Fine‐tuning Dynamic Cross–linking for Enhanced 3D Bioprinting of Hyaluronic Acid Hydrogels

3D bioprinting of stem cells shows promise for medical applications, but the development of an efficient bioink remains a challenge. Recently, the emergence of dynamically cross–linked hydrogels has advanced this field to obtain self‐healing materials. However, more advanced bioinks are needed that display optimum gelling kinetics, viscoelasticity, shear‐thinning property, structural fidelity, and hold the printed structures sufficiently long enough that allow maturation of the new tissue. Here, a novel extracellular matrix‐based bioink for human mesenchymal stem cells (hMSCs) is presented. Hyaluronic acid (HA) is modified with cysteine and aldehyde functional groups, creating hydrogels with dual cross–linking of disulfide and thiazolidine products. The investigation demonstrates that this cross–linking significantly improves hydrogel stability and biological properties. The bioink exhibits fast gelation kinetics, shear‐thinning, shape‐maintaining properties, high cell survival after printing with >2‐fold increase in stemness marker (OCT3/4 and NANOG), and supports cell proliferation and migration. Disulfide cross–linking contributes to self‐healing and cell migration, while thiazolidine cross–linking reduces gelation time, enhances long‐term stability, and supports cell proliferation. Overall, the HA‐based bioink fulfills the requirements for successful 3D printing of stem cells, providing a promising solution for cell therapy and regenerative medicine.


Introduction
3D bioprinting has advanced the field of tissue engineering by allowing spatiotemporal control over cellular organization mimicking the native tissue architecture.Despite great advancement in the field, an optimum bioink that provide excellent shearthinning property to reduce the shear stress experienced by the encapsulated stem cells are still elusive. [1,2]For the successful formulation of a bioink, the hydrogel must display a) optimum viscoelastic and shearthinning properties; b) an optimum processing window where the gel would crosslink upon passage through the nozzle without clogging c) limited toxicity to the encapsulated cells; d) prevent anoikis, i.e., minimize cell death after the cells are detached from the surface and before being cross-linked, as this is a critical time when cells are particularly vulnerable to stress and damage; e) transparency, particularly when using imaging techniques such as microscopy to visualize the printed structure.The most common bioink constituent that successfully demonstrates these vital obligations are composite materials that contain either gelatin methacrylate (GelMA) that allows UV-based cross-linking or the use of alginatecontaining gels that facilitate Ca 2+-mediated cross-linking. [1]Although these fabrication technologies significantly enhanced the field, UV-and Ca 2+ -based cross-linking are not ideal.Hydrogels developed by photo-cross-linking can result in cell toxicity, or mutagenesis to cells and display anisotropy due to inhomogeneous exposure to light. [3,4]Ionic cross-linking with Ca 2+ , on the other hand, is not suitable for long-term cultures as these bonds are unstable in the biological milieu resulting in the dissolution of the gels, as a result, additional covalent cross-linking of alginate gels are pursued. [5]Ca 2+ being an important signaling molecule could also alter the biological properties of the stem cells. [6]Recently, several non-UV-based and non-Ca 2+ based bioink fabrication strategies have developed, [7] however, there is a great need to develop new cross-linking strategies that are gentle to cells and are versatile, which would be easily adapted to different bioink compositions for cell therapy.
The utilization of 3D bioprinting for stem cell delivery has significantly improved the capacity to administer stem cells in a precise and controlled manner, using bioinks that provide a defined microenvironment that can modulate stem cell differentiation and stimulate tissue regeneration, either independently or in combination with gene therapy. [8,9]However, many aspects concerning stem cell-based therapy should be earnestly considered.One of the major concerns regarding stem cell-derived therapies is the survival and engraftment of transplanted cells into the host tissue.Overall, ≈1-20% of the transplanted cells survive, considerably limiting their therapeutic potential. [10]Silencing specific genes such as tissue factor genes in mesenchymal stromal cells (MSCs) [11] or polyelectrolyte coating of stem cells could prevent thrombotic activation and improve survival after implantation. [12]he use of hydrogels as stem cell carriers has also been reported as a promising approach to overcome these limitations by improving the cell survival and engraftment of the implanted stem cells. [1,13]Nevertheless, a successful injectable and printable hydrogel for stem cell delivery should fulfill all the requirements during formulation, injection, post-injection, and long-term survival phases. [10]or designing biomaterials for stem cell delivery, extracellular matrix (ECM) mimetic hydrogels have gained attention since they mimic the mechanical and biochemical properties of the stem cell niche microenvironment.16] For example, we have recently developed an HA hydrogel system to sequester recombinant bone morphogenetic protein-2 (rhBMP-2) to induce bone formation in rat models. [17,18]HA hydrogels can be synthesized using various chemical and physical cross-linking techniques, which allows for the fabrication of materials with different physical and mechanical properties.Nevertheless, hydrogels produced through dynamic cross-linking methods such as Schiff base, oxime and hydrazone, disulfide exchange, Diels-Alder, and boronic ester formation exhibit superior viscoelastic properties that could be necessary for the design of bioinks. [19]Dynamic cross-linking also facilitates shear-thinning and self-healing properties that safeguard cells during the injection phase against shear stress. [14,19]In one study, Wang et al. [20] demonstrated the use of HA hydrogel produced by dynamic hydrazone linkages for shear-thinning and self-healing properties, allowing successful delivery of cells through needles.Unfortunately, such dynamic bonds can make the hydrogel prone to fast erosion and viscous flow over time, making them challenging to use as a bioink.However, the incorporation of adhesive dopamine molecules on the HA backbone or the addition of rheological modifiers such as collagen and high molecular weight HA could partially resolve this issue. [7]The stability of hydrazone linkages could be improved by designing a hydrazide moiety having a delocalized positive charge on the neighboring amino residue [21] while the dynamic nature of these linkages could be improved by employing small molecule catalysts that can increase the printability of the hydrogel. [22]Other dynamic crosslinkages such as disulfide exchange have also gained attention for 3D printing and cell delivery since disulfide-cross-linked hydrogels can be dissociated by reductive catalysts such as glutathione, which is secreted by the cells. [23]Such disulfide-cross-linked hydrogels were shown to induce cell-mediated gradual degradation of the gels, as a result of endogenous glutathione produced by the cells. [24,25]However, hydrogels produced solely by disulfide cross-linking are prone to slow gelation kinetics at physiological pH and poor long-term mechanical stability. [26]One of the remedies to accelerate the gelation kinetics at physiological pH is to introduce electron-withdrawing groups at the -position of the thiol.We have recently demonstrated that introducing an amino group at the -position (e.g., cysteine modification) effectively lowers the pK a of the thiol group to 7.0, enabling hydrogel formation under physiological conditions. [27]Due to the susceptibility of disulfide linkages to cleavage by cell-produced factors, the long-term stability of the hydrogel in the presence of cells is limited.Therefore, an additional secondary cross-linking strategy is required.Thiazolidine chemistry is one such promising strategy that allows catalyst-free condensation reaction between 1,2 aminothiol moiety (or N-terminal cysteine derivative) and aldehyde-modified polymer to form a stable five-membered thiazolidine ring.We have previously shown that such a reaction can yield a fast and efficient thiazolidine product at both acidic and neutral pH, within 5 min, resulting in a stable product. [28]Despite the selectivity and efficiency of this reaction under physiological conditions, its potential for hydrogel formation remains largely unexplored.
In this study, we present the first thiazolidine and disulfide cross-linked HA-hydrogel to develop an ECM mimetic hydrogel that addresses the key challenges associated with stem cell survival upon injection, that can potentially be used for in vivo delivery of stem cells as well as for 3D printing of cells with defined architecture.To achieve this, we have conjugated cysteine derivative with HA carboxylate groups to create a disulfidecross-linked hydrogel at physiological conditions.The excess cysteine residues were reacted with aldehyde-modified HA to develop a thiazolidine cross-linked product.Such modification has been found to accelerate the hydrogel formation kinetics and increase the overall stability of the cell-encapsulated hydrogels over time.To optimize the composition of the hydrogel for 3D bioprinting, we have compared disulfide and thiazolidine cross-linked hydrogels with different ratios of cysteine and aldehyde residues to obtain gels with fast initial gelling kinetics with shear-thinning and self-healing properties for stem cell delivery.

