Gellan Gum Based Thiol-Ene Hydrogels with Tunable Properties for Use As Tissue Engineering Scaffolds Zihao Xu Iowa State University

Iowa State University Capstones, Theses and Graduate Theses and Dissertations Dissertations

2018 Gellan gum based thiol-ene hydrogels with tunable properties for use as tissue engineering scaffolds Zihao Xu Iowa State University

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Recommended Citation Xu, Zihao, "Gellan gum based thiol-ene hydrogels with tunable properties for use as tissue engineering scaffolds" (2018). Graduate Theses and Dissertations. 16493. https://lib.dr.iastate.edu/etd/16493

This Thesis is brought to you for free and open access by the Iowa State University Capstones, Theses and Dissertations at Iowa State University Digital Repository. It has been accepted for inclusion in Graduate Theses and Dissertations by an authorized administrator of Iowa State University Digital Repository. For more information, please contact [email protected]. Gellan gum based thiol-ene hydrogels with tunable properties for use as tissue engineering scaffolds

Zihao Xu

A thesis submitted to the graduate faculty

in partial fulfillment of the requirements for the degree of

MASTER OF SCIENCE

Major: Materials Science and Engineering

Program of Study Committee: Kaitlin Bratlie, Co-major Professor Shan Jiang, Co-major Professor Surya Mallapragada

The student author, whose presentation of the scholarship herein was approved by the program of study committee, is solely responsible for the content of this thesis. The Graduate College will ensure this thesis is globally accessible and will not permit alterations after a degree is conferred.

Iowa State University

Ames, Iowa

2018

TABLE OF CONTENTS

Page

LIST OF FIGURES ...... iv

LIST OF TABLES ...... v

NOMENCLATURE ...... vi

ACKNOWLEDGMENTS ...... vii

ABSTRACT ...... viii

CHAPTER 1. CLICK CHEMISTRY HYDROGELS IN TISSUE ENGINEERING APPLICATIONS ...... 1 1. Introduction ...... 1 2. Hydrogel crosslinking mechanisms ...... 2 2.1. Click chemistry crosslinking ...... 3 2.1.1. Diels-Alder reaction ...... 5 2.1.2. Azide and alkyne cycloaddition ...... 7 2.1.3. Chain growth mechanism ...... 10 2.1.4. Thiol-ene chemistry ...... 12 3. Click chemistry hydrogels ...... 18 3.1. Synthetic polymer based hydrogels ...... 18 3.1.1. PEG ...... 19 3.1.2. PVA ...... 22 3.2. Natural polymer based hydrogels ...... 23 3.2.1. Chitosan ...... 23 3.2.2. Gelatin ...... 25 3.2.3. Hyaluronic acid hydrogels ...... 27 4. Hydrogels fabricated using click chemistry as tissue scaffolds ...... 28 5. Conclusions and future perspectives ...... 30

CHAPTER 2. GELLAN GUM BASED THIOL-ENE HYDROGELS WITH TUNABLE PROPERTIES FOR USE AS TISSUE ENGINEERING SCAFFOLDS...... 32 1. Introduction ...... 32 2. Methods and materials ...... 34 2.1. Materials ...... 34 2.2. Gellan gum modification ...... 34 2.3. Characterization of modification of gellan gum ...... 34 2.4. Synthesis of hydrogels ...... 35 2.5. Compressive modulus ...... 36 2.6. Swelling ratio ...... 36 2.7. In vitro degradation profile ...... 36 2.8. Cell Culture ...... 37 2.9. Cell Proliferation...... 37 iii

2.10. Statistics and data analysis ...... 38 3. Results and discussion ...... 38 3.1. Gellan gum modification and characterization ...... 38 3.2. Compression modulus...... 39 3.3. Swelling ratio ...... 43 3.4. In vitro degradation ...... 46 3.5. Cell Proliferation...... 48 4. Conclusion ...... 51

CHAPTER 3. EFFECT OF SUBSTRATE HYDROPHOBICITY ON CELL PROLIFERATION AND POLARIZATION ON GELLAN GUM BASED HYDROGEL...... 53 1. Introduction ...... 53 2. Methods and materials ...... 55 2.1. Materials ...... 55 2.2. Hydrophobic modification ...... 55 2.3. Characterization of the modified gellan gum ...... 56 2.4. Hydrogel fabrication ...... 56 2.5. Compression modulus...... 57 2.6. Swelling ...... 57 2.7. Water contact angle ...... 57 2.8. Cell culture ...... 58 2.9. Cell proliferation ...... 58 2.10. Griess assay ...... 59 2.11. Urea assay ...... 59 2.12. Statistics and data analysis ...... 59 3. Results and discussion ...... 60 3.1. LMGG hydrophobic modification and characterization ...... 60 3.2. Compression modulus...... 61 3.3. Swelling of hydrogels ...... 63 3.4. Water contact angle ...... 64 3.5. Cell proliferation ...... 66 3.6. Cell polarization ...... 68 4. Conclusion ...... 72

REFERENCES ...... 73 iv

LIST OF FIGURES

Page

Figure 1. Crosslinking mechanism of representative functional groups ...... 3

Figure 2. Effect of PEGDA molecular weight and bi-functional thiol content on equilibrium shear moduli...... 16

Figure 3. Illustration of defects in dithiol crosslinking and effect of excessive thiol during thiol-ene reaction...... 17

Figure 4. Synthesis and characterization of methacrylated gellan gum...... 39

Figure 5. Mechanical properties of modified gels. Young’s moduli of low and high gellan gum hydrogels crosslinked through the different mechanisms...... 40

Figure 6. Swelling behavior of gellan gum hydrogels. Hydrogels crosslinked through the different mechanisms were swelled in PBS...... 44

Figure 7. Gellan gum hydrogel degradation ...... 47

Figure 8. Cytocompatibility of gellan gum hydrogels. NIH/3T3 fibroblasts were seeded on hydrogels and tissue culture plastic (control)...... 50

Figure 9. Schematic illustration of the synthesis and characterization of hydrophobic modified LMGG...... 61

Figure 10. Mechanical properties of LMGG and modified LMGG hydrogels...... 62

Figure 11. Swelling behavior of LMGG and modified LMGG hydrogels ...... 64

Figure 12. Water contact angle measurement of LMGG and modified LMGG hydrogels...... 65

Figure 13. Cell proliferation of macrophages seeded on hydrogels...... 67

Figure 14. Representative stitched fluorescent images of live (green) and dead (red) macrophage cells growth and proliferation on hydrogels after 48 h culture ...... 68

Figure 15. Cytokine secretion of macrophages cultured on hydrogels...... 71 v

LIST OF TABLES

Page

Table 1. Summary of crosslinking mechanisms and properties ...... 4

Table 2. Summary of naturally derived polymer based hydrogels crosslinked through click chemistry orchain growth polymerization ...... 13

Table 3 Summary of synthetic polymer based hydrogels crosslinked through click chemistry or chain growth polymerization ...... 21

NOMENCLATURE

CM Dulbecco’s modified Eagle’s medium

DTT Dithiothreitol

FTIR Fourier transform infrared spectroscopy

GG Gellan gum

NMR Nuclear magnetic resonance spectroscopy

PBS Phosphate Buffered Saline

TCP Tissue culture plastic

vii

ACKNOWLEDGMENTS

I would like to thank my major professor, Kaitlin Bratlie, and my committee members, Prof. Jiang and Prof. Mallapragada, for their guidance and support throughout the course of this research.

In addition, I would also like to thank my parents who support me both mentally and financially during my master program. Moreover, I am really grateful to all my colleagues as well. Hannah Bygd and Lida Zhu helped me a lot in starting this project.

Zhuqing Li, Anuraag Boddupalli, Yanhua Huang provided me valuable ideas in lab discussion. These excellent colleagues made my master program a wonderful experience.

This work was founded by National Science Foundation under Grant No. CBET

1227867, the Roy J. Carver Charitable Trust Grant No. 13-4265, the Mike and Denise

Mack faculty fellowship to KB, and the start-up funding to SJ.

viii

ABSTRACT

Gellan gum is a naturally occurring polymer that can crosslink in the presence of divalent cations to form biocompatible hydrogels. However, physically crosslinked gellan gum hydrogels lose stability under physiological conditions, which substantially limits the applications of these hydrogels in vivo. In order to improve the mechanical strength, we incorporated methacrylate into gellan gum and chemically crosslinked the hydrogel through three polymerization methods: step growth through thiol-ene photoclick chemistry, chain growth via photopolymerization, and mixed model in which both mechanisms were employed. Methacrylation was confirmed and quantified by proton nuclear magnetic resonance (1H NMR) and Fourier transform infrared spectroscopy

(FTIR). The mechanical property and chemistry of the crosslinked gels were systematically explored by varying the reaction conditions. The swelling ratios of the hydrogels were correlated with the compression moduli and affected by the addition of calcium. In vitro enzymatic degradation rate was found dependent on the degree of methacrylation. NIH/3T3 fibroblast cell proliferation and morphology were related to substrate stiffness with high stiffness leading generally to higher proliferation. The proliferation is further affected by the thiol-ene ratios. We then further modified methacrylate Gellan gum with alkane bromide to increase hydrophobicity. Cell attachment on resultant hydrogels were assessed and imaged. Cytokine release was also measured with comparison to pristine methacrylated Gellan gum based hydrogels. The results suggest that a hydrogel platform based on gellan gum can offer versatile chemical modifications and tunable mechanical properties for a variety of biomaterials applications, such as the wound healing scaffold. 1

CHAPTER 1. CLICK CHEMISTRY HYDROGELS IN TISSUE ENGINEERING APPLICATIONS

1. Introduction

Tissue engineering is an evolving field in which the areas of biology, engineering, and materials science work symbiotically to develop new scaffolds for treating diseased, damaged, or absent tissue. This growing area has been evaluated at $17 billion and is expected to reach $56.9 billion by 2019.1 Hydrogels are particularly apt candidates for tissue engineered scaffolds owing to their tunable materials properties and their similarity to the extracellular matrix (ECM).2 Both synthetic and naturally derived polymers can be used to fabricate hydrogels. Commonly used synthetic polymers include poly (ethylene glycol)

(PEG) and poly (vinyl alcohol) (PVA) while naturally derived polymers can be polysaccharide-based (e.g., hyaluronic acid, chitosan, and gellan gum), protein-based (e.g., collagen, fibrin, and gelatin), or a combination of proteins and polysaccharides (proteoglycan and glycoprotein).2,3 Synthetic polymers are highly plastic, which enables them to be processed and sterilize easily due to their resistance to elevated temperatures. The degradation rate and mechanical properties of synthetic polymers can be tuned by modifying the Flory interaction parameter by changing the repeat unit of the polymer or through changing monomer ratios. Crosslinking density is another engineering knob that can be adjusted to alter hydrogel physical properties.4,5 Degradation byproducts of synthetic polymers are typically acidic and can lead to secondary acute inflammation.6 Naturally derived polymers, on the other hand, provide adhesive surfaces for cell attachment and their degradation products are generally non-toxic, but are limited by poor mechanical performance and accelerated degradation.3

Hydrogels can be crosslinked through chemical bonds; permanent or temporary physical entanglements; or secondary interactions, such as hydrogen bonds.2 Covalent crosslinks are generally stable in physiological environments.7 Changing the ratios of reactants in synthetic hydrogels enables tunable mechanical properties.8 However, toxic chemicals used in fabricating these hydrogels can reduce their biocompatibility.2 Physically crosslinked hydrogels exhibit reversible sol-gel transitions. These hydrogels are often sensitive to temperature or ion concentration. For thermally sensitive hydrogels, hydrogen bonds are broken at elevated temperatures, which leads to a phase transition. Under physiological conditions, monovalent ions can replace the stronger multivalent ionic interactions, resulting in a loss of stability.7 As a result, the diversity of chemical and physical properties of chemically crosslinked hydrogels meet the requirements of tissue scaffolds, such as appropriate cellular response, structural integrity, biodegradability, biocompatibility, and solute transport.2

Several chemical crosslinking methods have been developed for the purpose of fabricating hydrogels for tissue engineering. Herein we describe mechanisms, features, and current applications of four click based crosslinking methods, and provide insight on synthetic and natural polymer hydrogels fabricated using click chemistry.

2. Hydrogel crosslinking mechanisms

As mentioned in previous section, physically crosslinked hydrogels lose their stability in physiological conditions while chemical crosslinking provides broad control over resultant hydrogel properties. In addition to designing hydrogel physical and chemical properties, their gelation point or crosslinking time is crucial in their fabrication. Hydrogels fabricated in situ must be rapidly formed to minimize surgical invasion,9 to adapt to complicated defect sites,10 and to be injected at the desired site.11 3

2.1. Click chemistry crosslinking

Click chemistry is a class of reactions first described by Sharpless that are able to rapidly fabricate hydrogels.12 His group defined a stringent set of criteria for classifying a process as a click reaction. Click reactions must be “modular, wide in scope, give very high yields, generate only inoffensive byproducts that can be removed by nonchromatographic methods, and be stereospecific.”13 Click reactions have a high thermodynamic driving force

(>20 kcal/mol), which enables these reactions to proceed rapidly to completion with high selectivity.13 These features make click reactions ideally suited for biological applications.12

The following sections will discuss three methods of crosslinking hydrogels using click chemistry: Diels-Alder reaction, azide and alkyne cycloaddition, and thiol-ene reactions.

Along with photoinitiated chain growth polymerization, which also proceeds rapidly. (Table

1 and Figure 1)

Figure 1. Crosslinking mechanism of representative functional groups (A) Diels-Alder reaction, (B) CuAAC and SPAAC reaction, (C) Photoinitiated chain growth and (D) Thiol- ene

Table 1. Summary of crosslinking mechanisms and properties

Crosslinking Category Initiation Functional groups conjugates Advantages Disadvantages mechanism Diels-Alder  No coupling agent or  Long gelation time (1.5-24h)16 Click chemistry Simple mixing Diene+alkene catalyst14  Inability to be injected in vivo17 e.g. furan and maleimide  Thermally reversible15  Poor solubility of functional  Easy to incorporate drugs groups into network12 Azide-alkyne Cu(I) catalyzed  Shortened gelation time For CuAAC reactions cycloaddition Click chemistry (CuAAC) Azide+alkyne  injectable18  Copper is cytotoxic19 e.g. azide and alkyne end group  Poor diffusion of copper causes ring strain promoted Azide+alkyne ring  No cytotoxic metal catalyst heterogeneity in hydrogel (SPAAC) e.g. azide and cyclooctyne  Shortened gelling point structure20  Injectable19 For all AAC reactions  Spatial and temporal control of hydrogel formation is not

achievable 4  Gelation time is relatively long for cell encapsulation Thiol-ene  Rapid reaction rate21  Radicals in initiation step are chemistry Click chemistry Free-radicals Alkene+thiol  Spatial and temporal control toxic to encapsulated cells2224 e.g. methacrylate group and of gel formation22  Thiyl radicals are insufficient to dithiothreitol  Insensitivity to oxygen23 initiate polymerization, need addition of small molecule (NVP) to accelerate25 Photoinitiated  Rapid kinetics  Reduced biocompatibility when chain growth Free radical Free-radicals Alkene  Spatial and temporal control using UV sensitive initiator24 polymerization polymerization e.g. methacrylate of gel formation26  Inhibited by oxygen26  Produce highly heterogeneous network27

2.1.1. Diels-Alder reaction

Diels-Alder cycloaddition is a click reaction in which a diene is conjugated to a substituted alkene without the addition of a catalyst.28,29 This reaction is widely used in biomedical applications.29 Under aqueous condition Diels-Alder chemistry can be used to fabricate hydrogels for tissue engineering applications due to its site-specific and efficient reactivity, biocompatibility, and irreversibility at body temperature.30

The gelation time of Diels-Alder reactions can be influenced by pH. The most commonly used diene and substituted alkene for synthesizing hydrogels via Diels-Alder reaction are furan and maleimide.31–34 Hydrogels formed using furan and maleimide functionalized 8-armed PEG exhibited an increase in gelation time when pH was increased

(87.8 min at pH 3 and 97.2 min at pH 7.4). This trend discontinued when the pH was increased above 7.4, which enhances hydrolysis of the maleimide group and prevents gelation.35 These relatively long gelation times precludes using these gels in an injectable format.17 To address this problem, changing the ratios of furan and maleimide30,36 and using enzymatic crosslinking17 were employed. Enzymatic crosslinking involves a two-step synthetic method. The first step involved mixing a furan-grafted hyaluronic acid (HA) that was further functionalized by tyramine with maleimide-functionalized PEG and hydrogen peroxide/horse radish peroxidase to trigger enzymatic crosslinking within 350 seconds. The second step was the Diels-Alder reaction, which was completed over the following 24 h.

Cytocompatibility of the HA/PEG hydrogel was tested on mouse embryonic carcinoma- derived clonal cell line ATD5. Encapsulated cells remained viable after 7 days in culture.17

The same group optimized the gelation time by varying the degree of furan substitution and using different molar ratios of furan:maleimide. A high degree of HA functionalization with 6 furan (75.9  5.4%) and an equimolar furan:maleimide ratio yielded the shortest gelation time at 37C (51 min).36

The mechanical properties of Diels-Alder hydrogels can be adjusted by forming interpenetrating networks and changing the molar ratios of the diene and alkene.3,33 Furan- functionalized HA and gelatin were mixed, then reacted with maleimide-terminated PEG to form a pre-hydrogel. Subsequent conjugation with chondroitin through EDC/NHS amidation resulted in a triple component interpenetrating hydrogel with a higher compression modulus

(2.2 kPa at furan:maleimide 5:1) than the two component pre-hydrogel (1.3 kPa at furan:maleimide 5:1). The molar ratio of furan:maleimide was also influential on the hydrogel compression modulus. Lower furan to maleimide ratio yielded a stiffer final hydrogel (2.2 kPa at furan:maleimide 5:1 and 1.6 kPa furan:maleimide 10:1).3 In another study, furan and HA functionalized with a similar amount of maleimide were mixed at different volume ratios and crosslinked through Diels-Alder chemistry. The compressive moduli of resultant hydrogels were dependent on the volume ratios of furan and maleimide modified HA. The highest modulus was ~20 kPa when furan to maleimide was 1:2. No statistical difference was observed for 2:1 and 1:1 furan:maleimide ratios hydrogels (all around 12-14 kPa).30

The advantage of Diels-Alder reaction in crosslinking hydrogels for tissue engineering application is its elimination of coupling agent or catalyst, which may introduce cytotoxic small molecules.32 In addition to the site-specific crosslinking and the thermal reversibility of this reaction, drugs can be incorporated into the hydrogel network for sustained release.12,30,37 The main obstacles of using Diels-Alder chemistry to fabricate hydrogels are the long gelation time, inability to be injected in vivo,17 and poor solubility of 7 the functional groups.38 These draw backs restrict application of the Diels-Alder reaction under physiological conditions.38 In order to address this issue, a more efficient crosslinking mechanism, azide-alkyne cycloaddition, was used to fabricate hydrogels in tissue engineering applications.

2.1.2. Azide and alkyne cycloaddition

Azide-alkyne cycloaddition (AAC), which produces 1, 2, 3-triazoles, is an effective approach for preparing well defined polymer networks due to the high chemical potential energy of the reaction. The resulting polymer is stable in the presence of common solvents and its biorthogonal nature renders it biologically inert.39,40 One of the drawbacks of azide- alkyne cycloaddition is that its relatively high activation barrier necessitates prolonged heating for unactivated alkynes.41 Copper (I) or ruthenium (II) catalyst can make the alkynes electron deficient, which can accelerate the low reaction rate. Since Ru-catalyzed azide- alkyne cycloaddition (RuAAC) is more sensitive to solvents than copper-catalyzed cycloaddition (CuAAC), the latter is more commonly used.41 Sun et al. used tetrakis (2- propynyloxymethyl) methane as tetra-alkyne crosslinker and PEG-diazide modified with the photosensitive moiety spiropyran to synthesize stiff (49.8 MPa break stress under compression in 70% water content) hydrogels. The gelation point was reached within 80 seconds and was continued for another 24 h to ensure complete conversion.42

The fast reaction rate of CuAAC results in rapid increases in regional viscosity and restricts diffusion of copper (I) into unreacted regions, which results in inhomogeneous hydrogel networks. This issue was solved through in situ reduction of copper (II).39,43 A copper (II) containing precursor solution was homogeneously mixed with the thermal initiator azobisisobutyronitrile (AIBN), which decomposes at 65C and reduces copper (II) to 8 copper (I). The reaction proceeded for 24 h and unreacted azide and alkyne groups were not detected using FTIR spectroscopy throughout the gel cylinder. Fabricating the more homogeneous hydrogel also improved the mechanical properties. The 70% water content gels had a higher break stress of 60.5 MPa under compression compared with 1.28 and 1.17 MPa for gels fabricated using traditional CuSO4/NaSac CuAAC and CuBr/N, N, N’, N’’, N’’- pentamethyldiethylenetriamine CuAAC, in which the catalyst was poorly dispersed.39 Cell viability of HL-7702 human liver cells cultured on these hydrogels for 24 h was found to exceed 90%.39

Relatively high activation temperatures and long gelation times pose potential barriers for in situ fabrication of tissue scaffolds using CuAAC. Photoinitiated copper (II) reduction is a promising way to provide temporal and spatial control of CuAAC hydrogel formation.43,44

The visible light photoinitiator Irgacure 819 was used to trigger the reduction of the CuCl2/

N, N, N′, N′′, N′′-pentamethyldiethylenetriamine complex to polymerize resin. This method resulted in 70% monomer conversion after 95 minutes of irradiation.44 Anseth and coworkers used Irgacure 2959 to reduce copper (II) in situ and trigger CuAAC conjugation to synthesize

PEG hydrogels functionalized with azide and alkyne. These hydrogels were formed within 4 minutes, as determined by rheology.43 This method can be used to pattern hydrogels. Azide functionalized fluorescent molecules were conjugated to the free alkyne on the hydrogels.

