In Vitro and In Vivo Analysis of Visible Light Crosslinkable Gelatin Methacryloyl (GelMA) Hydrogels

Introduction

Hydrogels have been widely used for tissue engineering and regenerative medicine as artificial scaffolds, mimicking the extracellular matrix (ECM) surrounding the cells¹. Their high water content, tunable chemical and physical properties, and ability to encapsulate cells, bio-macromolecules (e.g., peptides/proteins, nucleotides, antibodies), and therapeutic agents open a variety of potential applications^1–4. In particular, biopolymers that allow in-situ sol-gel transition could be used to develop injectable materials, which can be delivered locally in a minimally invasive and cost-effective manner⁵. Injectable hydrogels made of natural or synthetic polymers have been widely reported, and have been shown to exhibit both advantages and drawbacks^{5, 6}. For instance, natural polysaccharide-based hydrogels such as hyaluronic acid, chitosan, alginate, and chondroitin sulfate have been extensively used due to their biocompatibility and biodegradability, but are limited by their low mechanical properties, as well as difficulties in tuning their degradation rate and biological functions^{4, 6}. On the other hand, synthetic polymer-based hydrogels such as poly(ethylene glycol) (PEG) and poly(vinyl alcohol) (PVA) have high mechanical strength with tunable microstructure, but often suffer from the lack of biological moieties^{4, 6}. Thus, there is a need for the development of hydrogels that can combine the advantages of both natural and synthetic polymers⁷.

Light-induced photocrosslinking has been widely used for the fabrication of cell-laden injectable hydrogels with rapid gelation time and tunable physical properties⁸. Ultraviolet (UV) light is commonly used for photocrosslinking of such materials, as well as commercial products for dental applications. We have recently reported the engineering of highly elastic hydrogels via UV light-mediated crosslinking of chemically modified tropoelastin or elastin-like polypeptides for wound healing and cardiac tissue engineering^9–12. These UV crosslinked hydrogels exhibited high cytocompatibility for three-dimensional (3D) encapsulated cells in vitro, as well as a negligible immune responses after subcutaneous implantation in rats¹³. In addition, UV crosslinked gelatin methacryloyl (GelMA)-based hydrogels have been used as 3D cell-laden hydrogels to promote in vitro vascularization and tissue formation^14–16. Although, we have shown that injection of cell-laden GelMA hydrogels in rats led to vascular network formation and tissue healing¹⁷, in situ UV crosslinking of GelMA prepolymer may negatively impact the normal function of sensitive tissues such as the myocardium. The clinical translation of UV crosslinkable hydrogels as injectable materials for soft tissue regeneration has been limited due to the biosafety concerns associated with the use of UV light. UV light may generate reactive oxygen species (ROS), which may induce endogenous oxidative damage to DNA (e. g., dimerization or hydroxylation of bases) and cause accelerated tissue aging or cancer¹⁸. Other studies have reported that oxidative DNA alterations could be induced by UV (i.e., 337 nm)¹⁹, and that DNA damage was dramatically decreased above 470 nm²⁰. Furthermore, recent studies have demonstrated that UV light could directly induce alterations in the DNA, and lead to immunosuppression in vivo^{21, 22}. In another study, Federovich et al. demonstrated that the combination of both UV light and the photoinitiator Irgacure 2959 resulted in harmful effects on cell viability²³. However, it was shown that the photopolymerization process in the presence of hydrogel precursors led to higher cell viability. This observation suggested that free radicals generated during photopolymerization were consumed during the crosslinking process, reducing the damage to cellular components²³. These studies together demonstrated that the development of hydrogels that utilize visible light, with safer photoinitiators as the energy source for the crosslinking reaction, can enhance the utility of photocrosslinkable hydrogels for tissue engineering applications. Additionally, visible light has been shown to penetrate tissues at higher depths and with lower energy, as compared to UV light²⁴. This characteristic makes visible light suitable for developing in situ injectable materials, for minimally invasive procedures.

