Novel Approaches for Patterning Hierarchical Hydrogels

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Novel Approaches for Patterning

Hierarchical Hydrogels

Submitted by

Shuang Wang

B. Eng. (Hons)

M. Sc.

Submitted in fulfilment of the requirements for the degree of

IF80 Master of Philosophy

Chemistry, Physics and Mechanical Engineering (CPME)

Faculty of Science and Engineering (SEF)

Queensland University of Technology

2018

Abstract

Synthetic hydrogels featuring tunable biological functionalities and hierarchical structures are of compelling interest as scaffolds for tissue engineering applications.

With the expectation of regulating cell fate within the soft materials, many efforts have been placed on creating niches that can mimic the complexity of the native extracellular matrix, where biological cues are presented and mass transportation is facilitated by interconnected pore network. In this study, sacrificial moulding process was used to produce porous hydrogels, while two patterning approaches were developed to site-specifically immobilize molecules inside the hydrogels via thiol-

Michael addition, resembling natural extracellular matrix networks in terms of geometrical interconnectivity and cell-guidance functionalization. The simple approaches allow reproducible control over the size and architecture of the channels, as well as the spatial distribution and concentration of the patterning molecules, enabling controlled study of cell-substrate behaviour.

Keywords

Hydrogel, fugitive ink, poly(ethylene glycol), spatial patterning, transfer molecules, fused deposition modelling, melt electrospinning writing, template, scaffold, adhesion peptide, perfusion, NIH-3T3 cells.

Table of Contents

Abstract ...... i Keywords ...... ii Table of Contents ...... iii List of Figures and Tables ...... vi Publication and Conference Presentation...... xii List of Abbreviations ...... xiii Statement of Original Authorship ...... xvi Acknowledgements ...... xvii Chapter 1 Introduction ...... 1 1.1 Background ...... 1 1.2 Thesis Overview ...... 3 Chapter 2 Literature Review ...... 5 2.1 Introduction ...... 5 2.2 Patterning Biochemical Signals ...... 7 2.2.1 The Need for Modifying Biosignals ...... 7 2.2.2 Chemistry Strategies for Gelation and Patterning ...... 8 2.2.3 Surface Patterning ...... 11 2.2.4 Three-dimensional Patterning ...... 15 2.3 Patterning Complex Geometry ...... 19 2.3.1 The Need and Challenges for Structural Manipulation ...... 19 2.3.2 Soft Lithography ...... 20 2.3.3 Template Moulding ...... 21 2.3.4 Site-specific Photoablation ...... 24 2.3.5 Layer-by-layer Photopatterning ...... 25 2.4 Integrating Structural and Biochemical Features ...... 27 2.4.1 Patterned Microwells and Microgrooves ...... 27 2.4.2 Patterned Channels ...... 28

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2.4.3 Multi-compartmental Hydrogels ...... 30 2.5 Overview and Perspectives ...... 31 Chapter 3 Spatial Patterning of Hydrogels via 3D Covalent Transfer Stamping from a Fugitive Ink ...... 33 Abstract ...... 33 3.1 Introduction ...... 33 3.2 Materials and Methods ...... 35 3.2.1 Materials ...... 35 3.2.2 Methods ...... 36 3.3 Results and Discussion ...... 42 3.3.1 The Thiol-Michael Addition Reactions ...... 42 3.3.2 Fabrication of the Fugitive Ink Stamp ...... 47 3.3.3 The Release Profile of CPM from the CPM/PCL Stamps ...... 49 3.3.4 The 3D Covalent Transfer Stamping (3D-CTS) Process ...... 51 3.3.5 Characterization of the Patterned Hydrogels ...... 53 3.4 Evaluation of the Approach ...... 56 3.4.1 Advantages of the Approach ...... 56 3.4.2 Limitations of the Approach ...... 57 3.5 Conclusions ...... 57 Chapter 4 Multi-scalable Peptide-modified Channels in Hydrogels for Cell Study .. 59 Abstract ...... 59 4.1 Introduction ...... 59 4.2 Materials and Methods ...... 61 4.2.1 Materials ...... 61 4.2.2 Methods ...... 62 4.3 Results and Discussion ...... 70 4.3.1 Fabricating PCL Template using FDM and MEW ...... 70 4.3.2 Characterization of the PEG-4Mal/PEG-4SH Hydrogels ...... 71 4.3.3 The Efficiency of the Peptide Patterning ...... 74 4.3.4 Characterization of the Patterned Channels ...... 76 4.3.5 Cell Culture in the Channels ...... 79 4.4 Evaluation of the Approach ...... 83 4.4.1 Advantages of the Approach ...... 83 4.4.2 Limitations of the Approach ...... 84 iv

4.5 Conclusions ...... 84 Chapter 5 Conclusions and Future Directions ...... 85 Reference...... 88

List of Figures and Tables

Figure 2.1 Comparison between natural ECM and synthetic substrates

Table 2.1 Chemical reactions for hydrogel synthesis and patterning

Figure 2.2 Orthogonal chemistries for the creation of diverse biochemical patterns

Figure 2.3 (a) Conventional μCP process (up) and trans-print process using a PVA film (down). (b) Polyacrylamide patterning from the patterned PNIPAM brushes grafted onto a glass template

Figure 2.4 Fluorescence image of a PEGDA hydrogel patterned with ACRL-PEG-

RGDS (in green) and ACRL-PEG-REDV (in red) (a). HDFs only bound to RGDS patterned region but not to REDV patterned regions (b). Phase contrast image of multilayer 3D patterns generated by sequential patterning of a PEGDA precursor solution on PEGDA hydrogel base

Figure 2.5 Thiol-ene reaction for post network formation (a) to introduce three different fluorescently labelled peptide sequences (b) and micrometre-scale spatial patterning (c) within PEG gel

Figure 2.6 Generating adhesive biochemical channels in agarose hydrogel matrices by localization of photo-deprotection reaction and photo-coupling

Figure 2.7 (a) Schematic overview showing 3D printed carbohydrate-glass lattice served as sacrificial element to yield vascular architecture in a monolithic construct;

(b) Microscopic images showing an open perfusable channel in the fibrin gel upon dissolving the glass filament; (c) Confocal microscopy showing matrix with varied

crosslinking mechanisms (red beads), encapsulated cells (10T1/2, green), and the vascular lumen (blue beads)

Figure 2.8 (a) Schematic diagram of the dynamic optical projection stereolithography; (b–e) SEM images of the fabricated PEGDA microwells and micro-architectures; (f and g) Confocal z-projection image and 3D reconstruction of

NIH-3T3 cell culture after 4 days in the microwells

Figure 2.9 (a) Confocal z-stacks showing photo-coupling reaction of fluorescently labelled peptide initiated by multiphoton visible light and subsequent degradation locally with multiphoton ultraviolet light; (b) NIH-3T3 cells migrated to RGD patterned channels (c) Cell outgrowth controlled in three-dimensional functionalized channels

Figure 2.10 Preparation of multi-compartmental hydrogel microparticles using sequential electrospinning and photopatterning

Figure 3.1 Infrared spectra of PCL, 4-arm PEG vinyl sulfone and 4-arm PEG thiol

Figure 3.2 A photograph of the BioExtruder used in this study (a), a Schematic representation of the FDM process that produces laydawn patterns of PCL (b), and a photograph of a typical 4-layer PCL printed structure with a filament diameter of

250 μm, laydown pattern of 0/90o, and filament spacing of 1.6 mm (c)

Table 3.1 Total amount and absorbed density of CPM onto PCL stamp

Figure 3.3 Three regions of interest used to monitor the change of mean fluorescence intensity over time

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Scheme 3.1 Two pathways of thiol-Michael addition reaction: base-catalysed and nucleophile-catalysed mechanisms

Scheme 3.2 Reactivity of commonly utilized vinyl groups in thiol-Michael addition reactions

Figure 3.4 Typical plots of the storage modulus of the stoichiometric precursor from rheological measurements and gelation time under different pH conditions (inset)

Figure 3.5 Typical plots of the storage modulus of the non-stoichiometric precursor from rheological measurements under pH = 7.2 and gelation time of the two precursors (inset)

Scheme 3.3 PEG-4VS crosslinked with PEG-4SH via thiol-Michael addition reaction

Figure 3.6 The fluorescence stereomicroscope images of the CPM/PCL stamps (a) and corresponding lateral images of the PCL filaments (b) using 1 mM (left), 3 mM

(middle), and 5 mM (right) soaking solution.

Scheme 3.4 Reaction scheme showing CPM fluorescence increases upon reacting with thiols

Figure 3.7 Concentration calibration plots of CPM reacted with 10 mM PEG-4SH to create the fluorescent adduct and PEG-4VS control

Figure 3.8 Release profile of CPM into PBS from 1 mM and 3 mM CPM/PCL stamp

Figure 3.9 Bright-field microscope image (1×) of the 1 mM CPM/PCL (left) and 3 mM CPM/PCL stamps (right) embedded in PEG hydrogels (a) and a schematic of the process of 3D-CTS (b)

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Figure 3.10 Representative fluorescence stereomicroscope images (1×) after 0 h (a),

3 h (b) and 6 h (c)

Figure 3.11 The change in mean fluorescence intensity with time of the three ROIs of the CPM/PCL stamp

Figure 3.12 (a) A scheme representing a section view of the sacrificial moulding process after 3D-CTS. (b) A typical bright-field microscope image (3×, inset 1×) of the patterned hydrogel after removing the CPM/PCL stamp. (c–e) The fluorescence stereomicroscope images (3×, inset 1×) of the patterned hydrogels using (c) 1 mM and (d) 3 mM CPM/PCL stamp with 2 h transfer time, and (e) 3 mM CPM/PCL stamp with 6 h transfer time

Figure 3.13 (a) z-stack confocal laser scanning microscope (CLSM) image and corresponding scanning electron microscope (SEM) image of the sacrificial template

(a, inset), and CLSM images of two typical lateral sections (b and c) of the patterned hydrogels using 3 mM CPM/PCL stamp and 2 h transfer time

Figure 3.14 The fluorescence stereomicroscope images (3×, inset 1×) of the patterned hydrogel using 3 mM CPM/PCL stamp for 6 h before (a) and after DMSO treatment (b)

Figure 3.15 Intensity profiles across the channels of the patterned hydrogels using 3 mM CPM/PCL stamp for 2 h (a) and 6 h (b), respectively

Figure 3.16 (a–d) Angle view of the SEM images of the PCL sacrificial templates utilized as fugitive ink. (e–h) Corresponding top view of the fluorescence stereomicroscope images (3×, inset 1×) of the patterned hydrogels with transferring time of 2 h via 3D-CTS and sacrificial moulding.

Figure 4.1 Desktop FlashForge Dreamer 3D printer (a) and custom-built MEW machine (b)

Figure 4.2 Adjustable diameters of the FDM (a) and electrospun PCL filaments (b) by adjusting the nozzle travel speed (a) and the collector speed (b)

Figure 4.3 Stability of CRGDSGK peptide in acetic acid buffer (pH = 4)

Scheme 4.1 Chemical structures of CRGDSGK and FITC-tagged CRGDSGK

Figure 4.4 Fluorescence standard curve of the peptide mixture

Figure 4.5 Standard curve of the peptide mixture

Figure 4.6 Filament diameter adjusted by nozzle travel speed for FDM (a) and collector speed for MEW (b)

Scheme 4.2 PEG-4Mal crosslinked with PEG-4SH via thiol-Michael addition reaction

Table 4.1 Swelling properties of the hydrogels with different solid contents

Figure 4.7 Typical stress-strain curves and the Young’s modulus (inset table) of the hydrogels with different solid contents from uniaxial compression testings

Figure 4.8 Schematic overview of the sacrificial process (a), microscopic images (3×) of the hydrogel with PCL filament (b) and after the removal of the filament (c)

Scheme 4.3 Covalently tethered peptide patterning established by perfusion

Figure 4.9 Fluorescence images of the hydrogel with peptide mixture perfused in the channel for 10 (a) and 30 min (b)

Figure 4.10 Unreacted peptide mixture after 30 min of perfusion in the channel

Figure 4.11 Fluorescence stereomicroscope images (3×) of straight patterned channels with different sizes (a), 2D (b) and 3D (c) patterns (1.5×) and z-stack confocal laser scanning microscope (CLSM) image (10×) of two interconnected channels in (c)

Figure 4.12 (a) Fluorescence microscope images (a left, 3×) and confocal images of the cross-section of the patterned channels (a right, 10×) with different solid contents and perfused peptide mixture concentrations after thorough washing of the non- specifically bound molecules; (b) Gradient intensity profile of the patterned channels; (c) Penetration depth measured as the full width at half maximum

Figure 4.13 Microscopic images (10×) of NIH-3T3 cells cultured in the channels of

5% (a), 10% (b) and 15% (c) hydrogels conjugated with CRGDSGK peptide (1% of

Maleimide groups) and on the well plate (d) at day 7

Figure 4.14 Microscopic images (4×) of NIH-3T3 cells cultured in the channels of

10% hydrogel conjugated with 0, 0.2 mM and 2 mM peptide

Figure 4.15 Confocal laser scanning microscopy of NIH-3T3 cells stained with

Alexa Fluor 488 phalloidin and DAPI on day 7 and 14 in 5 groups of hydrogels.

Pictures show maximum intensity projection of confocal frames representing a height of 600 µm of the channel. Blue: nuclei; red: actin cytoskeleton

Figure 4.16 Cell densities on the channels as determined from DAPI counting

Publication and Conference Presentation

Peer-reviewed journal article

Shuang Wang, Paul D. Dalton*, Tim R. Dargaville*, Spatial Patterning of Hydrogels via 3D Covalent Transfer Stamping from a Fugitive Ink, Macromol. Rapid

Commun., 2018, 39, 1700564/1–5

Conference Oral Presentation

16/02/2016: Nanotechnology and Molecular Science HDR Symposium, Patterning

Hierarchical Hydrogels Based on Poly(2-oxazoline) via 3D Sacrificial Moulding,

Brisbane, Australia.

19/04/2017: Australasian Society of Biomaterials and Tissue Engineering (ASBTE)

Conference, Spatial Patterning of Poly(ethylene glycol) Hydrogels via 3D Covalent

Transfer Stamping from a Sacrificial Mould, Canberra, Australia.