Synthesis and Characterization of HA Derivatives
To design bioink for delivering stem cells, we envisioned employing disulfide and thiazolidine chemistry as these two reactions could be controlled by adjusting the concentration of the reactive functional groups.For this purpose, we aimed to develop HA aldehyde and HA cysteine derivatives, with a degree of modification of ≈10% with respect to disaccharide repeat units.This level of modification was deemed acceptable since a lower degree could result in a weak hydrogel, whereas excessively high modification may lead to faster reaction kinetics, limiting efficient mixing of the biopolymers having reactive groups.High chemical modification of HA may also change the cell-biomaterial interactions.To modify HA, we employed carbodiimide coupling chemistry with N-hydroxybenzotriazole (HOBt) as the nucleophilic catalyst to prepare HA-Cys and HA-Ald using our previously optimized protocol. [27,29]Of note, the aldehyde moiety was incorporated at the side chain without compromising the integrity of the sugar rings during the periodate oxidation step (Figure 1a).The degree of aldehyde modification on the HA backbone was quantified by 1 H NMR by performing a secondary reaction with tetra-butyl carbazate followed by sodium cyanoborohydride reduction as previously reported. [29] 1H NMR analysis of the ensuing product showed a tetra-butyl peak at 1.4 ppm which was compared with the N-acetyl signal at 1.9 ppm from HA to determine the degree of aldehyde modification which was estimated to be 10% (Figure S1, Supporting Information).The degree of cysteine modification was confirmed by validating the 1 H NMR signals corresponding to the methine (-CHNH 2 , 4.3 ppm) and methylene (-CH 2 SH, 2.8 ppm) protons.This result was also corroborated using Ellman's assay which indicated the degree of chemical modification to be ≈10% (Figure S1, Supporting Information).
The introduction of cysteine groups onto the HA backbone is expected to favor disulfide bond formation under physiological conditions, due to the presence of electron-withdrawing amino groups at the -position, facilitating the deprotonation of the thiol group.This reaction, however, is not fast with the pseudo-firstorder reaction rate k 1 = 5.04 × 10 −4 min −1 at pD 7.4. [27]On the other hand, the click-type condensation reaction between cysteine and aldehyde to yield thiazolidine product is reasonably fast at physiological pH. [28]To determine the reaction rate of thiazolidine formation, we performed a model reaction with acetaldehyde and L-cysteine at pD 7.4. 1 H NMR analysis of the product indicated complete conversion to thiazolidine product within 5 min (Figure 1b), as a result, the rate of conversion could not be determined under our experimental conditions.Because of such fast reaction kinetics, the hemiacetal intermediate also could not be observed within the NMR time scale.Our observed results are contrary to the previous report that suggested the formation of thiohemiacetal intermediate and slow conversion to thiazolidine product even at a 1:10 ratio of formaldehyde:cysteine substrate. [30]We believe the reaction with acetaldehyde and cysteine undergoes a concerted nucleophilic addition reaction of thiol moiety followed by the elimination of water molecule as a result of protonated amino residue in close proximity (facilitating proton transfer) and positive inductive effective of the methyl group.We therefore envisioned capturing the reaction intermediate by snap-freezing the reaction mixture in liquid nitrogen and monitoring the reaction mixture by 1 H NMR analysis.This however did not show any difference in NMR spectra.To trap the reaction intermediate, we froze the cysteine solution in the NMR tube and added acetaldehyde solution that diffuses slowly to the cysteine solution.Interestingly, this resulted in the observation of a thiohemiacetal signal at 1.4 ppm that slowly converted to thiazolidine product with time (Figure 1b, and full spectra in Figures S2-S4, Supporting Information).Our results unequivocally show that the cysteine derivatives indeed follow a concerted addition reaction to yield thiazolidine product without any observable thio-hemiacetal intermediate at room temperature, unlike reported recently. [31]Our result also corroborates with previous reports that suggest direct thiazolidine intermediate formation with higher-order aldehyde derivatives. [28,30]

Design of HA Hydrogels with Disulfide and Thiazolidine Chemistry
The distinct variance in reaction rates between disulfide and thiazolidine formation presents an exceptional opportunity to customize materials that possess both stability and dynamic properties.This can be achieved by adjusting the ratios of HA modified with cysteine and aldehyde residues in the reaction mixture.Such control over the reaction kinetics is particularly fascinating in the context of developing bioinks for stem cell delivery, where hydrogels with precise characteristics within a narrow range of gelling kinetics are required.It is crucial to strike the right balance, as excessively fast or slow reactions impede the efficient printing of the hydrogel scaffold.Hence, the ability to fine-tune the material properties allows for the creation of bioinks that exhibit optimal gelling kinetics, ensuring successful spatiotemporal control of stem cells in the printed matrix.We therefore evaluated five different ratios of HA-Cys to HA-Ald ratio (HA-Cys:HA-Ald) namely, 100:00, 75:25, 60:40, 50:50, and 40:60 (Figure 2a).The 100:00 ratio should theoretically correspond to 100% disulfide formation (abbreviated as "D-gel"), the 75:25 ratio on the other hand should correspond to 50% disulfide and 50% thiazolidine products (abbreviated as "DT-gel"), while 50:50 ratio should theoretically correspond to 100% thiazolidine formation (abbreviated as "T-gel").Although the actual types of cross-linked products may be influenced by various factors, including the mixing and arrangement of polymer chains and the proximity of functional groups, theoretical estimations can provide valuable indications for developing bioinks with optimal properties.

Time-Sweep Indicating Gelling Kinetics
To create injectable bioinks for delivering stem cells, the gelation kinetics need to be meticulously adjusted.In this study, we investigated the rate of gelation by vial inversion and quantitatively using time-sweep measurements with a rheometer.The vial inversion experiments indicated that HA hydrogel containing thiazolidine linkages (DT-gel and T-gel) formed a gel immediately after mixing the components, whereas the D-gel formed the hydrogel much slower (Figure 2b).To evaluate the precise time of cross-linking, we determined the crossover point of storage modulus (G′) and loss modulus (G″) as the gelling point (Figure 2c).The gelation data over a longer time is also presented in the Supporting Information (Figures S5-S7, Supporting Information).Interestingly, for the hydrogels with thiazolidine chemistry (DTgel and T-gel), G′ was >G″ immediately after mixing the two components which suggested that the gelling kinetics was too fast to be determined by the rheometer.
In contrast, for the disulfide-bonded hydrogel (D-gel), it took 74 min for G′ -G″ crossover, indicating a slower rate of disulfide bond formation compared to thiazolidine linkages at physiological pH. [27,28]The fast kinetics of thiazolidine gels corroborate with   the peptide-based hydrogel reported earlier that demonstrated thiazolidine cross-linked product formation within 30 sec. [32]

