Development and Systematic Characterization of GelMA/Alginate/PEGDMA/Xanthan Gum Hydrogel Bioink System for Extrusion Bioprinting

Abstract

Gelatin methacryloyl (GelMA)/alginate-based hydrogels have shown great promise in bioprinting, but their printability is limited at room temperature. In this paper, we present our development of a room temperature printable hydrogel bioink by introducing polyethylene glycol dimethacrylate (PEGDMA) and xanthan gum into the GelMA/alginate system. The inclusion of PEGDMA facilitates tuning of the hydrogel’s mechanical property, while xanthan gum improves the viscosity of the hydrogel system and allows easy extrusion at room temperature. To fine-tune the mechanical and degradation properties, methacrylated xanthan gum was synthesized and chemically crosslinked to the system. We systematically characterized this hydrogel with attention to printability, strut size, mechanical property, degradation and cytocompatibility, and achieved a broad range of compression modulus (~10 to 100 kPa) and degradation profile (100% degradation by 24 hrs to 40% by 2 weeks). Moreover, xanthan gum demonstrated solubility in ionic solutions such as cell culture medium, which is essential for biocompatibility. Live/dead staining showed that cell viability in the printed hydrogels was over 90% for 7 days. Metabolic activity analysis demonstrated excellent cell proliferation and survival within 4 weeks of incubation. In summary, the newly developed hydrogel system has demonstrated distinct features including extrusion printability, widely tunable mechanical property and degradation, ionic solubility, and cytocompatibility. It offers great flexibility in bioprinting and tissue engineering.

Introduction

Three decades ago tissue engineering research emerged as a means to synthesize tissue for organ replacement by combining extracellular matrix (ECM)-like biomaterials, cells, and cell signaling molecules (e.g.growth factors) to address the shortage of organ transplants. Despite considerable progress in the field, the gap between the number of patients on the waiting list for transplantation and the number of available transplants has grown larger over the years [1]. There are recent examples of less vascularized tissues being produced, such as cornea [2] and cartilage [3], that are well-established and show promising clinical applications. For example, the FDA has approved a hydrogel corneal inlay for improved vision in 2016 [4]. In the same year, a cartilage treatment procedure, referred to as Matrix-Associated Autologous Chondrocyte Implantation (MACI), was also approved by the FDA [5]. However, the engineering of other tissues, such as heart, liver, kidney, or musculoskeletal tissues that have multiple cell types, dense vasculature, and complex structures has yet to be successful [6, 7].

3D bioprinting [8–10] is a powerful tool for engineering implantable grafts by enabling spatially biomimetic presentation of ECM-like biomaterials, growth factors, cells, and embedded vasculature. Advancements have led to replicating tissue-like functions as an alternative to tissues and organ transplants [11, 12]. In particular, for syringe-based microextrusion bioprinting, bioinks - or extrudable ECM-like biomaterials or scaffolds - are a key factor for achieving the biomimetic structural complexity and functionality of native tissue alternatives. These bioinks not only require desired mechanical properties but also proper rheological and biological properties [13, 14]. Among them, Gelatin-Methacryloyl (GelMA)-based bioink has demonstrated great potential in providing a favorable environment for cells [15–19]. However, pure GelMA ink lacks structural stability and does not maintain its shape immediately following printing. Therefore, it needs to be augmented with other components to improve its mechanical property and printability [20, 21]. Alginate on the other hand, is a natural linear polysaccharide that has been widely used as a bioink component due to its fast gelation property, which leads to better extrusion and structural integrity post-printing [22–25]. For example, J. Jia et al. have bioprinted alginate onto a calcium substrate to achieve a lattice structure [26], while S. Hong et al. have developed a highly stretchable printable hydrogel using alginate combined with poly(ethylene glycol) diacrylate (PEGDA) [27]. Additionally, injectable alginate hydrogels with shear-thinning properties were formulated using divalent cation salts of low solubility (e.g. CaSO₄) [28, 29]. However, natural alginate lacks binding sites for cell attachment and needs to be used in combination with other materials. Therefore, alginate and GelMA have been combined to improve biocompatibility and mechanical property simultaneously [13]. They can be extruded in situ in a core-shell pattern with coaxial needles [30–32] or mixed in various ratios as a bioink blend before extrusion [33–35]. Despite numerous advancements in the fabrication and application of GelMA/alginate-based bioinks, the ideal hydrogel composition that takes into consideration printability, structural integrity, and cell compatibility is yet to be formulated. Moreover, the composition may also need further tuning to accommodate different applications.