Patterns were formed 50 s after UV irradiation.43

In addition to synthetic challenges, copper is cytotoxic and increasing copper concentration decreases cell viability.39 As such, metal-free azide-alkyne cycloaddition was developed to improve the biocompatibility of resultant hydrogels. Strain promoted azide- alkyne cycloaddition (SPAAC) is a promising method which uses ring strain to open the 9 carbon-carbon triple bond rather than a metal catalyst.10,11,19,45–60 Among those strain bearing alkyne rings, 4-dibenzocyclooctynol,10,48,51,53,55,59,60 oxanorbornadiene,19 and cyclooctyne,11,45–47,49,52,56–58,61,62 are most commonly used due to their optimized synthesis and grafting procedure11,59. Synthetic polymer based SPAAC hydrogels can be fabricated by mixing 4-dibenzocyclooctynol modified PEG and glycerol exytholate triazide. The gel point was reached within 5 minutes through gentle mixing and an equilibrium storage modulus was reached after 2.5 h.10 Naturally derived polymers have also been crosslinked using SPAAC.

By reacting azide functionalized HA and oxanorbornadiene grafted chitosan at equimolar oxanorbornadiene:azide ratios, hydrogels were formed in 23 minutes and the resultant gel had a higher compressive modulus than other oxanorbornadiene:azide ratios (41 kPa at 2% w/v polymer at equimolar ratio and 6 kPa at oxanorbornenadiene:azide 9:1 molar ratio).19

Cytocompatibility of equimolar SPAAC hydrogel was verified by in situ injection and gel formation in Balb/c mice. No significant inflammatory reactions were observed over 7 days.19.

While gelation time of hydrogel fabricated through AAC reaction is shortened when compared to Diels-Alder reaction, spatial and temporal control of hydrogel formation is not achievable under this mechanism. Once the reagents in a AAC reaction spatially approach one another, the reaction is triggered.63 Spatial and temporal control enables initiation of a reaction on demand, which is critical when materials must be prepared after mixing of components but before a reactions occurs.64 Moreover, hydrogels used for tissue engineering applications are oftentimes required to be fabricated with controlled shape and size and even more rapid gelation kinetics such that encapsulated cell remain viable during this process.2665

Photo-initiated crosslinking of hydrogels was subsequently developed to provide control over 10 when and where hydrogel crosslink. With the use of a photoinitiator, gelation reaction can be triggered immediately by light and completed within minutes.25

2.1.3. Chain growth mechanism

Chain growth polymerized hydrogels represent a class of hydrogels that are crosslinked via free radical polymerization of vinyl groups. When initiated by a photoinitiator, chain growth has several advantages over the above discussed AAC reaction in fabricating hydrogels. First, chain growth polymerization allows spatiotemporal control of gelation kinetics and more rapid encapsulation of cells, which imparts minimal damage to the cells.66–68 Second, gelation time of photo crosslinked hydrogels ranges from seconds to a few minutes at physiological temperatures and has a low activation energy.69 These features make chain growth crosslinking a viable method for fabricating hydrogels for use in tissue engineering applications.66

Naturally derived polymers resemble the ECM in terms of chemical versatility, controlled degradability, and biocompatibility, all of which make these polymers suitable tissue engineering scaffolds.70 Among these, kappa-carrageenan,70 chitosan,66,68,69,71 gelatin,67,72 gellan gum,73 fucoidan74, laminarin,75 chitin,76 alginate,77 and hyaluronic acid65 have been successfully grafted with carbon-carbon double bond bearing moieties and have formed hydrogels via photocrosslinking. (Table 2) Kappa carrageenan has been grafted with methacrylate groups by reacting with methacrylic anhydride Hydrogels were subsequently fabricated by exposure to UV light for 40 s using Irgacure 2959 as the photoinitiator.

Resultant hydrogels exhibited high swelling ratios (~65 in Dulbecco’s phosphate buffered saline and ~75 in cell culture medium). Encapsulated NIH/3T3 cells in these gels showed high viability for long time periods of up to 21 days.70 Gellan gum can be methacrylated in a similar fashion and hydrogels were photocrosslinked using methyl benzoylformate as the 11 photoinitiator. In this method, hydrogels were formed within 6 minutes. Encapsulated L929 fibroblasts dispersed evenly throughout the gels remained viable for up to 21 days.73

However, UV exposure may pose potential cell damage and reduce biocompatibility of the resultant hydrogels. This issue was solved by substituting UV sensitive initiators with a redox initiation system. Glycidyl methacrylated alginate was crosslinked via a persulfate-N’, N’,

N’, N’-tetramethylethylenediamine redox system. The resultant hydrogels were fabricated at body temperature within 20 minutes. Encapsulated human umbilical vein endothelial cells

(HUVEC) showed over 80% cell viability after 48 h of incubation.77

The mechanical properties of hydrogels crosslinked through chain growth polymerization can be tuned through the incorporation of a secondary network.

Copolymerizing natural polymers with either other type natural polymers, 65,67,78,79 or with synthetic polymer networks66,68,72 can introduce interpenetrating networks to the system. For example, PEG-diacrylate (PEGDA) and methacrylated gelatin were used to fabricate a composite hydrogel. Increasing the PEGDA content increased hydrogel stiffness and stability in 2.5 U/ml collagenase.72 When gelatin content was fixed at 10%, increasing PEG content from 5% to 10% increased compressive modulus from ~0.02 MPa to 0.04 MPa. Hydrogels completely degraded within 24 h for 5% PEG but remained at ~50% weight for 10% PEG hydrogels.80 In addition to natural and synthetic composite polymers, double network hydrogels can be fabricated with two naturally derived polymers. Gelatin and gellan gum were methacrylated through using methacrylic anhydride and formed a double network hydrogel. This double network hydrogel was much stiffer than the single network hydrogel

(~110 kPa for double network hydrogel, ~50 kPa for gelatin only, and ~10 kPa for gellan gum only hydrogel) .81 12

Unfortunately, classical photopolymerization is plagued by several critical problems: inhibition by oxygen; complicated volume relaxation and stress development; complex polymerization kinetics; and the formation of highly heterogeneous polymers/networks.82,83

These drawbacks restrict fabrication and precise control over the properties of chain growth hydrogels. Many of these drawbacks can be overcome using thiol-ene click chemistry.

2.1.4. Thiol-ene chemistry

Thiol-ene chemistry refers to a radical mediated reaction between thiol and alkene groups. This class of reaction has several attributes: high reaction yield, rapid reaction rate, low concentration of initiator required for the initiation step, insensitivity to oxygen or water, biocompatibility, and orthogonal polymeric network formation.82,86–88 The mechanism for this reaction is a combination of step growth and chain growth polymerizations. Radicals are first generated by the initiator, which abstracts protons from the thiol groups to form thiyl radicals. This thiyl radical activates the carbon-carbon double bond by forming a carbon based radical. The carbon based radical can either undergo a chain transfer reaction with another thiol group to generate a new thiyl radical or propagate through carbon-carbon double bonds.89 By using multi-functionalized thiol molecules as the crosslinker, carbon- carbon double bond bearing molecules can be covalently linked.

Table 2. Summary of naturally derived polymer based hydrogels crosslinked through click chemistry orchain growth polymerization

Natural Origin Modification methods Initiation Cytocompatibility polymers -carrageenan Seaweed Methacrylic acid, substitute hydroxyl groups Irgacure2959, UV Encapsulated hMSC, 21d, 75% viable70 Chitosan Chitin shells of Maleic anhydride, substitute amino groups Irgacure2959, UV BAEC cells seeded on top, 24h, most crustaceans cells were viable84 Glycidyl methacrylate, substitute amino Irgacure2959, UV Chondrocyte seeded on top, 7d, most groups cells were viable71 Gelatin Denatured Methacrylic anhydride, substitute amino Irgacure2959, UV NHDF with hydrogels, 24 or 72h, >80% collagen groups viable84

Encapsulated NIH/3T3, 6 or 48h, >80% 13 80 viable Gellan gum Sphingomonas Methacrylic anhydride, substitute hydroxyl Methyl benzoylformate, UV Encapsulated L929, 21d, viable and elodea groups maintain round shape73 Fucoidan Brown algae Methacrylic anhydride, substitute hydroxyl Eosin-Y, visible light L929 cells seeded on surface, 7d, groups spindle shape74 Laminarin Brown algae Glycidyl methacrylate, substitute hydroxyl Irgacure2959, UV Encapsulated human adipose stem cells, groups 7d, >90%85 Chitin Cell walls of Sodium monochloroacetate, methacrylic Irgacure2959, UV HeLa cells seeded at bottom of fungi, anhydride, substitute hydroxyl groups hydrogel, 48h, similar continuous exoskeletons proliferation with control76 Alginate Brown algae Glycidyl methacrylate, substitute hydroxyl Ammomium Encapsulated human endothelial cells, groups persulfate/tetraethylethylenediamine, 96h,~80%77 60℃ Hyaluronic ECM Methacrlate anhydride, substitute hydroxyl Irgacure2959, UV Encapsulated human endothelial cells, acid groups 7d65

Photoinitiators are a commonly used source of radicals due to their facile control of both step-growth kinetics and termination of reaction by modulating light irradiation.86

Lithium acylphophinate (LAP) and 2-hydroxy-4-(2-hydroxyethoxy) -2-methylpropiophenone

(Irgacure 2959) are widely used cleavable (type I) photoinitiators for thiol-ene reactions.87,88,90–95 LAP results in more rapid polymerization than Irgacure 2959.87 This type of photoinitiator splits to form radicals under long wavelength UV irradiation (λ = 365 nm,

10 mW cm-2).93 However, as mentioned previously, UV light and radicals generated by the cleavable initiator may cause cellular damage.91,96

Visible light sensitive, non-cleavable photoinitiators (type II) were developed to mitigate the cellular damage caused by UV light. Eosin-Y and rose Bengal are commonly used visible light photoinitiators which abstract protons from amine-bearing co-initiators

(e.g. triethanolamine, TEA) or dithiothreitol directly to generate radicals. Those radicals react with thiol groups to produce thiyl radicals.25,97 However, thiyl radicals might have insufficient reactivity to initiate polymerization, or they may not propagate quickly enough for crosslinking.25 N-vinylpyrrolidone (NVP) was introduced to accelerate gelation kinetics and to adjust the stiffness of hydrogels.25 Chien-Chi Lin and coworkers demonstrated that by adding 0.1% vol NVP, the gelation point decreased from 431 s in the absence of NVP to 231 s. Increasing the concentration of NVP to 1% vol further dropped the gelation time to 92 s.25

The same group compared the effects of the type I photoinitiator LAP with the type II eosin-

Y on hydrogel formation and cell response. Less Eosin-Y was required (0.1 mM) compared to LAP (4.0 mM) in order to fabricate hydrogels with similar shear moduli (15 kPa at 8% polymer content) using norbornene modified gelatin and thiol terminated 4 arm PEG.91

Control over crosslinking density of thiol-ene hydrogels can be done not only by changing polymer content,87,88,98 but also by adjusting the thiol-ene ratio.91,92,96,99,100

Increasing polymer content directly enhances hydrogel crosslinking while the thiol-ene ratio has a non-linear relationship with hydrogel crosslinking density. A parabolic relationship was found between bifunctional thiol concentration in precursor solution and resultant hydrogel shear moduli for PEGDA hydrogels.92 (Figure 2A and B) The initial increase in thiol content increased the thiyl radical concentration and promoted gelation. Further increases to the dithiol crosslinker concentration resulted in increased chain transfer and terminated homopolymerization.92 When vinyl groups were grafted on the polymer but were not able to homopolymerize, the thiol-ene ratio still exhibited a parabolic relationship with crosslinking density.100 (Figure 2C and D) This relationship was observed for norbornene, a special carbon-carbon double bond that is not able to homopolymerize due to steric hindrance, was grafted onto HA.100 Non-ideal crosslinking such as the formation of loops and dangling structures resulting in optimal crosslinking at a thiol-ene ratio at 0.6 – 0.8, which is lower than the anticipated maximum of equimolar.100,101 (Figure 3)

Figure 2. Effect of PEGDA molecular weight and bi-functional thiol content on equilibrium shear moduli. PEGDA macromers were used at: (A) 10 wt% and (B) 15 wt %. All hydrogels were formed with 0.1 M eosin-Y and 0.1% NVP and with 5 min of visible light exposure. Nonlinear curve fitting was conducted using parabolic relationships as a function of bifunctional thiol content. Reprinted with permission of Wiley.92 (C) Compressive modulus (Ec) as a function of XDTT (NorHA weight percent of 4%, irradiation time of 45 min) (D) Ec as a function of XDTT at an irradiation time of 30 x and NorHA weight percent of 4%. The precursor solutions for C and D contained 0.05 wt% I2959 and were irradiated at 10 mW/cm2 of UV light. The error bars in C and D are one standard deviation (n  3). Reprinted with permission from Gramlick, Kim, Burdick, Click chemistry hydrogels in tissue engineering applications, 9803, 2013, 100 with permission from Elsevier.

Figure 3. Illustration of defects in dithiol crosslinking and effect of excessive thiol during thiol-ene reaction. Adapted with permission from 101. Copyright 2006 American Chemical Society. Thiol-yne chemistry is similar to thiol-ene chemistry in which thiol groups react with carbon-carbon triple bonds,102,103 thus increasing functionality and crosslinking density.88

Daniele et al. created an interpenetrating network hydrogel by coupling a thiol-yne crosslinked PEG network and a methacrylated gelatin network.88 The resultant hydrogel exhibited an increased elastic modulus (327.7 kPa) compared to hydrogels crosslinked with thiol-ene chemistry (247.2 kPa).88 The extra functionality of –yne over –ene means that additional functional groups can be incorporated into the hydrogels.104 Hyper-branched poly

(ether amine) functionalized with alkyne was crosslinked with thiol-containing polyhedral oligomeric silsesquioxane under UV light with I907 as the photoinitiator.104 By maintaining excess alkyne compared to thiol, azide containing fluorescent molecules were conjugated to hydrogel through CuAAC chemistry.104 The electron density of carbon-carbon triple bonds

18 compared with double bonds enables thiol-yne reaction to be engaged without UV light and metal catalysts under mild conditions. Tetra alkyne and tetra thiol functionalized PEG was reported to form an organogel in the presence of trimethylamine at 80℃.105 When the alkyne to thiol ratio was 1:3 the shear modulus was 27000 kPa at 100 rad/s.105

3. Click chemistry hydrogels

Click chemistry provides a facile way to fabricate hydrogels used in biomedical applications and offers a spatiotemporal control over network evolution.87 Hydrogels crosslinked through click chemistry can be classified according to the origin of the polymer network, as either synthetic, natural polymer based hydrogels, or a combination of the two, termed hybrid hydrogels. As stated above, synthetic polymers are more amenable to sterilization and fabrication but lack cell binding motifs. Natural polymers are generally more cytocompatible than synthetic polymers, but generally have poorer mechanical properties.4

The following sections will discuss properties and applications of click chemistry hydrogels from a materials selection perspective.

3.1. Synthetic polymer based hydrogels

Synthetic polymers are generally less biocompatible when compared to natural derived polymers,4 as toxic chemicals and harsh synthetic chemistry are generally involved in fabricating synthetic hydrogels.106,107 This requires careful elimination of contaminants and unreacted regents present during synthesis.107 Synthetic polymers are easier to produce in large scales, with highly tunable and consistent properties. Moreover, synthetic polymers are more resilient to increased temperature and can be more easily sterilized.4 These features mean that synthetic polymers are more amenable to chemical modification.107 Herein, we describe methods to fabricate synthetic hydrogels and strategies to modulate resultant hydrogel mechanical properties as well as approaches to increase biocompatibility. (Table 3)

3.1.1. PEG

PEG and its multi-armed derivatives are commercially available hydrophilic polymers with low toxicity and highly controllable chemical and mechanical properties.93

Terminal hydroxyl groups on PEG can be easily functionalized with a variety of carbon- carbon double bond bearing moieties and subsequently crosslinked to form hydrogels.25,88,93,108,109 Norbornene, acrylate, methacrylate, and acrylamide are widely used to provide vinyl groups for thiol-ene crosslinking using multi-thiol crosslinkers.25 Gelation kinetics of tetra-armed PEG terminated with acrylate, acrylamide, and methacrylate under the same crosslinking conditions have shown that acrylate terminated tetra-armed PEG had the fastest gelation point at 92  11 s while the methacrylate terminated tetra-armed PEG had the slowest gelation at 275  11 s.25

PEG contains no biological motifs for cell or protein recognition and typically requires the introduction of peptides to become bioactive. Alternatively, hybrid hydrogels of

PEG and naturally derived polymers can introduce cell adhesion motifs.97,108 PEGDA was grafted to gelatin and L-cysteine, which contains a thiol moiety. Hydrogels were then fabricated using thiol-ene chemistry.93 NHDF fibroblast cells seeded on these hydrogels exhibited enhanced attachment with increasing gelatin content.93 In another study, Anseth and coworkers crosslinked tetra-norbornene terminated PEG with a chymotrypsin-degradable peptide through thiol-ene chemistry to form hydrogels with equimolar thiol-ene ratio.