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List of Abbreviations

3D-CTS: 3-dimensional covalent transfer stamping

ABD: albumin binding domain

BSA: bovine serum albumin

CLSM: confocal laser scanning microscope

CM: cardiomyocyte

CNTF: ciliary neurotrophic factor

CPM: 7-diethylamino-3-(4’maleimidylphenyl)-4-methylcoumarin

DAPI: 4’6-diamidino-2-phenylindole

DMEM: Dulbecco’s modified eagle medium

EC: endothelial cell

ECM: extracellular matrix

EGF: epidermal growth factor

FBR: foreign body reaction

FCS: fetal calf serum

FDM: fused deposition modelling

FGF2: fibroblast growth factor-2

FI: Fluorescence intensity

FITC: fluorescein isothiocyanate

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GelMA: gelatin methacrylate

HA: hyaluronic acid

HSA: human serum albumin

HUVEC : human umbilical vein endothelial cell

IPN: interpenetrating polymer network iRP: indirect rapid prototyping

MEW: melt electrospinning writing

MMP: matrix metalloproteinase

MSC: mesenchymal stem cell

PCL: poly(ε-caprolactone)

PDGF-BB: platelet derived growth factor-B

PDMS: polydimethylsiloxane

PDLGA: poly(D-lactide-co-glycolide)

PEG: poly(ethylene glycol)

PEGDA: PEG-diacrylate

PEG-4SH: 4-arm PEG thiol

PEG-4VS: 4-arm PEG vinyl sulfone

PEG-PLLA-DA: poly(ethylene glycol)-co(L-lactide) diacrylate

PVA: poly(vinyl alcohol)

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RGD: Arginine-glycine-aspartic acid

SEM: scanning electron microscopy

SHH: sonic hedgehog

SPA: single-photon absorption

SPAAC: strain promoted azide-alkyne cycloadditions

TPA: two-photon absorption

UV: ultraviolet

VEGF-A: vascular endothelial growth factor A

Statement of Original Authorship

The work contained in this thesis has not been previously submitted to meet

requirements for an award at this or any other higher education institution. To the

best of my knowledge and belief, the thesis contains no materials previously

published or written by another person except where due reference is made.

Signature

QUT Verified Signature

Date

September 2018

xvi

Acknowledgements

I would like to thank my principle supervisor A/Prof Tim Dargaville, for your ongoing support and encouragement throughout the years. Thank you for providing me with the life changing opportunity to study in your group. I’m deeply grateful for your professional instructions that make my journey much easier and fulfilled. I thank my co-supervisors Dr. Danica Hickey and Prof Damien Harkin for the support and illuminative suggestions. I thank Prof Paul Dalton for your insightful advices and encouragement in my project. I would also like to extend my gratitude to Dr

Aurelien Forget, for your encouragement and useful discussions from your professional biology aspects.

I sincerely appreciate the support and friendship from my fellow QUT students, including Jodie Haigh, Emily Grundgeiger, Rebecca McMaster, Eleonore Bolle,

Samantha Catt, Deanna Nicdao, Nick Huettner, David van der Heide, Mariah

Sarwat, Camille Fromageot and Christoph Meinert. Furthermore, I would like to thank the QUT staff members including Dr Sanjleena Singh, Dr Chris East, Felicity

Lawrence, Parvathi Haridas, and Andi, for their generous help on the training and use of the instruments.

Last but not least, I would like to thank my families. To my husband Peng Wang, your faithful love and enduring support make my life filled with hope and joy. Thank you to my parents for their endless love and support throughout my life.

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Chapter 1 Introduction

1.1 Background

Extracellular matrix (ECM) cues are important regulators of cell function and fate, including migration, adhesion, differentiation and communication. To mimic the functions of ECM, signalling components such as proteoglycans and proteins have been delivered onto biomaterial scaffolds as in vitro culture models. For example, vascular endothelial growth factor A (VEGF-A) controls blood vessel formation and function by directing endothelial cells (ECs) growth[1]. Fibronectin peptide fragment arginine-glycine-aspartic acid (RGD) enhances cellular attachment, based on its interaction with the integrins presented on the cell surface[2]. Being able to tailor the biochemical environment of a biomaterial is essential for guiding cellular behaviour, elucidating cell-cell and cell-substrate interaction and ultimately engineering complex tissues.

Native ECM gives structural support to the cells and contains conductive pores to facilitate nutrient/waste flow and tissue growth. It is important to introduce porosity on the scale of proper cellular processes within the in vitro scaffold[3]. The viability of cells in a scaffold requires sufficient delivery of oxygen and nutrients from the surrounding growth media and removal of the metabolic products from the cells through the pores. In addition, in vivo experiments showed that polymer with interconnected pores was able to integrate with the surrounding tissue with little or no foreign body reaction (FBR), while the same polymer solid implant triggered classic FBR[4]. The ability to control scaffold inner architectures ensures a certainty

on physical properties of the substrates to improve cell viability and study cell behaviour.

Hydrogels are hydrophilic polymeric networks capable of imbibing high weight fractions of water, giving them similar physical properties to natural ECM.

Hydrogel-based systems have emerged as powerful tools to mimic relevant static and temporal environment of the native ECM, serving not only as a structural support for cell growth, but also a biological niche where cells can respond to a number of physical and chemical cues. This requirement inspires innovations of patterning hydrogels with physically hierarchical structures (macro and micro features) and chemically regulating cues. With the ability of control at varied levels, patterned hydrogels have been used as in vitro culture models to understand how individual or a combination of extracellular cues influence cell behaviour. Among synthetic hydrogels, poly(ethylene glycol) (PEG) – based hydrogels have been established to take the lead for tissue engineering application. PEG exhibits biocompatibility and resistance to protein adsorption and cell adhesion, thus providing non-toxic substrates with little immunogenicity and a stealthy matrix on which desirable bio- functionalities can be built. The intrinsic structure of PEG allows end functionalization that endows a number of crosslinking mechanisms and molecule modifications.

Additive manufacturing (or 3D printing) is any of various processes used to make three-dimensional structures from electronic data sources. Poly(ε-caprolactone)

(PCL) has been extensively applied for 3D printing[5]. An interesting application for these structures is creating a negative copy of the part within a hydrogel matrix using sacrificial technique. Sacrificial template is encapsulated in the hydrogel precursor and then removed after the gel cures, leaving hollow structures that have

complementary shapes and sizes of the template. In recent decades, both fused deposition modelling (FDM) printing and melt electrospinning writing (MEW) are utilized to make sacrificial templates, with the MEW technique producing fibres of

1–2 orders of magnitude smaller in diameter than those possible by FDM. The presence of ordered, multi scalable channels within the hydrogels grants the soft materials interconnecting, geometrically hierarchical features, which is important for mimicking the transport and mechanical functions that exist in native ECM.

Although much interest has arisen to manipulate the presentation of biochemical cues in hydrogels, approaches for producing selective patterned hydrogels have been limited to the use of sophisticated chemistries and access to confocal lasers. Simpler and more translatable strategies are needed to create patterned hydrogels with desirable chemical and physical features that can be served as cell interactive substrates.

1.2 Thesis Overview

This study combines the concepts of 3D sacrificial moulding and reactive molecule transferring to produce hierarchical hydrogels with patterned channels within synthetic hydrogels. Two approaches were developed and investigated to selectively pattern hydrophobic and hydrophilic molecules respectively onto the channel lumen.

Chapter 2 provides an overview of the emerging cell-responsive hydrogel platforms with biochemical and geometrical patterns that mimic the properties of native ECM.

The chemistries used to crosslink and functionalize the hydrogel, the strategies developed to create geometric features, and the functions of the hydrogel models with biophysical and biochemical cues, are discussed.

Chapter 3 describes the development of a new method, coined ‘3D covalent transfer stamping (3D-CTS)’, for making 3D patterned hydrogels using a sacrificial template containing model molecules containing reactive groups which can be bound to the hydrogel during crosslinking. This method allows the creation of hydrogels with an inner architecture with hydrophobic molecules selectively patterned only onto the channel walls.

Chapter 4 describes a novel method for the fabrication of peptide coated channels within PEG hydrogels. Using sacrificial templates fabricated by additive manufacturing techniques, namely FDM and MEW, channels with diameters from tens to hundreds of micrometres were produced. The reactive peptide molecules perfused in the channels rapidly react with the residual functional groups at the hydrogel interface. The process was demonstrated to be cost efficient due to a reduction in peptide amount and experiment duration. Proof of principle experiments with the mouse fibroblast cell line (NIH-3T3 cells) showed preferred cell attachment and proliferation on channels coated with higher concentrations of peptide. The novel method developed here has the potential to be used for engineering of various in vitro culture models.

Chapter 5 summarizes the advantages and limitations of the two approaches for hydrogel patterning. Alternative approaches with improved biocompatibility were described, and future directions for possible applications of these synthetic models were discussed.

Chapter 2 Literature Review

2.1 Introduction

Hydrogels are natural or synthetic crosslinked, hydrophilic polymeric networks that swell extensively in water. With a variety of fabrication techniques, hydrogels can be cast to any shape and size[6], with usual forms such as nanofibers[7], films[8, 9], microparticles[10], and circular disks[11] produced. Hydrogels feature tunable physical, chemical, and biological properties that determine the mass transport efficiency and cell-matrix interactions, thus having potential use in biomedical applications. The pioneering work of the synthesis of poly(2- hydroxyethylmethacrylate) [(poly(HEMA)] hydrogels in the 1960s was soon patented for soft contact lenses[12], followed by a variety of gel systems with different network-forming mechanisms being investigated and documented for biomedical applications[13]. The building blocks of hydrogels are based on a range of polymer compositions derived from natural sources (collagen, gelatin, fibrin, hyaluronic acid, agarose, alginate, heparin and chitosan) to synthetic polymers

(polyacrylamide, poly(ethylene glycol), poly(vinyl alcohol), poly(lactic-co-glycolic acid), poly(2-oxazoline), and polypeptides)[14-16], and the combination of the two categories[17]. These works built the essential toolbox of modular building blocks from where mechanical and chemical versatility from hydrogels are available.

In many typical applications[18] such as contact lenses, biological adhesives[19], would dressings[20], and soft tissue fillers[21], hydrogels are relatively homogeneous and inert scaffolds that passively function as a waterish supporting structure. In these applications hydrogels are required to have proper mechanical strength, hydrophilicity, permeability, biocompatibility, and degradability[14].

However, for the emerging applications in drug delivery, regenerative medicine, cell research and cell therapy, and tissue engineering, more requirements regarding physical and biological attributes of hydrogels need to be satisfied beyond the above basic criteria. When serving as a habitat for cells, hydrogel scaffolds mimicking extracellular matrix (ECM) are used to facilitate cell attachment, migration, and proliferation (Figure 2.1). Tunable heterogeneity of the hydrogels enables the understanding of cell responses to their biomimetic 3D environment for the purpose of better facilitating complex tissue formation[22, 23]. Inspired by the natural design of ECM that gives structural and biochemical support to the cells, research focus has moved towards rational incorporation of extracellular cues[24] and dynamic design with spatial and temporal complexity via various patterning approaches[25, 26]. For example, light-based coupling/ablation is a straightforward and biocompatible method that enables real-time spatiotemporal change of the microenvironment; bioprinting is capable of direct deposition of different materials, which has been comprehensively reviewed elsewhere[23, 27]; and subtractive moulding is an indirect approach that produce negative copies of the template[28]. The combined present-day technologies open up new avenues in developing biomaterials with functionalities closer to those offered by organisms, leading up to the successful development and fabrication of well-defined hydrogels with intricate structures that significantly advance the promising technology of cell research and tissue engineering.

Figure 2.1 Comparison between natural ECM and synthetic substrates.

Reproduced from [29].

2.2 Patterning Biochemical Signals

2.2.1 The Need for Modifying Biosignals

Both natural and synthetic hydrogels have been explored to study cell-interactive properties including cell adhesion, migration, and differentiation. Naturally derived hydrogels possess inherent cell-responsive characteristics and biodegradability.

Gelatin methacrylate (GelMA), for example, is a photo-polymerizable hydrogel that retains cell binding motifs such as RGD and matrix metalloproteinase (MMP)- sensitive degradation sites. When GelMA was micropatterned on PEG surfaces, cells adhered, proliferated, and migrated on the surface of the cell-responsive micropatterns[30]. However, naturally derived hydrogels have limitations due to batch to batch variation, lack of mechanical strength, and uncontrollable structure and degradation. Moreover, because of the large number of variables in natural hydrogels, it is difficult to adequately control biological and therefore elucidate specific cell responses in a complex 3D environment. On the contrary, the structure

and degradation of synthetic hydrogels are controllable and provide batch to batch consistency[31]. Moreover, synthetic hydrogels can be designed to contain specific cell adhesion ligands and degrading motifs[32]. Thus, the cell-substrate environment can be readily adapted to address a particular hypothesis. The ability to control the dose and distribution of the biosignals facilitates a better understanding of their impact and affords greater control in experimentation design. RGD improves the viability of cardiomyocytes in a dose dependent manner, and this effect is particularly pronounced in degrading hydrogel[33]. A study showed NIH-3T3 cells not only adhere on covalently RGD gradient in PEG hydrogels, but also stretched, aligned, and preferentially migrated along the direction of the gradient, indicating the influence of the concentration and distribution of the biosignals on cell activities[34].

Overall, the fabrication of desired hydrogel matrices with specific bioactivity has been popular in these years to systematically study cell behaviour.

2.2.2 Chemistry Strategies for Gelation and Patterning

Hydrogel precursors crosslink upon initiating chemical stimuli to certain functional groups or exposing to physical changes. Photo-curable gels are prevailing due to the ease of forming a gel and post-patterning the gel upon exposure to ultraviolet (UV) light utilizing the unreacted groups. Materials modified with end- or side photo- crosslinkable functional groups such as (meth) acrylates have been extensively investigated, both of natural and synthetic sources[35]. Cytocompatible photo- initiators include type I Irgacure 2959 (I-2959) and lithium arylphosphinate (LAP), and type II eosin-Y[36]. I-2959 possesses low molar absorptivity at 365 nm and a near zero molar absorptivity at the range of visible light wavelength range, thus exhibiting lower efficiency. Recently, visible-light photo-initiators have been increasingly used in cell-laden hydrogels with superior cytocompatibility and high

efficiency[37]. Alternatively, Michael addition crosslinking of dextran[38, 39] or

PEG macromers[40, 41] containing acrylates, maleimides, vinyl sulfone end groups with nucleophilic thiols undergo rapid, quantitative reaction under physiological conditions without the aid of light or initiator. The reaction forms hydrogels that contain excess functional groups for efficient chemical functionalization even under off-stoichiometric conditions[42]. Relevant chemistries used in hydrogel synthesis and patterning have been discussed in literature [43] and were summarized in Table

2.1. Especially, cell-encapsulated hydrogel matrixes require non-toxic nature from reactions such as thiol-based and strain-induced coupling reactions.

Table 2.1 Chemical reactions for hydrogel synthesis and patterning

Reaction type Functional groups Reaction conditions Gelation Biocom- By-product

time patibility*

Thiol-Michael Thiol & electron- Alkaline buffer Seconds +/+ Potential reaction[39, 44] deficient ene/yne to minutes oligomer

Thiol-ene/yne radical Thiol & ene/yne Photo-initiator Seconds +/+ Potential reaction[10, 45] &light to minutes oligomer

Azide-alkyne Azide & strained Ambient condition ca. 2 min +/+ None cycloaddtion[46] cyclooctyne

Diels-Alder Furan & maleimide Heat ca. days –/+ None reaction[47, 48]

Alkene-tetrazine Tetrazine & strained Ambient condition ca. 5 min +/+ N2 inverse electron- cycloalkene demand Diels-Alder addition[49, 50]

Oxime click Primary amine & Acid & catalyst ca. 10 min –/+ None reaction[51, 52] aldehyde

Tetrazole-alkene photo Tetrazole & alkene UV 50 s to 5 +/+ N2 reaction[53] min * During and after gelation.