Eosin Y is a Food and Drug Administration (FDA)-approved photoinitiator excited by visible light (450 – 550 nm) (Figure 1a–b), which was used in the photocrosslinkable lung sealant FocalSeal® (Genzyme Biosurgical, Cambridge, MA). Eosin Y has been used in many thiol-ene crosslinking reactions to fabricate hydrogel-coated surfaces²⁵, mesenchymal stem cell-laden hydrogels²⁶, and growth factor delivery systems²⁷. Other photoinitiators such as camphorquinone, fluorescein, and riboflavin have also been used to initiate photocrosslinking upon exposure to blue light¹⁸. However, the use of prolonged irradiation to obtain hydrogels with appropriate mechanical strength with these photoinitiators may lead to toxic effects on the encapsulated cells, when using them as cell-laden injectable hydrogels²⁸.

Several groups have reported the application of injectable hydrogels for cardiac tissue regeneration. For example, an extracellular matrix (ECM)-derived hydrogel was used as an injectable material for the treatment of infarcted myocardial tissue²⁹. The injection of this hydrogel into infarcted pig hearts led to increased cell infiltration due to tissue self-assembling into the porous scaffold. In addition, histological analysis showed that a layer of endocardium was developed due to injection of the hydrogel²⁹. Other studies reported the use of UV light crosslinkable hydrogels for the repair of anterolateral cardiac incisions. These results showed an increase in wall thickness of the left ventricle, which enhanced the survival of both viable and infarcted cardiac tissues^30–32.

Here, we report the synthesis of a gelatin-based injectable hydrogel for myocardial repair via visible-light-initiated photocrosslinking, using Eosin Y as a photoinitiator, Triethanolamine (TEA) as a co-initiator and N-vinylcaprolactam (VC) as a co-monomer. The GelMA prepolymer solution was photopolymerized upon exposure to visible light to form hydrogels with tunable physical properties. We evaluated the mechanical properties, porosity, and swelling behavior of the engineered hydrogels, as well as their in vitro and in vivo biocompatibility. We also investigated the feasibility of hydrogel delivery to infarcted hearts in vivo. It is hypothesized that the engineered material can be used as a safe and fast polymerizable hydrogel with controlled physical properties. In addition, they could also be used for the repair of a variety of soft tissues such as heart, blood vessels, and skin, and may facilitate the clinical translation of injectable hydrogels for soft tissue regeneration.

Results and Discussion

In this study, we engineered injectable visible light crosslinked GelMA hydrogels for cardiac tissue repair. These hydrogels were formed by the photochemical reaction of methacryloyl units, available on GelMA chains and activated by Eosin Y, which absorbs light between 450 to 550 nm (Figure 1a–b)^{1, 3}. Photocrosslinking via UV light (250 nm < λ < 400 nm) relies on the use of photoinitiators such as Irgacure 2959^{26, 27}. However, alternative photoinitiators such as VA-086, that absorb safer UV ranges (≈405 nm) than Irgacure 2959 and lead to enhanced cytocompatibility have been recently described³³. UV light photons cause dissociation of the photoinitiator into two radicals, which cause radical formation on the methacryloyl groups. Since radicals are unstable entities, they quickly combine with each other in their vicinity, leading to a crosslinked network. An alternative to UV photoinitiation is using Eosin Y in combination with visible light (400 nm < λ < 700 nm). Visible light excites Eosin Y from the ground state into a triplet state. This then extracts hydrogen atoms from amine-functionalized co-initiators, such as TEA. The deprotonated TEA radical then initiates the formation of a radical center on the methacryloyl groups of GelMA. Visible light mediated polymerization often requires co-monomers, such as VC, to generate an adequate number of radicals to accelerate the gelation process (Figure 1)³⁴. This increased gelation rate in the presence of VC could be due to an increased vinyl group concentration, and/or rapid diffusion of radical bearing VC²⁶. In addition, photoinitiation via Eosin Y could be visually confirmed by a change in color from red to yellow, which occurs after the formation of activated Eosin Y (Figure 1c).

It has been reported that in an ideal and fully crosslinked hydrogel network with a certain prepolymer molecular weight and functionality, the crosslinking density of the hydrogel should not be affected by the photoinitiator and co-initiator concentrations³⁵. Therefore, the physical properties of the hydrogel, such as pore characteristics, swelling ratios, and mechanical properties should also remain constant. However, in a visible light system, which is based on free radical chain reaction polymerization, an ideal fully crosslinked hydrogel network cannot be obtained due to the presence of unreacted functionalities, dangling polymers, and primary cycles³⁵. This behavior is particularly observed at lower prepolymer or initiator concentrations³⁶. Therefore, in our system, since the hydrogel network was not fully crosslinked even after 180 sec light exposure time, the physical properties of the engineered hydrogels were dependent on TEA, Eosin Y, and VC concentrations.

