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A simple and high-resolution stereolithography-based 3D bioprinting system using visible light

crosslinkable bioinks

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2015 Biofabrication 7 045009

(http://iopscience.iop.org/1758-5090/7/4/045009)

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Biofabrication 7 (2015) 045009 doi:10.1088/1758-5090/7/4/045009

PAPER

A simple and high-resolution stereolithography-based 3D bioprintingsystem using visible light crosslinkable bioinks

ZongjieWang1, RaafaAbdulla2, BenjaminParker1, Roya Samanipour1, SanjoyGhosh2 andKeekyoungKim1

1 School of Engineering, University of British Columbia, Kelowna, BC,V1V1V7, Canada2 IrvingK. Barber School of Arts and Sciences, University of British Columbia, Kelowna, BC,V1V 1V7, Canada

E-mail: keekyoung.kim@ubc.ca

Keywords: bioprinting, stereolithography, biomaterials, visible light crosslinking

AbstractBioprinting is a rapidly developing technique for biofabrication. Because of its high resolution and theability to print living cells, bioprinting has beenwidely used in artificial tissue and organ generation aswell asmicroscale living cell deposition. In this paper, we present a low-cost stereolithography-basedbioprinting system that uses visible light crosslinkable bioinks. This low-cost stereolithography systemwas built around a commercial projector with a simplewater filter to prevent harmful infraredradiation from the projector. The visible light crosslinkingwas achieved by using amixture ofpolyethylene glycol diacrylate (PEGDA) and gelatinmethacrylate (GelMA)hydrogel with eosin Ybased photoinitiator. Three different concentrations of hydrogelmixtures (10%PEG, 5%PEG+5%GelMA, and 2.5%PEG+7.5%GelMA, all w/v)were studiedwith the presented systems. Themechanical properties andmicrostructure of the developed bioinkweremeasured and discussed indetail. Several cell-free hydrogel patterns were generated to demonstrate the resolution of the solution.Experimental results withNIH3T3fibroblast cells show that this system can produce a highly vertical3D structure with 50μmresolution and 85%cell viability for at least five days. The developed systemprovides a low-cost visible light stereolithography solution and has the potential to bewidely used intissue engineering and bioengineering formicroscale cell patterning.

1. Introduction

Tissue engineering aims to use a combination ofcells, materials, engineering and biochemical factorsto generate biologically functional tissues and organs[1]. Due to developments in nanotechnology andmicrofabrication, many fabrication techniques cancreate customized micrometer scale scaffolds andcontrol the distribution of cells more accurately thanbefore, which addresses many of the previouschallenges in tissue engineering [2, 3]. Among all thetechniques that have been developed in the lastdecade, bioprinting is one of themost promising andadvanced fabrication techniques. The aim of bio-printing is to use small units of living cell encapsu-lated biomaterials (usually hydrogels) to form thedesired tissue-like structure. Bioprinting can bedistinguished from the conventional 3D printingtechniques that have been used to print temporarilyscaffolds for surgery [4], as bioprinting is capable of

depositing cell-laden hydrogels in vitro rather thanthe scaffolding material itself. The advantages ofbioprinting include high controllability of cell dis-tribution, high resolution of cell deposition, andcost-effectiveness compared to other cell patterningtechniques with the help of expensive ‘clean roommicrofabrication’. Bioprinting has been utilized toregenerate artificial bone structure [5], cartilage [6],liver [7], and skin [8], as well as to investigate tumorgrowth [9], mimic vascular networks [10, 11], andmanipulate stem cell differentiation [12]. Aside fromthe applications for tissue engineering, bioprinting isalso employed to coat proteins [13] and deposit cellsfor drug delivery tests [14, 15]. The aforementionedsuccessful applications for bioprinting prove itsversatility and feasibility. However, traditional bio-printing techniques still have many inherent short-comings. In order to discuss those deficiencies, it isnecessary to understand the working principles ofmainstream bioprinting techniques.