In this study we have developed a novel GelMA/alginate-based hydrogel system by incorporating polyethylene glycol dimethacrylate (PEGDMA) and xanthan gum to improve extrusion printability and to control mechanical properties. Like PEGDA, PEGDMA is a linear polymer and has been frequently used in combination with other bioinks to improve their mechanical property [36]. Increased viscosity, often tuned by viscosity enhancers, is an important parameter of bioink that enhances printability and printing resolution.[37] However, it also results in an increase in shear force, causing more cellular injury and death [38, 39]. Among the viscosity enhancers, Laponite nanosilica has been widely used due to its inert characteristic and shear-thinning property, which are favorable for bioprinting [40, 41]. As an alternative, we have investigated using xanthan gum, which is a polysaccharide widely used in the food industry [42] as a viscosity enhancer. In addition to providing the shear-thinning property [43, 44], we observed other distinct features: bioinertness, tunable mechanical property, tunable degradation profile, and solubility in ionic liquids such as culture medium, PBS or Ca²⁺ solutions. Solubility in biological buffers or medium is important for biological applications [45]; therefore, such a feature makes xanthan gum a promising additive for applications involving cells or gels with Ca²⁺ for alginate crosslinking. With the addition of xanthan gum, we achieved printability at room temperature, in contrast to many other GelMA/alginate-based bioinks that require specific temperature control for printing [46, 47]. Moreover, we systematically characterized our hydrogel with respect to its printability, printing resolution, compressive modulus, degradation, and cell viability and proliferation. Despite establishing biocompatibility, the xanthan gum-incorporated hydrogel system presented the weaker compressive modulus. We, therefore, further synthesized methacrylated xanthan gum and incorporated it into an alginate/GelMA/PEGDMA hydrogel to investigate the effect of chemically crosslinked gum on the compressive modulus of a printed hydrogel. Representative 3D geometries were printed with the newly developed hydrogel for demonstration purposes. We believe the improved hydrogel system will be a promising candidate for bioinks for printing complex tissues and that this systematic characterization study will provide a roadmap for the development of future GelMA/alginate-based hydrogels.

Results and Discussions:

Hydrogel Extrusional Printability and Print Quality

We combined GelMA, alginate, and PEGDMA with viscosity enhancers for optimal printability. Specifically, we used xanthan gum as a novel viscosity enhancer, which can be further crosslinked into the GelMA/alginate/PEGDMA hydrogel system by methacrylation. Figure 1 shows the schematic of the methacrylated xanthan gum (M-Gum) incorporated hydrogel system. The system is comprised of a dual crosslinking mechanism where alginate is crosslinked by Ca²⁺ while other components including GelMA, PEGDMA, and M-Gum are crosslinked by methacrylate groups. Each component has its own specific functional role: GelMA supports cell attachment and growth; alginate provides structural integrity; PEGDMA improves mechanical property; and xanthan gum improves printability. Good printability and print quality ensure the delivery of the as-designed bio-scaffold. To achieve this, we added xanthan gum or Laponite to the hydrogel system as a viscosity enhancer. These viscosity enhancers also contribute to the shear-thinning property of the hydrogel system, which allows the hydrogel to be extrudable under minimum pressure while maintaining its structural integrity after being extruded. To characterize the printability, we printed hydrogel samples with 3% xanthan gum using various printing pressures and nozzle heights. Here, nozzle height refers to the distance between the nozzle and the substrate. It is worth mentioning that despite the tremendous amount of study on extrusion-based printing, the quantitative definition of printability is still ambiguous, and different quantitative measurements have been used to describe printability [48, 49]. In this work, we specifically used separation distances (D_s), reported by Y. He et al., as the quantitative representation of printability [50]. D_s and strut sizes (Figure 2). Representative measurements of D_s are included in supplementary Figure S1 to demonstrate the measurement protocol. D_s (Fig. 2a) is mainly dependent upon air pressure and is a commonly used benchmark to quantify printability [50], with D_s between 5–30 mm accepted as optimal for printing. Therefore, we can infer that a pressure under 10 psi is optimal to achieve such D_s. In following experiments, D_s would be recalibrated due to changes in material composition. Fig. 2b shows the measurement of strut size relative to the nozzle height over the substrate. When the nozzle was initially in close proximity to the substrate, the hydrogel extruded against itself resulting in a thicker strut. As the nozzle was raised, the strut size decreased and then stabilized. As the nozzle height increased, the hydrogel being extruded from the nozzle broke contact from the substrate. Results from Figure 2 provide a range of printing pressures and nozzle heights for optimal printability. The printability was good at room temperature due to the addition of xanthan gum even though GelMA requires specific temperature control to be properly printed [46, 47].

Figure 1.