Encapsulated hMSCs showed 95% cell viability 24 h post polymerization and cells encapsulated in the hydrogel were able to spread.87

In addition to increasing cytocompatibility, the mechanical properties of PEG based thiol-ene hydrogels can be engineered by forming hybrid hydrogels.54,84,110–113 PEGDA was

20 copolymerized with maleic functionalized chitosan and gelled through photocrosslinking using Irgacure 2959 as the initiator.84 The compression moduli of the resultant hydrogels were found to increase with increasing PEGDA mass ratio to chitosan.84 The highest compression modulus was 560.4  18.1 kPa at PEGDA to chitosan 5:1 and lowest was 61 

1.9 kPa when PEGDA:chitosan was 2:1.84 In addition to photocrosslinking, SPAAC has also been used in fabricating PEG-based hybrid hydrogels.54 Propiolic acid ester-functionalized

PEG and azide functionalized chitosan were mixed to form hydrogels at alkyne to azide ratio of 1:2.54 The shear modulus of the gel was 44.2  1.2 kPa, which was higher than the reported chitosan/glycerol phosphate system (5 kPa).54 Diels-Alder reactions can provide an additional method for synthesizing PEG-based hybrid hydrogels. Furan functionalized HA and maleimide terminated PEG were mixed at variable furan to maleimide ratios from 1:1 to

9:1.113 The compression moduli of the gels decreased with increasing furan to maleimide ratio.113 The highest compression modulus was 75 kPa for equimolar furan:maleimide gels and the lowest modulus was 15 kPa for furan:maleimide 9:1.113

Table 3 Summary of synthetic polymer based hydrogels crosslinked through click chemistry or chain growth polymerization

Syntheti c Crosslinking Modification method Crosslinker Biocompatibility polymer mechanism s Human epithelial cells(MDA-MB-231), on top, 14d, 98% bis-maleimide functionalized furan modified HA viability32 alanine and maleimide, furfurylamine grafted Diels-Alder Rabbit mesenchymal stem cells, on top, 24h, >80% viability114 maleimide functionalized chondrotin sulfate furylamine and tyramine Mouse embryonic carcinoma clonal cells (ATDC-5), maleimide functionalized substituted HA encapsulated, 7d, proliferate and viable115 Tetrakis (2- C sodium azide, azide Pig subcutaneous implantation, 2w, no obvious inflammation in propynyloxymethyl) u functionalized surrounding tissue42 methane A Tetrakis (2- A diazide functionalized propynyloxymethyl) HL-7702 cells, in gel treated media, 24h, 90% viability39

C 21 azide-alkyne methane PEG S propiolic acid, 3 arm azide azide functionalized Human bone marrow mesenchymal stem cells, encapsulated, P functinalized chitosan 24h, 95% viability54 A 2-((2-(cyclooct-2- Rat subcutaneous injection,3d, macrophage disappeared and scar A ynyloxy)ethoxy)methyl)oxirane, azide functionalized PEG emerged56 C cyclooctyne functinalized MDA-MB-231 cells, encapsulated, 5d, similar viability to those diacrylate PEG N-acryloyl glycinamide cultured in collagen matrix116 chain growth maleic functionalized Bovine aortic endothelial cells (BAEC), on top, 24h, mostly diacrylate PEG chitosan viable visually84 5-norbornene-2-carboxylic acid, dithiol functionalized PEG Chondrocytes, encapsulated, 24h, mostly viable visually117 tetranorbornene terminated PEG thiol-ene Norbornene anhydride, 4 arm cysteine peptide (thiol MIN-6 cells, encapsulated, 16h, 93% viability118 norbornene terminated PEG bearing peptide) norbornene functionalized PVA, thiol-ene functionalized PVAs L929 cells, PVA treated media, 30min, ~80%119 thiol functionalized PVA PVA azide-modified PVA and alkyne- azide-alkyne modified PVAs Not conducted120 modified PVA

3.1.2. PVA

Poly (vinyl alcohol) (PVA) is a water soluble synthetic polymer derived from alcoholysis of poly (vinyl acetate). PVA is non-toxic with LD50 (oral, rat) values between 10 and 215 g/kg.121 PVA can be modified through its secondary hydroxyl groups, increasing its chemical diversity.121 Crosslinking PVA can be done physically through repeated freeze-and- thaw cycles or chemically either through direct crosslinking of hydroxyl groups or incorporation of polymerizable moieties.121 Grafting PVA with allyl- and norbornene-based carbon-carbon double bond moieties and crosslinking with DTT is one way of chemically fabricating PVA hydrogels. Modified PVA hydrogels can be crosslinked using Irgacure 2959 as the photoinitiator. A kinetic study showed that using 60 mol % thiol resulted in the highest crosslinking density, 40 mol % thiol had highest photoreactivity, and 80 mol % thiol had the lowest gelation time.121 PVA grafted with allylic double bonds had higher reactivity at low to medium macromer content (<5 wt-%) while at higher macromer content (10-15 wt-%) norbornene-modified PVA was more reactive, demonstrating again that the ratio of thiol to vinyl groups is an important optimization parameter in thiol-ene reactions.121 Other than thiol-ene reactions, PVA-based hydrogel can also be crosslinked through CuAAC reaction.120

Azide-functionalized and alkyne-functionalized PVA were reported to crosslink with copper

(I) as a catalyst within 1 min in either DMSO or aqueous media. The shear moduli were adjusted by changing functional group concentration with highest 16.9 kPa at 29.2 mM and lowest 3.4 kPa at 14.4 mM120

PVA can be used to tune the mechanical properties of naturally derived hydrogels due to its high mechanical strength.122 PVA and starch modified with carbon-carbon double bonds were synthesized using maleic anhydride. Hydrogels were then formed using hydrogen sulfide as the crosslinker. The crosslinking density could be tuned by adjusting the molecular

23 weight of PVA. By increasing the intrinsic viscosity of PVA from 1.99 to 2.99 dL/g, the storage modulus increased from 12.7 to 50.9 MPa and the swelling ratio decreased from ~10 to ~5.122 In another study, PVA and chitosan were crosslinked by adding acrylic acid (AA).

The chain growth hydrogel was initiated by UV light (250-420 nm wavelength) without initiator. The hydrogel swelling behavior, which is indictor of crosslinking density, showed a parabolic trend with increasing PVA/chitosan to AA ratio.123 The swelling ratio initially decreased with increasing AA content, with further increasing AA content leading to higher swelling ratios, the range was from ~100 to ~200.123 The highest crosslinking density was observed when PVA/chitosan functional groups were in equality with AA functional group.123

3.2. Natural polymer based hydrogels

Natural polymers are derived from living sources and can range from bioactive to bioinert.4 Bioactive polymers can significantly modulate inflammatory responses while bioinert materials are ‘invisible’ to immune system and have low protein and cell adhesion.4

The following sections will discuss thiol-ene hydrogels fabricated from commonly used natural polymers with different origin, charge, and functionalities.

3.2.1. Chitosan

Chitosan is a linear polysaccharide that is obtained from partial deacetylation of the chitin shells of crustaceans.85 Chitosan is enzymatically degradable with a degree of deacetylation (DD) ranging from 15 to 85%. When its molecular weight is over 20 kDa, it is insoluble in physiological fluids.85 Chitosan is most common naturally occurring polysaccharide that is positively charged. The positive charge arises from amino groups and enables chemical modification.78 The pore size of chitosan hydrogels is influenced by the molecular weight of the polymer. Increasing the molecular weight of chitosan from 499 to

1360 kDa results in a more compact structure with fewer pores as visualized by SEM.124 DD is another factor that can change hydrogel morphology. Hydrogels networks became more compressed when DD increased from 75.4 to 85.5%.124

The mechanical properties of chitosan hydrogels can be tuned by copolymerizing with degradable synthetic polymers.78 Poly-lactic acid (PLA) grafted chitosan (310 kDa, 75%

DD) with different mass ratios was subsequently grafted with methacrylate groups.

Hydrogels were formed via chain growth crosslinking using a photoinitiator. The compression moduli of the hydrogels decreased with increasing PLA mass ratio (35 kPa without PLA and 15 kPa with chitosan:PLA 1:1).125 The reduced mechanical property was attributed to PLA acting as a plasticizer that reduced intermolecular interactions, such as hydrogen bonding.78 Cytotoxicity of the hydrogels was assessed using W-20-17 pre- osteoblast mouse bone marrow stromal cells and C2C12 mouse myoblast cells. Cells coated with these hydrogels were viable and proliferated after 3 days culture.78 Another way to tune the mechanical properties of chitosan hydrogels is by changing the degree of methacrylation.

By changing the ratio of methacrylate groups to unsubstituted amino groups on chitosan from

0.5 to 2, the shear moduli increased from < 125 Pa to over 175 Pa.126 Encapsulated chondrocytes in methacrylated chitosan hydrogels remained viable and proliferated after 7 days as visualized by fluorescence microscopy, which indicated that this hydrogel has potential as an injectable cartilage scaffold.126

The degradation rate of chitosan hydrogels is highly dependent on the crosslinking density and enzyme concentration. Chitosan hydrogels crosslinked with 5% initiator completely degraded in 1 mg/ml lysozyme within 8 days, while stiffer hydrogels formed using twice the concentration of initiator maintained 66% of its original mass after 18 d.127

Chitin is the precursor to chitosan and degrades via the same mechanism as chitosan in lysozyme. Chitin was modified to contain carbon-carbon double bonds and crosslinked to form hydrogels. These chitin hydrogels completely degraded in 50 mg/ml lysozyme within

10 h while 50% mass remained after 60 h in 1 mg/ml lysozyme.76

3.2.2. Gelatin

Gelatin is a natural biomacromolecule derived from denatured collagen. It has higher water solubility and lower immunogenicity compared to collagen.83 Gelatin contains an RGD peptide sequence that promotes cell attachment through integrin binding.83 In addition to the cell-adhesive sites on gelatin, protease-sensitive sites (such as the peptide sequence

GPQG↓IWGQ) are also present, which renders gelatin enzymatically degradable. Similar to chitosan, crosslinking gelatin through thiol-ene chemistry requires the introduction of carbon-carbon double bonds. Carbic anhydride and methacrylate anhydride were used to introduce norbornene and methacrylate group, respectively, on gelatin both through amidation.83

Mechanical properties of gelatin hydrogels can be tuned by the crosslinking mechanism and through copolymerization. Carbic anhydride can be used to modify gelatin, yielding norbornene-functionalized gelatin (GelNB). Hydrogels were fabricated using DTT as a crosslinker and their properties were compared with chain growth hydrogels formed by methacrylated gelatin (GelMA) with a similar degree of carbon-carbon double bond substitution. Chain growth GelMA hydrogels exhibited a higher shear modulus (0.9 kPa) than thiol-ene GelNB hydrogels (0.4 kPa). Cell viability of encapsulated hMSCs was higher in the thiol-ene hydrogels (97% viable in 4% GelNB hydrogels) compared with chain growth

GelMA (85% viable in 4% GelMA hydrogels) after 1 day incubation.83 Other than crosslinking mechanism, copolymerizing gelatin with synthetic polymers is another approach

26 to tune the mechanical properties of gelatin hydrogels.55 Alkyne-functionalized PEG was used to copolymerize azide functionalized gelatin through an SPAAC reaction.55 The shear moduli of these hydrogels were compared with hydrogels fabricated solely from gelatin.

Hydrogels fabricated with PEG exhibited a higher shear modulus (1.2  0.1 kPa) than pure gelatin hydrogel (0.8  0.1 kPa).55 Copolymerizing with naturally derived polymers is method to tune the crosslinking density of gelatin-based hydrogel.128 Methacrylated chitosan was copolymerized with methacrylated gelatin using Irgacure 2959 as photoinitiator through chain growth mechanism.128 By controlling the ratio between methacrylated chitosan and gelatin from 2:1 to 1:2, the pore size characterized by SEM decreased from 30 휇m to 2

휇m.128 In addition to the crosslinking mechanism, initiator efficiency is another direct factor that controls crosslinking density. The effectiveness of visible light sensitive eosin-Y and UV light initiator LAP were compared for hydrogels composed of GelNB and 4 arm thiol- functionalized PEG. Eosin-Y (0.1 mM) was more effective than LAP (4 mM) in crosslinking

GelNB hydrogels with similar shear moduli. Hydrogels fabricated using both initiators degraded within 240 min in the presence of 10 U/ml of collagenase.91

Swelling behavior is an important property for hydrogels used in biomedical applications. Tissue can be compressed and body fluid flow can be impinged if implanted hydrogels swell. This decreases both safety and patient comfort.102 Controlling swelling behavior of gelatin-based hydrogels can be done through the incorporation of a thermosensitive crosslinker. Non-swelling hydrogels were fabricated using thiolated gelatin and alkyne functionalized poly (ethylene glycol)-poly (propylene glycol)- poly (ethylene glycol)as a thermosensitive crosslinker.102 The hydrophobic block exhibited lower critical solution temperature behavior at 15-17 C and restricted hydrogel water uptake at

27 physiological temperature. The swelling ratios of the hydrogels ranged from 1.5 to 2.7 and decreased with increasing temperature.129 Incorporating a second component into gelatin hydrogels is another method to adjust swelling behavior.65 Methacrylated HA copolymerized with methacrylated gelatin through photo-initiated chain growth mechanism decreases the swelling ratio (~30 in the absence of HA, ~20 in the presence of 1% HA) when gelatin concentration was fixed at 3%.65

3.2.3. Hyaluronic acid hydrogels

HA is a non-sulfated glycosaminoglycan found in the ECM. It can be degraded by hyaluronidase secreted from various cells including human adipose derived stem cells,130 tumor cells,131 and fibroblasts.132 Cells can interact with HA through surface receptors such as CD44.130 The presence of hydroxyl and amide groups allows chemical modification of

HA.95 UV irradiation time was found to influence HA hydrogel mechanical properties. HA was functionalized with a norbornene group and subsequently crosslinked using DTT. At fixed thiol-ene ratios, increasing the UV irradiation time increases the hydrogel compression modulus. The highest compression modulus reached ~70 kPa at 20 min irradiation time when the thiol to carbon-carbon double bond ratio was ~0.8 and polymer content was 4% compared to ~20 kPa within 10 s.100 Other than UV irradiation time, the pH of the hydrogel precursor solution is also responsible for tuning stiffness. Methacrylated HA and thiolated heparin (Hep) were mixed at a 1:1 molar ratio of thiol: acrylate at 2% polymer content.

Gelation time decreased with increasing pH. The storage modulus of increased with increasing polymer content.130 The pH sensitivity of thiol-ene gelation was attributed to enhanced electron transfer from the initiator at higher pHs.24 Adipose derived mesenchymal stem cells showed similar attachment (~100%) to the HA-Hep hydrogels compared with tissue culture plastic 3 h after seeding.130

4. Hydrogels fabricated using click chemistry as tissue scaffolds

Click chemistry provides an easy way to fabricate hydrogels by simple mixing of reagents or exposure to UV light and can control where and when hydrogels are formed.

These features allow hydrogels to be used as scaffolds for tissue engineering. Hydrogel tissue scaffolds should have mechanical and structural properties similar to the ECM of the implantation site.2 By changing the crosslinking density, crosslinker length, degree of branching, and molecular weight of the precursor, mechanical properties can be carefully engineered.2 A wide array of implantation strategies from in situ formation to direct implantation of fully realized scaffolds have been implemented using these materials as scaffolds. Alkyne functionalized chitosan and azide functionalized HA was injected bilaterally into the dorsal subcutaneous region of Balb/c mice and formed hydrogel through metal free AAC reaction.19 Hydrogels maintained a circular shape after 7 days and the gel- tissue interface showed a lack of neutrophils and macrophages infiltration,19 indicating biocompatibility. The Diels-Alder reaction is another click chemistry that commonly used to crosslink hydrogels in situ. Furfurylamine functionalized chondroitin sulfate and maleimide terminated PEG were mixed to yield a mechanical stiff hydrogel (E~25 MPa).114 These two components can be injected into bone defect sites to form stiff hydrogels to enhance bone repair.114 Histological evaluation of defects site that were treated with hydrogels showed formation of new bone tissue after 14 days, which is faster than in mice without the hydrogel scaffold.114

Implantation under the skin restricts in situ fabrication of photocrosslinked hydrogels, such as thiol-ene and chain growth hydrogels that use photoinitiators. Substituting photoinitiators with thermal initiators is one of the method to address this problem.77

Glycidyl methacrylate grafted alginate mixed with a thermal initiation system (APS and

TEMED) were injected into Sprague-Dawley rats.77 Histological analysis showed few neutrophils and macrophages after two weeks, which indicated a mild inflammatory response and good biocompatibility. Direct implantation of preformed hydrogel scaffolds is another method to address the challenge of using photopolymerized gels. Methacrylated gellan gum hydrogels crosslinked through chain growth polymerization and ionic interaction were separately implanted into adult female Lewis rats for 18 days. Hematoxylin & eosin (H&E) staining showed moderate infiltration of macrophages and fibroblasts at day 10, which qualitatively decreased at day 18. Connective thin fibrous tissue was observed surrounding both hydrogels, demonstrating the biocompatibility of these gels.73 Photocrosslinked PEG hydrogels copolymerized with hydrogen bond forming N-acryloyl glycinamide (NAGA) was implanted into back of C57BL/6 mice and the fibrotic response was measured 4 weeks post- implantation. These hydrogels did induce a fibrotic response, however the density of deposited collagen surrounding the PEG-based hydrogels was lower compared to control hydrogel made from poly (2-hydroxyethyl methacrylate).116 The fibrotic response was related to NAGA content. Low (4%) and high (16%) NAGA content caused more fibrosis than PEG hydrogels with medium NAGA content (8%).116 The high fibrosis for the low NAGA content

PEG hydrogel was likely caused by degradation of the hydrogel and ensuing cell infiltration.116 The elevated fibrosis for the high NAGA content PEG hydrogel remains an open question. Both the low biocompatibility of stiff NAGA and the low biocompatibility of stiff PEG hydrogels were possible reasons.116

The host response of hydrogel tissue scaffold is affected by many factors. Hydrogel scaffolds incorporated with growth factors are able to accelerate tissue healing. Azide functionalized chitosan was photo-crosslinked directly in growth factor FGF-2 solution under

UV exposure without initiator.133 Resultant hydrogels were applied to full thickness circular skin wounds on the back of healing-impaired diabetic (db/db) mice.133 The healing process was faster on the wound site with hydrogels incorporated with FGF-2 (80% wound closure within 12 days) than for control hydrogels without FGF-2 (80% wound closure after 20 days).133 Crosslinking mechanism also influences the host response. Azide and cyclooctyne functionalized HA were injected subcutaneously into ICR mice and crosslinked through the

SPAAC mechanism.45 H&E staining images showed slightly more inflammation from

SPAAC crosslinked HA hydrogels than HA hydrogels crosslinked by a Schiff base reaction.45

5. Conclusions and future perspectives

Hydrogels are biocompatible materials due to their structural similarities with the

ECM and the inherent hydrophilicity of the polymer network.2 Their gelation time and network crosslinking density can be tuned through the selection of materials, optimization of the ratios of the reagent, and varying crosslinking mechanisms. The tunability of hydrogels properties means that they can be engineered for a wide variety of biomedical applications and that they have the potential to be used as in situ forming scaffolds.2

In this review, we discussed four diverse crosslinking mechanisms for fabrication of hydrogels. Advantages and drawbacks of each mechanism were discussed in detail. We also gave a general introduction to click chemistry hydrogel fabrication from the perspective of materials section. Representative materials from both the synthetic and naturally derived polymer families were discussed to demonstrate the feasibility of using hydrogels, along with their features and applications. Finally, we provide examples of hydrogel scaffolds synthesized through click chemistry and their specific features.

The future success of hydrogels as tissue scaffolds is highly dependent on rational design of the physical, chemical, and biological properties of these gels in order to better mimic the structural and biological properties of natural ECM into which it will be implanted.134 These properties can be tuned by careful selection of crosslinking mechanisms, modulating ratios of reactive functional groups, as well as polymer content in the hydrogel. It is also important for future click chemistry hydrogels to reduce the cytotoxicity of reagents and optimize fabrication procedures. Improving cytocompatibility during fabrication through methods such as using visible light initiators to substitute for UV sensitive initiators is critically important for cell-laden hydrogels. For hydrogels designed for use as tissue scaffolds, fibrosis is a determining factor for success. Surface modifications and mechanical properties should be carefully engineered to reduce fibrous encapsulation. Growth factors can be chemically incorporated into hydrogel scaffold structure that released upon degradation of the hydrogel. For hydrogels designed for cells to attach, the incorporation of cell-adhesive peptide sequences into hydrogel network should be investigated, which not only enhance cell binding but also can modulate enzymatic degradation. In this way, new tissue regenerated on the scaffold while the scaffold is degrading at a compatible rate. For hydrogels designed to study cell spreading, peptide patterning both on the surface and in the bulk of the hydrogel should be studied. In order to address these challenges, interdisciplinary cooperation between various fields including materials science, polymer chemistry, and cell biology will be essential to develop ideal tissue scaffolds.135

CHAPTER 2. GELLAN GUM BASED THIOL-ENE HYDROGELS WITH TUNABLE PROPERTIES FOR USE AS TISSUE ENGINEERING SCAFFOLDS

1. Introduction

The overriding design principles for tissue engineered scaffolds are that the scaffold should best simulate the properties of the native tissue.4 Namely, the scaffolds should have similar mechanical properties to the tissue it will replace, the scaffold is biocompatible, and the degradation rate of the scaffold matches those of the surrounding tissue.4 The native extracellular matrix (ECM) is a scaffold for cells to attach and proliferate. The ECM has a structure with high water content and a 3D network which can be mimicked by hydrogels.

Therefore, research efforts on designing hydrogels have been focused on improving biocompatibility and mechanical properties to match native tissues.

