New Visible-Light Photoinitiating System for Improved Print Fidelity in

Article pubs.acs.org/journal/abseba

New Visible-Light Photoinitiating System for Improved Print Fidelity in Gelatin-Based Bioinks Khoon S. Lim,*,† Benjamin S. Schon,† Naveen V. Mekhileri,† Gabriella C. J. Brown,† Catherine M. Chia,† Sujay Prabakar,‡,§ Gary J. Hooper,† and Tim B. F. Woodfield†,‡ †

Christchurch Regenerative Medicine and Tissue Engineering (CReaTE) Group, Department of Orthopaedics Surgery and Musculoskeletal Medicine, University of Otago Christchurch, Christchurch 8011, New Zealand ‡ The MacDiarmid Institute for Advanced Materials and Nanotechnology, Victoria University of Wellington, Wellington 6140, New Zealand § LASRA, Fitzherbert Science Centre, Manawatu-Wanganui, Wellington 6140, New Zealand S Supporting Information *

ABSTRACT: Oxygen inhibition is a phenomenon that directly impacts the print fidelity of 3D biofabricated and photopolymerized hydrogel constructs. It typically results in the undesirable physical collapse of fabricated constructs due to impaired cross-linking, and is an issue that generally remains unreported in the literature. In this study, we describe a systematic approach to minimizing oxygen inhibition in photopolymerized gelatin-methacryloyl (Gel-MA)-based hydrogel constructs, by comparing a new visible-light initiating system, Vis + ruthenium (Ru)/sodium persulfate (SPS) to more conventionally adopted ultraviolet (UV) + Irgacure 2959 system. For both systems, increasing photoinitiator concentration and light irradiation intensity successfully reduced oxygen inhibition. However, the UV + I2959 system was detrimental to cells at both high I2959 concentrations and UV light irradiation intensities. The Vis + Ru/SPS system yielded better cell cytocompatibility, where encapsulated cells remained >85% viable even at high Ru/SPS concentrations and visible-light irradiation intensities for up to 21 days, further highlighting the potential of this system to biofabricate cell-laden constructs with high shape fidelity, cell viability, and metabolic activity. KEYWORDS: biofabrication, hydrogels, visible light, cell encapsulation, gelatin

1. INTRODUCTION Tissue engineering and regenerative medicine (TERM) strategies based on combining cells in tissue engineering scaffolds have been widely researched as potential solutions to replace or repair damaged and diseased tissues.1−4 These strategies have been employed to engineer various tissues such as bone, skin or cartilage.4−6 However, one current major unmet challenge in TERM is the need for personalization, where different patients require dedicated personalized engineered tissue constructs depending on the size and shape of the targeted tissue defect.7−9 As such, several research groups have explored combining emerging biofabrication or bioprinting technologies with TERM, where materials and/or cells are deposited layer-by-layer from three-dimensional (3D) imaging data collected from patients, to engineer tissue constructs that are patient specific.10,11 Biofabrication, which enables precise control over the deposition of cells and biomaterials with the aid of a computer, has shown great promise in fabricating constructs of complex and organized designs that can mimic the native tissue organization.10,12−14 There are several different types of © XXXX American Chemical Society

biofabrication techniques such as laser-assisted printing, inkjet printing, microvalve printing and extrusion printing, where the latter two are generally more promising in building large constructs more clinically relevant for tissue engineering.12 However, these techniques require specialized biomaterials (bioinks), which generally have specific rheological properties that allow printing of constructs with high shape fidelity, as well as being cyto-compatible to support cell survival, function, and phenotype. Hydrogels, which are composed of highly hydrated polymeric networks, have shown great promise as bioinks due to their structural similarity to the native extracellular matrix.15−17 Among all the different hydrogel materials, gelatin hydrogels have shown huge potential as bioinks for biofabrication of a variety of tissue such as liver, skin, cancer models and cartilage.1,18,19 Gelatin, being a product of collagen hydroSpecial Issue: 3D Bioprinting Received: March 15, 2016 Accepted: August 1, 2016

A

DOI: 10.1021/acsbiomaterials.6b00149 ACS Biomater. Sci. Eng. XXXX, XXX, XXX−XXX

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ACS Biomaterials Science & Engineering

Figure 1. Schematic of photopolymerization of Gel-MA hydrogels: (A) generation of radicals for covalent cross-linking of Gel-MA hydrogels; (B) radicals quenched by oxygen in the system led to incomplete cross-linking of Gel-MA macromers.

fidelity.35−37 Studer et al. have previously reported that once quenched by oxygen, the free radicals formed by the photolysis of the photoinitiators are converted into peroxyl radicals that do not react to unsaturated ester bonds such as methacryloyl, acrylates and methacrylate groups. These peroxyl radicals will instead react readily with protons (H+) in the system, forming either hydroxyperoxides or alcohols, which hinder the formation of covalent cross-links (Figure 1B).36,37 Although oxygen inhibition directly affects the print fidelity, this issue has been often neglected and unreported in the literature. Maintenance of shape fidelity is critical for successful biofabrication of tissue engineered constructs where, by definition, the accurate placement of cells and/or bioactive factors in 3D is necessary.14 Therefore, biofabrication without maintenance of shape fidelity, could perhaps be more accurately considered as a complicated method for simply casting hydrogels. There have been suggestions to combat oxygen inhibition by performing radical photopolymerization under inert conditions using nitrogen or carbon dioxide to displace oxygen. However, these inert conditions are not ideal for crosslinking cell-laden constructs, as this is likely to negatively affect cell viability and presents challenges for implementation in some cases. To the best of our knowledge, there has not been a systematic approach reported aimed at studying and/or minimizing the effect of oxygen inhibition in biofabricated Gel-MA hydrogels. Therefore, the aim of this paper was to biofabricate cell-laden Gel-MA-based constructs with both high shape fidelity and maximum cell viability, using systematic approaches such as increasing initiator concentrations and light intensity to minimize oxygen inhibition. A secondary aim was to compare the conventionally used UV + I2959 system to the alternative Vis + Ru/SPS system.

lysation, is not only water-soluble but also contains various peptide sequences, such as RGD, that are known to support cell adhesion and proliferation, and thus is highly favorable as a target hydrogel or bioink for tissue engineering approaches.20 In terms of biofabrication, gelatin is commonly functionalized with methacryloyl groups (Gel-MA), enabling formation of irreversible covalent cross-links using light activated radical polymerization for shape preservation post printing.1,18,21,22 These 3D plotted/extruded Gel-MA hydrogel constructs containing photoinitiators are often subjected to light curing for shape preservation. In the presence of light, the photoinitiators are able to absorb photons then dissociate into radicals, where these radicals will propagate through the methacryloyl groups, forming covalent kinetic chains to crosslink the polymers (Figure 1A). In the past decade, the most popular system reported to cross-link Gel-MA is the combination of UV light (320−365 nm) and the photoinitiator, 1-[4-(2-hydroxyethoxy)-phenyl]-2hydroxy-2-methyl-1-propane-1-one (Irgacure 2959).21,23−25 However, the use of UV light has shown that it can influence chromosomal and genetic instability in cells.26 In terms of photopolymerization, UV is also known to have limited penetration depth which might affect the overall polymerization efficiency for large constructs.27,28 Therefore, other photoinitiators that absorb in the visible light (Vis) range may offer significant advantages for tissue engineering applications. To date, a number of visible light photoinitiating systems have been investigated to fabricate cell-laden hydrogels, and include: camphorquinone,29−31 eosin Y,32 fluorescein,29 lithium phenyl2,4,6-trimethylbenzoylphosphinate (LAP),33 riboflavin,29 and rose bengal.34 Of these, only eosin Y and LAP have been applied to Gel-MA. Furthermore, disadvantages surrounding the use of these initiators are their poor water solubility,31 complex synthesis route,33 limited photoreactivity, and cytotoxicity.29,31,34 Recently, Lim et al. have applied a new light activated radical polymerization system consisting of water-soluble photoinitiators, ruthenium (Ru), and sodium persulfate (SPS) compounds, capable of absorbing photons in the visible-light range. However, one major problem observed with photo-crosslinking Gel-MA hydrogel constructs is oxygen inhibition, where oxygen can rapidly scavenge the radicals required for the crosslinking, causing incomplete or insufficient formation of crosslinks to maintain the constructs’ shape and hence shape

