these spontaneous assemblies and interactions, numerous efficient crosslinking 10 strategies have been developed to engineer protein-based hydrogels under mild physiological conditions (Figure 1-4).44-47 However, there are prerequisites to consider when choosing interaction pairs. Firstly, suitable interaction pairs for hydrogel network formation should be relatively small in size, have good solubility and low steric hindrance. Secondly, high affinity and strong mechanical stability ensure optimal mechanical strength of the hydrogels. Thirdly, highly dynamic binding/unbinding processes may introduce sheer thinning or self-healing properties, but the hydrogels may tend to be unstable and weak. A successful hydrogel example which makes use of a protein-protein interaction was reported by Lin and co-workers based on 4-arm star-like tetrameric recombinant proteins.48 The formation of the protein hydrogel network was instantly completed through the specific non-covalent protein-peptide interaction by simply mixing the recombinant proteins ULD-TIP1 and ULD-GGGWRESAI (Figure 1-5). The resulting storage modulus from rheology tests was ~ 10 Pa of a 2 wt% resultant gel, suggesting the formation of a soft gel. 4-arm protein-based networks can be also realized by using the spontaneous reconstitution of the split green fluorescent protein (sGFP).49 Combining sGFP with SpyCatcher-SpyTag chemistry, which is described later in this section, (SpyCatcher)4-GFP and (SpyTag)2 led to the creation of hydrogels with storage modulus of ~ 1 kPa. Moreover, photo-responsive proteins are of special interest for hydrogel preparation, as they can serve as crosslinkers, allowing reversible control of sol-gel transition. Gelation strategies promoted by photo-responsive proteins are divergent, mainly including photoinduced association/dissociation of protein oligomers (CarHc, Dronpa145N, UVR8, Cph1),50-53 protein photolysis (PhoCl),47 and photoinduced conformational switching (LOV2).54 The photoresponsive hydrogels based on the first method exhibit reversible changes in mechanical properties or sol-gel transitions, but the gels typically have fast erosion rate due to the unstable physical crosslinks. The 11 second mechanism, such as PhoCl, can be programmed to photoweaken the hydrogels, resulting from the irreversible photocleavage of the proteins. The third method contributes to favorable photo-sensitivity in terms of kinetics. LOV2 based hydrogels showed fast and reversibly dynamic change of mechanical properties upon cyclic light exposure.55 Figure 1-5 4-arm star-like tetrameric recombinant protein-based hydrogel based on specific protein-protein interaction. A) Schematic three-dimensional networks for hydrogelation via the specific interaction between TIP1 and hexapeptide WRESAI. B) Rheological results for the gel sample made of equal molar ratio of ULD-TIP-1 and ULD-GGGWRESA. 48 Copyright 2016, Elsevier B.V. Most protein-protein interactions provide physical crosslinking methods. These hydrogels exhibit injectable and self-healing behaviors but weak in stiffness and strength, due to the weak crosslinking interactions commonly driven by electrostatic forces, hydrogen bonding and the hydrophobic effect. Thus, covalent bonding is an alternative approach to achieve strong and stable hydrogel materials. A variety of genetically encoded click chemistries have been discovered, such as the SpyCatcher-SpyTag system, the SnoopCatcher-SnoopTag system, Halo-tag chemistry, SNAP-tag chemistry, and CLIP-tag chemistry.56 These peptide/protein reactions can spontaneously form covalent linkages with high efficiency and selectivity under mild physiological conditions.56 Particularly, the SpyCatcher-SpyTag system has been widely applied 12 for entirely protein-based hydrogel engineering.49, 51-52, 57-60 SpyCatcher (Sc) and SpyTag (St) are two split peptide fragments derived from the second immunoglobulin-like collagen adhesion domain CnaB2 of Streptococcus pyogenes. This pair of genetically encoded reactive partners spontaneously form covalent isopeptide linkages and reconstitute the intact folded CnaB2 domain without additional chemical reagents or catalysts (Figure 1-4).61-64 Recently, the Tirrell group have reported covalently crosslinked hydrogels by employing St-Sc chemistry.57 St and Sc domains were placed within elastin-like proteins (ELPs), while the control group had an Asp-117-to-Ala mutation in one of the St domains. The mutated St only enabled noncovalent recognition between St and Sc. The rheological study confirmed gel formation in both groups with storage modulus ranging from 30-300 Pa. However, the slower erosion rate and higher storage modulus of the wild type Sc group indicated the covalent nature of the St-Sc network. Furthermore, Li and colleagues have demonstrated the utility of this reaction to engineer globular protein-based hydrogels for 3D cell encapsulation and protein release.58 They constructed two tandem modular proteins TNfn3-(GB1-Sc)3 and (GB1-St)4. The GB1 domains ensured good solubility and yield of multidomain expression in bacteria, while the TNfn3 domains contained an RGD sequence responsible for cell adhesion. At an Sc:St stoichiometry of 1:1, the reaction between these two multifunctional polyproteins led to the formation of a protein network, as TNfn3-(GB1-Sc)3 and (GB1-St)4 were considered tri- and tetra-functional. Rheological characterization of the 10-15 wt% protein mixture showed an increase in stiffness, indicating the gelation process with final storage moduli range of 300 to 600 Pa. The mechanical results of the weak St-Sc gels were similar as the St-Sc-ELPs gel system from Tirrell group.57 They suggested that the relatively low storage modulus may be attributed to the incomplete crosslinking of St-Sc, due to high viscosity of the highly concentrated protein mixture and possible steric hindrance. It 13 is critical to overcome this potential limitation and to tune the mechanical properties of St-Sc protein hydrogel system for biological applications. 1.1.2 Synthetic polypeptide-based hydrogels Using structural proteins as building blocks in hydrogels might give rise to certain disadvantages. Poor solubility and low stability of folded domains may be problematic during protein purification and the gelation process. Steric hindrance of bulky domains might also lead to inefficient crosslinking. As an alternative approach, synthetic polypeptides reveal their own advantages. The genetically encodable sequence of polypeptides provides precise control over hydrophobicity, net charge, as well as desirable biofunctions for biomaterial applications. Flexible peptide chains lower the effect of steric hindrance and introduce physical entanglements in gel networks, which can be critical for improving mechanical properties. Gelation of peptides is similar to the self-assembly of low molecular weight gelators, which are commonly used to form polymer-based hydrogels.65-68 These short polypeptides or oligopeptides self-assemble into micro-bundles/fibrils and then crosslink into entangled networks of hydrogels. According to their gelation mechanisms and network architectures, peptide-based gelators can be generally classified into four categories: ionic-complementary peptides, coiled-coil peptides, amphiphilic peptides, and β-sheet forming peptides. Typical ways to trigger self-assembly of peptide gelators include temperature, pH, solvent, enzymatic reactions and chemical agents.69-70 Many peptide-based hydrogels undergo reversible sol-gel transitions in response to these external stimuli. However, reversible peptide-based hydrogels often show low mechanical stability and deficient elasticity due to weak physical crosslinking networks. Biomedical applications of hydrogels require crosslinking strategies to 14 stabilize the hydrogels for in vivo uses. Thus, researchers have devoted significant efforts to improving mechanical performance and stability of peptide hydrogels. 1.1.2.1 Elastin-like polypeptide-based (ELP) hydrogels Elastin is an essential natural protein building up ECM for cellular interactions/regulation and forming elastic tissues like skin, blood vessels and bladder.71-72 Elastin provides ECM with elasticity and mechanical energy storage, allowing for extension upon stretching and recovery after stress release. However, natural elastin has hardly been used for hydrogel fabrication due to issues such as contamination, calcification and low solubility during synthesis and purification.73 Natural elastin has nevertheless inspired the design of many ELPs for biomaterial applications with the inherent properties of natural elastin, but with decent solubility and large yield through bacterial expression. These ELPs contain the characteristic pentapeptide repeating unit, valine-proline-glycine-X-glycine, in which X can be any natural amino acid residue except proline.72 One of the substantial traits of ELPs is temperature-dependence. The critical temperature of the phase transition of ELPs can be tuned by the choices of X residues as well as the number of pentapeptide repeats in the ELP sequence. At a low temperature, the environmental water surrounds the hydrophobic residues, and ELPs are soluble with a random coiled configuration. Above the inverse transition temperature, solubility of ELPs decreases due to the less-ordered surrounding water, triggering hydrophobic collapse and aggregation of ELPs.74-76 This characteristic of ELPs makes it possible to engineer recombinant protein-based hydrogels with similar behaviors as the thermoresponsive polymer Poly(N-isopropylacrylamide) (PNIPAM)-based hydrogels. The inverse phase transition of ELPs provides a thermosresponsive gelation approach. The Chilkoti and the Urry groups have extensively studied the responsive behaviors and the biomedical 15 applications based on ELPs.77-80 Although ELP-based materials were demonstrated with low cytotoxicity, the weak mechanical properties were still a limitation. A major approach to enhance the mechanical performance is to introduce chemical crosslinking into the ELP sequence, which provides well-defined locations and a quantifiable amount of added functionalized groups.81-87 These chemically active groups can be provided by recombinant incorporation of amino acids within the ELP sequence. Figure 1-6 ELP-PEG-based rubber like hydrogels. A) The chemical structure of ELP (VPGVG) repeats (repeating units n= 50, 75, 100) and 4-arm PEG/NHS. The C- and N- terminal amine on the ELP are designed for the PEG crosslinking reaction. B) The amine groups on both sides of ELP reacts with the 4-arm PEG crosslinker to achieve the hydrogel network. When the temperature is lower than the inverse transition temperature of ELP, the hydrogel is highly swollen and the chains act as a loose spring with low extensibility. However, the hydrogel shrinks at a temperature higher than its inverse transition temperature (37 ℃), resulting from the contraction of ELPs. The compact ELPs allow the gels to undergo large deformation. C) The reversible chain interaction contributes to hysteresis during cyclic tensile test.88 Copyright 2016, American Chemical Society. The Khademhosseini group described a photocrosslinked ELP-based hydrogel that consisted of 70 repeats of the ELP pentapeptide sequence and a flanking sequence (Lys-Cys-Thr-16 Ser).89 The flanking sequence containing cysteine residues provided thiol groups to form disulfide bonds as chemical crosslinking sites, resulting in promotion of chain extension and improvement in the mechanical properties of the ELP gel. The resulting hydrogel showed excellent extensibility (a 400% strain at fracture). The tensile strength and compression modulus of the hydrogel increased with the protein concentrations because of additional crosslinkable groups. Another way is to introduce chemical modifications in ELPs. Desai et al. designed ELPs composed of 50 to 100 pentapeptide units of “VPGVG” as well as lysine residues at the N- and C-termini.88 They employed 4-armed polyethylene glycol/ N-hydroxysuccinimide esters PEG/NHS as crosslinkers reacting specifically with the amine groups in the lysine side-chain to achieve end-to-end tethered chains (Figure 1-6). At low temperatures, ELP chains in the gel network behaved like loose springs with low extensibility, while ELPs contracted into tightly packed coils at a temperature higher than 37 ℃ allowing for extreme deformation. Additionally, these near ideal networks with few ineffective free ends greatly enhanced mechanical behaviors of the material. The hydrogels were demonstrated with extremely high stretchability (~1500%), elastic modulus of 20 kPa and high resilience. Moreover, control over ELP sizes not only changed the transition temperatures, but also helped regulate extensibility of the ELP gels by tuning the intercross-link distance. 1.1.2.2 Hydrogels based on coiled-coil domains. Self-assembly of synthetic peptides presents the opportunity to bridge nano-to-millimeter scale regimes for rational design of protein-based biomaterials. Indeed, a variety of peptide-based hydrogels via non-covalent peptides interactions have been produced.90-97 Much of these efforts have been devoted to α-helical gelling systems, which are based on the formation of a coiled-coil structure. Structurally, coiled-coils are right- or left-handed superhelices composed of two or more strands of α-helices. The primary sequence of coiled-coil motifs typically involves a characteristic 17 7-residue heptad repeating unit (abcdefg), shown in Figure 1-7A.98 Position a and d are occupied by hydrophobic residues, providing interhelical hydrophobic interaction. The hydrophobic interface between two helices promotes aggregation into a coiled-coil dimer. Charged amino acids located at position e and g contribute to pH-dependent stability through electrostatic interactions. Residues b, c and f are usually charged/neutral amino acids, which are exposed on the surfaces of the structure. By putting responsive residues at positions a, d, e, g, the resulting coiled-coil self-assemblies can be used as physical crosslinking sites for hydrogel formation, which can be responsive to a range of external stimuli, including environmental pH, temperature, ionic strength, and chemical denaturants (Figure 1-7B).99-102 The pioneering studies from the Tirrell group employed coiled-coil formation of a triblock protein for hydrogel fabrication. The triblock protein AC10A was designed with a middle random coil domain (C10) flanked by two associative leucine zipper domains (A) at both ends. Self-assembly of the leucine zipper A domains resulted in coiled-coil-rich soft hydrogels. The soft gels could dissociate into solutions at high pH and high temperatures. They also found that the erosion rate of these hydrogels could be tuned by controlling selective molecular recognition, discrete aggregation number and orientational discrimination of the coiled-coil protein domains. Besides the gel-sol transition, the self-healing property is another interesting characteristic of these physically associating gels.103-104 Pine and colleagues demonstrated hydrogels composed of amphiphilic block copolypeptides.105-106 The gelation depended on the charged hydrophilic segments as well as the formation of stable chain conformations, namely both the α-helical and β-strand structures. Additionally, it was found that the hydrogel was able to restore 80-90% of its initial stiffness in a short period (~10 sec) of time, after a breakdown under high shear amplitude (1000%). This remarkable speed of elastic recovery was mainly attributed to fast reformation of 18 packed α-helices in the dynamic network. Additionally, coiled coils can also behave as crosslinkers in synthetic polymer-based hydrogels to modulate the mechanical properties. Yu et al. reported a PEG-based hydrogel modified with collagen mimetic peptides as a 3D scaffold. The collagen mimetic peptides not only provided cell adhesion motifs, but also served as physical cross-links. The covalently linked peptides underwent a thermal-sensitive conformation transition and folded into triple helices, promoting the gelation process. The elasticity and stiffness of the matrices could be controlled by temperature and the concentrations of additional free collagen mimetic peptides.107 The dynamic mechanical properties of coiled-coil and other self-assembling peptide hydrogels can be well-characterized by rheological methods. Figure 1-7C shows typical dynamic moduli curves for such materials.98 An oscillatory frequency sweep reveals the responsiveness of the hydrogels to mechanical shear. The storage modulus G’ is observed to plateau at high frequency, while a peak in the loss modulus G” and a crossover point appear upon decreasing frequency. The crossover where G’ meets G” indicates the threshold of physical gelation, marking the transition from prevalently viscous to elastic properties.105-106, 108-109 The frequency of the crossover provides a precise measure for the dynamics and relaxation behaviors of the physical crosslinking sites. A high crossover frequency implies highly dynamical interconnections but unstable hydrogel networks. For these associating peptide hydrogels, the plateau of storage moduli G’ are commonly in the range of 100-5000 Pa, and the crossover frequencies are usually within a range of 0.01-1 rad/s. Although coiled-coil and other associating peptide hydrogels show relatively low stability and weak mechanical properties, the self-healing and sheer thinning properties are widely utilized for drug delivery, protein release and injectable printing systems.19, 110-111 19 Figure 1-7 Hydrogels based on coil-coiled domains. A) Schematic representation of the coiled-coil motif. B) The formation of coiled-coil structure generates a hydrogel network. The sol-gel transition can be reversibly governed by external stimuli. C) Representative dynamic mechanical properties of self-assembling peptide hydrogels. 98 Copyright 2019, American Chemical Society. 1.1.2.3 Hydrogels based on β-sheet motifs In addition to the coiled-coil, β-sheet forming peptides are another ubiquitous structure utilized in hydrogels. This motif has been found in various natural proteins such as silk fibres and amyloid aggregates that cause neurodegenerative disorders.112 The formation of β-sheet bundles conventionally involves the assembly of alternating hydrophobic and hydrophilic amino acids driven by hydrophobic interactions and π-π stacking of the aromatic residues. The charged amino acids form hydrogen bonds with surrounding water, while the nonpolar residues are sheltered in the middle. Such amphiphilic alignment leads to the formation of β-sheet bilayers where lateral aggregation occurs to obtain fibrils/fibers (Figure 1-8A).113 The Scherman group designed three pentapeptide sequences composed of aliphatic isoleucine residues (Ile, I) and aspartic acid (Asp, D).114 Upon the addition of protons, charge recognition and hydrogen bonding promoted β-sheet self-assembly and the gelation process. The gels demonstrated rapid self-healing and shear 20 thinning behaviors. Also, it was found that the stiffness of the hydrogels could be modulated in a range of 2-200 kPa by altering the charge distribution in the peptide sequences as well as the gelation concentration. The hydrogels stiffening mechanism was uncovered byscanning electron micrographs of gel microstructures, which showed the hydrogels contained high-aspect-ratio nanofibers and further entangled structures (Figure 1-8B). Distinct from the self-associating peptides discussed in the previous section (1.1.2.2), β-sheet-rich materials display much stronger and tougher mechanical behaviors, correlating with the packed fibrillar structures, ranging across length scales from nano-to micrometers. Another example of β-sheet crosslinked hydrogels was constructed by tandem repetition of the disordered squid ring teeth protein, and displayed elastic moduli on the order of megapascal.115 The monomer of the squid ring teeth protein is comprised of an N-terminal disordered peptide segment and a C-terminal β-sheet-forming sequence. The physically crosslinked hydrogels induced by β-sheet assembly contained amorphous regions and small β-sheet crystallites, where the crystalline β-sheet domains were responsible for the mechanical strength and the disordered chains contributed to energy dissipation. The mechanical tests indicated gel toughening at high tensile strain due to strain-induced β-sheet alignment. β-sheet assembly can also function as a strong physical cross-linker in polymer networks and improves mechanical performance of polymer networks. Stevens and coworkers presented a peptide-polymer hybrid hydrogel by grafting β-sheet motifs into poly(γ-glutamic acid) polymer networks.116 The nanofiber-forming β-sheet peptides contained an N-terminal azide and two Asp flanking a central Ile region (N3-D2I4D2). Copper-catalyzed click chemistry was used to modify the peptides onto the polymer backbones. The authors demonstrated that the mechanical properties of the hybrid gels could be tailored across an order of magnitude (10-200 kPa) by altering the concentration of the hybrid peptide-polymer and grafting density of the β-sheet peptide. Moreover, 21 this polymer-peptide hydrogel showed self-healing and biodegradable properties, providing a potential platform for tissue engineering. Besides these hydrogels based on β-sheet forming oligopeptides, silk fibroin is a natural structural protein that form β-sheet-rich fibrils responsible for formation of physical hydrogels.117 Silk fibroin hydrogels exhibit supreme mechanical properties. This will be discussed in the next section (1.1.3) by providing more details on this category of the natural protein-based hydrogels. Figure 1-8 Packed fibrillar network structures made of β-sheet motifs. A) Illustrative scheme of formation of β-sheet based fibrillar hydrogel network. The peptide chains self-assemble and pack into β-sheets, followed by formation of fibrils and fibres, and then form a crosslinked fibril gel network.113 Copyright 2016, Adapted with permission from The Royal Society of Chemistry. B) Scanning electron microscope images of the pentapeptide β-sheet-based gel networks. The networks are composed of high aspect ratio entangled nanofibers.113 Copyright 2018, Wiley-VCH Verlag GmbH & Co. KGaA. 22 1.1.3 Natural protein-based hydrogels 1.1.3.1 Silk-based hydrogels Silk is a natural fibrous protein material produced by insects such as silkworms and spiders.118-119 Silk proteins show variation in composition, structure, and properties from different species. Silk fibroin produced by the silkworm Bombyx mori is the most extensively studied silk protein. The raw silkworm silk consists of two primary proteins, fibroin and sericin. Sericin coats onto the surface of the central fibroin, binding two parallel fibroin fibers together. Silk fibroin is comprised of a glycoprotein, a light and a heavy chain polypeptide, which are linked by a disulfide bond. The heavy chain (~390 kDa) primarily contains the hydrophobic repeat sequence GAGAGS, contributing to strong intra- and intermolecular interactions and highly ordered β-sheet crystalline regions.120 The hydrophilic light chain forms disordered amorphous regions serving as linkers to the crystalline regions. The β-sheet crystallites are associated with strength and stiffness,121 while the amorphous regions contribute to the extensibility of silk. Incorporation of each component results in outstanding strength and elasticity comparable to those of silk protein fibers (Young’s modulus 1-17 GPa).122-126 Predominance of β-sheet assembly promotes gelation of silk fibroin above its critical gelling concentration. Rapid gelation of silk fibroin-based hydrogels have been achieved by lowering pH,127-128 changing temperature,129-130 applying shearing forces,131-133 adjusting ionic strength or chemical crosslinking.134-135 Hydrogels made of silk fibroin alone or combined with natural peptides/proteins have been validated to be cytocompatible for biomedical applications. Motta et al. developed injectable 3D hydrogel scaffolds to direct differentiation of neural stem cells.136 The silk fibroin-based hydrogels were functionalized using covalently linked laminin-derived peptides (IKVAV), which are responsible for cell adhesion and neurite outgrowth.137-138 23 The IKVAV-modified silk fibroin hydrogels were prepared by sonication-induced gelation. The resulting gels showed a stiffness comparable to brain tissue (~1 kPa), and supported cell viability and neuron differentiation. Cell encapsulation matrices with a larger tunable range of stiffness were achieved by a combination of silk fibroin/collagen. Kim and colleagues reported a composite silk-collagen hydrogel for cell encapsulation with collagen providing cell adhesion motifs and silk fibroin enhancing the mechanical strength. The elastic modulus of such matrices could be tailored in a range from 0.1 to 20 kPa by adjusting the ratio and concentration of the two components.139 Moreover, silk fibroin-based hydrogels with extremely high toughness have been achieved through gel networks reinforced by horseradish peroxidase crosslinking, which introduces chemical crosslinking sites and double-network (Section1.2.2.4). The Kundu and the Yang groups reported silk fibroin/polyacrylamide hybrid hydrogels respectively, demonstrating improved mechanical strength as well as satisfactory biocompatibility for biomedical applications.140-141 In another study from the Kundu group, the silk fibroin was blended with polyvinyl alcohol in a 1:1 molar ratio, leading to an increase in the Young’s modulus from 50 MPa (fibroin only gels) to 146.7 MPa, which arose because of additional elastic polymer networks and chemical crosslinking sites.142 Owing to its unique mechanical strength and biocompatibility, silk fibroin is an interesting polymeric biomaterial that has been extensively studied for controlled release, tissue engineering and regeneration.143 However, the lack of cell adhesion motifs and bioactive cues requires silk fibroin matrices to be further modified. For cell-ECM studies and tissue engineering, proteins derived from natural ECMs, such as collagen and gelatin, are relatively more accessible. 24 1.1.3.2 Type I Collagen-based hydrogels Collagen is the most abundant structural protein found in assorted connective tissues in mammals, such as skin, cartilage and tendons. It is one of the main constituents in natural ECMs, comprising 25-35% of total protein in vivo.144-146 There are over 30 kinds of collagen, identified by variations in sequence, structure, function and distribution.147 Type I collagen is the most prevalent type and has been broadly used to synthesize artificial scaffolds, owing to its fibril-forming structure and abundant sources. Type I collagen consists of N- and C-terminal non-helical motifs and three α-chains intertwined into a right-handed triple helix in the middle.148-149 The formation of the triple helix is attributed to the presence of (Glycine-X-Y) repeats, where X and Y often are proline, hydroxyproline or glycine. It is essential to have glycine, proline, and hydroxyproline in their designated positions with the correct configuration to form and stabilize the structure.150 The peptides on both ends also help stabilize the helical structure by providing intermolecular cross-links. The triple helices are able to form aggregates resulting in mechanically stable collagen fibrils/fibers.151-152 Collagen fibrils in a neutralized aqueous solution incubated at 37 ℃ self-assemble and crosslink into hydrogel networks. In addition to the fibrous structures of the collagen hydrogels, several functional sites are found within collagen, such as integrin-cell binding sites, matrix metalloproteinases cleavage sites, intermolecular crosslinking sites and metal-binding sites.146, 153 The combination of all these features makes collagen a good candidate to prepare biomaterials for tissue engineering. However, mechanical weakness, thermal instability and rapid degradation are the limitations of collagen hydrogels. Thus, recent studies have focused on chemical modifications, which have been validated to be effective in reinforcing the strength of collagen hydrogels. The Chaikof group introduced photochemically cross-linkable acrylate moieties into native collagen, facilitating rapid and controllable matrix protein cross-linking.154 25 The triple-helical content of the collagen remained unchanged after the chemical modification and crosslinking reactions, but the denaturation temperature was increased beyond the physiological range. Also, the crosslinked collagen hydrogels showed a 12-fold increase in shear modulus (from 13.5 kPa to 162 kPa), while maintaining cell viability. Similarly, chemical modifications of collagen, such as the use of oxidized dextran, glutaraldehyde and carbodiimide, successfully improved the mechanical strength and the thermostability of collagen hydrogels.155-159 1.1.3.3 Gelatin-based hydrogels Gelatin is a protein mixture produced by acid or alkaline hydrolysis of collagen, where the triple helix of collagen is denatured into single-strand gelatin molecules. Gelatin adopts a random coil conformation in warm water, but undergoes a coil-helix transition when cooled. Above a critical gelling concentration, gelation of gelatin occurs, resulting in a thermo-reversible physical gel. Rheological results illustrated that the melting temperature and equilibrium modulus of gelatin gels are closely related to the concentration and the molecular weight. Also, such physical gels display both elastic solid and viscoelastic solution behaviors.160-161 However, the fast erosion rate often limits the use of gelatin gels. For long-term biomedical applications, it is necessary to consider the instability of the gelatin gels, which can be solved with additional crosslinks. Chemical derivatization and polymerization approaches have been used successfully to reinforce the properties of gelatin gels.162-165 Besides chemical treatments, non-toxic enzymatic crosslinking methods also successfully lead to hydrogel fabrication. Enzymes, such as transglutaminase and tyrosinase, introduce isopeptide bridges among the gelatin chains, acting as chemical crosslinks.166-169 Moreover, Zhou and coworkers reported a solvent displacement-induced toughening strategy. The 26 gelatin hydrogels showed significant toughening by displacing the water with glycerol, obtaining a strong glycerol-based gelatin organohydrogel.170 Although chemical treatments can improve the stiffness of gelatin gels, these typically involve complex processes as well as cytotoxic chemical residues. Alternatively, double-network approaches can also be used to accomplish mechanical enhancement, biocompatibility, biodegradability of gelatin hydrogels, by associating gelatin with natural polymer/proteins. The Dong group presented a tough silk-gelatin hydrogel prepared via enzymatically and physically double-crosslinked networks.171 The silk-gelatin hydrogels showed controlled degradation rate and suitable compressive and tensile moduli of 250 kPa and 210 kPa, respectively. It was evident that the composite hydrogels promoted proliferation and differentiation of bone marrow-derived stem cells for the potential application of bone regeneration. Annabi et al. demonstrated another biomedical application by incorporating methacrylate gelatin with methacryloyl-substituted recombinant human tropoelastin.172 Two ECM-derived biopolymers were simply mixed to obtain a sprayable pre-gel. Upon visible light-induced crosslinking, a class of tropoelastin/gelatin hydrogels were achieved with tunable elastic modulus, degradation rate and adhesive strength by varying gel concentration and the ratio of two components. This hydrogel-based bio-adhesive showed potential for clinical treatment of chronic wound healing. 