Schematic of the GelMA/alginate hydrogel system incorporated with methacrylated xantham gum and PEGDMA. Alginate forms physically crosslinked networks via calcium ions, while GelMA, PEGDMA, and methacrylated xanthan gum (M-Gum) form the chemically crosslinked networks via methacrylate group. Non-methacrylated xanthan gum becomes physically trapped within the networks formed by other physically or chemically crosslinked components.

Figure 2.

Printability characterization of the hydrogel system. (a) Characterization of separation distance (D_s) under different pressures (nozzle height fixed at 0.4 mm). (b) Characterization of strut size at different nozzle heights (pressure fixed at 10 psi). For the printability experiment 2.5% alginate + 10% PEGDMA + 3% xanthan gum was used. (c, d) Rheological measurements for hydrogels with 10% GelMA + 2% PEGDMA + 1.25% Alginate + 3% xanthan gum / 3%Laponite / none (control).

To futher quantify the effects of viscosity enhancers on the printability, we also performed rheological measurements for the hydrogel system, including storage and loss modulus, and viscosity measurements for hydrogels with xanthan gum or Laponite, or no additive as our control (Figure 2c–d). The storage and loss moduli are 1080.6±7.7 Pa and 245.6±3.5 Pa for xanthan gum, 2534.6±22.1 Pa and 672.9±3.9 Pa for Laponite, 115.1±2.2 Pa and 65.7±2.5 Pa for control groups, respectively. Meanwhile, as the shear rate increased from 0.3 1/s to 12 1/s, the viscosity of xanthan gum, Laponite and control group reduced from 1330.9, 2534.9, and 157.6 Pa·s to 44.6, 78.4 and 20.1 Pa·s, respectively. Both storage and loss moduli, and the viscosity measurements showed shear-thinning property for both xanthan gum and Laponite groups. Very low modulus and little shear-thinning property were observed in the control group, which resulted in poor printability. Therefore, we weren’t able to achieve quality prints with the control group (Figure S2). This indicates the addition of viscosity enhancers did significantly improve the bioink printability. Notably, the modulus and shear-thinning property of the Laponite group is higher than that of the xanthan gum group, which correlates with our mechanical measurement. We also quantified the printability following approaches described in a review article [49]. Specifically, we first evaluated the Pr value, defined by Ouyang et al. [51], for a scaffold printed as shown in Figure 3a. As a result, the Pr value was 0.966 ± 0.071, demonstrating improved printability compared to Pr value of 1 in ideal condition. Second, we printed 1 cm square scaffolds with different heights of 2 and 5 mm and measured the tangent value of the side angle of the scaffold as a means of evaluating the structural integrity, following Gao et al.’s approach [52]. As a result, the angles for scaffolds with a height of 2 mm and 5 mm were 83.3° and 86.1° (comparing to ideal value of 90°), resulting in tangent values of 8.52 and 14.67, respectively. Both the measured values and pictures (Figure S3) showed structural integrity was maintained even when stacking multiple layers of the hydrogel.

Figure 3.

Printed demonstrations with 10% GelMA, 2% PEGDMA, 1.25% Alginate and 3% xanthan gum (a) a mesh structure and (b)(c) an anatomical ear using the developed hydrogel system. The side view of the ear in (c) shows the capability of printing hanging structures using the hydrogel. Scale bar: (a) 2 mm, (b) 1 cm.

Print resolution is a key criteria for evaluating print quality. To achieve optimal print resolution, we performed a systematic characterization for hydrogels with different PEGDMA and GelMA compositions (10% PEGDMA, 5% GelMA + 2% PEGDMA or 10% GelMA + 2% PEGDMA), each with varying concentrations of sodium alginate (1.25% or 2.5%) and viscosity enhancers (xanthan gum or Laponite, 3% or 5%). For each condition, we also varied the driving pressure (2–15 psi) and printer head traveling speed (10–30 mm/s), which culminated in 16 data sets presented in Figure S4–S7. The characterization is summarized in Table 1. The two values in each element of the table represent pressure range (unit: psi) and strut size range (unit: μm), respectively. We measured the strut sizes with respect to varying material compositions, nozzle heights, and printing pressure. As a result, we report a range as opposed to a single value for the sizes; however, the minimum value can be viewed as the representative resolution. From the data, an increase in concentration of any component correlates with an increase in required printing pressure due to the higher viscosity. For printing resolution, we can see that as the viscosity enhancers’ concentrations increase, the minimum strut size decreases. Meanwhile, alginate and GelMA concentrations have little effect on the minimum strut size. There is no significant difference between the xanthan gum and Laponite groups in terms of strut size and required pressure. With the current experimental setup, the minimum strut width achieved was 300 um, which is within the upper limit necessary to supply nutrients to cells in large latticed porous structures [53]. The minimal printable strut size is also comparable with existing work which requires temperature control [47]. Overall, we believe that such systematic characterization provides a comprehensive view of how each parameter affects the print quality, thus paving the way for research on hydrogel development and bioprinting.