Previous studies have investigated the effects of photoinitiator and co-initiator concentrations on the physical properties of visible light crosslinked hydrogels^{25, 26, 36, 37}. For example, Bahney et al. investigated the effect of Eosin Y (0.001–0.1 mM) and TEA (0.01–1.5 %) concentrations on the mechanical properties of visible light crosslinked poly(ethylene glycol) diacrylate (PEGDA) hydrogels³⁷. Their results demonstrated that the dynamic modulus of the resulting PEGDA hydrogels ranged from 25 to 125 kPa at 120 sec light exposure time³⁷. In another study, Shih et al. also showed that by increasing the Eosin Y concentration from 0.1 to 2.0 mM, the swelling ratio of a visible light crosslinked PEG-based hydrogel increased approximately 2 fold at 240 sec light exposure time²⁵. Here, we investigated the effects of these variables on the physical properties (i.e., porosity and swellability) of the engineered GelMA hydrogels (Figure 2). The porosity of a hydrogel is key for its interaction with cells and the surrounding tissues^{2, 38–40}. The modulation of the micro-architectural properties and the porosity of a hydrogel could be used to deliver bioactive cues to cells growing within the engineered constructs^{2, 39}. Additionally, increased porosity promotes cellular penetration and new tissue formation within the 3D structure of the hydrogels². For these experiments, we used a fixed concentration of 0.1 mM Eosin Y. Eosin Y regenerates during light exposure, whereas TEA radicals and VC are consumed by the crosslinking of GelMA, as shown in Figure 1a. Therefore, we anticipated that the swelling ratio and pore size are more dependent on the concentration of TEA and VC than Eosin Y. Our results demonstrated that the apparent pore size in the center area of GelMA hydrogels decreased from 77.1 ± 9.9 μm to 41.0 ± 8.4 μm and 25.6 ± 5.5 μm as the VC concentration increased from 0.5 %(w/v) to 1 %(w/v) and 1.5 %(w/v), respectively, for hydrogels with 10 %(w/v) GelMA, 1 %(w/v) TEA, and 0.1 mM Eosin Y (Figure 2a–c, g). We also investigated the changes in hydrogel porosity by varying the concentration of TEA (Figure S1, 2h). However, we did not observe any statistically significant differences in pore size between different TEA concentrations. Similarly, the pore sizes on the surface of the hydrogels could also be tuned by varying the concentration of VC (Figure 2g). It is important to note that the apparent porosity of the hydrogel could be due to lyophilization of the hydrogels prior to SEM analysis. Previous studies have shown that the pore size of hydrogels could be affected by the freezing temperature before lyophilization². In our study, all samples were processed following a methodology reported previously to compare the average pore size through SEM analysis, using different formulations of UV crosslinked GelMA hydrogels¹⁵. In general, GelMA hydrogels crosslinked via visible light showed a pore size distribution similar to that of UV crosslinked GelMA hydrogels. The effects of VC and TEA concentrations on the maximum swelling ratio of the engineered hydrogels were also investigated (Figure 2i–j). As shown in Figure 2i, the swelling ratio decreased significantly from 8.2 ± 0.36 % to 6.9 ± 0.55 % when the concentration of VC increased from 1.0 %(w/v) to 1.5 %(w/v). In addition, increasing TEA concentrations from 0.5 %(w/v) to 1.0 %(w/v), and 1.5 %(w/v) reduced the swelling ratios from 13.4 ± 1 % to 8.2 ± 0.3 % and 7.1 ± 0.4 %, respectively (Figure 2j). For all these hydrogels, 10 %(w/v) GelMA and 0.1 mM Eosin Y were used. These results suggested that lowering the concentration of the co-initiator and the co-monomer may also lower the crosslinking density and lead to higher swelling capacity.