Gellan gum is a natural polysaccharide derived from bacteria that has potential applications in this area owing to its biocompatibility and flexibility for modification. The repeating unit of gellan gum is a tetrasaccharide comprised of L-rhamnose, D-glucuronic acid, and two D-glucose residues. As with other polysaccharides, lysozyme secreted by monocytes and neutrophils136 can degrade the backbone.137 The kinetics of degradation can be controlled by altering the degree of crosslinking as well as the type of crosslinking bond, which can be achieved through altering physical or chemical linkages.7 As with changing degradation profiles, the mechanical properties can also be altered through modifying the type and degree of crosslinks.7 The mechanical property requirements for scaffolds depend on the type of tissue, with softer (<2 kPa) substrates yielding the desired outcomes for liver and harder (>40 kPa) implants required for bone tissue engineering.4

Cations can be used to crosslink gellan gum gels by screening charges on the molecules and promoting physical crosslinks. More specifically, multi-valent cations bridges

33 carboxyl groups and reduce electrical repulsion.138 While monovalent cations can also screen the charges on the gellan molecules, the interaction is much weaker than the divalent cations.138 Therefore, physically crosslinked gellan gum hydrogels will lose their stability in physiological conditions as the divalent cations are replaced by monovalent cations that are present at much higher concentrations in vivo.7

One way to improve the stability of gellan gum hydrogels is to covalently crosslink the gel.81,139,140 Carbodiimide amidation and esterification are commonly employed to conjugate the abundant carboxylic acid or hydroxyl groups present on the polysaccharide to a moiety capable of chemical crosslinking. Glycidyl methacrylate141 and methacrylic anhydride73,81,140 have been successfully used to methacrylate gellan gum.127,142 Enzymatic degradation of the hydrogel is retained as lysozyme attacks the backbone and can be tuned based on the degree of modification with acrylates and thiol concentrations. Likewise, the mechanical properties of these gels can be similarly tuned.89

Thiol-ene photoclick chemistry provides another free radical mediated method to synthesize hydrogels using crosslinkers containing multiple thiol groups.[18-20] Specifically, a free-radical abstracts a proton from a thiol, generating a thiyl radical. Subsequent propagation steps occur in a step growth mechanism between the thiyl radical and the vinyl group.143 In this work, we examine the relationship between mechanical properties and crosslinking mechanism with cell proliferation using covalently crosslinked gellan gum.

Through introducing carbon-carbon double bonds to the gellan gum, the polymer can be photocrosslinked through chain growth, step growth, or a combination of the two, termed mixed model, to form gels with varying compressive moduli. Thiol-ene step growth crosslinking is not oxygen inhibited, therefore chain growth crosslinking can be controlled

34 through purging oxygen from the precursor solution. By adjusting the crosslinking conditions, the mechanical properties of the gel were systematically varied. The degradation profiles of these polymers in the presence of lysozyme was also investigated. Correlations between mechanical properties and cell proliferation were explored.

2. Methods and materials

2.1.Materials

Gellan gum was purchased from Spectrum Chemical Mfg. Corp. (New Brunswick,

NJ); methacrylate anhydride (MA) was obtained from BeanTown Chemical (Hudson, NH);

2-hydroxy-4′-(2-hydroxyethoxy)-2-methylpropiophenone (Irgacure 2959) was obtained from

Sigma-Aldrich (St. Louis, MO); dithiothreitol (DTT) was supplied by VWR Chemical

(Batavia, IL). Fresh deionized (DI) water (Milli-Q Nanopure, Thermo Scientific, Waltham,

MA) was used throughout this study.

2.2.Gellan gum modification

Methacrylated gellan gum was synthesized via substitution of the hydroxyl groups in the gellan gum repeating units with MA. Briefly, 1 g gellan gum was dissolved in 100 ml DI water in a round bottom flask and heated to 90°C for 30 minutes under constant stirring. The mixture was slowly cooled to 60°C and either 3.5 ml or 8.5 ml MA was added for low modified gellan gum (LMGG) or high modified gellan gum (HMGG), respectively. The pH was maintained between 8 and 9.5. After 4 h, the product was dialyzed (12,000-14,000 molecular weight cutoff) against DI water for at least 3 days. The DI water was refreshed daily. The final product was lyophilized (Labconco, Kansas City, MO, 4.5 L).

2.3.Characterization of modification of gellan gum

To quantify the percent modification of synthesized LMGG and HMGG, NMR was used to characterize the polymers. The modified polymers (10 mg) were dissolved in 1 ml

1 D2O. The H-NMR spectra were recorded on a Bruker Avance III spectrometer with a sweep width of 6602.1 Hz, a 90° pulse, and an acquisition time of 2.48s. All spectra were obtained at room temperature. A total of 16 repetitive scans with 64 points were acquired and the data were processed in MNova with 128k points, zero filling, and exponential line broadening of

1.0 Hz. The methyl group on the rhamnose structural unit on gellan gum was used as a reference (δ =1.45ppm). The degree of substitution (DS) was calculated using equation 1. I I DB CHl3rham

hH hH DS = DB CH3rham h OHmonomer (1) I I Where DB is the integration of double bond proton peak, CHl3rham is the

hH hH integration of the reference peak, DB and CH3rham are the number of protons in the double

h bond and in the methyl group on the rhamnose, and OHmonomer is the number of reactive hydroxyl sites on gellan gum.

2.4.Synthesis of hydrogels

In 1 ml DI water, 10 mg modified gellan gum was dissolved at 55°C. To this mixture,

1mg Irgacure 2959 was added. For mixed model LMGG hydrogels, either 3.4 or 6.8 휇l 10 wt-% DTT was added for thiol-ene ratio 0.5 or 1, then 2 휇l of 0.5 M CaCl2 were added to solution before degassing for 20 minutes. For step growth and mixed model HMGG hydrogels, 10 or 20 l of 10 wt-% DTT was added such that the thiol to alkene mole ratio was 0.5 or 1. To these mixtures, either 0 or 6 휇l of 0.5 M CaCl2 was added to yield a final concentration of 1 mM Ca2+. For chain growth hydrogels, DTT was not added. Chemically crosslinked hydrogels were obtained by exposure of the precursor solution to UV light (15

W, 365 nm, UVP, Upland, CA) for 10 minutes. Both chain growth and mixed model solutions were degassed prior to crosslinking.

2.5.Compressive modulus

The compressive modulus for each experimental condition was obtained through manual compression measurements. Obtained hydrogel pegs (n = 3) approximately 15  15 

6 mm3 were placed between two microscope slides and measured as weight was added. A stress-strain curve was created using distances between the microscope slides measured through Image J (NIH, Bethesda, MD). The linear regions of the curve under 5-15% strain were used to report the compressive moduli.

2.6. Swelling ratio

The swelling ratio was measured for the hydrogels in DI water. Each experimental condition was tested at a concentration of 1% w/v modified gellan gum. The obtained hydrogels weighed ( w0 ). The hydrogels were subsequently immersed in DI water until equilibrium was reached and weighed ( ws ). The wet swelling ratio ( S ) was calculated using equation 2.

w - w S = s 0 ´100% w0 (2)

2.7.In vitro degradation profile

The effect of the different crosslinking mechanisms on the degradation profiles of

LMGG and HMGG hydrogels was examined. Hydrogels (n = 3) were immersed in 0.5 mg/ml lysozyme and in phosphate buffered saline (PBS, diluted from 10 solution to 0.1 M, pH 7.4, Fisher Scientific, Pittsburgh, PA) and were weighed daily.

2.8.Cell Culture

NIH/3T3 fibroblasts (ATCC, Manassas, VA) were cultured in complete medium

(CM, Dulbecco’s modified Eagle’s medium supplemented with 10% bovine calf serum, 100

U/L penicillin, and 100 µg/l streptomycin) at 37°C in 5% CO2. Cells are passaged every three to five days through 0.025% trypsin-EDTA (ethylenediaminetetraacetic acid; Mediatech,

Tewksbury, MA) detachment and subcultured at 6.7× 103 cells/cm2.

2.9.Cell Proliferation

Cell proliferation was tested on all gelation formulations for both LMGG and

HMGG. A solution of 10 mg/ml LMGG or HMGG in CM was prepared at 37°C. Solutions for the mixed and chain growth gelation procedures were degassed for 15 minutes. To each well, 100 µL of each modified gellan gum solution was added to 48-well plates (KSE

Scientific, Durham NC). Hydrogels were exposed to UV light for 10 minutes. Cells (1.25 ×

105 cells/cm2) were seeded on of each modified gellan gum hydrogel and 200 µl CM without phenol red was added the plate was incubated for 48h. Live and dead controls were generated by culturing cells without gel. After 48h, the dead control supernatant was aspirated and replaced with 95% ethanol for 10 minutes. After this incubation period, the supernatant in each well was aspirated. To each well, 150 µl of working solution (2 µM calcein AM

(AnaSpec, Fremont, CA) and 7.5 µM 7-aminoactinomycin D (Tonbo Biosciences, San

Diego, CA) in PBS) was added and the plates were incubated to 30 – 40 minutes at 37°C in 5% CO2. Live cells were quantified using an excitation/emission of 485/590 nm. Dead cells were measured using an excitation/emission of 528/645 nm using a plate reader

(BioTek Synergy HT Multidetection Microplate Rreader BioTek, Winooski, VT).

Cells stained using the live/dead assay were imaged using an EVOS® FLoid®

Imaging Station (Life Technologies, Grand Island, NY) using the red channel

(excitation/emission 586/646 nm) and the green channel (482/532 nm).

2.10. Statistics and data analysis

Statistical analysis was performed using JMP statistical software. Statistical significance of the mean comparisons was determined by a two-way ANOVA. Pair-wise comparisons were analyzed with Tukey’s honest significance difference test. Differences were considered statistically significant for p < 0.05.

3. Results and discussion

3.1.Gellan gum modification and characterization

Methacrylation of gellan gum is schematically illustrated in Figure. 4A. The NMR spectrum of methacrylated gellan gum is shown in Fig. 4B. The 1H-NMR spectra of unmodified gellan gum (Figure. 4B top) and modified gellan gum (Figure. 4B middle and bottom) were obtained at 25℃. All three spectra showed the characteristic methyl peak from rhamnose (훿 = 1.3 ppm). Methacrylation was verified and quantified by the characteristic carbon-carbon double bond peaks (훿 = 5.7 ppm, 훿 = 6.2 ppm) and a peak corresponding to the methyl group on the newly methacrylated moiety (훿 = 1.95 ppm). The degree of modification was calculated by comparing the integrated proton peaks from the methyl group on the methacrylate and the methyl group on rhamnose. The average number of methacrylate group per LMGG repeating unit is 0.5 ± 0.06 (Figure. 4B middle) while this increased to 2.0

± 0.11 for HMGG (Figure. 4B bottom).

Methacrylation of gellan gum was further confirmed by FTIR (Figure. 4C). The carbon-carbon double bond peak is expected to appear at 1640 cm-1, which overlaps with other vibrational modes in gellan gum. However, the characteristic peak at 1720 cm-1

39 corresponding to the carbonyl present on the methacrylate group appeared after chemical modification and grew in intensity with increasing degree of methacrylation. Resultant gels are labelled by numbers according to their unique formula shown under each graph. Gel 1-4 are LMGG gels while 5-14 are HMGG gels.

Figure 4. Synthesis and characterization of methacrylated gellan gum. (A) Schematic illustration of the methacrylation reaction. (B) 1H NMR of gellan gum (top) and low

(middle) and high (bottom) modified gellan gum. (C) FTIR spectra of gellan gum and low and high modified gellan gum.

3.2.Compression modulus

The compressive moduli of the 14 different formulations of hydrogels was obtained by analyzing stress-strain curves in the region of 5% - 15% strain. The modulus of the hydrogels ranged from 6.4 to 17.2 kPa, with the LMGG gels ranging from 6.4 to 11.3 kPa

40 and HMGG gels ranging from 7.4 to 17.2 kPa (Figure. 5). The increased mechanical strength in HMGG is expected due to a higher degree of modification and crosslinking.

Figure 5. Mechanical properties of modified gels. Young’s moduli of low and high gellan gum hydrogels crosslinked through the different mechanisms. Data represents the mean ± SD (standard deviation). n = 5. Statistical analysis through two-way ANOVA and Tukey’s HSD post-hoc test. Bars with the same letter (A-G) are not statistically different (p < 0.05).

The mechanical properties of the gellan gum hydrogels synthesized here were altered by three factors: the presence of calcium, the chemical crosslinking mechanism, and the thiol-ene ratio. Calcium was found to either increase or have no significant influence on the compressive modulus of the gels. LMGG gels were not able to form in the absence of

41 calcium. For HMGG gels, the addition of calcium had little effect on the hydrogels crosslinked through mixed model. Hydrogels crosslinked through the step growth mechanism of equimolar thiol-ene together with calcium exhibited an increased compression modulus

(9.9  0.6 kPa) compared to their counterparts in the absence of calcium (7.3  0.8 kPa). A similar trend was observed when half-molar equivalents of thiol were added with the modulus increasing from 9.8  0.7 to 12.8  0.9 kPa. The compressive modulus of the gels formed using chain growth crosslinking was also improved through the addition of calcium

(14.6  0.8 kPa to 16.5  1.0 kPa).

Divalent cations such as calcium form bridges between carboxyl groups and reduce electrical repulsion, resulting in physical crosslinks.144 These junctions allow gellan gum chains to bundle to form helical structures145 and may increase mechanical properties.

However, the interaction between calcium cations with the gellan gum hydroxyl and carboxyl groups reduces the hydrogen bonding with water,146 therefore decreased the strength between the network and water matrix. As a result, the compressive modulus was not significantly influenced for some hydrogels.

The crosslinking mechanisms significantly affected the mechanical properties for all gels studied here. For LMGG gels, compression modulus was increased from 6.4  0.7 to 8.1

 0.7 kPa when the crosslinking mechanism was changed from step growth to mixed model using equimolar thiol-ene gels. The same result was observed for half molar thiol-ene LMGG gels, where the modulus increased from 9.6  0.7 to 11.3  0.8 kPa. Similarly, for HMGG gels the compression modulus increased from 7.4  0.8 to 12.4  0.8 kPa in the absence of calcium and from 9.9  0.6 to 13.0  0.7 kPa in its presence for equimolar thiol-ene gels.

This trend was also observed for the half molar thiol-ene hydrogels, in which the

42 compression modulus changes from 9.8  0.7 to 16.4  0.9 kPa in the absence of calcium and from 12.8  0.9 to 17.2  0.1 kPa in the presence of calcium. Surprisingly, the chain growth crosslinked gels had the highest compression modulus of 14.6  0.8 and 16.5  1.0 kPa, in the absence and presence of calcium, respectively.

Two reasons attribute to the high compression moduli of chain growth hydrogels compared with the mixed model and step growth hydrogels in this study. First, homopolymerization of acrylate yields heterogeneous chains and causes a high degree of chain entanglements, which contributed to a higher modulus.83 Second, there is a parabolic relationship between the thiol-ene ratio and resultant hydrogel shear modulus, with the shear modulus initially increasing, then decreasing with increasing thiol content.92 The two ratios used may not be the optimal ratio to yield the highest compressive modulus. The higher moduli observed in mixed model hydrogels compared with step growth counterparts were caused by the oxygen level of the precursor solutions. The presence of oxygen in precursor solution inhibited radical polymerization and reduced crosslinking density.90,117 Dissolved oxygen in the mixed model hydrogels precursor solution was reduced via degassing that promoted crosslinking. The higher modulus of chain growth hydrogel over step growth hydrogel has been previously observed by Lin and coworkers.83 In their study chain growth methacrylated gelatin exhibited almost double shear modulus (0.9 kPa) when compared to step growth norbornene functionalized gelatin (0.4kPa).83

The thiol-ene ratio also plays an important role in modulating the compressive modulus of the hydrogels. Equimolar thiol-ene hydrogels were initially expected to yield the highest compression modulus since each thiol group was expected to react with one vinyl group and consume all available vinyls.100 However, carbon-carbon bonds can homo-

43 polymerize resulting in excessive thiols, which react to form dangling dithiol bonds.83,100,101

In addition, increased dithiol results in increased chain transfer, which terminates chain propagation and reduces the degree of polymerization.89,92,147 For these reasons, half-molar thiol-ene hydrogels had higher compressive moduli compared with equimolar gels. For

LMGG hydrogels the half-molar thiol-ene gels yielded higher compressive moduli than the equimolar gels for both step-growth and mixed model crosslinking (8.1  0.7 kPa and 11.3 

0.8 kPa vs. 6.4  0.7 kPa and 9.6  0.7 kPa, respectively). This trend continued for HMGG gels, with the half-molar thiol-ene gels having higher moduli than their equimolar counterparts. This relationship between the thiol content and crosslinking density for step growth and mixed model gels has been observed previously.92,100 Homopolymerization of norbornene is difficult due to steric hindrance and was still found to have the highest crosslinking density at a lower than equimolar thiol-ene ratio.100 Defects in crosslinking such as formation of loops and dangling structures are believed to be responsible for this phenomenon.101

3.3.Swelling ratio

The wet swelling ratio was measured in DI water as an indirect method for assessing crosslinking density, as a high swelling ratio generally corresponds to a low crosslinking density and vice versa. In this work, the wet swelling ratio of the 14 types of gels (Figure 6) were negatively correlated with compressive moduli (R = -0.55). The wet swelling ratio of the gels ranged from 2.46  0.11 to 4.28  0.14 for LMGG gels and from 0.98  0.01 to 1.95

 0.34 for HMGG gels. A higher compressive modulus generally resulted in less water uptake. In addition to crosslinking density, the formation of ionic bonds and the hydrophilicity of the network also influences swelling.

Figure 6. Swelling behavior of gellan gum hydrogels. Hydrogels crosslinked through the different mechanisms were swelled in PBS. Data represents the mean ± SD. n = 3. Statistical analysis through two-way ANOVA and Tukey’s HSD post-hoc test. Bars with the same letter (A-J) are not statistically different (p<0.05).

Calcium ions not only form ionic crosslinking and lowering the swelling ratio, but also reduce water uptake. Since all LMGG hydrogels were strengthened with calcium, this discussion will focus on HMGG hydrogels. The influence of calcium on the swelling ratio was significant for thiol crosslinked gels. For equimolar thiol-ene ratio hydrogels, the step growth gels decreased from 1.94  0.34 to 1.29  0.058 when calcium was added. Similar decreases in the swelling ratio were observed for equimolar mixed model hydrogels when

45 calcium was added (1.73  0.13 to 1.26  0.035). This trend continued for half-molar thiol- ene gels formed through step growth crosslinking (1.57  0.033 in the absence and 1.07 

0.030 in the presence of calcium) and mixed model crosslinking (1.46  0.031 in the absence and 1.09  0.044 in the presence of calcium). Chain growth gels, on the other hand, did not have significant differences with or without calcium (0.98  0.009 and 1.1  0.030, respectively). This could be caused by the rigid network of chain growth crosslinks, which may have redcued the effect of the calcium physical crosslinks.87

Ionic bonds can influence the swelling ratio through multiple mechanisms. First, the formation of ionic bonds increases the crosslinking density and reduces the swelling capacity.7 Second, interactions between the carboxylic acids on gellan gum and calcium cations decrease the available hydrogen bonding with water,146 and decrease the amount of water associated with the hydrogel network. Third, compared with covalently crosslinked polymeric networks, ionic bonds are less flexible and extendable, which reduces their ability to swell.70

All LMGG gels showed significantly higher swelling ratios than HMGG gels, which were all strengthened with ionic crosslinks. For LMGG gels, the compressive modulus was negatively correlated with the swelling ratio (R = -0.97). Hydrophilicity affects the swelling response of hydrogels since hydrophilic networks taking up more water.148 LMGG has one methacrylate groups on every two tetrasaccharide repeating units, while HMGG has two methacrylates on each repeat unit. This suggests that LMGG will have a stronger affinity for water than HMGG.

The swelling ratio of gellan gum hydrogels crosslinked via chain growth polymerization has been studied previously. Dual crosslinked gellan gum hydrogel (chain

46 growth and ionic crosslinking) exhibited significant lower swelling ratio (~100% of original weight), which is similar to results presented here.7 The same group also investigated gellan gum swelling behavior in PBS. Dual crosslinked gellan gum hydrogels shrank when immersed in PBS solution due to formation more helix structure and crosslinks.7 Reis and coworkers found 2500% water uptake of photocrosslinked gellan gum hydrogel in PBS after

30 days.141 The same group further investigated and compared water uptake of photocrosslinked and ionically crosslinked gellan gum hydrogels, where both showed

~2500% water uptake after 90 days.73 However, the swelling ratio of gellan gum hydrogels in both DI water and PBS were highly dependent on the degree of methacrylation and the final concentration of cations. Due to this reason, results here cannot be directly compared with the literature.