Table 1.

Characterization of the printing resolution, represented by the strut sizes at varying pressures and hydrogel compositions. In each element in the table, the values represent the pressure range (unit: psi) and strut size range (unit: μm).

Pressure (psi) Strut size (μm)		Viscosity Enhancers
		3% Gum	5% Gum	3% Laponite	5% Laponite
Material Compositions	1.25% Alginate 10% PEGDMA	2.5–3.5 490–1030	9–11 300–750	3–5 440–1320	4–6 370–940
	2.5% Alginate 10% PEGDMA	8–10 510–820	12–14 400–720	5–7 500–1130	5–7 460–960
	1.25% Alginate 5% GelMA 2% PEGDMA	4–6 490–910	7–9 380–870	6–8 420–960	6–8 400–1060
	1.25% Alginate 10% GelMA 2% PEGDMA	10–12 480–930	12–14 350–790	10–12 390–650	13–15 300–770

To further demonstrate the printability, print quality, and the structural integrity of the hydrogel system, we printed a mesh structure (a commonly used design for scaffolds) and an anatomical human ear (as a representation of native tissues) with our hydrogel containing 10% GelMA, 2% PEGDMA, 1.25% Alginate and 3% xanthan gum (Figure 3). In Figure 3a, the printed mesh structure contains 6 layers of crosshatch patterns with alternating horizontal and vertical patterns. The overlapping layers demonstrate the precision and structural integrity of the hydrogel. In Figure 3b, the printed anatomical ear was able to be held and manipulated by a tweezer, demonstrating its structural integrity. From Figure 3c, the side view shows that during printing, the upper part of the ear was printed as a non-supported, overhanging structure, which again demonstrates structural integrity and printability of our hydrogel system. A video of the printing process of the anatomical ear is also provided in supplemental information Video SV1.

Mechanical Property

To be utilized in surgical implantation, the hydrogel needs to possess proper mechanical property, both during the printing process and after it has been printed. To be an implantable material, the hydrogel needs to match the diverse mechanical properties of native tissues [54] and simulate the surrounding matrix characteristics essential for stem cell differentiation [55]. Throughout printing, the hydrogel needs to maintain structural integrity to support its own weight. Post-print crosslinking further improves the mechanical property to achieve the desired modulus within the range of native tissues. We printed a series of samples with varying hydrogel compositions (5% or 10% GelMA, 2%, 3% or 5% xanthan gum/Laponite, 0% or 2% PEGDMA), and different porosities (50% or 0%, 0% being completely filled), at a size of 1 cm × 1 cm × 1 mm. These concentrations were selected based on prior studies [56, 57]. Particularly for the viscosity enhancers, our preliminary study showed that much lower or much higher concentrations led to poor printatbility and therefore was not included in the main study. The compressive modulus of each sample was then characterized via a compression test. We have summarized all measurements in Table 2, from which the addition of Laponite could increase the compressive modulus, as reported in other literature [58]. Notably, an increase in uncrosslinked xanthan gum concentration decreased the compressive modulus of the hydrogel, which may be due to its inability to bind covalently to the hydrogel network and consequently not contribute to network elasticity. To improve the compressive modulus, we synthesized methacrylated xanthan gum and incorporated it into the alginate/GelMA/PEGDMA hydrogel to investigate the effect of its chemical crosslinking on the compressive modulus of the printed hydrogel. The ninhydrin test showed that the molar ratio of incorporated methacrylate groups to the total number of gum repeating units after the methacrylation reaction was 1 / 3.2. As a result, we have improved the compression modulus of hydrogel with 10% GelMA, 3% gum from 51.25 kPa to 87.63 kPa (Table 2 and Figure S8). Moreover, to expand the compressive modulus lower limits, we also printed and measured the compression modulus for a composition without PEGDMA, as PEGDMA is expected to play an important role in determining the compressive modulus. The resulting modulus without PEGDMA was reduced from 51.25 kPa to 18.22 kPa. Combining PEGDMA with the gum’s inert property and the new crosslinking mechanism, we can tune the compressive modulus to a desired range according to the targeted tissue. In general the measured compressive modulus of hydrogels with Laponite ranged from 100–300 kPa, which is comparable to existing literature [47], while the strength of hydrogels with xanthan gum ranged lower at 10–100 kPa. The use of xanthan gum would extend the applicable compressive modulus range and would be more appropriate for softer tissues such as live, kidney and skin, whose modulus range from 1–100 kPa [59–61]. Notably, while hydrogels with Laponite exhibit higher mechanical properties, they are still not comparable to rigid tissues such as bone, tendon, etc. whose moduli range from hundreds of MPa to GPa [61, 62]. In such cases, rigid polymeric scaffolds may be used in combination with the soft hydrogels. We also compared the compressive modulus range of our proposed hydrogel system with existing work [19, 63–67] (Figure S9). Results showed that our hydrogel system can be tuned with a broad range of modulus which is suitable for broader applications. Lastly, from Table 2, compressive modulus was affected by changes in porosity, but remained unaffected by changes in GelMA concentration.