1.1.3.4 Hydrogels made of other natural proteins In addition to the structural proteins mentioned above, here I briefly introduce several innovative designs of hydrogels made of other natural proteins. Bovine serum albumin (BSA) is a natural globular protein which has found plenty of applications, due to its low cost, high solubility, ligand-binding properties and stable structure.173-174 Khoury and Popa reported a smart BSA-based 27 hydrogel with programmable wide-range stiffness (10-64 kPa).175-176 Hydrogels were first formed from BSA solutions, where photo-crosslinking was applied to yield dityrosine adducts. These negatively charged BSA gels (pH 7.4) were subsequently moved to polyelectrolyte solutions, such as polyethyleneimine and poly-L-lysine. Protein-polymer hybrid hydrogels resulted from the physical absorption of polyelectrolytes into the BSA networks. Benefiting from polymer reinforcement and the BSA unfolding-refolding response to chemical denaturants, the approach enabled a reversible shape-memory behavior for electrolyte-treated BSA hydrogels. Besides the application of nonspecific binding and cross-linkable tyrosine residues, the exposed lysine residues of BSA provide the opportunity for gelation via glutaraldehyde crosslinking.177-178 In the presence of the enzymes glucose oxidise and catalase, glutaraldehyde can react with the lysine residues of BSA to form covalent bonds. This is a reversible reaction, resulting in a self-healing property of the glutaraldehyde cross-linked BSA hydrogels.179 Casein is the main protein component found in mammalian milk and comprises 80% of the protein found in bovine milk. Casein is actually a class of related phosphoproteins, including αs1-, αs2-, β-, and κ-caseins in an approximately ratio of 4:1:3.5:1.5. Caseins can self-assemble into micelles in the presence of calcium phosphate.180-181 However, the exact nature of the casein micellar structure is still under debate. One of the prevalent models suggests that the amphiphilic κ-caseins are found on the micelle surface, leaving the other three types of casein to be located in the interior.181-183 Moreover, acidification of casein micelles decreases electrostatic repulsion, which induces gelation of the micelles through physical interactions.181 Upon shearing, casein micelles show the capacity to dissipate energy by friction between casein micelles and rearranging its submicelles. Therefore, casein micelles are a good candidate for applications in the fabrication of tough hydrogels. Nam and Sun engineered a highly stretchable and notch-insensitive hydrogel 28 consisting of casein and polyacrylamide. The gel was prepared by a free radical polymerization of acrylamide and acid-coagulation of casein micelles. Although pure casein and pure polyacrylamide hydrogels are mechanically brittle, the shape recovery of the polyacrylamide network and the energy dissipation of the casein micelles both contribute to mechanical enhancement of the hybrid hydrogel. The resultant hybrid hydrogel exhibited deformation over 3500% strain under uniaxial tension with fracture energy of approximately 3000 J/m2.184 A hydrogel with similar casein-polyacrylamide design was demonstrated by the Gao group, showing a repeatable adhesive property at the surface of diverse materials with adequate adhesiveness and mechanical strength.185 Figure 1-9 Adhesion and cohesion of mussel byssal plaque. A) Mussel byssi attach on the underwater surface. Cartoon illustration of the six main types of adhesive-related Mfps making up the byssal plaque. B) Mussel adhesion and cohesion mechanism. Copyright 2020, WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim. 29 Mussel foot proteins (Mfp) are found in the family of marine mussels Mytilidae, which are found in cold sea water world-wide.186-187 Marine mussels strongly attach to various surfaces, withstanding the mechanical stresses that arise from tough underwater conditions. The strong adhesion of mussels to surfaces can be attributed to a bundle of sticky hair-like threads secreted by mussels, known as a byssus. Over 25 different proteins have been identified in byssi, out of which six main types of Mfps make up the byssal plaque (Figure 1-9A).188-189 These plaque Mfps share a similar abundance of the post-translationally modified amino acid 3,4-dihydroxyphenyl-L-alanine (DOPA). DOPA is crucial for mussel adhesion by forming both adhesive protein-surface bonding (bidentate hydrogen bonding, metal complexation, and hydrophobic interactions) and cohesive protein-protein interactions (bis- and tris-DOPA-Fe3+ coordinative complexes, DOPA-DOPA hydrogel bonding, aryloxyl radicalization, π-π stacking, and cation-π interactions), as shown in Figure 1-9B.190-191 The catechol group of DOPA readily undergoes chemically or enzymatically induced oxidation to form dopaquinone, which can trigger further protein crosslinking.192 These findings have inspired the design of diverse DOPA-based synthetic biomaterials for anti-corrosion coatings, saltwater anchors and adhesives.193-195 Benefiting from large-scale expression in Escherichia coli, DOPA-containing recombinant mussel foot proteins have been recently synthesized for hydrogel preparations. Cha and colleagues presented an adhesive hydrogel based on a recombinant mussel adhesive protein comprised of 12 repeats of type-1 Mfp. After enzymatic conversion of the tyrosine residues to DOPA, the gelation could be successfully achieved by using the DOPA-Fe3+ coordinative crosslinking or catecholic covalent cross-linking of DOPA. The mussel protein gels made by two different crosslinking methods showed distinct characteristics as bulk adhesives. The Fe3+-mediated noncovalent gels exhibited a fast gelation rate, deformable viscoelasticity and a self-healing behavior, with an adhesive strength 30 of ~130 kPa. On the other hand, the quinone-mediated gels had a higher adhesive strength (~200 kPa), owing to covalent cohesive bonding. In addition to the two crosslinking strategies of Mfps mentioned above, the Sun group employed type 3 Mfp since it possesses abundant tyrosine residues (20 mol%) that can undergo highly efficient dityrosine photochemical cross-linking, instead of oxidizing tyrosine to DOPA. The storage modulus of the resulting hydrogels could be tuned over a wide range of 1-9 kPa by controlling the protein concentration and dityrosine crosslinking density. Combining SpyTag/SpyCatcher chemistry, they were able to decorate the Mfp-based networks with biofunctional proteins for cell adhesion and proliferation studies.196 In summary, natural proteins are a rich source for building potential biomaterials (Table 1-1). However, many naturally sourced proteins show poor solubility, low yields, unstable structures, or are difficult to purify. Although promising, many challenges still remain in this area. It is essential to understand the sequence-structure-function relationship for biomaterial design. Table 1-1 Summary of the advantages and disadvantages of natural protein-based hydrogels in Section1.1.3. Building blocks Gelation Advantages Disadvantages Silk β-sheet assembly Rapid gelation, high mechanical strength, biocompatibility Lack of cell adhesion motifs Type I Collagen Formation of triple helices Functional binding sites, cell adhesion Mechanical weakness, thermal instability, rapid degradation Gelatin Coil-helix transition Cell adhesion, biocompatibility, thermo-reversible Mechanical weakness, thermal instability, rapid degradation BSA Additional chemical crosslinking Dynamic folding-unfolding, negatively charged Mechanical weakness, requiring additional chemical crosslinking and modification Casein Additional chemical crosslinking Highly stretchable, notch-insensitive Low stiffness Mfps DOPA-based crosslinking Fast gelation rate, self-healing, high adhesive strength Protein precipitation 31 1.2 Rubber elasticity theory: design principle of protein-based elastomer mechanics 1.2.1 Classic rubber elasticity theory Section1.1 covers the up-to-date mainstream studies of protein-/peptide-based hydrogels, including network engineering, gelation principles, and mechanical reinforcement methods for potential applications. Scientists have devoted great efforts to achieving desirable mechanical properties of hydrogels, because these are essential in many applications. For example, tissue engineering requires that the artificial materials match the stiffness and elasticity of disparate tissues, which can be as soft as brain tissues or as stiff as bones. Mechanical signals play important roles in governing cell differentiation and proliferation behaviors via cell-ECM interactions.197-200 It is evident that the design of molecular structures plays a key role in controlling functional and mechanical characteristics of materials. Typical hydrogels are made of water-swollen polymer networks, which are structurally and thermodynamically similar to elastomer/rubber networks. Thus, mechanical properties of hydrogels can be understood by using the theory of rubber elasticity. Rubber consists of crosslinked amorphous polymer chains above its glass transition temperature, displaying substantial stretchability and reversible deformation. Rubber networks allow for fast response to external stress with rearrangement of the polymer chains. Kuhn, Guth and Mark proposed the first model to describe the relationships between macroscopic deformation, chain extension and entropy reduction.201-203 Upon stretching, the polymer chains and displacement of crosslinks undergo affine deformation, meaning the chains are proportionally extended with the macroscopic strain (Figure 1-10A). In 1940s, Treloar, Guth and James clarified the quantitative relations between network extension and entropy reduction.204-208 As the polymer chain is extended, the number of chain configurations and conformations decrease. On the basis 32 of the second law of thermodynamics, Wall and Flory summarized the quantities for the thermodynamic equations of state for rubber elasticity:209-211 Equation 1-1 ' = )*+*, -.,0 − 2 )*3*,-.,0 Equation 1-2 ' = )*+*, -.,0 + 2 )*'*2-5,0 Equation 1-1 depicts that the retraction force ( f) and entropy (S) are related for an isovolumic (V) ideal elastomer. T is the absolute temperature of the elastomer system. It is noted that the bonds are not affected by change in length (l) and the internal energy is independent of chain conformations. Thus, the internal energy term (67875 9.,0) can be assumed to be zero. In brief, the retractive force of an elastomer in response to a tensile or shear deformation arises through a decrease in entropy, as the coiled network chains are stretched along the force direction. Upon removal the external stress, deformation recovery of the elastomer is driven by increasing in entropy toward the most probable end-to-end distance of the chains in a rapid fashion. As derived by Wall, Equation 1-2 illustrates the dependence of the retractive force and the temperature of an elastomer. For an isovolumic ideal elastomer under a fixed extension, the retractive force increases linearly with the temperature. The slope of the f-T curve is proportional to −67;759.,0 or 67<7.95,0. 33 Figure 1-10 The rubber elasticity theory based on affine deformation and statistical thermodynamics. A) Relaxation and extension of a polymer chain segment in an elastomer network, that undergoes affine deformation upon an external force. B) Typical experimental stress-strain behavior of a lightly crosslinked elastomer is compared to the statistical thermodynamic prediction by using Equation 1-3.212 Benefiting from a statistical thermodynamic approach, the relationship between an elastomer network and its retractive stress (σ) during simple extension has been developed with an equation of state for macroscopic networks: Equation 1-3 = = >?@)A − BAC- where λ is the extension ratio, and RT is the gas constant times the absolute temperature. The quantity n describes the number of active network strands per unit volume of the network and can also be written as n=ρ/Mc, where ρ is the density of polymer and Mc is the molecular weight between cross-links, describing mesh size or crosslinking density of the network. Equation 1-3 is based on the following assumptions: i) an ideal network deforms in an affine manner with constant molar volume; ii) all chains in the network equally contribute to elasticity of the elastomer; iii) the end-to-end distance of the polymer segments in the network follows Gaussian distribution; and iv) 34 the internal energy is constant. Comparing the experimental stress-strain curve with the predicted curve (Figure 1-10B),213-214 Equation 1-3 shows the best fit at low strain (λ<1.5), where the deformation is small, and the end-to end distance still follows Gaussian distribution. However, at a large strain or in a real elastomer network, the experimental data deviates from the theoretical curve, as the affine deformation and the Gaussian distribution are no longer valid to describe the elasticity of polymer networks. At an even higher strain, the extensibility of the polymer network is limited by the finite extension of polymers or possible strain-induced crystallinity, leading to a fast increase in the force. 215 To better describe the polymer elasticity, several polymer chain models were put forward.216-217 For example, Kuhn and Grün elucidated the stretch of finite extensibility of freely jointed chain using the Langevin distribution.218 The freely jointed chain model is used to describes flexible polymers. Each monomer is considered as a rigid rod with a fixed length (b), and there are no correlation or interactions among monomers. Over the full range of possible extensions, the retractive force (f) and polymer extension can be described as: Equation 1-4 D = EF@G HIB(KLMNOG ) Equation 1-5 ℒ(P) = KLMNQR = coth P − 1P Equation 1-6 P = ℒIX(KLMNQR ) 35 where ℒIX is the inverse Langevin function (Equation 1-5), β is defined as Equation 1-6,216 kB is the Boltzmann constant, T is absolute temperature, "#$% represents the end-to-end distance of a freely jointed chain, and N is the number of the monomers. Equation 1-4 can be employed to account for the deviation of stress-strain behavior at high strains, where the non-Gaussian nature of the chains stretch must be taken into consideration. In addition, the worm-like chain model is widely used to describe semi-flexible polymers, such as double-stranded DNA. It can also be used to describe unstructured RNA and proteins. The elastic response of a polymer from extension can be expressed as: Equation 1-7 ' = YZ2[ × ]14 61 − _`a9Ib + _`a − 14c where p is persistence length, and Lc is the contour length of the polymer. The force (f) linearly increases with the end-to-end distance of the worm-like chain (γ) at a low strain, obeying Hooke’s law. Owning to the limit of chain extensibility, the polymer chain reaches its contour length at an infinite force. Both the freely jointed chain and the worm-like chain models can be used to describe different elasticity of real polymers.216 Additionally, Equation 1-3 has other limitations, such as types of crosslinks and solvent quality. Specifically, only permanent crosslinks are applicable to the stress-strain behavior using Equation 1-3. On the other hand, transient crosslinks are one of the causes that the experimental curve derivates from Equation 1-3 at intermediate strains (Figure 1-10). These transient crosslinks are correlated with the energy dissipation and viscoelasticity properties of the network in response to deformation stimuli. Therefore, Equation 1-3 is generally applied for covalently linked networks, instead of elastomers made of reversible physical crosslinks or chain entanglements, such as coiled-coil protein-based hydrogels and DNA Olympic hydrogels.97, 103, 219 36 Moreover, the presence of solvent in the networks may lead to deviation from Equation 1-3. Mixing of polymer and solvent directly leads to an increase in entropy, favoring network swelling. However, there is also a reduction in the number of chain configurations upon swelling, resulting in a negative entropy change and unfavorable swelling. Besides, solvent quality critically effects on intermolecular interactions in the networks. With a good solvent, interactions between polymer segments and solvent molecules are thermodynamically favorable, and polymer segments are well separated along the chain. With a poor solvent, polymer-polymer interactions are preferred, leading to backfolding and contraction of polymer coils. In other words, solvents critically effect on the distribution of the end-to-end distance even for the same elastomer network, limiting the use of Equation 1-3. Succinctly, the network stress-strain equation of state is not universal. Many limitations have to be taken into account. Further refinements or adaptations are required for real elastic networks. 1.2.2 Methods to improve mechanical properties of protein-based hydrogels Hydrogels resemble natural rubbers in their network structures and elastic behavior. Due to the high water content, hydrogels commonly are mechanically weaker than classic polymer-based elastomers. The well-studied rubber elasticity theory provides mathematical guidance for designing the network structure required to achieve the desired elastic modulus of hydrogels. Based on Equation 1-3, the number of active network chain segments per unit volume (n) indicates the crosslinking density of elastomer networks. The crosslinking density directly determines the macroscopic mechanical performance. Many strategies have been validated to reinforce mechanical properties of hydrogels by increasing crosslinking density of their networks.220-221 In 37 this section, emerging methods and current challenges in the mechanical improvement of protein-/peptide-based hydrogels will be discussed by comparing with polymer-based hydrogels. 1.2.2.1 Increase of polymer concentration The modulus of polymer-based hydrogels is determined by the crosslinking density of the networks. The crosslinking density in turn is closely correlated with the polymer concentration. Hence, the most direct and effective method to enhance the degree of crosslinking in a network is by increasing the concentration of polymer chains per unit volume, as more linkable groups in the pregel come into close contact for the crosslinking reaction, as shown in Figure 1-11A. Shi and Mao developed a host-guest polymer-based hydrogel system. It was observed that increasing the overall host-guest polymer concentration from 0.07 to 0.14 mmol/mL resulted in a 100-fold increase in the storage modulus up to 141 kPa.222 Similar results have been found in many physically or chemically crosslinked polymer-based hydrogels, such as polyacrylamide hydrogels, Gellan Gum hydrogels, agar gels and alginate scaffolds.223-227 The mechanical properties of protein-/peptide-based hydrogels were studied via a similar strategy. The Tirrell group engineered random triblock proteins which assembled into hydrogels through aggregation of leucine zipper domains.108 The storage modulus of the hydrogels showed a protein concentration dependence. Increasing the protein concentration from 5 to 9 % (w/v) resulted in an increase in the storage modulus from 150 to 1400 Pa. In addition to disordered peptide-based hydrogels, the mechanical properties of structural protein-based hydrogels can also be regulated by varying the concentration. Fibrinogen is the essential fibrous protein related to blood clotting. The self-assembly of fibrin fibers promotes physical gelation. Stegemann, Weisel and Nagaswami found that the fibrin gel structure was influenced by the concentration of 38 fibrinogen, calcium and thrombin (the enzyme that catalyzes the microfibril formation).228-229 As the fibrinogen concentration was increased, tightened networks were formed with higher material modulus. It seems that facile adjustments to the protein-/peptide concentration is able to directly control material properties. However, it should be noted that highly concentrated proteins may become problematic for hydrogel preparation due to high viscosity and possible aggregation. Also, structural proteins are varied in size and shape. It is unclear whether increasing protein concentration is a universal solution for mechanical strengthening of all gel categories. Figure 1-11 Schematics of hydrogel mechanical enhancement approaches. A) An enhanced gel network with concentrated polymers. B) Chemical functionalization bringing in additional crosslinks in a gel network (the yellow dots). C) An entangled network formed by polymers with high molecular weight. D) A representative inter-crosslinked double network. E) A crystallite reinforced network. The green areas represent crystalline regions. F) A nanoclay composite hydrogel network. The blue ellipses represent nanoclay particles. 1.2.2.2 Network Functionalization with cross-linkable groups Given low yields and high cost of protein purification, an alternative method is to introduce cross-linking groups into the gelation system, instead of simply increasing the gelation 39 concentration. In polymer-based hydrogels, enhancement of mechanics can be achieved by additional crosslinkers (Figure 1-11B). For example, the Engler group designed polyacrylamide hydrogel substrates with tunable mechanical properties by adjusting the relative ratio of acrylamide and bis-acrylamide.227 Polyacrylamide hydrogels are prepared by mixing acrylamide monomers with bis-acrylamide, where bis-acrylamide acts as crosslinkers. Their results indicated that increasing the proportion of crosslinker from 0.03-0.3% resulted in a stiffened elastic modulus of 3 kPa to 35 kPa at a fixed 10% acrylamide concentration. Similar to polymer gelation which uses small crosslinker molecules, natural soy protein hydrogels were fabricated with the addition of linear PEG. Both ends of the carbonate-activated PEG served as crosslinkers by releasing p-nitrophenol.230 It was found that the soy protein hydrogels exhibited improved elastic modulus with more PEG molecules conjugating in the elastic network. It is noted that the stiffened gels also showed lower swelling ratio and relatively lower maximum deformation strain. Additionally, Elvin and coworkers reported a lysine derivatization method to enhance the mechanical properties of gelatin hydrogels by using Bolton-Hunter reagent.164 The lysine residues of gelatin were derivatized with phenolic (tyrosine-like) functional groups, where the phenol groups could form covalent crosslinks by employing a ruthenium complex-catalyzed photochemical reaction. The hydrogels made of derivatized gelatin showed enhanced elastic modulus. The lysine derivatization method has also been applied to improve the mechanical properties of globular polyprotein (GB1)8-based hydrogels by forming additional dityrosine-like adducts.231 The degree of crosslinking of hydrogels could be further adjusted by controlling the phenolic content of (GB1)8, leading to a tunable elastic modulus ranging from 13 to 65 kPa. 40 Besides adding chemical crosslinkers, engineering recombinant polyproteins that contain diverse assembling domains or physical coordination sites is another successful means to achieve extraordinary mechanical properties in hydrogels. Gonzalez et al. engineered tough and stretchable hydrogels by using intrinsically unstructured ELP fusion proteins.232 Incorporation of other non-ELP motifs allows the protein to maintain the thermal responsiveness of ELP as well as additional biofunctions. In this study, a zinc-binding motif found in matric metalloproteinases was used to introduce metal reversible coordination into the networks.233 Additionally, a tripeptide (GKG) was inserted so that the lysine can serve as a covalent crosslinking site. These two different types of crosslinking protein domains were genetically encoded and interspersed with periodic repeats of ELP domains respectively, and became the metal crosslinkable protein (MCP) and the covalently crosslinkable protein (CCP) ( Figure 1-12). The protein mixture of CCPs and MCPs were covalently crosslinked into a hydrogel in the presence of the chemical crosslinker tetrakis(hydroxymethyl)phosphonium chloride (THPC). The MCPs in the network were subsequently crosslinked via metal coordination by immersing the gel in a zinc solution. These features and the strong hydrophobic interactions of ELPs yielded hydrogels that were shown high strength high strength (>2.5 MPa), toughness (>1300 Jm-2) and stretchability (>500%). Distinct from the unfolding of globular proteins, the mechanical energy dissipation of such hydrogels was attributed to reversible metal coordination crosslinks and the hydrophobic interactions of ELPs, while the CCP networks provide elasticity. Non-covalent crosslinks mechanically strengthen elastomers in a dynamic manner. Reversible binding/unbinding provides the elastomers with excellent extensibility and load-bearing properties. However, physical crosslinks that are too dynamic may cause instability and mechanical weakening of the hydrogels. Unlike the physical crosslinking reinforcement, chemical 41 decoration with crosslinking groups enhances the mechanics of hydrogels effectively, but it typically brings on other changes in material properties, such as faster gelling rate, lower swelling ratio and smaller porosity. Furthermore, the high density of covalent crosslinking sites spread across short polymer segments constrains intermolecular movement, which largely reduces the stretchability and energy dissipative ability of the elastomer network. Thus, an excess of crosslinking sites may lead to stiff but brittle hydrogels. Figure 1-12 Schematics of ELP-CCP-MCP tough hydrogels. A) Intrinsically unstructured protein building blocks: ELP, metal coordinating domain (MCP 1 and 2), and covalently cross-linkable domain (CCP). B) Schematics of the elastic network formed by covalent crosslinking and the dissipating network resulting from metal coordination in the presence of zinc ions. The tough protein hydrogel forms based on interpenetration of both networks. C) Upon stretching and relaxation, the hydrophobic interaction of ELPs can dissociate and reform reversibly. The metal coordination breaks within the MCP network under further stretching providing the energy dissipative property, while the covalently crosslinked network remains elastic throughout.232 Copyright 2017, WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim. 1.2.2.3 Increase of the degree of polymerization From the perspective of polymer physics, both the length of building blocks and the overall molecular weight of a polymer impact on the mechanics of polymeric materials. Linear polymers 42 with high molecular weight form entanglements, acting as physical crosslinking sites to reinforce the elastomer network in (Figure 1-11C). Moreover, in these cases, the number of inactive free ends is relatively small, which can be considered as a reduction of network defects, resulting in improved modulus and extensibility.234 For example, Killion et al. investigated the mechanics of a polymer hydrogel made of poly(ethylene glycol) dimethacrylate (PEGDMA) by manipulating the molecular weight of the polymer.235 The use of 1000 Da PEGDMA compared to 600, 400, and 200 Da PEGDMA resulted in stronger hydrogels, but all gels showed a similar swelling ratio. Similar results can be found in peptide-/protein-based hydrogels. The Chilkoti group studied the swelling and the mechanical behavior of ELP hydrogels formed using ELPs of different molecular weights.81 The genetic constructs VPGKG(VPGVG)n were synthesized with molecular weights ranging from 3.1 kDa to 93.4 kDa. Rheological studies indicated improved stiffness of the ELP gels (from 1.6 to 15 kPa at 37 ℃) with increasing numbers of ELP repeating units. On the other hand, the gels made of low molecular weight ELP showed higher swelling ratios than the high molecular weight ELP gels. In addition to ELP-based hydrogels, Pena-Francesch engineered tandem repeats of the disordered squid ring teeth protein, demonstrating that the tensile modulus (from 1 to 2.5 MPa) increased with the number of repeating building blocks (from 4 to 25).236 Inspired by the strength of muscle tissues, a representative protein-based polymer is the giant muscle protein titin, containing more than 100 tandem modular globular domains with a molecular weight of more than 3.5 MDa.237-239 Wang et al. reported an efficient protein fragment reconstitution strategy based on two split fragments GN and GC for globular protein polymerization. GN and GC are derived from GL5CC, a variant of GB1 where residues 42/44 were mutated to cysteines.240 Under oxidizing conditions, GN and GC form disulfide bonds and reconstitute into a native folded GB1 with covalent connection. Starting from protein macromonomers GN-I27-GC, 43 they were able to engineer protein polymers with molecular weights up to ~0.5 MDa. Rheology measurements indicated that the gelation process was complete in less than two minutes via a Ru(II)-mediated photocrosslinking strategy. The storage moduli were 0.6 kPa, 2 kPa and 11 kPa for 1%, 2% and 5% (w/v) hydrogels, respectively. Additionally, this ultra-high molecular weight polyprotein allowed them to fabricate hydrogels at a protein concentration as low as 1%, in contrast to other globular polyproteins where the critical gelling concentration must be beyond 5%. Such a protein fragment reconstitution strategy for protein polymerization provided a new approach for protein-based biomaterials and functional assemblies with improved physical and mechanical properties. However, there are also challenges in using high molecular weight proteins/peptides for hydrogel preparation. Firstly, large efforts in protein engineering are required to construct such long repetitive sequences. Secondly, the pre-gel solutions of protein/peptide with high molecular weight normally are viscous due to the entanglements of free long chains, which makes transferring and handling difficult. Thirdly, for self-assembling proteins (such as β-sheet forming domains or coiled-coil motifs), unwanted aggregation and precipitation may occur during protein purification or in pregel solutions. 1.2.2.4 Double networks A double network (DN) hydrogel is comprised of two heterogeneous polymer networks, which are interpenetrating or covalently crosslinked together (Figure 1-11D). The Gong group first developed an innovative approach to achieve material having both high water content as well as excellent mechanical performance. They suggested that a successful DN hydrogel normally consists of a high relative molecular weight neutral polymer network incorporated with a swollen 44 polyelectrolyte network, where the interpenetration of two polymers largely increases the network density and results in them interlocking with each other. The resultant hydrogels with the optimal mixing ratio ameliorated mechanical properties, compared to each individual component.241 Among all the polymer pairs they studied so far, the DN gel made of poly (2-acrylamido, 2-methyl, 1-propanesulfonic acid) (PAMPS) polyelectrolyte and polyacrylamide (PAAm) neutral polymer showed outstanding properties with a nominal stress up to 17.2 MPa and vertical compressive strain of ~92%, while the individual PAMPS gel and PAAm gel broke at stress of 0.4 MPa (41%) and 0.8 MPa (84%), respectively.242 In addition to the synthetic polymer-based hydrogel, the Gong group reported another DN hydrogel made of bacterial-cellulose/gelatin showing extremely high mechanical strength.243 Bacterial cellulose consists of a hydrophobic stratified fibrous structure, which has a high tensile modulus (2.9 MPa) along the fiber-layer direction but a low compression modulus (0.007 MPa) perpendicular to the stratified layers. Bacterial cellulose hydrogels show an irreversible shrinking behavior upon compression. In contrast, gelatin gels are mechanically brittle but their networks retain water and recover from repeated compression. These two mechanically contrasting biopolymers were chemically incorporated and crosslinked into a double network. The resultant hydrogel overcame drawbacks from both components, displaying an impressive compression modulus (1.7 MPa) and reversible swelling. In summary, all successful DN hydrogels are formed by choosing suitable polymer pairs and developing an incorporation method of the two heterogenous networks. Additionally, the ratio of each network component in the DN significantly affects the final gel mechanics. Tuning the ratio is an inevitable step to obtaining optimal gels. 45 1.2.2.5 Mechanical reinforcement by forming nanostructures Incorporating nanostructures is an alternative approach for producing tougher hydrogels. Such reinforcements can be found in natural protein-based materials, such as collagen matrices and silk protein fibers. As discussed in the Section 1.1.3, the gelation and strength of collagen hydrogels is related to the formation of nanofibers/collagen fibrils (Figure 1-11E). In contrast, β-sheet crystallites give rise to the exceptional mechanical strength and extensibility of silk. This is analogous to the combination of amorphous and crystalline regions found in high density polyethylene (HDPE) and polyurethane elastomers, where the degree of crystallinity largely impacts on polymer stiffening and stabilization.244-246 Yuan et al. demonstrated a gel-to-crystal phase transition of a β-sheet dipeptide (diphenylalanine) system. Characterization of the morphology of this system showed that the microstructure of molecular assembly changed from ultrafine fiber to wider and longer nanofibers during the phase transition, leading to a ten-fold increase in storage modulus.247 The Zhao lab developed gel strengthening approaches by mimicking muscle training. The aligned nanofibrillar architectures and nanocrystalline domains of polyvinyl alcohol hydrogels could be achieved by mechanical training, where the gels showed enhanced Young’s modulus and fatigue resistance.248 1.2.2.6 Incorporation of nanocomposites Besides the reinforcement introduced through nanostructures within gel networks, incorporation of nanocomposites into main hydrogel networks is another approach that results in improved mechanical strength with high extensibility due to the high functionality of the exfoliated nanosized fillers. One of the examples is a composite hydrogel made of nanoplatelets (inorganic nanoclay) and radical polymerized cross-linked networks. Due to the charged surface,249 clay 46 nanosheets absorb the polymerization initiator onto their surfaces, leading to polymerization starting from the clay surfaces. Thus, one nanoclay particle links many polymer chains, acting as a crosslinker and enhancing the macroscopic mechanical strength (Figure 1-11F).250-252 In addition to nanoclay composite hydrogels, carbonaceous nanocomposite hydrogels show exceptional mechanical properties as well. Graphene oxide (GO) as a functionalized carbon allotrope has been incorporated into silk-fibroin matrices, demonstrating improved mechanical properties.253-257 The Choudhury group designed tough silk fibroin/GO nanocomposite hydrogels by using the ruthenium-based photochemical crosslinking method.258 GO is capable of forming interactions with silk fibroin domains via hydrophobic β-sheet assembly, hydrophilic interactions, π-π stacking and potential electrostatic interactions. After irradiation with visible light for 2 mins, the silk fibroin/GO mixture formed an adhesive and soft hydrogel, while the control gel (silk fibroin only) was observed to be non-adhesive. However, further exposure to the light resulted in a tacky hydrogel that underwent a physicochemical transformation due to increases in crosslinking density, becoming tough and non-adhesive, with a tensile strength as high as ~8 MPa. Nanocomposite hydrogels exhibit excellent toughness and decent swelling ability, but the gels typically involve complex preparation procedures, as nanocomponents tend to aggregate, causing by uneven dispersity or a high local concentration. 