2. MATERIALS AND METHODS 2.1. Materials. Gelatin (porcine skin, type A, 300g bloom strength), phosphate buffered saline (PBS), methacrylic anhydride, cellulose dialysis membrane (10 kDa molecular weight cutoff), tris(2,2bipyridyl)dichlororuthenium(II) hexahydrate (Ru), sodium persulfate (SPS), L-ascorbic acid-2-phosphate (AsAp), calcein-AM, and propidium iodide (PI) were purchased from Sigma-Aldrich. Dulbecco’s modified Eagle’s medium (DMEM) high glucose, minimum essential medium alpha (α-MEM), 4-(2-hydroxyethyl)-1-piperazine-ethanesulfonic acid (HEPES), fetal calf serum (FCS), 0.25% trypsin/EDTA, nonessential amino acids (NEAA), and penicillin-streptomycin (PS) were purchased from Thermo Fisher Scientific. Fibroblasts growth B


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at 15 min for all experiments. In a first experiment, the light intensity was kept at 3 mW/cm2 whereas the photoinitiator concentrations were varied from 0.05 to 0.5 wt % and 0.2/2 to 2/20 (mM/mM) for I2959 and Ru/SPS respectively. In a second experiment the light intensity was varied from 3 to 100 mW/cm2 while keeping the photoinitiator concentrations constant at 0.05 wt % and 0.2/2 (mM/mM) for I2959 and Ru/SPS, respectively. The resultant light irradiation dosage corresponding to the intensity and exposure time used is outlined in Table 1 below.

factor 2 (FGF-2) was purchased from Millipore. Medical grade silicone sheets were obtained from BioPlexus (Ventura, USA). Bovine collagen type 1 (Col) solution was prepared and gifted by the New Zealand Leather and Shoe Research Association (LASRA). 2.2. Synthesis of Gelatin-Methacryloyl. Gelatin was dissolved in PBS at a 10 wt % concentration. 0.6 g of methacrylic anhydride per gram of gelatin was added to the gelatin solution, and left to react for 1 h at 50 °C under constant stirring, followed by dialysis against deionized water for 3 days to remove unreacted methacrylic anhydride. The purified Gel-MA solution was filtered through a 0.22 μm sterile filter then lyophilized under sterile conditions. The degree of methacrylation was quantified to be 60% using 1H-proton nuclear magnetic resonance (Bruker Advance 400 MHz). 2.3. Preparation of Gel-MA/Col Macromers and Rheology Measurement. Dried sterile Gel-MA (20 wt %) was dissolved in sterile PBS at 37 °C and left to cool overnight at room temperature. Collagen type 1 solution (1.2 wt %) was added to the Gel-MA solution at a 1:1 ratio. The final concentration of macromer solution used in this study was 10 wt % Gel-MA and 0.6 wt % collagen (Gel-MA/Col). Rheology was performed at 20.5 °C on an AR-G2 rheometer (TA Instruments) fitted with a cone-and-plate attachment. The viscosity and shear stress of the macromer solutions were measured over a shear rate sweep from 1 to 2000 s−1. 2.4. Casting of Gel-MA/Col hydrogel discs. Prior to crosslinking, photoinitiators were added to the Gel-MA/Col macromer solution, the mixture was scooped into silicone molds (Ø5 mm × 1 mm) on a glass slide and then irradiated under light (OmniCure S1500, Excelitas Technologies), with the surfaces of the gels exposed to oxygen. For UV cross-linking, light of wavelength 300−400 nm was used in combination with I2959. For visible-light cross-linking, the light of wavelength 400−450 nm was used in combination of Ru/SPS. 2.5. Mass Loss and Swelling Studies of Gel-MA/Col Hydrogels. Mass loss and swelling studies were performed to study the effect of incorporating collagen into Gel-MA hydrogels on the overall hydrogel cross-linking kinetics. Gel-MA hydrogels with and without collagen were initially cross-linked using UV + 0.05 wt % I2959 or Vis +0.2/2 Ru/SPS (mM/mM), where the samples were irradiated for 15 min at 3 mW/cm2 of light as per previously reported in literature.38,39 The samples were polymerized in a closed environment to minimize oxygen inhibition, where as described by Loessner et al., all the air bubbles in the macromer were first removed, and then a glass slide was used to cover the macromer solution to prevent influx of oxygen during the photo-cross-linking process.40 All samples were weighed immediately after cross-linking for the initial wet mass (minitial, t0), and three samples were lyophilized to obtain their dry weights (mdry, t=0). The actual macromer fraction was calculated based on the equation below mdry, t = 0 actual macromer fraction = m initial, t = 0 (1)

Table 1. Irradiation Conditions Used to Vary Irradiation Dosages

sol fraction =

m q = swollen mdry

m initial,dry − mdry m initial,dry

exposure time (seconds)

irradiation dosage (J/cm2)

3 30 50 100

900 900 900 900

3 27 45 90

Controls used for this study were samples irradiated in a closed environment where the irradiation conditions were 15 min of 3 mW/ cm2 UV or Vis light combined with 0.05 wt % I2959 or 0.2/2 Ru/SPS (mM/mM) respectively. Sample thickness was measured using vernier calipers. The level of oxygen inhibition was characterized as the percentage (%) change in thickness before (t0) and after equilibrium swelling (ts). The percentage change in thickness is given by the equation below:

percentage change in thickness =

t0 − ts 100% t0

(5)

2.7. Cell Encapsulation in Gel-MA/Col Hydrogels. Breast adenocarcinoma cells (MCF-7) were encapsulated within the GelMA/Col hydrogel discs at a concentration of 1 × 106 cells/mL. In brief, the cells were mixed with the Gel-MA/Col macromer and photoinitiators, then irradiated under the test conditions. MCF-7 laden gels were cultured in DMEM supplemented with 10% v/v FCS and 1% v/v PS. After 1 day, samples were stained with 1 μg/mL of calceinAM (live/green) and PI (dead/red) for 10 min, then imaged using a fluorescence microscope (Zeiss Axio Imager Z1). The number of live and dead cells were quantified using ImageJ software (version 1.46, National Institutes of Health) and cell viability was evaluated as follows:

viability =

no. of live cells 100% no. of live cells + no. of dead cells

(6)

2.8. Biofabrication of Cell-Laden Gel-MA/Col Constructs. 3D plotted hydrogel constructs consisting of Gel-MA/Col + UV or Vis photoinitiators were fabricated using a BioScaffolder (SYS+ENG, Germany). The bioink was extruded from a computer controlled syringe dispenser using a needle (inner diameter 300 μm), at a temperature of 20.5 °C, XY-plane speed of 500 mm/min, Z-speed of 800 mm/min, auger speed of 4.1 rpm, and a fiber spacing of 1.5 mm, in a repeating 0−90° pattern. Constructs were imaged and the diameter of fibers were measured using ImageJ. The effect of oxygen inhibition on shape fidelity was characterized as the change in fiber diameter before (d0) and after equilibrium swelling (ds), and is given by the equation