Table 2.

Compressive modulus of scaffolds printed with the proposed hydrogel with different compositions. Except for the group without PEGDMA, all other groups have 2% PEGDMA added to the hydrogel system.

Compressive modulus (kPa)		Viscosity Enhancers
		2% Gum	3% Gum	5% Gum	2% Laponite	3% Laponite	5% Laponite	3% M-Gum
Material Compositions	10% GelMA 0 % porosity No PEGDMA	28.25 ± 5.83	18.22 ± 0.78	10.92 ± 2.28	8.63 ± 1.12	29.42 ± 4.25	67.49 ± 7.96	N/A
	5% GelMA 50% porosity	30.66 ± 3.22	18.2 ± 0.14	14.5 ± 3.25	10.18 ± 2.01	23.7 ± 4.95	53.5 ± 14.0	N/A
	5% GelMA 0% porosity	90.73 ± 10.77	53.35 ± 7.42	24.5 ± 3.39	121.19 ± 2.98	156.1 ± 1.41	306.55 ± 1.34	N/A
	10% GelMA 50% porosity	17.93 ± 5.64	25.95 ± 0.21	9.80 ± 0.14	21.22 ± 8.61	37.65 ± 6.72	83.2 ± 6.93	N/A
	10% GelMA 0% porosity	129.51 ± 18.15	51.25 ± 7.14	38.45 ± 6.58	145.69 ± 13.61	153.45 ± 8.70	360.75 ± 69.93	87.63 ± 5.25

Degradation

As implants, the hydrogel scaffold needs to possess an appropriately tuned degradation rate to allow for cell ingrowth and vascularization in different applications. To characterize the degradation of our proposed hydrogel system, we performed a degradation study for hydrogels with different concentrations of xanthan gum (methacrylated or unmethacrylated) or Laponite. We tested degradation in collagenase solution and measured the remaining mass over time as shown in Figure 4. Hydrogels without the addition of PEGDMA showed the fastest degradation. For groups with PEGDMA, hydrogels with unmodified gum degraded completely (i.e. no bulk material could be collected, see supplementary Figure S10) within a week in collagenase solution. Hydrogels with M-Gum degraded more slowly, with more than 80% of the hydrogel degrading over two weeks. On the other hand, about 40%−50% of hydrogels with Laponite degraded in collagenase over the two weeks. Higher concentrations of Laponite reduced the degradation rate. They were much slower compared to original GelMA/alginate/PEGDMA systems reported by literature [20, 67, 68], while the addition of non-modified xanthan gum didn’t increase degradation time. This further shows that the non-modified xanthan gum did not contribute to the crosslinking of the hydrogel network, whereas methacrylation of xanthan gum allowed for crosslinking into the system. These results are also consistent with the compressive modulus characterization showing that higher compressive modulus of the hydrogel results in slower degradation.

Figure 4.

Degradation profile of printed hydrogels in 1μg/mL collagenase solution by measuring the dry weight after submerging the samples in the solutions. The weights are normalized to the original weights after samples are printed. For this degradation experiment GelMA/alginate concentration was kept at 10% / 1.25%.