1.3 Goals of the thesis Globular proteins have attracted considerable interest as hydrogel building blocks, owing to their diverse biofunctions, genetically encodable sequence, and dynamic folding-unfolding. However, how the bulky sizes of folded globular domains impact the elastic networks remains poorly understood. The nature of the networks in turn also affects the physical and mechanical 47 performance of the hydrogels. Moreover, the toughness and the strength of current protein-/peptide-based hydrogels are inferior to polymer elastomers, critically limiting biomedical applications. Compared to polymer hydrogels, globular proteins impact differently on mechanical and physical properties of the hydrogels, as these bulky domains have shorter effective length, larger chain rigidity, and greater steric hindrance. Despite these differences, unfolded polyproteins structurally resemble synthetic polymers, bridging coiled polymers and structural polyproteins. Based on the similarity, this thesis aims at exploring the contribution of structural domains to the elastomer networks, and developing methods to precisely control the mechanical properties of globular polyprotein-based hydrogels. We also aim to use mechanical reinforcement of globular protein-based hydrogels. In Chapter 2, the goal is to develop a quantitative tool, allowing us to predict the mechanical properties of protein hydrogels. Characterizing the mechanical properties of synthetic polymer-based hydrogels has been largely based on the classic theory of rubber elasticity. However, the theory cannot be directly applied to structural protein hydrogel networks. Thus, photochemically dityrosine crosslinked protein hydrogels are engineered by utilizing tandem modular elastomeric proteins (GB1-Resilin)n. At the unfolding transition in the hydrogel networks, achieved by use of chemical denaturants, it is proposed that unfolded polyproteins structurally resemble synthetic polymer networks, where the rubber elasticity theory can be used. Based on the Young’s moduli measured from tensile tests on the native and denatured states of the protein hydrogels, it is possible to establish a quantitative connection between the native and denatured protein networks, as the chemical crosslinking sites remain intact during the folding-unfolding 48 transitions. This approach may help tailor mechanical properties of protein hydrogels in a rational and predictable fashion. Inspired by the conclusions of the study in Chapter 2, cyclic conformational changes of proteins (folding-unfolding) in the gel network leads to elongation and contraction of the biomaterials. Such mechanochemical cycles of hydrogels lead to mechanical work generation due to protein folding in the protein hydrogels, behaving as a soft actuator. Similarly, mechanochemical cycles of protein domains and mechanical work generated by protein folding can be investigated at a single molecule level by using optical tweezers. The goal of Chapter 3 is to develop a hydrogel actuator based on the mechanical energy output of protein folding as a molecular motor. Furthermore, it is proposed that there are close correlations between single-molecule and macroscopic mechanochemical cycles and mechanochemical actuation. The goal of Chapter 4 and Chapter 5 is to design mechanically reinforced hydrogels using globular polyproteins. Guided by the results in Chapter 2, tough elastomers generally have both high crosslinking density and elastomeric regions in the networks. The former leads to good material strength and stiffness, while the latter contributes to decent extensibility. However, folded proteins may be subject to steric hindrance, leading to inefficient photochemical crosslinking. Thus, crosslinking denatured proteins and then renaturing them is an alternative approach to improve the crosslinking density. In addition, due to the long chains of unfolded proteins, polypeptide chains can cross each other, resulting in entanglements. Chain entanglements can effectively increase network density, as they act as physical crosslinks. Further renaturation of the crosslinked network leads to a contraction of the hydrogel due to protein refolding and hydrophobic collapse. The detailed content is illustrated in Chapter 4. 49 Another way to enhance the mechanical properties of protein hydrogels is to adopt a double network strategy. In Chapter 5, alginate is chosen as the secondary network due to its ionically cross-linkable chain blocks. More importantly, the carboxyl groups located on the polymer backbone can be modified into phenol moieties.259 Alginate with phenol moieties can covalently conjugate with primary protein networks by forming phenol-tyrosine adducts, leading to tough double network hydrogels with enhanced network density. Taking advantage of precise mechanical control and the dynamic behaviors of mutually exclusive protein-based hydrogels, Chapter 6 shows an application of protein-based hydrogels as an ECM for cell spreading studies and mechanochemical studies. We demonstrate that human lung fibroblasts dynamically change their morphology in a fully reversible fashion in response to a cyclic change of the hydrogel stiffness. 50 Chapter 2: Towards Quantitative Prediction of Mechanical Properties of Tandem Modular Elastomeric Protein-based Hydrogels 2.1 Synopsis Elastomeric proteins made of tandemly arranged individually folded globular domains have been used as building blocks to construct protein hydrogels with tailored mechanical properties. Characterizing the mechanical properties of synthetic polymer-based hydrogels have been largely based on the classic theory of rubber elasticity. However, classical rubber elasticity theory cannot be directly used to describe the mechanical properties of such protein hydrogels, due to the fact that only part of the structure of such folded globular domains contributes to the elastic properties of the hydrogel network. Here we present a length equivalent molecular weight approach to adapt the classic rubber elasticity theory to such elastomeric protein-based hydrogels, fulfilling the practical purpose of predicting the mechanical properties of protein hydrogels semi-quantitatively. The validity of this approach was tested using different elastomeric protein-based hydrogels. Our results not only allowed us to quantitatively predict/explain the mechanical properties of such elastomeric protein-based hydrogels upon folding-unfolding transitions in the hydrogel network, but may also help tailor mechanical properties of protein hydrogels in a rational fashion for various applications. *A version of Chapter 2 has been published as [Linglan Fu] and Hongbin Li. (2020) Toward Quantitative Prediction of the Mechanical Properties of Tandem Modular Elastomeric Protein-Based Hydrogels. Macromolecules. 53 (12): 4704-4710. This chapter is adapted with permission of American Chemical Society, Copyright 2020. 51 2.2 Introduction Due to their potential biomedical applications, recombinant protein-based hydrogels have attracted considerable interest over the last two decades.260-263 Compared to synthetic polymers, recombinant proteins as building blocks for the construction of hydrogels have many unique features and advantages, such as their well-defined and tunable primary sequence, unique secondary and tertiary structure, and monodispersed molecular weight. These features make it possible to tune the mechanical and functional properties of such protein hydrogels by programming the protein sequence. Intrinsically disordered elastomeric proteins, such as elastin and resilin, have been widely used in the construction of protein hydrogels.81, 264-266 Due to the random coil structure of such elastomeric proteins, the elastic properties of such protein hydrogels can be readily explained and predicted by the classical rubber elasticity theory, assuming that the random coil sequences can be described by Gaussian chain statistics: 207, 212, 267-270 Equation 2-1 = = d)A − BAC- = e?@fg )A − BAC- where σ is the stress, h = ij.k" is the shear modulus, ρ is the density of the polymer in the network and is equal to the dried mass divided by hydrogel volume, lm is the molecular weight between two neighboring crosslinking points, R is the gas constant and λ is the extension ratio. The Young’s modulus E measured from tensile testing is defined as n = 7o7p . When λ=1, n = 3h . This relationship is derived based on the assumption that the Poisson’s ratio of the network is 0.5. Considering loops and free ends of the network strands,271-274 52 Equation 2-2 h = rs2lm t1 − 2lml v where M is the number average molecular weight of the polymer, and the second term removes the free ends that do not contribute to the elasticity of the network. It is of note that the number of crosslinking points per chain Nc=M/lm. Equation 2-2 has been used successfully in describing the elastic properties of a range of intrinsically disordered protein-based hydrogels. 212, 267, 270, 275 Recently, we and others have started to employ tandem modular protein-based elastomeric proteins as building blocks to construct protein hydrogels.29, 33, 42, 69, 276-277 Elastomeric proteins made of individually folded globular domains interspersed with random coil sequences are present in a wide range of biological systems, where they serve as elastic connections under biological settings to bear and transmit mechanical forces, and determine the mechanical properties of the constituting tissues.237, 278-279 The giant muscle protein titin is the best-known example of such elastomeric proteins.237 One of the unique characteristics of such elastomeric proteins is that the folded globular domains can unfold under a mechanical stretching force to effectively dissipate energy, making such elastomeric proteins effective shock absorbers.280-282 Artificial elastomeric proteins made of folded globular domains have been employed as building blocks to engineer novel protein hydrogels. GB1 (G) and Resilin (R) proteins were used as building blocks to develop muscle-mimetic soft materials in our previous work. GB1, the B1 IgG binding domain of streptococcal protein G, is a small folded globular domain with a four-stranded β-sheet packed against an α-helix (Figure 2-1A). Its mechanical properties have been well-characterized at the single molecule level.29, 283 Resilin is an intrinsically disordered elastomeric protein found in the cuticle of most insects,275 and has been used to construct protein elastomers. Here we used its 15 amino acid (aa) residue-long consensus sequence as the building block.275, 284 29 GR-based 53 hydrogels showed comparable mechanical properties to muscle tissues with excellent elasticity, fast hysteresis recovery and high resilience at low strain.29 The rational tuning of their mechanical properties has been realized by tailoring the mechanical stability of the globular domains and the strength of the crosslinking points.29, 32-34, 231, 285-287 By controlling the folded-unfolded states of the globular domains via folding switches, it has become possible to engineer dynamical protein hydrogels whose mechanical properties can be dynamically tuned in response to environmental stimuli.43, 288-290 These novel protein hydrogels have opened up new opportunities to design and engineer biomaterials for potential applications in mechanobiology and biomedical engineering, and led to the efforts to develop models to explain their mechanical properties.32-33, 291-292 Despite these advances, it is challenging to use the classical rubber elasticity theory to quantitatively predict the mechanical properties of these hydrogels, as only part of a folded globular domain is subject to the stretching force and contributes to the network elasticity. As such, the folded globular structure of the protein domains effectively changes the relationship between the length of the protein polymer and its molecular weight. Applying Equation 2-2 to tandem modular protein-based hydrogels would yield an unphysical Mc, leading to a Nc that is greater than that of the available crosslinking sites per protein polymer chain. Moreover, folded globular domains have a persistence length different from disordered polypeptide chains. To properly address these challenges, it is necessary to treat the tandem modular elastomeric protein as a polymer chain of different stiffness and develop an analytical or numerical solution, as those developed for biological gels made of filamentous proteins, such as actin.293-294 However, such a solution is yet to be achieved. 54 From a practical point of view, here we present a simple method to adapt the classical rubber elasticity theory to tandem modular protein-based hydrogels to fulfill the practical need of predicting elastic properties of tandem modular elastomeric protein-based hydrogels quantitatively. Figure 2-1 Schematics of the length-equivalent molecular weight approach. A) Schematics of the GR repeating unit. GB1 is a 56 aa long, folded globular domain (PDB code 1PGA), and resilin is a disordered polypeptide. In GB1, only the two terminal β-strands are experiencing the stretching force in an elastic network. The distance between its N and C-termini Ld is 2.6 nm. Thus, the folded GB1 domain can be approximated as a polypeptide chain of 7.2 aa (Ld/La, La is the length of a single aa residue) with a length equivalent molecule weight Me of 790 Da. The GR repeating unit has a Me of 2882 Da. B) Schematics of the tandem modular protein-based hydrogel network. #$%%%% is the network strand molecular weight. 2.3 Materials and methods 2.3.1 Protein engineering We constructed the following polyproteins for our experiments: (GR)4, (GR)8, (GR)12, (GR)12-SpyTag, (GR)12-SpyCatcher, (GR)16 and (GB1)8. All polyproteins were constructed using 55 standard molecular biology methods. The genes of (GR)4, (GB1)8 and (FL)8 were constructed as described.29, 34, 276 We modified pQE80L and pET29b to obtain pQEcc and pET29bMod through alteration of the original multiple cloning sites with a synthetic oligonucleotide sequence (important parts of cloning sites maps of the plasmids and the construction steps are shown in Figure 2-2). The plasmid pQE80L-(GR)4 was digested by BamHI and KpnI to create the insert (GR)4, and the insert coupled with vector pQE-(GB)4 digested by BglII and KpnI to form pQE80L-(GR)8. The vector pQE80L-(GR)8 digested by BglII and KpnI from the pQE80L-(GR)8 was ligated with the insert (GR)4 prepared above, obtaining pQE80L-(GR)12, and used in a similar manner to create pQE80L-(GR)16 from the plasmid pQE80L-(GR)12 and the insert (GR)4. However, truncation of GR polyprotein was found in protein expressed from the original pQE80L vector carrying an N-terminal His-tag. To isolate only the fully transcribed proteins, we transferred the target sequence into a pET29bMod vector, which carries a C-terminal His-tag. First, the stop codon before KpnI in the gene of our constructs was removed by the following procedure: the (GR)4 insert was isolated by digesting pQE80L-(GR)4 with EcoRI and BglII, and then ligated into pQEcc vector which was also cut by EcoRI and BglII, yielding pQEcc-(GR)4; then the (GR)4 insert was prepared by BamHI and KpnI digestion, and cloned into vector pET29b (modified) digested with BamHI and KpnI to form pET29bMod-(GR)4. The plasmids pET29bMod-(GR)8, pET29bMod-(GR)12, pET29bMod-(GR)12-SpyTag, pET29bMod-(GR)12-SpyCatcher and pET29bMod-(GR)16 were all constructed using similar steps. The plasmids pET29bMod-(GR)4, pET29bMod-(GR)8, pET29bMod-(GR)12, pET29bMod-(GR)12-SpyTag, pET29bMod-(GR)12-SpyCatcher and pET29bMod-(GR)16 were transformed into E. coli strain BL21(DE3) competent cells. Seeding culture was allowed to grow overnight in 10 mL 2.5% Luria-Bertani broth (LB) medium containing 50 mg/L kanamycin. The 56 overnight culture grew in 1L sterile LB medium at 225 rpm and 37℃ for another 3 hours, and protein expression was induced with 1mM isopropyl-1-β-D-thiogalactoside (IPTG) at 37℃ for 4 hours. The cells were harvested by centrifugation at 4000 rpm for 10 mins at 4℃, and then frozen at -80℃. At a later point in time, cells were thawed and lysed by incubation with 1mg/mL lysozyme for 30 mins. Nucleic acids were removed by adding 0.1 mg/mL DNase and RNase. The supernatant containing soluble protein was collected after centrifuging the cell mixture at 12000 rpm for 60 mins. The soluble protein was purified by Co2+ affinity column. Yields of (GR)4, (GR)8, (GR)12-SpyTag, (GR)12-SpyCatcher, (GR)12, (GR)16 were approximately 40 mg/L of bacterial culture. Purified proteins were dialyzed against deionized water for 2 days to remove NaCl, imidazole, and phosphate. Then the protein solution was filtered and lyophilized. To engineer (GR)24, (GR)12-SpyTag was reacted with (GR)12-SpyCatcher at a 1:1 molar ratio in phosphate buffered saline (PBS). SpyTag can bind and react with SpyCatcher to form a covalent isopeptide bond linkage. The use of SpyTag and SpyCatcher chemistry allowed us construct (GR)24 much more efficiently, as bacterial expression of (GR)24 resulted in a large amount of truncated product and low yield. The constituted SpyTag-SpyCatcher complex contains one exposed tyrosine residue, which can participate in dityrosine crosslinking. Figure 2-3 shows the SDS-PAGE of the engineered polyproteins. The solubility of the purified polyproteins (GR)n is at least 300 mg/mL in PBS buffer at pH 7.4. At 300 mg/mL, no precipitation or aggregation was observed in the (GR)n solutions. This high solubility is consistent with the excellent solubility of GB1, for which GB1 is often used as a fusion protein to help increase the protein expression yield for other proteins.295 57 Figure 2-2 Engineering the gene of desired polyprotein through multiple steps. pQE80L-(GR)n constructs are from our previous work carrying an N-terminus His-tag and a C-terminus stop codon.29 To obtain polyprotein with monodispersity, the desired polyprotein genes were transferred into pQEcc to remove the C-terminus stop codon. Then the constructs were ligated into pET29b Modified with a C-terminus His-tag. 58 2.3.2 Hydrogel preparation and dimension measurements Gelation of the polyprotein was based on the photochemical crosslinking strategy as described.29, 275, 296 To prepare a ring-shaped, 20% hydrogel for tensile tests, 20 mg of lyophilized protein was redissolved in 93.7 μL of PBS, with 5 μL of ammonium persulfate (APS, 1M) and 1.3 μL of tris(bipyridine)ruthenium(II) chloride (20 mM) with final concentrations of 50 mM and 260 μM, respectively. The protein solution was quickly transferred to a custom-made plexiglass ring-shaped mold (din=8 mm, dout=10 mm, h= 3 mm), and was exposed under a 200W fiber-optic white light source from a height of 10 cm above the mold for 10 min. The ring sample was carefully taken out of the mold, and the initial weight of the freshly prepared sample was measured before swelling. Native gel rings were stored in filtered PBS to achieve equilibrium swelling. To prepare denatured samples, the native ring samples stored in PBS were transferred and immersed in 7 M guanidine hydrochloride in PBS for over 3 hrs to allow complete denaturation and to achieve swelling equilibrium. The outer diameter, thickness, width and weight of all ring-shaped equilibrated swollen samples were measured before tensile tests. To measure the weight of the dry gel, the samples were lyophilized at -80℃ for 2 days. The hydrogel swelling ratio (r) was determined by the difference in mass between the freshly prepared gel (mi) and the equilibrated gel (ms) in native or denatured states, Equation 2-3 " = $!%$"$" × 100%. 59 Figure 2-3 Monodispersed molecular weight of GB1-Resilin (GR)n tandem modular polyproteins. A) 10% SDS-PAGE gel of the purified (GR)4, (GR)12, (GR)16 elution fraction from Co2+ affinity column expressed using the C-terminal His-tag plasmid. B) 10% SDS-PAGE gel of (GR)12-SpyTag-SpyCatcher-(GR)12 (1:1 molar ratio mix), purified (GR)12-SpyTag, (GR)12-SpyCatcher expressed using the C-terminal His-tag plasmid. The size of a GR-unit is ~8.3 kDa, Sc is 12.3 kDa, and St is 1.6 kDa. 2.3.3 Tensile tests Tensile tests were used to characterize the mechanical properties of hydrogels following well-established protocols for soft hydrogel samples.297 An Instron-5500R tensometer with a custom-made force gauge was used to perform tensile tests. Unless stated otherwise, native and denatured ring samples were tested in PBS and 7 M guanidine-HCl (GdHCl) in PBS respectively at an extension rate of 25 mm/min at a constant temperature of 25℃. Fully equilibrated ring-shaped hydrogel samples were used for tensile testing and were always immersed in the corresponding testing buffer throughout the testing (Figure 2-4). The test specimen was hung between a hook (radius 1.5 mm) and a rod (radius 1.5 mm). The hook is connected to a 5N force transducer, which is fixed on the top of the machine frame. The force transducer is driven to detect the force applied on the sample during the stretching and relaxation cycles. The rod can be moved to put the ring-60 shaped sample on, and it was set on the platform that supported the buffer container. The platform is connected to the strain gauge, which can accurately move vertically at constant speed to the setting distance, while the buffer container and the rod move simultaneously. Young’s modulus is derived from the stress-strain ratio from each stretching-relaxation curve at 0-5% strain. Figure 2-4 Schematics of an Instron-5500R tensometer. A ring-shaped hydrogel sample is hung on a hook and a rod, and is immersed in a testing buffer. The hook is connected with a force transducer, and the rod is connected on a strain platform controlled by strain gauge. 2.4 Results and discussion 2.4.1 Length-equivalent molecular weight for folded globular domains At a constant temperature, the shear modulus G of the polymer network is mainly affected by the number of network strands per unit volume (+/-.////), where -.//// is the number average molecular weight of a network strand.29 To adapt the classical rubber elasticity theory to a tandem modular protein-based network, we used the (GR)12-based protein hydrogel as a model system to illustrate the concept of length equivalent molecular weight for folded globular domains (Me). (GR)12 is an engineered artificial tandem modular elastomeric protein that is composed of an alternating folded globular domain GB1 (G) and the random coil sequence resilin (R). (GR)12 61 hydrogel was constructed using the well-developed [Ru(bpy)3]2+-mediated photochemical crosslinking reaction,296, 298 where two tyrosine residues in proximity can be crosslinked into a dityrosine adduct that serves as a chemical crosslinking point. Both GB1 and resilin contain tyrosine residues that can be crosslinked. Figure 2-1 shows the schematics of the GR repeating unit and the hydrogel network and Figure 2-5A shows the photographs of 20% (GR)12-based protein hydrogels in their native and denatured states. In the hydrogel network, the intrinsically disordered resilin chain in its entirety contributes to the network elasticity. However, due to the folded nature of GB1, only β strands 1 and 4 in GB1 are force-bearing and the rest of the polypeptide chain of GB1 does not contribute to the network elasticity. Thus, the real molecular weight of GB1 cannot be directly used in Equation 2-2. To mitigate this problem, we introduce the “length equivalent molecular weight” (Me) for folded globular domains and the whole elastomeric protein. Since the effective chain length of a GB1 in a network strand is the distance between its N- and C-termini (2.6 nm, measured from its three dimensional structure, PDB code: 1PGA)299, the length equivalent molecular weight Me of GB1 is 794 Da, equivalent to that of a 7.2 aa long disordered polypeptide chain (2.6nm/0.36 nm/aa=7.2 aa, where 0.36 nm is the length of a single amino acid residue300-301). Therefore, the Me for a GR repeating unit can be calculated as follows: Equation 2-4 -1 = 2 34355 + 7558 × -55 where Ld is the effective length of the folded protein, Laa is the average length of an amino acid residue (0.36 nm),300-301 Naa is the number of amino acid residues in the disordered polypeptide chain, and -55///// is the average Mw of an amino acid residue (110 Da). Therefore, the Me for a GR repeating unit is 2882 Da, instead of its real Mw of 8250 Da. And the Me for (GR)12 is 35 kDa 62 versus the real Mw of 99 kDa. With this adaptation, the classical rubber elasticity theory could be used to predict the mechanical properties of (GR)12-based protein hydrogels on a semi-quantitative basis. 2.4.2 Length equivalent molecular weight approach adequately describes the Young’s modulus of (GR)12 hydrogels To validate the length equivalent molecular weight approach, we investigated the tensile properties of (GR)12-based hydrogels to determine the shear modulus G. There are two ways of determining the shear modulus G from stress-strain curves: one is to plot σ versus λ-(1/λ2). According to 9 = : ;< − >?#@ = ABCD$ ;< − >?#@, σ versus λ-(1/λ2) should yield a straight line with a slope equal to G for an ideal elastomer. The second method is to directly calculate G using the relationship G=E/3, where E=∂σ/∂λ when λ=1. In our experiments, E is determined where λ ranges from 1 to 1.05 (corresponding to strain of 0-5%).207, 275 For our native (GR)n hydrogels, the linear region is up to a λ of ~1.1. The two methods yielded values of the shear modulus G that are similar. Figure 2-5B shows the stress-strain curves of 20% (GR)12-based hydrogels in PBS. Hysteresis was observed between the stretching and relaxation curves at large strains, which was largely due to the force-induced unfolding of GB1 domains.29 The hysteresis could be recovered rapidly, and consecutive stretching-relaxation curves were largely superimposable. The Young’s modulus E of the hydrogel is 9.7 kPa. Using Equation 2-2, we measured a length equivalent network strand molecular weight (-.1/////) of 15.3 kDa. This translates into a Nc of 2.3, the number of crosslinking points per protein polymer chain (-1/-.1/////). 63 Figure 2-5 Physical dimensions and mechanical properties native and denatured (GR)12 hydrogel rings. A) Photographs of the 20% (GR)12 hydrogels in native and denatured states. Evidently, the denatured (GR)12 hydrogel swells much more than the native hydrogel. B) Representative tensile stress-strain curves of 20% (GR)12 hydrogel in its native state (in PBS) and denatured state (in 7M GdHCl). The folded GB1 domains can be chemically unfolded by chemical denaturants, such as GdHCl. In the presence of 7 M GdHCl, GB1 domains lose their secondary and tertiary structures and become random coils. Therefore, the denatured (GR)12 hydrogel network is composed of random coil-like polypeptide chains, similar to synthetic polymer networks. As such, Equation 2-2 can be directly applied to the denatured (GR)12 hydrogel network with the real Mw of (GR)12. However, before and after denaturation of the same hydrogel, the number of dityrosine crosslinking points per chain in the hydrogel network remained unchanged. As a result, the unfolding of folded GB1 domains led to a large increase of the chain length between the two neighboring crosslinking points and a decrease of the crosslinking density per unit volume in the hydrogel network. Thus, a reduction of the Young’s modulus as well as much greater swelling of the denatured hydrogel are expected. Indeed, the denatured hydrogel showed a greater degree of swelling (Figure 2-5A), and the Young’s modulus of the denatured samples was dramatically lower than their native ones (Figure 2-5B, 2.6 kPa for the denatured hydrogel versus 9.9 kPa for 64 the native hydrogel). Using Equation 2-2, we estimated a -.1///// of 41.4 kDa for the denatured hydrogel, which gave rise to an Nc of 2.4. The values of Nc determined from the native and denatured (GR)12 hydrogels are almost the same (Table 2-1). Table 2-1. Average results applying length equivalent molecular weight on (GR)12, (GB1)8, (FL)8 hydrogels. Young’s modulus (E), molecular weight (Mw), length equivalent molecular weight (Me), length equivalent %&'''' (%&(''''') and Nc of 20% (GR)12, (GB1)8 and (FL)8 hydrogels. Data are presented as average ± standard deviation. For each data, n=5. ρ mg/cm3 E (kPa) Mw (kDa) Me (kDa) %&( (kDa) Nc (GR)12 (20%) native 158.5 ± 12.8 9.7 ± 0.9 99.0 34.9 15.3 ± 0.2 2.3 denatured 61.6 ± 1.3 1.9 ± 0.4 99.0 99.0 41.4 ± 1.3 2.4 (GB1)8 (20%) native 105.5 ± 20.1 4.8 ± 1.1 51.0 8.3 4.0 ± 0.0 2.1 denatured 40.1 ± 4.0 0.4 ± 0.3 51.0 51.0 24.7 ± 0.5 2.1 (FL)8 (20%) native 220.0 ± 4.6 4.1 ± 0.4 73.9 7.3 3.6 ± 0.0 2.0 denatured 17.6 ± 1.9 0.3 ± 0.0 73.9 73.9 33.9 ± 0.5 2.2 To further check the validity of this approach, we also tested protein hydrogels constructed from polyprotein (GB1)8 and (FL)8. (FL)8 is made of 8 tandem repeats of a small de novo designed globular ferredoxin-like protein (FL) and has been used to engineer highly elastic and tough protein hydrogels.34 For the same protein, the results revealed that the values of Nc determined for the native and denatured states are largely consistent with each other. These results provide a validation for the length equivalent molecular weight approach for tandem modular protein-based hydrogels. 65 2.4.3 Length equivalent molecular weight approach can be used to predict the change of the mechanical properties of dynamic protein hydrogels The unique property of tandem modular protein-based hydrogels is that the unfolding of folded globular domains can provide a powerful means to tune the hydrogel stiffness by controlling the folded state of the globular domains in the hydrogel network. Having demonstrated the validity of the length equivalent molecular approach, we further used this approach to semi-quantitatively predict the change of Young’s modulus of dynamic protein hydrogels constructed from elastomeric proteins containing a folding switch. As Nc=M/Mc, Equation 2-2 can be rewritten as: Equation 2-5 G= ABCE$D ;1 − FE$@ Since the number of crosslinking points Nc per chain remains constant before and after the complete unfolding of globular domains, the change of Young’s modulus can be readily estimated from ρ/M of the elastomeric protein before and after the unfolding. If ρ of the hydrogel network does not change significantly, the change of the modulus can be estimated from the change of M alone. 66 Figure 2-6 Using length equivalent molecular weight approach to predict the change of the mechanical properties of MEP-based dynamic hydrogels. A) Schematic showing the MEP-based, redox-responsive folding switch. B) Schematics of the MEP-based dynamic protein hydrogel. A) and B) are adapted from Ref. 43 with permission, Copyright 2014, WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim. In our previous work, we engineered an elastomeric protein G-R-(G-MEP-R)2, where MEP is a protein folding switch where a guest domain is inserted into a host domain. The host domain can change its folded state in response to the change of redox potential (Figure 2-6).43 In the oxidizing condition, the host domain of MEP is folded with a distance of 2.6 nm. In the reduced state, the host domain is unfolded and the whole MEP is ~30 nm long. Using the Me approach, the Me of G-R-(G-MEP-R)2 in the oxidized state is ~10.4 kDa, while its Me is 27.1 kDa in the reduced state. Assuming ρ does not change significantly upon reduction, the length equivalent molecular weight approach predicted an ~3 fold change in Young’s modulus of G-R-(G-MEP-R)2 hydrogels 67 upon reduction (Table 2-1). This prediction is in close agreement with the experimental results (2.8 kPa in the reduced state versus 9.0 kPa in the oxidized state for 20% hydrogel).43 It is of note that although the length change of MEP before and after reduction is about 10 fold, yet only a 3 fold change in Young’s modulus was achieved, which is much smaller than that observed for (GB1)8 and (FL)8. This reduced change of Young’s modulus is due to the presence of G and R sequences in the polyprotein whose length did not change, serving as an element damping the change caused by MEP. This insight is particular useful in helping design new dynamic protein hydrogels whose mechanical properties are tuned using protein unfolding as a stimulus. Table 2-2 Prediction of MEP-based dynamic hydrogel (Eoxidized/Ereduced) based on Me.  