The remaining samples were then submerged in a bath of PBS and incubated at 37 °C. Samples were removed from the incubator after 1 day, blotted dry, and weighed (mswollen). The swollen samples were then freeze-dried and weighed again (mdry). The sol fraction was defined as the mass loss after 1 day and was calculated as eq 3 below. The mass swelling ratio (q) was calculated as eq 4 below. m initial, dry = m initial (actual macromer fraction)

intensity (mW/cm2)

(2)

100%

percentage change in fiber diameter =

(3)

d0 − ds 100% d0

(7)

Micro-CT analysis was performed on samples before and after equilibrium swelling. Samples were embedded in oil and scanned using a Bruker SkyScan 1272. For comparing utility of UV and Vis photoinitiator systems as bioinks for biofabrication of cell-laden hydrogel constructs, human articular chondrocytes (HACs) and human bone marrow-derived mesenchymal stromal cells (MSCs) were isolated as described previously.41 Expanded HACs at passage 2 or MSCs at passage 1 were incorporated into sterile Gel-MA/Col

(4)

2.6. Percentage Change in Thickness of Gel-MA/Col Hydrogel Discs. Hydrogel discs were prepared as outlined in section 2.4. To study the effect of oxygen inhibition on the hydrogel cross-linking kinetics, the constructs were irradiated with the surfaces of the constructs exposed to oxygen. The irradiation time was kept constant C


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Figure 2. 3D plotting of Gel-MA based bioink: (A) Gel-MA; (B) Gel-MA/Col; (C) inset of Gel-MA/Col. Scale bar = 1 mm. macromer and photoinitiator solutions at a concentration of 1 × 106 cells/mlL, then 3D bioprinted as described above. HAC laden bioprinted constructs were subsequently cultured in DMEM supplemented with 10% v/v FCS, 1% v/v PS, 10 mM HEPES, 0.1 mM NEAA and 0.1 mM AsAp for up to 21 days. 3D bioprinted constructs encapsulated with MSCs were cultured in α-MEM supplemented with 10% v/v FCS and 1% v/v PS for up to 21 days. Long-term viability of HACs and MScs in bioprinted constructs was quantified at days 1, 7, 14, and 21 as described in Section 2.7 above. An AlamarBlue assay was also performed at days 1, 7, 14, and 21 to determine the metabolic activity of HACs and MSCs according to the manufacturer’s protocol. Cell-laden bioprinted constructs were incubated in medium containing 10% (v/v) AlamarBlue reagent for 12 h. After incubation, the fluorescence was measured using excitation wavelength of 540−570m (peak excitation at 570 nm). The total metabolic activity of the bioprinted cell-laden constructs was evaluated by the measured relative fluorescence units (RFU) normalized to the wet weight of the constructs at each time point. 2.9. Statistical Analysis. All results (n = 3) were analyzed using a two-way ANOVA model and Tukey’s multiple comparisons tests. The models were constructed using GraphPad Prism (GraphPad Software, version 6). A p < 0.05 was considered as statistically significant.

of collagen on rheological dynamics of Gel-MA. In this study, it was shown that large, porous and multilayered constructs (e.g., 10 layers) with complex 3D architecture based on computer aided design (CAD) models of a dome and human nose were able to be biofabricated with high shape fidelity using this formulation of 10 wt % Gel-MA + 0.6 wt % collagen (Figure 3B, D).

3. RESULTS AND DISCUSSION 3.1. Biofabrication of Gel-MA/Col Hydrogels. To study the effect of oxygen inhibition on biofabricated Gel-MA based constructs, it was essential to first define a bioink with high print fidelity for further analysis. In this study, it was shown that using 10 wt % Gel-MA alone as the bioink, it was not possible to extrude defined fibers for successful layer-by-layer fabrication of a large construct (Figure 2A). However, incorporation of 0.6 wt % collagen successfully enhanced the extrudability of the bioink, where distinct fibers could be plotted accurately with high fidelity in multiple layers to form constructs with ordered 3D architecture (Figure 2B, C). This result is in accordance with the literature where it was previously reported that polysaccharides such as gellan gum and hyaluronic acid were required to alter the rheological behavior of Gel-MA for optimum printability.18,35 It was speculated that these polysaccharides altered the pseudoplasticity and yield stress of Gel-MA by modulating the amount of physical crosslinks formed during the printing process.35 Similarly, it was hypothesized that in this study, addition of collagen 1 into GelMA has a similar effect as the polysaccharides, where collagen was able to obstruct assembly of Gel-MA polymer chains, resulted in improved control over the flow and shear thinning properties. Rheological analysis of Gel-MA and Gel-MA/Col macromers confirmed this hypothesis, where it was observed that collagen I reduced the viscosity and shear stress of Gel-MA (Figure S1). Ongoing experiments are currently being conducted to further characterize and understand the effect

Figure 3. Biofabrication of photo-cross-linked, complex, thick hydrogel constructs consisting of completely interconnecting 3D pore network using Gel-MA/Col bioinks: (A) CAD model of designed dome structure; (B) cross-section of biofabricated dome structure (max width: 17 mm; max height: 8 mm); (C) CAD model of designed nose structure; (D) biofabricated nose structure (max length: 17 mm; max height: 11 mm). Scale bar = 2 mm.

3.2. Physico-chemical Properties of Gel-MA/Col Hydrogels. One important consideration is that while the addition of collagen to Gel-MA altered the rheological behavior of the resultant macromer solution, there may have also been a negative effect on the cross-linking efficiency of the system. As such, we conducted a mass loss and swelling study using cast hydrogel discs to evaluate the cross-linking efficiency of both the UV + I2959 and Vis + Ru/SPS system. In this study, the photopolymerization process was conducted in a closed environment to minimize oxygen inhibition, where the hydrogel macromer solution was covered with a glass slide to prevent influx of oxygen during the cross-linking process.40 As listed in Table 2, the sol fraction values, which are defined as the amount of non-cross-linked polymer after the polymerization process, are not statistically different for all samples examined (sol fraction ∼10−11%; p = 0.995). This result indicates that although addition of collagen affected the D