RECEIVED

12 June 2015

REVISED

5November 2015

ACCEPTED FOR PUBLICATION

23November 2015

PUBLISHED

22December 2015

© 2015 IOPPublishing Ltd

Currently, there are three mainstream bioprintingtechniques: inkjet printing, laser-assisted printing, andextrusion-based printing [16, 17]. Inkjet printing isvery similar to the conventional inkjet 2D printing[18]. The major difference is that inkjet bioprintingutilizes bioinks, the cell-laden prepolymer solution, toreplace the normal 3D printing inks. However, thistechnique has proven difficult in printing vertical 3Dstructures because the inkjet printing head cannotgenerate a continuous flow [16]. This inherent dis-advantage significantly limits the application of inkjetprinting systems in bioprinting. To address this issue,extrusion printing uses a pump or a piston to con-tinuously dispense the bioinks in the form of a highvertical cylinder. However, due to the shear stressapplied to cells during extrusion, the viability of extru-sion printing is relatively low, usually lower than 80%[17]. Laser-assisted printing uses a high-intensity laserto deposit the bioinks for cell patterning. Because thereis no force directly applied to cells during printing,laser-assisted printing can preserve very high cell via-bility. However, major problems associated with thislaser-assisted printer are its expensive laser sourcesand the complexity of laser pulse control [19, 20].

A stereolithography-based technique was recentlymodified to fabricate cell-free [21, 22] scaffolds as wellas cell-laden scaffolds [23–26]. The stereolithography-based bioprinting system uses digital micromirrorarrays to control the light intensity of each pixel forprinting areas in which light-sensitive polymer mate-rials are polymerized by the light. Stereolithographyhas many advantages as compared to the techniquesmentioned above. First, stereolithography can printthe light-sensitive hydrogels layer by layer rather thanin straps or droplets. No matter how complex or largethe layer is, the printing time for each layer is the same.The total printing time depends only on the thicknessof the structure. The printing time under stereo-lithography has been reported to be around 30 min[23]. Thus, stereolithography can significantly reduceprinting time. Also, stereolithography is a nozzle-freeprinting technique, which results in cell viabilityhigher than 90% and resolution down to 200 μm [26].Table 1 summarizes a brief comparison of inkjet,laser-assisted, extrusion, and stereolithography bio-printing systems. It demonstrates that stereo-lithography is a very competitive technique forbioprinting due to its high resolution, high speed, andhigh cell viability.

However, current stereolithography bioprintingsolutions have several limitations. Chen et al utilizedan ultra-violet (UV) light source to polymerize gelatinmethacrylate hydrogel (GelMA) with Irgacure 2959photoinitiator [23–25]. UV light has been reported todamage cell DNA [27, 28], and even induce cancer ofthe skin [27, 29]. Moreover, UV light is also harmful tothe DMD array itself [30]. Tuan et al adopted a com-mercial visible light stereolithography system to poly-merize cell-laden polyethylene glycol diacrylate(PEGDA) hydrogel with lithium phenyl-2,4,6-tri-methylbenzoylphosphinate (LAP) photoinitiator [26].The LAP is still a UV-sensitive photoinitiator, thoughit can be crosslinked by a near-UV blue light. Becausethe system is commercially available but not fully cus-tomized to 3D bioprinting, controlling the system andthe printing process may not be suitable for envir-onmentally sensitive cells such as stem cells. More-over, previous studies showed that near-UV blue light(400 nm–490 nm wavelength) is toxic to mammalcells and disruptive to cellular processes [31, 32].

In this paper, we present a custom-built visiblelight based stereolithography bioprinting systemwhich consists of a commercial beam projector andbioinks, a mixture of PEGDA, GelMA, and eosin Ybased photoinitiator. To our knowledge, we demon-strate the first detailed schematic of a visible lightbased stereolithography system and reveal the neces-sity of an infrared ray (IR) filtering water filter to thesystem. By using a commercially available projector,the cost of the whole bioprinting system has been sig-nificantly reduced. Experimental results withNIH3T3cells demonstrate that this proposed low-cost systemcan support the bioprinting of visible light curablehydrogels with 50 μm resolution and high cell viability(∼85%) for at least five days.