Me (kDa) Predicted Eoxidized/Ereduced E (kPa) Experimental Eoxidized/Ereduced G-R-(G-MEP-R)2 oxidized 10.4 2.6 9.0±0.9 3.2 reduced 27.1 2.8±0.4 2.4.4 Mechanical properties of (GR)-based hydrogels The length equivalent molecular weight approach also provides a useful means to quantitatively investigate and understand the mechanical properties of (GR)-based hydrogels in a systematic manner. In synthetic polymer based hydrogels or a rubber network, mechanical properties can be significantly improved by increasing the average molecular weight of polymer chains.302 We first investigated the effect of molecular weight (or chain length) on the elastic properties of (GR)n-hydrogels. For this purpose, we engineered a series of polyproteins (GR)n, where n=4, 8, 12, 16 and 24, using protein engineering techniques (Figure 2-2 and Figure 2-3) and carried out tensile testing (Figure 2-7A). As shown in Figure 2-7B, the increase of the chain length led to significant 68 increases in the Young’s modulus of the hydrogels. In this vein, the lower molecular weight polyproteins, such as (GR)4 and (GR)8, can only be largely regarded as oligomers, and their resultant protein hydrogels showed relatively low Young’s modulus. Calculation using Equation 2-2 revealed that the number of crosslinking sites per chain (Nc) increased from 2.0 for (GR)4 to 3.2 for (GR)24, clearly indicating that the increase of the chain length effectively increased the number of crosslinking points per protein polymer chain. Assuming that the crosslinking efficiency is the same for GR monomers, higher molecular weight/longer protein polymer chain lengths will lead to a higher number of crosslinking points per chain when the net GR monomer concentration is the same for different (GR)n. This finding is consistent with prior work on synthetic polymer hydrogels as well as protein biomaterials.212, 302-303 Figure 2-7 Molecular weight of globular polyproteins in crosslinked network contributes to mechanical properties. A) Stress-strain curves of 20% GR)n-based hydrogels. B) The Young’s modulus and the number of crosslinking points per protein chain of (GR)n-based hydrogels depend on the degree of polymerization of (GR)n-based elastomeric proteins. The error bars indicate the standard deviation. For each data point, the number of replicates is 5. We then investigated the effect of protein concentration on the mechanical properties of the protein hydrogels of (GR)12 (Figure 2-8). In synthetic polymer hydrogels, increasing polymer concentration often leads to an increase of Young’s modulus. However, in (GR)12 hydrogels, we 69 did not observe this effect. Instead, we observed that the (GR)12 protein concentration did not have a significant effect on the Young’s modulus of both native and denatured hydrogels within a protein concentration range of 10% to 20%. To investigate the origin of this unusual behavior, we measured their equilibrium swelling ratio and calculated their average crosslinking points per chain Nc. Interestingly, the swelling ratio of the protein hydrogels also showed behaviors different from typical synthetic polymer hydrogels, as its swelling ratio increased with the protein concentration. Also, the average number of crosslinking points per chain Nc decreased with the increasing of protein concentration, suggesting that fewer crosslinking sites were formed at higher protein concentration. One possible reason for these unusual behaviors is the steric hindrance caused by the bulky folded GB1 domains. Different from disordered polypeptide or synthetic polymers, bulky folded GB1 domains may restrain the formation of crosslinks between tyrosine residues at higher protein concentrations. This result suggested that increasing protein concentration may not lead to noticeable improvement on Young’s modulus of (GR)12 hydrogels. The reduced number of crosslinking sites per chain led to increased swelling, effectively offsetting the effect caused by the increased protein concentration. However, this effect may be protein-specific, as other protein hydrogels, such as those constructed from FnIII domain,285 showed an increase of the Young’s modulus with the increase of protein concentration. 2.5 Conclusions By introducing a novel approach, the length equivalent molecular weight for folded globular domains, we have adapted the classical rubber elasticity theory to describe the elastic properties of tandem modular protein-based protein hydrogels. This new approach has not only allowed us to quantitatively predict/explain the mechanical properties of such elastomeric protein-70 based hydrogels, but may also provide a guide to the rational engineering of elastomeric proteins to achieve tailored, dynamically tunable mechanical properties of protein hydrogels for applications in mechanobiology and biomaterials. Figure 2-8 Native and denatured (GR)12 hydrogels at protein concentration ranging from 10% to 20%. A) Young’s modulus of native and denatured (GR)12 hydrogels at protein concentration ranging from 10% to 20%. Student t-test showed that the difference in Young’s moduli was statistically not significant (p>0.05). B) Crosslinks numbers per (GR)12 polyprotein chain (Nc) obtained from the length equivalent molecular weight approach. The difference for Nc at different concentrations is statistically not significant. C) Swelling ratio of (GR)12 hydrogels as a function of protein concentration. Swelling ratio of (GR)12 increases in native and denatured states increases with increasing protein concentrations. Statistical significance is indicated (***: p<0.001; *: p<0.05; n.s.: p>0.05, no statistical significance). Error bars indicate standard deviation. 71 Chapter 3: Harvesting Mechanical Work from Folding-based Protein Engines: From Single-Molecule Mechanochemical Cycles to Macroscopic Devices 3.1 Synopsis Mechanochemical coupling cycles underlie the work generation mechanisms of biological systems, and are realized by highly regulated conformational changes of the protein machineries. However, it has been challenging to utilize protein conformational changes to do mechanical work at a macroscopic level in biomaterials, and it remains elusive to construct macroscopic mechanochemical devices based on molecular-level mechanochemical coupling systems. Here we demonstrate that protein folding can be utilized to realize a protein’s mechanochemical cycles at both single-molecule and macroscopic levels. Our results demonstrate, for the first time, the successful harnessing of mechanical work generated by protein folding in a macroscopic protein hydrogel device, and the work generated by protein folding compares favorably with the energy output of molecular motors. Our work bridges a gap between single molecule and macroscopic levels, and paves the way to utilizing proteins as building blocks to design protein-based artificial muscles and soft actuators. *A version of Chapter 3 has been published as [Linglan Fu], Han Wang and Hongbin Li. (2019) Harvesting Mechanical Work From Folding-Based Protein Engines: From Single-Molecule Mechanochemical Cycles to Macroscopic Devices. CCS Chemistry 1(1) 138-147. Open access with permission of Chinese Chemical Society, Copyright 2020. 72 3.2 Introduction Mechanochemistry refers to the mutual influence between mechanical and chemical phenomena at a molecular level, including chemical reactions induced by mechanical forces and mechanical output triggered by chemical input.304 Mechanochemical coupling cycles underlie the work generation mechanisms of biological systems,305 ranging from muscle contraction,306-307 adenine triphosphate (ATP) synthesis308 to cargo transport.309-310 Effective free energy transduction, a defining feature of such mechanochemical coupling processes, is realized by the folding and assembly of protein machineries and their highly regulated conformational changes. During this process, the system is converted from a high free energy state to a low free energy state and performs work. For example, key steps involved in the working cycle of the molecular motor myosin are conformational changes of the myosin motor domain, which result from the binding and hydrolysis of ATP and the subsequent release of ADP, and lead to the generation of the power stroke and mechanical work.311 In addition, the folding of immunoglobulin-like domains in the giant muscle protein titin can generate work to assist muscle active contraction.312-313 However, it remains elusive to construct macroscopic mechanochemical devices from such molecular-level mechanochemical coupling systems. Here we demonstrate the first realization of such a macroscopic mechanochemical device based on the mechanochemical coupling of proteins. Folding-unfolding is the ultimate conformational change in proteins, resulting in a large change in free energy, effective chain length, flexibility and mechanical stability of proteins.282 Using a well-characterized artificial elastomeric protein domain GB1,29, 283 here we demonstrate the mechanochemical cycles of proteins at both single-molecule and macroscopic levels. We were able to lengthen and contract individual GB1 polyprotein chains by switching GB1 between its folded and unfolded conformations. The protein polymer chain was found to contract against an 73 external force acting on the protein chain, thus generating mechanical work at the single molecule level. At the macroscopic level, when GB1 is switched from its unfolded to folded conformation, protein hydrogels engineered from GB1 polyprotein can contract against a considerable external load, generating mechanical work that can be harnessed and utilized at the macroscopic scale. Our results demonstrate for the first time mechanochemical energy conversion for proteins in both single molecule and macroscopic devices, and open a new avenue towards utilizing proteins as building blocks to designing protein-based soft actuators for various potential applications. The development of single molecule force spectroscopy techniques has allowed the measurement of force-induced unfolding-folding transitions at the single molecule level in great detail.314-316 It has been shown that upon stretching, proteins can unfold under a stretching force, leading to an increase in the length of the protein and a decrease in the mechanical resistance to stretching. Upon relaxation back to zero force, unfolded proteins can refold to regain their mechanical stability. Some proteins can refold even against a stretching force on the order of a few piconewton, leading to the shortening of the protein chain length and generation of mechanical work against a stretching force.317-321 If the mechanical work generated during protein folding at the molecular level could be harnessed at the single-molecule as well as macroscopic levels via a mechanochemical coupling cycle, a new molecular mechanism for generating actuation and mechanical work at both single-molecule and macroscopic levels should become feasible, and so are developing mechanochemical devices at both scales. By combining single molecule force spectroscopy, protein and hydrogel engineering, here we demonstrate the proof-of-principle of realizing such a mechanochemical operating cycle at both single molecule and macroscopic level, and harvesting the mechanical work from protein folding at two distinct length scales. 74 3.3 Materials and methods 3.3.1 Protein engineering The genes of polyproteins (GB1)4, GRG5RG4R and ferredoxin-like protein (FL)8 were constructed following standard and well-established molecular biology methods as reported before.29, 34, 283 Polyprotein genes were inserted into pQE80L vector for protein expression in E. coli strain DH5α. Protein expression and purification were performed using the same procedure as described in Chapter 2 (2.3.1). 3.3.2 Optical trapping experiments Optical trapping experiments were carried out on a MiniTweezers setup, which has been described previously.322-323 The construction of dsDNA handles and protein-DNA chimera was carried out following well-established protocols as described.322 A streptavidin modified polystyrene bead (with a diameter of ~ 1 µm, purchased from Spherotech) was sucked by a micro pipette tip, and an anti-digoxigenin beads (with a diameter of ~ 2 µm, purchased from Spherotech) sparsely covered with protein-DNA chimera was trapped by a laser beam and brought into close proximity to build a bead-DNA-protein dumbbell. Experiments were performed in PBS buffer. Data analysis was done using a custom program written in Igor Pro. 3.3.3 Hydrogel preparation and dimension measurements 20% GRG5RG4R and (FL)8 hydrogel rings are prepared and measured as described in Chapter 2 (2.3.2). 75 3.3.4 Tensile Tests Tensile tests were performed on an Instron-5500R tensometer with a custom-made force gauge in PBS or 7 M GdHCl PBS at a constant temperature (25 ℃). For technical reasons, ring-shaped specimens were used.29, 34, 283 An extension rate of 25 mm/min was used in our experiments. The stress at 15% strain is taken as the Young’s modulus of the sample. Experimental details and the Intron setup are shown in Section 2.3.3. To measure refolding and work-generation process of denatured hydrogel rings, the denatured ring sample was first stretched by a fixed load or to a fixed extension. During the refolding process, the extension of the hydrogel under the constant load or the load generated in the hydrogel under the constant extension was monitored. 3.3.5 Calculation of the work generated by the folding of proteins at the single molecule level At a single molecule level, the work generated by protein folding in a constant force experiment can be calculated as: F0×DL×Pf, where F0 is the constant force, DL is the contracted length of the protein polymer chain upon protein folding, and Pf is the folding probability of the protein. Both DL and Pf depend on the bias force F0.324-326 DL(F0) can be calculated by solving the worm-like chain (WLC) model of polymer elasticity:326 Equation 3-1 GH = IJKL × M14 21 − ∆3(GH)∆3P 8%F + ∆3(GH)∆3P − 14Q where kB is the Boltzmann constant, T is absolute temperature, p is persistence length, and ΔLc is the contour length increment upon protein unfolding/folding. For unfolded GB1, ΔLc is 18 nm and p is 0.6 nm. The folding probability Pf(F0) at a given force F0 can be calculated as: 76 Equation 3-2 RS(GH) = TUL (∆:H − GUIJK )/ W1 + TUL 2∆:H − GUIJK 8X where ΔG0 is the free energy difference between the unfolded and native states of the protein at zero force, and x is the distance between the native and unfolded states of protein (x=Δxf+Δxu), Δxu is the distance between the native and transition state, and Δxf is the distance between the unfolded and the transition state. For GB1, ΔG0 equals to 24.8 kJ/mol,327-328 and x is 3 nm283. It is of note that in these calculations, the constant force condition is used. Under such a condition, all protein domains in the polyprotein chain behave independently and thus the number of folded domains in the polyprotein only affects the total amount of work being generated; yet the work generated per domain remains the same. 3.4 Results and discussion 3.4.1 Mechanochemical coupling of GB1 at the single molecule level Here we choose a well-characterized artificial elastomeric protein GB1 as our model system.283 GB1 is a small globular protein with an a-helix packing against a b-sheet and its mechanical properties have been investigated in great detail using single molecule atomic force microscopy.283 However, the folding of GB1 has not been observed directly using force spectroscopy techniques. To directly examine whether GB1 domain can refold against a stretching force, we constructed DNA-(GB1)4-DNA chimera, which consists of four tandem repeats of GB1 and is flanked by two dsDNA handles, and used optical tweezers to directly characterize the mechanical unfolding and folding of GB1. Figure 3-1 shows the representative consecutive stretching-relaxation curves of DNA-(GB1)4-DNA at a pulling velocity of 50 nm/s. In the 77 stretching traces, unfolding events of GB1 occurred at forces that are larger than ~40 pN and showed a contour length increment DLc of ~18 nm. Due to its high mechanical stability, in most pulling curves only one or two GB1 were observed to unfold. This unfolding force measured in the optical tweezers experiment is significantly lower than that of GB1 measured in AFM experiments.283 This can be explained by the fact that the unfolding force depends on the loading rate. The effective loading rate in the optical tweezers experiments is significantly lower than that in AFM, arising from the lower spring constant of the optical tweezers and the much longer spacer (dsDNA handles). Upon relaxation, the unfolded GB1 domains were observed to refold individually against a stretching force of up to ~7 pN, leading to the contraction of the protein polymer chain against the stretching force, and the generation of the mechanical work. Figure 3-1 Force-distance curves of (GB1)4 show that GB1 can refold against a stretching force. Stretching dsDNA-(GB1)4-dsDNA resulted in the mechanical unfolding of GB1 domains at forces typically higher than 40 pN at a pulling speed of 20 nm/s. Upon relaxation, the unfolded GB1 domains were observed to refold against a stretching force of ~7 pN, resulting in the shortening of the protein-DNA chain and the generation of mechanical work. Due to its high mechanical stability, only one or two GB1 domains were observed to unfold in constant velocity experiments. Red traces are the raw data and blue traces are filtered data to 100 Hz. 78 Figure 3-2 Mechanochemical cycles lead to the harvesting of mechanical work generated by the folding of proteins. A) A gedanken experiment shows the realization of the mechanochemical coupling and mechanochemical cycles to harness the mechanical work generated from protein folding. In state 1, a folded globular protein polymer is pre-stretched by a bias weight F0, at which the unfolded protein can refold and which is often smaller than F1/2 (the force where the folding probability Pf is 0.5). Then a denaturing stimulus, being chemical, mechanical or optical, is applied to trigger the unfolding of folded proteins in the polymer chain, leading to the lengthening of the protein polymer chain and bringing the system to state 2. After the system reaches equilibrium, the denaturing stimulus is removed and a renaturing stimulus is applied. Against the bias force F0, the unfolded proteins will refold, leading to the contraction of the protein polymer chain and the generation of work )* × ∆-. After reaching equilibrium, the system is back to the original state (state 1’), and a new cycle is ready to start again. Repeating these mechanochemical cycles, the work generated by protein folding can be harnessed at both single molecule and macroscopic levels. B) Realization of the mechanochemical cycles at the single molecule level. At state 1, the folded protein-DNA polymer is prestretched with a bias force of 7.5 pN to a length L1. Then a denaturing stimulus, a high stretching force of 45 pN, is applied. The folded protein-DNA polymer extends to L2 following the worm-like chain model of polymer elasticity. At this high stretching force, GB1 domains unfold, leading to the lengthening of the system to L3 and bringing the system to State 2. Then the high denaturing force is removed and renaturing force is applied (7.5 pN). Following WLC, the unfolded polymer shortens to L4. Then unfolded GB1 domains refold one by one, generating the characteristic staircase-like contraction of the polymer chain. After all GB1 domains have refolded, the system is back to the original state (State 1’) with a length of L1. During the folding process, the system generates work of F0×(L4-L1), which is 2.3×10-19 J. The shaded area corresponds to the work being generated. 79 How could we make use of such a protein polymer system to harness the generated mechanical work at the single molecule level? Figure 3-2A shows a gedanken experiment to illustrate the mechanochemical operating cycle to harness the work generated by the folding of individual proteins. The protein polymer is first pre-stretched with a weight F0 (State 1), which is small enough not to trigger protein unfolding but allow for unfolded protein to refold. Then, a denaturing condition would be applied, being thermal, chemical or mechanical, to switch the globular domains in the protein polymer chain from their folded states to unfolded states. Due to protein unfolding and the increase in contour length of the protein polymer chain, the protein polymer elongates (State 2). Switching the condition back to the renaturing condition, unfolded domains would be able to refold and the protein polymer contracts by DL, lifting the weight F0 thereby doing mechanical work F0×DL and bringing the system back to its original state (State 1’). Switching to the denaturing condition would allow the cycle to restart. Realizing this cycle at the single molecule level would allow the harnessing of the mechanical work generated at the single molecule level, while realizing this cycle at the macroscopic level would allow the utilization of protein folding to generate useful work at the macroscopic scale. At the single molecule level, it is challenging to realize such an operating cycle using a chemical/thermal denaturation method in an optical trap setup. However, mechanical denaturation would readily enable this cycle, as shown in Figure 3-2B. The folded GB1 protein was first stretched by a bias force F0 of ~7.5 pN to an extension of L1, and then a mechanical denaturation force was applied (F1=45 pN) to stretch the folded protein to the extension of L2, subsequent mechanical unfolding will lead to the extension to L3. Upon removing the denaturation force F1, the original bias force F0 is applied and the unfolded protein contracts to an extension of L4 due to 80 elastic recoiling of the unfolded protein chain, which can be described by the worm-like chain (WLC) model of polymer elasticity.326 Under the bias force of F0, the unfolded GB1 refolds, leading to the contraction of the protein polymer chain against F0 and thereby generating mechanical work of F0×(L4-L1). And the system is brought back to its original state 1 and a new cycle can start again. During the folding process, the system generated mechanical work of ~6×10-20 J per GB1 domain (7.5 pN×8 nm, where 8 nm is the length change upon the folding of a GB1 domain at 7.5 pN). This experiment clearly demonstrates the feasibility of harnessing mechanical work generated by protein folding at the single molecule level through a mechanochemical operating cycle. The work generated during this operating cycle consists of contributions from two processes, one is the refolding of GB1 and the other one is the elastic recoiling of the unfolded GB1. The work generated by the folding of GB1 is substantial, comparing favorably with the work generated by most molecular motors, such as myosin306 and kinesin,329 thereby making GB1 a folding-driven molecular engine. Harnessing such mechanical work at the macroscopic level would enable the generation of useful mechanical work in the macroscopic world, and provide new mechanisms for actuation. 3.4.2 Mechanochemical coupling at the macroscopic level As a proof-of-principle, we used protein hydrogels to demonstrate the realization of such mechanochemical coupling cycles at the macroscopic scale, where the switching of protein folding and unfolding is realized via chemical denaturation and renaturation. The protein hydrogel constructed from GRG5RG4R, where G corresponds to the GB1 domain and R corresponds to a 15-residue consensus sequence from the elastomeric protein resilin from Drosophila,275 was employed. This protein hydrogel, which was engineered to mimic the passive elastic properties of 81 muscles and has been studied in detail,29 was constructed by using a well-established ruthenium-catalyzed photo-crosslinking chemistry.296, 298 The GRG5RG4R hydrogel shows drastically different mechanical properties in its folded and denatured states (Figure 3-3A). In its native state, the 20% GRG5RG4R hydrogel displays a Young’s modulus of ~30 kPa in PBS. In contrast, the same protein hydrogel in the denatured state is much softer, displaying a Young’s modulus of ~3 kPa. This large decrease in the Young’s modulus is due to the unfolding of GB1 domains, where the length of the unfolded GB1 domain is about ten times longer than the folded one, leading to a large increase of the effective polymer chain length between two neighboring crosslinking points. This sharp difference in the mechanical properties of the protein hydrogel in its two distinct states allowed us to readily realize the mechanochemical operating cycle (Figure 3-3B). First, a bias load of 5 g was applied to stretch the hydrogel from its original length of 11 mm to 17 mm in its native environment (PBS). To switch GB1 from its folded to unfolded state, we applied chemical denaturation (7 M GdHCl). After equilibration, the protein hydrogel extended significantly to ~32 mm under the bias load of 5 g. Then, to switch GB1 from its unfolded state back to folded state, the buffer was changed back to the renaturing buffer PBS. We observed that the hydrogel contracted towards its original length in PBS and lifted the 5g weight (which is 50 times heavier than the hydrogel itself) by ~ 15 mm, producing mechanical work of ~0.68 mJ. Repeating this cycle (i.e. cycling between 7 M GdHCl and PBS) leads to the generation mechanical work in a repetitive fashion (Figure 3-3B). Figure 3-4A and B show the time course of the evolution of extension of the hydrogel under a constant load as well as the stress-extension curves during this process. The mechanochemical coupling of the GRG5RG4R hydrogel at a fixed length is also evident (Figure 3-4C and D). When the protein hydrogel was held at a constant extension, changing the buffer from 7 M GdHCl to PBS resulted 82 in a continuous increase of the tension in the protein hydrogel from ~1 g to ~12 g. Moreover, the switching between folded and denatured state of the hydrogel is fully reversible, as evidenced by the length and the Young’s modulus of the protein hydrogel measured in the tensile testing (Figure 3-3C). Hence, the mechanochemical cycle observed here is also fully reversible (Figure 3-3C), and no obvious creeping effect was observed. These results clearly demonstrate that switching the conformations of the globular protein GB1 in the protein hydrogels via a chemical means leads to the effective mechanochemical coupling of the protein hydrogel and the generation of mechanical work, revealing the feasibility of harnessing the mechanical work generated by protein folding at both single molecule and macroscopic levels via mechanochemical coupling. It is of note that this protein hydrogel system displays a work density of ~5.9 kJ/m3, which is comparable to that of skeletal muscles powered by ATP.330 In this example, chemical denaturation of proteins was used to realize the mechanochemical cycles in the hydrogels. In principle, mechanical denaturation, similar to that in single molecule force spectroscopy experiments, can also be used to realize the mechanochemical cycles, albeit much less efficient. Previous work showed that stretching protein-based hydrogels to high strain can trigger the unfolding of some folded protein domains, leading to the observed hysteresis between the stretching and relaxation curves.29, 33-34, 286 Even under residual stress, protein hydrogels can partially recover the hysteresis, suggesting that unfolded proteins can refold under residual stress and thus generate work. However, mechanical stretching can only lead to the unfolding of a small fraction of protein domains in hydrogels, thus mechanical denaturation is a less efficient way in which to realize the mechanochemical cycles. 83 Figure 3-3 Mechanochemical cycles at the macroscopic level lead to the harnessing of mechanical work generated by protein folding in protein hydrogels. A) Stress-extension curves of 20% protein hydrogels based on the polyprotein GRG5RG4R. In the folded state, GRG5RG4R-hydrogel shows a Young’s modulus of ~30 kPa; in the unfolded state, the hydrogel becomes much softer with a Young’s modulus of ~3 kPa. The dimension of the denatured hydrogel is much larger than the folded one. B) Realization of mechanochemical cycles at the macrocopic level leads to the generation of mechanical work. At state 1, the hydrogel is prestreched with a bias weight of 5 g in PBS, where GB1 domains are folded. Then a denaturing stimuli is applied by switching the buffer to 7 M GdHCl, leading the unfolding of GB1 domains and lengthening of the hydrogel under the weight of 5 g (State 2). After reaching equilibrium, the denaturing stimuli GdHCl is replaced by a renaturing stimuli (PBS). The hydrogel contracts and lifting the 5g weight by ~15 mm and generating mechanical work of ~0.68 mJ. This value is the net work generated by the protein hydrogel after substracting the work produced by the buoyance force during this process. Reapeating this cycle leads to the generation of mechanical work in a repetititive fashion. The inset shows the force-extension curve, where the shaded area indicated the work being generated when the hydrogel is placed under a renaturing condition. C) The macroscopic mechanochemical cycles are reversible. The Young’s modulus and the length of the hydrogels show reversible changes in response to the change of the denaturing and renaturing stimuli, as measured in the tensile testing experiments at a constant load of 3 g. 84 Figure 3-4 Dynamic response of GRG5RG4R-based hydrogels during mechano-chemical coupling. A and C) Dynamic response of the 15% GRG5RG4R-based hydrogels under a constant load (A) and constant extension condition (C) when the buffer is switched from PBS to 7 M GdHCl, respectively. Red lines are exponential fits to the experimental data (blue symbols). The estimated rate constants are 1.0×10-3 s-1 (A), and 3.2 x 10-3 s-1 (C), respectively. B and D) Time evolution of load-extension curves of the 15% GRG5RG4R-based hydrogels under a constant load (B) and constant extension condition (D) when buffer is switched from PBS to 7 M GdHCl, respectively. 3.4.3 Bridging the gap between the mechanochemical coupling at single molecule level and macroscopic level The nature of the mechanochemical coupling at both single molecule and macroscopic levels is common and lies in the folding transitions of globular protein GB1 against a stretching force. This common nature gives rise to the unique yet similar dependence of the generated work on the bias force at both single-molecule and macroscopic scales. At the single molecule level, the work generated by protein folding is F0×DL×Pf, where F0 is the bias force, DL is the contracted length of the protein polymer chain upon protein folding, and Pf is the folding probability of the 85 protein. Both DL and Pf depend on the bias force F0.324-326 As shown in Figure 3-5, the quantitative relationship between the bias force and mechanical work output can be predicted based on the free energy profile of the mechanical folding and unfolding of GB1. The work generated by protein folding reaches a maximum at a particular bias force Fc (which is smaller than F1/2, a force at which the folding probability Pf is 50%.). When the bias force is smaller than Fc, the output of the generated mechanical work shows a monotonic increase with the bias force, in good agreement with the experimental data measured in our optical trapping experiments. Figure 3-5 Mechanical work generated by the folding of GB1 domains at the single molecule level. A) Force dependency of the the length change DL of the protein polymer chain upon GB1 folding (solid line: WLC prediction, symbols: experimental data). B) Force dependency of the mechanical work generated by the folding of GB1 domains. Red line is the folding probability of a GB1 domain at different forces, which is calculated based on the free energy difference between the folded and unfolded states of GB1 (DG=10 kBT) and the distance between the folded and unfolded states (Dxu+ Dxf = 3 nm). Blue line is the work generated by the folding of a GB1 domain at different forces. The work is calculated using the following relationship: w= F0DLPf , where F0 is the bias force against which GB1 domain folds, DL is the length change of the protein polymer at the force F0, and Pf is the folding probability of GB1 at the force F0. Fc indicates the force at which the generated work is maximum, and F1/2 indicated the force at which the folding probability Pf is 0.5. Symbols are experimental data measured from optical tweezers experiments. 86 Figure 3-6 The output of the mechanical work can be tuned by the bias load in mechanochemical cycles. A and B) Tensile extension of the protein hydrogel rings in their native (black) and denatured states (red). (A: 20% GRG5RG4R hydrogel; B: 20% (FL)8 hydrogel). C and D) The load dependency of the length contraction of the hydrogels upon switching from denaturing to renaturing condition. (C: GRG5RG4R; D: (FL)8). The contraction is measured from the length of the hydrogels in their native and denatured states in the tensile testing (length in A and B). E and F), The mechanical work generated in the mechanochemical cycles depends on the bias load (E: GRG5RG4R; F: (FL)8). At the macroscopic level, the load dependency of the contracted length (Figure 3-6A) and the generated work (Figure 3-6B) of the GRG5RG4R protein hydrogel upon switching from denaturing to renaturing condition is evidenced from tensile testing. The work output increased from ~0.04 mJ to 0.76 mJ when the load was increased from 0.5 g to 4.5 g (Figure 3-6A and B). 