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significant change in thickness after equilibrium swelling, decreasing by 45% and 32% change in thicknesses for both UV + I2959 and Vis + Ru/SPS systems, respectively. This decrease in thickness was likely due to incomplete cross-linking of the Gel-MA/Col hydrogel precursor solution, where sol fraction analysis showed that high percentage change in thickness values also correspond to high sol fraction values (Figure S2). This observation indicates that the polymerization reaction was likely impaired by oxygen inhibition during the photo-cross-linking process. One of the possible mechanisms contributing to the change in thickness or fiber dimension may be due to the physically entrapped collagen leaching out from the gels. However, given that a small volume of collagen (total 0.6 wt %) was incorporated in Gel-MA to achieve the desired rheological behavior for biofabrication, it is therefore unlikely to have had a significant contribution to reducing gel thicknesses. The same trend was observed for both UV + I2959 and Vis + Ru/SPS systems where increasing photoinitiator concentration successfully reduced the overall change of thickness. Furthermore, it was shown that a minimum photoinitiator concentration of 0.25 wt % I2959 and 1/10 Ru/SPS (mM/ mM) were required to produce hydrogels that were not significantly different in overall thickness to the controls. This result was expected as it was hypothesized that increasing the initiator concentration would subsequently increase the number of radicals generated for the polymerization reaction.36,43 This radical increment was able to compensate for the radicals quenched by oxygen, permitting cross-linking of the Gel-MA macromers. A similar approach has been implemented by Melchels et al. where biofabricated GelMA/gellan gum constructs were UV irradiated while being immersed in a photoinitiator bath to prevent loss of radicals due to oxygen inhibition.35 However, as the ultimate goal is to use this bioink for fabrication of cell-laden constructs, it is important to know if cells can survive these photopolymerization conditions. It has been previously reported that high photoinitiator concentrations are generally toxic because of the generation of more detrimental reaction products such as superoxide radicals (O2··) during the polymerization process.16,44 Therefore, in this study, we encapsulated human breast adenocarcinoma (MCF-7) cells into Gel-MA/Col hydrogels to evaluate cell viability. Live−dead images showed that increasing I2959 concentration significantly increased the number of dead cells in the hydrogels (Figure 5B−E), where 0.5 wt % I2959 resulted in only 50% cell viability after 1 day. However, increasing the Ru/ SPS concentration did not result in a similar detrimental effect on the cell viability. As illustrated in Figure 5A and Figure 5F− I, there was no significant decrease in cell viability with increasing Ru/SPS photoinitiator concentration, with cell viability remaining >80% across all samples. As the numbers of moles of radicals generated in both systems are identical, this dissimilarity between the UV + I2959 and Vis + Ru/SPS systems might be due to the different radical generation mechanism. For example, the I2959 photoinitiator, once exposed to UV, absorbs photons and dissociates into radicals immediately in a simple one-way approach.45,46 However, on the other hand, the Vis + Ru/SPS is a two-step approach, where Ru was first photoexcited from Ru2+ to Ru3+ by absorbing photons in the visible-light range, followed by donating electrons to SPS.20,47 SPS then dissociates into sulfate anions and radicals which propagates through the functional methacryloyl moieties, forming covalent cross-links in the

Table 2. Physico-chemical Properties of Gel-MA and GelMA/Col Gels Fabricated Using UV + I2959 and Vis + Ru/ SPS initiating system

sample

UV + I2959

Gel-MA Gel-MA/Col Gel-MA Gel-MA/Col

vis + Ru/SPS

sol fraction (%) 10.7 10.8 10.6 10.3

± ± ± ±

3.1 4.2 3.9 3.0

mass swelling ratio, q 10.1 9.54 9.63 9.94

± ± ± ±

1.21 0.31 0.56 0.45

rheological behavior of Gel-MA, it did not affect the ability of radicals to propagate through the system to form covalent cross-links. The mass swelling ratio values were also not significantly different for all conditions examined (q ≈ 9−10, p = 0.512), further confirming that the collagen did not affect the cross-linking process, and that both UV + I2959 and Vis + Ru/ SPS behave similarly in terms of the kinetics of the radical initiated polymerization process. These results demonstrated that incorporation of collagen into Gel-MA improved the print fidelity of Gel-MA without affecting the physicochemical properties of the fabricated gels. Therefore, only Gel-MA/Col gels were examined for the remainder of the study. 3.3. Oxygen Inhibition in Gel-MA + Collagen 1 Gels. 3.3.1. Increasing Initiator Concentrations. The effect of oxygen inhibition on photo-cross-linked Gel-MA/Col hydrogels was studied by photocuring disc constructs with the surfaces exposed to oxygen. In a first experiment, the constructs were exposed to 15 min of light, both kept at an intensity of 3 mW/cm2 as per previously reported,15,17,42 where the concentration of photoinitiators were varied accordingly. The level of oxygen inhibition was then evaluated by measuring the overall change in thickness (or height) of the gel constructs after equilibrium swelling. It was observed that for both the UV and visible light system, control samples fabricated without exposure to oxygen showed less than a 10% change in thickness (Figure 4). However, when the constructs were photopolymerized in the presence of oxygen with the same light intensity and initiator concentration, the samples exhibited a

Figure 4. Percentage (%) change in thickness of Gel-MA/Col hydrogels at different initiator concentrations. Light intensity and exposure time were kept constant at 3 mW/cm2 and 15 min. *Significant differences between columns below each end of lines (p < 0.05). Inset shows thickness of disc constructs after equilibrium swelling, scale bar = 1 mm. E


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Figure 6. Percentage (%) change in thickness of Gel-MA/Col hydrogels at different light intensity. Photoinitiator concentration was kept at 0.05 wt % I2959 and 0.2/2 Ru/SPS (mM/mM) for both UV and Vis, respectively. All samples were irradiated for 15 min. *Significant differences between columns below each end of lines (p < 0.05). Inset shows thickness of disc constructs after equilibrium swelling, scale bar = 1 mm.

Figure 5. (A) Cell viability and (B−I) live dead images of MCF-7 encapsulated in Gel-MA/Col hydrogels fabricated with varying photoinitiator concentrations. Both light intensity and exposure time were kept at 3 mW/cm2 and 15 min. Scale bar = 100 μm. *Significant differences between columns below each end of lines (p < 0.05).

process. As such, it was hypothesized that the rate of sulfate radical generation was dependent on the photoactivation of Ru, which might be slower than the UV + I2959. Previous work conducted in our lab has confirmed that the UV + I2959 system does have a faster polymerization rate where complete crosslinking of Gel-MA macromers was obtained after 0.5 min of UV exposure as compared to 3 min for Vis + Ru/SPS (data not shown). 3.3.2. Increasing Light Intensity. Another approach to minimize oxygen inhibition is to increase the light intensity. It was hypothesized that a higher activation energy for the photoinitiators will increase the rate of radical production to offset the rate of radicals quenched by oxygen present in the system.37 In this study, the initiator concentrations were kept minimal at 0.05 wt % I2959 and 0.2/2 Ru/SPS (mM/mM) for UV and Vis, respectively. As expected, it was found that for both systems, increasing the light intensity from 3−100 mW/ cm2 successfully minimized oxygen inhibition as shown by the minimal change in thickness after equilibrium swelling (Figure 6). Furthermore, at a low light intensity of 3 mW/cm2, the visible-light system yielded hydrogels with a percentage change in thickness that was significantly lower to the UV system (p < 0.0001), which once again may have been due to the different rates at which radicals were generated. Cell encapsulation studies showed that there was a significant decrease in cell viability with increasing UV light intensity (Figure 7A), where 100 mW/cm2 of UV resulted in approximately 45% viability. Live−dead images also showed that increasing UV irradiation intensity resulted in an increase in the number of dead cells in the hydrogel constructs (Figure 7B−E). In contrast, there was no significant change in cell viability with increasing visible light intensity from 3−100 mW/ cm2 (Figure 7A), with equally high numbers of live cells observed in live−dead images across all light intensities (Figure 7F−I). As compared to the approach of increasing initiator

Figure 7. (A) Cell viability and (B−I) live dead images of MCF-7 encapsulated in Gel-MA/Col hydrogels fabricated with varying light intensity. Photoinitiator concentrations were kept constant at 0.05 wt % I2959 and 0.2/2 Ru/SPS for UV and Vis, respectively. All samples were irradiated for 15 min. Scale bar = 100 μm. *Significant differences between columns below each end of lines (p < 0.05).

concentrations to minimize oxygen inhibition (Section 3.3.1), our data suggests that increasing light intensity may be a more sensible and clinically relevant approach as the cells generally have better cell viability. It has been previously reported in the literature that UV can damage DNA and chromosomal stability in cells, further inducing apoptosis resulting in cell death.48 On the other hand, the visible light system showed significantly higher cell viability (∼90%) even at high intensities of 50 and 100 mW/cm2, further highlighting the potential of this system for biofabrication of cell-laden constructs. F