2.Materials andmethods

2.1. Preparation and characterization of hydrogelmaterialsThe preparation process of the visible light cross-linkable PEG solution was adopted from [33]. Briefly,we mixed 10% w/v PEGDA solution (Mn 700) in PBSwith 0.01 mM eosin Y disodium salt (eosin Y), 0.1%w/v triethanolamine (TEA), and 37 nM 1-vinyl-2pyrrolidinone (NVP). GelMA was synthesized by theprocess previously reported [34]. 5 g gelatin wasdissolved in 50 ml dimethyl sulfoxide solvent at 50 °Cwith stirring. Then, 0.3 g 4-dimethylaminopyridine

Table 1.Comparison of existing bioprinting techniques.

Inkjet Extrusion Laser assisted Stereolithography

Viability (%) [17, 23, 24, 26] >85 40–80 >95 >85

Speed fast slow medium fast

Cost low medium high low

Vertical structure poor good medium good

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(DMAP) was added to the mixture and dissolved.Subsequently, 2 ml of glycidyl methacrylate was addedto the mixture, and the mixture was stirred for twodays at 50 °C. The mixture was then dialyzed withreverse osmosis (RO) water at room temperature forfive days. The water was changed twice a day. Afterdialysis, a freeze-dried sample was achieved vialyophilization. Then, 10% w/v GelMA macromeresolution was prepared with eosin Y based photoinitia-tor. Finally, the PEG andGelMA solutions weremixedto obtain visible light crosslinkable PEG–GelMAhybrid hydrogels (5% PEG+5% GelMA and 2.5%PEG+7.5% GelMA). All the materials used abovewere purchased from Sigma Aldrich, St. Louis,MO,USA.

To test mechanical properties, 5 ml of each PEG–GelMA prepolymer solution was pipetted into a petridish 6 cm in diameter and exposed to the developedvisible light system for 12 min to crosslink the hydro-gel. Five cylindrical specimens (12.7 mm in diameter)from each type of PEG–GelMA hydrogel were pun-ched from the petri dish. The compressive Young’smodulus of the each sample was tested by a dynamicmechanical analysis (DMA) instrument (Q800, TAInstruments, New Castle, DE, USA). The compressivemodulus was determined as the slope of the linearregion between strains from5% to 15%.

For the mass swelling ratio test, six cylindrical spe-cimens were prepared using the same method asdescribed above, and the residual liquid of the sampleswas removed with a paper tissue. Then, the swollenweight of the sample was measured with a precisionbalance (Sartorius, Mississauga, ON, Canada). Subse-quently, these samples were lyophilized at −40 °C forfive days to determine the dry weight of the samples.The mass swelling ratio was given by the followingformula:

mass swelling ratioswollen weight of the sample

dry weight of the sample.=

We examined the microstructure of the samplescoated with 10 nm of gold–palladium (Au–Pd) alloyusing sputtering. Images of the microstructure of eachtype of mixture were taken using a scanning electronmicroscope (Mira3 XMU, TESCAN, Brno, CzechRepublic).

2.2. Cell culturingNIH 3T3 fibroblasts were cultured at 37 °C and 5%CO2. The cellmedium consists ofDulbecco’smodifiedEagle medium (DMEM) with 10% v/v fetal bovineserum and 1% v/v penicillin–streptomycin. The cellmediumwas changed every two days. All the materialsused in this subsection were purchased from LifeTechnologies, Grand Island,NY,USA.