87 Currently, the maximum work output from our hydrogels is limited by the breaking strain of the hydrogel in the tensile testing, which is likely limited by the macroscopic defects in the hydrogel. We estimated that the average force experienced by individual GRG5RG4R molecules in the hydrogel network caused by the bias load is likely to be significantly smaller than the F1/2 of GB1, which is estimated to be ~13 pN. These results clearly indicate the close correlation between single molecule and macroscopic mechanochemical cycles and mechanochemical actuation, and strongly suggest that the same physiochemical principles govern the behaviors at these two distinct length scales. Figure 3-7 Dynamic response of the 20% GRG5RG4R-based hydrogel at zero load when the buffer is switched from PBS to 7 M GdHCl. The red line is the single exponential fit to the experimental data (blue symbols) with a rate constant k of 7.6×10-4 s-1. It is important to note that despite this close correlation between single-molecule and macroscopic mechanochemical cycles, the dynamic response of the mechanochemical operating cycle shows large differences between single-molecule and macroscopic ones. The dynamic response of the mechanochemical cycle is slow, showing a first order kinetics with a rate constant k of 1.0×10-3 s-1 (Figure 3-4A). Even in the absence of an external load, the dynamic response of 88 the protein hydrogel in the macroscopic mechanochemical cycles is also very slow (with a rate constant of 7.6×10-4 s-1) (Figure 3-7), significantly slower than the folding rate constant of GB1 at zero force (which is 720 s-1 as measured in single molecule AFM283 and chemical denaturation experiments).328 327 This large difference is likely due to the slow mass transport in the protein hydrogels during buffer exchange. 331-333 3.5 Conclusions Free energy transduction allows living organisms to access available energy and convert it into those needed to power and propagate life.334 Through free energy transduction, not only active machineries, such as molecular motors, can power biological movement, but also “passive” biomacromolecules, including proteins,317-321 DNA/RNA335-336 and polysaccharides,337 as well as synthetic polymers338-339 can also undergo conformation changes (such as folding) against a stretching force to perform work at the single molecule level. However, making use of such conformational changes, in other words free energy transduction processes, to do mechanical work at the macroscopic level has been challenging. Using a small artificial elastomeric protein GB1 as a model system, here we demonstrate for the first time that a large protein conformation change, i.e. folding-unfolding, can be controlled at both single molecule and macroscopic levels via mechanochemical coupling cycles, and be utilized to generate mechanical work that can be harnessed at the macroscopic level. This is the first realization of such a mechanochemical operating cycle at both levels, bridging a gap that researchers have long been attempted to fill. For example, pioneering studies demonstrated the successful realization of such mechanochemical operating cycles either at the single molecule level, such as the optomechanical cycle in single azobenzene-containing polymers,338 or at the macroscopic level, such as the mechanochemical 89 coupling in elastin-based protein hydrogels,340 but bridging the gap between these two distinct levels proved challenging and remained ellusive.341 Our study enables a novel approach to directly connect single molecule mechanochemical events with macroscopic mechanical work generation. It is of note that at the macroscopic scale, the work generation in protein hydrogels involves two intertwined processes, protein folding and elastic recoiling of unfolded polypeptide chains, which manifests itself as shrinking. In chemically denatured protein hydrogels, the swelling force of the hydrogel network will lead to the stretching of the unfolded protein chain and the swelling the hydrogel; under renaturing conditions, proteins fold accompanied by elastic recoiling of the polypeptide chain and shrinking of the hydrogel network. In this vein, the work generation during the shrinking process of hydrogels made of temperature responsive polymers (such as pNIPAM (poly-N-isopropylacrylamide)) and polypeptide chain (such as elastin), which is caused by the temperature induced collapse of the polymer network, shares similar features with the work generation process reported here for protein folding, although the temperature driven collapse of polymer/polypeptide chains is much less specific than protein folding. Moreover, the controlled mechanochemical coupling cycle of GB1, provides a new avenue towards engineering protein-based artificial muscles and soft actuators. Since the mechanism for work generation is protein folding-based mechanochemical coupling, this mechanism should be general and be applicable to a wide variety of proteins. It should be feasible to synergistically fine-tune the nanomechanical properties of individual proteins (including folding properties) at the single molecule level and protein hydrogel mechanics at the macroscopic level in order to further tune the mechanical work generation in protein hydrogels. For example, using a ferredoxin-like protein (FL) as building blocks,34 we constructed protein hydrogels that show much improved mechanical work generation based on protein folding. Compared with GRG5RG4R-based 90 hydrogels, the output of the mechanical work of FR-based hydrogels reaches 1 mJ ((Figure 3-6C and D) under a much smaller load. Our results open up a new avenue towards engineering protein-based materials and actuators, just like the way how nature constructs and utilizes the great variety of protein-based mechanochemical machineries, such as molecular motors, to accomplish a wide variety of biological functions. These novel materials should find applications as soft actuators for various potential applications. 91 Chapter 4: Converting Muscle-mimetic Biomaterials to Cartilage-like Materials 4.1 Synopsis Load-bearing tissues, such as muscles and cartilage, exhibit mechanical properties that often combine high elasticity, high toughness and fast recovery, despite their different stiffnesses (~100 kPa for muscles and one to several MPa for cartilage). 342-348 The advance in protein engineering and protein mechanics has made it possible to engineer protein-based biomaterials to mimic soft load-bearing tissues, such as muscles.29, 33, 287 However, it is challenging to develop protein biomaterials to achieve the mechanical properties exhibited by stiff tissues, such as articular cartilage,347, 349 or to develop stiff synthetic ECMs for cartilage stem/progenitor cell differentiation350. By employing physical entanglements207 of protein chains and force-induced protein unfolding34, 282, here we report the engineering of a highly tough and stiff protein hydrogel to mimic articular cartilage. By crosslinking an engineered artificial elastomeric protein from its unfolded state, we introduced chain entanglement into the hydrogel network. Upon renaturation, the entangled protein chain network and forced protein unfolding resulted in this single network protein hydrogel with superb mechanical properties in both tensile and compression tests, showing a Young’s modulus of ~0.7 MPa and toughness of 250 kJ/m3 in tensile testing; and ~1.7 MPa in compressive modulus and a toughness of 3.2 MJ/m3. The energy dissipation in both tensile and compression tests is reversible and the hydrogel can recover its mechanical properties rapidly. Moreover, this hydrogel can withstand compression stress of >60 MPa without failure, which represents the highest compressive strength achieved by a hydrogel to the best of our knowledge. These properties are comparable to those of articular cartilage, making this protein hydrogel a novel cartilage-mimetic biomaterial. Our study opened up a new potential avenue towards 92 engineering protein hydrogel-based substitutes for articular cartilage, and may also help develop protein biomaterials with superb mechanical properties for applications in soft actuators and robotics. 4.2 Introduction Many load-bearing tissues, ranging from muscles to cartilage, are protein-based biomaterials that exhibit finely regulated mechanical properties to uniquely suit their biological functions.284, 342 To engineer biomimetics of these biological tissues, protein-based hydrogels have been widely explored.351 Protein hydrogels are generally soft, with a Young’s modulus up to ~100 kPa.29, 275 Thus, current protein hydrogel technologies have achieved considerable success in mimicking softer tissues264, 275, 351, such as muscles.29, 33-34 However, stiffer tissues often have a modulus on the order of megapascal (MPa), which is challenging to achieve using current protein hydrogel technology. For example, an articular cartilage is a superb load-bearing material made of collagen and proteoglycans. It has a modulus ranging from one to several MPa, can withstand a load up to a hundred MPa and sustain millions of cycles of loading-unloading without much wear or fatigue, showing fast recovery in its shape and mechanical properties.342, 348, 352 Developing protein hydrogels that can combine these seemingly mutually incompatible properties (high stiffness, high toughness and fast recovery) remains challenging. In order to engineer protein hydrogels with such a combination of mechanical features, higher crosslinking density and an efficient microscopic mechanism for energy dissipation and recovery are essential. In nature, articular cartilage possesses this unique combination of mechanical features through the use of protein fiber-polysaccharides composites.353 However, this approach is difficult to design from bottom up.354-355 On the other hand, muscles may provide an 93 inspiration for an alternative bottom-up design approach for engineering cartilage-like protein-based materials. Forced-unfolding and refolding of globular proteins is a mechanism employed in muscles to effectively dissipate energy upon overstretching, and allow muscles to recover upon relaxation.280, 282, 344 This mechanism has been used to engineer protein-hydrogels to mimic the passive elastic properties of muscles.29, 33-34, 287 We reason that if an efficient method can be developed to significantly improve the stiffness of muscle-mimetic biomaterials, one should be able to take advantage of the forced-unfolding as an efficient energy dissipation mechanism to convert muscle-mimetic biomaterials into biomaterials with mechanical properties resembling those of articular cartilage. Chain entanglement is an important strengthening mechanism in polymeric materials.207 Due to their long chains, polymer chains in the network can cross each other, resulting in chain entanglement. Chain entanglement effectively increases the crosslinking density in the polymer network and leads to improvements in Young’s modulus.207 However, in muscle fibers, the muscle protein titin are organized as parallel bundles without chain entanglement.343-344 In the engineered muscle mimetic protein hydrogels constructed from tandem modular protein-based elastomeric proteins, no chain entanglement is present either, due to the short contour length of such elastomeric proteins (about 10 to 40 nm in length).29, 34 Here we explore the feasibility to impart soft, muscle-mimetic protein biomaterial with chain entanglement to significantly increase its stiffness. For this purpose, we used the elastomeric protein (FL)8 as a model system. FL is a de novo designed ferredoxin-like globular protein.356 Single molecule optical tweezers experiments showed that FL is mechanically labile, and unfolds and refolds readily at ~5 pN, significantly lower than most folded elastomeric protein domains (Figure 4-1).34 Protein hydrogels constructed from the elastomeric protein (FL)8 were used to engineer highly stretchy 94 and tough protein hydrogels, in which the forced-unfolding of FL domains serves as a highly effective means in dissipating energy in the hydrogel. 34 However, the Young’s modulus of the (FL)8 hydrogel is only ~20 kPa.34 We sought to use (FL)8 as a model system for enhancing its mechanical stiffness. The molecular weight of (FL)8 is ~80 kDa, but its contour length in its native folded state is only ~10 nm. Thus, there is no chain entanglement in the native (FL)8 hydrogels. However, in its unfolded state, (FL)8 is ~260 nm long, showing the characteristic length of a polymer chain. In the concentrated solution of unfolded (FL)8, the unfolded polypeptide chains will overlap and likely entangle. We reason that if unfolded (FL)8 is crosslinked from its concentrated solution, chain entanglement could be introduced/retained in the crosslinked hydrogel network. Figure 4-1 Representative force-distance curves of FL domain at a pulling speed of 50 nm/s. The unfolding-refolding of FL occurred at ~ 5 pN, making FL a mechanically labile protein. The inset shows the three-dimensional structure of FL (PDB code: 2KL8). 95 4.3 Materials and methods 4.3.1 Protein engineering FL domain is a redesigned variant of Di-I_5 (PDB code: 2KL8).34, 356 The amino acid sequence of FL is given in Appendix A.2. The genes of the polyproteins (FL)8, (FLM23C)8, (FL)16, (NuG2)8, (GB1)8, GRG5RG4R, N4RN4RNR and (GA)8 were constructed following standard and well-established molecular biology methods as reported previously.29, 34, 276, 357-358 BSA lyophilized powder was purchased from Sigma. Polyprotein genes were inserted into the vector pQE80L for protein expression in E. coli strain DH5α. Protein expression and purification were performed using the same procedure as described in Chapter 2 (2.3.1). 4.3.2 Hydrogel preparation and dimension measurements Lyophilized native (FL)8 protein was processed using two different gelation methods. The NC (native crosslinking) hydrogels were prepared in native state by dissolving and gelating proteins in 1x PBS. After gelation, N-NC hydrogels were soaked in PBS, while the D-NC hydrogels were stored in 1x PBS containing 7M GdHCl until swelling equilibrium was achieved. The DC (denatured crosslinking) hydrogels (D-DC and N-DC) were prepared by dissolving the lyophilized (FL)8 in 7M GdHCl for 2 hrs before use. The denatured protein solution was crosslinked into hydrogels and equilibrated in 7M GdHCl to obtain D-DC hydrogels, while N-DC hydrogels were renatured in PBS on a rocker by changing a large amount of fresh PBS ten times over the course of 1 day until reaching equilibrium. Gelation of (FL)8, (FLM23C)8 and (FL)16 were based on the photochemical crosslinking strategy described previously.29, 34, 275, 296 A typical crosslinking reaction mixture contained 200 mg/mL of polyprotein, 50 mM ammonium persulfate (APS) and 200 μM [Ru(bpy)3]Cl2. The 96 protein mixture was quickly cast into a custom-made plexiglass ring-shaped mold (din=8 mm, dout=10 mm, h= 3 mm), and was exposed to a 200 W fiber optical white light source placed 10 cm above the mold for 10 min at room temperature. After gelation was complete, the hydrogel sample was carefully taken out of the mold. After the ring-shaped hydrogels were stored in the desired buffers for 3 hrs (N-DC gels for 24 hrs), the outer diameter, thickness, width and weight of all ring-shape equilibrated swollen/shrunken samples were measured before tensile tests. For compressive tests and SEM imaging, the hydrogels were prepared in a cylindrical shape following the same gelation procedures. 4.3.3 Swelling ratio and water content measurements For swelling ratio and water content measurements, ring-shaped hydrogels were weighed immediately after being taken out of the mold, and the weight was recorded as mi. The swollen weight ms was recorded after gently removing excess buffer from equilibrated hydrogels. To measure the dry weight of the gels (mdry), the hydrogels were firstly immersed in deionized water to remove extra salts, then lyophilized for 2 days and dried in a 70 ℃ incubator for 1 day. The swelling ratio (r) of the hydrogels was calculated using the following formula, " = $!%$"$" × 100%. The water content (w) was determined by using the weight of the dried gel (mdry) and the equilibrated gel (ms) and the following formula (measured specimens, n=4): Y = (1 − $/01$! ) × 100%. 97 4.3.4 Tensile tests Tensile tests were performed using an Instron-5500R tensometer with a custom-made force gauge and 5-N load transducer. The ring-shaped hydrogel specimen was stretched and relaxed in PBS (N-DC and N-NC) or 7 M GdHCl in PBS (D-DC and D-NC) at constant temperature (25 ℃) without special preconditioning. The stress was calculated by dividing the load by the initial cross-sectional area of the hydrogel sample. The Young’s modulus, breaking strain, and energy dissipation were measured using an extension rate of 25 mm/min. The stress at 15% strain is taken as the Young’s modulus of the sample. Toughness was determined by integrating stress-strain curves where specimens were loaded directly to failure. Energy dissipation was calculated by integrating the loop area between stretching and relaxing stress-strain curves. In hysteresis recovery experiments, a pulling rate of 200 mm/min was used. The same ring sample was stretched and relaxed at various time intervals (Measured specimens, n=18) 4.3.5 Compressive tests Uniaxial compression tests were performed on cylindrical hydrogels that were swollen to equilibrium using the Instron-5500R with 5000-N load transducer in air at room temperature. The dimensions (height (h0) and diameter (d0)) of the equilibrated N-DC and N-NC (FL)8 hydrogel samples are: h0: 3.0 mm and d0: 6.5 mm for the N-DC hydrogel, and h0: 5.0 mm and d0: 5.0 mm for the N-NC hydrogel. The gel was put on the lower plate, while the upper plate approached the sample slowly until a rise in force was detected, indicating contact between the plate and the gel. The stress was calculated by dividing the load by the initial cross-sectional area of the hydrogel sample. Unless a different rate is stated, the gel was compressed and relaxed at a compression speed of 2 mm/min. No water was squeezed out of the gels during compression. The compressive 98 modulus was measured at a strain of 10-20 %. The maximum stress and strain were determined at failure points, where the first crack in the gel was observed. Energy dissipation was calculated by integrating the loop area between compressing and relaxing stress-strain curves. In hysteresis recovery experiments, a compression rate of 100 mm/min was used. In fatigue experiments, a compression rate of 20 min/min and 100 mm/min were used respectively. (Measured specimens, n=7). 4.3.6 Scanning electron microscopy (SEM) imaging 20 % (w/v) D-NC and N-NC (FL)8 hydrogel samples were prepared for SEM imaging using a Hitachi S4700 scanning electron microscope. The samples were shock-frozen in liquid nitrogen, and quickly transferred to a freeze drier where they were lyophilized for 24 hrs. Lyophilized samples were then carefully fractured in liquid nitrogen, and fixed on aluminum stubs. The sample surface was coated by 8 nm of gold prior to SEM measurements. 4.3.7 Characterization of dityrosine cross-links in hydrogel by acid hydrolysis-fluorescence method The degree of dityrosine crosslinking in both NC and DC (FL)8 hydrogels was characterized following a well-established fluorimetry method.231 Dityrosine emits at a wavelength of 410 nm when excited at 315 nm. For quantification of the dityrosine and dityrosine-like compounds generated in NC and DC (FL)8 hydrogels, we followed a well-established fluorescence standard curve method.231 Typically, 20 % (w/v) hydrogel samples (∼25 mg) were reacted with 100 μL HCl (6 N) in a sealed 1.5 mL centrifuge tube in a metal heat block at 105 °C for 2 hrs to achieve full hydrolysis of the peptide bonds. Then, 100 μL of acid hydrolysis product 99 was transferred into a new 1.5 mL centrifuge tube and neutralized by 10 μL NaOH (5 M). Next, 100 mM Na2CO3-NaHCO3 buffer (pH 9.9) was added to a final volume of 1 mL. Fluorescence spectra of the samples were measured by a Varian Cary Eclipse fluorescence spectrophotometer. According to the fluorescence-concentration standard calibration curve of dityrosine, the yield of dityrosine and dityrosine-like products in the hydrogel was then determined. (Measured specimens, n=8) 4.3.8 Cysteine shotgun fluorescence labeling by IAEDANS and fluorescence measurements DC and NC (FLM23C)8 hydrogels for cysteine shotgun labeling were prepared with the same protein concentration and gel preparation procedures as the wild-type (FL)8. The labeling reaction was performed in the dark at room temperature for 3 hrs in PBS buffer (pH 7.4) containing 5 mM TCEP and 2 mM 5-((2-[(iodoacetyl)amino]ethyl)amino)naphthalene-1-sulfonic acid (IAEDANS). As a control, D-DC and D-NC (FLM23C)8 hydrogels incubated in 7 M GdHCl (containing 5 mM TCEP, pH 7.4) were also labeled using IAEDANS. Then all hydrogels were transferred to PBS buffer containing 5 mM β-mercaptoethanol to quench the reaction. To remove excess labeling dye, additional PBS buffer (containing 5 mM β-mercaptoethanol) was added, and changed five times over the course of 5 hrs until fluorescence could no longer be detected in the buffer solution. To quantify the fluorescence intensity of IAEDANS labeled hydrogels, we digested the hydrogels with trypsin at 37 °C for 5 hrs. The digestion reaction contained 5 % trypsin (relative to the hydrogel weight), 25 mM NH4HCO3, 10 mM CaCl2, 1 M GdHCl and 10 mM dithiothreitol. Unlabeled (FLM23C)8 hydrogel was digested in the same way to serve as a negative control. After 100 digestion, 50 μL of the digested mixture was diluted to 3 mL using PBS buffer. The fluorescence emission spectrum was measured using a Varian Cary Eclipse fluorescence spectrometer using an excitation wavelength of 336 nm and an emission wavelength of 490 nm. An IAEDANS standard calibration curve was also created, over the concentration range of 0-60 μM (Measured specimens, n=8). 4.3.9 Single-molecule optical tweezers measurements. Single-molecule optical tweezers measurements were carried out using a MiniTweezers setup as previously.322, 359 Sample preparation including the protein–DNA construct formation and force measurement protocols was adapted from protocols described previously.322 Force-distance curves of the protein–DNA construct were obtained using a constant velocity pulling protocol. 4.3.10 Hydrogel biocompatibility measurement, LIVE/DEAD viability assay. Human lung fibroblasts (HLF) and 3T3 cells were first cultured in T25 tissue culture flasks with DMEM (Thermofisher Scientific) supplemented with 10% (v/v) fetal bovine serum (FBS) and 1% (v/v) Penicillin/Streptomycin mix, respectively. The disc-shaped hydrogels were used for 2D cell culture. For cell adhesion, an RGD-containing peptide sequence (RGD, Appendix A.8) was constructed into (FL)8. The N-DC (FL)4-RGD-(FL)4-RGD hydrogels were washed with a large amount of 1x PBS and preconditioned with Dulbecco's Modified Eagle's medium (DMEM). For cell spreading experiments, HLF and 3T3 cells were seeded respectively on top of the pre-treated D-NC gels with 60% confluence in each well of a 6-well plate at 37 ℃ for 24 hours. 101 A LIVE/DEAD Viability Kit for Mammalian Cells (Thermofisher Scientific) was used to ascertain the viability of cells on hydrogels. LIVE/DEAD Assay staining solution was prepared with 2 μM calcein acetoxymethyl and 4 μM ethidium homodimer-1 in sterile 1x PBS. Culturing medium was removed from the cells. Hydrogel samples were then incubated in the staining solution for 30 mins at room temperature. After staining, hydrogels were washed with 1x PBS 3 times. Then, the hydrogels were imaged using an Olympus Fluorescence Microscope at a magnification of x10. Live cells were captured with GFP filter and colored green in images (Ex/Em: 494/517 nm). Dead cells were imaged with RFP and colored red (Ex/Em: 528/ 617 nm). Figure 4-2 Schematics of the NC-(FL)8 hydrogels. A) The elastomeric protein (FL)8 was first dissolved in PBS at a high concentration (~200 mg/mL). B) Upon photochemical crosslinking, (FL)8 were crosslinked into a hydrogel network without chain entanglements, due to the short length of folded (FL)8, resulting in the N-NC hydrogel. C) When denatured in GdHCl, the (FL)8 in the hydrogel network unfolded, and behaved as random coils. The resultant D-NC hydrogel is also free of chain entanglement. 4.4 Results and discussion 4.4.1 A denaturing-crosslinking strategy reinforces globular protein-based hydrogels To construct (FL)8 hydrogels, the well-developed [Ru(bpy)3]2+-mediated photochemical crosslinking strategy was employed. Two tyrosine residues in proximity can crosslink into a 102 dityrosine adduct.29, 275, 296, 360 We used the DC (denatured crosslinking) method to construct the denatured (FL)8 hydrogels, in which the unfolded (FL)8 was photochemically crosslinked into hydrogels directly from its concentrated solution (20%, 200 mg/mL) in 7 M guanidinium hydrochloride (GdHCl). The as-prepared hydrogel was then equilibrated in 7 M GdHCl to obtain the denatured DC hydrogel (referred to as the D-DC hydrogel). As a control, we also constructed a denatured (FL)8 hydrogel that is free of chain entanglement using the NC (native crosslinking) method (Figure 4-2). In this method, native (FL)8 solution (200 mg/mL) in phosphate buffered saline (PBS) was first crosslinked into a hydrogel, which was free of chain entanglement due to the short length of native (FL)8. Then the prepared hydrogel was denatured and equilibrated in 7M GdHCl to obtain the denatured NC hydrogel (referred to as the D-NC hydrogel). Figure 4-3A shows the photographs of both D-DC and D-NC (FL)8 hydrogels prepared using the same ring-shaped mold as well as their stress-strain curves. The D-DC hydrogel was self-standing and swelled to a lesser degree than the D-NC hydrogel, while the D-NC hydrogel ring-shaped sample collapsed onto itself. The D-DC hydrogel displayed a Young’s modulus of 56 kPa, significantly higher than that of D-NC hydrogel (~1 kPa). 103 Figure 4-3 Physical entanglements enhanced the stiffness of the (FL)8 hydrogels. A) Stress-strain curves of D-DC (red) and D-NC (blue) (FL)8 hydrogels (200 mg/mL). Inset is the zoomed-in view of the stress-strain curve of the D-NC hydrogel. The D-DC hydrogel showed a Young’s modulus of ~50 kPa, significantly higher than that of D-NC hydrogel (~1 kPa). Top panel shows the photographs of both hydrogels after being equilibrated in 7M GdHCl. The D-DC hydrogel is self-standing and swells to a lesser degree than the D-NC hydrogel. B) Stress-strain curves of N-DC (red) and N-NC (blue) (FL)8 hydrogels (200 mg/mL). Inset is the zoomed-in view of the stress-strain curve of the N-NC hydrogel. The N-DC hydrogel showed a Young’s modulus of ~0.7 MPa, significantly higher than that of N-NC hydrogel (~20 kPa). The N-DC hydrogel ruptured at ~100% strain. Top panel shows the photographs of both hydrogels equilibrated in PBS. The N-DC hydrogel is translucent, while the N-NC hydrogel is opaque. 104 Figure 4-4 Fluorescence spectra of (FL)8 hydrogels. A) Fluorescence spectra of acid hydrolyzed 20% N-DC and N-NC (FL)8 hydrogels prepared from the same weight of lyophilized (FL)8 proteins. The fluorescence signal at 410 nm arises from the dityrosine fluorescence. B) Fluorescence spectrum of IAEDANS labeled 20% (FL-M23C)8 hydrogel. Dotted lines are Gaussian fits to the two fluorescence peaks, one is the dityrosine fluorescence at 410 nm, and the other one is the IAEDANS fluorescence at 490 nm. According to the classical rubber elasticity theory (Equation 2-1),207 G=ρRT/-.////, where G is the shear modulus, ρ is the density of the polymer in the network, -.//// is the average molecular weight between two neighboring crosslinking points, R is the gas constant and T is temperature. The stiffness (Young’s modulus) of a hydrogel network is determined from the crosslinking density (-.////). Thus, our results suggested a much higher effective crosslinking density for the D-105 DC gel than the D-NC gel. To examine if these two preparation methods led to a different number of dityrosine crosslinking points in the (FL)8 hydrogel, we quantified the number of dityrosine adducts by measuring the characteristic dityrosine fluorescence (Figure 4-4A).231, 275 Our results showed that both hydrogels contained roughly the same number of dityrosine crosslinking points (~17% of the total number of tyrosine residues in FL domains were cross-linked into dityrosine adducts). With the same number of crosslinking points and the same unfolded state of the polypeptide chain, the higher chemical crosslinking density of the swollen D-DC hydrogel should originate from additional effective crosslinking points resulting from the chain entanglements of the unfolded (FL)8 polypeptide chains in the D-DC hydrogel network. Evidently, the chain entanglement significantly enhanced the stiffness of the denatured (FL)8 hydrogels. Figure 4-5 DC (FL)8 hydrogels can be cycled between its N-DC and D-DC states reversibly. The swelling ratio (r) is relative to the initial weight of freshly prepared hydrogel (Section 4.3.3). Error bars indicate standard deviation. For each data point, n=4. Chain entanglement in a chemically crosslinked hydrogel network cannot be removed without disrupting the chemical crosslinking network. Thus, chain entanglement will be retained in the hydrogel network even if the polypeptide chains undergo conformational changes, such as folding. Taking advantage of this unique topological property, we “renatured” the D-DC hydrogel in PBS buffer to allow the refolding of FL domain to obtain the native DC (N-DC) (FL)8 hydrogel. 106 Figure 4-6 Scanning electron microscopy images of the N-DC (A) and N-NC (B) (FL)8 hydrogels. Both hydrogels showed porous network structures, however, the pore size of the N-DC hydrogel (~2 μm) is significantly smaller than that of the N-NC hydrogel (~20 μm). We observed that after renaturing in PBS, the N-DC hydrogel shrunk dramatically compared with the D-DC hydrogel, and changed from being transparent to largely translucent (Figure 4-3B). In contrast, upon renaturation, the N-NC hydrogel shrunk and became opaque. The swelling ratio of the N-DC hydrogel was smaller than that of the N-NC hydrogel. Moreover, both hydrogels can be cycled between their native and denatured states (N-DC to D-DC, N-NC to D-NC) for many cycles without noticeable change in their respective physical appearances and properties (Figure 4-5). The shrinkage is likely due to the refolding of some FL domains in PBS, which is accompanied by a significant shortening of the contour length of the polyprotein (from 260 nm to 10 nm). Microscopically, both N-DC and N-NC hydrogels showed microporous structures, but the mesh size of the N-DC hydrogel was significantly smaller than that of the N-NC hydrogel, consistent with the smaller swelling ratio of the N-DC hydrogel (Figure 4-6). 107 Figure 4-7 Schematics of the chain entangled network structure of the N-DC (FL)8 hydrogel. A-D) Schematics of the preparation of D-DC and N-DC (FL)8 hydrogels. The elastomeric protein (FL)8 was first dissolved in PBS at high concentration (~200 mg/mL) (A). When denatured in GdHCl, the unfolded (FL)8 polypeptide chain behaved as random coils, which overlapped with one another due to the high protein concentration, leading to possible chain entanglements (B). Upon photochemical crosslinking, chain entanglements were retained or created, leading to a network of entangled polypeptide chains, in the D-DC hydrogel (C). Entangled chains are highlighted in dashed squares. Lower left bottom shows a zoomed-in view of one such chain entanglement. When the D-DC hydrogel was equilibrated in PBS, some FL domains refolded, while others underwent hydrophobic collapse. In the renatured hydrogel (N-DC hydrogel), the chain entanglements remained, which are highlighted in dashed squares (D). For clarity, different molecules in B-D) are colored differently. E-F) Photographs of (FL-M23C)8 N-DC hydrogels. E) a (FL-M23C)8 N-DC hydrogel under UV illumination light. The blue fluorescence was from the dityrosine crosslinking points; F) a (FL-M23C)8 N-DC hydrogel under UV-illumination after labeling with IAEDANS. The cyan fluorescence was from the labeling of the exposed cysteine residues, and indicated that some FL domains were unfolded in the hydrogel. 108 During renaturation, due to the presence of chain entanglement and the shortening of the polypeptide chain, it is expected that some FL domains would not be able to fold into their native state. Instead, they likely assumed a hydrophobically collapsed state and/or random coil state, as PBS is a poor solvent for an unfolded FL domain. Therefore, the N-DC (FL)8 hydrogel network assumed a single network structure consisting of folded FL domains as well as unfolded ones. Figure 4-7 shows the schematics of the N-DC (FL)8 hydrogel network structure. Figure 4-8 Typical tensile stress-strain curves of 20% N-DC (FM23C)8 hydrogels. The mechanical properties of (FM23C)8 hydrogels are similar to those of (FL)8 hydrogels. The Young’s modulus is 0.89 MPa. Inset shows a photograph of an N-DC (FM23C)8 hydrogel. To verify the existence of the unfolded FL domains in the N-DC hydrogel, we used a well-established cysteine shotgun labeling approach, which allows for labeling of only solvent-exposed cysteine residues using the thiol reactive fluorescent dye IAEDANS.361 For this, we used a cysteine variant FL-M23C, where the buried residue Met23 in FL was mutated to Cys. Cys23 is sequestered in the hydrophobic core of the folded FL and can only be labeled with IAEDANS when FL-M23C is unfolded.231 The (FLM23C)8-based hydrogels showed similar physical and mechanical properties as (FL)8 (Figure 4-8). After IAEDANS labeling in PBS, the N-DC hydrogel showed the characteristic cyan fluorescence of IAEDANS under UV illumination (Figure 4-7E-F and Figure 109 4-4B). Quantitative analysis showed that ~50% of the FL domains were unfolded in the N-DC hydrogel. In contrast, only ~23% of the FL domains were unfolded in the N-NC hydrogel.231 These results indicated that the N-DC hydrogel is indeed a single network hydrogel consisting of folded and unfolded FL domains. Figure 4-9 Mechanical properties of N-DC and D-DC (FL)8 hydrogels in tensile testing. A) Young’s modulus and breaking strain of N-DC and D-DC (FL)8 hydrogels. B) Toughness and swelling ratio of N-DC and D-DC (FL)8 hydrogels. It is evident that the N-DC hydrogel exhibited much higher Young’s modulus and higher toughness than the N-NC hydrogel. 