3.4.In vitro degradation

Lysozyme is an antibacterial enzyme released by macrophages during the wound healing process,149 which breaks down the cell walls of bacteria and releases N-acetyl-D- glucosamine.150 Previous studies have shown that polysaccharides can undergo enzymatic hydrolysis in the presence of lysozyme.127,151–154 It is believed that lysozyme breaks ether bonds connecting the structural backbone of the polysaccharide.155 The in vitro degradation profile of hydrogels in this study was obtained by exposing the hydrogels to PBS or 0.5 mg/ml lysozyme (Figure. 7). Hydrogels immersed in pH 7.4 PBS were stable, with ≥ 80% mass remaining after 42 days (Figure 7A). This weight loss likely results from dehydration of the gel as the ionic concentration of the PBS solution is higher than that in the hydrogels.127 In lysozyme, all gels degraded to 20% of their initial weight after 16 days

(Figure 7B), with HMGG gels degrading slightly faster than the LMGG gels. The differences in degradation rates for LMGG and HMGG hydrogels may result from

47 differences in network hydrophilicity. The HMGG backbone contains more hydrophobic methacrylate groups than LMGG. These methacrylate groups can form hydrophobic kinetic chains during crosslinking,75,117 which may retard diffusion of lysozyme solution into the hydrogel.

Figure 7Gellan gum hydrogel degradation. Hydrogels formed through the different crosslinking methods were immersed in (A) PBS and (B) 0.5 mg/ml lysozyme at 37℃. Data represents the mean ± SD. n = 3.

The stability of gellan gum hydrogel in PBS was confirmed by several studies. Reis and coworkers immersed gellan gum hydrogel for 30d in PBS and hydrogels showed less than 20% weight loss.141 In a later study, the same group did PBS stability test for both ionic crosslinked and photocrosslinked gellan gum gels up to 90d, nearly 20% weight loss was

48 observed.73 The stability of another polysaccharide, chitosan, in PBS was also found to retain

80% of its original weight after 18 days, 127 which is in line with our findings.

The degradation rate of polysaccharide hydrogels in lysozyme is highly dependent on both the concentration of lysozyme and the type of polysaccharide. Photocrosslinked chitosan hydrogels were found to completely disintegrated in 1 mg/ml lysozyme after 8 days.127 Methacrylated chitin hydrogels retained 50% of their original weight after 60 h in

1mg/ml lysozyme while completely degrading in 50 mg/ml lysozyme within 10 h.76 The density of crosslinks in methacrylated chitin gels also alters the degradation rate, with a higher crosslinking density leading to a slower degradation rate.76 These results combined with observations in our study provide clues for controlling the degradation rates for future application of using hydrogels as tissue engineering scaffolds.

3.5.Cell Proliferation

Cell proliferation of NIH/3T3 fibroblasts seeded on 14 various substrates was visualized by live/dead fluorescence staining (Figure 8A). No dead cells were observed for any of the hydrogels. All LMGG hydrogels and HMGG hydrogels synthesized only via chain growth showed the highest cell proliferation. LMGG mixed model gels showed lower cell proliferation than LMGG step growth gels. Cell proliferation on HMGG hydrogels was correlated well with the hydrogel compressive modulus (R = 0.80). The softest HMGG gels – those formed using equimolar thiol-ene ratios – yielded the lowest cell proliferation (4.6 

0.5% in the absence and 5.7  0.4% in the presence of calcium). The stiffest HMGG substrates – those formed through chain growth crosslinking – had the highest cell proliferation (46.5  5.7% in the absence and 52.9  5.8 in the presence of calcium).

It is known that mechanical signals play an important role in myofibroblast differentiation, activation, and adhesion.156 Fibroblasts differentiate into a contractile myofibroblast phenotype when exposed to substrates with elastic moduli corresponding to pathologically stiff fibrotic tissue with microfilament bundles or mature focal adhesions,157 while those cultured on softer substrate show little development of stress fibers.158,159

Fluorescent images of cells stained with live-and-dead assay were used to visualize cell morphology. Here, fibroblasts seeded on the stiffest hydrogel (gel 5) had some stress fibers and microfilament bundles (Figure 8A). Cells seeded on softer substrates (gel 9 and gel 11) formed clusters of cells that did not spread on the surfaces. Conversely, cells seeded on stiffer hydrogels (gel 12 and gel 13) spread more evenly across the surface of the gel.

In addition to mechanical stress, surface chemistry can influence fibroblast morphology and differentiation. As shown in Figure 8, Fibroblasts cultured on half-molar thiol-ene step growth HMGG hydrogels in the absence of calcium showed microfilament bundles even though it was softer than the similarly formed chain growth gel 6.

Microfilament bundles were only observed on substrates fabricated with half-molar thiol-ene ratios. These microfilaments disappeared when the modulus of gel 11 was increased via either more chemical crosslinking (gel 12) or ionic crosslinking (gel 13).

Several studies have demonstrated cell adhesion, spreading and proliferation depends on surface chemistry, topology, and wettability. Combinatorial approaches for developing non-fouling biomaterials have revealed that chemistry plays a role in cell activation160 and mitigating the foreign body response.161 A multi-arm PEG-norbornene macromer was developed via step growth and the resultant thiol-ene hydrogel contained more homogeneous networks. Encapsulated MIN6 β-cells had significantly higher proliferation in the hydrogel

50 formed with thiol-ene chemistry. Additionally, the encapsulated cells in the thiol-ene hydrogels grew in clusters.118 Here, the effect of the thiol content was found to be the major parameter responsible for fibroblast morphology differences. Cell behaviors in relation to thiol-ene group have not been studied extensively. More experiments are necessary to decouple the effect of surface chemistry and mechanical properties on fibroblast differentiation and morphology.

Figure 8. Cytocompatibility of gellan gum hydrogels. NIH/3T3 fibroblasts were seeded on hydrogels and tissue culture plastic (control). (A) Representative micrographs of live (green) and dead (red) cells cultured for 48 h. (B) Quantification of live and dead cells. Date represents the mean±SD. n = 4. Statistical analysis through two-way ANOVA and Tukey’s HSD post-hoc test. Bars with the same letter (A-G) are not statistically different (p<0.05)

Cell response and proliferation on biomaterial surfaces are crucial to biomedical applications and tissue engineering. No dead cells were observed. Cell proliferation data indicates that both LMGG and HMGG hydrogels are biocompatible and bioactive. In addition, cell proliferation of HMGG was controlled, in part, by the tunable mechanical properties of the hydrogel. The design principles of tissue engineering constructs greatly

51 depend on the native mechanical properties of the organ in which the biomaterial will be implanted. The range of compressive modulus from 6.4 to 17.2 kPa of LMGG and HMGG present a potential to be applied as dermal implants. In wound healing process, fibroblasts differentiate into myofibroblasts and deposit newly synthesized ECM 2 weeks after injury.162

Cell proliferation, activation of growth factors, and mechanics of the ECM induce fibrosis.163

Higher elastic moduli (16-20 kPa) of granulation tissue during the repairing stage leads to a significantly higher level of fibroblast differentiation.164 Skin stiffness changes during the proliferation and regenerative stages and becomes stiffer in fibrotic tissue (~50kPa).

Therefore, a soft tissue scaffold for dermal should be in a relatively lower range of 1-20kPa.

164 Our hydrogels have the potential to be applied as biologically active wound dressings and mimic the mechanical properties and microenvironment of the ECM.

4. Conclusion

We successfully modified gellan gum by reacting it with methacrylic anhydride. Both low modified and high modified gellan gum hydrogels were synthesized by varying the amount of methacrylic anhydride. Taking advantage of the thiol-ene click chemistry, hydrogels were fabricated through chain growth, step growth, and mixed model mechanisms.

We demonstrated that the addition of calcium and thiol-ene ratio can significantly alter the compressive modulus of the resultant hydrogels. Swelling ratios of hydrogels were negatively correlated with hydrogel compression modulus. The formation of physical crosslinks through the addition of calcium reduced network flexibility and suppressed hydrogel swelling. NIH/3T3 fibroblasts proliferated on the hydrogels and had a positive correlation with substrate stiffness. Cell morphology and spreading were found to be related to substrate chemistry. Over all, this study demonstrated the ability to tune mechanical properties of gellan gum hydrogel through chemical modification and addition of divalent

52 ions. Data presented here will help build a gellan gum platform for future applications in tissue engineering.

CHAPTER 3. EFFECT OF SUBSTRATE HYDROPHOBICITY ON CELL PROLIFERATION AND POLARIZATION ON GELLAN GUM BASED HYDROGEL

1. Introduction

Hydrogels are frequently used as wound dressing materials owing to their structural similarity to ECM.165 When hydrogels are applied to the wound site, they are able to maintain a moist environment that facilitates the wound healing process by preventing cell apoptosis caused by dehydration and by promoting angiogenesis by providing biomimetic structures for cells to adhere and proliferate.166 Furthermore, enzymatic degradation of hydrogels fabricated from natural derived polymers allows ingrowth of new tissue.63,167

However, wound healing is a complex series of processes and involves activation of various intercellular and intracellular pathways.168,169 Specific design of hydrogels used as wound dressings is necessary to better regulate the wound healing process. Generally, wounds are healed by three distinct and overlapping stages: inflammation, new tissue formation, and remodeling.168 Specifically, hemostasis is achieved by aggregation of platelets during the initial inflammatory stage. Subsequently, neutrophils and macrophages differentiated from monocyte are recruited to sterilize the wound site.169 In the new tissue formation stage, new substrates for cell migration formed near capillaries associate with macrophage and fibroblast. Fibroblasts are stimulated by macrophages and can differentiate into myofibroblasts, which are responsible for closing the edges of the wound site and producing ECM, in the later part of this stage.169 In the last stage, macrophages, fibroblasts and endothelial cells secrete matrix metalloproteinase to remodel the acellular matrix from type III to type I collagen.169

Macrophages play a crucial role in the wound healing process and determine outcome of the foreign body response to implanted biomaterials.170 Macrophages are heterogeneous cells with a spectrum of phenotypes representing distinct sublineages.171 In the presence of specific microenvironmental signals, macrophages are able to switch from one phenotype to another, which means these cells are plastic.171 In general, macrophages classically activated

(M1) by interferon-γ (IFN- γ) or liposaccharides (LPS) represent the pro-inflammatory end of the spectrum.170 These cells are marked by the secretion of tumor necrosis factor (TNF-α) and reactive nitrogen intermediated (RNIs).170 On the other end of spectrum, alternatively activated macrophages (M2) are polarized using interleukin-4 (IL-4) and are pro- angiogenic.170 These cells are characterized by increased levels of IL-10 and transforming growth factor (TGF-β) secretion.170 A balance between these macrophage phenotypes is required for the proper healing from an injury or biomaterial integration.172

Macrophage attachment and polarization on implanted biomaterials can be regulated by surface microstructure and wettability.172 Micro-rough surfaces stimulate more anti- inflammatory cytokines than smooth surfaces.172 In another study, macrophages prefer to migrate onto more hydrophobic substrata.173 It was also found that macrophage apoptosis increased on hydrophilic surfaces when compared to polyethylene terephthalate substrates.174

Upon adhesion t biomaterial substrates, macrophages can release a large number of pro- inflammatory cytokines and oxygen radicals that contribute to cell activation and inflammation.175 These studies indicate that the surface wettability of wound dressings should be carefully designed to favor macrophage attachment and modulate macrophage polarization during wound healing process. However, macrophage adhesion on chemically modified hydrogel surface is poorly understood.

Here, we further modified LMGG with 1-bromohexane and 1-bromoundecane to increase the hydrophobicity of the naturally derived polymer. The wettability was controlled by the length of the hydrophobic branch. Hydrogels were subsequently fabricated through step-growth crosslinking. Physical properties of the hydrogels were characterized.

Macrophages of both phenotypes, along with naïve macrophages were seeded on the resultant hydrogels and their proliferation as well their cytokine secretion was measured.

2. Methods and materials

2.1.Materials

Gellan gum was purchased from Spectrum Chemical Mfg. Corp. (New Brunswick,

NJ); methacrylate anhydride (MA) was obtained from BeanTown Chemical (Hudson, NH);

2-hydroxy-4′-(2-hydroxyethoxy)-2-methylpropiophenone (Irgacure 2959) was obtained from

Sigma-Aldrich (St. Louis, MO); dithiothreitol (DTT) was supplied by VWR Chemical

(Batavia, IL). 1-bromohexane and 1-bromoundecane were obtained from Sigma-Aldrich (St.

Louis, MO); dimethyl sulfoxide (DMSO) was from Fishier Scientific (Fair Lawn, NJ); complete medium (CM, Dulbecco’s modified Eagle’s medium (DMEM, Mediatech, Inc.,

Manassas, VA)); fresh deionized (DI) water (Milli-Q Nanopure, Thermo Scientific,

Waltham, MA) was used throughout this study.

2.2.Hydrophobic modification

Low modified gellan gum (LMGG) was synthesized according to an established protocol mentioned in the above chapter. The obtained LMGG (200 mg) was dissolved in 20 ml DMSO at 60℃ under constant stirring for 1 h to obtain a viscous, homogeneous solution.

To this solution, either 4.55 휇L of 1-bromohexane (0.11 molar ratio to LMGG repeat units) or 8.22 휇L (0.125 molar ratio to LMGG repeat units) of 1-bromouhdecane was added.

Potassium carbonate (10% w/w, 20mg) was dispersed in reaction mixture to enhance the

56 reaction. The reaction was maintained at 60℃ for 12 h and the product was obtained and purified by precipitation followed washing 5 times in 95% ethanol. The purified product was re-dissolved in DI-water and freeze-dried. The final product was a white, cotton-like materials and was stored in a desiccator, protected from light. Modified LMGG was labeled according to the molar ratio of hydrophobic modifier used in the reaction and the chain length of the modifier (H011-C6 and H0125-C11).

2.3.Characterization of the modified gellan gum

To quantify the percent modification of modified LMGG, NMR was used to characterize the polymers. The modified polymers (10 mg) were dissolved in 1 ml D2O. The

1H-NMR spectra were recorded on a Bruker Avance III spectrometer with a sweep width of

6602.1 Hz, a 90° pulse, and an acquisition time of 2.48s. All spectra were obtained at room temperature. A total of 16 repetitive scans with 64 points were acquired and the data were processed in MNova with 128k points, zero filling, and exponential line broadening of 1.0

Hz. The methyl group of methacrylate group on LMGG was used as a reference (훿 = 1.95 ppm). The degree of substitution (DS) was calculated by comparing integral area of methyl group peak at end of hydrophobic branch (훿 = 0.86 ppm) to reference peak and normalized to degree of methacrylate group substitution (0.45) using equation 1.

퐼 DS= 푏푟푎푛푐ℎ 푚푒푡ℎ푦푙 ×0.45 (1) 퐼푚푒푡ℎ푎푐푟푦푙푎푡푒 푚푒푡ℎ푦푙

2.4.Hydrogel fabrication

Hydrogels were fabricated with LMGG, H011-C6, and H0125-C11 and crosslinked using thiol-ene chemistry. Briefly, 10mg LMGG, H011-C6 or H0125-C11 was dissolved in

DI water. To this homogeneous solution, 5 휇L Irgacure 2959 ethanol solution (20% wt) was added with a final 10% initiator concentration. 3.4휇L DTT water solution (10% wt) and 2 휇L

0.5 M calcium chloride were subsequently added. After gentle mixing of reagents, hydrogels were fabricated by putting mixture under UV (15 W, 365 nm, UVP, Upland, CA) light for 10 min.

2.5.Compression modulus

The compressive modulus for each experimental condition was obtained through manual compression measurements. Obtained hydrogel pegs (n = 3) approximately 15  15 

2.6.Swelling

w - w S = s 0 ´100% w0 (2)

2.7.Water contact angle

Surface wettability was measured directly by dropping DI water onto hydrogel coated glass slides. The glass slides were coated with poly-L-lysine (PLL) to create a positively charged layer. Then, 50 휇L negatively charged hydrogel precursor solution was coated on the

PLL coated glass slide and cured under UV irradiation. DI water (10 휇L) was dropped on each hydrogel sample (n = 5). The water contact angle was analyzed using Image J (NIH,

Bethesda, MD).

2.8.Cell culture

RAW 264.7 cells (ATCC, Manassas, VA) were cultured in CM at 37°C in 5% CO2.

Every three to five days the cells were passaged using a cell scraper to detach cells and subcultured at 6.7 × 103 to 2.7 × 104 cells/cm2

2.9.Cell proliferation

Cell proliferation of RAW 264.7 cells in contact with hydrogel substrates was quantified and visualized by live/dead assay. Briefly, to each well of 48-well plate coated with 200 휇L hydrogel, cells were seeded at 125,000 cells/cm2 and 250 휇L CM was added.

Macrophages were activated with 25 ng/mL IL-4 (M (IL-4)), eBioscience Inc., San

Diego,CA) or 500ng/mL LPS (M (LPS)), Sigma Aldrich, St. Louis, MO). Naïve cells were not activated. The plates were incubated for 48 h after which supernatant from the dead controls wre aspired and replaced with 70% ethanol for 20 min. To each well, 150 휇L of live/dead working solution (2 µM calcein AM (AnaSpec, Fremont, CA) and 7.5 µM 7- aminoactinomycin D (Tonbo Biosciences, San Diego, CA) in PBS) was added. The plates were incubated for 30-40 min at 37℃. Live and dead cells were quantified using an excitation/emission of 485/590 nm and 528/645 nm, respectively, using a plate reader

(BioTek Synergy HT Multidetection Microplate Rreader, BioTek, Winooski, VT). RAW

264.7 cells were also fluorescently imaged using a Leica DM i8 microscope (Leica, Wetzlar,

Germany)

2.10. Griess assay

Nitrite production was measured as an M (LPS) indicator by using the supernatant collected from sample wells of 24-well plate after 48 h incubation. A nitrite standard curve was generated using a serial dilution of 100 휇M sodium nitrite with volume of 150 휇L in a 96 well plate. Samples (150 휇L) were also added to the plate. To each well, 130 휇L DI water and 20 휇L Griess reagent (Acros Technologies) were added and the plate was incubated for

20 min. The plate was read at 448 nm with 690 nm as a reference.

2.11. Urea assay

Arginase activity was measured as an M (IL-4) indicator through conversion of L- arginine into urea. The cells obtained in the Griess assay were washed with 400 휇L PBS. The cells were then lysed by incubating at 4℃ for 10 min in 100 휇L cell lysis buffer (150 휇L protease inhibitor cocktail (Amresco, Solon, OH) and 15 휇L Triton X-100 (Acros Organics,

Elgin, IL) diluted to 15 mL with DI water). The lysate solution (25 휇L) was transferred to a

96 well plate along with 25 휇L of 10 mM MnCl2 (Fisher) and 50 mM Tris solution (Fisher).

The plates were then incubated at 55℃ for 10 min to activate the enzymes. Next, 50 휇L of 1

M arginine (pH=9.7) was added to each well and the plate was incubated at 37℃ for 20 h. To a new 96 well plate, 5 휇L of solution was transferred and a colorimetric response was obtained by adding 200 휇L mixture of solution 1 (1.2g o-phthaldiadehyde (Alfa Aesar, Ward

Hill, MA), 1L H 2 O, and 500 µL HCl (Fisher)) and solution 2 (0.6 g N-(naphthyl) ethylenediamine dihydrochloride (Acros Organics) at a 1:2 ratio. The plates were read at 520 nm with a reference at 630 nm.

2.12. Statistics and data analysis

Statistical analysis was performed using JMP statistical software. Statistical

60 significance of the mean comparisons was determined by a two-way ANOVA. Pair-wise comparisons were analyzed with Tukey’s honest significance difference test. Differences were considered statistically significant for p < 0.05.

3. Results and discussion

3.1.LMGG hydrophobic modification and characterization

The hydrophobic modification of LMGG is schematically illustrated in Figure 9. The

1H-NMR spectra of unmodified LMGG (Figure 9 top) and modified LMGG (Figure 9 middle and bottom) were obtained at 25℃. All three spectra exhibited a methyl peak from the methacrylate group on LMGG (훿 = 1.95 ppm) and a methyl peak from the LMGG rhamnose structural unit (훿 = 1.3 ppm). Since the methyl peak on the LMGG rhamnose structural unit was overlapping with the peaks arising from the hydrophobic modifiers, the methyl peak from the methacrylate group was used as a reference. The degree of 1- bromohexane and 1-bromoundecane substitution (DS) was calculate by comparing the integrated peak area from the methyl peak at the terminus of the hydrophobic modifier to the methyl peak arising from methacrylate group on LMGG. This was normalized to the degree of methacrylate group substitution, which was 0.45 in this study. The DS of 1-bromohexane and 1-bromoundecane are equivalent with ~0.045 hydrophobic modifiers attached to each

LMGG repeating unit.

Similar degrees of substitution were achieved by controlling the molar ratio of the modifiers with respect to the LMGG repeating units. In this study, a lower molar ratio of 1- bromohexane (0.11) was required to yield same substitution as 1-bromoundecane (0.125).