2.3.Design of bioprinting systemA beam projector (HD6510BD, Acer, Taipei, Taiwan)was connected to the computer and employed as the

stereolithography projection device. The projectorwas placed 10 cm away from the petri dish (6 cm indiameter). Between the petri dish and projector, therewas a 4 cm thick water filter to block the harmfulinfrared radiation (IR) generated by the projector. Aschematic diagramof the proposed system is presentedin figure 1(a). During printing, the displayed pattern ofthe projector was controlled with the computer via theHDMI port. The desired pattern was designed by theuser and was saved as an STL (stereolithography) file.Then, the STL file was sliced with a gap of 100 μm inthe Z direction via Freesteel Z-level slicer software(version 1.5, Freesteel, Liverpool, UK). The slices oflayers were converted to binary color patterns.Exposed areas were projected with white color whilethe remaining parts were projected with black color.Finally, such slices were projected by the projectorlayer by layer to achieve the pattern of hydrogels.

A typical bioprinting process is presented infigure 1(b). In the pre-printing stage, the STL filescontained desired pattern structures, which werefirst designed, then sliced to form a group of layeredimages with binary color. In this paper, three differ-ent patterns were printed: the logo of the Universityof British Columbia (UBC), a mesh pattern with var-ious line widths to verify the resolution, and a meshwith uniform line width for cell encapsulationexperiments.

For the mesh pattern without cells and the UBClogo, 287 μl (approximately equal to 100 μm volumeheight in a petri dish 6 cm in diameter) of eosin Ybased 10% PEG hydrogel prepolymer was evenlyadded to the petri dish with a pipette. The printingtime was two minutes per layer, and we printed eightlayers. The images taken of the projected UBC logopattern during the printing process are shown infigure 1(b). Immediately after printing, the unpoly-merized part of the solution was removed to exposethe printed pattern, and the sample was colored withfood dye for clear visualization of the pattern. To printthe cell-encapsulated mesh pattern, 287 μl of theaforementioned hydrogels (10% PEG, 5% PEG+5%GelMA, or 2.5% PEG+7.5% GelMA) mixed withNIH 3T3 fibroblasts (5×106 cells ml−1) was evenlyadded to the petri dishwith a pipette.

2.4. Assessment of cell viabilityIn order to verify that the printing process is biocom-patible, we checked the cell viability immediately afterthe printing. Printed samples were washed three timeswith PBS and treated with a live/dead assay (Biotium,Hayward, CA, USA) for 60 min. Subsequently, weobserved the assayed samples under a confocal fluor-escent microscope (FV1000, Olympus, Tokyo, Japan).To investigate long-term cell viability, we cultured thesamples for five days. On day 5, we repeated the live/dead cell viability assessment.

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We used a 10× objective and two fluorescentchannels and one phase contrast channel to capturethe microscope image. To avoid the crosstalk betweenlaser signals, sequential imaging modes were used totake groups of confocal fluorescent images with a20 μm step in the Z direction. Fluoview ASW software(version 3.1a, Olympus, Tokyo, Japan) was used tostack the fluorescent and phase contrast images takenby themicroscope. Its optional 3Dmodule was used torender and reconstruct the 3D distribution of cells. Toanalyze the cell viability, the images taken by themicroscope were converted to 16 bit gray value for-mat, and the cell number was counted briefly with theparticle counting function (Otsu Method) in ImageJ(NIH, Bethesda, MD, USA). Finally, the cell viabilitywas calculated by the following formula:

cell viabilitynumber of live cells

number of all cells.=

2.5. Statistical analysisA one-way analysis of variance (ANOVA analysis)function in MATLAB 2014b (MathWorks, Natick,MA, USA) was used to statistically analyze the data ofmechanical properties, swelling ratio, and cell viabi-lity. Results are shown as an average±standarddeviation.