Cell Viability and Proliferation

Cell viability and proliferation are among the most important features of bioinks. To verify that our improved hydrogel system is cell compatible, we encapsulated hMSCs in three different groups: 0% PEGDMA (10% GelMA + 1.25 % alginate + 0% PEGDMA + 3% gum), Gum (10% GelMA + 1.25 % alginate + 2% PEGDMA + 3% gum), M-Gum (10% GelMA + 1.25 % alginate + 2% PEGDMA + 3% crosslinked gum). Figure 5 shows the viability test results of hydrogels with xanthan gum for one week, including representative fluorescent pictures for the group without PEGDMA (Figure 5a–d) and the quantified results from MATLAB. Note that the live/dead images have a slight green background, which is possibly due to the absorption of dye in the material. During image processing we increased the contrast to eliminate this background effect. It can be seen that the three groups of hydrogels have similar cell viability. All groups showed decreased viability at Day 1 (88.9%−98.7% viability at Day 0, 74.4%−88.7% at Day 1), while the overall viability after Day 4 of the hydrogel system is over 90% (Figure 5e, 96.1%−98.6% at Day 4, 94.7%−99.8% at Day 7), demonstrating cytocompatibility of the proposed hydrogel with gum. Therefore, we conclude that the high overall polymer content didn’t affect cell viability. We have also tested the hydrogel system with Laponite using the same protocol, substituting the Laponite solution with water as a medium. This was due to Laponite’s insolubility in culture medium or PBS. This rendered all cells unstainable (Figure S11), which we attribute to the suboptimal osmotic pressure of the solution that lysed the cells. Such results suggest that under a simple one-step dissolution process, xanthan gum, as a novel viscosity enhancer, may provide better biocompatibility with its ability to encapsulate the cells. Alternative approaches for cell encapsulation using Laponite may involve dissolving a high concentration of Laponite in milli-Q water and diluting with a high concentration PBS. We were able to achieve viable cells in the following proliferation experiment for Laponite group with this approach. However, this adds complexity to the hydrogel preparation.

Figure 5.

Viability results for hMSCs in hydrogel system at (a) Day 0, (b) Day 1, (c) Day 4 and (d) Day 7, for the group with 0% PEGDMA + 3% gum, (e) quantified viability for all three groups, including 0% PEGDMA + 3% gum, 2% PEGDMA + 3% gum, and 2% PEGDMA + 3% M-gum, (f) MTS assay results for both xanthan gum group (2% PEGDMA + 3% gum) and Laponite group (2% PEGDMA + 3% Laponite). Scale bar: 100 μm. For all groups GelMA/alginate concentration was kept at 10%/1.25%.

We also performed a 4-week proliferation study with MTS metabolic assay (Figure 5f) for both the xanthan gum and Laponite groups. The results show a significant increase in cell metabolic activity from Day 0 to Day 3, which indicates cell proliferation. This increase is also comparable to existing literature using GelMA/Laponite-based bioink [69]. Xanthan gum groups provided slightly better cell proliferation compared to Laponite groups, which is likely due to better degradability of the xanthan gum group. The metabolic activity increase stabilized over the longer term of more than one week, which is possibly due to the condensed gel not leaving much room for cells to spread out. In future applications where cell enrichment is needed, we may consider reducing the overall gel content, which sacrifices some mechanical property but allows the cells to proliferate more.

Conclusion:

In this work we developed and characterized a new GelMA/alginate/PEGDMA/xanthan gum-based hydrogel system for bioprinting by incorporating xanthan gum as a bioink additive, and the resultant product has demonstrated printability and biocompatibility. We designed the hydrogel system with each component serving a specific role: GelMA for biocompatibility, shear-thinning alginate for structural integrity, PEGDMA for tunable mechanical property, and xanthan gum as viscosity enhancers for improving printability. In addition, we have performed systematic characterizations of the hydrogel system documenting printability, strut size, compressive modulus, degradation and cytocompatibility, and have tuned its composition and crosslinking mechanism to achieve a broad range of compressive modulus and degradation profile. The newly developed GelMA/alginate-based hydrogel in this study offers several distinct features including: (1) xanthan gum as an additive acts as a viscosity enhancer with a shear-thinning property which improved the printability, allowing for direct extrusion printing of GelMA hydrogel at room temperature. (2) Crosslinking the gum into the hydrogel system with methacrylic modification improves and controls the mechanical property of the printed hydrogel. By tuning the crosslinking and hydrogel composition, we were also able to control the degradation profile, which can be optimized for in vivo experiments. (3) The gum-modified hydrogel provides excellent biocompatibility, as shown by both the viability and proliferation studies. The solubility of gum in ionic solutions like cell culture medium allows for cell encapsulation with culture medium as the hydrogel solvent, which ensures cell viability both during and post printing. Moreover, as Laponite is a composite nanomaterial, we believe xanthan gum would generally be safer to use as a general polysaccharide. For example, Laponite is known to induce osteogenic differentiation for MSCs [70], possibly due to the presence of positive ions in Laponite [71], which could be unwanted in applications for other tissues. We believe the developed hydrogel has potential as a bioprinting material and that the systematic characterization of its property and printing conditions will establish the framework for future hydrogel bioprinting research studies. Future plans of this work could involve further increasing the mechanical property range by introducing components with different molecular weight, as well as further validate the biocompatibility and biofunctionality of the material.