Stiffness and Swelling of HA Hydrogels
Since hydrogel stiffness is one of the determining factors that influence stem cell survival upon injection as well as adhesion, spreading, and differentiation, we set to determine how these properties are affected by different contributions of disulfide and thiazolidine linkages.For this purpose, we evaluated G′ of different materials through an amplitude-sweep test.Figure 3a presents the G′ of each hydrogel in relation to oscillation strain, while the stiffness of each hydrogel measured at 1% strain is presented in Figure 3b.The rheological characterization of different hydrogels demonstrated that with a higher percentage of disulfide modifications, a higher G′ value was observed as compared to gels with more thiazolidine linkages (D-gel = 3028 Pa, DT-gel = 2040 Pa, T-gel = 1128 Pa), however, no significant differences were detected in G″ (Figure S8, Supporting Information).The observed differences in G′ may be attributed to the differences in homogeneity of the hydrogels as a result of differences in reaction rates resulting in improper mixing of the gel components, as observed previously. [33]The T-gels with a 50:50 ratio of HA-Cys to HA-Ald, form cross-links very rapidly (not measurable by time-sweep experiment), as a result, the gel components do not get enough time to mix properly, causing decreases in mechanical stiffness.Increasing the thiazolidine content further by decreasing the cysteine component will further influence the cross-linking rate leading to a more heterogeneous gel with excess of free aldehyde groups.We have previously shown that the hydrogel gelling kinetics could also influence its swelling behavior [33] and therefore we determined the swelling percentage of these gels in PBS at 37 °C.Additionally, the type of cross-linking and homogeneity of hydrogels can influence the pore size and swelling ratio of the hydrogels.Our hydrogel swelling study indicated that the disulfide cross-linked D-gel demonstrated <30% swelling in PBS after 24 h, whereas the hydrogel with only thiazolidine linkages (T-gel) swelled twice as much (Figure 3c).When this experiment was continued for 28 days, all the hydrogels except the one with the 40:60 ratio of HA-Cys:HA-Ald were stable in PBS.
We therefore determined the average mesh size () and average molecular weight between cross-links (Mc) based on rubber elastic theory that can be applied to hydrogels that have elastic characteristics (Figure 3d).These theoretical calculations indicate that the hydrogels containing more disulfide crosslinking possess smaller pores with lower molecular weight between cross-links.For example, in D-gel the pore size is 11 nm, while in the sample formed by thiazolidine chemistry (T-gel), the pore size is 15 nm.This result suggests that more efficient cross-linking (e.g., D-gel) with reduced pore size can prevent absorption of water molecules, minimizing the swelling characteristics.

Shear-Thinning Properties of HA Hydrogels
We further investigated the impact of thiazolidine and disulfide linkages on the shear-thinning properties of HA hydrogels.In shear-thinning materials, the viscosity decreases in response to shear stress, allowing easy injection of stem cells through a needle, and minimizing toxicity due to the shear stress.Shearthinning hydrogels that are based on dynamic-covalent interaction which can reform after the removal of shear stress, can also display self-healing characteristics.A combination of shearthinning and self-healing behavior allows better retention of cellloaded materials post-injection, without clogging the needle or damaging the cells, which are desirable behaviors for developing bioinks. [34]To investigate the shear-thinning behavior of D-, DT-, and T-gels, the viscosity of gels was measured with respect to shear rate after 120 min of pH adjustment to 7.4 and mixing of gel components.The results, shown in Figure 3e, indicate that the viscosity of hydrogels decreases with increasing shear rate, demonstrating their shear-thinning property and suitability for injection.However, at a low shear rate of 0.1 s −1 , hydrogels containing thiazolidine linkages (DT-and T-gels) have a higher initial viscosity compared to those containing only disulfide bonds (D-gel).This is due to the differences in reaction rates of thiazolidine and disulfide formation as discussed earlier.As illustrated in Figure 3f the formation of thiazolidine cross-links provides an initial viscosity for the hydrogels, allowing them to keep cells inside the gel and hold the structure, while disulfide bonds form gradually after injection, increasing the final strength of the structure, providing a suitable bed for the delivery of cells through injection.

Self-Healing Properties of HA Hydrogels
Next, we investigated the effect of the presence of dynamic and non-dynamic cross-links in our hydrogel system on self-healing properties.To assess this, we performed injections of the hydrogels through a 22G needle at different time points after pH adjustment (7.4) and mixing the gel components (Figure 4a).Hydrogels injected after longer time points allow better curing of the materials resulting in the formation of a more rigid network.This was evident from the visual examination of the extruded materials.Injection of the hydrogels at longer time points (24 h), resulted in a decrease in transparency of the extruded gels, indicating deformation and crushing of the hydrogels.To further assess the potential of broken bonds to reform, we subjected the extruded gels at various time points to self-heal by placing them between two glass slides and maintaining them in a humidified environment at 37 °C for 24 h.As illustrated in Figure 4b the hydrogels formulated with thiazolidine chemistry (T-gel) were unable to fully regain their initial state when injected even after 10 min.However, hydrogels containing disulfide bonds (Dand DT-gels) exhibited a certain degree of recovery.Notably, injection of gels after 120-min and 24 h time points, demonstrated that only the hydrogels created with HA-Cys (D-gel) possessed the ability to recover and heal following extrusion.This material (Dgel) showed improved tolerance toward tensile stress posts selfhealing after 24 h as compared to the thiazolidine cross-linked material (Figure 4c).This particular gel also exhibited greater transparency compared to the thiazolidine-cross-linked hydrogel (T-gel), indicating the re-formation of covalent disulfide bonds in this sample (Figure 4d).Conversely, the thiazolidine-crosslinked T-gel displayed a poor self-healing property, which is expected due to its limited potential for the retro-thiazolidine reaction as illustrated in Figure 4e.Our results suggest that although both types of cross-links were susceptible to being broken upon extrusion through the needle, it is predominantly the disulfide bonds that exhibited the ability to reform and partially heal the hydrogel, particularly in the case of gels that had been allowed to gel for 24 h.This is presumably due to the dynamic nature of the disulfide bonds that allow the disulfide exchange reaction in the presence of thiolate ions in the vicinity. [14,35]o further prove the self-healing characteristics of our hydrogels, we performed an additional experiment involving two pieces of hydrogel that could self-heal/repair into one piece.These experiments also corroborated our previous observation that the hydrogels containing disulfide linkages (D-and DT-gels) were able to recover into one piece after incubation of the two pieces for 2 h at 37 °C (Figure S9, Supporting Information).The dynamic nature of the thiol-exchange reaction facilitated efficient repair of the gels as a result the D-gel completely healed while the DT-gel indicated partial healing, however, the T-gel did not heal after 2 h (data not shown).

Enzymatic and Non-Enzymatic Degradation of HA Hydrogels
HA is known to get degraded in vivo in the presence of a ubiquitous enzyme hyaluronidase.Disulfide cross-linking on such materials allows further degradation in the presence of cellproduced reducing agents such as glutathione.Degradation of materials could also be induced by treating them with acidic or alkaline solutions.We, therefore, investigated the stability of different cross-linked hydrogels under different conditions.In this regard, we first performed stability studies with thiazolidine and disulfide cross-linked materials under acidic and basic conditions (Figure 4f).Interestingly, these experiments revealed that hydrogels with disulfide and thiazolidine linkages (Dgel, DT-gel, and T-gel) are stable in both acidic and basic conditions, although all groups exhibited higher swelling in basic conditions.This observation is in line with a previous report that suggested that chemically-cross-linked HA hydrogels are more prone to degradation at higher pH, leading to disintegration and dissolution of the gel, regardless of the cross-linking chemistry. [36]It is generally believed that the thiazolidine reaction requires acidic pH, and the obtained product is also susceptible to hydrolysis under acidic conditions. [37]However, in our experiment, all the hydrogels were formed at physiological pH and they also demonstrated stability at acidic, neutral, and basic pH levels.
We then assessed the degradation rate of hydrogels in the presence of glutathione and hyaluronidase, the two natural molecules that are known to cleave disulfide bonds and  1→4 glycosidic linkages of the HA backbone, respectively.The weight change of the hydrogels over time was measured and expressed as a percentage of the initial weight.The hydrogel mass obtained after the initial setting served as the 100% reference.Disulfide-crosslinked hydrogels (D-and DT-gel) are expected to undergo disintegration in the presence of disulfide-reducing agents with increasing content of disulfide linkages.As anticipated, the D-and DTgel demonstrated higher swelling and faster degradation in the presence of glutathione after 8 and 72 h, respectively (Figure 4g).As anticipated, all the hydrogel groups (D-, DT-, and T-gels) degraded in the presence of hyaluronidase with ≈40 times the concentration of the enzyme in human plasma (Figure 4h).The hydrogel weight after 168 h incubation in hyaluronidase solution dropped to 50, 51, and 46% of initial weight for D-, DT-, and Tgels, in order.Surprisingly, even though D-gel showed the lowest swelling ratio in PBS, it experienced more swelling in the presence of hyaluronidase compared to DT-, and T-gels.Although the exact reason for such an observation could not be determined, the crystal structure of hyaluronidase indicates the presence of a cysteine-rich region in the C-terminal domain that can cleave the disulfide linkages. [38]We believe some of the free cysteine residues of hyaluronidase contribute to the cleavage of disulfide linkages resulting in the degradation of the hydrogel matrix.