3. Results and discussion

The photocrosslinking of hydrogels is popular inbioprinting. Previous studies have already utilizedPEG [6, 35] and GelMA [10, 23, 24, 36] with various

Figure 1.Bioprinting system’s configuration and process. (a) Schematic diagramof developed visible light stereolithography-basedbioprinting system.Controlled by the computer, the projector generated binary color lights on every pixel based on a user-definedpattern. The patternwas transferred to the bioinks (red areas in the picture) layer by layer by using high-intensity white light to patternvisible light crosslinkable hydrogels. (b)Bioprinting procedure of the developed system. In the pre-printing stage, the desired patternwas designed via CAD software and sliced to a layerfile. During the printing process, the patternwas transferred to the visible lightcrosslinkable inks layer by layer. In post-printing, the uncrosslinked inkswere removed, and the patternwaswashed for furtherculturing.

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photoinitiators as bioinks for bioprinting. In tissueengineering, most of the widely used photoinitiatorswork within the range of UV wavelength. Detailedinformation regarding the absorbing peak of thephotoinitiators is presented in table 2. The eosin Ybased photoinitiator is designed for crosslinkinghydrogels in the range of green light (around 514 nm).Considering the potential harmful effects of UV light,such as DNA damage to cells [27, 28], cancer [27, 29],and the negative effects of near-UV blue light [31, 32],eosin Y based visible light polymerization of hydrogelhas an advantage in maintaining cell function. Eosin Yhas also been reported to be less toxic than Irgacure2959 [33]. Taking into account all of its advantages, theeosin Y based photoinitiator is an excellent choice forbioprinting systems. Although PEG is widely used inbioprinting and has excellent mechanical properties,PEG-encapsulated cell viability has been observed todecrease significantly after two days culturing [37].Hutson et al demonstrated that, after mixing withGelMA hydrogel, the cell viability of hybrid PEG–GelMA improved [37]. Motivated by this method, weinvestigated the feasibility of using hybrid PEG–GelMA hydrogel with an eosin Y based photoinitiatorfor visible light stereolithography-based bioprinting.

3.1. Characterization of bioinksMechanical properties of the hydrogels play an impor-tant role in cell functions and differentiation. Asreported in [38], the elasticity of thematrices can directthe fate of stem cells. Bahney et al reported that, afterchanging the photoinitiator from Irgacure 2959 toeosin Y, themechanical properties of the PEGDAwereimproved [33]. Thus, it is necessary to measure themechanical properties of PEG–GelMA hybrid hydro-gels crosslinked by the visible light from a beamprojector. We utilized the DMA instrument to deter-mine the compressive Young’s modulus of the hybridhydrogels, as presented in figures 2(a) and (b). Becausethe molecular weight of the PEG we used is small(∼700 Da), the 10% PEG hydrogel had a high com-pressive Young’s modulus (∼200 kPa). After mixing itwith GelMA, the Young’s modulus of the hydrogelswas decreased significantly, because the GelMA is arelatively ‘soft’ hydrogel [34]. However, compared tothe 10% GelMA (15 kPa) with Irgacure 2959 photo-initiator, the Young’s modulus of 5% PEG+5%GelMAwith eosin Y based photoinitiator is four timeshigher (60 kPa). The compressive Young’s modulus of2.5% PEG+7.5% GelMA is 20 kPa, which is one-third of that of 5% PEG+5% GelMA. Therefore, the

Table 2.Comparison ofwidely used photoinitiators in tissue engineering.

Chemical name Abbreviation Absorbing peak (nm) Sources

1-[4-(2-hydroxyethoxy)-phenyl]-2-hydroxy-2-methyl-1-propanone Irgacure 2959 257 [49, 50]Lithiumphenyl-2,4,6-trimethylbenzoylphosphinate LAP 375 [50]2,2′-azobis[2-methyl-n-(2-hydroxyethyl)propionamide] VA-086 385 [51]2′,4′,5′,7′-tetrabromofluorescein disodium salt eosin Y 514 [33, 50]

Figure 2.Mechanical properties,mass swelling ratio, andmicrostructures of eosin Y initiated PEG and PEG+GelMA. (a) Strain–stress curve. (b)Compressive Young’smodulus (*p<0.0001, n=5). (c)Equilibrium swelling properties of visible light crosslinkedPEG and PEG andGelMA (*p<0.05, **p<0.001, n=6). (d) SEMpicture of 10%PEG (scale bar: 20 μm). (b) SEMpicture of 5%PEG+5%GelMA (scale bar: 100 μm). (c) SEMpicture of 2.5%PEG+7.5%GelMA (scale bar: 100 μm).