Materials and Methods:

Materials

Laponite was provided by courtesy of BYK Inc. (Germany). Xanthan gum, Methacrylic anhydride, gelatin (Type A, gel strength 300), N-(3-Dimethylaminopropyl)-N′-ethylcarbodiimide hydrochloride (EDC), N-Hydroxysulfosuccinimide (NHS), 2-(N-Morpholino)ethanesulfonic acid (MES), and N-(3-Aminopropyl)methacrylamide hydrochloride (APMA) were purchased from Sigma-Aldrich (St. Louis, MO). PEGDMA 1k was purchased from Polysciences Inc. (Warrington, PA)

Synthesis of GelMA and methacrylated Xanthan gum

To synthesize GelMA, gelatin was dissolved in DI water (10% w/v) at 50 °C. Methacrylic anhydride was added to gelatin solution at a molar ratio of 100:1 (methacrylic anhydride: gelatin), and the solution was allowed to react under stirring for 1 hr at 50 °C [72]. The mixture was then diluted 5X with DI water and dialyzed against DI water using a dialysis tube (Spectrum Laboratories, Rancho Dominquez, CA) with 6–8 kDa molecular weight cutoff for 3 days at 40 °C. The solution was then freeze-dried and stored at −80 °C. The degree of substitution of resulted GelMA was 73.2% [17]. To synthesize methacrylated xanthan gum, 1 gr gum was dissolved in 100 mL MES buffer (100 mM). 5 mL MES buffer containing 50 mg EDC and 50 mg NHS was then added to the gum solution. After 1 hr reaction at room temperature, 75 mg APMA in 1mL MES was added to the gum solution and allowed to react for 2 hr at room temperature. The solution was then dialyzed against DI water using a dialysis tube (Spectrum Laboratories, Rancho Dominquez, CA) with 6–8 kDa molecular weight cutoff for 3 days at ambient temperature, lyophilized, and stored at −80°C.

To evaluate the methacrylation reaction yield of the xanthan gum, the concentration of unreacted APMA after the reaction was measured using ninhydrin assay as described elsewhere [73]. Briefly, ninhydrin (Sigma Aldrich) was dissolved in ethanol to make a 2% (wt/v) ninhydrin reagent. Then 40 μL of the ninhydrin reagent was added to 200 μL of the APMA solution after the reaction with gum. After mixing, the solution was heated to 90°C in a capped tube for 8 min and the absorbance was read at 570 nm using a SpectraMax M2 plate reader (Molecular Devices LLC). The concentration of unreacted APMA in the solution was calculated using a calibration curve made for the absorbance of solutions with known concentrations of APMA.

Hydrogel preparation and crosslinking

6% or 10% viscosity enhancer solution was prepared by dissolving Laponite or xanthan gum in DI water. Separately, the hydrogel solution was prepared by step-by-step dissolving of 2% 2-hydroxy-4’-(2-hydroxyethoxy)-2-methylpropiophenone (Igracure 2959), 20% GelMA, 4% PEGDMA, and 2.5% or 5% alginate in DI water. The viscosity enhancer solution and the hydrogel solution were then degassed and mixed at a 1:1 ratio. At the last step, CaSO₄ slurry was added to the mixture at a final concentration of 1mg/mL, forming a partially crosslinked hydrogel system. For the strut size characterization, besides using 20% GelMA and 4% PEGDMA, 10% GelMA and 4% PEGDMA or 20% PEGDMA were also used for the hydrogel system. After printing, hydrogel samples were crosslinked under 4 mW/cm² UV lamp at 365 nm wavelength for 2 min. The samples were then fully-crosslinked in 1% CaCl₂ solution for 2 min.

Hydrogel printing and characterization

The prepared hydrogel system was loaded into a 3cc syringe (Nordson) with a 25 gauge syringe tip (250 μm inner diameter), mounted on a three-dimensional translation stage (ATOM). A compressed air controller (Nordson) was used to actuate the printing of the hydrogel at constant pressure. The movement of the stage was controlled by software (Repetier), and the extrusion was coordinated by a pump switch. The separation distance (D_s), which is the length of hydrogel filament when separating from the nozzle, was measured by analyzing the recorded video of extruded hydrogel ink and measuring the filament length at the frame before the filament break (number of samples = 3). The nozzle height was characterized by gradually raising the nozzle from the substrate and measuring the extruded strut sizes via a microscope (DinoLite). The printing parameters’ effect on strut size was characterized under different nozzle movement speed and air pressure. For strut size characterization, hydrogel systems with different concentrations of sodium alginate (1.25% or 2.5%), xanthan gum, and Laponite (3% or 5%) were prepared and measured. Rheological measurements were performed using an ARES-G2 rheometer (TA Instrument, DE). Three groups were tested including 3% xanthan gum, 3% Laponite, or no additive (control). Other concentrations were kept at 10% GelMA, 1.25% alginate and 2% PEGDMA. Specifically, storage and loss moduli were measured according to a standard oscillatory time sweep protocol, with the step time ranging from 0 to 120 s. The shear-thinning property was evaluated by a flow ramp test with the shear rate changing from 0.3 to 12 1/s. The stress and viscosity were then measured and calculated for the evaluation of the shear-thinning property.