The various ways that biomolecules are incorporated to polymeric networks include utilizing covalent attachment through available functional groups[46] and high- affinity immobilization through specific binding[54]. Biomolecules with fluorophore are frequently used to pattern hydrogels for the ease of visualization, making the patterning quantifiable. Functional groups for binding are either residuals directly from hydrogel formation[55], or being revealed by specific triggers[56]. For example, a copolymer containing furan-protected maleimide was employed to form hydrogel using a radical thiol-ene reaction, followed by thermal activation of the masked maleimide groups to undergo nucleophilic thiol-Michael addition reaction for micropatterning[45]. Reaction orthogonality by introducing various reactive groups enables independent control over different reaction mechanisms in a single system. As shown in Figure 2.2, orthogonal reactions for gelation and patterning allow manipulating spatial and temporal features of the hydrogel environment and avoid confounding associations from each other or the culture media[57, 58].

Figure 2.2 Orthogonal chemistries for the creation of diverse biochemical patterns.

Reproduced from [58].

2.2.3 Surface Patterning

Two dimensional (2D) culture systems exclude the constraint effects on the cells in a

3D niche and have been used to identify the specific functionality of each biochemical component in a simple way. The modification of hydrogel surfaces regarding chemistry and topography have attracted much research interest for the simplicity of the set-up and the need to break down complicated systems into manageable ones. Functional surfaces immobilized with specific biomolecules in a defined region were fabricated to tailor cell microenvironment and study the fundamental cellular response to environmental cues. For example, fibroblasts spread and grew along the direction of the patterned fibronectin line, cells ‘bridging’ occurred between the line patterns and increased as the pattern spacing decreased[59]. Surface patterning strategies are based on contact technique of soft lithography, and non-contact technique of photolithography.

2.2.3.1 Soft lithography

The microcontact printing (μCP) technique was developed by Whitesides and coworkers for printing alkanethiols on the surface of gold using a polydimethylsiloxane (PDMS) stamp[60]. It was later derived a set of variants which collectively terms “soft lithography”[61]. Some studies extended μCP to immobilize biomolecules on soft hydrogels using a micro-structured stamp through covalent bond[62], specific affinity[63], or non-specific interactions[64-66] with features on the order of several micrometers. For example, a study conducted by Burnham et al.[62] showed that iodoacetyl biotin and streptavidin could be immobilized on the polyacrylamide hydrogel through reacting with sulfhydryl groups created by treatment with a reducing agent via μCP, generating well-resolved feature with a width of 2 μm. In another paper, a gel surface functionalized with streptavidin was

prepared by photo-polymerizing acrylamide in the presence of streptavidin- acrylamide, enabling the binding of biotinylated human plasma fibronectin and mouse laminin via μCP[63].

Due to the soft and tacky nature of the target hydrogels, conventional μCP may cause deformation of the substrates and distortion of the patterns, thus making protein patterning not as straightforward as on solid substrates. To overcome the problem, researchers have adjusted the physiochemical properties of either stamp or substrates, which are demonstrated by the following examples that create surface- bound patterns through non-specific adsorption. In the first example, a simple trans- print method using a water-soluble polyvinyl alcohol (PVA) film was developed to pattern fibronectin and collagen on polyacrylamide hydrogel[64]. The pattern was firstly transferred from PDMS stamp to a thin PVA film, which used as a trans-print media by placing in conformal contact with the hydrogel substrate (Figure 2.3a). The reproducible and high-resolution pattern was proven to be beneficial to the attachment and alignment of human mesenchymal stem cells along the pattern direction. The second strategy realized protein (fibronectin and fibrinogen) micropattern transferred and physically embedded onto polyacrylamide hydrogels using a glass templates bearing patterned hydrophobic poly(N-isopropyl acrylamide)

(PNIPAM) brushes[65]. The main advantage of this method is its excellent reusability due to the robust covalent bonded polymer brush on the glass template

(Figure 2.3b). In another example, Castaño et al.[66] employed PDMS stamps featuring lines to pattern several proteins including streptavidin, laminin, and fibronectin on freeze-dried MatrigelTM and gelatin substrates. After reconstitution and during cell culture, the micropattern remained and guided NIH-3T3 cellular alignment, even though the proteins were not chemically patterned but rather

immobilized by physical adsorption that diffused 8 μm into the hydrogel.

Furthermore, beating foci were observed earlier when culturing human embryonic stem cells on fibronectin-patterned hydrogel than on non-patterned substrates.

Figure 2.3 (a) Conventional μCP process (up) and trans-print process using a PVA film (down). (b) Polyacrylamide patterning from the patterned PNIPAM brushes grafted onto a glass template. Reproduced from [64] and [65].

Although μCP techniques are powerful tools for surface patterning, there are inherent limitations including the uncontrollable amount of the transferred materials and alignment steps required for patterning multiple biomolecules. Zhang et al.[67] combined microfluidic patterning with μCP, demonstrating the capability of patterning multiple biomolecular ink solutions and generating gradients using multiple channel design. Specifically, microfluidic channels were formed by reversibly sealed to PDMS stamp using a modified polycarbonate track-etched PCTE membrane, where biotinylated biomolecule solution diffused through to the polyacrylamide gel bearing streptavidin. The use of PCTE membrane prevented fluid leakage during patterning and flow disruption after patterning when gel substrate

was removed, offering good reusability and reproducibility of the pattern transfer.

However, the resolution of the patterns was in the order of hundreds of micrometres.

Despite these efforts, limitation of soft lithography is obvious: when hydrogels possess high aspect ratios (height to width ratio), the quality of the features reduces, and it is difficult to retain the integrity of the hydrogel structure during the removal of the master.

2.2.3.2 Photolithography

Photopatterning in the presence of a photomask is a widely used technique that produces micro-/nanopatterned immobilization of biomolecules with high fidelity[52, 53]. Although the topographic height can reach hundreds of micrometres, which is related to the transparency of the hydrogel, the activity of the functional groups for photolithography, irradiation time and light source, the regional immobilization is regarded as surface patterning.

With hydrogel networks formed from active components, post functionalization of bioactive factors to the residual functional groups via radical thiolation or inverse electron-demand Diels−Alder addition is a versatile, straightforward way to a spatially controlled surface patterning[49, 52]. These reactions often proceed under physiological conditions, enabling the encapsulation of cells. Utilizing multiple steps of photomask patterning, different biomolecules can be separately immobilized on the hydrogel surface in a shape-controlled manner. The group of West[68] showed that multiple peptides could be sequentially immobilized on the surface of a photoactive PEG-diacrylate (PEGDA) hydrogel in distinct patterns defined by photolithographic masks and UV light (Figure 2.4a). Human dermal fibroblasts

(HDFs) were found to adhere to and spread on the RGDS patterned regions, but not

to unpatterned regions or regions patterned with the endothelial cell-specific adhesion peptide REDV (Figure 2.4b). Topographical structures with feature heights were also generated by selective curing the PEGDA precursor layered onto the hydrogel surface using UV light for multiple times, making the first step towards creating intricate patterning in 3D (Figure 2.4c). The minimum feature size obtained was relatively large, reaching 40 μm. This feature size can be reduced by using a higher resolution printer in mask fabrication and a collimated light source.

Figure 2.4 Fluorescence image of a PEGDA hydrogel patterned with ACRL-PEG-

2.2.4 Three-dimensional Patterning

Building 3D biomaterials is more difficult than 2D surface patterning, however more appealing due to the resemblance of natural tissue in vivo. For the aim of cell guidance in 3D, it is important to deliver and localize biomolecules into hydrogels in a site-controlled manner. Several microfabrication methods such as photolithography, microfluidics[69] and 3D printing have been exploited, among which laser based methodologies being the most commonly explored[24]. The use of advanced confocal laser light controlled by software opened up new avenues to

patterning biomolecules such as cell adhesive peptides[55] and growth factors[70] into hydrogels in a spatiotemporal controllable manner by manipulating the focal point of the laser light[71]. These scaffolds are being advanced for tissue engineering and also being used as tools to answer more fundamental questions. However, in many cases, integrated strategies were utilized to realize the establishment of a complex system.

For optically transparent, photoactive hydrogel matrices, photolithography allows for rapid patterning in three dimensions to create complex bioactive features. The functional biomolecules are allowed to diffuse throughout the entire hydrogel before photo immobilizing them spatially. Using sequential click reactions, one for gelation and another for post modification (Figure 2.5a), the conjugation of biomolecules at specific locations within the gel was achieved by controlling the focal point of the laser light in three dimensions using a confocal microscope [55], providing tailorable synthetic microenvironment for cell studies.

Figure 2.5 Thiol-ene reaction for post network formation (a) to introduce three different fluorescently labelled peptide sequences (b) and micrometre-scale spatial patterning (c) within PEG gel. Reproduced from [55].

The feasibility of 3D biochemical and biomechanical patterning PEGDA adapting single-photon absorption (SPA) and two-photon absorption (TPA) was examined

and compared by the group of West[72]. While pattern complexity using SPA limits to topographical microstructures on a surface and axially regular features with the utilization of a mask, TPA photolithography is capable of focusing on a confined volume, therefore creating freeform 3D patterns without using a physical mask. A combining approach of photo-conjugation and orthogonal physical binding was reported by Shoichet and coworkers[73]. Using two-photon photochemistry, maleimide–barnase and maleimide–streptavidin were sequentially immobilized before stem-cell differentiation factors sonic hedgehog (SHH) and ciliary neurotrophic factor (CNTF) independently bound to barnase and streptavidin, respectively. This method facilitates spatial immobilization of multiple growth factors and thereby specifically used for guiding stem/progenitor cell differentiation.

Other works utilized the photocleavage of labile moieties, thus activating immobilization sites that enable conjugation of desired biomolecules and allowing spatiotemporal control over the 3D distribution of the bioactive molecules. The group of Shoichet pioneered this 3D-patterning approach in 2004, using chemically patterned cylindrical volumes in agarose gel fabricated using a UV laser, as shown in

Figure 2.6[56]. Briefly, photolabile 2-nitrobenzyl groups were cleaved by a laser beam with a diameter of ~150 μm, leaving free sulphydryl groups within the irradiated volumes (ca. 150 μm in diameter and 2–3 mm in depth) that were readily modified by maleimide-terminated oligopeptides via Michael-type addition. Through site-specific cleavage and modification, oligopeptide channels were fabricated and verified to be effective in cellular guidance and neurite outgrowth[74]. Later in 2008, the group used a multiphoton laser scanning lithography and a 6-bromo-7- hydroxycoumarin (Bhr)-based protecting group that showed a higher efficiency of laser deprotection, creating multiple immobilized concentration gradients of model

thiols with high spatial resolution within the agarose hydrogels[75]. The same photochemistry method was used to conjugate gradients of human epidermal growth factor (EGF) into a hybrid hydrogel consisting of the natural proteoglycan hyaluronic acid (HA) and the synthetic PEG[76], created an EGF concentration gradient of 25 to 250 nM over 150-μm depth. These nitrobenzyl and coumarin derivatives were utilized to orthogonally and sequentially release multiple proteins on demand by wavelength-selective photocleavage, creating a dynamic presence of biomolecules[77]. The technique was extended by taking advantage of the femtomolar physical binding partners, human serum albumin (HSA) and albumin binding domain (ABD), following a successive chemical bonding of fibroblast growth factor-2 (FGF2) and HSA to the photo-exposed agarose thiols in agarose.

The FGF2-ABD fusion protein formed a stable complex with the immobilized

HSA[78].

Figure 2.6 Generating adhesive biochemical channels in agarose hydrogel matrices by localization of photo-deprotection reaction and photo-coupling. Reproduced from

[56].

Gradient hydrogels have been developed to mimic the spatiotemporal differences of multiple cues in tissue. Gradient hydrogels are produced by physical diffusion of

biomolecules before UV immobilization[79] and microfluidic mixing[80]. These methods, however, are restricted to a single gradient distribution without much manipulation. Wei et al.[81] developed a modular gradient hydrogel constructs utilizing a self-healing network via Schiff base reaction. Encapsulated sarcoma cells respond to the gradient cues of RGD peptides and MMP-sensitive crosslinkers in the hydrogel, enabling guidance of cellular responses to their microenvironment. Tabata and Lutolf[82] reported a perfusion-free microchip concept for the in vitro perturbation of stem cell fate and long range cell-cell communication. Stem cells encapsulated in the hydrogel compartment respond to diffusion-driven leukemia inhibitory factor (LIF) gradients created at the flank open reservoir. This technique afforded a considerable reduction in biomolecule consumption (tens of microliters) to establish stable gradients compared with traditional fluidic-based system

(typically required hundreds of microliters to a few millilitres)[83].

2.3 Patterning Complex Geometry

2.3.1 The Need and Challenges for Structural Manipulation

The lack of perfused vascular networks within hydrogels may cause the formation of a necrotic core because diffusion-driven mass transportation via hydrogel meshes alone is insufficient[84]. In hydrogels with channels, cell viability and differentiation is greater than in bulk hydrogels, presumably due to improved mass transport[85].

Channel geometry exerts an impact on cell adhesion for cell sorting in microfluidic devices – channels with curved turns provide more uniform and predictable cell adhesion comparable to that in straight channels, while channels with sharp turns were observed to lack uniform adhesion in turn regions[86].

The challenge of creating interconnectivity in soft hydrogels while maintaining the structural integrity has been resolved using traditional strategies[87]. These methods include gas forming[88], freeze-drying[89, 90], particle leaching[91-93], electrospinning[94], and cryogelation[95], which produce porosity with sizes ranging from several to hundreds of micrometres. Indeed, these porous hydrogels possess advantages such as a high surface area for mass transport, the ease of processing and the compatibility with a variety of polymers. However, due to the irregular morphology and batch-to-batch variations, precise spatial control of scaffold interior architectural details such as scale, geometry and interconnectivity is hard to achieve.

In the very recent years, a growing number of studies using rational tools have been reported for producing scalable pores with precisely defined dimensions.

2.3.2 Soft Lithography

Surface topography can be produced by micromoulding techniques. For example,

Annabi and colleges[96] fabricated hydrogel films with grooves by photo- crosslinking methacrylated tropoelastin between a micropatterned PDMS mould and a glass slide. These elastic human protein-based gel substrates with high-resolution patterns (20×20 μm and 50×50 μm channel width and spacing) were found to promote cardiomyocyte (CM) attachment, spreading, and elongation. Based on the surface topographic patterning, a trough-containing slab was bonded to a second slab to make closed channels. Using this method, Zheng et al. [97] produced microfluidic vascular network in type I collagen that allowed seeding of endothelial cells (ECs), defining a network of endothelialized channels that showed barrier function. This in vitro microvessel model was then used to investigate the interactions with perivascular cells and blood components.

Using microreplication method, He et al.[98] fabricated a PDMS mould with negative pattern of leaf venation. The open vessels on agarose gels made from the

PDMS mould had diameters from 1 mm to 30 μm. Endothelial cells in the vessels maintained high viability, proliferation and spreading over the 3 days. On the basis of soft lithography, Jiang et al.[99] used a structured collagen fibril hydrogel filled with liquid mould to bond another layer of gel to construct vascular networks through fibrillogenesis across the hydrogel interface. Multilayer microchannels were built by bonding more layers that mimic mass transfer-based physiological functions[100]. However, alignment issue between the layers arising from multiple iterations hinders the broad application of this approach. Besides, unlike native branching architecture with cylindrical channels, fluidic networks via soft lithography typically yield uniform channel dimensions with right-angle cross- sections that cause heterogeneous wall shear stresses, which subsequently influence fluid mechanical cues and cell patterns[86].