Characterization of Encapsulated Cells in Disulfide and
Thiazolidine HA Hydrogels

Biocompatibility of Hydrogels
To evaluate the biocompatibility of the materials, a human osteosarcoma cell line (MG63) and hMSCs were encapsulated within the hydrogels, and Live/Dead staining was carried out.The fluorescent microscopic images (Figure S10, Supporting Information) demonstrate that all compositions of HA-Cys to HA-Ald hydrogel display excellent biocompatibility when cultured with MG63 cells.In addition, confocal microscopic images of hMSCs encapsulated in the hydrogels after 1 day demonstrated that more than 95% of the cells were viable in all the groups (Figure 5a).This evaluation of the hydrogels showed high biocompatibility of all the hydrogel groups as there was no significant difference in the viability of cells in the D, DT, and T-gel groups after 1, 3, and 7 days of encapsulation (Figure 5b).This indicates that disulfide and thiazolidine cross-linking do not induce any detrimental effects on encapsulated cells.

Cell-Matrix Interactions on Matrix Stiffness and Cell Proliferation
Encapsulation of cells within the hydrogel matrix is known to influence the mechanics and degradation kinetics of the hydrogel.This is attributed to several factors such as the cell-secreted factors (e.g., matrix metalloproteases) and the soluble factors present in the cell culture medium which modulate the stability of the cross-links and the viscoelastic properties of the hydrogel. [38]To ascertain the impact of encapsulated cells on disulfide and thiazolidine cross-links, we measured the stiffness of hydrogels in the presence or absence of the encapsulated cells.Our findings reveal an interesting trend (Figure 5c), D-gel exhibited a significant drop in storage modulus value from 2203 ± 413 Pa without cells to 706 ± 245 Pa with cells after 3 days of cell culture.In contrast, no noticeable change was observed in the hydrogel formed by thiazolidine linkages (T-gel), with the storage modulus remaining relatively stable at 790 ± 177 Pa and 866 ± 176 Pa in the absence or presence of the encapsulated cells, respectively.Moreover, in the DT-gel, the storage modulus decreased from 1600 ± 360 Pa to 886 ± 221 Pa, indicating that cells have a more pronounced effect on the mechanical properties of disulfide-cross-linked than thiazolidine-cross-linked gels.This behavior stems from the cleavability of disulfide bonds in the presence of glutathione.Glutathione, a tripeptide composed of L--glutamyl-L-cysteinyl-glycine, is synthesized in the cytosol from glutamate, cysteine, and glycine, which is secreted by cells.The intracellular concentration of glutathione is in the millimolar range, which enables the formation of native disulfide bonds  in the endoplasmic reticulum through a complex process involving the disulfide-bond formation and isomerization of nonnative disulfide bonds. [39]As a result, when cells are encapsulated within a disulfide-bonded matrix, the glutathione produced by the cells can interact with the matrix bonds to reduce them, leading to changes in the stability and stiffness of the hydrogel.Thus, the cell-mediated matrix remodeling by the cleavage of disulfide bonds could be proportional to cell proliferation.To ascertain this factor, we quantified the cell proliferation inside the hydrogels using a metabolic PrestoBlue assay. [40]All the hydrogels showed high viability of hMSCs (≈95%) after 1 day of culture, and after 3 days, we observed a comparable cell proliferation between all the groups (Figure 5d).However, after 7 days of 3D culture, this number dropped in the disulfide-cross-linked (D-gel), while it continued the increase in the T-gels.These results indicate that the hydrogel cross-linked by thiazolidine chemistry (T-gels) provided superior culture conditions for stem cells when compared with D-gels.A similar trend was also observed for MG63 encapsulated hydrogel; however, the proliferation rate of cancer cells was higher than the encapsulated hMSCs, indicating the differences in the rate of cell division or proliferation between MG63 and hMSCs (Figure S10, Supporting Information).The observed difference in cell proliferation between D-gels and T-gels may be attributed to several underlying factors, such as the cellular influence on the matrix, glutathione-mediated degradation of crosslinks, and the stiffness characteristics of the hydrogels.Notably, the higher stability of the T-gels in the presence of cells may be attributed to the higher stability of the thiazolidine linkage and higher retention of cells within the gel.Additionally, the T-gels displayed a higher storage modulus (or matrix stiffness) when compared to D-gels, which could be another factor for the higher proliferation of hMSCs as stiffer gels are reported to show higher proliferation and viability of hMSCs. [41]us, by carefully tuning the ratio of disulfide and thiazolidine cross-links in HA hydrogel, we can achieve an optimal balance between the two types of cross-links, resulting in a biomimetic matrix with enhanced proliferation of stem cells with controlled degradation and release.It is worth noting that manipulating the composition of cross-links in HA hydrogel can provide a promising platform for developing advanced biomaterials with tailored mechanical properties with tunable degradation and controlled release capabilities, which could significantly impact the field of tissue engineering and regenerative medicine.

Migration and Morphology of the Encapsulated Cells
Cell migration is a crucial process for stem cells in 3D hydrogels, enabling them to interact with their microenvironment and respond to necessary cues for survival and function.However, successful 3D cell migration depends not only on the cell type and its inherent ability to adapt to environmental changes but also on the intrinsic properties of the hydrogel, such as stiffness, pore size, and adhesiveness. [42]Limited migration of cells within a living scaffold hinders the formation of functional tissue structures and leads to graft failure or poor clinical outcomes. [10]Thus, understanding the factors that regulate cell migration in 3D hydrogels is critical for developing effective strategies to enhance the survival and function of stem cells for tissue engineering applications.In order to investigate the potential of disulfide and thiazolidine-bonded hydrogels to support cell migration in 3D environments, we conducted a chemotaxis-based assay.Here, we fabricated the hydrogels in the presence of fetal bovine serum (FBS) without cells, and thereafter the center of the gel was carefully removed using a biopsy punch, and the space was filled with hydrogels containing cells but without any FBS (Figure 6a).
After 24 and 48 h, the cells were stained with Hoechst fluorescent dye, and their movement from the center to the wall of the plate was evaluated using a fluorescent microscope to assess the gradient of cell migration.As anticipated, the cell migration was more prominent in D-gels when compared to T-gels (Figure 6b,c).Surprisingly, the migration of cells in the T-and DT-gels were significantly restricted compared to D-gel.The difference in cell migration observed between D-gel and T-gel hydrogels developed through disulfide and thiazolidine cross-linking could be attributed to the same reasons responsible for cell proliferation, namely, glutathione-mediated degradation of cross-links and its impact on the stiffness characteristics of the hydrogels.The disruption of disulfide bonds will also lead to an increase in porosity, facilitating cell migration.In contrast, thiazolidine linkages will remain intact in the presence of glutathione produced by the cells.This effect was balanced in the DT-gel where the thiazolidine linkages and disulfide groups displayed differential stability thereby limiting the migration of cells particularly after 24 h of cell encapsulation.These results suggest that while cells may disrupt some of the disulfide cross-linking which can facilitate cell migration, the thiazolidine cross-linking can hold the hydrogel structure and restrict cell movement.Such hydrogels with tunable porosity and cell migration capability will be advantageous for cell delivery applications where the release of cells needs to be precisely regulated.
Another aspect that can indicate cellular behavior in a 3D matrix is the morphology of the encapsulated cells, which can reveal cell differentiation and integrin-mediated interaction with the hydrogel matrix.Entrapment of cell-produced factors by the matrix remodels the scaffold and alters cell behavior.Stem cell morphology also affects the property of the engineered tissue, e.g., stem cells that are elongated and aligned interact efficiently with other cells within the hydrogel, thereby facilitating the formation of functional tissues. [43,44]To investigate the impact of disulfide and thiazolidine cross-linked HA hydrogels on cell behavior, DAPI/phalloidin staining was used to visualize the nucleus and actin filaments of encapsulated cells.The resulting confocal microscopy images revealed that cell morphology remained rounded and did not differ between the D, DT, and T-gel groups of hydrogels (Figure S11, Supporting Information).HA primarily interacts with two cell-surface receptors, cluster determinant 44 (CD44) and receptor for hyaluronatemediated motility (RHAMM).CD44 is involved in tissue organization by facilitating ECM remodeling, cell-cell interactions, and cell-matrix interactions. [45]It also has been shown that CD44knockout mice exhibited decreased cell migration, motility, and ECM turnover. [46]Despite playing a crucial role in cellular behavior, HA does not possess integrin binding sites, which limits its ability to support cell spreading.However, this can be advantageous as rounded stem cells retain their stemness and preserve their undifferentiated state when compared to elongated and polarized cells. [44]o investigate if the morphology of the cells and cell adhesion of the encapsulated stem within the matrix can be altered, we added gelatin to the pre-solution of the HA hydrogel components.Gelatin contains the RGD (arginine-glycine-aspartic acid) sequence, which is a well-known cell adhesion motif.This sequence is known to bind to integrin receptors on the surface of cells, promoting the relationships between the cells and the surrounding ECM. [47]The RGD sequence is often added to biomaterials to enhance their ability to support cell attachment and proliferation.The confocal microscopic images of the cell-encapsulated hydrogels revealed that despite the presence of gelatin in the hydrogels, the cells remained rounded in the hydrogel formed by thiazolidine chemistry (Figure 6d).However, some degree of cell elongation was observed in the hydrogels containing disulfide-linkages, with the most significant elongation seen in the hydrogel formed solely with disulfide bonds (D-gel).This can be due to greater cell migration in disulfide-bonded hydrogels in a short time, which enables cells to move and interact with the RGD sequence of gelatin.In addition, it may stem from the initial stiffness of hydrogels that can impact cell morphology, [36] since disulfide-cross-linked hydrogel (D-gel) has a storage module 1.5 and 2.4 higher than DT and T-gels, respectively.