4.4.2 Tough N-DC (FL)8 hydrogels display excellent tensile properties We then carried out tensile and compression tests to measure the mechanical properties of the N-DC (FL)8 hydrogel. A typical stress-strain curve of the N-DC hydrogel in PBS buffer is shown in Figure 4-3B, where the N-DC hydrogel was stretched to failure. For comparison, the stress-strain curve of the N-NC hydrogel is also shown. It is evident that the stiffness of the N-DC hydrogel increased dramatically compared with that of the N-NC and D-DC hydrogels. The N-DC 110 hydrogel displayed a Young’s modulus of 0.7 ± 0.1 MPa (n=18) (Figure 4-3B and Figure 4-9A), significantly higher than that of protein-based hydrogels reported in the literature.29, 33, 264, 275, 362 The breaking strain of the N-DC hydrogel is 107 ± 14%, indicative of its good stretchability (Figure 4-9A). Similar Young’s modulus and breaking strain were also observed for 15% N-DC (FL)8 hydrogel as well as N-DC hydrogels constructed from (FL)16. However, 10% N-DC (FL)8 hydrogel showed similar Young’s modulus but much smaller breaking strain (50%) (Figure 4-10B). Figure 4-10 Mechanical properties of N-DC (FL)8 and (FL)16 hydrogels at different protein concentrations. A-C) Young’s modulus, breaking strain and toughness were measured from tensile tests on ring-shaped samples. D) The swelling ratio (r) is relative to the initial weight of freshly prepared hydrogel (Section 4.3.3). Error bars indicate standard deviation. For each data point, n=5. Moreover, during cyclic stretching-relaxation experiments, the N-DC hydrogel exhibited a large hysteresis, indicative of a significant amount of energy dissipation at high strains during stretching (Figure 4-11A). The average energy dissipation of the N-DC hydrogel is 250 ± 68 kJ/m3, demonstrating its superb toughness (Figure 4-9). Moreover, the energy dissipation of the N-DC (FL)8 hydrogel was fully reversible, and the hysteresis could be rapidly recovered once the 111 hydrogel was relaxed to zero strain (Figure 4-11B). After being stretched to 60% strain and then relaxed to zero, 70% of the hysteresis was recovered right after the hydrogel was relaxed back to zero strain, and the rest of the 30% hysteresis was recovered more slowly following a double-exponential kinetics, with a rate constant k1 of 0.05 ± 0.02 s-1 and k2 of (1.7 ± 0.3)×10-3 s-1, respectively (Figure 4-11C). This reversible hysteresis is likely due to the forced-unfolding, and subsequent refolding of the FL domains in the hydrogel network, consistent with the fast refolding rate of FL domain at zero force.34 Evidently, the N-DC (FL)8 hydrogels displayed tensile mechanical properties that uniquely combined a high Young’s modulus (~0.7 MPa), high toughness as well as fast recovery, a combination that is difficult to achieve as these individual properties are often mutually incompatible. Additionally, the high Young’s modulus and toughness of this single network protein hydrogel are amongst the highest of reported engineered protein hydrogels, and even compare favorably with those of some specially designed synthetic polymer hydrogels with unique network structures, such as double network hydrogels,241-242 or polymer composite hydrogels.354 112 Figure 4-11 The N-DC (FL)8 hydrogels can dissipate a large amount of energy. A) Stretching-relaxation stress-strain curves of the N-DC (FL)8 hydrogel. A large hysteresis was present in the stretching and relaxation curves, indicative of large energy dissipation. B) The hysteresis between stretching and relaxation curves can be recovered rapidly. The hydrogel was first stretched to ~60% strain and then relaxed to zero strain. After waiting for certain time Δt, the hydrogel was subject to the stretching-relaxation cycle again. The hysteresis recovery can be directly observed. C) The kinetics of the hysteresis recovery in N-DC (FL)8 hydrogel. ~70% of the hysteresis can be recovered rapidly within a few seconds, and the rest 30% hysteresis can be recovered following a double-exponential kinetics. The red line is a double exponential fit to the data, with a rate constant k1 of 0.05 ± 0.02 s-1 and k2 of (1.7 ± 0.3)×10-3 s-1, respectively. 113 Figure 4-12 N-DC (FL)8 hydrogels displayed superb compressive mechanical properties. A) Photographs show that the N-DC (FL)8 hydrogel can resist cutting with a sharp scalpel. B) Compressive stress-strain curves of the N-DC (red) and N-NC (FL)8 (blue) hydrogels. Inset is a zoom view of the stress-strain curves of N-NC hydrogel. The N-DC hydrogel can be compressed to more than 80% strain and sustain a compressive stress of >70 MPa without failure. A large hysteresis was present between the loading and unloading curves, indicating that a large amount of energy was dissipated. C) Consecutive compression-unloading curve of the N-DC hydrogel. The hysteresis grows with the increasing of the strain. The toughness of the hydrogel is ~3.2 MJ/m3. The inset is a zoom view of the stress-strain curves at lower strain. D) Consecutive compression-unloading cycles show that the hysteresis of the N-DC hydrogel can be recovered rapidly. Inset shows the hysteresis recovery kinetics of the hydrogel. ~65% hysteresis can be recovered right after unloading, and the rest hysteresis can be recovered following a double exponential kinetics, with k1 of 0.10 ± 0.02 s-1 and k2 of (2.0 ± 0.3)×10-3 s-1. E) Consecutive loading-unloading curves of the N-DC (FL)8 hydrogel at a frequency of ~0.37 Hz. The pulling speed was 100 mm/min. In each cycle, the hydrogel was stretched to 60% strain and subsequently relaxed to zero strain. After 100 cycles, the hydrogel displayed little fatigue, and the stress of the hydrogel at 60% strain retained ~90% of the original stress in the first cycle. 114 4.4.3 Remarkable compressive performance of N-DC hydrogels In addition to these superb tensile properties, the N-DC (FL)8 hydrogel demonstrated even more striking compressive mechanical properties. We found that the N-DC (FL)8 hydrogel is super tough and can resist slicing with a sharp scalpel, despite the fact that it contains ~60% water (Figure 4-12A). To quantify the compressive mechanical properties of the N-DC (FL)8 hydrogel, standard compression tests were carried out (Figure 4-12B). The stress-strain curves showed that the N-DC (FL)8 hydrogels displayed a compressive modulus of ~1.7 MPa at 10-20% strain. In comparison, the N-NC (FL)8 hydrogels only showed a compressive modulus of ~50 kPa (Figure 4-12B, inset), again revealing the significant enhancement effect of chain entanglement on the stiffness of the (FL)8 hydrogels. As the strain was increased to 75%, the stress of the N-DC (FL)8 hydrogel showed a much more pronounced increase of the modulus and reached a value of ~10 MPa. Strikingly, the N-DC FL hydrogel could be compressed to more than 80% strain and sustain a compressive stress as high as 75 MPa without fracture (Figure 4-13A). The average compressive strength of the N-DC (FL)8 hydrogel is 68 ± 12 MPa (n=7), and the average failure strain of 82 ± 3%. At the failure strain, a small crack often started to appear in the hydrogel sample. However, the failure was not brittle, as evidenced by the subsequent stress-strain curves of the same hydrogel sample (Figure 4-13B). These results demonstrated the super high compressive strength of the N-DC (FL)8 hydrogel. To the best of our knowledge, the compressive strength of the N-DC (FL)8 hydrogel is the highest among the hydrogels reported in the literature, and compares favorably with that of articular cartilage.345, 348, 353 As a comparison, the super tough double network polymer hydrogels typically fractured at a stress of no more than 20 MPa.241-242 115 Figure 4-13 Compression tests and stress-strain curves of a N-DC (FL)8 hydrogel compressed to failure. A) Photographs of the N-DC (FL)8 hydrogel under compression. After unloading, the hydrogel recovered its shape rapidly. B) Cracks were observed right after the failure. Subsequent compression led to the propagation of the crack. Inset shows the photographs of the hydrogel right after failure (1st cycle) and after three more consecutive compression-unloading cycles (4th cycle). In the compression-unloading cycles, a large hysteresis was observed (Figure 4-12C), indicative of a large amount of energy being dissipated during compression. The hysteresis increased as the strain increased. At 80% strain, the hydrogel exhibited a toughness of 3.2 ± 0.6 MJ/m3 (n=7). Moreover, the (FL)8 hydrogel displayed superb recovery properties (Figure 4-12D). At lower strain (<40%), the hydrogel recovered its original dimensions right at the end of the 116 loading-unloading cycles, suggesting that the recovery of the hydrogel is fast. Even at a large compression strain (>60%), ~65% of the hysteresis can be recovered right after unloading. The remaining 35% hysteresis could be recovered within an hour following double exponential kinetics (Figure 4-12D). Furthermore, after 100 consecutive loading-unloading cycles at a frequency of 0.37 Hz and a strain of 50%, the consecutive loading-unloading curves did not show much change, and the stress of the N-DC (FL)8 hydrogel at 60% strain retained ~90% of its original stress (Figure 4-12E). Similar results were obtained at a frequency of 0.67 Hz and 0.08 Hz (Figure 4-14), demonstrating that the N-DC (FL)8 hydrogel can quickly recover its mechanical properties and did not show much fatigue. Figure 4-14 Consecutive compression-unloading curves of a N-DC (FL)8 hydrogel at a frequency of 0.08 Hz (A) and 0.67 Hz (B). The loading rate was 20 mm/min in A) and 200 mm/min in B), respectively. Collectively, these results revealed that the N-DC (FL)8 hydrogels are mechanically strong and tough, and can recover their shape and mechanical properties rapidly and do not show much mechanical fatigue. Moreover, these protein hydrogels showed excellent long term stability: after being stored in PBS (with 0.2‰ NaN3) for over six months, their physical shape and mechanical 117 properties remained largely unchanged. These exceptional mechanical properties and their unique integration in one material are rare for protein hydrogels, and compare favorably with those of polymer hydrogels with special network structure (Table 4-1). 118 Table 4-1 Mechanical properties of the state-of-the-art protein hydrogels and polymer hydrogels. Tensile Compression Young’s modulus Breaking stress Breaking strain Compressive modulus Breaking stress Breaking strain Reference Cartilage Articular cartilage 1-10 MPa 1-20 MPa 0.5-10 MPa 10-50 MPa 342, 347-348 Current work (FL)8 0.7 MPa 0.6 MPa 107% 1.7 MPa 68 MPa 82% Protein Hydrogels (GRG5RG4R 50 kPa 70 kPa 100% 29 (FL)8 16 kPa 35 kPa 450% 34 Bovine fibrinogen 70 kPa 46 kPa 62% 298 Gelatin 70 kPa 175 kPa 0.5 MPa 363 Gelatin 67 kPa 91 kPa 145% 164 Regenerated silk fibroin 50 kPa 20 kPa 362 Gel1 crosslinked with 4-arm PEG 10-150 kPa 20 kPa 16% 33 (GB1)8-RX3 65 kPa 110 kPa 80% 231 Regenerated silk fibroin (alcohol treated) 2 MPa 0.7 MPa 110% 0.3-3 MPa 0.5-1.7 MPa ~60% 134 Regenerated silk fibroin (alcohol treated double network) 5.8 MPa 12.4 MPa 80% 364 Polymer hydrogels Double network PAMPS-PMAAm 0.33 MPa 17 MPa 92% 242 Bacterial cellulose-PDMAAm 1.6 MPa 2.9 MPa 50% 243 PAMPS-collagen 2.9 MPa 53% 242 Cojoined network Chitosan-gelatin-phytate 2.5 MPa 4.3 MPa 350% 6.6 MPa 64 MPa 87% 365 Slide-ring network NIPA–AAcNa–HPR-CD 43 kPa 40 kPa 450% 366 119 Additionally, we used representative cell lines derived from mice and human to test the cytocompatibility of the N-DC gels. The cell culture on the N-DC (FL)4-RGD-(FL)4-RGD gels showed excellent cell viability results (Figure 4-15). These properties closely reproduced many features of articular cartilage, and thus make the N-DC (FL)8 hydrogels a cartilage-mimetic protein biomaterial. Figure 4-15 Cell viability tests on N-DC hydrogels. Live/Dead staining after cell culture on 20 wt% N-DC (FL)4-RGD-(FL)4-RDG hydrogels for 48 hrs. Live cells are in green and dead cells are in red. A) Human lung fibroblasts cell culture showed cell viability of 94.5 ± 5.2% on the gel. B) The cell viability of 3T3 cells culture on the gel is 98.7 ± 7.3%. These superb mechanical properties of N-DC (FL)8 hydrogels are likely the result of a combination of factors, including chain entanglement, folded and hydrophobically collapsed FL domains in the hydrogel network, as well as the forced unfolding and refolding of FL domains. These unique features likely make this DC method for protein hydrogelation broadly applicable. Indeed, as shown in Table 4-2, in protein hydrogels constructed from a range of elastomeric proteins, which range from all α proteins to α/β proteins, we significantly enhanced their stiffness 120 via this DC hydrogelation method and improved their Young’s modulus to several hundred kPa, depending on the inherent tyrosine content of the elastomeric protein. Similar enhancement was also achieved in the compressive modulus of these protein hydrogels (Table 4-2). These results demonstrate the generality of this novel method. Table 4-2 Enhancement of mechanical properties via the DC hydrogelation method. (NuG2)8, a polyprotein made of eight tandem repeats of the protein NuG2.357 (GB1)8, a polyprotein made of eight tandem repeats of the protein GB1.283 GRG5RG4R,G represents GB1 domain, R represents the 15 residue consensus sequence of resilin.29. NRN4RN4R, N represent NuG2 domain, R represents the 15 residue consensus sequence of resilin. BSA: bovine serum albumin.176. (GA)8, a polyprotein made of eight tandem repeats of the protein GA.358 E: tensile modulus; Y: compressive modulus. Tensile Compression N-DC hydrogel N-NC hydrogel N-DC hydrogel N-NC hydrogel Protein EN-DC (kPa) EN-NC (kPa) Enhancement EN-DC/EN-NC YN-DC (kPa) YN-NC (kPa) Enhancement YN-DC/YN-NC (FL)8 700 ± 110 16 ± 2 44 1614 ± 201 40 ± 3 40 (NuG2)8 620 ± 93 11± 3 56 1086 ± 31 55 ± 11 20 (GB1)8 227 ± 30 5 ± 1 47 339 ± 36 68 ± 11 5.0 GRG5RG4R 123 ± 53 33 ± 4 3.8 220 ± 31 20 ± 4 11 N4RN4RNR 126 ± 21 22 ± 3 5.7 592 ± 148 114 ± 25 5.2 BSA 236 ± 5 7 ± 0 36 776 ± 173 20 ± 4 40 (GA)8 438 ± 55 60 ± 9 7.3 1704 ± 248 95 ± 17 18 4.5 Conclusions Stiff biological tissues, such as cartilage, tendons and ligaments, often integrate seemingly mutually incompatible mechanical properties into themselves.342 Mimicking such properties in synthetic hydrogels has been challenging because optimizing one property often comes at the expense of another. To alleviate this issue and achieve high stiffness and high toughness, polymer 121 hydrogels of designed network structures and polymer composite hydrogels have been developed,241-242, 365-370 such as double network hydrogels,241-242 conjoined network hydrogels365 and slide-ring hydrogels366-367. Sacrificial bonds or weak secondary networks that can be ruptured are often introduced into the hydrogel as an energy dissipation mechanism.241, 371-372 Although high stiffness and high toughness have been achieved in some of these hydrogels, slow recovery and mechanical fatigue are often present, due to the slow dynamics of these sacrificial bonds/secondary networks. The achieved compressive strength also remained lower than that of articular cartilage. Although proteins are attractive building blocks to construct soft biomaterials, it has been challenging to construct stiff protein-based biomaterials to mimic stiff tissues. To learn from biology, here we demonstrated a novel approach that combines forced-unfolding of globular proteins and chain entanglement to enable the engineering of strong and tough protein hydrogels. On the one hand, forced-unfolding of globular proteins provides an efficient mechanism for energy dissipation, and the ability to refold allows the hydrogel to recover its mechanical properties rapidly and minimize mechanical fatigue. On the other hand, we are the first to introduce the denaturing-crosslinking gelation method, where chain entanglement allows globular protein-based hydrogels to achieve high stiffness. In so doing, our results demonstrated that the engineered protein hydrogels successfully integrate high stiffness, high toughness, ultrahigh compressive strength and fast recovery into one protein biomaterial, thus converting a muscle-like soft biomaterial to a stiff material exhibiting mechanical properties that mimic cartilage. Although the engineered protein hydrogel has a single network structure, the chain entanglement effectively increased the crosslinking density and imparted high stiffness to the hydrogel network. At the same time, the chain entanglement prevented some proteins to fold, providing a mechanism to lessen the rigidifying effect on the network due to protein folding, which often leads to the brittleness of 122 the network. The incorporated mechanically labile domains allowed the energy to be dissipated effectively in the network and prevented stress concentration. The fast folding of the protein domains can then enable the hydrogel to recover rapidly upon relaxation. These effects work cooperatively to allow the integration of high stiffness, high toughness, fast recovery and high compressive strength into protein hydrogels. To the best of our knowledge, the N-DC (FL)8 hydrogel displays the highest toughness and stiffness among the globular protein-based hydrogels that have been reported. By expanding the range of mechanical properties that protein hydrogels can achieve, our results demonstrate the feasibility to engineering protein-based hydrogels to mimic stiff tissues like cartilage, and may open up a new promising avenue for developing novel protein-based biomaterials for applications in fields ranging from cartilage repair and engineering to soft robotics and actuators. 123 Chapter 5: Super Tough and Strong Protein-Alginate Double-network Hydrogel 5.1 Synopsis Globular protein-based hydrogels are promising biomaterials. Their biomedical applications are often limited by their relatively poor mechanical properties. Our previous work demonstrated a denatured-crosslinking-renaturation (N-DC) method via which the mechanical properties of a protein hydrogel made of (FL)8 were significantly improved (Young’s modulus increased from 20 kPa to 0.7 MPa). Following this strategy, here we synthesized a double-network hydrogel with the assistance of an ionically crosslinked network to further improve the mechanical properties of the (FL)8 gels. A biocompatible natural polymer alginate was used to fulfill this purpose. To enable alginate to covalently bond with the protein network, the alginate was conjugated with tyramine using aqueous-phase carbodiimide activation chemistry. When (FL)8 and tyramine-derivatized alginate were photochemically crosslinked into conjoined hydrogels, the hybrid gels showed much improved and tunable Young’s modulus (up to ~1.0 MPa), extensibility (up to 230%), and toughness (up to ~1.02 MJ/m3), which are much higher than the mechanical properties of the individual network components. The superb mechanical properties of this novel network design expand the range of characteristics that can be achieved by protein-containing materials, providing great potential for such protein/alginate hybrid hydrogels as biocompatible materials in the biomedical field. 5.2 Introduction The articular cartilage is a resilient elastic tissue that covers and protects the joints in the human body. However, articular cartilage generally shows limited self-repairing ability.373 As such, 124 articular cartilage injuries and joint diseases are a significant health issue, affecting millions of people annually worldwide.374-375 Although the current therapies are focused on regeneration of the chondrocytes and surrounding matrix transplantation,376-379 it still remains challenging to engineer a biomechanically suitable scaffold for cartilage regeneration studies, due to the load-bearing nature and superb toughness of cartilage.380 Protein-based biomaterials are one of the promising candidates for tissue engineering and regenerative medicine. However, the use of protein-based biomaterials for artificial cartilage engineering is limited, due to their relatively poor mechanical properties. Our previous study reported a cartilage-like hydrogel made of an engineered artificial elastomeric protein (FL)8. By crosslinking (FL)8 from its unfolded state, we were able to introduce chain entanglement into the hydrogel network. Upon renaturation, the entangled and refolded protein chains significantly enhanced the mechanical properties of the protein hydrogel. Such an entirely protein-based hydrogel showed a tensile modulus of ~0.7 MPa and toughness of 250 kJ/m3, properties that are within the range of cartilage tissues (stiffness ranging from 0.5 MPa to several MPa, and toughness of 83-141 kJ/m3).342-348, 381-382 Comparing to the native (FL)8 hydrogel, the denatured-crosslinking method effectively enhances the tensile modulus (FL)8 by a factor of 35 (from ~20 kPa to ~0.7 MPa). However, the extensibility of the (FL)8 hydrogels decreased from ~350% to ~107%. In this work, we aimed at further improving toughness and elasticity of such a hydrogel, providing a wide range of mechanical strength for cartilage-like biomaterials. 125 Figure 5-1 Schematic design of the Alginate-(FL)8 double-network hydrogel. A) A physically crosslinked alginate network via ionic bonding. B) A natural alginate polymer chain consists of mannuronate and guluronate residues, in which guluronic acid can form ionic interactions in the presence of calcium ions. C) Alginate-tyramine coupling. By using carbodiimide activation chemistry, carboxyl groups of alginate can form covalent conjugation with tyramine. D) The native and chemical denatured (FL)8 protein networks. E) The protein network and the alginate network form dityrosine-like covalent coupling by using Ru2+-catalyzed photo-chemistry. The denatured-crosslinking-renaturation approach is used as described in Chapter 4. 126 Introducing energy dissipative mechanisms can enhance the toughness of a rubber-like network. For example, a synthetic polymer hydrogel with extremely high fracture strength was achieved by using a double-network strategy, in which a long-chain network is interpenetrated within another short-chain network.241-242 The long-chain network is responsible for the stiffness, while the short-chain component ruptures and dissipates energy upon stretching. However, the rupture of the short-chain network causes unrecoverable damage. To overcome this drawback, a secondary network with recoverable energy dissipative mechanisms, which rely on non-covalent bonding, is preferred.183, 232, 370 Alginate is a natural polysaccharide extracted from the algal cell wall. Due to its abundance, low cost and biocompatibility,383-384 alginate is widely used in biomedical applications. Alginate can rapidly form an ionically crosslinked hydrogel in the presence of calcium ions, which makes alginate a suitable candidate as an energy dissipative network (Figure 5-1A and B). The Suo group reported a highly stretchable and tough alginate-polyacrylamide double-network hydrogel.371 The resultant hydrogel showed significant improvement in both stiffness and extensibility, compared to the single-network alginate or polyacrylamide hydrogel. The enhanced toughness was attributed to the reversible unzipping-rezipping of ionic crosslinks provided by the alginate network. Inspired by this prior work, we employed the double-network approach to combine a strong covalent network with a weak physically crosslinked one. We report the engineering of a protein-alginate hybrid hydrogel by forming both ionically and covalently cross-linked networks. We reason that the additional ionic network of alginate could serve as an extensible component. On the one hand, alginate has a large molecular weight (~90 to 190 kDa). It is likely that the alginate chains could form physical entanglements and decrease network defects. On the other hand, alginate introduces ionic crosslinking in the presence of calcium ions, serving as energy dissipative 127 units. To connect the two networks with controllable crosslinking density, we derivatized alginate with phenolic groups (Figure 5-1C).259 The tyrosine residue of protein building blocks and the phenol groups from the modified alginate could undergo Ru2+-catalyzed photo-chemistry to form a covalently crosslinked conjoined double-network (Figure 5-1E).34, 296 Benefiting from the “denatured-crosslinking” method and the additional ionic crosslinking, we successfully toughened the single-network (FL)8 gel with a tunable range of mechanical behaviors. 5.3 Materials and methods 5.3.1 Protein engineering The constructs and the lyophilized polyproteins of (FL)8, (NuG2)8, and (GB1)8 were obtained using protein engineering as described in Chapter 4 (4.3.1). Protein expression and purification were performed using the same procedure as described in Chapter 2 (2.3.1). 5.3.2 Alginate derivatization with phenol moieties Sodium alginate, tyramine hydrochloride, N-(3-dimethylaminopropyl)-N’-ethylcarbodiimide hydrochloride (EDC) and N-hydroxysuccinimide (NHS) were purchased from Sigma-Aldrich. Considering solubility and workable viscosity of pregel solutions, we chose sodium alginate with a molecular weight ~120,000-190,000 g/mol, and a ratio of mannuronic acid to guluronic acid residues of 1.56. Alginate with phenol groups was synthesized via conjugation of alginate and tyramine, using carbodiimide activation chemistry (Figure 5-1C).259 The sodium alginate was dissolved in 100 mM morpholinoethanesulfonic acid buffer (MES) to a concentration of 0.5 w/v% (pH 6.0). The alginate solution was mixed with EDC and NHS stirring at 25 ℃ for 1 hr. To achieve different functionalization density of phenol groups, different concentrations of EDC and NHS were used in the reaction (Table 5-1). Then tyramine hydrochloride was added into 128 the reaction mixture with an alginate/tyramine weight ratio of 1:9, gently stirring for 20 hrs at 25 ℃. After reaction, the reaction mixture was dialyzed against deionized water for 2 days to remove small molecule reactants and salts, and then lyophilized. Table 5-1 Concentration of alginate, EDC, NHS and tyramine for the synthesis of alginate-tyramine conjugates Samples Alginate w/v% EDC (mM) NHS (mM) Tyramine wt% Alg1 0.5 6.3 3.1 4.5 AlgL 0.5 12.6 6.3 4.5 AlgM 0.5 31.5 15.8 4.5 AlgH 0.5 63.0 31.5 4.5 Alg5 0.5 126.0 63.0 4.5 5.3.3 Characterization of phenolic density of the modified alginate. The degree of grafting of the derivatized alginate was determined by UV-Vis absorbance at 275 nm, using a Cary 4000 UV-Vis spectrophotometer. The unmodified alginate and the derivatized alginate samples (AlgL, AlgM, and AlgH) were dissolved in deionized water with a concentration of 0.5 wt%. To calculate the degree of grafting, a standard UV absorbance-concentration calibration curve was plotted using the absorbance of tyramine hydrochloride solutions (from 0.005 wt% to 0.05 wt% in deionized water), detected at 275 nm. 5.3.4 Preparation of alginate hydrogels and double-network hydrogels. After derivatization, the lyophilized alginate samples were re-dissolved into 100 mM tris(hydroxymethyl)aminomethane buffer (Tris) with a concentration of 1 w/v%. A physical alginate hydrogel was prepared by mixing the derivatized alginate solution with a 200 mM CaCl2 solution at an equal volume. The chemically crosslinked alginate and the double-network alginate-129 protein hydrogels were fabricated by following the Ru2+-mediated photochemical crosslinking protocol.34, 296 To prepare a photocrosslinked alginate gel, the redissolved alginate solution (1 w/v% in Tris) was mixed with APS and [Ru(II)(bpy)3]Cl2 at final concentrations of 50 mM and 260 μM respectively. To prepare a protein-alginate double-network hydrogel, 20 mg of the lyophilized polyprotein was dissolved in 100 μL 7 M GdHCl containing 1 w/v% derivatized alginate, 50 mM APS and 260 μM [Ru(II)(bpy)3]Cl2. The ring shaped hydrogels were prepared by transferring the pregel mix into a custom-made ring-shaped mold. The pregel solutions were then irradiated under a 200W fiber optic white light source from the height of 10 cm above the samples for 10 min. Then the hydrogels were stored in Tris buffer to reach swelling equilibrium. To achieve the secondary physical crosslinking networks, the hydrogels were transferred into a Tris solution containing 25 mM CaCl2. The hydrogel dimensions were measured before tensile tests. The hydrogel dimension and water content (w) were measured as described in Sections 4.3.2 and 4.3.3. 5.3.5 Tensile tests Tensile tests were performed on an Instron-5500R tensometer with a custom-made force gauge at constant temperature (25 ℃). The D-DC (denaturing-crosslinking-denatured), N-DC (denaturing-crosslinking-renatured), and I-N-DC (denaturing-crosslinking-renatured with Ca2+) ring-shaped samples were tested in Tris with 7M GdHCl, Tris, and 25 mM CaCl2 in Tris (pH 7.4), respectively. To remove Ca2+ from I-N-DC hydrogels, the samples were treated with 25 mM EDTA in Tris (pH7.4) for 4 hrs. An extension rate of 25 mm/min was used in our experiments. The stress at 15% strain was taken as the Young’s modulus of the sample. Toughness was determined by integrating stress-strain curves where specimens were loaded directly to failure. (Measured specimens, n=4) 130 5.3.6 Rheological tests All rheology measurements were performed on a TA Instruments Discovery HR-2 equipped with an 8 mm flat plate at 25 ℃. The gelation of the alginate samples (1 w/v%), (FL)8 (20 w/v%), and the D-DC (FL)8/Alg were measure by using oscillation-time sweep under fixed strain (1%) and frequency (10 Hz). The strain and frequency were chosen to be within the linear viscoelastic regime. The storage and loss modulus were recorded over a 1260 sec period, where the light was turned on at the 60th sec. An oscillation-strain sweep from 1% to 1000% followed, in which the frequency was fixed at 10 Hz. The breaking strains were measured from the onset points of strain-sweep curves, where the storage modulus started to dramatically decrease with the increasing of shear strain. (Measured specimens, n=4) 5.4 Results and discussion 5.4.1 Physical and chemical gelation of derivatized alginate Alginate is a linear block copolymer consisting of D-mannuronic acid (M block) and L-guluronic acid (G block). In the presence of divalent cations, the alginate solution immediately forms a hydrogel via ionic interaction between the carboxyl groups in the G blocks and these cations. To covalently interconnect alginate with proteins, we used Ru2+-mediated photocrosslinking chemistry, which allows the crosslinking of two phenolic groups into a dityrosine adduct. For this, we introduced phenolic groups into alginate by derivatizing alginate using tyramine via EDC chemistry. EDC first activates the carboxylic acid group in alginate to form an active ester intermediate, which is stabilized by NHS. The intermediate is easily replaced by the nucleophile from the amine group of tyramine, forming an amide bond with the original 131 carboxyl group. Three different modification densities of phenolic groups were prepared by tuning the molar ratios of EDC and NHS in the reaction (Table 5-1). The derivatized products were labeled as AlgL (alginate with low tyramine modification density), AlgM (alginate with medium tyramine modification density) and AlgH (alginate with high tyramine modification density). The density of the modifications was determined by UV-Vis. Compared to the unmodified alginate, the derivatized alginate samples showed a characteristic absorbance of phenol at 275 nm, indicating the successful conjugation of tyramine to alginate (Figure 5-2). The degree of grafting of AlgL, AlgM and AlgH, which were determined using UV-Vis spectroscopy, were 13.4%, 29.5% and 41.9% of the total carboxyl groups in alginate, respectively. Figure 5-2 UV-Vis spectra are used to determine phenolic density of the modified alginate. A) UV-Vis absorbance spectra of derivatized alginate samples and unmodified alginate (0.5 wt%). B) Standard UV absorbance-concentration curve of tyramine hydrochloride in deionized water. The absorbance of tyramine compound at 275 nm is linearly proportional to its concentration within the concentration range of 0.005 wt% to 0.03 wt%. This standard curve is used to determine the concentration of phenol groups in derivatized alginate. To achieve a double-network hydrogel, we carried out a Ru2+-catalyzed photochemical crosslinking reaction with the mixture of the derivatized alginate and polyproteins, in which the phenol groups on alginate reacted with the tyrosine residues in the polyproteins to form dityrosine-like adducts (Figure 5-1E). In this way, we could take advantage of the Ca2+ triggered physical 132 crosslinking to form the physical gel, via which the intermolecular physical crosslinks in alginate gels can break and reform reversibly during loading and relaxing, serving as a secondary recoverable energy-dissipating mechanism. Figure 5-3 Physical and chemical gelation of alginate with phenol moieties. A) A photograph of physically crosslinked AlgL, AlgM and AlgH hydrogels (1 wt%), obtained by mixing with Ca2+ cations. B) Time-sweep rheological characterization of alginate gelation (1 wt%). The light source was turned on at 60 secs to activate the photochemical crosslinking reaction, followed with increasing in storage modulus of derivatized alginate samples. The photo-gelation reached equilibrium after 10 mins. The chemically crosslinked AlgL, AlgM and AlgH showed storage moduli of 0.4 kPa, 2.8 kPa and 9.4 kPa, respectively. The storage modulus and loss modulus of the unmodified alginate remain unchanged with the identical reaction condition. C) Rheology measurement of strain-sweep of AlgL, AlgM and AlgH hydrogels. The onset point of each hydrogel sample indicates the breaking strain of the gels, where the storage modulus starts to dramatically decrease with the increasing of shear strain. The photo-crosslinked AlgL, AlgM and AlgH hydrogels displayed breaking strains of 77%, 105% and 30%, respectively. 