The mechanical properties of photocrosslinked GelMA hydrogels were also tuned by changing the concentrations of TEA and VC (Figure 3). Our results showed that GelMA hydrogels exhibited tunable compressive moduli in the range of 5 – 56.5 kPa, and tensile moduli ranging from 5 to 22.7 kPa, depending on the VC, TEA, and Eosin Y concentrations (Figure 3). The compressive moduli of 10 %(w/v) GelMA hydrogels, formed by using 1.0 %(w/v) TEA, and 0.1 mM Eosin Y decreased from 26.9 ± 9.9 kPa to 17.2 ± 5.7 kPa and 10.3 ± 5.2 kPa when the VC concentrations were lowered from 1.5 %(w/v) to 1 %(w/v) and 0.5 %(w/v), respectively (Figure 3a–c). In addition, the tensile moduli of the hydrogels decreased from 15.1 ± 4.0 kPa to 12.9 ± 1.7 kPa and 6.3 ± 0.6 kPa as the VC concentrations were lowered from 1.5 %(w/v) to 1 %(w/v) and 0.5 %(w/v), respectively, at a constant 1.0 %(w/v) TEA and 0.1 mM Eosin Y concentration (Figure 3d–f). Increasing VC and TEA concentrations influences the dynamics of the reaction and thus, the number of crosslinking sites may increase at a constant light exposure time (i.e., 180 sec). This enhanced crosslinking density may result in an increase in hydrogel stiffness, as well as a decrease in the swelling ratio and pore sizes³⁴. These results are in agreement with our previous work based on engineering UV light crosslinked GelMA and methacryloyl-substituted tropoelastin (MeTro) hydrogels^{9, 15}. In this study, the mechanical properties of the hydrogels were dependent on the prepolymer and initiator concentrations, as well as light exposure time. We also investigated the effect of varying the concentration of Eosin Y on the mechanical properties of the engineered GelMA hydrogels. Our results showed that the highest compressive modulus was obtained at 0.1 mM Eosin Y for 10 % GelMA hydrogels with 1 %(w/v) VC and 1 %(w/v) TEA (Figure S2). The mechanical properties of the engineered GelMA hydrogels (6–56.5 kPa) in this study were in the range of those reported previously for UV crosslinked GelMA hydrogels (5–50 kPa)¹⁵. In addition, the mechanical strength was comparable to that of native heart tissue (i.e., Young’s modulus: 10–50 kPa)^{41, 42}. These observations suggest that GelMA hydrogels should integrate well with native tissue, without causing failure due to the mechanical mismatch. Based on the mechanical characterization, the optimized hydrogel formulation of 10 %(w/v) GelMA, 1.5 %(w/v) TEA, 1 %(w/v) VC and 0.1 mM Eosin Y was used for all in vitro, ex vivo and in vivo experiments. In particular, the light exposure time (i.e., 180 sec) for hydrogels with 1 mm thickness synthesized for mechanical characterization, ex vivo and in vivo experiments, was lowered to 20 sec for the 150 μm-thick hydrogels used for in vitro experiments.

Engineering cardiac tissue constructs requires a matrix/scaffold for maintenance of cell viability, organization, and tissue formation²⁷. Therefore, we investigated the suitability of the engineered visible light crosslinked GelMA hydrogel as a matrix to support cardiomyocyte (CM) growth. We first investigated the in vitro cytocompatibility of GelMA hydrogels using surface seeding (2D, Figure S3a–f) and 3D encapsulation (Figure S3g–l) of NIH-3T3 fibroblasts. We used a standard live/dead assay to calculate cell viability, by determining the percentage of live cells seeded in 2D (Figure S3c) and 3D (Figure S3i) cultures with respect to total cells. These results demonstrated that cell viability remained >90 % for both 2D and 3D cultures. In addition, fibroblasts grown in both 2D (Figure S3d–f) and 3D (Figure S3j–k) cultures could spread and proliferate for up to 7 days as confirmed by nuclei and F-actin staining. These results demonstrated that visible light crosslinked hydrogels could efficiently support the growth, spread, and proliferation of NIH 3T3 fibroblasts grown in both 2D and 3D cultures. We then evaluated the ability of the engineered GelMA hydrogels to support the growth and contractile function of primary rat CMs in vitro. We have previously shown that UV crosslinked GelMA hydrogels could be used to engineer cardiac tissue constructs^{11, 43}. Although these GelMA-based hydrogels were shown to support the growth of CM in vitro, direct in vivo delivery could potentially impair cardiac function due to tissue damage caused by UV light. Therefore, we aimed to address this limitation by using visible light to form a highly biocompatible cell-laden hydrogel that could preserve the contractile function of CMs. For these experiments, we also used a live/dead assay and DAPI/F-actin staining to evaluate cell viability (Figure 4a–d) and spread (Figure 4e–h) of CMs on the surface of the engineered GelMA hydrogels. Our results demonstrated that the viability (Figure 4d) and the number of cells on the surface of GelMA hydrogels (Figure 4h) increased significantly at day 5, compared to day 1 post seeding. In particular, the cell number on the surface of GelMA hydrogels increased from 92 cells/mm² on day 1 to 205 cells/mm² on day 5 (Figure 4h). However, this observation could be due to the proliferation of cardiac fibroblasts that could not be removed after CM isolation.