2.3.3 Template Moulding

Template moulding uses customer-designed scaffolds fabricated with a material of choice as the sacrificial element. After the removal of the templates, a negative structure of the template generates within the gels. Indicated as “indirect rapid prototyping (iRP)” by Dubruel[27], template moulding using a wide range of materials has been extensively studied in recent years. The key points for a successful template moulding process are the build-up and removal of the sacrificial element–both primarily rest with the physical properties of the sacrificial material.

Straight tubular channels with a diameter of ~100 μm were yielded in collagen gel using removable steel needles. On the surface of these open channels, endothelial cells could form a confluent layer that exhibited high barrier function and quick

inflammatory stimuli[101]. These simple needle-moulded channels were adopted by

Chen and coworkers for the computational and experimental modelling of shear stress threshold on angiogenic sprouting[102]. Another work out of Harvard/MIT used bioprinted agarose fibres as templates. After the GelMA or PEGDA precursor cured under UV light, the fibres could be easily aspirated with a gentle vacuum, leaving microchannels with diameters ranging from 150 to 750 μm[85]. Golden and

Tien fabricated planar and stacks of planar microfluidic networks in collagen gel by melting micromoulded gelatin template at 37 and flushing away the liquid gelatin, producing channels as narrow as 6 μm[103]. Using a similar principle, perfused vascular channels were fabricated by printing collagen precursor around solidified gelatin pattern and washing out the liquefied gelatin at elevated temperature[104], or using moulded alginate lattice that is reversibly ionic crosslinking and dissolvable[105], or by scarifying thermos-responsive solvent-spun PNIPAM microfibers[106]. While these methods are convenient and solvent-free, controllable complex stereo structures are not possible.

This limitation can be addressed by using 3D printed templates as the formed channels replicate the features of the sacrificial template. Melt electrospinning PCL templating was demonstrated for fabrication of 3D interconnected channels with a size of ca. 20 μm in poly(2-oxazoline) hydrogels, where fine detail of the PCL template was reproduced[107]. However, the use of acetone to dissolve the template might limit the applications where cells need to be in situ encapsulated. The group of

Lewis[108] used a thermally reversible gelation of Pluronic F127 that experienced a gel-to-fluid transition at 4 , such that the printed structure became liquid and could be easily removed to form 3D interconnected microchannels in GelMA. A suspension of the human umbilical vein endothelial cell line (HUVEC) injected into

the channels was highly viable and assembled to a confluent layer. However, the toxicity of Pluronic F127 and the low temperature required for the removal of it potentially damage encapsulated cells. Miller et al.[84] developed an elegant method using 3D-printed carbohydrate glass lattice as the sacrificial element. The hydrophilic template was protectively coated and removed with culture media, leaving open channels in a cell-laden matrix (Figure 2.7a and b). This simple approach was demonstrated to be compatible with a variety of cell types and different gelation mechanisms (Figure 2.7c). Later the printing parameter of carbohydrate glass was optimized to minimize drooping and pattern filaments in free space for building a microfluidic device[109].

Figure 2.7 (a) Schematic overview showing 3D printed carbohydrate-glass lattice served as sacrificial element to yield vascular architecture in a monolithic construct;

(b) Microscopic images showing an open perfusable channel in the fibrin gel upon

dissolving the glass filament; (c) Confocal microscopy showing matrix with varied crosslinking mechanisms (red beads), encapsulated cells (10T1/2, green), and the vascular lumen (blue beads). Reproduced from [84].

Compared to small carbohydrate molecules, water-dissolvable polymers are more of interest for the reduced osmotic pressure, yet relevant studies are rather rare except for the use of water-soluble poly(vinyl alcohol) (PVA). By casting silicone elastomer around a 3D printed sacrificial water-soluble PVA mould, a 3D silicone scaffold replica was formed and subsequently impregnated with hydrogel. The interpenetrating polymer network (IPN) scaffold was found to support growth of human mesenchymal stem cells with high cell viability, metabolic activity and well spread morphology. In addition, sustained release of doxycycline with biological activity from the gel network was confirmed, demonstrating its potential in directing stem cell differentiation[110]. Another example used paraffin as coating of PVA to prevent deformation of the channels[111]. For the same purpose, plus allowing dissolved carbohydrates to be flowed out without passing through the bulk of the gel, a thin layer of poly(D-lactide-co-glycolide) (PDLGA) was coated on the carbohydrate-glass lattice in Miller’s work[84]. Neither study mentioned the removal of the coating nor to what extent would the coating inhibit diffusive transport between the channels and the bulk gel.

2.3.4 Site-specific Photoablation

To make photo-cleavable hydrogels, photodegradable groups were incorporated into the cross-linked hydrophilic polymer network. Upon light irradiation, crosslinking density of hydrogels can be tuned by cleavage of photodegradable moieties or crosslinkers, bringing about the macroscopic changes such as Young’s modulus, swelling degree, and water diffusion[70, 112]. Similar to the effects of photo-

conjugation, laser-induced photocleavage is confined to a particular volume where a high level of fidelity of erosion is afforded, creating pre-defined microchannels and shapes with micrometre-scale resolution[113] and enabling directional growth guidance of neurites[114]. Recently, Deforest et al.[115] developed microchannel networks in photosensitive gels with sizes spanning those of native human vasculature using multiphoton lithography-assisted photo-scission. Endothelium coated channels were readily generated in the presence of encapsulated stromal cells.

2.3.5 Layer-by-layer Photopatterning

Photolithography can solidify certain volume of precursor material at a time. When the exposed volume is confined and patterned layer by layer, 3D scaffolds with fine microstructures can be fabricated. An early study[116] used stereolithography to photo-crosslink bioactive PEG-dimethacrylate solution to a movable platform, one layer at a time, producing internal branched channels with diameters of 1 mm. Over

87% of the in vitro encapsulated human dermal fibroblasts were found to be viable after 24 h following fabrication. Poly(ethylene glycol)-co(L-lactide) diacrylate

(PEG-PLLA-DA) was patterned within multilayer PEGDA hydrogels using photolithography. Upon exposure in high pH environment, PEG-PLLA-DA parts degraded within 3 days in bulk PEGDA hydrogel to form microchannels[117].

Another photopatterning approach used manually operated adjustment of spacers and masks, patterns of multiple cell types in macroporous hydrogel were fabricated using an apparatus with a height-adjustable chamber, where PEGDA prepolymer solution was injected and partially cured under UV through different masks in a layer-by- layer fashion. The placement of cells could be 3D controlled throughout a thick construct and cell viability in the multilayered hydrogel microstructures was demonstrated[118]. Using a similar apparatus refined by the same group, hepatocytes

were incorporated in PEG hydrogels with defined architectures as a culture model[119]. Compared with unpatterned hydrogels, the multilayered cellular hydrogels minimized barriers to mass transport, thus improved the viability of these highly metabolic cells. However, due to the repeated labour and the consequent problems of inefficiency and consistent fidelity, this approach is not widely used though it’s capable of control 3D architectures.

Recent layer-by-layer photopatterning turns to rely on projection stereolithography

(PSL), a rapid lithography technique that produces controllable geometries with fine resolution. Dynamic photomasks are formed by projected UV light from digital light processing device controlled by computer aided design (CAD), thus selectively curing a photo-reactive material using sequential projections (Figure 2.8a). Gauvin et al.[120] produced scaffolds with woodpile and hexagonal geometries using gelatin methacrylate (GelMA). The interconnected pores within the structures allowed uniform distribution and proliferation of the human umbilical vein endothelial cell line (HUVEC) to produce high cell density and confluency. Zhang et al.[121] reported well-defined microwells with varied sizes and architectures (Figure 2.8 b– e), that support cell adhesion and proliferation and the features significantly affected cell alignment (Figure 2.8f and g). To avoid the use of cell-damaging UV, a visible light-based PSL system was developed to fabricate PEG hydrogels with precise geometries and internal architectures[122]. The authors demonstrated efficient in situ encapsulation and long-term viability of human adipose-derived stem cells.

NIH-3T3 cell culture after 4 days in the microwells. Reproduced from [121].

2.4 Integrating Structural and Biochemical Features

Very recently, elaborate designed building blocks that allow independent manipulation of chemical and physical cues are becoming more emphasized[24, 58].

This trend is because native ECM is dynamic space where various specialized proteins and proteoglycans present beyond its critical role of structurally connecting cells to form tissues. Building complex scaffolds with biochemical and biophysical cues within synthetic polymers is not as straightforward as for generating a single feature. Thus, established patterning techniques have been combined to integrate biochemical and biophysical cues in a single scaffold.

2.4.1 Patterned Microwells and Microgrooves

Combining soft lithographic and micromoulding technique, a model of hydrogel matrix patterned with microgrooves was established. The integrated structural and biochemical patterns were used to study the aggregation and coalescence of

endothelial cells and co-culture configuration with fibroblasts[123]. A microwell platform with photo-patterned heparin was developed for controlled cellular spheroid formation and defined signal presentation to regulate tissue formation[124]. Lutolf et al.[125] assessed the haematopoietic stem cell behaviours in response to a series of tethered and secreted molecules on PEG gel microwells fabricated using multistep soft lithography process. They found the self-renewal of the cells were induced by specific proteins such as Wnt3a and neural cadherin and stated this approach could be applied to elucidating interfacial effects of any single/combinatorial regulators on any cell types on substrates with modular stiffness.

2.4.2 Patterned Channels

The group of Anseth[126] developed photo-cleavable hydrogels that directed cell migration along the channel length. Moreover, the photodegradable linkages were also exploited to modify the chemical environment by cleaving tethered, biologically active functionalities upon irradiation, creating a controlled release of ligand presentation environment. In another outstanding work published in 2011 by the same group[46], a hydrogel system that enables independent manipulation of chemical and physical cues was designed by orthogonal photo-conjugation and photocleavage reactions. Tetra-functionalized PEG and bisazide peptide formed hydrogel network through a strain promoted azide-alkyne cycloadditions (SPAAC) reaction. The network bears vinyl groups that are readily patterned with thiol- containing biomolecules within user-defined regions on exposure to visible light. It was proved that the patterned concentration of a fluorescent RGD peptide relied on the dose of visible light and photo-initiator (eosin Y) amount, and multiple types of biomolecules could be localized by repeating the patterning process. An independent photo-degradation using multiphoton laser light rendered precise cleavage of the

nitrobenzyl ether moieties from peptide linker, removing the peptide from selected regions and forming physical pores with defined dimension (Figure 2.9a).

Manipulating of the cell migration was demonstrated by NIH-3T3 cells encapsulated within the hydrogel, which were found to migrate only down the physical channel functionalized with RGD (Figure 2.9b and c). Using similar chemistries, reversible presentation of biomolecules can be spatiotemporally controlled and for the cellular microenvironment to be programmed on demand[127, 128].

patterned channels (c) Cell outgrowth controlled in three-dimensional functionalized channels. Reproduced from [46].

Through microfluidic approach, platelet derived growth factor-B (PDGF-BB) gradient was established by specific binding around the hydrogel gel channel, functioning as an in vitro model to study the mesenchymal stem cell (MSC) behaviour within the perivascular niche. It was found the encapsulated MSCs showed locally restricted morphogenetic response to the matrix-tethered PDGF-BB gradients, while soluble PDGF-BB resulted in MSCs respond in a dose-dependent manner[54]. The microfluidic approach was also used to coat GelMA on channels of

PDMS device and this cell culture substrate facilitates attachment, alignment and beating of primary cardiomyocytes[129].

2.4.3 Multi-compartmental Hydrogels

Scaffolds with multiple functionalities have the ability to control cell behaviour through various mechanical and chemical cues. Multi-compartmental hydrogel microstructures based on composites and microgels were fabricated. Multi-layered fibre matrices containing different components were deposited using electrospinning before casting PEG precursor solution and irradiating UV using a photomask (Figure

2.10). By varying the molecules within each electrospun layer, biochemical properties can be tuned precisely, demonstrating potential applications in controlling release and binding different targets[130].

Figure 2.10 Preparation of multi-compartmental hydrogel microparticles using sequential electrospinning and photopatterning. Reproduced from [130].

Microgels can be used as building blocks which assembled to a heterogeneous 3D architecture. Diverse techniques including emulsification, micromoulding, photolithography and microfluidics[131, 132] have been developed to fabricate microgels[133]. Self-assembly, manual assembly or robotic assembly of the building blocks are utilized to provide delicate functionalities, based on the surface properties and geometry of the microgels. Chiang et al. [134] developed heterogeneous hydrogel architectures using adjustable microgel building blocks by crosslinking and assembly of microgels on an electromicrofluidic platform, which is a promising alternative technology to 3D bioprinting for 2D and 3D cell patterning.

2.5 Overview and Perspectives

Synthetic hydrogels provide a versatile stage where different biochemical and architecture features can be designed and looked into separately or integrated [135].

The patterning strategies aimed at endowing hydrogels with controllable physical and biological attributes to emulate in vivo cell function. Until now, some progresses

have been made in temporal and spatial manipulation of cell microenvironment within hydrogels, but mostly represent simplified models illustrating cell-ECM interactions. With the continued development of chemical tools and manufacturing techniques, future research will highlight on capturing the dynamics of ECM[136] and engineering sophisticated hydrogel models[137] for specific biomedical applications[138]. It is envisioned that hydrogels will be engineered using integrated design strategies, allowing incorporation of multiple attributes that offer tunable and reversible chemical and physical cues to be activated or deactivated. This will enable suitable systems emulating native ECM by which to benefit the field of tissue engineering.

This literature review summarized the techniques used to create structural and biochemical features, and how these features were utilized to study cell behaviour.

Photopatterning is the most applicable technique that is capable of precise patterning, yet the requirements of complex chemistry and instrument, and the inefficient use of the expensive biomolecules, are its intrinsic limitations. This thesis explored alternative hydrogel patterning approaches that are simple, efficient, and cost- effective, thus addressing some of the limitations of the present techniques.

Chapter 3 Spatial Patterning of Hydrogels via 3D Covalent

Transfer Stamping from a Fugitive Ink

This chapter is based on the publication: Shuang Wang, Paul D. Dalton*, Tim R.

Dargaville*, Spatial Patterning of Hydrogels via 3D Covalent Transfer Stamping from a Fugitive Ink, Macromol. Rapid Commun., 2018, 39, 1700564/1–5

Abstract

In this study, a novel hydrogel patterning approach is introduced using 3- dimensional covalent transfer stamping (3D-CTS) from a fugitive ink. A model transfer molecule (7-diethylamino-3-(4’maleimidylphenyl)-4-methylcoumarin

(CPM)) was absorbed into a 3-dimensional fugitive ink stamp made from poly(caprolactone), and then encapsulated in a poly(ethylene glycol) (PEG) hydrogel. As the CPM diffuses from the fugitive ink to the curing hydrogel it was selectively covalently bound at the boundary of the ink and the hydrogel. Removal of the fugitive ink by solvent exchange led to a negative copy of interconnected channels patterned with the localized transfer of the molecules at the hydrogel interface. Our results suggest that small molecules can be patterned in porous hydrogels in a spatially controllable manner and that the transferred amount can be tuned.