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mechanical properties of GelMA can be greatlyimproved by adding PEG.

The swelling ratio of the hydrogels is an essentialaspect of tissue engineering since it affects variousparameters including surface properties and mobilityand solute diffusion [39]. The swelling degree ofhydrogels depends on the pore size of polymer and theinteraction between the solvent and polymer [40]. Thecalculated mass swelling ratio of PEG–GelMA hydro-gels is shown in figure 2(c). The higher GelMA con-centration was found to increase the swelling ratio.However, compared to the 10 w/v%GelMA hydrogelreported in [34], the average swelling ratio of the threedifferent compositions of hybrid PEG–GelMA hydro-gels in this paper was significantly lower (∼6 versus∼10). A possible reason could be that the pore size ofthe hybrid hydrogel network became much smallerbecause of the lowmolecular weight of PEGhydrogel.

To further investigate the changes to mechanicalproperties, the microstructures of three hydrogelcombinations were analyzed using SEM. The SEMimages of the hydrogel microstructures are presentedin figures 2(d)–(f). The microstructure of 10% PEG isvery dense; no pores larger than 5 μm were observed(figure 2(d)). The dense structure resulted in a highercompressive Young’s modulus. However, the densestructure may generate negative effects in cell cultur-ing because it would be hard for the biomolecules incell media to diffuse into the hydrogel. Conversely, themicrostructures of 5% PEG+5% GelMA and 2.5%PEG+7.5% GelMA contained many pores with dia-meters between 50 and 100 μm (figures 2(e) and (f)).

Although the large pore size results in a decrease ofmechanical properties, it creates better environmentsfor cells to live and spread. The large pore size has beenreported to help cells spread and proliferate [41]. Themicroenvironment of the hybrid PEG–GelMA wasproduced with significantly larger pores to improvethe cell culturing.

3.2. Capability of bioprinting systemA mesh pattern with 500 μm line width and aminiaturized UBC logo were printed to show thecapability of this developed printing system forcustomized printing. Images of the patterns coloredwith food dye are presented in figures 3(a) and (b). Theprinted pattern demonstrates the ability of the bio-printing system to print a complex pattern with highaccuracy. Although the presented mesh patterns had a1.5 cm×1.5 cm area and finished within 16 min(2 min per layer), larger sized patterns of the samethickness can be done at the same time since stereo-lithography can print a whole patterned surface atonce. Therefore, the printing time depends only on thethickness of the pattern. In contrast to the inkjet andextrusion printing based on ‘dot’ printing (i.e. [42]),stereolithography is much faster—especially for pat-terns with a large in-plane area.

In order to check the minimum resolution of thedeveloped system, a mesh pattern with different linewidths was designed. The minimum resolution hadthe line width of one pixel. When the projector wasplaced 10 cm away from the petri dish, a one pixelwidth line generated 50 μmwide lines (figure 3(c)). In

Figure 3.Hydrogel patterns fabricatedwith developed visible light stereolithography. (a)UBC logo; (b) 4×4 mesh (scale bar: 1 cmfor the top and 2mm for the bottom). (c)Theminimumwidth can be fabricated via the developed stereolithography system (scale bar:100 μm). (d)Amesh pattern of a one-pixel vertical line and two- and five-pixel horizontal lines under amicroscope (scale bar:200 μm).

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addition, the width of the pattern is a linear function ofthe number of pixels, as seen in figure 3(d). This rela-tionship provides a guideline for pattern design. Notethat the texture on the surface of the PEGDA resultedfrom the photoinitiated patterning process. Lin et alalso observed a similar surface texture on 4000 DaPEGDAafter patterning [26].