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likely present in 3D plotted hydrogels compared to their cast hydrogel counterparts, which subsequently quenched the radicals required for cross-linking. It was previously reported by Decker et al. that larger surface area leads to faster oxygen diffusion and higher oxygen concentration in hydrogel discs, which in turn elevates oxygen inhibition of the photopolymerization process.37 In our case, only constructs crosslinked with 1/10 mM/mM of Ru/SPS or 0.05 wt % I2959 with irradiation intensities greater than 30 mW/cm2, successfully produced hydrogels that remained stable after equilibrium swelling (Table 2). It should be noted that all biofabricated constructs had fiber diameters ranging between 500−550 μm prior to swelling (Figure 8A, C, E, G). For the UV cross-linked constructs, a significant change in fiber diameter was observed post swelling (Figures 8B, D, F, H). Constructs irradiated with 30 mW/cm2 of UV had a 72 ± 1.1% change in fiber diameter (Figure 8I), whereas further increasing the UV intensity to 50 mW/cm2 successfully reduced the change in fiber diameter to 58 ± 0.8% (p < 0.0001). In contrast, constructs photopolymerized using

Moreover, it was also emphasized that for the Vis + Ru/SPS system, increasing both the initiator concentration and light intensity did not pose any detrimental effect on cell viability (Figure 7F−I). These results suggest that the Vis + Ru/SPS system has potential as an alternative route for biofabrication of cell-laden constructs with minimum effects of oxygen inhibition and maximum cell survival. 3.4. Oxygen Inhibition in 3D Biofabricated Gel-MA/ Col Constructs. Gel-MA/Col constructs with varying concentrations of Ru/SPS (0.2/2, 0.5/5 and 1/10 mM/mM) were 3D plotted by extruding and assembling fibers in a layerby-layer fashion, with subsequent photo-cross-linking using a range of visible light intensities (3, 30, 50 mW/cm2). In terms of the UV + I2959 system, only 0.05 wt % I2959 was examined as previous results (section 3.3.1) showed that increasing I2959 concentration significantly reduced cell viability, and hence was not suitable for fabrication of cell-laden constructs. The photopolymerized constructs were then allowed to reach equilibrium swelling (1 day), where the fiber diameter of constructs were measured. The percentage change in fiber diameter was quantified as a direct measure of shape fidelity. It was observed that for UV + 0.05 wt % I2959, the construct irradiated with 3 mW/cm2 of UV dissolved after 1 day (Table 3). This result further highlighted that the shape fidelity of the Table 3. Preservation of Gel-MA/Col Constructs Fabricated Using Varying Light Irradiation Intensity and Ru/SPS Concentration after Equilibrium Swellinga light source

I2959 (wt %)

UV

0.05

vis

Ru/SPS (mM/mM)

0.2/2

0.5/5

1/10

light intensity (mW/cm2)

after equilibrium swelling

3 30 50 3 30 50 3 30 50 3 30 50

no gel gelb gelb no gel no gel no gel no gel no gel no gel no gel gelb gelb

a Irradiation time was kept constant at 15 min. bStable constructs remained after equilibrium swelling.

3D plotted Gel-MA constructs using UV + I2959 were highly affected by oxygen inhibition, and remains an issue to be addressed. Higher intensities of UV (30 and 50 mW/cm2) were required to fabricate constructs that were chemically stable after equilibrium swelling. For the visible light system, at low Ru/ SPS concentrations (0.2/2 and 0.5/5 mM/mM), no hydrogel remained after 1 day for all the intensities (3, 30, and 50 mW/ cm2) examined (Table 3). These results also suggested that a higher level of oxygen inhibition occurred in the 3D plotted constructs as compared to the cast hydrogel discs, and can be explained by the larger surface area of the extruded fibers. As the porous 3D plotted constructs were assembled by layering fibers in an ordered manner, these fibers present a large surface area as well as interconnected pore network to allow maximum diffusion of oxygen throughout the constructs. Therefore, during photopolymerization, higher amounts of oxygen molecules were

Figure 8. Biofabricated and photopolymerized Gel-MA/Col constructs before and after equilibrium swelling, scale bar = 1 mm: (A) UV, 30 mW/cm2, t = 0; (B) UV, 30 mW/cm2, t = 1 day; (C) Vis, 30 mW/cm2, t = 0; (D) Vis, 30 mW/cm2, t = 1day; E) UV, 50 mW/cm2, t = 0; (F) UV, 50 mW/cm2, t = 1 day; (G) Vis, 50 mW/cm2, t = 0; (H) Vis, 50 mW/cm2, t = 1 day. Photoinitiators concentrations were 0.05 wt % I2959 or 1/10 Ru/SPS (mM/mM), respectively. All constructs were irradiated for 15 min. (I) % Change in fiber diameter. *Significant differences between columns below each end of lines (p < 0.05). μCT images of Gel-MA/Col constructs: (J) UV, 50 mW/cm2, t = 1 day; (K) Vis, 50 mW/cm2, t = 1 day. G


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Figure 9. Live dead images of cells encapsulated in 3D bioprinted Gel-MA/Col hydrogel constructs after 1 day in culture, scale bar = 200 μm. (A) HACs encapsulated at 30 mW/cm2 UV; (B) HAC encapsulated at 50 mW/cm2 UV; C) HAC encapsulated at 30 mW/cm2 visible light; (D) HAC encapsulated at 50 mW/cm2 visible light; (E) MSC encapsulated at 30 mW/cm2 UV; (F) MSC encapsulated at 50 mW/cm2 UV; (G) MSC encapsulated at 30 mW/cm2 visible light; H) MSC encapsulated at 50 mW/cm2 visible light. Photoinitiator concentrations were 0.05 wt % I2959 or 1/10 Ru/SPS (mM/mM), respectively, and exposure time was kept at 15 min.

literature where Colosi et al. and Bertasonni et al. both showed that the viability of human umbilical vein endothelial cells (HUVEC) and liver hepatocellular carcinoma cells (HepG2) bioprinted in Gel-MA gels were 75% and 80% respectively, although different cross-linking conditions were employed.49,50 However, in this study, we observed that bioprinting of porous, cell-laden constructs cross-linked using 30 mW/cm2 of UV resulted in poor shape fidelity and chemical stability, and as a result, the constructs completely eroded within 3 days of in vitro culture, likely due to increased levels of oxygen inhibition and incomplete cross-linking (Figure 9A, E). This observation is similar to a study published by Ouyang et al. where 3D plotted photopolymerizable hyaluronic acid based constructs irradiated at 15 mW/cm2 of UV (0.05 wt % I2959 and 5 min exposure time) under ambient air had poor shape stability and eroded over time, further emphasizing the issue of oxygen inhibition in photo-cross-linked 3D bioprinted constructs.51 For above reasons, no long-term culture of 3D bioprinted, cell laden constructs cross-linked using 3 and 30 mW/cm2 of UV was possible. Further increasing the light intensity to 50 mW/ cm2 resulted in constructs of much higher shape fidelity but significantly lower cell viability of both HAC (63 ± 3.4%, p < 0.001, Figure 9A, B) and MSC (67 ± 5.7, p < 0.0001, Figure 9E, F), compared to samples cured using 30 mW/cm2 of UV light. Increasing the light intensity while keeping the exposure time constant, corresponds to a significant increment in the total irradiation dosage (intensity x time, Table 1). Colosi et al. previously reported that increasing UV irradiation dosage from 90 mJ/cm2 (15 s of 6 mW/cm2 UV) to 540 mJ/cm2 (60 s of 6 mW/cm2 UV) resulted in a detrimental decrease in HUVEC viability from 75 to 30% in bioprinted Gel-MA/alginate constructs.50 Similarly, Billiet et al. also showed that increasing irradiation dosage from 1350 to 5400 mJ/cm2 significantly decreased viability of HepG2 cells in bioprinted Gel-MA constructs from 89 to 56%.52 It is to be noted that in the bioprinting field, different irradiation conditions such as various UV intensity (2.6−15 mW/cm2), exposure time (15 s−15 min), and different initiator concentrations (0.05−0.5 wt % I2959) have been used.18,19,35,49−51 Therefore, it is challenging