3.1 Introduction

Hydrogels are an important class of biomaterials based on their ability to imbibe high weight fractions of water[12]. In the past decade, synthetic hydrogels have been frequently explored as tissue engineering scaffolds to mimic the function of extracellular matrix (ECM), which provides structural and biochemical support for

cells to function[25, 136]. One major requirement for thick tissue/scaffold constructs is the sufficient transport of nutrients and waste products such that cells remain viable. An obvious approach, analogous to the circulatory system found in native tissue, is the introduction of connected channels. Prevailing strategies used to create channels in hydrogels and elastomers have included using fugitive inks based on sugars[84, 109] and thermos-responsive polymers[108], physically extracted templates[85], and chemical photoablation[114]. Of these subtractive techniques, sacrificial moulding by dissolution of fugitive inks has attracted much attention for use with hydrogels because of its simplicity, speed, reproducibility[84], and ability to create channels with dimensions ranging from tens of micrometres[107] to millimetres in scale[111].

While physical patterning is useful for introducing channels, it does not address the need for chemical patterning when aiming to create ECM-like hydrogel structures complete with spatially-defined biochemical cues[135, 136]. Strategies for covalent chemical patterning of biomolecules to hydrogels (without introducing porosity) often exploit photochemical reaction[26]. Luo and Shoichet[56] created pillars of immobilized peptides with agarose hydrogels by using photo-labile linkers that yielded thiols under a focused laser. These spatially-revolved thiols were then conjugated with maleimide-peptides to create the biochemical pillars for guided axonal growth. Several other strategies have been used to control the distribution of biomolecules on or within hydrogel including physical photochemical masks[49,

139], diffusion gradients[79], and graded two-photon laser patterning[140]. More intricate multi-material patterning is also possible using stereolithography with a confocal laser to pattern in 3-dimensions (3D) in a spatially and temporally controllable manner[55, 75]. A recent review by Yanagawa et al. has summarized

these and other techniques for hydrogel microfabrication of 3D constructs and comment on the need for cheap and scalable technologies[28].

It is hypothesized that if a hydrogel were cured around a fugitive ink coated with molecules containing reactive groups then both the physical and chemical patterning could be combined to create hydrogels containing channels with the lumen selectively patterned – a process coined, ‘3D covalent transfer stamping’ (3D-CTS).

Herein, a proof-of-principle is provided that hydrogels containing channels decorated with fluorescent molecules are feasible using this method combining physical and chemical patterning into one procedure.

3.2 Materials and Methods

3.2.1 Materials

TM Poly(ε-caprolactone) (PCL) with Mw 50,000 (Capa 6500C) in the form of 3 mm pellets was purchased from Perstorp Caprolactones (Cheshire, United Kingdom). 1H

NMR (600 MHz, CDCl3)/ppm: δ 1.4 (m, –O–CH2–CH2–CH2–CH2–CH2–C(=O)–),

1.5–1.7 (m, –O–CH2–CH2–CH2–CH2–CH2–C(=O)–), 2.3 (t, –O–CH2–CH2–CH2–

CH2–CH2–C(=O)–), 4.1 (t, –O–CH2–CH2–CH2–CH2–CH2–C(=O)–). The polymer has a Mn of 45911 with a polydispersity of 1.7 as measured by GPC using polystyrene standards for calibration. 4-arm PEG thiol (PEG-4SH, Mw = 5,000, purity 98.4%) was obtained from Jenkem Technology Co., Ltd. (Beijing, China). 1H

NMR (600 MHz, CDCl3)/ppm: δ 1.6 (t, 1H, –SH), 2.7 (q, 2H, –CH2–SH), 3.6 (m,

1 152H, –OCH2–CH2O–). Mn estimated by H NMR measurements was 6820 and polydispersity was 1.5 as determined by GPC measurement. 4-arm PEG vinyl sulfone (PEG-4VS, Mw = 20,000) was obtained from Jenkem Technology Co., Ltd.

1 (Beijing, China). H NMR (600 MHz, CDCl3)/ppm: δ 3.6 (m, 512H, –OCH2–CH2O–

), 6.1 (d, 1H, –SO2CH=CH2), 6.4 (d, 1H, –SO2CH=CH2), 6.8 (dd, 1H, –

1 SO2CH=CH2). Mn estimated by H NMR measurements was 22892 and polydispersity was 1.1 as determined by GPC measurement. IR absorption bands conformed to the structures of the polymers (Figure 3.1).

Figure 3.1 Infrared spectra of PCL, 4-arm PEG vinyl sulfone and 4-arm PEG thiol.

7-Diethylamino-3-(4’-maleimideylphenyl)-4-methylcoumarin (CPM, 402.44 g/mol, purity 99%) was purchased from Setareh Biotech (Eugene, the USA). Acetone

(analytical grade) was from Thermo Fisher Scientific. PBS buffer (pH = 7.2) was used throughout the experiments.

3.2.2 Methods

Design and fabrication of the PCL scaffolds: A rectangular prism designed in a

CAD software (SolidWorks, Dassault Systèmes S. A.) was imported to open-source

3D slicer software, where layer thickness, lay-down pattern, and filament spacing could be defined, thus generating a G code file that instructed the movement of the stage collector. The PCL scaffolds were fabricated via the additive fabrication technique of fused deposition modelling (FDM) by a custom-made BioExtruder device built by Paulo Bartolo of Leiria, Portugal. PCL pellets were loaded into the

material chamber of the BioExtruder. A plunger was inserted into the chamber to assist the flow of polymer into the rotating extrusion screw by pneumatic air pressure. The temperature of the material chamber and the rotating screw was set at

105 °C to melt the PCL. By applying adjustable voltage (4 – 7 V) to the rotating screw, consistent PCL filament flow through a needle was extruded before printing was initiated. Deposition of the PCL filament solidified on the moving stage collector to form the predefined pattern. Using needles with different gauges, scaffolds with filament sizes of 180, 250 and 400 μm were fabricated. The 3D struts were removed from the stage and cut to small pieces using scalpel for further use.

Figure 3.2 A photograph of the BioExtruder used in this study (a), a Schematic representation of the FDM process that produces laydawn patterns of PCL (b), and a

photograph of a typical 4-layer PCL printed structure with a filament diameter of

250 μm, laydown pattern of 0/90o, and filament spacing of 1.6 mm (c).

Morphology observation of the filament integrity and interlayer adhesion was performed using scanning electron microscopy (SEM, JEOL JSM-6360A, Tokyo,

Japan). A 10-nm thick gold was sputtered onto PCL scaffolds using a Leica EM

SCD005 gold coater. PCL scaffolds were placed on a pre-tilted stub with an angle of

70⁰ to take the side view of the scaffolds. The images were taken with accelerating voltage of 10 kV using the secondary electron image (SEI) detector.

Fabrication of the CPM/PCL templates: The CPM fused PCL templates (1 mM

CPM/PCL, 3 mM CPM/PCL and 5 mM CPM/PCL stamps) were obtained by soaking the PCL scaffolds in 500 μL of DMSO solution containing 1 mM, 3 mM or

5 mM CPM for 2 h. PCL templates absorbed CPM from DMSO due to their similar hydrophobicity, yet PCL itself neither dissolved nor swelled in DMSO. After 2 h, the templates were briefly washed with DMSO and water and blotted dry, thus removing any CPM attached on the filament surface.

Preparation of the precursor solution: Fresh solutions of 7 mM PEG-4VS and 21 mM PEG-4SH were prepared in PBS buffer before use. Certain molar ratios of PEG-

4VS to PEG-4SH were mixed and topped up with PBS to reach a solid content (w/v) of 10%. It was briefly vortexed and centrifuged to mix the components evenly and avoid any bubbles.

Characterization of the hydrogels: Rheology was performed on an Anton Paar

MCR302 rheometer in an oscillating mode using a 25 mm parallel measuring plates at 25 °C. Data points were collected every 2 s with constant shear strain of 0.05 and constant frequency of 0.5 Hz. A volume of 55 μL of freshly prepared precursor

solution was pipetted onto the lower glass plate of the rheometer and the top plate lowered to a gap of 0.1 mm. The handling time between preparing the precursor and starting the rheological test was 1.5 min. The tests were run in triplicate without gas flow to minimize dehydration during the tests.

Swelling ratios of cured hydrogels were measured by immersing them in ddH2O for

24 h at room temperature to reach equilibrium swelling (msw), and then dried to constant mass at 60 °C (mdry). The mass swelling ratio (Q) and water content were determined using the following equations:

Q = (msw−mdry)/mdry

Water content (%) = 100*(msw−mdry)/ msw

CPM Absorbed onto and release from CPM/PCL templates:

A 1.24 mM CPM stock solution was prepared by dissolving 0.5 mg CPM in 1 mL

DMSO. This stock solution was diluted into PBS to make working solutions of 100 nM to 1000 nM. To obtain a calibration curve, a blank sample containing 50 μL of

20 μM PEG-4SH and 50 μL of PBS was used. The standards containing 50 μL of 20

μM PEG-4SH and 50 μL of CPM working solution were incubated at room temperature for 1 h to get stable fluorescence reading. The value of the blank was automatically subtracted and each standard was done in triplicates. The release of

CPM was measured by immersing the 1 mM CPM/PCL and 3 mM CPM/PCL stamps (with mass of 6.0 ± 0.2 mg and surface area of 56.2 mm2) in 1 mL PBS. At various time intervals, 50 μL of the release solution was removed, mixed with 50 μL

PBS solution of 20 μM PEG-4SH, and incubated in a black plate at room temperature for 1 h before reading the fluorescence.

The amount of absorbed CPM onto PCL stamps was measured by dissolving the

CPM/PCL structures in 10 mM PEG-4SH in 90% acetone and comparing the fluorescence to a calibration curve of CPM solutions diluted in 10 mM PEG-4SH in

90% acetone. The blank used was 50 μL of 20 μM PEG-4SH in PBS + 50 μL of 90% acetone. It was noted the presence of PCL in the system did not influence the fluorescence intensity, whether with or without CPM.

The total amount of CPM absorbed into the templates was calculated based on the equation from the calibration curve:

Fluorescence intensity (FI) = –389.7 + 33.4 c, R2 = 0.9893

Triplicate samples were tested to calculate the amount of absorbed CPM onto PCL stamps (Table 3.1) and to plot the release profile of CPM from the CPM/PCL template.

Table 3.1 Total amount and absorbed density of CPM onto PCL stamp

Template Total amount of CPM absorbed Absorbed density of CPM onto

onto PCL (nmol/mg PCL) PCL stamp (nmol/mm2)

1 mM CPM/PCL 0.05 0.005

3 mM CPM/PCL 0.10 0.010

5 mM CPM/PCL 0.17 0.018

Covalent transfer stamping process: The 1 mM CPM/PCL and 3 mM CPM/PCL stamps were placed side-by-side on a hydrophobic glass slide pre-treated with

Sigmacote and 30 μL of the precursor solution was pipetted onto each template. A hydrophobic glass coverslip pre-treated with Sigmacote was put on the 1.5 mm thick spacer, forming a round disk of polymer solution between the glass slide and

coverslip. Every 1 h, microscope images were taken and for each gel the mean fluorescence intensity of three regions of interest (Figure 3.3) was measured using

NIS Elements microscope imaging software until 6 h.

Figure 3.3 Three regions of interest used to monitor the change of mean fluorescence intensity over time. Scale bar = 1 mm.

Sacrificial moulding process: After transferring for 2 h or 6 h, the hydrogels were soaked in 50 mL of 90% acetone. The CPM/PCL templates were removed by applying ultrasonication for 10 min. Another 10-min sonication in fresh 50 mL of

90% acetone removed any PCL and unbounded CPM in the hydrogel. The patterned hydrogels were left in 50 mL of water for 2 d before taking further characterizations.

To confirm the covalent bonding between CPM and the hydrogel, the patterned hydrogel using 3 mM CPM/PCL stamp for 6 h was further soaked in 10 mL DMSO for 24 h and transferred back into water for 2 d before taking fluorescence images and checking intensity change.

Fluorescence measurements: The fluorescence readings were obtained by using a

POLARstar OPTIMA Microplate Reader (BMG Labtech, Germany) in 96-well black microplates, using a 355-nm broad band excitation filter and a 460-nm emission filter to cover the 387-nm excitation and 463-nm emission of CPM.

Microscopy: The microscope images were obtained on a Nikon SMZ25 fluorescence stereomicroscope under DAPI mode or bright field mode. The confocal laser scanning microscope (CLSM) images were obtained on a Nikon A1R confocal microscope. Large images the lateral sections were stitched from the scanned small images. The z-stack images were scanned in the range of width 1260 × height 1260 × depth 260 μm, with a step of 10 μm. Microscopy settings were identical in order to make comparison among samples.

3.3 Results and Discussion

3.3.1 The Thiol-Michael Addition Reactions

Mechanism of the thiol-Michael addition reactions

Due to the selectivity and favourable kinetics under physiological conditions, thiol-

Michael addition reactions have been documented to extensive applications in molecule synthesis, polymer crosslinking, and biomolecule functionalization. A thiolate anion generated under alkaline conditions attacks the electron-deficient vinyl group to yield a thioether adduct. The nucleophilic attack, which is rate-limiting step of the thiol-Michael addition, depends on the basicity of the catalyst, acidity and steric accessibility of the thiol and the electrophilicity of the vinyl[141, 142], through both base-catalysed mechanism and nucleophile-catalysed mechanism pathways

(Scheme 3.1).

Scheme 3.1 Two pathways of thiol-Michael addition reaction: base-catalysed and

nucleophile-catalysed mechanisms. Reproduced from [143].

Typical enes used in thiol-Michael addition reactions are acrylates (A), methacrylates (MA), vinyl sulfones (VS), and maleimides (MAL). Depending on the electron withdrawing group coupled to the vinyl group, the electron-deficiency of the

C=C bond varies towards its susceptibility in the thiol-Michael addition reactions

(Scheme 3.2). 4-arm PEG-MAL (PEG-4MAL) was found to have significantly faster crosslinking reaction and bioligand incorporation than those of PEG-4VS and

PEG-4A. Both PEG-4MAL and PEG-4VS can form gels at low polymer weight percentage of 4%[40]. In another study, vinyl sulfones show higher reactivity and selectivity towards thiols in the presence of acrylates. In terms of the stability of the thiol-Michael addition product, vinyl sulfones result stable thioether sulfone bond, while the counterparts of acrylates and maleimides form thioether ester or succinimide bonds that are relatively susceptible to hydrolytic degradation[144].

Scheme 3.2 Reactivity of commonly utilized vinyl groups in thiol-Michael addition

reactions. Reproduced from [143].