This is attributed to effect of molecular size on reactivity. As previously reported, increasing molecular size decreases reactivity.176

Figure 9. Schematic illustration of the synthesis and characterization of hydrophobic modified LMGG. 1H-NMR of LMGG (top), H011-C6 (middle), and H0125-C11(bottom)

3.2.Compression modulus

The compression moduli of the hydrogels fabricated here were obtained by analyzing the stress-strain curve in the 5%-10% strain region. Unmodified LMGG served as a comparison. The step growth half molar thiol-ene ratio from chapter 2 (gel3) was chosen for this study due to its ease of fabrication. As illustrated in Figure 10, all the hydrogels studied here have statistically similar compression moduli. The unmodified LMGG was 9.6 ± 0.73 kPa. Gels fabricated from 1-bromohexane and 1-bromoundecane were very similar

(9.6±0.68 and 9.4±0.74 kPa, respectively).

Figure 10. Mechanical properties of LMGG and modified LMGG hydrogels. Data represents the mean±SD (standard deviation). n=5.Statistical analysis through two-way ANOVA and Tukey's HSD post-hoc test.

Several studies have reported that substrate stiffness has a significant influence on macrophage proliferation and activation.177–179 Macrophages were found to preferentially locate on stiff areas of PEGDA hydrogels with patterned elasticity.177 In another study, macrophage morphology was found be influenced by substrates. The shape of macrophages changed from round to flattened when the substrate stiffness increased from 0.1 kPa to 3

MPa.178 This trend was also observed when modulus of substrate increased from 130 kPa to

240 kPa.179 The same study also found that LPS activation of macrophage was significantly reduced on soft hydrogels.179 Since the aim of this study is to investigate how substrate wettability affects macrophage proliferation and polarization, we aimed to fabricate LMGG and modified LMGG hydrogels with similar stiffness.

Hydrogel stiffness was affect by two factors in this study: chemical crosslinking density and intermolecular interactions. Since the concentration of the functional groups remains unchanged for all LMGG and hydrophobic modified LMGG hydrogels, the hydrophobic modification could be influential on intermolecular interactions. A study on branched chitosan hydrogels showed that hydrogen bonds were disrupted when hydroxyl and amino groups were substituted, causing a reduction in intermolecular interactions.78 In order to have similar compressive moduli for all hydrogels studied here, the degree of substitution of hydrophobic modifier was kept low.

3.3.Swelling of hydrogels

The wet swelling ratio was measured in DI water as an indirect method to ascertain crosslinking density and the intrinsic hydrophobicity of hydrogel material. The swelling ratio of the unmodified hydrogel was 2.67±0.087. The swelling ratios of of the hydrophobic modified LMGG decreased significantly. Gel3-H011-C6 had a swelling ratio 1.50±0.056 and gel3-H0125-C11 was even lower at 1.45±0.019 (Figure 11). Crosslinking density is a factor that can alter hydrogel swelling if the same polymer is used. However, hydrophobic modifications alter the properties of polymer forming the network. Thus, the swelling ratio of the modified hydrogels decreased for two reasons. First, substitution of hydrophobic modifiers reduces the amount of hydroxyl groups available to form hydrogen bonds, which decreases the amount of water associated with the hydrogel.180 Second, the increasing hydrophobicity of the network decreases its ability to take up water.181 Since macrophages will be seeded on the surface of the hydrogels, the bulk swelling ratio may not reflect surface wettability directly. The water contact angle was subsequently measured.

Figure 11. Swelling behavior of LMGG and modified LMGG hydrogels. Data represents the mean±SD (standard deviation). n=5.Statistical analysis through two-way ANOVA and Tukey's HSD post-hoc test. * p<0.05, n.s. no statistical difference

3.4.Water contact angle

The wettability of the hydrogels was measured using water contact angles (Figure

12). Unmodified LMGG was relatively hydrophilic (37.3±1.3°). As was expected, the hydrophobically modified hydrogels had an increased water contact angle. Gel3-H011-C6 was 58.6±1.1° and gel3-H0125-C11 was 69.2±2.3°.

Figure 12. Water contact angle measurement of LMGG and modified LMGG hydrogels. Data represents the mean±SD (standard deviation). n=5.Statistical analysis through two-way ANOVA and Tukey's HSD post-hoc test. * p<0.05

Hydrophobic branches with either 6 or 11 carbons will collapse and form a hydrophobic phase in the hydrated state. We hypothesized that the reduction in wettability o the hydrogels were caused by two factors. First, the hydrophobicity of the hydrogel network was increased by attaching hydrophobic modifiers, which reduced the amount of hydroxyl groups able to associate with water.180 Second, phase segregation induced by the hydrophobic branch and hydrophilic LMGG might alter surface topology by increasing surface roughness.88 As reported previously, water repellency on rough surfaces may be attributed to a Wenzel wetting state in which water droplets pin the surface in a wet-contact mode, or Cassie-Baxter wetting state in which open air pockets form between the water droplet and the surface. An intermediate wetting state between the Wenzel and Cassie-Baxter

66 states is also possible.182–184 However, in this study more experiments are needed to elucidate the mechanism responsible for changing the water contact angle.

3.5.Cell proliferation

Naïve, M(IL-4), and M(LPS) macrophages were seeded on the three hydrogels fabricated here. Quantitative macrophage proliferation on hydrogels was obtained by comparing the live signal on the hydrogels to the live signal on tissue culture plastic (TCP) control. Macrophages of all phenotypes proliferated more on the hydrophobic LMGG hydrogels (Figure 13). Specifically, the naïve macrophage proliferation on gel 3 was 31±2% and remained unchanged on gel3-H011-C6 (30±2%) but increased significantly on gel3-

H0125-C11 (35±1%). This increase in macrophage proliferation was more obvious for

M(LPS) cells, which increased from 28±2% to 33±2%, and was able to reach 42±3% on gel3-H0125-C11. The same trend was observed for M(IL-4) cells, which increased from

19±3% to 25±2% on gel3-H011-C6 and 34±3% on gel3-H0125-C11. Fluorescent images

(Figure 14) of cell density and spreading on the hydrogel surface are in line with these findings.

Macrophage attachment on substrates is influenced by multiple factors. Several studies have revealed that substrate wettability and topology could effectively modulate cell attachment and proliferation.173,175,185 Macrophages were found to have a strong preference to reside on more hydrophobic and more roughened substrates than hydrophilic, smooth surfaces.173 Directional locomotion of macrophage stimulated by hydrophobic surfaces was a possible explanation for this observation, however, a more detailed mechanism is necessary to fully answer this question.173 In addition, cell proliferation might not be influenced by the cell-substrate interaction solely. Protein adsorption is believed to be another factor that

67 regulates cell adhesion.186 The type, quantity, and conformation of adsorbed proteins depends upon surface chemistry and can affect cell adhesion behavior.175 As suggested by several previous studies, maximal fibronectin adsorption was found on moderately wettable substrates,187 while albumin exhibited an growing adsorption trend when substrate contact angle increased up to 80°.186 In our case, increased serum protein adsorption, such as albumin and fibronectin present in cell culture media, might contribute to the phenomena we observed.

Figure 13. Cell proliferation of macrophages seeded on hydrogels. Data represents the mean±SD (standard deviation). n=5. Statistical analysis through two-way ANOVA and Tukey's HSD post-hoc test. p<0.05 (*: gel3-H0125-C11 vs. gel3, @: gel3-H0125-C11 vs. gel3- H011-C6 and &: gel3-H011-C6 vs. gel3)

Figure 14. Representative stitched fluorescent images of live (green) and dead (red) macrophage cells growth and proliferation on hydrogels after 48 h culture

3.6.Cell polarization

Nitrite production evaluated by a Griess assay and was used as a marker of M1 phenotype (Figure 15A). Generally, nitrite production from M(LPS) cells seeded on hydrogels increased with increasing substrate hydrophobicity and ranged from 0.82±0.04

mol/L on unmodified gel 3 to 2.1±0.078 mol/L on gel3-H0125-C11. The production of nitrite from M (LPS) on gel3-H011-C6 was 1.8±0.025 mol/L, significantly higher than on

69 gel3 but lower than on gel3-H011-C11. Nitrites produced from M(0) and M(IL-4) on all three hydrogels showed no significant differences and were 0.44-0.62 mol/L.

Arginase activity, a marker of M2 macrophages, was quantified by a urea assay

(Figure 15B). Arginase activity for M(IL-4) cells increased with increasing hydrophobicity.

M(IL-4) cells had the lowest arginase activity on gel3 (54.8±8.1 mg/dl) and was higher on gel3-H011-C6 (66.6±6.2 mg/dl), although not significantly higher. Arginase activity for

M(IL-4) cells was highest on gel3-H0125-C11 though not significantly different when compared to M(IL-4) cells seeded on gel3-H011-C6. M(0) and M(IL-4) cells had low arginase activity on all gels and were not significant different for any of the cells gels.

(~10mg/dl from M(0) and ~15mg/dl from M(LPS), respectively). Since both arginase and inducible nitrite oxide synthase (iNOS) compete for arginine secreted by macrophage as a precursor, urea: nitrites ratios are commonly used as a functional readout in approximating the M1/M2 polarization170 (Figure 15C). Here, we use the ratio between urea and nitrite to approximate the urea: nitrite rations. The ratio between M(IL-4) urea and nitrite production was expressed as mg/mol and decreased with increased surface hydrophobicity.

Specifically, arginase: iNOS for M(IL-4) on gel 3 was1673±137 mg/mol and maintain a similar level on gel 3-H011-C6 (1569±214 mg/mol), but decreased significantly on gel 3-

H0125-C11 (1186±188 mg/mol). The arginase: iNOS ratio for M(0) and M(LPS) cells on all three gels fell in range of 243±57– 149±24 mg/mol and 189±18– 69±17 mg/mol, respectively.

Increased arginase activity is associated with pro-angiogenesis.170 On the other hand, nitrites are the stable form of NO and result from increased levels of iNOS. High iNOS levels indicates a M1 phenotype and pro-inflammatory immune response.170 Both enzymes compete

70 for arginine as a precursor.188 M1 macrophages use iNOS to breakdown arginine into reactive nitrogen intermediates (RNI) which further produce NO to kill microbes.189 M2 macrophages use arginase to convert arginine into ornithine and polyamines, which are important in collagen metabolism.170 The ratio of urea to nitrites was used to reflect the extent of macrophage activation in either direction.190 Our data showed a decrease in urea:nitrite for all macrophage phenotypes with increasing substrate hydrophobicity. This suggests that hydrophobicity dampens the pro-angiogenic response and may be useful in applications where a pro-inflammatory response is crucial.190 One of such application is at the early stage of the wound healing process where monocytes are recruited to the wound site and differentiate to macrophages. These macrophages are first activated toward the M1 phenotype and initiate inflammation to kill bacteria.168–170,172 A decreased inflammatory response could negatively affect the host’s ability to fight off infection and could influence the generation of de novo tissue.172 However, continued classical activation of macrophages can lead to chronic inflammation and result in the breakdown of healthy tissue surrounding the injury site.172 Therefore, balancing macrophage phenotypes is extremely important in designing wound dressings.

Our results agree with several previous studies which also observed that increased wettability resulted in increased anti-inflammatory cytokine compared to hydrophobic substrates.172,191,192 One group indicated that wettability has a stronger immunomodulatory effect than surface roughness.172 However, these studies used titanium surfaces, which are much stiffer than our susbtrates.

Figure 15. Cytokine secretion of macrophages cultured on hydrogels. (A) Nitrates secreted by naïve, LPS and IL-4 activated macrophages seeded on LMGG and modified LMGG hydrogels, (B) Urea production of naïve, LPS and IL-4 activated macrophages seeded on LMGG and modified LMGG hydrogels, (C) urea: nitrite ratios for all macrophage phenotypes seeded on LMGG and modified LMGG hydrogels. Data represents the mean±SD (standard deviation). n=5.Statistical analysis through two-way ANOVA and Tukey's HSD post-hoc test. p<0.05 (*: gel3-H0125-C11 vs. gel3, @: gel3-H0125-C11 vs. gel3-H011-C6 and &: gel3-H011-C6 vs. gel3)

4. Conclusion

In this study, we provide a strategy to modify the wettability of LMGG gels. The physical properties of the hydrogels, including compression modulus, swelling ratio, and water contact angle, were measured and compared to the unmodified LMGG. M(LPS),

M(IL-4), and M(0) macrophages were seeded on these hydrogels. The macrophages exhibited increasing cell proliferation with decreasing surface wettability. Macrophage activation was altered by the hydrogel substrates with increased nitrite secreted on more hydrophobic substrates. This work as a whole presents a novel approach to regulate macrophage responses through changing hydrogel wettability and offers a new design principle in fabricating hydrogel wound dressings.

REFERENCES

1. Jeevane, Y. Tissue Engineering and Regeneration: Technologies and Global Markets. BCC research (2014).

2. Saul JM, Williams DF. Hydrogels in Regenerative Medicine. Handb Polym Appl Med Med Devices. 2013:279-302. doi:10.1016/B978-0-323-22805-3.00012-8.

3. Yu F, Cao X, Zeng L, Zhang Q, Chen X. An interpenetrating HA / G / CS biomimic hydrogel via Diels – Alder click chemistry for cartilage tissue engineering. Carbohydr Polym. 2013;97(1):188-195. doi:10.1016/j.carbpol.2013.04.046.

4. Boddupalli A, Zhu L, Bratlie KM. Methods for Implant Acceptance and Wound Healing: Material Selection and Implant Location Modulate Macrophage and Fibroblast Phenotypes. Adv Healthc Mater. 2016:2575-2594. doi:10.1002/adhm.201600532.

5. Emileh A, Vasheghani-farahani E, Imani M. Swelling behavior, mechanical properties and network parameters of pH- and temperature-sensitive hydrogels of poly (( 2- dimethyl amino ) ethyl methacrylate- co -butyl methacrylate ). Eur Polym J. 2007;43:1986-1995. doi:10.1016/j.eurpolymj.2007.02.002.

6. Filion TM, Xu J, Prasad ML, Song J. In vivo tissue responses to thermal-responsive shape memory polymer nanocomposites. Biomaterials. 2011;32(4):985-991. doi:10.1016/j.biomaterials.2010.10.012.

7. Coutinho DF, Sant S V., Shin H, et al. Modified Gellan Gum hydrogels with tunable physical and mechanical properties. Biomaterials. 2010;31(29):7494-7502. doi:10.1016/j.biomaterials.2010.06.035.

8. Seidlits SK, Khaing ZZ, Petersen RR, et al. The effects of hyaluronic acid hydrogels with tunable mechanical properties on neural progenitor cell differentiation. Biomaterials. 2010;31(14):3930-3940. doi:10.1016/j.biomaterials.2010.01.125.

9. Li F, He J, Zhang M, Ni P. A pH-sensitive and biodegradable supramolecular hydrogel constructed from a PEGylated polyphosphoester-doxorubicin prodrug and α- cyclodextrin. Polym Chem. 2015;6(28):5009-5014. doi:10.1039/C5PY00620A.

10. Zheng J, Smith Callahan LA, Hao J, et al. Strain-promoted cross-linking of PEG-based hydrogels via copper- free cycloaddition. ACS Macro Lett. 2012;1(8):1071-1073. doi:10.1021/mz3003775.

11. Jiang H, Qin S, Dong H, et al. An injectable and fast-degradable poly(ethylene glycol) hydrogel fabricated via bioorthogonal strain-promoted azide-alkyne cycloaddition click chemistry. Soft Matter. 2015;11(30):6029-6036. doi:10.1039/C5SM00508F.

12. Koehler KC, Anseth KS, Bowman CN. Diels − Alder Mediated Controlled Release from a Poly ( ethylene glycol ) Based Hydrogel. 2013.

13. Kolb HC, Finn MG, Sharpless KB. Click Chemistry : Diverse Chemical Function from a Few Good Reactions.

14. Tan H, Rubin JP, Marra KG. Direct Synthesis of Biodegradable Polysaccharide Derivative Hydrogels Through Aqueous Diels-Alder Chemistry a. 2011:905-911. doi:10.1002/marc.201100125.

15. Wei H, Yang Z, Zheng L, Shen Y. Thermosensitive hydrogels synthesized by fast Diels – Alder reaction in water. Polymer (Guildf). 2009;50(13):2836-2840. doi:10.1016/j.polymer.2009.04.032.

16. Fisher SA, Anandakumaran PN, Owen SC, Shoichet MS. Tuning the Microenvironment : Click-Crosslinked Hyaluronic Acid-Based Hydrogels Provide a Platform for Studying Breast Cancer Cell Invasion. 2015:7163-7172. doi:10.1002/adfm.201502778.

17. Yu F, Cao X, Li Y, Zeng L, Yuan B, Chen X. An injectable hyaluronic acid/PEG hydrogel for cartilage tissue engineering formed by integrating enzymatic crosslinking and Diels–Alder “click chemistry.” Polym Chem. 2014;5(3):1082-1090. doi:10.1039/C3PY00869J.

18. Li F, He J, Zhang M, Ni P. A pH-sensitive and biodegradable supramolecular hydrogel constructed from a PEGylated polyphosphoester-doxorubicin prodrug and α- cyclodextrin. Polym Chem. 2015;6(28):5009-5014. doi:10.1039/C5PY00620A.

19. Fan M, Ma Y, Mao J, Zhang Z, Tan H. Cytocompatible in situ forming chitosan/hyaluronan hydrogels via a metal-free click chemistry for soft tissue engineering. Acta Biomater. 2015;20:60-68. doi:10.1016/j.actbio.2015.03.033.

20. Li KW, Cen L, Zhou C, et al. Well-Defined Poly(ethylene glycol) Hydrogels with Enhanced Mechanical Performance Prepared by Thermally Induced Copper-Catalyzed Azide-Alkyne Cycloaddition. Macromol Mater Eng. 2016;301(11):1374-1382. doi:10.1002/mame.201600222.

21. Reddy SK, Anseth KS, Bowman CN. Modeling of network degradation in mixed step- chain growth polymerizations. Polymer (Guildf). 2005;46(12):4212-4222. doi:10.1016/j.polymer.2005.02.050.

22. Rydholm AE, Bowman CN, Anseth KS. Degradable thiol-acrylate photopolymers: Polymerization and degradation behavior of an in situ forming biomaterial. Biomaterials. 2005;26(22):4495-4506. doi:10.1016/j.biomaterials.2004.11.046.

23. Chalovich JM, Eisenberg E. NIH Public Access. Biophys Chem. 2005;257(5):2432- 2437. doi:10.1016/j.immuni.2010.12.017.Two-stage.

24. Fu A, Gwon K, Kim M, Tae G, Kornfield JA. Visible-light-initiated thiol-acrylate photopolymerization of heparin-based hydrogels. Biomacromolecules. 2015;16(2):497-506. doi:10.1021/bm501543a.

25. Hao Y, Shih H, Muňoz Z, Kemp A, Lin CC. Visible light cured thiol-vinyl hydrogels with tunable degradation for 3D cell culture. Acta Biomater. 2014;10(1):104-114. doi:10.1016/j.actbio.2013.08.044.

26. Roberts JJ, Bryant SJ. Comparison of photopolymerizable thiol-ene PEG and acrylate- based PEG hydrogels for cartilage development. Biomaterials. 2013;34(38):9969- 9979. doi:10.1016/j.biomaterials.2013.09.020.

27. Lee S, Tong X, Yang F. The effects of varying poly(ethylene glycol) hydrogel crosslinking density and the crosslinking mechanism on protein accumulation in three- dimensional hydrogels. Acta Biomater. 2014;10(10):4167-4174. doi:10.1016/j.actbio.2014.05.023.

28. Nicolaou KC, Snyder SA, Montagnon T, Vassilikogiannakis G. The Diels ± Alder Reaction in Total Synthesis.

29. Kirchhof S, Brandl FP, Goepferich AM. linking mechanism for degradable poly ( ethylene glycol ). 2013:4855-4864. doi:10.1039/c3tb20831a.

30. Fan M, Ma Y, Zhang Z, Mao J, Tan H, Hu X. Biodegradable hyaluronic acid hydrogels to control release of dexamethasone through aqueous Diels – Alder chemistry for adipose tissue engineering. 2015;56:311-317. doi:10.1016/j.msec.2015.04.004.

31. Bai X, Lü S, Cao Z, et al. Self-reinforcing injectable hydrogel with both high water content and mechanical strength for bone repair. Chem Eng J. 2016;288:546-556. doi:10.1016/j.cej.2015.12.021.

32. Nimmo CM, Owen SC, Shoichet MS. Diels - Alder Click Cross-Linked Hyaluronic Acid Hydrogels for Tissue Engineering. 2011:824-830.