the Vis + Ru/SPS had statistically (p < 0.0001) lower percentage change in fiber diameter than the UV + I2959 system. Both 30 and 50 mW/cm2 of visible-light successfully produced gels that were not only chemically and mechanically stable, but exhibited less than 15% change in fiber diameter. As shown in Figure 8J, K, μCT images also revealed that the gelMA/Col construct irradiated using 50 mW/cm2 of UV had fibers that were noticeably and uniformly thinner than the visible light irradiated samples in the x-, y-, and z-planes. These results highlight that the visible light system adopted in this study is less prone to the detrimental effects of oxygen inhibition and offers an alternative approach for biofabrication of Gel-MA hydrogel constructs with high shape fidelity. To further compare the utility of UV and Vis photoinitiator systems as bioinks for 3D biofabrication and tissue engineering of cell-laden hydrogel constructs using clinically relevant cell sources, human articular chondrocytes (HAC) and human bone marrow-derived mesenchymal stromal cells (MSC) were successfully encapsulated into porous, 3D bioprinted Gel-MA/ Col constructs with high print fidelity (Figure 9). Postfabrication, constructs were cross-linked using either UV (0.05 wt % I2959) or Vis (1/10 mM/mM Ru/SPS) for 15 min at a light intensity of 3, 30, or 50 mW/cm2. A recent protocol published by Loessner et al. showed that for encapsulation of HAC in 10 wt % Gel-MA discs using the UV + I2959 system, an intensity of 2.6 mW/cm2, 15 min exposure time and 0.05 wt % I2959 initiator concentration were recommended.40 However, in our case, no gel was formed for the 3D plotted cell-laden constructs after cross-linking using UV intensity of 3 mW/cm2, 15 min exposure time and 0.05 wt % I2959 initiator concentration conditions. Once again, these results indicate that a higher level of oxygen inhibition likely occurred in the porous 3D plotted constructs as compared to the cast hydrogel discs, similar to the results obtained previously for non-cellladen constructs (see Table 3). Further increasing the UV intensity to 30 mW/cm2 (15 min of exposure time) successfully yielded cross-linked bioprinted cell-laden constructs with good cell viability after 1 day, as observed in both the HAC (76 ± 2.1%) and MSC (79 ± 6.1%) laden constructs (Figure 9A, E). These values are similar to previous studies reported in the H


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Figure 10. Viability and metabolic acitivity of HAC and MSC encapsulated in 3D bioprinted Gel-MA/Col hydrogel constructs at 1, 7, 14, and 21 days. (A) HAC viability; (B) MSC viability; (C) fluorescence intensity of bioprinted HAC constructs normalized to hydrogel wet weight; (D) fluorescence intensity of bioprinted MSC constructs normalized to hydrogel wet weight. Photoinitiators concentrations were 0.05 wt % I2959 or 1/ 10 Ru/SPS (mM/mM), respectively. Both UV and visible light irradiation intensities were kept at 50 mW/cm2 and exposure time was kept at 15 min. *Significant differences between columns below each end of lines (p < 0.05).

viable over the 3 weeks culture period. Moreover, the viability of HAC and MSC was significantly higher in the visible light polymerized 3D bioprinted constructs (Figure 10A, B) across all time points (1, 7, 14, and 21 days). Metabolic activity (AlamarBlue assay) of bioprinted HAC and MSC increased from day 1 to day 21 for both cell types, and also for both photopolymerization systems (UV + I2959 and Vis + Ru/SPS respectively, Figure 10C, D). This observation indicated that the cells were able to grow and proliferate in the Gel-MA/Col constructs. Moreover, cell-laden HAC and MSC constructs photoirradiated using the visible light system had significantly higher metabolic activity across all time points (1, 7, 14, and 21 days) examined, compared to the UV cross-linking system (Figure 10C, D). Once again, this result showed that the visible light system is an attractive approach for 3D bioprinting of cellladen constructs with high cell viability and metabolic activity. For the above reasons, we therefore suggest that researchers investigating the use of Gel-MA hydrogels for biofabrication of viable, cell laden constructs with high print fidelity for in vitro and/or in vivo characterization may wish to consider the use of alternative visible-light photoinitiating systems (vis + Ru/SPS) described herein.

to directly compare irradiation conditions between studies reported in the literature. In contrast, 3D bioprinted Gel-MA/Col constructs photocross-linked using the visible light system had significantly higher cell viability and shape fidelity to their UV counterparts for both HAC and MSC (Figure 9). Moreover, increasing visible light intensity from 30 to 50 mW/cm2 did not significantly affect viability of HAC and MSC encapsulated in bioprinted gel-MA/Col constructs, where high cell viability (>80%) was achieved. This result is similar to cell encapsulation studies conducted in section 3.3.2, where increasing visible light intensity did not affect the viability of MCF-7 cells encapsulated in the Gel-MA/Col casted discs (Figure 7). In order to further evaluate the potential of this system for tissue engineering and regenerative medicine applications, 3D bioprinted HAC and MSC laden constructs photo-cross-linked using 50 mW/cm2 of UV or visible light were cultured for a longer period of time (3 weeks), where the cell viability and metabolic activity were examined. The viability of both HAC and MSC was shown to increase significantly in Gel-MA/Col constructs photo-cross-linked using UV from 1 day to 21 days of culture (Figure 10A). This observation was similar to a previous study by Bertassoni et al. where hepatocellular carcinoma cells (HepG2) bioprinted in UV cross-linked Gel-MA constructs showed an increase in cell viability over time.49 However, no change in cell viability was observed in constructs cross-linked using visible light, where both HAC and MSC laden constructs remained >80%

4. CONCLUSIONS We have described the development and characterization of an alternative bioink and photoinitiator system suitable for biofabrication, with an improved print window and favorable cross-linking method. Incorporating small concentrations of I