3D Printing of Stem Cell-Laden Hydrogels
In the next step, we explored the potential of different hydrogels as a bioink for 3D bioprinting applications.In recent years, 3D bioprinting has revolutionized the field of tissue engineering as it allows the creation of complex and functional tissues that can be tailored with precise shape, size, and anatomical features. [48,49]oreover, 3D bioprinted tissues can serve as realistic models for drug testing, reducing the need for animal testing, and with the ability to create patient-specific tissues and organs, 3D bioprinting holds promise for personalized treatment approaches and therapies.However, one of the key challenges in the field of extrusion-based bioprinting is to design bioinks that allow controlled delivery of stem cells without injury or alteration of cellular function.
For 3D printing different gel compositions, we chose a needle with a gauge size of 25G, and the pressure was adjusted according to the hydrogel viscosity.The 3D rectangular prism model with the dimension of 10 mm × 10 mm × 1 mm and an infill ratio of 40% were printed.In this study, four layers of the structure were printed for each gel composition.Extrusion printing of T-gel displayed limited printability with a printable time of <10 min before clogging the needle (Figure 7a).Conversely, D-gel had a more extended printable window, but the gel lacked sufficient viscosity to maintain the matrix stability, however, after 150 min of pH adjustment, the gel could be printed with reasonable fidelity.Such an extended time could compromise the viability of stem cells due to inadequate oxygen and nutrient supply.Interestingly, the DT-gel displayed an optimum processing window of 60 min and yielded the most precise structure, as observed in the microscopic images (Figure 7a).These results suggest that the combination of DT-gel can provide a superior bioink compared to hydrogels with only thiazolidine or disulfide cross-linking.The rapid gelation of thiazolidine provides the necessary initial viscosity to the gels, while the gradual formation of disulfide allows for a more extended printing time and elasticity.
We designed another structure with DT-gel having a dimension of 20 mm × 20 mm × 1 mm with an infill ratio of 5% by printing four layers.As illustrated in Figure 7b the printed structure maintained its stability even after 12 and 24 h of printing, with no discernible alterations in its dimensions or overall structure.Additionally, subsequent examination revealed that when the printed structures were immersed in PBS for a duration of 1 h, 24 h, and 14 days, there were no significant swelling or changes in sample dimensions.This underscores the unique potential of the hydrogel to print diverse structures with high fidelity while maintaining structural stability in PBS.
Furthermore, to showcase the robust capabilities of DT-gel for printing larger structures with multiple layers, two additional designs were made with a dimension of 10 mm × 10 mm × 5 mm (infill ratio of 5%), and 20 mm × 20 mm × 2.5 mm (infill ratio of 5%), respectively.As depicted in Figure 7c,d, these designs are well accommodated to yield higher-order structures with a 20-layer (≈5 mm in height) and a 10-layer printed structure (≈2.5 mm in height).Swelling of the 10-layer printed structure in PBS over 21 days revealed that the hydrogels maintained their structural integrity with minimal change in overall dimension (Figure 7e).
In order to assess the efficacy of the hydrogel in promoting cell survival following extrusion, hMSCs were embedded within the gel prior to printing.The resulting printed structure displayed a uniform distribution of cells (Figure 8a,b).Furthermore, the viability of the cells post-printing was assessed using Live/Dead staining, and representative microscopic images from two distinct sites within the printed structure are presented in Figure 8c.These findings confirm that the DT-gel composition possesses ideal viscoelastic characteristics and is capable of preserving cell viability (with survival rates exceeding 85%).Thereafter, we compared the proliferation of cells within the bioprinted DT-gel, over a period of 7 days (Figure 8d).These experiments showed that the viability and proliferation of cells were higher in the group of encapsulated cells without extrusion compared to those in the 3D-printed samples, indicating the possibility of cell damage during the extrusion process.Nonetheless, the cells were able to recover their proliferative capacity, as evidenced by an increase in the proliferation rate from 85±6% to 119±9% between days 3 and 7.These findings suggest that while 3D printing can cause some initial damage to the cells, this hydrogel composition is still able to support their recovery and subsequent proliferation.Previous studies have suggested that cells exposed to syringe needle flow can experience mechanical shear and extensional forces that can damage the cell membrane.This effect is exacerbated when cells are injected in low-viscosity solutions such as saline or cell culture medium. [10]o assess the hydrogel's efficacy in shielding cells from shear forces during injection, a cell culture medium was employed as a comparison carrier for stem cells.The cells were injected separately via a 3D bioprinter using varying needle sizes and identical pressure.To measure the extent of damage to the plasma membrane upon injection, we monitored the release of lactate dehydrogenase (LDH) enzyme into the cell culture medium.LDH is a cytosolic enzyme that is released into the cell culture medium upon damage to the plasma membrane, which can be easily quantified using colorimetric assay.The results obtained from the LDH assay indicated differences in cell viability in three groups, namely, cells encapsulated in the DT-gel without injection, cells injected via cell culture media, and cells injected via DT-gel (Figure 8e).The group of encapsulated cells without injection exhibited the highest viability, while some cytotoxicity was observed in all injected samples.The percentage of dead cells extruded from a 22G needle was significantly lower than that of 25G and 27G needles, which possess smaller diameters.Injection of cells using a 25G needle resulted in cell death in all cases.Notably, using a smaller needle size (27G) showed a marked difference in cell viability between the hydrogel and cell culture media.Interestingly, the DT-gel could preserve 85±3% of the cells after injection through a 27G needle, while cell culture media only partially protected the cells upon injection (75±3%), demonstrating the hydrogel's excellent ability to shield cells from anoikis and further damage.
To assess the impact of both the hydrogel composition and the 3D bioprinting process on the stemness of the encapsulated hM-SCs, we quantified the expression levels of OCT3/4 and NANOG (Figure 8f).These pivotal transcription factors are responsible for preserving pluripotency and the self-renewal capacity of undifferentiated stem cells.We employed a quantitative polymerase chain reaction (qPCR) to measure the expression levels of these two markers.Interestingly, a comparison of 2D versus 3D (encapsulated as well as 3D printed) cell culture indicated that the overall expression of stemness markers increased in the 3D culture conditions.The relative expression of NANOG increased by 2.5-3.0 folds while the expression of OCT3/4 increased by 2.0 -2.4 folds in the 3D printed and encapsulated conditions, respectively (Figure 8g).The increase in stemness of hMSCs could be attributed to the dedifferentiation of stem cells due to hypoxia [50] and increased HA-CD44 interaction that promotes Rho/ROCK signaling in MSCs. [51]hus, we have discovered a unique combination of cellresponsive disulfide cross-linking in synergy with stable thiazolidine cross-linking to obtain a functional bioink that possesses optimum self-healing and injectability characteristics.This bioink formulation offers protection to the cells during the injection process and supports cell proliferation and migration.These results have major implications in the field of tissue engineering, where the development of bioinks that can support the viability and functionality of cells is critical for the successful fabrication of functional tissue constructs.