The higher resolution can be achieved by adjustingthe projector lens to focus clearly at a near distance (i.e.5 cm) and crosslinking the hydrogel. Compared toextrusion printing (usually with the resolutionbetween 200 μm and 500 μm [16, 43]), stereo-lithography techniques offer much more accuratemicroscale patterning. Thus, the quality of the printedpattern demonstrated that the developed stereo-lithography-based bioprinting system was able to pro-vide an advanced, high-resolution, rapid printingmethod for hydrogels.

3.3. 3D cell bioprinting experimentsLight directly from the beam projector is not suitablefor printing cells since the lamp generates B band(long-wavelength band) infrared radiation (IR).According to our experiments, at room temperature(∼25 °C), without the help of the water filter, thetemperature of bioinks rose to beyond 40 °C within

15 min. This high temperature caused significant celldeath (figure 4(a)). As reported in [44], long-wave-length IR can be absorbed by the skin, and theabsorbed energy is converted to heat inside the tissue.Because water itself absorbs emitted heat generated bylight, the water filter was used to investigate the cellresponse to strong light sources [45, 46]. For the firsttime, we employed a custom-built water filter in a 3Dbioprinting system to avoid the effect of IR.

To check the feasibility of both proposed systemsand hydrogels, NIH 3T3 fibroblast cells were encapsu-lated by the aforementioned three hydrogel combina-tions and printed by our visible light stereolithographybioprinting system. The total fabrication time for 10%PEG, 5% PEG+5% GelMA, and 2.5% PEG+7.5%GelMA was 6 min, 12 min, and 24 min, respectively.With the increment of GelMA concentration, it takesmore time to crosslink the hydrogel. The live/deadassayed images of cell-encapsulated mesh patterns of5% PEG+5% GelMA with a 500 μm line right afterprinting are presented infigures 4(b) and (c).

From the image, it can be observed that, immedi-ately after printing,most of the cells were alive, and thecell distribution in the pattern was uniform. The500 μm width patterns were sharp and clear. The cellviability results are presented in figure 4(d). The

Figure 4. 3DCell-laden bioprinting experiments. (a)Assayed images of the crosslinked hydrogel without the help of awaterfilter(scale bar: 100 μm). (b)Assayed images with phase contrast backgrounds of T-shape pattern based on 5%PEG+5%GelMAhydrogel on day 0 (scale bar: 200 μm). The sub-graphs at the right and bottom show the projection of thefluorescent image in theXandY directions. (b) 3D reconstructed image of the day 0 T-shape pattern present in (b) (scale bar: 200 μm). (d)Cell viability analysis(*p<0.001, n=10).

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general cell viability right after printing (day 0) wasaround 86%, 83%, and 80% for 10% PEG, 5%PEG+5% GelMA, and 2.5%+PEG 7.5% GelMArespectively, which were higher than most of thereported extrusion systems [17]. These results can beexplained by the fact that stereolithography is a nozzle-free printing technology andwe usually have high cellviability from nozzle-based printing technologies [17].The high cell viability also makes stereolithography-based systems a competitive technology for bioprint-ing. It is noted that Chen et al achieved about 90% via-bility using their UV stereolithography system [23–25]. One possible reason for slightly decreased cell via-bility could be the increment of crosslinking time ofthe visible light bioprinting system that was used inthis paper. The crosslinking time of the UV system is8 s per layer, while the average printing time of the cur-rent system is 4 min per layer. However, as previouslyreported, UV light is a potential source for inductionof DNA damage and cancer [27–29]. It is safer toemploy visible light considering the long-term stabi-lity and functionality of the cell. In comparison to a2500 lumen lamp used in the present projector, 4000lumen or higher power lamps can be used if a shortercrosslinking time is desired. However, the system’scost will triple, and thus lose the advantage of cost-effectiveness.