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(3) Yang, S.; Leong, K.-F.; Du, Z.; Chua, C.-K. The design of scaffolds for use in tissue engineering. Part I. Traditional factors. Tissue Eng. 2001, 7, 679−89. (4) Langer, R.; Vacanti, J. P. Tissue engineering. Science 1993, 260, 920−6. (5) Vacanti, J. P. Beyond transplantation. Third annual Samuel Jason Mixter Lecture. Arch. Surg. 1988, 123, 545−9. (6) Woodfield, T. B. F.; Malda, J.; de Wijn, J.; Péters, F.; Riesle, J.; van Blitterswijk, C. A. Design of porous scaffolds for cartilage tissue engineering using a three-dimensional fiber-deposition technique. Biomaterials 2004, 25, 4149−61. (7) Hochleitner, G.; Jüngst, T.; Brown, T. D.; Hahn, K.; Moseke, C.; Jakob, F.; et al. Additive manufacturing of scaffolds with sub-micron filaments via melt electrospinning writing. Biofabrication 2015, 7, 035002. (8) Costa, P. F.; Puga, A. M.; Díaz-Gomez, L.; Concheiro, A.; Busch, D. H.; Alvarez-Lorenzo, C. Additive manufacturing of scaffolds with dexamethasone controlled release for enhanced bone regeneration. Int. J. Pharm. 2015, 496, 541−50. (9) Costa, P. F.; Vaquette, C.; Baldwin, J.; Chhaya, M.; Gomes, M. E.; Reis, R. L.; et al. Biofabrication of customized bone grafts by combination of additive manufacturing and bioreactor knowhow. Biofabrication 2014, 6, 035006. (10) Giannitelli, S. M.; Mozetic, P.; Trombetta, M.; Rainer, A. Combined additive manufacturing approaches in tissue engineering. Acta Biomater. 2015, 24, 1−11. (11) Schon, B. S.; Hooper, G. J.; Woodfield, T. B. F. Modular Tissue Assembly Strategies for Biofabrication of Engineered Cartilage. Ann. Biomed. Eng. 2016, 1−15. (12) Bose, S.; Vahabzadeh, S.; Bandyopadhyay, A. Bone tissue engineering using 3D printing. Mater. Today 2013, 16, 496−504. (13) Mota, C.; Puppi, D.; Chiellini, F.; Chiellini, E. Additive manufacturing techniques for the production of tissue engineering constructs. J. Tissue Eng. Regener. Med. 2015, 9, 174−90. (14) Groll, J.; Boland, T.; Blunk, T.; Burdick, A. B.; Cho, D.-W.; Dalton, P. D.; Derby, B.; Forgacs, G.; Li, Q.; Mironov, V. A.; et al. Biofabrication: reappraising the definition of an evolving field. Biofabrication 2016, 8, 013001. (15) Lim, K. S.; Kundu, J.; Reeves, A.; Poole-Warren, L. A.; Kundu, S. C.; Martens, P. J. The influence of silkworm species on cellular interactions with novel PVA/silk sericin hydrogels. Macromol. Biosci. 2012, 12, 322−32. (16) Lim, K. S.; Ramaswamy, Y.; Roberts, J. J.; Alves, M.-H.; PooleWarren, L. A.; Martens, P. J. Promoting Cell Survival and Proliferation in Degradable Poly(vinyl alcohol)−Tyramine Hydrogels. Macromol. Biosci. 2015, 15, 1423−1432. (17) Nafea, E. H.; Poole-Warren, L. A.; Martens, P. J. Structural and permeability characterization of biosynthetic PVA hydrogels designed for cell-based therapy. J. Biomater. Sci., Polym. Ed. 2014, 25, 1771−90. (18) Schuurman, W.; Levett, P. A.; Pot, M. W.; van Weeren, P. R.; Dhert, W. J. A.; Hutmacher, D. W.; et al. Gelatin-Methacrylamide Hydrogels as Potential Biomaterials for Fabrication of TissueEngineered Cartilage Constructs. Macromol. Biosci. 2013, 13, 551−61. (19) Bertassoni, L. E.; Cecconi, M.; Manoharan, V.; Nikkhah, M.; Hjortnaes, J.; Cristino, A. L.; et al. Hydrogel bioprinted microchannel networks for vascularization of tissue engineering constructs. Lab Chip 2014, 14, 2202−11. (20) Lim, K. S.; Alves, M. H.; Poole-Warren, L. A.; Martens, P. J. Covalent incorporation of non-chemically modified gelatin into degradable PVA-tyramine hydrogels. Biomaterials 2013, 34, 7097−105. (21) Nichol, J. W.; Koshy, S. T.; Bae, H.; Hwang, C. M.; Yamanlar, S.; Khademhosseini, A. Cell-laden microengineered gelatin methacrylate hydrogels. Biomaterials 2010, 31, 5536−44. (22) Van Den Bulcke, A. I.; Bogdanov, B.; De Rooze, N.; Schacht, E. H.; Cornelissen, M.; Berghmans, H. Structural and Rheological Properties of Methacrylamide Modified Gelatin Hydrogels. Biomacromolecules 2000, 1, 31−8. (23) Bae, H.; Ahari, A. F.; Shin, H.; Nichol, J. W.; Hutson, C. B.; Masaeli, M.; et al. Cell-laden microengineered pullulan methacrylate

collagen I (0.6 wt %) altered the rheological behavior of GelMA to allow better control over the shape fidelity. Using this bioink, we have shown that in photopolymerizable constructs, oxygen inhibition has a detrimental effect on shape fidelity, which can be overcome by increasing UV or visible light photoinitiator concentration or light irradiation intensity. In comparison with the conventionally adopted UV + I2959 system, we demonstrated that the new visible light (Vis + Ru/ SPS) system resulted in significantly improved cell cytocompatibility, evidenced by higher cell viability at high Ru/SPS photoinitiator concentrations or visible light irradiation intensities. Porous biofabricated constructs photopolymerized using the Vis + Ru/SPS were less susceptible to the effects of oxygen inhibition and incomplete cross-linking, evidenced by significantly lower % change in fiber diameter compared to the UV + I2959 system post swelling. This study, therefore, highlights that the optimized visible light system is more suitable for bioprinting of cell-laden constructs with high shape fidelity and cell viability, and should be considered as a potential alternative bioink for researchers wishing to biofabricate complex, thick, cell-laden gelatin-based constructs.

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ASSOCIATED CONTENT

S Supporting Information *

The Supporting Information is available free of charge on the ACS Publications website at DOI: 10.1021/acsbiomaterials.6b00149. Viscosity and shear stress measurements of Gel-MA and Gel-MA/Col “bio-inks” (Figure S1); sol fraction values of Gel-MA/Col hydrogels polymerized using various initiator concentration of I2959 and Ru/SPS while exposed to oxygen (Figure S2); long-term survival of HAC encapsulated in Gel-MA/Col casted hydrogel constructs (Figure S3) (PDF)

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AUTHOR INFORMATION

Corresponding Author

*E-mail: [email protected]. Notes

The authors declare no competing financial interest.

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ACKNOWLEDGMENTS The authors acknowledge funding support from the Royal Society of New Zealand Rutherford Discovery Fellowship (T.W., RDF-UOO1204), Health Research Council of New Zealand (K.L., HRC15/483), Dr. Ferry Melchels and Dr. Isha Mutreja for scientific discussions, Dr. Elisabeth Phillips for MCF-7 cells, and the Imaging and Scanning platform of the Medical Technologies Centre of Research Excellence (MedTech CORE), as well as Dane Gerneke at the Auckland Bioengineering Institute for his technical assistance with the micro-CT scanning.