Conclusion
In conclusion, we present the development of a novel ECMmimetic bioink for 3D bioprinting.By utilizing dual crosslinking mechanisms involving disulfide and thiazolidine chemistry, we have achieved remarkable enhancements in hydrogel stability and biological properties.The combination of two different cross-linking methods in our bioink has exhibited exceptional characteristics, including optimal gelation kinetics, shearthinning behavior, and shape-maintaining capabilities.Our study Figure 7. 3D printing of hydrogels.a) A table presentation of 3D printed structures of different hydrogel compositions using an extrusion-based printer with a needle size of 25G.Representative digital images of the printed structures after mixing the gel components for different time intervals.The microscopic images of the printed samples were shown after 150 min for D-gel and 10 min for T-and DT-gels; b) A printed structure with 4 layers and its stability after 12 and 24 h after printing, followed by incubation in PBS for 1 h, 24 h, and 14 days; c) Digital images from printed structure with 20 layers; d and e) Digital images from a printed structure with 10 layers and its stability after 21 days of incubation in PBS, respectively.highlights the critical role of optimizing both dynamic and nondynamic cross-linking for developing optimal bioink.This optimization process has proven essential in attaining biocompatible bioink, as demonstrated by high cell viability post-printing, as well as providing support for cell proliferation and migration.Furthermore, our two-component system offers the unique opportunity of loading two different types of cells, such as cancer cells and immune cells, between printed layers.This capability allows for the exploration of crosstalk between these cell types, opening new avenues for studying intricate cellular in-teractions.Importantly, the strategy we have employed to incorporate cysteine and aldehyde functional groups can be readily adapted to modify other biopolymers, broadening the application of the dual cross-linking methodology to create bioinks using diverse biopolymers.
Our study highlights a remarkable stride in the evolution of novel ECM-mimetic bioinks for 3D bioprinting.Through optimization of both dynamic and non-dynamic cross-linking mechanisms, we have harnessed an equilibrium of properties that catapults our hydrogel to the forefront of chemically cross-linked bioinks.The optimized gelation kinetics, unparalleled self-healing capabilities, and favorable biological cues to support stem cells make these bioinks suitable for studying cell-matrix interactions.These results hold immense potential for bioprinting matrices for tissue engineering and regenerative medicine, offering the means to usher in an era of cuttingedge bioinks tailored for the most demanding 3D printing applications.
HA Modification-Synthesis of Aldehyde-Modified HA (HA-Ald): HA-Ald was synthesized following a previously published protocol from the group. [29]Briefly, 0.4 g of HA (1 mmol, 1 eq with respect to disaccharides) was dissolved in 100 mL of deionized water.To this solution, 0.135 g of HOBt (1 mmol, 1 eq) dissolved in 2 mL of DMSO was added to the HA mixture.Subsequently, 0.091 g of 3-amino-1,2-propanediol (1 mmol, 1 eq) was added to the reaction mixture and stirred until a homogeneous solution was obtained.The pH of the resulting solution was adjusted to 6.0, followed by the addition of 0.057 g of EDC (0.3 mmol, 0.3 eq) in two portions at 30-min intervals.The reaction mixture was stirred overnight.The resulting mixture was then dialyzed against a dilute HCl solution (pH 5) containing 0.1 m NaCl.This dialysis step was carried out twice, using a total volume of 2 L for each dialysis, and the duration was 48 h for each dialysis.Subsequently, the mixture was dialyzed in dilute HCl (pH 5) without NaCl, resulting in diol-modified HA.To obtain HA-Ald, the diol-modified HA was further treated with 0.213 g of NaIO 4 solution (1 mmol, 1 eq) dissolved in 0.5 mL of water that was added dropwise under stirring to the reaction mixture.After 10 min, 0.310 g of ethylene glycol (5 mmol, 5 eq) was added to quench the unreacted NaIO 4 .The reaction mixture was allowed to stir for an additional 1.5 h.The resultant solution was then dialyzed against deionized water using the same dialysis protocol described earlier (2 L of water, 48 h).Finally, the solution was lyophilized to obtain the aldehyde-modified HA (HA-Ald).The percentage of modification of HA was determined relative to the disaccharide repeat units of HA using 1 H NMR spectroscopy.This was achieved by treating HA-Ald with tertbutyl carbazate followed by NaCNBH 3 reduction.The modification level was determined by integrating the methyl peaks of the conjugated t-butyl carbazate relative to the methyl peak of the N-acetyl group of HA.More information regarding the NMR analysis is presented in the supporting information.
HA Modification-Synthesis of Cysteine-Modified HA (HA-Cys): To prepare the cysteine-modified HA (HA-Cys), the previously reported procedure as described below was followed. [27]Briefly, 0.4 g of HA (1 mmol, 1 eq with respect to disaccharides) and HOBt (0.135 g, 1 mmol, 1 eq) were dissolved in 50 mL of deionized water.To this solution, 0.310 g of 3,3′dithiodi(2-aminopropanehydrazide) (0.9 mmol, 0.9 eq) was added and stirred until a clear solution was obtained.The pH of the solution was adjusted to 4.5 using 1 m NaOH or HCl.Next, 0.042 g of EDC (0.22 mmol, 0.22 equivalent) was added in two portions while maintaining the pH at 4.5.The reaction mixture was stirred at room temperature for 12 h.Subsequently, DTT (0.771 g, 5 mmol, 5 eq) was added, and the reaction mixture was stirred overnight at room temperature.The resulting solution was diluted to a concentration of 4 mg mL −1 , and the pH was adjusted to 4 by adding 1 m HCl.It was then dialyzed against a dilute HCl solution (pH 5) containing NaCl (0.1 m) under nitrogen gas purging for 24 h, with three changes of dialyzing solvent.This was followed by dialysis against dilute HCl (pH 5) with nitrogen gas purging for an additional 24 h, again with three changes of dialyzing solvent.The resulting solution was finally lyophilized to obtain the cysteine-modified HA (HA-Cys).The incorporation of thiol functionality was confirmed through 1 H NMR analysis, which could be found in the supporting information.The degree of modification was also assessed using Ellman's test, employing 5,5′-dithiobis-(2-nitrobenzoic acid) reagent (DTNB).DTNB, containing an oxidizing disulfide bond, reacts with free thiols, resulting in the release of 5-thio-2nitrobenzoic acid (TNB) and the formation of a mixed disulfide. [52]The concentration of free thiols in HA-Cys was determined by measuring the absorption at 412 nm using a UV-vis spectrometer.
Hydrogels Formation and Characterization: In the experimental study, aldehyde-modified HA (HA-Ald) and cysteine-modified HA (HA-Cys) were utilized to investigate the influence of cross-linking chemistry on gelation kinetics and material properties.As a proof of concept for thiazolidine cross-linking, the reaction between acetaldehyde and L-cysteine was employed as the model system.The kinetics of thiazolidine formation were carefully monitored using detailed 1 H NMR analyses to gain insights into the gelation kinetics in the system, detailed results can be found in Supporting Information.To prepare hydrogels with different cross-linking patterns (thiazolidine/disulfide), the following procedure was followed.Initially, stock solutions of HA-Ald and HA-Cysteine in PBS at pH 7.4 were prepared separately.It was important to note that HA-Cysteine was initially dissolved in PBS at pH 5.0 to obtain a homogeneous solution, which was then adjusted to pH 7.4 using 1 m NaOH.To achieve varying distributions of thiazolidine and disulfide formation, HA-Cys and HA-Ald solutions were mixed and vortexed in different ratios (Cys:Ald) with the following compositions: 100:00 (D-gel), 75:25 (DT-gel), 60:40, 50:50 (T-gel), and 40:60.The final solid content of the hydrogels was set to 2%.Before conducting experiments involving cell encapsulation, the hydrogel materials underwent rigorous UV sterilization for 20 min and were subsequently dissolved in sterile solutions.
Rheological Properties: To evaluate the rheological properties of the hydrogels, 300 μL volumes were prepared in the form of cylinders with a diameter of 12 mm and left to cure for 24 h in a humid environment.The TA instruments, TRIOS Discovery HR 2 rheometer was then used to measure the storage modulus (G′) and loss modulus (G″) values through an amplitude-sweep, which was then plotted against the strain to demonstrate the viscoelastic shear behavior of the materials.Gelation time was also determined using the same rheometer, with all measurements conducted inside a humidity chamber to prevent hydrogel desiccation.After pH adjustment, the solutions of the HA derivatives (450 μL) were immediately transferred to the rheometer, and the gap was adjusted to a size of 500 μm.Using 20 mm parallel plate geometry (strain,1%; frequency, 0.5 Hz), the storage and loss modulus were recorded over time, and the gelation time was defined as the point at which G′ crossed G′′.In addition to time-sweep evaluation, the vial inversion method was also used to investigate the gelation time of hydrogels.For this experiment, HA-Cys and HA-Ald were stained with Toluidine blue for better visualization and after pH adjustment to 7.4, they were mixed in a glass vial.After 10 and 120 min of mixing, the vials were tilted to 45°, and digital images were taken to monitor the gelation of the samples.Furthermore, the shear-thinning and injectability behavior of the hydrogels were also monitored by measuring the viscosity (*) against different shear strains after 120 min of pH adjustment to 7.4 and mixing of the gel components (HA-Cys and HA-Ald).
To understand the relationship between the molecular structure and the rheological data, the average mesh size () and the average molecular weight between cross-links (Mc) were calculated using the following Equations (1) and ( 2) that applied to hydrogels that had elastic behavior. [40] = where NA is the Avogadro constant (6.022 × 1023 mol −1 ), R is the molar gas constant (8.314J mol −1 K −1 ), T is the temperature (298 K), c is the polymer concentration (kg m −3 ), and G′ is the hydrogel storage modulus (Pa).