After culturing for five days, the cell viability of10% PEG interestingly decreased to less than 60%. Apossible reason could be that the pore size of 10%PEGis too small, resulting in the difficulties noted in med-ium diffusion. Thus, the low molecular weightPEGDA is not suitable for long-term cell encapsula-tion and culturing due to its small pore size. Usinghighermolecular weight PEGDAmay improve the cellviability, since higher molecular weight PEGDA canform a relatively loose network with larger pore size ifthe crosslinking time remains the same [47]. Largerpore size can benefit the exchange of biomolecules andthus promote long-term cell viability [48]. However,PEGDA is not a good scaffolding material for tissueregeneration since it is non-degradable. Degradablehydrogels made partly from PEGDA, such as the poly(D,L-lactide)/poly(ethylene glycol)/poly(D,L-lactide)(PDLLA–PEG–PDLLA) [21], can be potentially usedwith eosin Y photoinitiator as visible light cross-linkable bioinks. However, the compatibility betweenPDLLA–PEG–PDLLA hydrogel and eosin Y requiresfurther studies.

The small pore size problem has been greatlyimproved by mixing GelMA with low molecularweight PEGDA, as shown in figure 2. The cell viabilityof 5% PEG+5% GelMA and 2.5% PEG +7.5%GelMA was largely unchanged between day 0 and day5, revealing the long-term biocompatibility of hybridPEG–GelMA hydrogel as bioinks. Therefore, thehybrid hydrogel is more cost-effective for tissue engi-neering applications than high molecular weightPEGDA. Our result is consistent with a previous study

regarding PEG dimethacrylate (PEGDMA)–GelMA byHutson et al [37]. They observed statistically sig-nificant improvements of cell viability after mixingGelMA hydrogel with pure PEGDMA hydrogel. Also,a higher percentage of GelMA had higher long-termcell viability although it was not statistically significant.Therefore, pure GelMA crosslinked by eosin Y photo-initiator will be potentially better for tissue regenera-tion considering the great biocompatibility anddegradability of GelMA [34, 37]. However, 10% w/vGelMAwith the current concentration of eosin Y pho-toinitiator was not crosslinked. We will further studythe feasibility of using various concentrations of eosinY as photoinitiator to crosslink a higher percentage ofGelMA in the near future.

In this study, we found that 5% PEG+5%GelMA hydrogel with eosin Y based photoinitiator is agreat bioink solution for a visible light stereo-lithography 3D bioprinting system considering long-term biocompatibility andmechanical properties. Thedeveloped visible light stereolithography technique,which includes both a low-cost system and hybridhydrogels, is an advanced and highly biocompatiblesolution for high-resolution and rapid 3Dbioprinting.

4. Conclusion

In this paper, we present the design of a novel visiblelight stereolithography 3D bioprinting system and acorresponding visible light crosslinkable bioink madefrom PEG and GelMA. The developed system hasmany advantages over existing systems. In contrast totraditional bioprinting techniques, our system is lowcost, but with a high printing speed and high resolu-tion. Compared to the reported UV and near-UVstereolithography system, the visible light is safer formaintaining long-term cell functionality. Themechanical properties, biocompatibility, and micro-structures of the proposed hydrogel solution werestudied in detail, and the key components in thesystem were illustrated. For the optimized 5%PEG+5% GelMA hydrogel, a Young’s modulus of60 kPa can be achieved with about 6 mass swellingratio and about 50 μmpore size. The developed systemcan reach 50 μm resolution with cell viability around85% both right after printing and after culturing for 5days, providing an excellent low-cost solution forbioprinting applications.

Acknowledgments

This work was supported by a Natural Sciences andEngineering Research Council of Canada (NSERC)Discovery Grant. SG was supported by the CanadianDiabetes Association’s Scholar award and an operatinggrant. The authors thank Dr Deborah Roberts forsharing her facilities.

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