133 Derivatization of alginate with tyramine consumes the carboxylic groups in both the G and M blocks of alginate. However, if a large portion of the carboxylic groups in the G blocks are consumed during derivatization, it is possible that the derivatized alginate may lose its ability to form a physically crosslinked hydrogel in response to Ca2+, as the ionic crosslinking of alginate is mainly through the coordination of Ca2+ by the G-block. To investigate this effect, we first examined if the derivatized alginate can form a Ca2+-triggered physical gel and a chemically crosslinked gel via Ru2+ catalyzed photocrosslinking chemistry. The physical gelation of the modified alginate was completed by simply exposing the alginate solution (1 wt%) to Ca2+. As shown in Figure 5-3A, all of the derivatized alginate solutions formed Ca2+-triggered physical hydrogels, indicating that the number of remaining unreacted carboxyl groups in the G blocks of alginate was sufficient to support physical crosslinking via ionic interactions in the presence of Ca2+. If tyramine is successfully conjugated with alginate, it is expected that the incorporated phenolic groups on the alginate chains can undergo Ru2+-mediated photochemical coupling, leading to the formation of covalently crosslinked hydrogel networks in addition to physical crosslinking. To test this possibility, we monitored the photochemical gelation of the derivatized alginate by using shear rheology. After the alginate was mixed with Ru2+ and APS, shear rheology experiments were carried out. At 60 s, the white light was turned on, and the derivatized alginate mixtures were exposed to illumination for 20 mins while the storage modulus (G’) and loss modulus (G’’) were monitored. Figure 5-3B shows the time-evolution of G’ and G” of the hydrogels at a fixed strain and frequency for 1% alginate samples. After exposure to the light, the storage moduli of AlgH, AlgM and AlgL indeed increased significantly, suggesting that these modified phenolic groups are highly reactive during the photochemical crosslinking reaction and 134 efficiently form cross-linking points. Upon reaching equilibrium, AlgL, AlgM and AlgH showed storage moduli of 0.4 kPa, 2.8 kPa and 9.4 kPa, respectively. In contrast, the mechanical properties (G’ and G”) of the physical gel formed by unmodified alginate remained unchanged as a function of time under the same white light illumination. Evidently, gel formation of the derivatized alginate demonstrated successful conjugation of alginate chains and tyramine. Moreover, the increase in gelling rates increased proportionally with the higher density of crosslinkable phenol groups: AlgH and AlgM showed faster increase in storage modulus during gelation as compared to AlgL. Moreover, the derivatization density of alginate was also indicated by the trend observed for the equilibrated stiffness (AlgH>AlgM>AlgL). The results obtained using amplitude sweeps also showed the longer linear viscoelastic region of AlgH and AlgM gels as compared to AlgL, indicating better physical stability of the AlgH and AlgM gels (Figure 5-3C) In addition, we found that when the modification density is lower than AlgL, the alginate could not form covalent gels (Alg1 in Table 5-1). And if the modification density higher than AlgH, the alginate could not form physical gels in the presence of Ca2+ either (Alg5). Thus, there is a range of grafting density of tyramine that will allow both physical and chemical crosslinking of alginate. Therefore, we chose AlgL, AlgH and AlgM to prepare the alginate-protein hybrid hydrogels. 135 5.4.2 Cooperation of alginate and polyproteins: mechanical improvement of a double-network hydrogel 5.4.2.1 Denatured-crosslinking (FL)8 covalently crosslinks with derivatized alginate (D-DC (FL)8/Alg hydrogel). Having confirmed the reactive physical and chemical crosslinking sites on the modified alginate chains, we attempted to employ the alginate to reinforce (FL)8 protein-based hydrogels, using the conjoined double-network strategy. The double-network hydrogels were prepared by mixing 1% derivatized alginate and 20% (FL)8 in 7 M GdHCl. After carrying out Ru2+-catalyzed photochemical crosslinking, the phenol groups from the modified alginate and the tyrosine residues of (FL)8 were coupled into dityrosine or dityrosine-like crosslinks, forming D-DC (denatured-crosslinking) (FL)8/Alg double-network gels. Figure 5-4 Mechanical properties of D-DC (FL)8/Alg double-network hydrogels. A)-B) The storage modulus and breaking strain of the derivatized alginate alone, N-NC and D-DC (FL)8 alone, and the D-DC (FL)8/Alg double-network gels (20 wt% (FL)8 and 1% modified alginate). The storage modulus was obtained from the equilibrated storage modulus in a time-sweep gelation test. The breaking strain was obtained from the strain of an onset point in the strain-sweep test of each hydrogel. Error bars indicate standard deviation. For each data point, n=6. We first carried out shear rheology experiments to investigate the mechanical properties of the D-DC (FL)8/Alg double-network gels, comparing to the gels made of photo-crosslinked 136 alginate alone and (FL)8 alone. As shown in Figure 5-4, the D-DC double-network gels showed a ~10-fold enhancement in storage modulus as compared to the single-network alginate. On the other hand, the storage modulus of the D-DC (FL)8/Alg gels stays in the similar range as the D-DC (FL)8 single-network gels. This can be explained by the fact that the storage modulus of the double network hydrogels was largely determined by the (FL)8 network (the concentration of (FL)8 is 20%, while alginate is 1%). Noticeably, the amplitude sweep of the double networks showed that the incorporation of alginate significantly enlarged the linear viscoelastic region of the hydrogels, compared to the individual components (Figure 5-4B). This result suggested that the double-network gels are more physically stable than single-network (FL)8 gels or alginate gels. 5.4.2.2 Renaturation of the denatured-crosslinking (FL)8 enhances double-network density (N-DC (FL)8/Alg). To further enhance the mechanical properties of the (FL)8/Alg double-network gels, we renatured the D-DC (FL)8/Alg hydrogel in Tris buffer to allow FL to refold. As the Tris buffer is a poor solvent for unfolded (FL)8, the refolding and hydrophobic effect of (FL)8 lead to contraction of the double network, increasing the network density, as well as resulting in significant shrinkage (Figure 5-5A). The renatured double-network gel is labelled as N-DC (FL)8/Alg. We first studied the tensile properties of the (FL)8/AlgM double network. After renaturation, the tensile modulus of N-DC (FL)8/AlgM gels showed an 8-fold increase from ~48 kPa to ~447 kPa. The breaking strain was increased from ~42% to ~95% (Figure 5-5B and Figure 5-6A). However, the stiffness of N-DC (FL)8/AlgM gels is lower than that of the single-network N-DC (FL)8 (Figure 5-5B and Figure 5-6B). It is likely that the alginate component and dityrosine-like crosslinks interfere with either the refolding process or the hydrophobic collapse of (FL)8, thus lowering the network density and tensile modulus of the N-DC double-network hydrogel. As a control, we also 137 characterized the mechanical properties of the double-network gel using 1% (wt%) unmodified alginate (unAlg) hydrogels (Figure 5-6C). The N-DC (FL)8/unAlg gels showed an average Young’s modulus of ~567 kPa and breaking strain of ~84%. The stiffness of N-DC (FL)8/unAlg gel was lower than that of the single-network of N-DC (FL)8. This result suggested that the presence of alginate chains in the hydrogel network may have an impeding effect on the renaturing process of (FL)8. The crosslinking of (FL)8 with AlgM may further restrain the folding of the protein network, leading to the observed lower modulus of the N-DC (FL)8/AlgM hydrogel than that of the N-DC (FL)8/unAlg hydrogel. 138 Figure 5-5 Mechanical performance and swelling behavior of (FL)8/Alg double network hydrogels. A) Water content of D-DC (FL)8/AlgM (green), N-DC (FL)8/AlgM (blue), and I-N-DC (FL)8/AlgM (red) double-network hydrogels. B) The Young’ modulus and breaking strain of N-DC (FL)8, D-DC (FL)8/AlgM, N-DC (FL)8/AlgM, and I-N-DC (FL)8/AlgM hydrogels. Compared to the N-DC (FL)8/AlgM gels. the ionical crosslinking significantly reinforced Young’s modulus and breaking strain of I-N-DC (FL)8/AlgM. C) The Young’s modulus and the breaking strain of I-N-DC (FL)8/Alg double-network hydrogels can be controlled by the modification density of tyramine on alginate. D) The representative tensile stress-strain curves of I-N-DC (FL)8/AlgH and N-DC (FL)8/AlgH hydrogel rings. E) The Young’s modulus of the N-DC (FL)8/AlgH hydrogels can be reversibly tuned between ~400 kPa and ~1 MPa by adding and removing Ca2. F) Young’s modulus of I-N-DC (FL)8/AlgM double-network hydrogel rings made of different concentration of (FL)8. 139 5.4.2.3 N-DC (FL)8/Alg in the presence of calcium ions further improved toughness of the double-network hydrogels (I-N-DC (FL)8/Alg). Owing to the remaining G blocks on the derivatized alginate chains, ionic crosslinks can be introduced into the double networks (FL)8/Alg hydrogels by exposing the hybrid gels to Ca2+. The G blocks of the derivatized alginate can form intermolecular ionic interactions with Ca2+, leading to the formation of the ionically crosslinked N-DC (FL)8/Alg hydrogels (I-N-DC). The mechanical properties of the I-N-DC (FL)8/Alg hydrogels were characterized by tensile tests. Figure 5-5C and Figure 5-6D show the tensile properties of the I-N-DC (FL)8/Alg double-network hydrogel rings. For comparison, the previously reported single-network N-DC (FL)8 gel (Chapter 4) is also shown. It is evident that the introduction of ionic crosslinking not only significantly increased the stiffness of the I-N-DC (FL)8/Alg hydrogel over the N-DC (FL)8 single-network gel, but also increased the breaking strain. It is possible that the ionically crosslinked network served as a reversible sacrificial network to prevent stress-concentration and improve the extensibility of the network. It is of note that the I-N-DC (FL)8/AlgL and I-N-DC (FL)8/AlgM hydrogels displayed a similar Young’s modulus as N-DC (FL)8, but a much higher extensibility. This combined effect leads to significantly increased toughness for the I-N-DC (FL)8/AlgL and I-N-DC (FL)8/AlgM hydrogels (~1.0 MJ/m3 and ~860 kJ/m3, respectively) over N-DC (FL)8 hydrogels. In contrast, the I-N-DC (FL)8/AlgH gel exhibited the highest Young’s modulus among the I-N-DC (FL)8/Alg hydrogels, but the lowest extensibility, leading to a toughness (~280 kJ/m3) that is similar to that of N-DC (FL)8 hydrogels. These results clearly indicate that the conjoined double-network strategy we employed here effectively improved the toughness of the single-network N-DC (FL)8 gels. By tuning the grafting density of phenol groups in the alginate chains, it is now possible to tune the 140 mechanical properties of (FL)8 hydrogels over a broad range. However, there is a trade-off between the Young’s modulus and extensibility of hydrogels, as the higher crosslinking density of the double network may increase the Young’s modules as the expense of stretchability. It is likely that excessive covalent crosslinks may constrain the intramolecular movement of the polymers in the networks, which typically contributes to energy dissipation. Figure 5-6 Representative stress-strain curves of (FL)8/Alg gels. A) Representative stress-strain curves of D-DC (FL)8/AlgM (grey) and N-DC (FL)8/AlgM (blue). B) Representative stress-strain curves of I-N-DC (FL)8/AlgM (red), N-DC (FL)8/AlgM (blue), N-DC (FL)8 alone (yellow) and N-NC (FL)8 alone (green). The I-N-DC (FL)8/AlgM hydrogels show improved tensile strength and extensibility, compared to N-DC (FL)8 alone reported previously. C) Representative stress-strain curves of N-DC (FL)8/unmodified alginate (blue), I-N-DC (FL)8/unmodified alginate (red). The N-DC 20%(FL)8/1% unmodified Alginate showed Young’s modulus 566.9±73.6 kPa with breaking strain of 84.0±9.4%. After equilibration in 25 mM Ca2+, the I-N-DC (FL)8/unmodified Alginate exhibited improved Young’s modulus of 944.0±71.8 kPa and breaking strain of 140.3±23.6% (n=3). D) The representative tensile stress-strain curves of I-N-DC (FL)8/Alg hydrogel rings, compared to N-DC (FL)8 alone reported previously. The Young’s modulus and the breaking strain of I-N-DC (FL)8/Alg double-network hydrogels are related to the modification density of tyramine on alginate. 141 The control N-DC (FL)8/unAlg gels also showed improved stiffness (~944 kPa) and extensibility (~140%) in the presence of Ca2+. Although the ionically crosslinked unmodified alginate can reinforce the mechanical performance of the N-DC (FL)8 network, the covalently conjoined double-network gels displayed a wide range of stiffness and stretchability when the density of phenol groups in the alginate component was varied. To better understand the enhancement mechanism of the I-N-DC (FL)8/Alg gels, the I-N-DC (FL)8/AlgM gels was used as a model system. We investigated the contribution of the additional ionic crosslinks provided by the derivatized alginate. After ionic crosslinking, the I-N-DC (FL)8/AlgM hydrogels became stiffer and more extensible than the N-DC (FL)8/AlgM rings (Figure 5-5A and Figure 5-6B). This may be indicative that the ionic crosslinking sites enhanced the double network. Additionally, the improved extensibility may be attributed to the weak but reversible physical interaction between the G blocks and Ca2+. The ionic crosslinks could dissipate energy by unzipping the network of covalent crosslinks during large deformation. The ionically reinforced network is supported by the lower water content of I-N-DC (FL)8/AlgM as compared with N-DC (FL)8/AlgM (Figure 5-5B). Additionally, the stiffness of the double-network (FL)8/Alg gels can be cyclically regulated between I-N-DC and N-DC states by adding or removing Ca2+. As shown in Figure 5-5D and E, the Young’s modulus of the N-DC (FL)8/AlgH hydrogels can be tuned between ~0.4 MPa and ~1 MPa, due to the reversible ionic bonding of the alginate constituent, demonstrating a wide tunable range in stiffness. Finally, we also studied the influence of protein concentrations (20%, 15% and 10%) on the mechanical properties of the double-network gels. For gels made of the same alginate, similar Young’s moduli were observed for the double network hydrogel, as shown in the Figure 5-5F. The 142 breaking strain exhibited a similar trend, except for I-N-DC (FL)8/AlgH gels, which became brittle when the protein concentrations were reduced (Table 5-2). Similarly, there was no significant difference in water content of the I-N-DC (FL)8/AlgH gels as a function of the protein concentration (Figure 5-5A). Table 5-2 Young’s modulus (E) and breaking strain of (FL)8/Alg double network hydrogels. (Average ± Standard deviation, n=4) I-N-DC AlgL (FL)8 E (kPa) Breaking strain 20% 624±147 228±28% 15% 620±50 203±61% 10% 542±134 200±52% AlgM (FL)8 E (kPa) Breaking strain 20% 664±88 198±25% 15% 730±119 201±15% 10% 750±53 196±13% AlgH (FL)8 E (kPa) Breaking strain 20% 1025±103 103.3±16% 15% 925±40 60.0±10% 10% 858±58 45.0±11% 5.5 Conclusions By chemically modifying alginate with phenol groups, the derivatized alginate was able to undergo both ionically and photochemically crosslinking reactions. This allowed us to engineer a double-network hydrogel by covalently interpenetrating polyproteins and alginate. By employing a denatured-crosslinking approach, the resultant gels were demonstrated to possess both high elasticity and high stiffness. Compared to each individual network component, the double network 143 hydrogel displayed significantly improved toughness. We suggest that the chemical crosslinks provided by both DC-(FL)8 and alginate contributed to the superb stiffness, while the ionic crosslinks from the alginate can also act as extensible components. Moreover, it is noted that the stiffness and the stretchability of the double-network hydrogels can be tailored over a wide range (Young’s modulus from ~0.6 kPa to ~1 MPa) by adjusting the derivatization density of the alginate. In addition, the reversible ionic bonding of the alginate component allows for cyclic tuning of the double-network gels over a wide range in stiffness. The stiffness of the double-network hydrogels was vastly improved over the single-network protein-based hydrogels. Such an improvement in the mechanical properties will make these hydrogels promising for biomedical applications. 144 Chapter 6: Dynamic Protein Hydrogels with Reversibly Tunable Stiffness Regulate Human Lung Fibroblasts Spreading Reversibly 6.1 Synopsis Protein-based hydrogels can be utilized for ECM by providing biocompatible and dynamic local microenvironments. Through cell-ECM interactions, these external mechanical signals are transduced to the cell’s interior, triggering biochemical signaling cascades that determine a wide variety of cellular behaviors, ranging from cell adhesion, cell spreading to cell differentiation. However, only a limited number of such hydrogels have been engineered allowing for investigating cell response to the change of hydrogel stiffness. Moreover, the stiffness of these hydrogels can only be switched uni-directionally and irreversibly. Here we report an entirely protein-based hydrogel that allows for a cyclic, reversible modulation of hydrogel mechanics over a relatively large range of Young’s moduli, and used it to investigate the cell spreading behaviors of human lung fibroblasts (HLF) in response to the cyclic change of the hydrogel Young’s modulus. *A version of Chapter 6 has been published as [Linglan Fu], Amanda Haage, Na Kong, Guy Tanentzapf, and Hongbin Li. (2019) Dynamic protein hydrogels with reversibly tunable stiffness regulate human lung fibroblast spreading reversibly. Chemical Communications. 55 (36): 5235-5238. Adapted with permission of The Royal Society of Chemistry, Copyright 2019. 145 This protein-based dynamic hydrogel was fabricated by using the mutually exclusive protein (MEP) GL5CC/I27 as a building block. MEP is a specially designed, redox-responsive protein-folding switch. In the oxidized state, the host domain GL5CC is folded; in the reduced state, the GL5CC unfolds and the effective chain length of the MEP changes from 2.6 to 26 nm, leading to a change in the polymer chain length between two crosslinking points.43 Building upon this approach, we engineered a dynamic protein hydrogel based on TNfn3-GB1-resilin-(GB1-GL5CC/I27-resilin)2. TNfn3 contains a RGD motif for cell adhesion and resilin provides covalent crosslinking residues. We found that this dynamic hydrogel could be reversibly tuned between a stiff and soft state via a redox reaction. Moreover, this protein hydrogel supports cell adhesion and proliferation. In response to a cyclic change of the hydrogel’s Young’s modulus between ~6 kPa and 20 kPa, HLF dynamically change their morphology in a fully reversible fashion, demonstrating the great potential utility of such dynamic protein hydrogels in mechanochemical studies. 6.2 Introduction Mechanobiology is a field that studies the roles played by local force stimuli and changes in the mechanics of cells or tissues in mediating cell morphology, cell or tissue development, cell differentiation, and physiology.385-386 Over the last two decades, the importance of mechanical cues presented by the local ECM, whether native or engineered artificial ECM, on cellular behaviors and function have been well recognized and documented.197, 387-388 The external mechanical signals are transduced to the cell interior through cell-EMC interactions, triggering biochemical signals that determine various cellular behaviors, such as cell adhesion, cell spreading, and cell differentiation. Understanding the specific effects of mechanical properties of the ECM on cell physiology is of critical importance, and has attracted considerable interest. 146 Most previous studies in this area were carried out utilizing static polymer hydrogels as an artificial biomimetic ECM, whose properties are static.388-391 Such polymer hydrogels lack the characteristic dynamic changes found in living biological systems. To better mimic a dynamic biological environment, dynamic polymer hydrogels whose properties can be regulated via external stimuli have attracted considerable interest, and significant progress has been achieved.40, 390, 392-394 Novel approaches have been developed to engineer hydrogels that can either stiffen or soften in response to external stimuli, such as light, pH, or temperature.40-41, 43, 52, 392-403. Although some dynamic hydrogels were tested for various cellular applications (including cell release and migration), only a limited number of such hydrogels were exploited to investigate cell response to a change in hydrogel stiffness.40, 399-402 Moreover, the stiffness of some hydrogels can only be switched uni-directionally and irreversibly. In living tissues, cells not only experience unidirectional, but also varying changes of matrix stiffness. For example, during the wound healing process, tissues will experience sequential stiffening and softening. Such varying changes can even be cyclic and repeated continuously. However, it is challenging to engineer dynamic hydrogels whose stiffness can be switched in a reversible fashion.40, 43, 401, 403-404 Recently, Anseth and coworkers reported a successful realization of one cycle of stiffening-softening-stiffening in a polymer hydrogel made via photocrosslinking and photolytic reactions, and investigated the change of hydrogel mechanics on the behavior of human mesenchymal stem cells.40 Using the light sensitive protein LOV domain, which can undergo reversible association-dissociation in response to light, DeForest and coworkers engineered a protein-poly(ethylene glycol) hybrid hydrogel,404 whose stiffness could be changed in a cyclic fashion (with a 15% difference in Young’s modulus between the softer and stiffer states), making it possible to study cellular response to a cyclic change of hydrogel stiffness. Despite this progress, engineering entirely 147 protein-based dynamic hydrogels for mechanobiological studies remains limited and underexplored. Here, we report an entirely protein-based hydrogel that allows for cyclic, reversible modulation of hydrogel mechanics over a relatively large range of Young’s moduli, and used it to investigate the cell spreading behavior of HLF in response to the cyclic change of the Young’s modulus of the hydrogel. In our previous work, we engineered a redox-responsive protein folding switch based on the so-called mutually exclusive protein (MEP), GL5CC-I27, and used it to engineer dynamic protein hydrogels whose Young’s modulus can be reversibly regulated by using redox conditions.43, 288 MEP is a special class of fusion protein, in which a guest domain with a large distance between its N- and C-termini is inserted into a small loop of the host domain.405-406 This creates structural incompatibility in the MEP, so that only one of the two domains can remain folded at a time. In the MEP GL5CC-I27, the guest protein I27 is inserted into a short loop of the host domain GL5CC.43, 405 In the oxidized state, the MEP exists as GL5(F)/I27(U), where GL5 is folded and I27 is unfolded, leading to a short effective chain length (Lec) between the N- and C-termini of the MEP; upon reduction, the MEP exists as GL5(U)/I27(F), resulting in a much larger Lec (Figure 6-1A). Using this MEP as a building block, we engineered a dynamic protein hydrogel based on GB1-resilin-(GB1-MEP-resilin)2 (GR(G-MEP-R)2).43, 288 The large change of Lec in the two conformational states of the MEP resulted in a large change in the effective crosslinking density of the hydrogel and the hydrogel stiffness (Figure 6-1). 148 Figure 6-1 Dynamic protein hydrogels constructed from MEP. A) Schematic of the redox-responsive MEP. In the oxidized state, the MEP has a shorter effective chain length (Lec); upon reduction, the MEP shows a longer Lec. B) Schematic of the protein hydrogel network constructed from FGR(G-MEP-R)2. Building upon this prior successful approach, we report here the engineering of a reversible protein hydrogel that incorporates the third FnIII domain from the ECM protein tenascin (TNfn3) as the cell adhesive ligand. We engineered the polyprotein FGR(G-MEP-R) 2, where F corresponds to TNfn3 (Appendix A.6). TNfn3, which contains an Arginine-Glycine-Aspartic acid (RGD) motif, is responsible for the cell adhesion properties of tenascin407 and has been used to engineer various protein hydrogels that can support cell adhesion and spreading.58, 285, 408 The polyprotein FGR(G-MEP-R)2 was overexpressed in E. coli strain DH5α with a good yield (~45 mg per liter of bacterial culture). 149 6.3 Materials and methods 6.3.1 Protein engineering The genes of FGR(G-MEP-R)2 and FRF4RF4R were constructed using standard molecular biology techniques following well-established protocols.43, 285 Protein sequences are shown in Appendix A. Protein expression was carried out in E. coli strain DH5α. Protein expression and purification were performed using the procedure described in Chapter 2 (2.3.1). 6.3.2 Hydrogel preparation for cell spreading and tensile tests Gelation of the polyprotein based on photochemical crosslinking strategy was done near a flame or in a biocabinet to ensure that the resulting hydrogels were sterile.296 To prepare 12% FGR(G-MEP-R)2 (120 mg/mL) or 7% FRF4RF4R (70 mg/mL) hydrogels with flat and large surface areas for cell spreading and imaging, lyophilized proteins were re-dissolved in sterile PBS mixed with APS (1M) and [Ru(II)(bpy)3]2+ (20 mM) with final concentrations of 50 mM and 260 μM, respectively. The mixture was quickly transferred onto a clean hydrophobic surface and covered with a sterile glass coverslip. The thickness of hydrogels was ~0.8 mm, and the area of the hydrogel matrices was ~0.5 cm2. Then, the sample was exposed under a 200W fiber optic white light source from the height of 10 cm above the coverslip for 10 min. The hydrogel was gently taken out of the mold and soaked into sterile PBS to get rid of unconsumed crosslinking reagents. Then the washed gels were transferred into 6-well plates in fresh DMEM medium under the desired redox conditions. Oxidized hydrogels were immersed in regular DMEM oxidized by air, whereas reduced hydrogels were treated with an additional solution of 10 mM glutathione (GSH), which is a reducing agent generally used for cell studies. 150 To prepare a ring-shaped 12% FGR(G-MEP-R)2 hydrogel for tensile tests, 12 mg of lyophilized protein was dissolved in 93.7 μL phosphate-buffered saline (PBS), with 5 μL APS (1M) and 1.3 μL [Ru(II)(bpy)3]2+ (20 mM) with final concentrations of 50 mM and 260 μM, respectively. The protein solution was quickly transferred to a custom-made plexiglass ring-shaped mold (din=8 mm, dout=10mm, h= 3mm), and was exposed under a 200W fiber optic white light source from the height of 10 cm above the mold for 10 min. The ring sample was carefully taken out of mold. The oxidized gel rings were stored in filtered PBS in presence of 10 mM H2O2 to achieve equilibrium swelling. To obtain reduced samples, the oxidized ring samples stored in PBS with H2O2 were transferred and immersed in PBS with additional 10 mM DTT for over 6 hours to allow complete reduction and to achieve the swelling equilibrium. 6.3.3 Characterization of dityrosine cross-links in hydrogel by acid hydrolysis-fluorescence method The degree of crosslinking of the FGR(G-MEP-R)2 hydrogels was characterized following a well-established fluorimetry method.231, 275 Typically, a 12 % hydrogel sample (∼10 mg) was reacted with 100 μL HCl (6 N) in a sealed 1.5 mL centrifuge tube in the metal heat block at 105 ℃ for 2 h to achieve full hydrolysis of the peptide bonds. Then, 100 μL of acid hydrolysis product was transferred into a new 1.5 mL centrifuge tube and neutralized by 10 μL NaOH (5 M). Next, 100 mM Na2CO3-NaHCO3 buffer (pH 9.9) was added to the tube to 1 mL final volume. Fluorescence spectra of the samples were measured, using the same procedure as described in Chapter 4 (4.3.7). 151 6.3.4 Cell spreading and redox switching HLF cells were first cultured in T75 tissue culture flasks with DMEM, supplemented with 10% (v/v) fetal bovine serum (FBS) and 1% (v/v) Penicillin/Streptomycin mix. Since the main purpose of this study was to investigate the effect of cyclic change of Young’s modulus on the spreading/adhesion behavior of HLF, only standard DMEM medium and 37 ℃ were used (with the exception that the reduced medium was DMEM supplemented with 10 mM GSH). For cell spreading experiments, 1.4x105 HLF cells were seeded on top of the pre-treated oxidized and reduced FGR(G-MEP-R)2 hydrogels in each well of a 6-well plate at 37 ℃ for 24 hours, with the FRF4RF4R hydrogels serving as controls. The hydrogels from DAY1 (Ox, Rd) without redox switching were fixed and stained after 24 hours incubation, using the protocol described below. The rest of the hydrogels underwent the first redox switching cycle by removing the old oxidized and reduced media and replacing with fresh reduced DMEM (10 mM GSH) and oxidized DMEM respectively (DAY2: Ox-Rd, Rd-Ox), and then incubated at 37 ℃ for another 24 hours. This was then repeated to obtain the second redox switching, DAY3 samples (Ox-Rd-Ox, Rd-Ox-Rd). To stain and image HLF cells, samples were first washed twice with pre-warmed sterilized PBS (pH 7.2, 37 ℃) to remove non-adherent cells before cell fixing. Adherent cells were stained with Alexa Fluor 488 phalloidin and 4′,6-diamidino-2-phenylindole (DAPI) following the manufacturer’s protocol (Thermo Fisher). Briefly, cells were fixed with a 4% paraformaldehyde solution in PBS for 15 mins at room temperature. Fixed cells were permeabilized by exposure to 0.1% Triton X-100 in PBS for 5 mins. Samples were then incubated with phalloidin for 20 mins and DAPI for 5 mins at room temperature. After staining, samples were washed with PBS to 152 remove excess staining dye. Images were taken at ×10 magnifications using a fluorescence microscope (Nikon, Eclipse Ti). To assess the cytocompatibility of FGR(G-MEP-R)2-based hydrogels, we used a Click-iT EdU assay to assess the hydrogels’ cytocompatibility. In this assay, the fluorescently labeled thymidine analogue EdU (5-ethynyl-2’-deoxyuridine) is incorporated into newly synthesized DNA efficiently, allowing for the fluorescent labeling of proliferating cells. HLF cells were plated on 7% FRF4RF4R hydrogels and both oxidized and reduced 12% FGR(G-MEP-R)2 hydrogels which were pre-treated for cell loading. We also cultured HLF on polystyrene tissue culture dishes as a control. The cells were seeded with 50% confluence and allowed to grow on hydrogels and control dish surfaces for 4 days before sequential treatment. Oxidized/reduced DMEM media mixed with 1X EdU working solution (10 μM EdU in DMSO) were added in the culturing wells to replace the regular DMEM under redox conditions for continuous incubation. Then the hydrogel samples were fixed and stained at two different lengths of incubation time, namely 24 hrs and 72 hrs. Staining and EdU detection of samples followed the standard Click-iT EdU Assay protocol. The samples were washed with pre-warmed PBS. The HLF cells were fixed by 3.7% formaldehyde followed by permeabilization in 0.5% TritonX-100, and then washed by 3% BSA in PBS. 200 μL of Click-iT reaction cocktail was added to each sample, incubating for 30 mins before washing with 3% BSA in PBS. For additional nucleus staining, DAPI was added afterwards. The samples were soaking-washed with PBS to remove extra staining dye before imaging. Images were collected using an Olympus FV1000 confocal microscope. 153 6.3.5 Tensile tests Tensile tests were performed on an Instron-5500R tensometer with a custom-made force gauge in PBS with 10 mM H2O2 or PBS with 10 mM DTT at constant temperature (25 ℃). For technical reasons, ring-shaped biomaterial specimens were used. An extension rate of 25 mm/min was used in our experiments. The stress at 15% strain is taken as the Young’s modulus of the sample. 