Characterizations of the hydrogels

PEG was chosen for this proof-of-principle study as it commonly used in tissue engineering and provides highly swollen and transparent gels[145, 146]. The gelation chemistry used in this study was the thiol-vinyl sulfone Michael-type crosslinking reaction using 4-arm PEGs to rapidly form hydrogels with stable covalent linkages[41]. The rheological tests of PEG-4VS/PEG-4SH (molar ratio 1:1) precursors (Figure 3.4) showed that the gelation time is greatly influenced by pH – the higher the pH is, the faster the crosslink proceeds. The gelation time under pH 7,

7.5, and 8 is 8.7 ± 0.3, 5.6 ± 0.7, and 2.1 ± 0.1 min, respectively. In addition, the storage modulus of the gel crosslinked at pH = 8.0 reached 6000 Pa after 10 min, while it took 16 min and 25 min for gels crosslinked at pH of 7.5 and 7.0 to reach a storage modulus of 1600 Pa. These results indicate at higher pH PEG-4VS/PEG-4SH system has a more rapid reaction rate and a greater crosslinking extent.

Figure 3.4 Typical plots of the storage modulus of the stoichiometric precursor from rheological measurements and gelation time under different pH conditions (inset).

Importantly, PEG hydrogels formed by thiol-Michael addition reactions will still cure under non-stoichiometric conditions. This allows the adjust of the reactive components for further introduction of functionality [42]. Either thiol or vinyl sulfone functionalized hydrogels can be synthesized by tailoring the feed ratio for post functionalization. As shown in Figure 3.5, for precursors with molar ratio of

PEG-4VS and PEG-4SH of 1:1.5 and 1.5:1, the storage modulus at 30 min is 2380

Pa and 1218 Pa, respectively, and the gelation time is 6.5 ± 1 and 10.0 ± 0.6 min, respectively. PEG-4SH acts as the crosslinker as it has a lower molecular weight (Mw

= 5,000) compared with PEG-4VS (Mw = 20,000). When PEG-4SH is in excess, the system forms a network with higher crosslinking density at a faster speed, thus leading to larger storage modulus and shorter gelation time.

Figure 3.5 Typical plots of the storage modulus of the non-stoichiometric precursor from rheological measurements under pH = 7.2 and gelation time of the two precursors (inset).

In this study, to allow covalent functionalization of the transfer molecule CPM to thiols, PEG-4VS was reacted with 1.5 molar equivalent of PEG-4SH so that free thiol residuals would present in the gel (Scheme 3.3). The obtained clear and transparent hydrogel had a mass swelling ratio (Q) of 22.4 ± 1.1 and % water content of 95.8 ± 0.2. Theoretically, the concentration of the unreacted thiols in the hydrogel was 7.3 mM.

Scheme 3.3 PEG-4VS crosslinked with PEG-4SH via thiol-Michael addition reaction.

3.3.2 Fabrication of the Fugitive Ink Stamp

To create the 3D fugitive ink stamp, PCL was printed via the additive fabrication technique of fused deposition modelling (FDM) by Bioextruder. First invented by

Scott Crump, FDM is the most widely used 3D printing process that builds up layer- by-layer 3D parts by heating and extruding thermoplastics. The BioExtruder used here is an additive biomanufacturing system proved to be capable of fabricating 3D porous PCL scaffolds with well-defined structural characteristics[147].

The PCL scaffold was then infused with a model transfer molecule, 7-diethylamino-

3-(4’maleimidylphenyl)-4-methylcoumarin (commonly referred to as ‘CPM’), by soaking in a solution of CPM dissolved in DMSO. As the concentration of the soaking solution increased, the overall amount of absorbed CPM increased and the penetration of CPM into the PCL templates was greater, as shown in Figure 3.6.

Figure 3.6 The fluorescence stereomicroscope images of the CPM/PCL stamps (a) and corresponding lateral images of the PCL filaments (b) using 1 mM (left), 3 mM

(middle), and 5 mM (right) soaking solution. Scale bar = 1000 μm (a) and 500 μm

(b).

The advantage of using CPM is that the coumarin group increases in fluorescence when the maleimide reacts with thiols via Michael addition so that the transfer reaction can be monitored[148]. CPM was reported as one of the first examples of thiol probes utilizing thiol addition to the maleimide moiety[149, 150]. The coumarin fluorescence is significantly quenched by the adjacent maleimide ring.

When the double bond in maleimide group becomes saturate to form a thioether by addition of a thiol, the parent coumarin fluorescence revived, producing a CPM-thiol adduct with excitation/emission maxima of 387/463 nm (Scheme 3.4). This blue fluorescent thiol-reactive dye is widely used to quantitate thiols in microplate reactions, cells and plasma without separation step[151, 152].

Scheme 3.4 Reaction scheme showing CPM fluorescence increases upon reacting with thiols.

3.3.3 The Release Profile of CPM from the CPM/PCL Stamps

The 3D-CTS process is dependent on the physical process of limited diffusion of

CPM from the CPM/PCL stamp into the hydrogel and therefore it is important to understand the release kinetics of the dye from the stamp. To achieve a repeatable process of 3D-CTS, the release profile of CPM from PCL soaked in a 1 mM and 3 mM solution was first studied. While the fluorescence intensity of CPM itself did not correlate with concentration, it was found that the fluorescence intensity of the CPM- thiol adduct showed linear correlation with the concentration of CPM over the range of 0–400 nM (Figure 3.7). It is notable that the CPM fluorescence did not increase in the control experiment where it was mixed with vinyl sulfone (expect no reaction) instead of PEG-4SH, indicating CPM was indeed covalently bonding to PEG-4SH.

Figure 3.7 Concentration calibration plots of CPM reacted with 10 mM PEG-4SH to

create the fluorescent adduct and PEG-4VS control.

Thus, the released CPM from the CPM/PCL stamp into PBS was quantitated converting it to the thiol-adduct with PEG-4SH then monitoring the fluorescence intensity over time using the regression linear equation FI = 45.3 + 11.1c (R2 =

0.998). As shown in Figure 3.8, the initial release of CPM from the 3 mM CPM/PCL stamp was relatively fast, and then reached a plateau of 7.6% after 5 h and remained constant up to 24 h. The amount of CPM absorbed onto the PCL stamp (Table 3.1) was determined to be 0.10 nmol/mg PCL. When soaking solution with lower concentration was used (1 mM) this dropped to 0.05 nmol/mg PCL, although the total percentage released increased to 12.2%, indicating a certain degree of irreversible absorption of CPM to PCL. The difference in the released percentages of

CPM is reasonable because of the greater penetration of CPM into the PCL filament when the concentration of the soaking solution increased (Figure 3.6), reducing diffusion out from PCL to the aqueous environment.

Figure 3.8 Release profile of CPM into PBS from 1 mM and 3 mM CPM/PCL stamp.

2.3 The 3D Covalent Transfer Stamping Process

3.3.4 The 3D Covalent Transfer Stamping (3D-CTS) Process

CPM/PCL stamps were immersed in PEG-4VS and PEG-4SH (1:1.5 mole ratio).

The bright-field image and a schematic of the transfer process are shown in Figure

3.9.

Figure 3.9 Bright-field microscope image (1×) of the 1 mM CPM/PCL (left) and 3 mM CPM/PCL stamps (right) embedded in PEG hydrogels (a) and a schematic of the process of 3D-CTS (b). Scale bar = 1 mm.

There is an increase over 6 hours in the mean fluorescence intensity of regions of interest. Select fluorescence microscopy images at 0, 3, 6 hours qualitatively show

the increase in fluorescence intensity as the CPM is transferred to the hydrogel from

PCL stamps soaked in 1 mM and 3 mM solutions of CPM (Figure 3.10). This phenomenon is due to the recovery of fluorescence in the coumarin moiety of CPM as the maleimide quenching group is lost during the reaction with thiols (Scheme

3.4). Although the PEG reached gelation within a few minutes, the fluorescence increased over a period of hours. As shown in Figure 11, for 3 mM CPM/PCL stamp, the mean fluorescence intensity reached a plateau after 4 h, which was in good agreement with the model release experiment in Figure 3.8, suggesting the available CPM has been exhausted after approximately 4 h. It is noticeable that the saturation in functionalization of the interior wall was only limited by the amount of

CPM released as there were high excess of unreacted thiol groups available. Controls without PEG-4SH showed no increase in fluorescence (Figure 11), confirming the change in fluorescence is due to the reaction of maleimide with thiol.

Figure 3.10 Representative fluorescence stereomicroscope images (1×) after 0 h (a),

3 h (b) and 6 h (c). Scale bar = 1 mm.

Figure 3.11 The change in mean fluorescence intensity with time of the three ROIs of the CPM/PCL stamp. Data is shown as mean ± standard deviation, n = 3.

3.3.5 Characterization of the Patterned Hydrogels

After gelation and transfer of the CPM to the hydrogel, the PCL was removed by immersing the composites in 90% acetone with gentle sonication then transferring back to water, leaving a negative copy of the PCL structure in the hydrogel (Figure

3.12 a and b). As can be seen from Figure 3.12 c–e, the amount of CPM transferred to hydrogel could be changed by using PCL templates soaked in varied concentrations of CPM solution, and by varying the transferring time.

To examine the fidelity of the channelling and 3D-CTS processes the fusion point between two PCL struts (Figure 3.13a, inset) and the corresponding negative copy within the hydrogel was examined (Figure 3.13 b and c). The figures clearly show the CPM was distributed homogeneously on the channel lumen and that the channels were interconnected at the points where the PCL fused during the FDM process.

Even after extensive washings in DMSO the CPM remained, again confirming the covalent nature of the patterning (Figure 3.14). This is significant as while non- covalent patterning has been recently reported by de Lange et al. [153] who made a hydrogel microarray with physically entrapped protein-coated microparticles in a porous hydrogel matrix, this is the first time covalent patterning of channel lumen has been reported without using complex photolabile chemicals and processes.

Figure 3.13 (a) z-stack confocal laser scanning microscope (CLSM) image and corresponding scanning electron microscope (SEM) image of the sacrificial template

(a, inset), and CLSM images of two typical lateral sections (b and c) of the patterned hydrogels using 3 mM CPM/PCL stamp and 2 h transfer time.

Figure 3.14 The fluorescence stereomicroscope images (3×, inset 1×) of the patterned hydrogel using 3 mM CPM/PCL stamp for 6 h before (a) and after DMSO treatment (b). Scale bar = 500 μm.

To investigate how far into the hydrogels the CPM migrated during transfer, we measured the distance from the channel for which an arbitrary 90% of the fluorescence intensity remained. Two transfer times (2 and 6 h) resulted in 3 and 20

μm penetration, respectively (Figure 3.15). The small migration distance indicated limited diffusion of CPM after desorbing from the template and its reaction with the thiol groups took place close to the interface of the template and the hydrogel.

Figure 3.15 Intensity profiles across the channels of the patterned hydrogels using 3 mM CPM/PCL stamp for 2 h (a) and 6 h (b), respectively.

A series of different patterns (Figure 3.16) were produced using PCL templates with different filament size, filament spacing, and laydown pattern (0/60o (Figure 3.16a) and 0/90o (Figure 3.16b–d)) fabricated from FDM process. The channels were defined by the structure and the size of the template, showing the fidelity of the sacrificial moulding process (Figure 3.16e–h). The conjugation of CPM was confined to pre-defined areas and was distributed homogeneously along the channels, showing the feasibility of the 3D-CTS approach.

(a) and (b) 180 μm; (c) 250 μm; (d) 400 μm. Filament spacing: (a) 1 mm; (b) 1 mm and 4 mm; (c) and (d) 1.6 mm.

3.4 Evaluation of the Approach

3.4.1 Advantages of the Approach

The approach combining sacrificial moulding and 3D covalent transfer stamping is a novel way to create hydrogels featuring hierarchical structures and molecular patterning. The same thiol-Michael addition was utilized in this study to form the

hydrogel and to covalently conjugate the patterning molecules, yet the chemistries for gelation and molecular conjugation are flexible and broad, depending on the kinetics of the relevant reactions. The sacrificial template is made from hydrophobic polymer PCL, which can endure long-term storage and transportation after printing.

We aimed to modify a thiol probe CPM, but theoretically any hydrophobic molecules would be able to be patterned on the channel wall. The physical and chemical patterning is produced with a simple, inexpensive, and scalable process, without the use of sophisticated chemistries and instrument.

3.4.2 Limitations of the Approach

The dissolution of the template was realized using acetone which makes the process non-biocompatible for in situ cell encapsulation. Transfer molecules must be stable during acetone treatment. The percentage of transferred molecules is low (ca. 10%) and the process is time-consuming as the transferring only relies on physical diffusion of the transfer molecules from the stamp to hydrogel.

3.5 Conclusions

A simple approach combining 3D-CTS and sacrificial moulding was developed to physically and chemically pattern hydrogels in a single step. While a model transfer molecule was used here the same approach could be applicable to a range of other molecules. The method avoided the use of UV light or lasers to generate the patterning and only a small amount of the transfer molecules was needed as the conjugation was defined to selected regions by the template. The simplicity and the efficient use of the transfer molecules are especially important when it comes to patterning expensive biomolecules such as peptides and proteins for studying cell-

substrate interactions and manipulating cell growth in tissue engineering applications.

Chapter 4 Multi-scalable Peptide-modified Channels in

Hydrogels for Cell Study

Abstract

It is important to understand cell-substrate interaction within a 3D model that has a specific tethered biosignal. In this study, cell-adhesive peptide is covalently attached onto the channel lumen by perfusing CRGDSGK peptide solution in the channels made from sacrificial moulding from PCL templates. The diameters of the channels depend on the adjustable sizes of the template resulting from additive manufacturing techniques, ranging from couples of to hundreds of diameters. It was found that

NIH-3T3 cells elongated and proliferated along the patterned channels compared with the control groups with unpatterned channels. This work demonstrated a feasible way to design a 3D environment with both physical channels and site- specific tethered biomolecules to study cell behaviour.

4.1 Introduction

It has long been recognized that cell behaviour is profoundly influenced by extracellular matrix (ECM)[154]. Hydrogels that biochemically and biophysically resembling nature ECM have been useful to study cell-cell and cell-matrix interactions[136]. Commonly used matrices derived from natural resources such as

Matrigel are complex and variable in composition from batch to batch[155]. Thus, designer scaffolds from synthetic hydrogels that are conductive to physical and biochemical manipulations have drawn much attention. Scaffolds with controlled characteristics provide important platforms to investigate the effects of biomolecule distribution and scaffold architecture on tissue regeneration, which in return benefits

the optimal design of the scaffolds[156]. Among the various biomaterials, modified poly(ethylene glycol) (PEG) hydrogels have been popular due to the flexibility of including specific functionality to their biologically inert nature, thereby allowing precise biological questions to be answered in basic research. Compared with physical adsorption, covalent immobilization of bioactive molecules to hydrogels is superior due to the endurance of attachment required in cell culture process. Photo- induced conjugation is a widely approach used to tethering biomolecules in a site- specific manner using a laser light. The group of Shoichet developed site-specific biomolecular modification of hydrogels through Michael-type addition with photo- exposed thiols upon precise cleavage of the nitrobenzyl[56] and coumarin moieties[75]. Despite these studies, the efficient use of biomolecules is still limited as large extent of unreacted biomolecules had to be washed away after the conjugation to the residue functional groups in the hydrogel.