33. Hsu Y, Masutani K, Yamaoka T, Kimura Y. Strengthening of hydrogels made from enantiomeric block copolymers of polylactide ( PLA ) and poly ( ethylene glycol ) ( PEG ) by the chain extending Diels e Alder reaction at the hydrophilic PEG terminals. Polymer (Guildf). 2015;67:157-166. doi:10.1016/j.polymer.2015.04.026.

34. Wei H, Yang Z, Chu H, Zhu J, Li Z, Cui J. Facile preparation of poly ( N - isopropylacrylamide ) -based hydrogels via aqueous Diels e Alder click reaction. Polymer (Guildf). 2010;51(8):1694-1702. doi:10.1016/j.polymer.2010.02.008.

35. Kirchhof S, Strasser A, Wittmann H, et al. mechanism of Diels – Alder hydrogels †. J Mater Chem B Mater Biol Med. 2014;3:449-457. doi:10.1039/C4TB01680G.

36. Yu F, Cao X, Li Y, et al. Diels–Alder crosslinked HA/PEG hydrogels with high elasticity and fatigue resistance for cell encapsulation and articular cartilage tissue repair. Polym Chem. 2014;5(17):5116-5123. doi:10.1039/C4PY00473F.

37. Guaresti O, González K, Eceiza A, Corcuera MA. Click gelatin hydrogels : Characterization and drug release behaviour. Mater Lett. 2016;182:134-137. doi:10.1016/j.matlet.2016.06.115.

38. Wei Z, Yang JH, Du XJ, et al. Dextran-Based Self-Healing Hydrogels Formed by Reversible Diels – Alder Reaction under Physiological Conditions. (Dcc):1464-1470.

39. Li KW, Cen L, Zhou C, et al. Well-Defined Poly(ethylene glycol) Hydrogels with Enhanced Mechanical Performance Prepared by Thermally Induced Copper-Catalyzed Azide–Alkyne Cycloaddition. Macromol Mater Eng. 2016;301(11):1374-1382. doi:10.1002/mame.201600222.

40. Finn PMG, Fokin V, Finn MG, Fokin V V. Applications of click chemistry themed issue Click chemistry : function follows form w. (4). doi:10.1039/c003740k.

41. Finn PMG, Fokin V, Hein JE, Fokin V V. Applications of click chemistry themed issue. (4). doi:10.1039/b904091a.

42. Sun Z, Liu S, Li K, Tan L, Cen L, Fu G. Well-defined and biocompatible hydrogels with toughening and reversible photoresponsive properties. Soft Matter. 2016;12(7):2192-2199. doi:10.1039/C5SM02129D.

43. Adzima BJ, Tao Y, Kloxin CJ, DeForest CA, Anseth KS, Bowman CN. Spatial and temporal control of the alkyne–azide cycloaddition by photoinitiated Cu(II) reduction. Nat Chem. 2011;3(3):258-261. doi:10.1038/nchem.980.

44. Gong T, Adzima BJ, Baker NH, Bowman CN. Photopolymerization reactions using the photoinitiated copper (I)-catalyzed azide-alkyne cycloaddition (CuAAC) reaction. Adv Mater. 2013;25(14):2024-2028. doi:10.1002/adma.201203815.

45. Takahashi A, Suzuki Y, Suhara T, et al. In situ cross-linkable hydrogel of hyaluronan produced via copper-free click chemistry. Biomacromolecules. 2013;14(10):3581- 3588. doi:10.1021/bm4009606.

46. Dey P, Schneider T, Chiappisi L, Gradzielski M, Schulze-Tanzil G, Haag R. Mimicking of Chondrocyte Microenvironment Using in Situ Forming Dendritic Polyglycerol Sulfate-Based Synthetic Polyanionic Hydrogels. Macromol Biosci. 2016;16(4):580-590. doi:10.1002/mabi.201500377.

47. Fu S, Dong H, Deng X, Zhuo R, Zhong Z. Injectable hyaluronic acid/poly(ethylene glycol) hydrogels crosslinked via strain-promoted azide-alkyne cycloaddition click reaction. Carbohydr Polym. 2017;169:332-340. doi:10.1016/j.carbpol.2017.04.028.

48. Brown TE, Marozas IA, Anseth KS. Amplified Photodegradation of Cell-Laden Hydrogels via an Addition???Fragmentation Chain Transfer Reaction. Adv Mater. 2017;29(11). doi:10.1002/adma.201605001.

49. Liu X, Miller AL, Fundora KA, Yaszemski MJ, Lu L. Poly(ϵ-caprolactone) Dendrimer Cross-Linked via Metal-Free Click Chemistry: Injectable Hydrophobic Platform for Tissue Engineering. ACS Macro Lett. 2016;5(11):1261-1265. doi:10.1021/acsmacrolett.6b00736.

50. Henise J, Hearn BR, Ashley GW, Santi D V. Biodegradable tetra-peg hydrogels as carriers for a releasable drug delivery system. Bioconjug Chem. 2015;26(2):270-278. doi:10.1021/bc5005476.

51. Xu J, Feng E, Song J. Bioorthogonally cross-linked hydrogel network with precisely controlled disintegration time over a broad range. J Am Chem Soc. 2014;136(11):4105-4108. doi:10.1021/ja4130862.

52. Madl CM, Katz LM, Heilshorn SC. Bio-Orthogonally Crosslinked, Engineered Protein Hydrogels with Tunable Mechanics and Biochemistry for Cell Encapsulation. Adv Funct Mater. 2016;26(21):3612-3620. doi:10.1002/adfm.201505329.

53. Barker K, Rastogi SK, Dominguez J, et al. Biodegradable DNA-enabled poly(ethylene glycol) hydrogels prepared by copper-free click chemistry. J Biomater Sci Polym Ed. 2016;27(1):22-39. doi:10.1080/09205063.2015.1103590.

54. Truong VX, Ablett MP, Gilbert HTJ, et al. In situ-forming robust chitosan- poly(ethylene glycol) hydrogels prepared by copper-free azide–alkyne click reaction for tissue engineering. Biomater Sci. 2014;2(2):167. doi:10.1039/c3bm60159e.

55. Truong VX, Tsang KM, Simon GP, et al. Photodegradable gelatin-based hydrogels prepared by bioorthogonal click chemistry for cell encapsulation and release. Biomacromolecules. 2015;16(7):2246-2253. doi:10.1021/acs.biomac.5b00706.

56. Su X, Bu L, Dong H, Fu S, Zhuo R, Zhong Z. An injectable PEG-based hydrogel synthesized by strain-promoted alkyne–azide cycloaddition for use as an embolic agent. RSC Adv. 2016;6:2904-2909. doi:10.1039/C5RA23551K.

57. Jonker AM, Bode SA, Kusters AH, Van Hest JCM, L??wik DWPM. Soft PEG- Hydrogels with Independently Tunable Stiffness and RGDS-Content for Cell Adhesion Studies. Macromol Biosci. 2015;15(10):1338-1347. doi:10.1002/mabi.201500110.

58. Dey P, Hemmati-Sadeghi S, Haag R. Hydrolytically degradable, dendritic polyglycerol sulfate based injectable hydrogels using strain promoted azide–alkyne cycloaddition reaction. Polym Chem. 2016:375-383. doi:10.1039/C5PY01326G.

59. Hodgson SM, Bakaic E, Stewart SA, Hoare T, Adronov A. Properties of Poly(ethylene glycol) Hydrogels Cross-Linked via Strain-Promoted Alkyne-Azide Cycloaddition (SPAAC). Biomacromolecules. 2016;17(3):1093-1100. doi:10.1021/acs.biomac.5b01711.

60. Xu J, Filion TM, Prifti F, Song J. Cytocompatible poly(ethylene glycol)-co- polycarbonate hydrogels cross-linked by copper-free, strain-promoted click chemistry. Chem - An Asian J. 2011;6(10):2730-2737. doi:10.1002/asia.201100411.

61. DeForest CA, Tirrell DA. A photoreversible protein-patterning approach for guiding stem cell fate in three-dimensional gels. Nat Mater. 2015;14(5):523-531. doi:10.1038/nmat4219.

62. Jiang Y, Chen J, Deng C, Suuronen EJ, Zhong Z. Click hydrogels, microgels and nanogels: Emerging platforms for drug delivery and tissue engineering. Biomaterials. 2014;35(18):4969-4985. doi:10.1016/j.biomaterials.2014.03.001.

63. Dey P, Hemmati-Sadeghi S, Haag R. Hydrolytically degradable, dendritic polyglycerol sulfate based injectable hydrogels using strain promoted azide–alkyne cycloaddition reaction. Polym Chem. 2016:375-383. doi:10.1039/C5PY01326G.

64. Adzima BJ, Tao Y, Kloxin CJ, DeForest CA, Anseth KS, Bowman CN. Spatial and temporal control of the alkyne–azide cycloaddition by photoinitiated Cu(II) reduction. Nat Chem. 2011;3(3):258-261. doi:10.1038/nchem.980.

65. Camci-Unal G, Cuttica D, Annabi N, Demarchi D, Khademhosseini A. Synthesis and characterization of hybrid hyaluronic acid-gelatin hydrogels. Biomacromolecules. 2013;14(4):1085-1092. doi:10.1021/bm3019856.

66. Zhong C, Wu J, Reinhart-King CA, Chu CC. Synthesis, characterization and cytotoxicity of photo-crosslinked maleic chitosan-polyethylene glycol diacrylate hybrid hydrogels. Acta Biomater. 2010;6(10):3908-3918. doi:10.1016/j.actbio.2010.04.011.

67. Saraiva SM, Miguel SP, Ribeiro MP, Coutinho P, Correia IJ. Synthesis and characterization of a photocrosslinkable chitosan–gelatin hydrogel aimed for tissue regeneration. RSC Adv. 2015;5(78):63478-63488. doi:10.1039/C5RA10638A.

68. Nguyen NT, Liu JH. Fabrication and characterization of poly(vinyl alcohol)/chitosan hydrogel thin films via UV irradiation. Eur Polym J. 2013;49(12):4201-4211. doi:10.1016/j.eurpolymj.2013.09.032.

69. Qi Z, Xu J, Wang Z, Nie J, Ma G. Preparation and properties of photo-crosslinkable hydrogel based on photopolymerizable chitosan derivative. Int J Biol Macromol. 2013;53:144-149. doi:10.1016/j.ijbiomac.2012.10.021.

70. Mihaila SM, Gaharwar AK, Reis RL, Marques AP, Gomes ME, Khademhosseini A. Photocrosslinkable kappa-carrageenan hydrogels for tissue engineering applications. Adv Healthc Mater. 2013;2(6):895-907. doi:10.1002/adhm.201200317.

71. Amsden BG, Sukarto A, Knight DK, Shapka SN. Methacrylated glycol chitosan as a photopolymerizable biomaterial. Biomacromolecules. 2007;8(12):3758-3766. doi:10.1021/bm700691e.

72. Hutson CB, Nichol JW, Aubin H, et al. Synthesis and Characterization of Tunable Poly(Ethylene Glycol): Gelatin Methacrylate Composite Hydrogels. Tissue Eng Part A. 2011;17(13-14):1713-1723. doi:10.1089/ten.tea.2010.0666.

73. Silva-Correia J, Zavan B, Vindigni V, et al. Biocompatibility evaluation of ionic- and photo-crosslinked methacrylated gellan gum hydrogels: In vitro and in vivo study. Adv Healthc Mater. 2013;2(4):568-575. doi:10.1002/adhm.201200256.

74. Reys LL, Silva SS, Soares Da Costa D, et al. Fucoidan Hydrogels Photo-Cross-Linked with Visible Radiation As Matrices for Cell Culture. ACS Biomater Sci Eng. 2016;2(7):1151-1161. doi:10.1021/acsbiomaterials.6b00180.

75. Cust??dio CA, Reis RL, Mano JF. Photo-Cross-Linked Laminarin-Based Hydrogels for Biomedical Applications. Biomacromolecules. 2016;17(5):1602-1609. doi:10.1021/acs.biomac.5b01736.

76. Kang W, Bi B, Zhuo R, Jiang X. Photocrosslinked methacrylated carboxymethyl chitin hydrogels with tunable degradation and mechanical behavior. Carbohydr Polym. 2017;160:18-25. doi:10.1016/j.carbpol.2016.12.032.

77. Wang X, Hao T, Qu J, Wang C, Chen H. Synthesis of thermal polymerizable alginate- GMA hydrogel for cell encapsulation. J Nanomater. 2015;2015. doi:10.1155/2015/970619.

78. Kim S, Kang Y, Mercado-Pag??n ??ngel E., Maloney WJ, Yang Y. In vitro evaluation of photo-crosslinkable chitosan-lactide hydrogels for bone tissue engineering. J Biomed Mater Res - Part B Appl Biomater. 2014;102(7):1393-1406. doi:10.1002/jbm.b.33118.

79. Han F, Yang X, Zhao J, Zhao Y, Yuan X. Photocrosslinked layered gelatin-chitosan hydrogel with graded compositions for osteochondral defect repair. J Mater Sci Mater Med. 2015;26(4). doi:10.1007/s10856-015-5489-0.

80. Hutson CB, Nichol JW, Aubin H, et al. Synthesis and Characterization of Tunable Poly(Ethylene Glycol): Gelatin Methacrylate Composite Hydrogels. Tissue Eng Part A. 2011;17(13-14):1713-1723. doi:10.1089/ten.tea.2010.0666.

81. Shin H, Olsen BD, Khademhosseini A. The mechanical properties and cytotoxicity of cell-laden double-network hydrogels based on photocrosslinkable gelatin and gellan gum biomacromolecules. Biomaterials. 2012;33(11):3143-3152. doi:10.1016/j.biomaterials.2011.12.050.

82. Hoyle CE, Bowman CN. Thiol-ene click chemistry. Angew Chemie - Int Ed. 2010;49(9):1540-1573. doi:10.1002/anie.200903924.

83. Munoz Z, Shih H, Lin C-C. Gelatin hydrogels formed by orthogonal thiol-norbornene photochemistry for cell encapsulation. Biomater Sci. 2014;2(8):1063-1072. doi:10.1039/c4bm00070f.

84. Zhong C, Wu J, Reinhart-King CA, Chu CC. Synthesis, characterization and cytotoxicity of photo-crosslinked maleic chitosan-polyethylene glycol diacrylate hybrid hydrogels. Acta Biomater. 2010;6(10):3908-3918. doi:10.1016/j.actbio.2010.04.011.

85. Mori T, Murakami M, Okumura M, Kadosawa T, Uede T. Mechanism of Macrophage Activation by Chitin Derivatives. 2004;1.

86. Lin CC, Ki CS, Shih H. Thiol-norbornene photoclick hydrogels for tissue engineering applications. J Appl Polym Sci. 2015;132(8):1-11. doi:10.1002/app.41563.

87. Fairbanks BD, Schwartz MP, Halevi AE, Nuttelman CR, Bowman CN, Anseth KS. A versatile synthetic extracellular matrix mimic via thiol-norbornene photopolymerization. Adv Mater. 2009;21(48):5005-5010. doi:10.1002/adma.200901808.

88. Daniele MA, Adams AA, Naciri J, North SH, Ligler FS. Interpenetrating networks based on gelatin methacrylamide and PEG formed using concurrent thiol click chemistries for hydrogel tissue engineering scaffolds. Biomaterials. 2014;35(6):1845- 1856. doi:10.1016/j.biomaterials.2013.11.009.

89. Rydholm AE, Bowman CN, Anseth KS. Degradable thiol-acrylate photopolymers: Polymerization and degradation behavior of an in situ forming biomaterial. Biomaterials. 2005;26(22):4495-4506. doi:10.1016/j.biomaterials.2004.11.046.

90. Ouyang L, Highley CB, Sun W, Burdick JA. A Generalizable Strategy for the 3D Bioprinting of Hydrogels from Nonviscous Photo-crosslinkable Inks. Adv Mater. 2017;29(8). doi:10.1002/adma.201604983.

91. Greene T, Lin TY, Andrisani OM, Lin CC. Comparative study of visible light polymerized gelatin hydrogels for 3D culture of hepatic progenitor cells. J Appl Polym Sci. 2016;44585:1-10. doi:10.1002/app.44585.

92. Hao Y, Lin CC. Degradable thiol-acrylate hydrogels as tunable matrices for three- dimensional hepatic culture. J Biomed Mater Res - Part A. 2014;102(11):3813-3827. doi:10.1002/jbm.a.35044.

93. Xu K, Fu Y, Chung W, et al. Thiol-ene-based biological/synthetic hybrid biomatrix for 3-D living cell culture. Acta Biomater. 2012;8(7):2504-2516. doi:10.1016/j.actbio.2012.03.049.

94. Toepke MW, Impellitteri NA, Theisen JM, Murphy WL. Characterization of thiol-ene crosslinked PEG hydrogels. Macromol Mater Eng. 2013;298(6):699-703. doi:10.1002/mame.201200119.

95. Mergy J, Fournier A, Hachet E, Auzély-Velty R. Modification of polysaccharides via thiol-ene chemistry: A versatile route to functional biomaterials. J Polym Sci Part A Polym Chem. 2012;50(19):4019-4028. doi:10.1002/pola.26201.

96. Emmakah AM, Arman HE, Bragg JC, et al. A fast-degrading thiol–acrylate based hydrogel for cranial regeneration. Biomed Mater. 2017;12(2):25011. doi:10.1088/1748-605X/aa5f3e.

97. Lin TY, Bragg JC, Lin CC. Designing Visible Light-Cured Thiol-Acrylate Hydrogels for Studying the HIPPO Pathway Activation in Hepatocellular Carcinoma Cells. Macromol Biosci. 2016;16(4):496-507. doi:10.1002/mabi.201500361.

98. Stichler S, Jungst T, Schamel M, et al. Thiol-ene Clickable Poly(glycidol) Hydrogels for Biofabrication. Ann Biomed Eng. 2016;45(1):1-13. doi:10.1007/s10439-016-1633- 3.

99. Li Y, Tan Y, Xu K, Lu C, Wang P. A biodegradable starch hydrogel synthesized via thiol-ene click chemistry. Polym Degrad Stab. 2016;137:75-82. doi:10.1016/j.polymdegradstab.2016.07.015.

100. Gramlich WM, Kim IL, Burdick JA. Synthesis and orthogonal photopatterning of hyaluronic acid hydrogels with thiol-norbornene chemistry. Biomaterials. 2013;34(38):9803-9811. doi:10.1016/j.biomaterials.2013.08.089.

101. Miquelard-Garnier G, Demeures S, Creton C, Hourdet D. Synthesis and rheological behavior of new hydrophobically modified hydrogels with tunable properties. Macromolecules. 2006;39(23):8128-8139. doi:10.1021/ma061361n.

102. Truong VX, Tsang KM, Forsythe JS. Nonswelling Click-Cross-Linked Gelatin and PEG Hydrogels with Tunable Properties Using Pluronic Linkers. Biomacromolecules. 2017:acs.biomac.6b01601. doi:10.1021/acs.biomac.6b01601.

103. Macdougall LJ, Truong VX, Dove AP. Efficient In Situ Nucleophilic Thiol-yne Click Chemistry for the Synthesis of Strong Hydrogel Materials with Tunable Properties. ACS Macro Lett. 2017:93-97. doi:10.1021/acsmacrolett.6b00857.

104. Xu Y, Xu H, Jiang X, Yin J. Versatile functionalization of the micropatterned hydrogel of hyperbranched poly(ether amine) based on “thiol-yne” chemistry. Adv Funct Mater. 2014;24(12):1679-1686. doi:10.1002/adfm.201302139.

105. Ganivada MN, Kumar P, Shunmugam R. A unique polymeric gel by thiol–alkyne click chemistry. RSC Adv. 2015;5(62):50001-50004. doi:10.1039/C5RA06339F.

106. Peppas BNA, Hilt JZ, Khademhosseini A, Langer R. Hydrogels in Biology and Medicine : From Molecular Principles to Bionanotechnology **. 2006:1345-1360. doi:10.1002/adma.200501612.

107. Saul JM, Williams DF. Hydrogels in Regenerative Medicine. In: Handbook of Polymer Applications in Medicine and Medical Devices. ; 2013:279-302. doi:10.1016/B978-0-323-22805-3.00012-8.

108. McCall JD, Anseth KS. Thiol-ene photopolymerizations provide a facile method to encapsulate proteins and maintain their bioactivity. Biomacromolecules. 2012;13(8):2410-2417. doi:10.1021/bm300671s.

109. Aimetti AA, Machen AJ, Anseth KS. Poly(ethylene glycol) hydrogels formed by thiol- ene photopolymerization for enzyme-responsive protein delivery. Biomaterials. 2009;30(30):6048-6054. doi:10.1016/j.biomaterials.2009.07.043.