Previous studies have demonstrated that CMs maintained in conventional 2D culture conditions tend to regress to a less mature phenotype, and lose the ability to respond to physiological stimuli^{44, 45}. Thus, we also investigated the ability of GelMA hydrogels to maintain the phenotype of primary rat CMs in vitro. First, we evaluated the expression of the CM marker sarcomeric α-actinin (SAA), using 2D cultures of primary CMs on GelMA hydrogels. Our results showed positive immunolocalization of SAA in CMs maintained on the engineered GelMA hydrogels after 7 days of culture (Figure 4i). In addition, fluorescent images showed that cells exhibited well-defined, cross-striated sarcomeric structures, which resembled those of native rat ventricular myocardium⁴⁶ (Figure 4i). Lastly, we evaluated the ability of visible light crosslinked GelMA hydrogels to support the contractile function of primary CMs growing on 2D cultures. These results showed that CMs exhibited synchronous contractions, when cultured on the surface of GelMA hydrogels (Supporting Information, Videos 1, 2). The beating frequency of CMs was quantified up to 10 days after seeding, and was shown to vary from 18–77 beats min⁻¹ (Figure 4j). Furthermore, the metabolic activity of cardiac cells on 2D cultures increased significantly at day 5 post seeding, as shown by a commercial PrestoBlue assay (Figure 4k). Taken together, these results demonstrated that visible light crosslinked GelMA hydrogels could be used to support the growth and preserve the contractile function and phenotype of primary CMs in vitro.

We then performed ex vivo tests to investigate the injectability of GelMA hydrogels, and the ability to photopolymerize the gel with visible light following injection into the myocardium (Figure 5). For this, 200 μL of GelMA prepolymer (10% (w/v)) were injected into the right ventricle of an excised, adult Sprague-Dawley rat heart, followed by 180 sec of exposure to visible light, at a distance of 1 cm. The presence of the photocrosslinked hydrogels within the explanted tissue was verified via immunofluorescent staining (nuclei - DAPI, α-smooth muscle actin (αSMA) - TRITC) of cryosectioned right ventricular tissue. Due to the strong auto-fluorescence of Eosin Y^{47, 48}, the engineered hydrogels appeared as large green fluorescent areas devoid of nuclear staining, which suggested that GelMA was successfully injected into the myocardium and photocrosslinked in situ (Figure 5a–d). In addition to the lack of cellularity, the structure and morphology of the injected GelMA is vastly different from healthy tissue (Figure 5a–d). Furthermore, the positive green fluorescence within the αSMA positive vascular structures suggested that GelMA was crosslinked within the blood vessels (Figure 5c). However, this was mainly because the injection was performed ex vivo, where the absence of blood flow allowed for the infiltration of small amounts of hydrogel precursor into the vasculature. The presence of GelMA hydrogel within the ventricular wall was further confirmed through H&E staining of tissue cryosections following ex vivo injection (Figure 5e–h). Histological assessment revealed acellular regions with increased matrix density, as compared to the surrounding healthy myocardium (Figure 5e–h). These results confirmed that visible light crosslinked GelMA materials can be injected and photocrosslinked in situ, demonstrating their suitability for cardiac tissue engineering applications.