The mechanical property was characterized by measuring the printed samples’ compressive modulus via compression testing (Instron), during which the maximum compression was set to 50%, with an incremental speed of 1%/sec and a maximum load of 100 N. Results were analyzed in MATLAB, and the compressive modulus was evaluated by the modulus at 10% compression.

Degradation measurement

Hydrogels with 10% GelMA, 4% PEGDMA, 1.25% Alginate, and 3% xanthan gum or Laponite were prepared following the aforementioned protocol. Hydrogel samples were then syringe-printed into a patch of 1 cm × 1 cm × 1 mm, followed by crosslinking according to the aforementioned protocol. The weight of the printed samples was recorded immediately after printing. Crosslinked samples were then collected into 24-well plates separately. 1 mL of 1μg/mL collagenase solution was added to each well. The samples were then incubated at 37 °C and collected on Day 0, 1, 4, 7, 14 for collagenase condition, respectively. 3 samples were collected at each data point and then freeze-dried to measure their dry weight. The dry weight was normalized to the printed weight and then calculated for weight loss as evaluation for degradation.

Cell viability study

Human Mesenchymal Stem Cells (hMSCs) were cultured in DMEM medium (Life Technologies, USA) supplemented with 10% fetal bovine serum (FBS, Life Technologies, USA) and 1% Penicillin and Streptomycin. For encapsulation of cells in the hydrogel, hMSCs were trypsinized and added to the hydrogel solution before mixing with the viscosity enhancer solution. During the encapsulated hydrogel preparation, we modified the protocol by splitting half of the GelMA content into the xanthan gum solution before mixing, so that the maximum concentration is reduced, avoiding cell damage. The rest of the steps remained the same as hydrogel preparation and printing, except that the DMEM medium was used as the solvent instead of DI water. For Laponite groups, the solvent for Laponite component was still DI water due to Laponite’s insolubility in medium or PBS. Final cell density of the hydrogel mix was 1×10⁶/mL. Note that Laponite was still dissolved in DI water due to its insolubility in the DMEM medium. After printing into patches of 1 cm × 1 cm × 0.5 mm and crosslinking using the same process described in hydrogel preparation and crosslinking, samples were transferred into a 24-well plate, with 1 ml DMEM medium added to each well for culturing. Samples were then analyzed on Day 0, 1, 4, and 7. Upon analyzing, samples were washed with medium, live/dead staining kit (Invitrogen^™, Thermo Fisher, Waltham, MA) was used to visualize live and dead cells. Results were obtained with a fluorescent microscope (Zeiss), and the viability was analyzed via MATLab scripts.

Cell proliferation

4-week cell proliferation was evaluated by both metabolic activity and DNA quantification. The hydrogel preparation and cell encapsulation were performed using the same protocol described in the viability study, except that for the Laponite group, Laponite was first dissolved in DI water, followed by mixing at a 9:1 ratio with 10X DMEM medium to achieve physiological osmotic pressure. hMSC encapsulated hydrogels were then printed into discs of 6 mm diameter and 0.45 mm thickness, and collected on Day 0, 3, 7, 14, 21 and 28. For metabolic activity, samples were transferred into new well plates with 500 μL DMEM medium in each well for each sample. 100 μL of MTS Assay Kit solution (Abcam, UK) was then added to each well and incubated for 3 hrs. The absorption of resulted solution was measured by a plate reader (SpectraMax iD3, Molecular Devices, CA).

Statistical analysis

For all experiments three replicates were tested. All the data was subjected to statistical analysis and reported as a mean ± standard deviation. Statistical differences (*p<0.05) were determined using two sample t-test between two groups for multiple comparisons in viability and proliferation data.

Highlights.

• A novel GelMA/alginate bioink system with the inclusion of PEGDMA and xanthan gum

• Precision printability without the need for temperature control

• Wide mechanical property range enabled by PEGDMA and methacrylated gum

• Optimal solubility in ionic solutions including culture media, enabling optimal cell viability