110. Cui J, Lackey MA, Tew GN, Crosby AJ. Mechanical properties of end-linked PEG/PDMS hydrogels. Macromolecules. 2012;45(15):6104-6110. doi:10.1021/ma300593g.

111. Mũnoz Z, Shih H, Lin C-C. Gelatin hydrogels formed by orthogonal thiol–norbornene photochemistry for cell encapsulation. Biomater Sci. 2014;2(8):1063. doi:10.1039/c4bm00070f.

112. Daniele MA, Adams AA, Naciri J, North SH, Ligler FS. Interpenetrating networks based on gelatin methacrylamide and PEG formed using concurrent thiol click chemistries for hydrogel tissue engineering scaffolds. Biomaterials. 2014;35(6):1845- 1856. doi:10.1016/j.biomaterials.2013.11.009.

113. Yu F, Cao X, Li Y, et al. Diels–Alder crosslinked HA/PEG hydrogels with high elasticity and fatigue resistance for cell encapsulation and articular cartilage tissue repair. Polym Chem. 2014;5(17):5116-5123. doi:10.1039/C4PY00473F.

114. Bai X, Lü S, Cao Z, et al. Self-reinforcing injectable hydrogel with both high water content and mechanical strength for bone repair. Chem Eng J. 2016;288:546-556. doi:10.1016/j.cej.2015.12.021.

115. Yu F, Cao X, Li Y, Zeng L, Yuan B, Chen X. An injectable hyaluronic acid/PEG hydrogel for cartilage tissue engineering formed by integrating enzymatic crosslinking and Diels–Alder “click chemistry.” Polym Chem. 2014;5(3):1082-1090. doi:10.1039/C3PY00869J.

116. Zhang Y, An D, Pardo Y, et al. High-water-content and Resilient PEG-containing Hydrogels with Low Fibrotic Response. Acta Biomater. 2017. doi:10.1016/j.actbio.2017.02.028.

117. Roberts JJ, Bryant SJ. Comparison of photopolymerizable thiol-ene PEG and acrylate- based PEG hydrogels for cartilage development. Biomaterials. 2013;34(38):9969- 9979. doi:10.1016/j.biomaterials.2013.09.020.

118. Lin CC, Raza A, Shih H. PEG hydrogels formed by thiol-ene photo-click chemistry and their effect on the formation and recovery of insulin-secreting cell spheroids. Biomaterials. 2011;32(36):9685-9695. doi:10.1016/j.biomaterials.2011.08.083.

119. Baudis S, Bomze D, Markovic M, Gruber P, Ovsianikov A, Liska R. Modular Material System for the Microfabrication of Biocompatible Hydrogels Based on Thiol – Ene-Modified Poly ( vinyl alcohol ). 2016:2060-2070. doi:10.1002/pola.28073.

120. Ossipov DA, Hilborn J. Poly(vinyl alcohol)-based hydrogels formed by “click chemistry.” Macromolecules. 2006;39(5):1709-1718. doi:10.1021/ma052545p.

121. Baudis S, Bomze D, Markovic M, Gruber P, Ovsianikov A, Liska R. Modular Material System for the Microfabrication of Biocompatible Hydrogels Based on Thiol – Ene-Modified Poly ( vinyl alcohol ). 2016:2060-2070. doi:10.1002/pola.28073.

122. Huang L, Xiao C. Formation of a tunable starch-based network by in situ incorporation of mercapto groups and a subsequent thiol-ene click reaction. 2013;(December 2011):427-431. doi:10.1002/pi.4327.

123. Nguyen NT, Liu JH. Fabrication and characterization of poly(vinyl alcohol)/chitosan hydrogel thin films via UV irradiation. Eur Polym J. 2013;49(12):4201-4211. doi:10.1016/j.eurpolymj.2013.09.032.

124. Zhou HY, Guang X, Kong M, Sheng C, Su D, Kennedy JF. Effect of molecular weight and degree of chitosan deacetylation on the preparation and characteristics of chitosan

thermosensitive hydrogel as a delivery system. 2008;73:265-273. doi:10.1016/j.carbpol.2007.11.026.

125. Kim S, Kang Y, Mercado-Pag??n ??ngel E., Maloney WJ, Yang Y. In vitro evaluation of photo-crosslinkable chitosan-lactide hydrogels for bone tissue engineering. J Biomed Mater Res - Part B Appl Biomater. 2014;102(7):1393-1406. doi:10.1002/jbm.b.33118.

126. Cho IS, Cho MO, Li Z, et al. Synthesis and characterization of a new photocrosslinkable glycol chitosan thermogel for biomedical applications. Carbohydr Polym. 2016;144:59-67. doi:10.1016/j.carbpol.2016.02.029.

127. Hong Y, Song H, Gong Y, Mao Z, Gao C, Shen J. Covalently crosslinked chitosan hydrogel: Properties of in vitro degradation and chondrocyte encapsulation. Acta Biomater. 2007;3(1):23-31. doi:10.1016/j.actbio.2006.06.007.

128. Saraiva SM, Miguel SP, Ribeiro MP, Coutinho P, Correia IJ. Synthesis and characterization of a photocrosslinkable chitosan–gelatin hydrogel aimed for tissue regeneration. RSC Adv. 2015;5(78):63478-63488. doi:10.1039/C5RA10638A.

129. Truong VX, Tsang KM, Forsythe JS. Nonswelling Click-Cross-Linked Gelatin and PEG Hydrogels with Tunable Properties Using Pluronic Linkers. Biomacromolecules. 2017:acs.biomac.6b01601. doi:10.1021/acs.biomac.6b01601.

130. Gwon K, Kim E, Tae G. Heparin-hyaluronic acid hydrogel in support of cellular activities of 3D encapsulated adipose derived stem cells. Acta Biomater. 2017;49:284- 295. doi:10.1016/j.actbio.2016.12.001.

131. Lokeshwar B, Lokeshwar BL, Pham T, Block L. Association of Elevated Levels of Hyaluronidase , a Matrix-degrading Enzyme , with Prostate Cancer Progression ’. 1996.

132. Orkint RW, Toole BP. Isolation and Characterization of Hyaluronidase from Cultures of Chick Embryo Skin- and Muscle-derived Fibroblasts*. 1980;2(3):1036-1042.

133. Obara K, Ishihara M, Ishizuka T, et al. Photocrosslinkable chitosan hydrogel containing fibroblast growth factor-2 stimulates wound healing in healing-impaired db/db mice. Biomaterials. 2003;24(20):3437-3444. doi:10.1016/S0142- 9612(03)00220-5.

134. Zhu J, Marchant RE. Design properties of hydrogel tissue-engineering scaffolds. Expert Rev Med Devices. 2011;8(5):607-626. doi:10.1586/erd.11.27.

135. Van Vlierberghe S, Dubruel P, Schacht E. Biopolymer-based hydrogels as scaffolds for tissue engineering applications: A review. Biomacromolecules. 2011;12(5):1387- 1408. doi:10.1021/bm200083n.

136. Gordon S, Todd J, Cohn ZA. In vitro synthesis and secretion of lysozyme by mononuclear phagocytes. J Exp Med. 1974;139:1228-1248.

137. Suri S, Banerjee R. In vitro evaluation of in situ gels as short term vitreous substitutes. J Biomed Mater part A. 2006;79A:650-664. doi:10.1002/jbm.a.30917.

138. Matsukawa S, Tang Z, Watanabe T. Hydrogen-bonding behavior of gellan in solution during structural change observed by 1H NMR and circular dichroism methods. Prog Colloid Polym Sci. 1999;114:15-24.

139. Hadjizadeh A, Doillon CJ. Directional migration of endothelial cells towards angiogenesis using polymer fibres in a 3D co-culture system. J Tissue Eng Regen Med. 2010;4(7):524-531. doi:10.1002/term.

140. Nichol JW, Koshy ST, Bae H, Hwang CM, Yamanlar S, Khademhosseini A. Cell- laden microengineered gelatin methacrylate hydrogels. Biomaterials. 2010;31(21):5536-5544. doi:10.1016/j.biomaterials.2010.03.064.

141. J. Silva-Correia1, J.M.Oliveira, S.G.Caridade, J.T.Oliveira, R. A.Sousa JFM, Reis and RL. gellan gum-based hydrogels for intervertebral disc tissue-engineering applications. J Tissue Eng Regen Med. 2010;4(7):524-531. doi:10.1002/term.363.

142. Shih H, Lin CC. Cross-linking and degradation of step-growth hydrogels formed by thiol-ene photoclick chemistry. Biomacromolecules. 2012;13(7):2003-2012. doi:10.1021/bm300752j.

143. Cramer NB, Bowman CN. Kinetics of thiol-ene and thiol-acrylate photopolymerizations with real-time Fourier transform infrared. J Polym Sci Part A Polym Chem. 2001;39(19):3311-3319. doi:10.1002/pola.1314.

144. Tako M, Teruya T, Tamaki Y, Konishi T. Molecular origin for rheological characteristics of native gellan gum. Colloid Polym Sci. 2009;287(12):1445-1454. doi:10.1007/s00396-009-2112-2.

145. Ferris CJ, Gilmore KJ, Wallace GG, Panhuis M in het. Modified gellan gum hydrogels for tissue engineering applications. Soft Matter. 2013;9(14):3705-3711. doi:10.1039/C3SM27389J.

146. Huang Y, Yu H, Xiao C. pH-sensitive cationic guar gum/poly (acrylic acid) polyelectrolyte hydrogels: Swelling and in vitro drug release. Carbohydr Polym. 2007;69(4):774-783. doi:10.1016/j.carbpol.2007.02.016.

147. Grover GN, Rao N, Christman KL, et al. proliferation tissue engineering A fast- degrading thiol – acrylate based hydrogel for cranial regeneration.

148. Qiu Y, Park K. Environment-sensitive hydrogels for drug delivery. 2001;53:321-339.

149. Frohm M, Gunne H, Bergman a C, et al. Biochemical and antibacterial analysis of human wound and blister fluid. Eur J Biochem. 1996;237(1):86-92. doi:DOI 10.1111/j.1432-1033.1996.0086n.x.

150. Charernsriwilaiwat N, Opanasopit P, Rojanarata T, Ngawhirunpat T. Lysozyme- loaded, electrospun chitosan-based nanofiber mats for wound healing. Int J Pharm. 2012;427(2):379-384. doi:10.1016/j.ijpharm.2012.02.010.

151. Pangburn SH, Trescony P V., Heller J. Lysozyme degradation of partially deacetylated chitin, its films and hydrogels. Biomaterials. 1982;3(2):105-108. doi:10.1016/0142- 9612(82)90043-6.

152. Tomihata K, Ikada Y. In vitro and in vivo degradation of films of chitin and its deacetylated derivatives. Biomaterials. 1997;18(7):567-575. doi:10.1016/S0142- 9612(96)00167-6.

153. De Jong SJ, Van Eerdenbrugh B, Van Nostrum CF, Kettenes-van den Bosch JJ, Hennink WE. Physically crosslinked dextran hydrogels by stereocomplex formation of lactic acid oligomers: Degradation and protein release behavior. J Control Release. 2001;71(3):261-275. doi:10.1016/S0168-3659(01)00228-0.

154. Jin R, Moreira Teixeira LS, Dijkstra PJ, et al. Injectable chitosan-based hydrogels for cartilage tissue engineering. Biomaterials. 2009;30(13):2544-2551. doi:10.1016/j.biomaterials.2009.01.020.

155. Blake ACCF, Johnson LN, Mair GA, North ACT, Phillips DC, Sarma VR. Crystallographic Studies of the Activity of Hen Egg-White Lysozyme Source : Proceedings of the Royal Society of London . Series B , Biological Sciences , Vol . 167 , No . 1009 , A Discussion on the Structure and Function of Lysozyme ( Apr . 18 , 1967 ), p. 2017;167(1009).

156. Hinz B. The myofibroblast: Paradigm for a mechanically active cell. J Biomech. 2010;43(1):146-155. doi:10.1016/j.jbiomech.2009.09.020.

157. Hinz B, Mastrangelo D, Iselin CE, Chaponnier C, Gabbiani G. Mechanical tension controls granulation tissue contractile activity and myofibroblast differentiation. Am J Pathol. 2001;159(3):1009-1020. doi:10.1016/S0002-9440(10)61776-2.

158. Balestrini JL, Chaudhry S, Sarrazy V, Koehler A, Hinz B. The mechanical memory of lung myofibroblasts. Integr Biol. 2012;4(4):410. doi:10.1039/c2ib00149g.

159. Discher DE, Janmey P, Wang Y-L. Tissue cells feel and respond to the stiffness of their substrate. Science. 2005;310(5751):1139-1143. doi:10.1126/science.1116995.

160. Brocchini S, James K, Tangpasuthadol V, Kohn J. Structure-property correlations in a combinatorial library of degradable biomaterials. J Biomed Mater Res. 1998;42(1):66- 75. doi:10.1002/(SICI)1097-4636(199810)42:1<66::AID-JBM9>3.0.CO;2-M.

161. Vegas AJ, Veiseh O, Doloff JC, et al. Combinatorial hydrogel library enables identification of materials that mitigate the foreign body response in primates. Nat Biotechnol. January 2016. doi:10.1038/nbt.3462.

162. Shevchenko R V., James SL, James SE. A review of tissue-engineered skin bioconstructs available for skin reconstruction. J R Soc Interface. 2010;7(43):229-258. doi:10.1098/rsif.2009.0403.

163. Wells RG. Tissue mechanics and fibrosis. Biochim Biophys Acta - Mol Basis Dis. 2013;1832(7):884-890. doi:10.1016/j.bbadis.2013.02.007.

164. Goffin JM, Pittet P, Csucs G, Lussi JW, Meister JJ, Hinz B. Focal adhesion size controls tension-dependent recruitment of ??-smooth muscle actin to stress fibers. J Cell Biol. 2006;172(2):259-268. doi:10.1083/jcb.200506179.

165. Rusanu A, Tama AI, Vulpe R, Rusu A. Biocompatible and Biodegradable Hydrogels Based on Chitosan and Gelatin with Potential Applications as Wound Dressings. 2017. doi:10.1166/jnn.2017.14298.

166. Huang X, Zhang Y, Zhang X, Xu L, Chen X, Wei S. Influence of radiation crosslinked carboxymethyl-chitosan/gelatin hydrogel on cutaneous wound healing. Mater Sci Eng C. 2013;33(8):4816-4824. doi:10.1016/j.msec.2013.07.044.

167. Mabry KM, Schroeder ME, Payne SZ, Anseth KS. Three-Dimensional High- Throughput Cell Encapsulation Platform to Study Changes in Cell-Matrix Interactions. ACS Appl Mater Interfaces. 2016;8(34):21914-21922. doi:10.1021/acsami.5b11359.

168. Broughton, II G, Janis JE, Attinger CE. The basic science of wound healing. Plast Reconstr Surg. 2006;117(7S):12-34. doi:10.1097/01.prs.0000225430.42531.c2.

169. Gurtner GC, Werner S, Barrandon Y, Longaker MT. Wound repair and regeneration. Nature. 2008;453(7193):314-321. doi:10.1038/nature07039.

170. Bygd HC, Forsmark KD, Bratlie KM. Altering invivo macrophage responses with modified polymer properties. Biomaterials. 2015;56:187-197. doi:10.1016/j.biomaterials.2015.03.042.

171. Bygd HC, Forsmark KD, Bratlie KM. The significance of macrophage phenotype in cancer and biomaterials. Clin Transl Med. 2014;3(1):62. doi:10.1186/s40169-014- 0041-2.

172. Hotchkiss KM, Reddy GB, Hyzy SL, Schwartz Z, Boyan BD, Olivares-Navarrete R. Titanium surface characteristics, including topography and wettability, alter macrophage activation. Acta Biomater. 2016;31:425-434. doi:10.1016/j.actbio.2015.12.003.

173. Rich A, Harris AK. Anomalous preferences of cultured macrophages for hydrophobic and roughened substrata. J Cell Sci. 1981;50:1-7. http://www.ncbi.nlm.nih.gov/pubmed/7033247.

174. Brodbeck WG, Patel J, Voskerician G, et al. Biomaterial adherent macrophage apoptosis is increased by hydrophilic and anionic substrates in vivo. Proc Natl Acad Sci U S A. 2002;99(16):10287-10292. doi:10.1073/pnas.162124199.

175. Shen M, Horbett TA. The effects of surface chemistry and adsorbed proteins on monocyte/macrophage adhesion to chemically modified polystyrene surfaces. J Biomed Mater Res. 2001;57(3):336-345. doi:10.1002/1097- 4636(20011205)57:3<336::AID-JBM1176>3.0.CO;2-E.

176. Ab L-. T h e kinetics o f polymerisation reactions. 1935;(178).

177. Nemir S, Hayenga HN, West JL. PEGDA hydrogels with patterned elasticity: Novel tools for the study of cell response to substrate rigidity. Biotechnol Bioeng. 2010;105(3):636-644. doi:10.1002/bit.22574.

178. Féréol S, Fodil R, Labat B, et al. Sensitivity of alveolar macrophages to substrate mechanical and adhesive properties. Cell Motil Cytoskeleton. 2006;63(6):321-340. doi:10.1002/cm.20130.

179. Blakney AK, Swartzlander MD, Bryant SJ. The effects of substrate stiffness on the in vitro activation of macrophages and in vivo host response to poly(ethylene glycol)- based hydrogels. J Biomed Mater Res. 2013;100(6):1375-1386. doi:10.1002/jbm.a.34104.The.

180. Huang Y, Yu H, Xiao C. pH-sensitive cationic guar gum/poly (acrylic acid) polyelectrolyte hydrogels: Swelling and in vitro drug release. Carbohydr Polym. 2007;69(4):774-783. doi:10.1016/j.carbpol.2007.02.016.

181. Qiu Y, Park K. Environment-sensitive hydrogels for drug delivery. Adv Drug Deliv Rev. 2012;64(SUPPL.):49-60. doi:10.1016/j.addr.2012.09.024.

182. Oliveira NM, Zhang YS, Ju J, et al. Hydrophobic Hydrogels: Toward Construction of Floating (Bio)microdevices. Chem Mater. 2016;28(11):3641-3648. doi:10.1021/acs.chemmater.5b04445.

183. Wang S, Jiang L. Definition of superhydrophobic states. Adv Mater. 2007;19(21):3423-3424. doi:10.1002/adma.200700934.

184. Xia F, Jiang L. Bio-inspired, smart, multiscale interfacial materials. Adv Mater. 2008;20(15):2842-2858. doi:10.1002/adma.200800836.

185. Harris CA, Resau JH, Hudson EA, et al. Effects of surface wettability, flow, and protein concentration on macrophage and astrocyte adhesion in an in vitro model of central nervous system catheter obstruction. J Biomed Mater Res - Part A. 2011;97 A(4):433-440. doi:10.1002/jbm.a.33078.

186. Wei J, Igarashi T, Okumori N, et al. Influence of surface wettability on competitive protein adsorption and initial attachment of osteoblasts. Biomed Mater. 2009;4(4). doi:10.1088/1748-6041/4/4/045002.

187. van Wachem PB, Hogt AH, Beugeling T, et al. Adhesion of cultured human endothelial cells onto methacrylate polymers with varying surface wettability and charge. Biomaterials. 1987;8(5):323-328. doi:10.1016/0142-9612(87)90001-9.

188. Mills CD, Lenz LL, Ley K. Macrophages at the fork in the road to health or disease. Front Immunol. 2015;6(FEB):1-6. doi:10.3389/fimmu.2015.00059.

189. Pekarova M, Lojek A. The crucial role of L-arginine in macrophage activation: What you need to know about it. Life Sci. 2015;137:44-48. doi:10.1016/j.lfs.2015.07.012.

190. Bygd HC, Bratlie KM. The effect of chemically modified alginates on macrophage phenotype and biomolecule transport. J Biomed Mater Res - Part A. 2016;104(7):1707-1719. doi:10.1002/jbm.a.35700.

191. Alfarsi MA, Hamlet SM, Ivanovski S. Titanium surface hydrophilicity modulates the human macrophage inflammatory cytokine response. J Biomed Mater Res - Part A. 2014;102(1):60-67. doi:10.1002/jbm.a.34666.

192. Hamlet S, Alfarsi M, George R, Ivanovski S. The effect of hydrophilic titanium surface modification on macrophage inflammatory cytokine gene expression. Clin Oral Implants Res. 2012;23(5):584-590. doi:10.1111/j.1600-0501.2011.02325.x.