Next, we aimed to investigate the regenerative potential and in vivo injectability of visible light crosslinked GelMA hydrogels, using a myocardial infarction (MI) model. MI was induced in 8-week old Sprague-Dawley rats via permanent ligation of the coronary artery. Next, 50 μL of prepolymer solution (10% (w/v)) were injected into 4 areas (200 μL total) within the infarcted region of the left ventricle, followed by photocrosslinking by exposure to visible light for 180 sec, at a distance of 1 cm (Figure 6a,b). Animals receiving saline injections (sham) were used as controls (Figure 6c,e). The efficient delivery of GelMA hydrogels in vivo following MI demonstrated the feasibility for injection and in situ gelation during open-heart surgeries (Figure 6d,f). Crosslinked GelMA was visible from the epicardial surface (Figure 6d) and was also visible as a change in ECM density in Masson’s trichrome images of hearts isolated immediately post-injection (Figure 6f). Animals were sacrificed 3 weeks after MI, and cardiac tissues were harvested and processed for histological assessment. As expected, the scar tissue on the free wall from the left ventricles of animals with MI became thin and collagenous, with little histological differences among the infarcted groups, as demonstrated by Masson’s trichrome staining (Figure 6g–i).

Differences were apparent with GelMA treatment in the immunofluorescent staining of the infarct region of the hearts. Due to the permanent ligation of the left coronary artery, significant decreases in cardiac α-actin (CAA)⁺ CMs were observed in the left ventricle of the sham and GelMA treated animals after 3 weeks (Figure 6 j–m). However, low numbers of CAA⁺ CMs and von Willebrand Factor (vWF)⁺ endothelial cells were identified on the inner region of the scar area in GelMA treated animals (Figure 6l,m). This observation suggested that the GelMA injection may alter the rate of remodeling of the damaged tissue, given that the sham did not show any positive staining for SAA. In addition, positive staining for vWF was only observed in the endothelium lining the inner ventricular surface (Figure 6k). Typically, the infarct region becomes inundated with activated cardiac fibroblasts (also called myofibroblasts), which are defined by expression of αSMA. The reduction in αSMA⁺ cells within the GelMA treated scar region, as compared to the sham scar region, suggested a reduction in myofibroblast activation, 3 weeks after MI (Figure 6 n,o). αSMA expression in GelMA treated tissues more closely resembles that of healthy tissues, with no positive αSMA staining outside of the vessel linings. In addition, our results also showed the maintenance of larger vessels normally present in healthy tissue, as opposed to smaller neo-vessels present in the remodeled infarcted region in control (sham) animals (Figure 6 n–p). Therefore, the delivery of GelMA hydrogels appeared to enhance CM viability, which might be due to the modulation of the remodeling response of the tissue (evidenced by reduced αSMA+ fibroblasts), as well as the maintenance of larger vessels in the infarcted region (Figure 6 p). However, further studies are necessary to fully determine the exact nature of these effects. Previous studies have demonstrated that upon MI, cardiac stem cells (CSCs) migrate to the site of injury and participate in the repair of the injured tissue^{49, 50}. In addition, we have also demonstrated previously that UV crosslinked GelMA-based hydrogels could support capillary formation in vivo after implantation¹⁴ and after injection⁵¹. Therefore, GelMA hydrogels could potentially be used to support homing and attachment of migratory CSCs and to promote vascularization in vivo, which would potentially promote regeneration after MI.

In summary, we demonstrated the ability to photocrosslink GelMA hydrogels in situ using visible light, following intra myocardial delivery. In vivo studies following MI confirmed that the delivery of photocrosslinkable GelMA hydrogels appears to be safe and feasible. However, more comprehensive experiments should be undertaken to optimize the delivery of the hydrogels and to assess the potential benefits of GelMA hydrogels for improving cardiac function after MI.

Conclusion

Here, we developed visible light crosslinked GelMA hydrogels with tunable physical properties. The use of visible light may help circumvent the potential biosafety concerns associated with the use of toxic initiators or UV light. The physical properties of the engineered hydrogels (i.e., mechanics, swelling behavior, and porosity) could be modulated by varying the concentrations of the photoinitiators in the hydrogel precursor solution. For example, the compressive modulus could be tuned in the range of 5–56.5 kPa, making it a suitable candidate for soft tissue engineering applications. In addition, the engineered hydrogels were shown to be cytocompatible, both in vitro and in vivo. The safe delivery and in situ polymerization of visible light crosslinked GelMA hydrogels make them remarkably suitable for tissue engineering applications and lithography-based platforms, where the use of UV light may potentially affect biological functions.

Supplementary Material