6.3.6 Scanning electron microscopy (SEM) imaging Each hydrogel was imaged using a Hitachi S4700 scanning electron microscope. A 12 % hydrogel sample was prepared in Eppendorf tubes, and stored at 4 °C for 24 hours to allow the mixture to completely gelate. The samples were then shock-frozen in -80 ℃ , and quickly transferred to a freeze drier where they were lyophilized for 12 hours. Lyophilized samples were then carefully fractured in liquid nitrogen, and fixed on aluminum stubs. The sample surface was coated by 5 nm of gold prior to SEM measurements. 6.4 Results and discussion 6.4.1 Dynamic protein hydrogel based on FGR(G-MEP-R)2 shows reversible redox-responsive microstructural, physical and mechanical properties To crosslink the aqueous solution of FGR(G-MEP-R)2 into a chemically crosslinked protein hydrogel, we used the well-developed Ru2+-mediated photochemical strategy, 29, 288, 296 which allows for the coupling of two tyrosine residues in close proximity into a dityrosine adduct. All the protein building blocks in the polyprotein FGR(G-MEP-R)2 contain exposed tyrosine residues, thus the crosslinking can occur along the polyprotein chain randomly. Under white light illumination, the aqueous solution of FGR(G-MEP-R)2 can be readily crosslinked into a 154 transparent hydrogel within tens of seconds, with approximately 5% of tyrosine crosslinked into dityrosine crosslinking points. Figure 6-2A shows a photograph of the 12% protein hydrogel prepared from the oxidized polyprotein. Since the oxidized MEP has a shorter Lec than the reduced one, we expect that the protein hydrogels will exhibit different physical and mechanical properties depending on their redox state. Indeed, upon reduction, the FGR(G-MEP-R)2 protein hydrogel exhibited a higher swelling ratio than the oxidized one (Figure 6-2B). Scanning electron microscopy experiments revealed a porous structure for the hydrogels in both oxidized and reduced states. However, the reduced hydrogels showed slight larger pores than the oxidized ones (Figure 6-2C-D). Figure 6-2 Swelling behavior and microstructure of oxidized and reduced FGR(G-MEP-R)2 hydrogels. A-B) Photographs of an oxidized (A) and reduced (B) FGR(G-MEP-R)2 hydrogel (12%). Reduced hydrogels showed increased swelling. C-D) SEM images of the oxidized (C) and reduced (D) 12% FGR(G-MEP-R)2 hydrogels. The scale bars represent 100 μm. The hydrogels showed a microporous crosslinking network structure. The pore size of the reduced hydrogels is on average larger than that of oxidized ones. We then carried out tensile tests to characterize the mechanical properties of the FGR(G-MEP-R)2 hydrogels. In the oxidized state, the 12% FGR(G-MEP-R)2 hydrogel exhibited a nonlinear stress-strain curve with a Young’s modulus of ~20 kPa. A small hysteresis was observed 155 between the stretching and relaxation curves, indicative of a high resilience of the hydrogel (~80%) (Figure 6-3A). The hydrogel was subsequently treated with a biological reducing agent GSH (10 mM). After reaching equilibrium, the reduced FGR(G-MEP-R)2 protein hydrogel showed a much reduced Young’s modulus (~6 kPa), a 3 fold change compared to the oxidized one. And the hysteresis between the stretching and relaxation cycles also increased. These results are consistent with the expected changes of the MEP conformation in the oxidized and reduced states: in the oxidized state, MEP is in the GL5CC(F)/I27(U) conformation with a shorter Lec. In the reduced state, the host domain GL5CC unfolds, leading to a significant increase of Lec. Since the crosslinking points in the hydrogel network do not change, the increasing of the Lec effectively reduced the crosslinking density of the hydrogel network and led to a decrease of Young’s modulus and subsequent increase of the swelling ratio. To further confirm that these changes results from the MEP and not the incorporated TNfn3 domain, we carried out similar experiments on the hydrogel based on the polyprotein FRF4RF4R, where F represents the TNfn domain. It is evident that the mechanical properties of the FRF4RF4R hydrogel do not change significantly regardless of whether the buffer contains the reducing agent GSH (Figure 6-3A inset) or not, supporting our conclusion that the changes in the mechanical properties of the FGR(G- MEP-R)2 hydrogel are the result of conformational changes of the MEP, regulated by the redox potential. 156 Figure 6-3 Reversible mechanical properties of FGR(G-MEP-R)2 hydrogels governed by redox stimuli. A) Stress-strain curves of FGR(G-MEP-R)2-based 12% hydrogels in different redox states (oxidized, reduced, re-oxidized). The stress-strain curves of FRF4RF4R-based hydrogels (inset) do not change with the changing redox conditions. B) The dynamic change of the Young’s modulus and swelling ratio of the 12% FGR(G-MEP-R)2 hydrogel in response to changes of the redox condition. The data are presented as average ± standard deviation (n=5). Having established that the FGR(G-MEP-R)2 hydrogel can respond to changes in redox potential, we further investigated if such changes in the physical and mechanical properties are reversible and can be repeated by cycling through oxidizing and reducing conditions. Starting from the oxidized state, we cycled the FGR(G-MEP-R)2 protein hydrogel between oxidized and reduced states and carried out tensile and swelling tests. In three subsequent oxidation/reduction cycles, we found that both the swelling ratio and Young’s modulus of the FGR(G-MEP-R)2 protein hydrogel can be cycled reversibly in response to the change of the redox potential (Figure 6-3B). Moreover, this reversible change does not depend on the initial state of the hydrogel. In other words, when 157 starting from the reduced state, the hydrogel showed the same reversible changes in swelling ratio and Young’s modulus. In addition, this hydrogel demonstrated good long-term reversibility. After being stored in PBS or cell culture medium for two weeks, this hydrogel showed reversibility in the mechanical and physical properties in response to the change of the redox condition, similar to those shown in Figure 6-3B. 6.4.2 The MEP hydrogels support cell proliferation Our results have demonstrated that the FGR(G-MEP-R)2 hydrogel system can serve as a redox-controlled, fully reversible hydrogel whose Young’s modulus can be dynamically and reversibly modulated, thus providing an ideal platform for us to investigate how cells respond to the reversible and cyclic changes of hydrogel mechanics. For this purpose, we first assessed the cytocompatibility of FGR(G-MEP-R)2-based hydrogels using the Click-iT EdU assay by culturing HLF on the hydrogels. After 24 hrs in culture, the cell proliferation ratio of HLF incubated on oxidized and reduced hydrogels is 5% and 8%, respectively. After 72 hrs, the cell proliferation ratio was 12% and 23% on oxidized and reduced cells, respectively (Figure 6-4). The proliferation ratio of HLF on hydrogels is comparable with that of HLF on polystyrene tissue culture dishes which was ~8%, indicating that the FGR(G-MEP-R)2-based hydrogels are cytocompatible. 158 Figure 6-4 HLF cells can proliferate on MEP hydrogels. All cell nuclei were stained using the blue fluorescent dye DAPI (4′,6-diamidino-2-phenylindole). Cell nuclei, newly grown after Click-iT Edu treatment, were fluorescently labelled by Alexa Fluor azide (red). The proliferation ratio was calculated from the percentage of the red fluorescent nuclei among the total number of cells. On reduced FGR(G-MEP-R)2 gels, the ratio of new cells to the initially seeded cell increases from ~8% to 23%. On oxidized FGR(G-MEP-R)2 gels, the cell proliferation ratio increases from ~5.0% to 12.0%, The scale bars represent 50 μm. 6.4.3 HLF cells dynamically change morphology in response to a cyclic mechanical change of the MEP hydrogel Having confirmed the cytocompatibility of FGR(G-MEP-R)2-based hydrogels, we then investigated if the cell spreading and adhesion behaviors of HLF can be modulated by reversible changes in hydrogel stiffness. Previous studies showed that the rigidity of the substrates affects the morphology, cytoskeletal structure, and adhesion of fibroblasts.387 It was shown that fibroblasts spread well on stiffer substrates (with a Young’s modulus higher than 10 kPa) but spread less or even remain round on softer matrices (with a Young’s modulus less than 1 kPa).409-410 Here we investigated the spreading of HLF in response to the cyclic change in the Young’s modulus of the 159 FGR(G-MEP-R)2-based dynamic hydrogel using standard mammalian cell culture conditions at 37 ℃. HLF were seeded onto both oxidized and reduced hydrogels (12%) on day 1. For the oxidizing condition, regular DMEM medium was used and the oxygen in the surrounding air was used as the oxidizing agent; for the reducing condition, DMEM medium supplemented with 10 mM GSH was used. As expected, HLF cells spread well on the oxidized hydrogel with a Young’s modulus of ~20 kPa. The actin filaments of HLF were stained by phalloidin , clearly showing the morphology of the cells. The results indicated that fibroblasts adhere well on the stiffer hydrogel surface. The average cell area of the spread fibroblasts was approximately 5000 μm2. In contrast, fibroblasts spread to a much lesser extent on the reduced hydrogel, with a Young’s modulus of ~6 kPa, and some cells remained round. The average cell area was only around 2500 μm2. These results are consistent with previous studies that showed fibroblasts tend to spread well on stiffer hydrogel surfaces.40, 394, 409-410 It is noted that the swelling behavior of a MEP gel is closely correlated with its stiffness. The swelling/shrinking of the gels might also contribute to the change of cell morphology, as the interactions of cell-ECM can be affected by the spatial changes.411 To confirm that the observed morphology differences of the fibroblasts are not due to the change of cell culture media, we carried out control experiments on FRF4RF4R hydrogels, as the stiffness of a FRF4RF4R hydrogel doesn’t change with redox conditions. However, the Young’s modulus of the FRF4RF4R gels can be varied by controlling the gel concentration (5%-20%), and resulting moduli ranged from ~5 kPa to ~20 kPa. We found that fibroblast spread equally well on these hydrogels in the regular, as well as the reducing medium (10 mM GSH), confirming that indeed the morphology change of fibroblasts on the FGR(G-MEP-R)2 hydrogel is due to the stiffness change, not the cell culture media (Figure 6-5 and Figure 6-6). 160 Figure 6-5 Control experiments were carried out with HLF cells spreading on 7% FRF4RF4R hydrogels. Control experiments were carried out at the same time as the same conditions and procedures as the experiments on 12% FGR(G-MEP-R)2. Redox conditions of the culture medium were switched alternatively every 24 hrs. There were no obvious change of cell area and cell circularity on control gels during the redox switching as MEP hydrogels. Statistical significance is indicated (n.s.: p>0.05, no statistical significance; *: p<0.05; ***: p<0.001). 161 Figure 6-6 Control HLF culturing experiments on FRF4RF4R hydrogels with different Young’s modulus. A) Young’s modulus of 20%, 7%, 5% FRF4RF4R hydrogels. Their Young’s moduli are 19.9 kPa, 13.7 kPa and 6.4 kPa respectively. Error bars indicate standard deviation. B-C) Cell spreading area and circularity of HLF on the reduced/oxidized FRF4RF4R hydrogels with different Young’s modulus. Both cell spreading area and circularity of HLF do not display changes in response to redox condition. Statistical significance is indicated (****: p<0.0001; ***: p<0.001; *: p<0.05; n.s.: p>0.05, no statistical significance). 162 Figure 6-7 Morphology of human lung fibroblasts can dynamically respond to the change of the hydrogel stiffness. A) Histograms of HLF cell area during a soft-stiff-soft-stiff switching cycle. On soft hydrogels, HLF cells spread to a lesser extent, with small cell areas. Upon switching to a stiff hydrogel state, HLF cells dynamically change their morphology and display much larger cell area. Insets show fluorescence images of HLF cells. Actin filaments were labelled by Alexa Fluor 488 phalloidin and DNA in the nucleus was stained by DAPI. The scale bars represent 50 μm. B) and E) Box plot of HLF cell area on protein hydrogels. C) and F) Box plot of HLF cell circularity on protein hydrogels with different Young’s modulus. In response to the changing Young’s modulus of the hydrogel, HLF cells only showed a small change in circularity. D) Histograms of the cell area of HLF during the two stiff-soft-stiff-soft switching cycles of change the stiffness of the hydrogels. On stiff hydrogels, HLF spread well with large cell areas. Upon switching to soft hydrogel state, HLF dynamically change their morphology and display much small cell area. Inset show fluorescence images of HLF. The scale bars represent 50 μm. Statistical significance is indicated (****: p<0.0001; ***: p<0.001; *: p<0.05; n.s.: p>0.05, no statistical significance). 163 To investigate if fibroblasts can dynamically respond to changes in the stiffness of the hydrogel substrate in a reversible fashion, we switched the cell culture medium between the oxidizing and reducing condition in a cyclic fashion. As shown in Figure 6-7, we found that HLF cells can dynamically respond to the changes of hydrogel mechanics in a reversible fashion, as a function of the redox state. This dynamic change was readily manifested as a change in spread area of the HLF cells. Upon switching from a reducing to oxidizing condition, HLF cells switched in morphology, going from a less spread to more spread state, leading to a significant increase in cell area (Figure 6-7A and B). In addition, the roundness of the HFL cells only decreased slightly (Figure 6-7C). Upon switching from an oxidizing to reducing condition, the opposite trend of morphology changes was clearly observed. From two complete cycles of switching between reducing and oxidizing conditions (RORO), it is evident that the morphology change of HLF cells is fully reversible in response to the change of the Young’s modulus of the hydrogel (Figure 6-7A and B). The change in the HLF roundness during the switching cycles was relatively small, and not statistically significant (Figure 6-7C). This result is different from that obtained on human mesenchymal stem cells using softer hydrogels (~500 Pa).40 A possible reason is that both the oxidized and reduced FGR(G-MEP-R)2 are relatively stiff (6 kPa versus 20 kPa). Importantly, the dynamic responsiveness of HLF cells to the change of hydrogel stiffness does not depend on the initial state of the hydrogel. Starting from stiff hydrogels, we observed a similar morphological change of HLF cells upon subsequent cycling between stiff and soft states (Figure 6-7 D-F). Moreover, in the control experiments of cultured HLF cells on the FRF4RF4R hydrogel, no obvious change was observed for the spreading area or cell circularity during the cyclic switching of the cell culture medium between oxidizing and reducing conditions (Figure 6-5 and Figure 6-6). Taken together, these results clearly demonstrated that the HLF cells can constantly sense the 164 hydrogel stiffness, and the mechano-response led to the significant change in cell spreading that was observed. This change in cell morphology is dynamic and fully reversible, in response to the change of the hydrogel stiffness. 6.5 Conclusions In summary, we have engineered a novel cell-adhesive, entirely protein-based dynamic hydrogel whose mechanical and physical properties can be switched reversibly by redox state. This protein hydrogel supports cell adhesion and proliferation. In response to a cyclic change of the hydrogel’s Young’s modulus between ~6 kPa and 20 kPa, HLF dynamically change their morphology in a fully reversible fashion, demonstrating the great potential utility of such dynamic protein hydrogels in mechanochemical studies. To the best of our knowledge, this hydrogel is the first reported example of an entirely protein-based dynamic hydrogel that has been used to investigate the effect of cyclic change of hydrogel mechanics on cell response. This hydrogel shows a 3-fold change in Young’s modulus between its two conformational states, thus offering a broader dynamic range of Young’s modulus for cell studies. It is worth noting that such hydrogels are relatively stiff in the spectrum of physiologically relevant stiffness. Expanding the range of Young’s modulus of these protein hydrogels will be the focus of our future work. Using protein-engineering techniques, it should be feasible to change the effective chain length of the two conformations of the MEP to further increase the dynamic range of the Young’s modulus of the protein hydrogels. Experimenting with the photocrosslinking conditions could also allow for further tailoring of the mechanical properties of the protein hydrogels. We anticipate that these protein hydrogels, with an expanded dynamic range of stiffness, will find applications in a wide 165 range of mechanobiology studies, including those that focus on how stem cells differentiate in response to cyclic changes in the mechanical properties of a protein hydrogel matrix. 166 Chapter 7: Conclusions and Outlook 7.1 Conclusions This thesis provides a comprehensive understanding of how hydrogel networks can be made using globular folded polyproteins. Using proteins offers many advantages when making biomaterials with the desired mechanical performance for particular applications. In the introductory chapter, pioneering and state-of-the-art studies on protein-based hydrogels are reviewed. The related crosslinking methods and the general mechanical behaviors of these biomaterials are also discussed. Moreover, classic rubber elasticity theory is introduced in Chapter 1. This framework provides a mathematic guide to understanding mechanical properties of a rubber-like network. Additionally, I comment on the strengths and the limitations of mechanical reinforcement approaches that have been currently used to construct strong and tough hydrogels. Chapter 2 describes a novel semi-quantitative tool to predict mechanical properties of globular folded polyprotein-based hydrogels. The length equivalent molecular weight approach adapts the classic rubber elasticity theory for protein-based hydrogels composed of globular folded domains. On the one hand, this semi-quantitative tool allows us to understand the influence of the globular folded domains on the mechanical behavior of hydrogels. On the other hand, this new method provides a guide for the rational design of globular polyprotein-based hydrogels with desired mechanical properties. Chapter 3 investigates mechanochemical coupling cycles of globular folded proteins at both the single-molecule and macroscopic levels, using the artificial elastomeric protein GB1 as a model system. We demonstrated for the first time that protein folding can act as a driving force, generating mechanical energy. In addition, the macroscopic mechanical energy generated from conformational change of globular proteins closely corresponds to single molecule 167 mechanochemical events. Moreover, the mechanical energy generated by globular protein folding/unfolding can be controlled reversibly at a macroscopic level by exchanging surrounding buffers. The controlled mechanochemical coupling cycle of GB1-based hydrogels provides a new avenue towards engineering protein-based soft actuators. In Chapter 4, we develop a novel denaturing-crosslinking hydrogelation method to synthesize a super tough protein-based hydrogel. (FL)8 polyproteins are crosslinked in their unfolded state at high concentration, which introduces chain entanglements into the gel network. Following renaturation, the shortening of the polypeptide chains and the presence of chain entanglements can significantly increase the effective network crosslinking density. The resultant hydrogel exhibits a Young’s modulus of ~0.7 MPa, a toughness of ~250 kJ/m3 in tensile testing, ~1.7 MPa in compressive modulus and finally, a toughness of ~3.2 MJ/m3. To the best of our knowledge, this gel represents the highest compressive strength (~68 MPa) achieved by a globular protein-based hydrogel. The high elasticity, high toughness, and fast recovery of the hydrogel resemble the load-bearing feature and high stiffness of cartilage tissues. This work not only breaks the mechanical limitations of globular protein-based hydrogels, but also provides a general strategy for mechanical enhancement for protein-based hydrogels. Based on the denatured-crosslinking hydrogelation method discussed in Chapter 4, experimental work in Chapter 5 further improves the tensile modulus and the toughness of the (FL)8 polyprotein hydrogel by employing conjoined double-network method. A natural polymer, alginate is used as a secondary network, which provides ionic crosslinkable blocks. The additional ionic crosslinks not only enhance the network density of the hydrogel, but also introduce a favorable energy dissipative mechanism. To covalently connect the protein and the alginate networks, the alginate is derivatized with tyramine using aqueous-phase carbodiimide activation 168 chemistry. By controlling the modification density of tyramine, we were able to tune the stiffness and extensibility of the protein-alginate hybrid hydrogels over a wide range. The resulting hybrid gels showed improved Young’s modulus (up to ~1.02 MPa), extensibility (up to 230%), and toughness (up to ~1.02 MJ/m3). This work broadens the range of mechanical properties that protein-based biomaterials can reach, providing a promising candidate for potential model system for cartilage tissue engineering. Chapter 6 demonstrates a novel cell-adhesive, entirely protein-based hydrogel that acts as an ECM for mechanochemical-induced cell spreading studies. By using a dynamic redox-responsive MEP protein as a building block, the mechanical and physical properties of the MEP hydrogels can be reversibly tuned by redox states. The MEP hydrogel exhibits a 3-fold change in Young’s modulus between its two conformational states. In response to the cyclic change of the hydrogel’s Young’s modulus, the morphology of HLF cells can be dynamically regulated in a fully reversible fashion, demonstrating the great potential utility of such dynamic protein hydrogels in cellular mechano-biology studies and regenerative biomedical applications. Overall, the novel materials developed in this thesis could potentially lead to model systems for mechanobiology studies and tissue engineering. Mechanical signals induce changes in functions of cells and tissues. For example, mechanical stress on cartilage and bones allows for body movement. Similarly, senses of hearing and touching are caused by forces.385 Thus, it is crucial to establish models and tissue substitutes for mechanobiological and biomedical research. Hydrogels are certainly suitable materials for cell scaffolds or tissue engineering, due to the resemblant elasticity and high water content. However, polymer hydrogels could lack dynamic features and biocompatibility as ECMs, and living tissues are mostly made of protein-based matrices. Therefore, protein-based hydrogels are promising candidates with designable 169 mechanical features. This thesis presents a semiquantitative tool to precise control over mechanical performance of protein-based hydrogels, leading to a significant mechanical breakthrough of protein-based hydrogels. This make it possible to use protein-based hydrogels as potential model systems for wide biomedical applications. 7.2 Outlook 7.2.1 Exploring smart protein-based hydrogels with actuation functions Recently, soft actuators have received particular attention and shown promising potentials in many fields, such as soft robots, artificial muscles, and sensors.412-413 As discussed in Chapter 3, we demonstrated an example of a protein hydrogel engine by exploiting mechanical energy generated from protein conformational change upon denaturation/renaturation. In principle, hydrogels with such “smart” behaviors are related to dynamic and reversible changes in the crosslinked networks which arise upon triggers from environmental stimuli, such as changes in ionic strength, heat, pH and the presence or absence of chemical reagents. In response to external stimuli, deformation and asymmetric swelling of a hydrogel can generate an actuation force, allowing for movement from a soft actuator. The Popa group reported cation-induced BSA hydrogels behaving as actuators.175 Several exposed amino acids of BSA could bind cations, such as histidine, tryptophan and cysteine residues. The hydrogels were programmed into varied shapes with significant changes in stiffness (of up to 17-fold), induced by adsorption of Zn2+ or Cu2+ cations. Upon removal of the cations from the hydrogels, they showed that these programmed materials could then recover their original shape, generating actuation forces. The generation of actuation was produced upon reversible ionic crosslinks. However, the primary building blocks, 170 which consist of BSA folded domains, remained unchanged, leading to relatively weak actuating responses. To design an actuator with an efficient energy output, a variety of functional proteins can be designed into gel networks to fulfill the purpose, such as conformational switchable domains, reversible protein-protein interactions, and proteins with physical crosslinking groups. The primary shape of an actuator can be fabricated by employing covalent crosslinking. Additional dynamic and reversible protein domains can act as actuating engines. The larger the change in the networks, the higher the efficiency in actuation. With the help of the semi-quantitative tool introduced in Chapter 2, it could be possible to engineer a protein actuator with programmable and controlled networks and swelling behaviors, allowing for efficient actuation. 7.2.2 Designing dynamic protein-based hydrogels supporting cell mechano-biology studies The cell-ECM interactions via receptors on cell surface result in intracellular signaling cascades that modulate a series of cell behaviors, including cell adhesion, migration, proliferation, and differentiation.414-415 In addition to the biochemical induction factors, cells also sense mechanical cues from the surrounding matrices. For instance, mesenchymal stem cells have been shown to be sensitive to the stiffness of the local microenvironment. Mesenchymal stem cells can recognize and adapt to the elasticity of their substrate, so that on brain-like soft matrices (elastic modulus, E < 1 kPa) they develop neural characteristics, while on rigid ones (E > 30 kPa) they tend toward myogenesis or osteogenesis.197, 388, 416 Hydrogels are the most extensively studied mimics to natural ECMs and cell scaffolds in vitro, due to their biocompatibility, controllable network architecture and high water content.417-418 For example, Se Heang Oh and colleagues designed polyvinyl alcohol (PVA)/hyaluronic acid 171 (HA) hydrogel with a wide-range stiffness gradient to investigate stem cell differentiation behavior.419 From 2D cell cultures using the stiffness-gradient PVA/HA hydrogel matrices, it was observed that human bone marrow mesenchymal stem cells have favorable stiffness ranges for induction of differentiation into specific cell types (~20 kPa for a nerve cell, ~40 kPa for a muscle cell, ~80 kPa for chondrocytes, and ~190 kPa for osteoblasts). Although artificial ECMs made of biocompatible polymers such as gelatin and collagen have been successfully developed for cell culture, the static elasticity of these materials limits the dynamic mechano-biology studies of cell-ECM interactions. As shown in Chapter 6, we provided a potential solution, using a novel protein-based hydrogel as a dynamic ECM. This dynamic protein gel not only supports cell adhesion and proliferation, but also regulates the cell spreading behavior in a reversible manner. However, the dynamic change of the gel stiffness only covers a range from ~6 to ~20 kPa, limiting the use as matrices for different cell types. Thus, a future direction is to design a dynamic protein-based hydrogels with larger responsive range of mechanical properties. Designable polyprotein constructs containing dynamic domains allow us to regulate protein contour length in gel networks. Upon precise control over elasticity of protein substrates with dynamic features, it is possible to use the protein-based hydrogels for mechano-signaling studies for various cell types, such as mechanical induced differentiation and self-renewal of stem cells. 7.2.3 Designing recombinant protein-based hydrogels as 3D ECMs Cell culture on 2D surfaces has predominated and greatly contributed to in vitro studies in cell biology. However, under some circumstances, 2D culture shows limitations in mimicking natural ECMs. For example, 2D models cannot accurately reproduce in vivo microenvironments for human breast epithelial cells, where they behave like tumor cells, whereas in 3D culture they 172 revert to normal growth behavior.420 Indeed, many researchers have shown that 3D cell cultures are more accurate in mimicking realistic biochemical and biomechanical microenvironments of natural ECMs.421-422 For instance, the interactions between β1-integrin and epidermal growth factor receptor signaling were identified in 3D culture but not found in 2D.423 Significant progress has been achieved in 3D cell encapsulation using polymeric soft materials. However, it still remains challenging to develop polymer matrices that promote preferable cell fates for tissue regeneration, which take into account biochemical properties, cell adherence, matrix elasticity, and macroporosity. This thesis has shown evidence that protein-based hydrogels are promising candidates for 3D cell culture, due to excellent biocompatibility, designable biochemical components and mechanical properties. Currently, the photo-crosslinked hydrogels we studied involved the use of the chemical reagent APS (50 mM), which is cytotoxic for 3D cell encapsulation. Although it was shown that cytotoxicity can be minimized by decreasing concentration of APS to a biocompatible level (> 5 mM),285 insufficient crosslinking reagent may sacrifice the stiffness of the protein substrates. As a future direction, alternative crosslinking strategies could be introduced to fabricate biocompatible protein 3D gel matrices. As mentioned in Section 1.1.1.3, the recombinant proteins containing SpyTag and SpyCatcher successfully formed mechanical stable hydrogels for the 3D cell encapsulation application.58 In addition to this approach, it could be worth exploring how genetically encoded protein fragment reconstitution, self-assembling peptides, and cysteine containing domains could be used as hydrogel building blocks for 3D culture. These crosslinking strategies could enable gelation of biocompatible hydrogels under physiological conditions. 173 7.2.4 Engineering protein-based biomaterials for bioprinting As described in Chapter 5, the protein-alginate hybrid network not only provides an approach for mechanical enhancement, but could also lead to the development of printable materials. Bioprinting technology aims at accurate assembly of living and non-living physiological materials. The mechanized and organized cell-laden complex can mimic the microenvironment precisely for further tissue maturation.424 The bioprinting approach has been driving major innovations in many applications, such as regenerative medicine, tissue engineering, and therapeutic discovery.425-426 Bioink formulation and process compatibility play key roles for bioprinting. Bioinks are typically composed of biocompatible ink solution and cells. For example, the Lee group developed a kidney ECM-derived bioink for 3D bioprinting of a human kidney prototype. Porcine whole kidney was decellularized and dissolved in acid. Then the kidney ECM-derived bioink was formulated by chemical modification with rheological properties and printability for the printing process. The results showed high viability and maturation of bioprinted human kidney cells. In addition, the bioprinted renal constructs exhibited structural and functional similarities with native renal tissue.426 This work showed great potential of 3D bioprinting of organs and tissue regeneration. However, the bioink still required living organs and complicated processing methods. Printable biopolymers are more sustainable and accessible materials to fulfill the purpose. In many cases, hydrogels with multiple crosslinking mechanisms are used, where the primary crosslinking provides printability, while the secondary crosslinking provides shape fidelity after printing. Alginate has been used extensively in bioprinting, due to its rapid gelation, low toxicity and high availability.427 Afghah et al. used a cell laden alginate gel-based nanoclay composite to bioprint representative structures, for complex and hollow tissues.428 Instant ionic 174 crosslinking of alginate enables printability of the pregel mixture with adaptable viscosity. 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Thesis/Dissertation
eng
Chemistry
Vancouver : University of British Columbia Library
Attribution-NonCommercial-NoDerivatives 4.0 International
Graduate
Engineering globular protein-based hydrogels with tunable mechanical properties
Text