Controlled porosity is of increasing interest in biomedical and tissue engineering applications. Sacrificial moulding using a 3D printed template is a widely used approach whereby defined architectures can be translated to the porous features within the hydrogels. Hydrogel precursor is casted around the template and allowed to cure before removing the template by varied techniques, including dissolution in solvents[84], aspiration by manual pulling or by vacuum[85], and liquefaction upon thermal change[108]. To make the channel diameters adjustable to certain applications, the dimensions of the channels attainable defined by the template were studied. Templates fabricated using fused deposition modelling (FDM) technique normally resulted channel diameters from 150 µm to millimetre scale. Lower limit of diameter was ca. 20 µm using sugar or Pluronic templates through small nozzles but surface aberration occurs due to loss of precision during deposition. Alternatively,

melt electrospinning writing (MEW) technique was used to produce fine filament with high surface smoothness which was transferable to hydrogel[107]. The attainable ranges of template diameters in a single print were investigated using

FDM and MEW techniques, thus creating channels with a wide range of adjustable sizes in hydrogels.

Recently, Brandenberg and Lutolf[112] demonstrated photo-patterned vessels using fluidic perfusion of a reactive dye that tethered to the channel wall via Michael addition. In this study, the perfusion method was adopted to fabricate biomolecules patterned channels in which NIH-3T3 cells were seeded and cultured. The aim of this study was to establish a novel hydrogel patterning approach by which adjustable sized channels decorated with biomolecules could be fabricated. The efficiency of the biomolecule usage, the ease of control over the sizes of the channels, the flexibility of the dose and distribution of biomolecules, as well as cell response to the channels, were investigated and discussed.

4.2 Materials and Methods

4.2.1 Materials

4-arm PEG maleimide (PEG-4Mal, Mw = 20,000) was obtained from Jenkem

1 Technology Co., Ltd. (Beijing, China). H NMR (600 MHz, CDCl3)/ppm: δ 3.6 (m,

1 –OCH2–CH2O–), 6.7 (s, CH2=CH2). Mn estimated by H NMR measurements was

22560 and polydispersity was 1.1 as determined by GPC measurement. Peptide with a sequence of CRGDSGK (Mw = 721.8) and fluorescein isothiocyanate (FITC)- tagged CRGDSGK (Mw = 1111.2) were ordered from Pepmic Co., Ltd (Suzhou,

China). PCL filament 1.75 mm was purchased from 3d4makers.com (Mn = 47,500, polydispersity = 1.8). GibcoTM Dulbecco’s Modified Eagle Medium (DMEM), L-

glutamine 200 mM (100×), Pen Strep, and 0.05% Trypsin-EDTA (1×) were sourced from ThermoFisher Scientific. Materials not mentioned here were from the same source in Chapter 3.

4.2.2 Methods

Fabrication of PCL structures using fused deposition modelling (FDM): PCL filament and scaffolds were printed using a FlashForge Dreamer 3D printer

(FlashForge Corporation, Figure 4.1a). The nozzle with an aperture size of 0.2 mm was purchased from Micro Swiss LLC. Temperatures were set at 115 for extrusion and 30 for the platform. The printed parts were collected on a piece of aluminium foil stick to the platform so that they can be peeled off with ease after printing. PCL filament of different diameter was obtained by adjusting the nozzle travel speed from 500 to 3000 mm/min with 500 mm/min intervals (Figure 4.2a, bottom to top). 2D and 3D structures were designed using Autodesk CAD software and the G-codes were generated using Simplify3D software.

Figure 4.1 Desktop FlashForge Dreamer 3D printer (a) and custom-built MEW machine (b).

Melt electrospinning writing (MEW) of PCL: In a 3 mL syringe was filled 2 g of medical grade PCL, and a plastic piston was plugged into the open end. The syringe

was heated at 65 overnight with tip pointed upwards to aggregate air close to the opening. The trapped air was released from the molten polymer by pushing the piston. The PCL-loaded syringe was allowed to cool down to room temperature and stored in a dry environment until use. A blunt 21G (inner diameter: 510 µm) syringe nozzle was affixed to the syringe before being mounted into the heating Teflon cylinder. In the MEW machine (Figure 4.1b), The PCL syringe was heated for at least half an hour to allow full melt of PCL at temperatures of 90 for the cylinder and 95 for the nozzle. The PCL melt was electrospun under pneumatic control (4 bar) on the collector at a distance of 6 mm with applied voltage of 5 kV. Linear PCL filaments of different diameters were obtained by adjusting the collector speed from

60 to 600 mm/min with 60 mm/min intervals (Figure 4.2b, bottom to top). In order to facilitate easy transfer and encapsulation of the PCL templates, cover glasses were placed on the collector and used to collect the filaments.

Figure 4.2 Adjustable diameters of the FDM (a) and electrospun PCL filaments (b) by adjusting the nozzle travel speed (a) and the collector speed (b).

Preparation of the precursor solution and the encapsulation of PCL in the gel:

Fresh solutions of 14 mM PEG-4Mal and 21 mM PEG-4SH were prepared in sodium acetate/acetic acid buffer (pH = 4) before use. PEG-4Mal to PEG-4SH were mixed at a molar ratio of 1.5:1 and topped up with the buffer to reach a solid content

(w/v) of 5, 10 and 15%. The precursor was briefly vortexed, centrifuged and quickly

dispensed around PCL template (filament or scaffold) hung between two spacers on a hydrophobic glass slide pre-treated with Sigmacote siliconizing solution. A hydrophobic glass coverslip pre-treated with Sigmacote was put on the spacer, forming a round disk of polymer solution between the glass slide and coverslip. The gelation was allowed to proceed in a humidified chamber for 30 min.

To conjugate the peptide throughout the hydrogel, PEG-4Mal was first mixed with

CRGDSGK peptide diluted with the same acetic acid buffer. The mixture was allowed to react for 1 h before PEG-4SH was added to form a gel. Ellman’s tests were performed to monitor the consumption of thiols in the mixture of PEG-4Mal and the peptide (thiols not detectable after 10 min), and in the peptide solution of the same concentration as in the mixture (Figure 4.3).

Figure 4.3 Stability of CRGDSGK peptide in acetic acid buffer (pH = 4).

Sacrificial moulding process: After 30 min of gelation, the hydrogels encapsulated with the straight PCL template were soaked in PBS and the template could be easily pulled out as the gels swelled in PBS. Hydrogels with complex structured PCL template were placed in 90% acetone and the template could be removed by

applying ultrasonication for 20 min. The hydrogels with channels were soaked in

PBS for 1 d before patterning with the peptide.

Characterization of the hydrogels: Swelling ratios of cured hydrogels were measured using the same methods described in Chapter 3. The swelling after gelation was calculated using the equation below:

Gel swelling after gelation = (water content after swelling – water content in the precursor) / water content in the precursor

Uniaxial compression was performed in air at 37 using an Instron universal testing system, fitted with a 5 kN load cell. The gel disk was compressed at a rate of

0.005 mm/min. Each sample was performed at least triplicate. Young’s modulus was calculated as the slope of the linear region (10–15% strain). Data were displayed as means ± standard deviation.

The preparation of peptide mixture: The chemical structures of the peptides were shown in Scheme 4.1. CRGDSGK peptide was dissolved in PBS as a 10% (w/v,

138.5 mM) solution. FITC-tagged CRGDSGK (1 mg) was first dissolved in 5.8 µL of DMSO and then 52.6 µL of PBS was slowly added to make up 15.4 mM stock solution. The aliquoted peptide solutions of 2 µL were stored in −80 freezer. To make 1 mM peptide mixture of CRGDSGK and FITC-tagged CRGDSGK (molar ratio 9:1), 2 µL of the two stock solutions were taken and diluted with 304 µL of

PBS.

Scheme 4.1 Chemical structures of CRGDSGK and FITC-tagged CRGDSGK.

Peptide patterning of the gels: The hydrogels with channels were placed on a glass slide. PBS in the channel was absorbed using Kimwipes and 1 mM peptide mixture was injected into the channel using an ultrafine needle. The hydrogels were left in a humidified, dark chamber for 30 min. The perfused peptide mixture was taken out using a pipette tip and its fluorescence intensity and residual thiol groups were measured. The gels were washed extensively with PBS and stored in PBS in dark.

Fluorescence calibration curve and measurements: The fluorescence readings were obtained by using a POLARstar OPTIMA Microplate Reader (BMG Labtech,

Germany) in 96-well black microplates, using a 485p excitation filter and a 520p emission filter to cover the 488-nm excitation and 525-nm emission of FITC.

A serial dilution of the solutions was made as peptide mixture standards and the fluorescence intensity vs. the concentration was plotted (Figure 4.4). Errors were calculated as the standard deviation from the mean values. The fluorescence standard

curve was used to obtain the concentration of the peptide mixture after perfusion in the channels.

Figure 4.4 Fluorescence standard curve of the peptide mixture.

Ellman’s test: Ellman’s reagent was used to quantify free thiols in solution. Dilution buffer consisted of 0.1 M Tris.HCl was prepared by mixing 5 mL of 1 M Tris.HCL

(pH = 8.0), 0.1 mL of 0.5 M EDTA and 44.9 mL of MilliQ water. Reaction buffer with a concentration of 2.5 mM of Ellman’s reagent was made by dissolving

Ellman’s reagent in dilution buffer. To a 96-well transparent plate, each containing

80 µL of Ellman’s reagent solution, 20 µL of each peptide mixture standards was added. The plate was incubated at dark for 30 min before measuring the optical density (OD) at 412 nm on the POLARstar OPTIMA Microplate Reader. The standard curve of the peptide mixture was plotted (Figure 4.5, linear range 10 − 200

µM) and used to quantify the residual thiols of the peptide mixture after perfusion.

The residual thiol concentration was shown as a percentage of the original.

Figure 4.5 Standard curve of the peptide mixture.

Microscopy: The microscope images were obtained on a Nikon SMZ25 fluorescence stereomicroscope under GFP mode or bright field mode. The confocal laser scanning microscope (CLSM) images were obtained on a Nikon A1R confocal microscope with a 10× objective. Each image acquired was 1024 × 1024 pixels. The z-stack images were scanned in the range of width 1260 × height 1260 × depth 260

μm, with a step of 10 μm. Microscopy settings were identical in order to make comparison among samples. The cell staining samples were obtained from CLSM by scanning across the channel every 10 µm and compressing all the images to a maximum intensity projection. ImageJ software was used to count the DAPI nuclei

(equals the number of cells) using at least five fields of view from different areas of three gels for each condition.

Cell culture: To make up culture media, 5 mL of Pen Strep, 5 mL of L-glutamine, and 25 mL of fetal calf serum (FCS) were added to DMEM. 5 mL of the culture media was taken into a T25 flask and underwent sterility check in 37 incubator.

A 1-mL vial of frozen NIH/3T3 mouse fibroblasts stored in liquid nitrogen was allowed to thaw and added to 9 mL of the culture media in a 50-mL Falcon tube. The

tube was centrifuged at 1000 rpm for 5 min to spin down the cells. After the cells pellet was resuspended in 10 mL of media, cell number was determined via the trypan blue exclusion test and a haemocytometer. Media was changed every three days and the cells were passaged when 80% confluency was reached. The passage number of the NIH-3T3 cells in this study is from 23 to 28.

Hydrogels preparation for cell seeding: For sterilization, the hydrogels were transferred to a 24-well plate and soaked in 2 mL/well 80% ethanol for 30 min. The hydrogels were washed for 3 times with sterilized PBS and soaked in 2 mL sterilized

PBS overnight. For acclimatization, the hydrogels were washed again with fresh sterilized PBS and soaked in 1 mL/well NIH-3T3 cell culture media. The plate was left in 37 incubator for at least 24 h before proceeding with cell seeding.

Cell seeding: Trypsin was used to detach NIH-3T3 cells from the flask. The cell number was counted before suspending the cells to a desired density to make the cell seeding density of 400 cells/mm2 channel surface area. The hydrogel was placed on a sterilized surface and a tip was used to pipette out the culture media in the channel.

The same volume of cell suspension was injected into the channel using a tip. The gel was carefully transferred to 24-well plate and kept humidified by applying 5 µL of culture media on the bottom and top of each gel. Cells were allowed to attach in static for 3 h before gently topping up with 1 mL media/well. The media was gently changed every 2 or 3 days. A sterilized blunt tweezer was used to handle the hydrogels. It is noticeable that a pipette tip was used to detach the gel right after soaking in ethanol, PBS or culture media, to avoid gels adhering to the bottom of the wells after long-term soaking.

Cell staining: The hydrogel samples cultured for a certain of time were transferred to a 24-well plate and washed with 1 mL of PBS (with Mg2+ and Ca2+) for two times.

To fix the cells, the hydrogel samples were treated with 4% paraformaldehyde (PFA) for 45 min at room temperature in a fume-hood followed by a wash with PBS. The samples were then incubated in 0.2% Triton X-100 for 1 h to permeabilize the cells.

After washing with PBS once, the samples were incubated in Alexa Fluor 488 phalloidin and DAPI (Invitrogen) in 1% bovine serum albumin (BSA)/PBS on a shaker protected from light. After staining the samples were washed extensively with

PBS and stored at 4 protected from light until confocal imaging.

4.3 Results and Discussion

4.3.1 Fabricating PCL Template using FDM and MEW

FDM can extrude thermoplastic material layer by layer onto a platform using temperature-controlled head. The filament diameters depend on the size of the extruding nozzle and the nozzle travel speed. Using a 200-µm nozzle, nozzle travel speed was tuned to establish a reproducible range of filament dimeters (170 – 530

µm, Figure 4.6a).

MEW is a manufacturing technology that combines the principles of electrospinning and additive manufacturing. It facilitates direct and accurate deposition of polymer fibres in the microscale range according to preprogrammed codes. By applying voltage between the spinneret and the collector, a stable, viscoelastic polymer jet is established and directly written on the moving collector instructed by a G-code. The filament diameter could be altered on demand within one construct by modifying printer parameters. In this study, collection speed was tuned to establish a

reproducible range of filament diameters (50 – 220 µm) using a single 21G syringe nozzle, as shown in Figure 4.6b.

Combining FDM and MEW techniques, multiscale filaments ranging from 50 µm to

500 µm could be printed and used as the channelling template in hydrogels.

Figure 4.6 Filament diameter adjusted by nozzle travel speed for FDM (a) and collector speed for MEW (b).

4.3.2 Characterization of the PEG-4Mal/PEG-4SH Hydrogels

To allow covalent functionalization of the peptide to maleimide, PEG-4SH was reacted with 1.5 molar equivalent of PEG-4Mal so that free maleimide residuals would present in the gel (Scheme 4.2). Theoretically, the concentrations of the unreacted maleimide groups in the 5%, 10% and 15% hydrogel were 2.9, 5.7 and 8.6 mM, respectively. It was found that the precursor made in PBS (pH = 7.3) formed gels instantly upon contact of the two compositions. Thus, sodium acetate/acetic acid with a low pH of 4 was used as the reaction buffer. The gels cured quickly – approximately less than 2 min for 5% solid content, 1 min for 10% solid content and

0.5 min for 15% solid content, respectively. The obtained hydrogels were clear and transparent and the swelling properties are listed in Table 4.1. The mass swelling ratio of the hydrogels decreased when the solid content increased due to the

increased crosslinking density. Interestingly, the gels further swelled to an equilibrium state after the crosslinking reaction, with the 15% gel reaching the maximum swelling of 12.6%. This swelling behaviour of the hydrogel facilitated manual removal of the template with simple linear structures from the hydrogel to form channels.