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REFERENCES

(1) Melchels, F. P. W.; Dhert, W. J. A.; Hutmacher, D. W.; Malda, J. Development and characterisation of a new bioink for additive tissue manufacturing. J. Mater. Chem. B 2014, 2, 2282−9. (2) Lu, L.; Arbit, H. M.; Herrick, J. L.; Segovis, S. G.; Maran, A.; Yaszemski, M. J. Tissue Engineered Constructs: Perspectives on Clinical Translation. Ann. Biomed. Eng. 2015, 43, 796−804. J


Article

ACS Biomaterials Science & Engineering hydrogels promote cell proliferation and 3D cluster formation. Soft Matter 2011, 7, 1903−11. (24) Bartnikowski, M.; Bartnikowski, N. J.; Woodruff, M. A.; Schrobback, K.; Klein, T. J. Protective effects of reactive functional groups on chondrocytes in photocrosslinkable hydrogel systems. Acta Biomater. 2015, 27, 66−76. (25) Klotz, B. J.; Gawlitta, D.; Rosenberg, A. J.; Malda, J.; Melchels, F. P. Gelatin-Methacryloyl Hydrogels: Towards Biofabrication-Based Tissue Repair. Trends Biotechnol. 2016, 34, 394−407. (26) Dahle, J.; Kvam, E.; Stokke, T. Bystander effects in UV-induced genomic instability: antioxidants inhibit delayed mutagenesis induced by ultraviolet A and B radiation. J. Carcinog. 2005, 4, 11. (27) Kim, S. H.; Chu, C. C. Pore structure analysis of swollen dextran-methacrylate hydrogels by SEM and mercury intrusion porosimetry. J. Biomed. Mater. Res. 2000, 53, 258−266. (28) Choi, B, Lee, M. Injectable Hydrogels for Articular Cartilage Regeneration. In Injectable Hydrogels for Regenerative Engineering; Lakshmi, S. N., Ed.; World Scientific Publishing, Singapore, 2016. (29) Hu, J.; Hou, Y.; Park, H.; Choi, B.; Hou, S.; Chung, A.; et al. Visible light crosslinkable chitosan hydrogels for tissue engineering. Acta Biomater. 2012, 8, 1730−8. (30) Jakubiak, J.; Allonas, X.; Fouassier, J. P.; Sionkowska, A.; Andrzejewska, E.; Linden, LÅ; et al. Camphorquinone−amines photoinitating systems for the initiation of free radical polymerization. Polymer 2003, 44, 5219−26. (31) Kamoun, E. A., Winkel, A., Eisenburger, M., Menzel, H. Carboxylated camphorquinone as visible-light photoinitiator for biomedical application: Synthesis, characterization, and application. Arabian J. Chem.. 2014, DOI: 10.1016/j.arabjc.2014.03.008. (32) Shih, H.; Lin, C. C. Visible-light-mediated thiol-ene hydrogelation using eosin-Y as the only photoinitiator. Macromol. Rapid Commun. 2013, 34, 269−73. (33) Fairbanks, B. D.; Schwartz, M. P.; Bowman, C. N.; Anseth, K. S. Photoinitiated polymerization of PEG-diacrylate with lithium phenyl2,4,6-trimethylbenzoylphosphinate: polymerization rate and cytocompatibility. Biomaterials 2009, 30, 6702−7. (34) Mazaki, T.; Shiozaki, Y.; Yamane, K.; Yoshida, A.; Nakamura, M.; Yoshida, Y.; Zhou, D.; Kitajima, T.; Tanaka, M.; Ito, Y.; Ozaki, T.; Matsukawa, A. A novel, visible light-induced, rapidly cross-linkable gelatin scaffold for osteochondral tissue engineering. Sci. Rep. 2014, 4, 4. (35) Melchels, F. P.; Dhert, W. J.; Hutmacher, D. W.; Malda, J. Development and characterisation of a new bioink for additive tissue manufacturing. J. Mater. Chem. B 2014, 2, 2282−9. (36) Studer, K.; Decker, C.; Beck, E.; Schwalm, R. Overcoming oxygen inhibition in UV-curing of acrylate coatings by carbon dioxide inerting, Part I. Prog. Org. Coat. 2003, 48, 92−100. (37) Decker, C.; Jenkins, A. D. Kinetic approach of oxygen inhibition in ultraviolet- and laser-induced polymerizations. Macromolecules 1985, 18, 1241−4. (38) Visser, J.; Levett, P. A.; te M?ller, N. C. R.; Besems, J.; Boere, K. W. M.; van Rijen, M. H. P.; de Grauw, J. C.; Dhert, W. J. A.; van Weeren, P. R.; Malda, J. Crosslinkable Hydrogels Derived from Cartilage, Meniscus, and Tendon Tissue. Tissue Eng., Part A 2015, 21, 1195−206. (39) Visser, J.; Melchels, F. P. W.; Jeon, J. E.; van Bussel, E. M.; Kimpton, L. S.; Byrne, H. M.; Dhert, W. J. A.; Dalton, P. D.; Hutmacher, D. W.; Malda, J. Reinforcement of hydrogels using threedimensionally printed microfibres. Nat. Commun. 2015, 6, 6933. (40) Loessner, D.; Meinert, C.; Kaemmerer, E.; Martine, L. C.; Yue, K.; Levett, P. A.; et al. Functionalization, preparation and use of cellladen gelatin methacryloyl-based hydrogels as modular tissue culture platforms. Nat. Protoc. 2016, 11, 727−46. (41) Schrobback, K.; Klein, T. J.; Woodfield, T. B. The importance of connexin hemichannels during chondroprogenitor cell differentiation in hydrogel versus microtissue culture models. Tissue Eng., Part A 2015, 21, 1785−94. (42) Young, C.; Rozario, K.; Serra, C.; Poole-Warren, L.; Martens, P. Poly (vinyl alcohol)-heparin biosynthetic microspheres produced by

microfluidics and ultraviolet photopolymerisation. Biomicrofluidics 2013, 7, 044109. (43) Soucek, M. D.; Ren, X. UV-curable coating technologies. RSC Smart Materials 2015, 15−48. (44) Bryant, S. J.; Nuttelman, C. R.; Anseth, K. S. Cytocompatibility of UV and visible light photoinitiating systems on cultured NIH/3T3 fibroblasts in vitro. J. Biomater. Sci., Polym. Ed. 2000, 11, 439−57. (45) Nicodemus, G.; Bryant, S. J. Cell encapsulation in biodegradable hydrogels for tissue engineering applications. Tissue Eng., Part B 2008, 14, 149−65. (46) Nuttelman, C. R.; Henry, S. M.; Anseth, K. S. Synthesis and characterization of photocrosslinkable, degradable poly(vinyl alcohol)based tissue engineering scaffolds. Biomaterials 2002, 23, 3617−26. (47) Fancy, D. A.; Denison, C.; Kim, K.; Xie, Y.; Holdeman, T.; Amini, F.; et al. Scope, limitations and mechanistic aspects of the photo-induced cross-linking of proteins by water-soluble metal complexes. Chem. Biol. 2000, 7, 697−708. (48) Parsons, P.; Goss, P. Chromosome damage and DNA repair induced in human fibroblasts by UV and cholesterol oxide. Immunol. Cell Biol. 1978, 56, 287−96. (49) Bertassoni, L. E.; Cardoso, J. C.; Manoharan, V.; Cristino, A. L.; Bhise, N. S.; Araujo, W. A.; Zorlutuna, P.; Vrana, N. E.; Ghaemmaghami, A. M.; Dokmeci, M. R.; Khademhosseini, A. Direct-write bioprinting of cell-laden methacrylated gelatin hydrogels. Biofabrication 2014, 6, 024105. (50) Colosi, C.; Shin, S. R.; Manoharan, V.; Massa, S.; Costantini, M.; Barbetta, A.; et al. Microfluidic Bioprinting of Heterogeneous 3D Tissue Constructs Using Low-Viscosity Bioink. Adv. Mater. 2016, 28, 677−84. (51) Ouyang, L.; Highley, C. B.; Rodell, C. B.; Sun, W.; Burdick, J. A. 3D Printing of Shear-Thinning Hyaluronic Acid Hydrogels with Secondary Cross-Linking. ACS Biomater. Sci. Eng. 2016, DOI: 10.1021/acsbiomaterials.6b00158. (52) Billiet, T.; Gevaert, E.; De Schryver, T.; Cornelissen, M.; Dubruel, P. The 3D printing of gelatin methacrylamide cell-laden tissue-engineered constructs with high cell viability. Biomaterials 2014, 35, 49−62.

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