Swelling and Stability of HA Hydrogels:
To study the stability and swelling characteristics of the hydrogels, prepared samples were subjected to different conditions including PBS pH 7.4, hyaluronidase (250 U mL −1 ) solution, and glutathione solution (2 mM) in PBS at 37 °C.Briefly, 250 μL gels were prepared in an 8 mm mold placed in a glass vial, and kept for 24 h to complete the gelation, and then, the initial weight of the hydrogels was recorded.Prior to mixing HA-Cys and HA-Ald, Toluidine blue was used to dye the HA chains for better visualization.The gels were then submerged in PBS with pH 7.4 for swelling study.To determine hydrogel degradation hyaluronidase and glutathione solution were employed.To observe the swelling and subsequent stability characteristics of the hydrogels, at different time points, the solution was taken out and gels were weighed, afterward, the buffer was replaced with a fresh solution.The remaining weight percentage was calculated by using the equation (3).Additionally, equation ( 4) was employed to calculate the swelling percentage for the samples.

Mass change (%) =
Measured weight Initial weight ( 3) Cell Culture: MG63 cells were purchased from the American type culture collection (ATCC).These cells were cultured in T-75 cell culture flasks in Dulbecco's Modified Eagle Medium (DMEM, GlutaMAX, Gibco) with 10% fetal bovine serum (Gibco) and 1% Penicillin-Streptomycin as an antibiotic (DMEM complete medium) in a cell culture incubator at 37 °C and 5% CO 2 .The medium was changed every alternate day.TrypLE Select (Gibco) was used to detach the cells from the flasks during passaging.Cells were passaged upon reaching ≈80% confluency.
hMSCs ( Donated by AO Foundation, Davos, Switzerland) were cultured in T-150 cell culture flasks in Dulbecco's Modified Eagle Medium (DMEM, Gibco) with 10% fetal bovine serum (Gibco) and 1% Penicillin-Streptomycin as an antibiotic in a cell culture incubator at 37 °C and 5% CO 2 .Cells from passages 2-7 were used for experiments.
Before encapsulation, cells were detached from the flasks using TrypLE Select, centrifuged, and resuspended in a complete medium, counted, and the required number of cells was further centrifuged, and the pellets were resuspended in HA-Cys solution after pH adjustment to 7.4.200 μL of hydrogels were formed by mixing HA-Cys with HA-Ald, and were incubated for 1 h at room temperature followed by the addition of 350 μL complete medium to each well in a 48-well plate.The medium was changed every alternate day or as otherwise indicated.
Cell Culture-Live/Dead staining of 3D Encapsulated Cells: To assess cell viability inside the gels, cells were encapsulated in 250 μL hydrogels at a concentration of 2 × 10 6 cells mL −1 for MG63 and 0.5 × 10 6 cells mL −1 for hMSCs in a 48-well plate and cultured for 1, 3, and 7 days with alternate-day medium changes.Live/Dead staining (Viability/Cytotoxicity Kit for mammalian cells, ThermoFisher) was performed to visualize cell viability over time.To conduct the staining, the hydrogels were washed twice with PBS after removing the medium from the wells.Then, a 300 μL Live/Dead staining solution containing 0.5 μL mL −1 Calcein AM and 2 μL mL −1 Ethidium homodimer in cell culture medium was added to the wells, and the plates were incubated at 37 °C for 1.5 h.After incubation, the hydrogels were washed two times with PBS and imaged using a Nikon Eclipse Ts 2 fluorescence microscope with a 10x objective and an LSM 700 confocal microscope for 3D images at 10× using z-stack images series.The obtained images were analyzed using ImageJ software, and cell viability was calculated by comparing the number of green cells (live) to the number of green and red cells (dead).
Cell Culture-Cell Proliferation in the Hydrogels: To monitor the metabolic activity of cells and their proliferation in 3D, a PrestoBlue (Thermofisher) assay was applied.When added to the cells, the PrestoBlue reagent was modified by the reducing environment of the viable cell and turned red in color, becoming highly fluorescent.In this study, cells were encapsulated in 200 μL gels in a 48-well plate and incubated for 1, 3, and 7 days.The proliferation number at each time point was obtained from fluorescent intensity measured by a plate reader.The proliferation percentage was calculated by comparing the values obtained at each time point to the value obtained immediately after cell encapsulation in the hydrogels (time point zero).
Cell Culture-Cell Migration in 3D Culture: To assess cell migration within the hydrogels, 120 μL of hydrogels in PBS containing 10% FBS were prepared in a 96-well plate and incubated overnight.A 2 mm diameter biopsy punch was used to remove a cylindrical portion from the center of each hydrogel, and the void was replaced with hydrogels containing 2 × 10 6 cells mL −1 .The gels were then incubated with 200 μL of cell culture media for 24 and 48 h.After removing the media, the gels were washed twice with PBS.Hoechst, a blue fluorescent dye that stains DNA, was applied to visualize the cell nuclei, and images were captured using fluorescent microscopy.The images were compared to those taken at time point zero (just after encapsulation and placement of the gel) to monitor cell migration over time.The diameter of cell migration was measured using Image J software.
Cell Culture-Immunostaining of Cell Cytoskeletal Organization: To examine the impact of disulfide and thiazolidine cross-links on the cytoskeletal organization (F-actin) of hMSCs, 6-diamidino-2-phenyl indole dihydrochloride (DAPI) and phalloidin staining (ThermoFisher) was used.Hydrogels were formed as previously described using 2% solid content of HA.Additionally, for the experiment containing gelatin in the hydrogels, 0.4% of gelatin was physically added to HA solution, and cells were encapsulated at a concentration of 0.5 × 10 6 cells mL −1 .After 1 day of incubation, the hydrogels were washed twice with PBS and fixed with 4% paraformaldehyde for 30 min.Then, cells were permeabilized by 0.1% Triton X-100 (Sigma) for 20 min, washed with PBS, and incubated in 0.5% bovine serum albumin in PBS for 1 h to reduce nonspecific background staining.After triple rinsing with PBS, the cells were stained with DAPI and rhodamine phalloidin solutions.DAPI stains cell nuclei, while rhodamine phalloidin stains actin filaments.Finally, the stained samples were observed under an LSM 700 confocal microscope at 20×, 40×, and 63×.
3D Printing of Hydrogels with Encapsulated Cells: To evaluate the printability of the hydrogels, the hydrogel components (HA-Cys and HA-Ald), were dissolved in PBS and the pH was adjusted to 7.4.These solutions were then combined into a single syringe and subjected to vigorous vortexing to ensure thorough mixing.Subsequently, the blended material was extruded through a needle into plastic printing cartridges provided by CELLINK Corporation, Sweden, to allow complete homogenization of the gel mixture.Subsequently, different designed structures with rectilinear infill patterns were printed with a pneumatic printhead using a 25G needle.The printing was performed with CELLINK BIO-X (CELLINK Corporation, Sweden) after 10, 60, and 150 min of keeping the gel components in the printing cartridges to obtain a suitable window of printing.To assess the injectability of hydrogels and the potential for safeguarding stem cells, a concentration of 1 × 10 6 cells mL −1 was encapsulated within the hydrogels and extruded the mixture through a needle.Prior to adding the cell culture on top of the printed constructs, the hydrogels were kept for 1 h, including the time that hydrogel was in the syringe before printing and the time spent during printing to avoid variations between different printed samples.The distribution of cells after printing was visualized using Hoechst staining.Additionally, Live/Dead staining was applied to observe the cell viability after extrusion.The extrusion was also performed using needles of varying gauges (22, 25, and 27 G) at a constant pressure of 300 kPa.Cell viability was measured after 24 h using the Lactate dehydrogenase (LDH) assay (Promega), a colorimetric assay that determines cellular cytotoxicity following extrusion.These findings were then compared with those from extruding cells using only cell culture media.
3D Printing of Hydrogels with Encapsulated Cells-Stemness of hMSCs After Printing: To observe the effects of the hydrogel and 3D printing on stemness of hMSCs, 1.5 × 10 6 hMSCs (passage number 2) cells mL −1 were encapsulated in DT-gel and added to a 48-well plate by pippeting or 3D printing (needle 25G).Then, the hydrogels were covered by cell culture media and incubated at 37 °C for three days.After 3 days, the hydrogels were dissolved by treating with 2000 U mL −1 hyaluronidase and 10 mM glutathione for 2 h.The cell suspension was isolated by centrifugation, lysed using lysis buffer, and the RNA extraction was performed using RNeasy Plus Mini Kit (Qiagen) according to the protocol.Thereafter, cDNA synthesis was performed which was then subjected to a qPCR to quantify the fold change in gene expression levels.Briefly, for qPCR reactions, the cDNA sample was added along with TaqMan Fast Advanced Master Mix (Thermo Fisher Scientific), nuclease-free water (Invitrogen), and TaqMan assay primers and subjected to the reaction process in a Bio-Rad CFX96 Real-time PCR machine as per the manufacturer's instruction.The expression levels of the following genes were analyzed using commercially available TaqMan primers: NANOG (Hs02387400), OCT4 (Hs04260367), and ACTB (Hs01060665) were considered as an internal control for qPCR.
Statistical Analysis: All the experiments were done at least in triplicates and the statistical significance between groups in various obtained results was determined using one-and two-way ANOVA analyses, using Graph-Pad Prism Software (V8), and p<0.05 was the statistical significance for all tests (**** = p < 0.0001, *** = p < 0.001, ** = p < 0.01, and * = p < 0.05).