Scheme 4.2 PEG-4Mal crosslinked with PEG-4SH via thiol-Michael addition reaction.

Table 4.1 Swelling properties of the hydrogels with different solid contents

Water content in Swelling ratio Water content Gel swelling after gelation Solid content (%) the precursor (%) (Q) after swelling (%) (%)

5 95 35.1±5.8 97.2±0.5 2.3

10 90 31.5±3.3 96.9±0.3 7.7

15 85 22.3±1.3 95.7±0.2 12.6

The in situ rheological tests of PEG-4Mal/PEG-4SH precursors were not feasible due to the fast gelation. Uniaxial compression testings were performed to evaluate the mechanical properties of the swollen hydrogels with different solid contents

(Figure 4.7). The Young’s modulus almost doubled when the solid content increased from 5% to 10% (Figure 4.7, inset table). However, when the solid content was 15%, the gelation rate is so rapid and uncontrollable that the modulus varies among samples due to the heterogeneity caused by varied crosslinking density.

Figure 4.7 Typical stress-strain curves and the Young’s modulus (inset table) of the

hydrogels with different solid contents from uniaxial compression testings.

To conjugate peptide in the hydrogel, PEG-4Mal was allowed reacted with

CRGDSGK diluted in acetate buffer for 1 h. Ellman’s test suggested that the thiol groups were below the detection limit (10 µM) after 10 min of the reaction, indicating a rapid conjugation of the peptide to maleimide groups. The concentration of the peptide dilution was monitored and found to be stable over 1 h, confirming the consumption of the peptide was due to the formation of maleimide-thiol adduct instead of that of disulfide. PEG-4Mal or CRGDSGK conjugated PEG-4Mal was then crosslinked with PEG-4SH around PCL template, followed by the swell of the hydrogel in PBS and manual removal of the template to form a channel (Figure 4.8).

90% acetone was used with gentle sonication to dissolve nonlinear 2D and 3D templates. The removal process yielded channels with high fidelity. The channel diameters were 1.5 times of the size of template fibre due to swell of the hydrogel after curing.

Figure 4.8 Schematic overview of the sacrificial process (a), microscopic images (3×)

of the hydrogel with PCL filament (b) and after the removal of the filament (c).

4.3.3 The Efficiency of the Peptide Patterning

By perfusing reactive thiols in the hydrogel channel, tethered gradients were established within the PEG-based hydrogels bearing free maleimide groups (Scheme

4.3). The peptide mixture (CRGDSGK and FITC-tagged CRGDSGK, molar ratio 9:1) was perfused in the channel by placing syringe tip into the inlet and then injecting the peptide solution into the channel. At 10 and 30 min, the fluorescence images were taken and compared (Figure 4.9). It was observed the fluorescence only appeared close to the channel lumen and extending the reaction time to 30 min did not cause any spread of the fluorescence, indicating the conjugation of peptide to the maleimide groups in the hydrogel was fast enough compared with the diffusion of the peptide and the reaction was completed within 10 min.

Scheme 4.3 Covalently tethered peptide patterning established by perfusion.

Figure 4.9 Fluorescence images of the hydrogel with peptide mixture perfused in the channel for 10 (a) and 30 min (b). Scale bar = 1 mm.

To understand the efficiency of the peptide conjugation, the perfusing peptide mixture was aspirated using a pipette tip from the hydrogel channel after 30 min of perfusion and tested. Firstly, the fluorescence of the perfusing peptide mixture was measured and converted to the concentration. As shown in Figure 4.10, the unreacted peptide mixture accounted for less than 1% of the initial concentration (1 mM).

Secondly, the thiol residuals (equal to the amount of peptide) in the perfusing peptide mixture were measured using Ellman’s test and the results were below detection

limit (10 µM as determined in Figure 4.5), indicating after 30 min of perfusion there was only less than 1% of the peptide in the perfusing solution, and more than 99% of the peptide had conjugated to the channel of the gels, which is in accordance with the results evaluated from the fluorescence method. The concentration of the peptide mixture in PBS was found to be stable over 30 min, demonstrating the thiols in the peptide conjugated with maleimides in the hydrogel rather than self-oxidation to form a disulfide. Both the fluorescence test and the Ellman’s test indicated a highly efficient peptide conjugation to the channel lumen of the gels in terms of reaction time and the extent of reaction.

Figure 4.10 Unreacted peptide mixture after 30 min of perfusion in the channel.

4.3.4 Characterization of the Patterned Channels

The patterned channels were imaged to examine the fidelity of the perfusion method.

Figure 4.11 clearly shows that the sizes of the patterned channels were adjustable from 64 to 770 µm using FDM and MEW PCL filaments. With PCL templates printed by FDM, 2D and 3D patterns could be obtained. The peptide was conjugated

homogeneously on the channel lumen and the interconnected channels with high fidelity.

Figure 4.11 Fluorescence stereomicroscope images (3×) of straight patterned channels with different sizes (a), 2D (b) and 3D (c) patterns (1.5×), and z-stack confocal laser scanning microscope (CLSM) image (10×) of two interconnected channels in (c).

To investigate the penetration depth of the peptide, the distance from the channel for which an arbitrary 50% of the fluorescence intensity remained was measured, as shown in Figure 4.12. Due to the lowest crosslinking density, the 5% gel had the longest penetration distance of 138 ± 6 μm, while into 10% and 15% gel the peptide penetrated shorter distance (103 ± 4 and 93 ± 8 μm, respectively). Decreasing thee concentration of the peptide mixture resulted in a migration distance of 17 ± 2 μm due to rapid consumption of the peptide. The small penetration distance indicated

fast conjugation reaction of the peptide to maleimide residuals in hydrogel compared with the diffusion of the hydrophilic peptide. For 10% gel, the concentrations of the peptide at the interface of channels patterned with 1 mM and 0.1 mM peptide mixture were approximately 2 and 0.2 mM.

4.3.5 Cell Culture in the Channels

To investigate the effect of macroscopic properties on cellular behaviour, NIH-3T3 cells were firstly cultured in the channels (600-µm diameter) of the hydrogels with different solid contents. Hydrogels with modified adhesion peptide RGD were found to promote adhesion and viability of adult human mesenchymal stem cells[157] and fibroblasts[158] through the expression of cell surface integrins compared with in the

RGD-free controls. In this study the adhesion peptide sequence of CRGDSGK was incorporated to the hydrogel network via thiol-Michael addition reaction to promote cell attachment. At the same peptide concentration, the best cell adhesion and elongation along the channel was achieved with 15% hydrogel, demonstrating a positive effect of stiffness on cell growth (Figure 4.13). The NIH-3T3 cells only attached and proliferated on the channel wall without invading into the bulk gel.

Considering the heterogeneity of 15% hydrogel, 10% hydrogels were used for the following cell culture experiments.

Figure 4.13 Microscopic images (10×) of NIH-3T3 cells cultured in the channels of

5% (a), 10% (b) and 15% (c) hydrogels conjugated with CRGDSGK peptide (1% of maleimide groups) and on the well plate (d) at day 7. Scale bar = 100 μm.

The effect of peptide concentration was investigated by culturing NIH-3T3 cells in the channels of 10% hydrogel conjugated with 0.2 and 2 mM peptide and with no peptide as control, as shown in Figure 4.14. In the control group, NIH-3T3 cells showed round-up morphology with few initial attachment at day 4, leading to the

least proliferation after 14 days. Low concentration of 0.2 mM peptide facilitated limited but improved cell adhesion, proliferation and elongation in morphology.

High concentration of 2 mM significantly improved cell adhesion and growth, indicating an increased cell response with increasing immobilized.

Figure 4.14 Microscopic images (4×) of NIH-3T3 cells cultured in the channels of

10% hydrogel conjugated with 0, 0.2 mM and 2 mM peptide. Scale bar = 500 μm.

The perfusion method greatly reduced the amount of peptide in order to achieve the same concentration of peptide on the interface of the channels. For the 30-µL hydrogel with a single 600-µm channel, conjugating 2 mM peptide throughout the gel required 60 nmol of peptide; while with the perfusion method only 3 nmol was needed. To demonstrate the effect of these two peptide tethering modes, 5 groups of hydrogels with single channel were made. It is noticeable that group 4 gels patterned

with 0.1 mM peptide have 0.2 mM peptide tethered on the channel interface, and group 5 gels patterned with 1 mM peptide have 2 mM peptide tethered on the channel interface. The NIH-3T3 cells cultured in the channels were stained at day 7 and day 14 with Alexa Fluor 488 phalloidin for cytoskeleton actin and 4’6- diamidino-2-phenylindole (DAPI) for nuclei. As shown in Figure 4.15, both clumps of cells and individual cells were observed in the channels of all the hydrogels. The cells maintained a round shape in the channel without peptide functionalization. On the contrary, actin was visualized as more spread on the peptide functionalized channels. In the meanwhile, the cells attached and proliferated on the gel channels functionalized with adhesion peptide in a dose-dependent manner, while the number of cells residing in the hydrogel without peptide was significantly less than that of the other groups (Figure 4.16). The results indicated that the density of adhesion peptide tethered to hydrogel dramatically influenced their availability and subsequent interactions with cells.

Figure 4.15 Confocal laser scanning microscopy of NIH-3T3 cells stained with

Alexa Fluor 488 phalloidin and DAPI on day 7 and 14 in 5 groups of hydrogels.

Pictures show maximum intensity projection of confocal frames representing a height of 600 µm of the channel. Blue: nuclei; red: actin cytoskeleton. Scale bar =

200 µm.

Figure 4.16 Cell densities on the channels as determined from DAPI counting.

4.4 Evaluation of the Approach

4.4.1 Advantages of the Approach

The scale of PCL template was extended using two additive fabrication techniques –

FDM and MEW, achieving diameters from tens of micrometres to hundreds of micrometres. The perfusion method to create patterned channels is simple, rapid and efficient with commercial available materials. The high transparency of PEG permits microscopic analysis of cells embedded within the hydrogels.

4.4.2 Limitations of the Approach

The perfusion of peptide in a small channel (less than 100 µm) or in a complex 3D structure was difficult using manually handled syringe needle – a pump-actuated fluidic system through a wider open inlet should resolve the difficulty in perfusion.

Maleimide groups undergo hydrolysis in the aqueous alkaline solution[159].

Although the hydrolysis does not affect the conjugation of peptide, the compressive modulus of the hydrogel was found to decrease to half the value over a month (data not shown), which may influence long-term cell culture. The tethered peptide may experience thiol exchange with protein such as BSA in culture media[160], causing fluctuation of the peptide density.

4.5 Conclusions

A hydrogel patterning approach was developed by simply perfusing peptide solution in the channels produced from sacrificial moulding. The diameters of the template were adjustable in the fabrication process and transferable to create channels in hydrogels. Adhesion peptide was conjugated to defined regions by the template with high fidelity and efficiency. The patterned channels were shown to support cell growth in a dose-dependent manner. This approach features the efficient use of peptide and the simplicity in manipulating physical and chemical characteristics, which facilitate powerful control in engineering tunable properties to study cell- substrate interactions.

Chapter 5 Conclusions and Future Directions

In this thesis, two approaches were developed to create multi-scalable channels decorated with tethered molecules. Thiol-Michael addition reactions through end- functionalized PEGs enable well-defined hydrogels to be obtained via efficient off stoichiometric crosslinking of commercial available polymers, leaving reactive groups for rapid, high-yielding post-polymerization modification with complementary reactive components. Using sacrificial PCL template fabricated by

FDM and MEW techniques, interconnected channels with different architectures and channel diameters from micrometres to millimetres could be obtained within the hydrogels. Patterning molecules can be efficiently incorporated site-specifically onto the channel wall of the hydrogel matrix. The 3D-CTS method described in Chapter 3 allows hydrophobic molecules to be transferred from the PCL template and covalently functionalized to the interior channel wall, while the perfusion method presented in Chapter 4 enables hydrophilic biomolecules to be transferred from the perfusion solution and covalently conjugated to the channel lumen. The methods that only require easy accessible materials and facilities are simple, and subject to finely tuned design for incorporating other types of biological cues onto channels of different architectures and sizes. In literatures, independent manipulation of the hydrogel architecture and chemistry has been achieved by performing successive patterning techniques or using photolithography that enables orthogonal photopatterning and photoablation reactions in multi-functionalized hydrogels. These approaches either require laborious, multi-step operations, or sophisticated chemistries and instruments.

Chapter 3 described a method of creating patterned channels by impregnating sacrificial template with a reactive model molecule to stamp the PEG hydrogel in three dimensions. The concept of 3D-CTS was developed based on the traditional

2D stamping strategy that uses a stamp to transfer molecules onto surfaces.

Furthermore, the stamp used here is readily removed after functionalizing the channel wall through covalent transfer, thus creating physical channels and chemical modification in a single procedure. However, the use of acetone to remove PCL template is not cytocompatible, making this approach not suitable for in situ encapsulation of cells. Fugitive ink made from water dissolvable polymer such as

PVA is potentially biocompatible for the reduced osmotic pressure to the encapsulated cells when dissolving in aqueous solution. However, fast dissolution of the polymer can cause deformation of the channels and subsequently anamorphic location of the transferring molecules. The conditions required to remove the sacrificial stamp would also limit the use of patterning biologically relevant proteins, limiting the use of this technique to small, hydrophobic molecules.

To broaden the applicable patterning molecules and channel architectures, Chapter 4 demonstrated the fabrication of multi-scalable sacrificial templates that are transferrable to create channel structures in hydrogels. Due to the swell of the hydrogel after curing, linear PCL template can be easily removed by manual extraction. Taking advantage of the rapid kinetics of nucleophilic thiol-Michael addition reaction, a bioactive adhesion peptide sequence was functionalized on the channel lumen with spatial control by perfusing the peptide solution in the channels.

Cell experiment showed the patterned hydrogel supported cell growth, with preferred adhesion and proliferation at higher peptide concentration. The potential thiol exchange which may cause fluctuation of the tethering peptide density could be

resolved by using maleimide-functionalized peptide to a thiol excessive hydrogel.

Acetate buffer (pH = 4) was used to avoid PEG-4Mal/PEG-4SH precursors from instant curing upon mixing, however at such low pH it is not possible to encapsulate cells.

In future a biocompatible process that enables in situ cell encapsulation can be established, by using a slower kinetics of PEG-4VS/PEG-4SH precursor that allow gelation under physiological conditions, followed by solvent-free removal of the linear PCL template and maleimide-functionalized peptide patterning by perfusion.

Further incorporation of MMP-sensitive moieties will allow the encapsulated cells to migrate in the hydrogel matrix. By manipulating the design of PCL template, multi- scalable channels with site-specific patterning of multiple biomolecules can be engineered according to needs. It is anticipated that the proposed hydrogel patterning methods would potentially lead to useful hydrogel biomaterials, and consequently make impact in the 3D culture of cells and biological tissues, such as cell positioning, as well as building up blood vessels, bioreactor and in vitro model to study spheroid formation and cell invasion.

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