Figure 1 .
Figure1.Synthetic strategy and cross-linking mechanism.a) Synthetic scheme for the synthesis of HA-Ald and HA-Cys using carbodiimide chemistry, by conjugating amino glycerol and cysteine derivatives on HA carboxylate; b) Mechanism and 1 H NMR analysis of thiazolidine and thiohemiacetal formation represented by the reaction between acetaldehyde and cysteine at room temperature (RT), and at cold conditions (i.e., after freezing the cysteine substrate followed by the addition of acetaldehyde solution).

Figure 2 .
Figure 2. Key findings related to hydrogel formation and kinetics of cross-linking reactions through disulfide and thiazolidine formation.a) Schematic representation of HA-Cys and HA-Ald and their reaction to form disulfide and thiazolidine linkages, demonstrating the contribution of individual crosslinking reactions by mixing different ratios of HA-Cys and HA-Ald; b) Representative images of gelation of hydrogels after 10 and 120 min by vial inversion method (The HA chains were stained with Toluidine blue for better visualization of hydrogels); c) Rheological measurements gained from time-sweep for D-gel, DT-gel, and T-gel hydrogels.* indicates a maximum of 80% of cross-linking can form in this hydrogel.

Figure 3 .
Figure 3.The effect of thiazolidine and disulfide linkages on rheological and physical properties of hydrogels.a) Rheological measurements obtained from amplitude-sweep for HA-Cys and HA-Ald based hydrogels; b) Storage modulus of hydrogels at oscillation strain of 1%; c) Swelling percentage of hydrogels in PBS over time at 37 °C over 28 days.The inset in the right corner shows the initial hydrogel swelling for the first 24 h; d) Theoretical measurement of mesh size () and average molecular weight between cross-links (Mc); e) Shear-thinning behavior of hydrogels obtained from viscosity measurement; f) The schematic representative of the contribution of thiazolidine and disulfide chemistries upon mixing of HA-Ald and HA-Cys and the impression on initial viscosity.

Figure 4 .
Figure 4. Role of disulfide and thiazolidine chemistry in self-healing and stability of hydrogels.a) Representative images of injectability D-, DT-, and T-gels after different time points in 25 mm glass slide; b) Images of hydrogels after extrusion through a needle and 24 h curing, representing the self-healing properties of different compositions; c) Reforming of disulfide bonds in D-gel group after deforming through extrusion, demonstrating the capability to tolerate the applied tensile force; d) The transparency of hydrogels in D-and T-gels after 24 h extrusion through a needle; e) Schematic illustration of self-healing properties of hydrogels, demonstrating cleavage of cross-links after injection and the ability of disulfide bonds to reform after a certain time; f) Representative images of stability of hydrogels in two different acidic (pH 4) and basic (pH 10) conditions; g and h) Degradation of hydrogels presented by the mass change in presence of glutathione and hyaluronidase respectively.

Figure 5 .
Figure 5. Biocompatibility and support of hydrogels for cell proliferation.a) 3D reconstruction of encapsulated hMSCs in hydrogels after 1 day, cells are stained with green (live) and red (dead) stains to determine cell viability; b) Viability of hMSCs in hydrogels over time calculated from microscopic images of Live/Dead stained hydrogels; c) Storage modulus of hydrogels after 3 days incubation in cell culture media with and without encapsulated cells; d) Proliferation of encapsulated cells over time determined by PrestoBlue assay.

Figure 6 .
Figure 6.Evaluation of morphology and migration of cells inside the hydrogels.a) Schematic presentation of migration assay in 3D based on a chemotaxis assay, replacing the punched area in hydrogel with encapsulated gel and gradient movement of cells toward the edges with a concentration of FBS; b) The diameter of cell movement after 24 and 48 h; c) Microscopic images of hydrogels after time pint zero (After replacing the missing part of hydrogel), and after 24 h incubation, cells were stained with Hoechst dye; d) Confocal microscopic images of hMSCs encapsulated in HA hydrogels containing gelatin after 24 h, red and blue stains indicate the cytoskeleton and nucleus of cells respectively.

Figure 8 .
Figure 8. 3D printing of hydrogels with encapsulated hMSCs.a and b) Microscopic image of printed Hoechst stained hMSCs using the hydrogel in DT-gel group; c) Microscopic images of hMSCs printed through DT-gel and stained by Live/Dead kit, presenting two different areas of structure with two magnifications; d) Proliferation rate of encapsulated cells in DT-gel and comparison between groups with injection and without; e) Comparison of cell viability after injection with cell culture media and DT-gel, using different needle sizes; f) Schematic representation of the impact of 2D and 3D on the stemness of hMSCs; g) Relative expression of NANOG and OCT3/4 as quantified by qPCR.