Development of biodegradable methacrylated guar gum 3D bioprinting bioinks for stem cell delivery and cartilage tissue engineering

Abstract

Three-dimensional (3D) bioprinting has revolutionized tissue engineering by precisely fabricating customized scaffolds that recapitulate native tissue architectures. This study introduces a photo-crosslinkable methacrylated guar gum (GG-MA) hydrogel as a tunable monophasic bioink for cartilage tissue engineering. By adjusting methacrylation degrees, GG-MA hydrogels achieved tailored mechanical strength (Young’s modulus: GG-MA2 = 0.184 MPa vs. GG-MA1 = 0.069 MPa), controlled degradation (61.41% vs. 90.71% mass loss over 60 days), and shear-thinning behavior suitable for extrusion bioprinting. Encapsulated with bone marrow mesenchymal stem cells (BMSCs), GG-MA2 scaffolds exhibited favorable biocompatibility, and promoted cell proliferation, cell migration, and chondrogenic differentiation of BMSCs, evidenced by promoting the secretion of extracellular matrix and upregulating gene expression of Collagen Type II Alpha 1 Chain (COL2A1), Aggrecan (ACAN), and SRY-box transcription factor 9 (SOX9). The novel 3D bioprinting GG-MA hydrogel scaffolds demonstrated significant potential as a versatile platform balancing biocompatibility, mechanical stability, and chondrogenic capacity for cartilage tissue engineering.

Graphical Abstract

Introduction

Articular cartilage defect, characterized by limited self-repair capacity, poses significant clinical challenges [1]. Current therapeutic approaches, such as bone marrow stimulation, cell transplantation, and osteochondral grafting, often fail to achieve satisfactory outcomes due to their inability to fully restore the structure and function of native cartilage [2]. Tissue engineering and regenerative medicine have emerged as promising fields to address these challenges [3]. Hydrogels, particularly, have garnered substantial attention as ideal biomaterials for cartilage regeneration owing to their excellent biocompatibility, tunable mechanical properties, and injectability [4]. These characteristics enable hydrogels to provide a favorable microenvironment for chondrocyte proliferation and differentiation, supporting the synthesis of cartilage matrix components [5]. As an emerging cornerstone of advanced tissue engineering, 3D bioprinting enables the architecturally precise fabrication of bioactive, cell-laden constructs that recapitulate native tissue complexity through spatially programmed biomaterial-cellular integration [6, 7]. This paradigm leverages optimized bioink systems to orchestrate cellular dynamics within engineered microenvironments, while computer-aided design capabilities facilitate patient-specific tissue analogs—a transformative convergence of fabrication fidelity, biological functionality, and personalized therapeutic potential that redefines the boundaries of regenerative medicine [8, 9].

The properties of bioinks are considered key factors in bioprinting parameters [10]. An ideal bioink should possess four main characteristics: biocompatibility, shear-thinning behavior, high viscosity gelation, and rapid crosslinking after printing [11]. Synthetic materials with excellent mechanical properties and printability are used in bioprinting, including polyethylene glycol diacrylate [12], Pluronics [13], polycaprolactone (PCL) [14], and polylactic acid (PLA) [15], which are suitable candidate materials. However, these synthetic polymers face challenges such as poor biocompatibility, low cell loading rates, and immune responses, which limit their application in tissue engineering [16]. Natural polymers such as hyaluronic acid, gelatin, collagen, alginate, and methylcellulose possess structures similar to cartilage extracellular matrix and exhibit good biocompatibility. However, their rheological properties limit their application, as low-viscosity pre-gel solutions cannot effectively maintain filamentous shapes in a monophasic system and are prone to deformation after printing [17]. Functionalized hydrogels incorporating various chemical groups (tyrosine [18], thiols [19], and vinyl sulfone [20], e.g.) enhance the printability and mechanical properties of these polymers. Methacrylic modifications are a common approach, providing polymers with ultraviolet (UV)-responsive crosslinking characteristics [21]. Photo-crosslinking, known for its high efficiency and fidelity, is currently the most widely used crosslinking method in 3D bioprinting. Methacrylate polymers can rapidly crosslink under UV irradiation, and their physicochemical properties can be adjusted by varying the degree of substitution (DS) of methacryl [22].

Guar gum is a natural polysaccharide known for its excellent viscosity, stability, and adjustable rheological properties [23]. It exhibits high hydrophilicity, dissolving completely in water at room temperature, and its viscosity can be adjusted by varying the concentration [24]. As a non-ionic polysaccharide, guar gum demonstrates good stability in aqueous solutions, remaining largely unaffected by changes in pH, ionic interaction, and temperature [25]. Notably, guar gum solutions exhibit characteristics of a non-Newtonian fluid with shear-thinning properties, making them suitable as inks for extrusion-based 3D-bioprinting [26]. Compared to conventional bioinks such as gelatin methacrylate (GelMA) and hyaluronic acid methacrylate (HAMA), guar gum methacrylate (GG-MA), modified via methacrylation, not only retains the advantageous properties of natural polymers, including excellent biocompatibility, controllable biodegradability, and non-immunogenicity, but also acquires enhanced mechanical strength and shape fidelity [27]. These improvements effectively address the inherent limitations of traditional bioinks, such as poor mechanical robustness, cumbersome degradation modulation, and inadequate cell adhesion [28]. The molecular structure of GG-MA, rich in hydroxyl groups, enables strong hydrogen bonding interactions with water molecules, leading to a significantly higher water absorption capacity compared to hyaluronic acid- or gelatin-based counterparts [29]. Moreover, GG-MA-based bioinks offer additional merits: abundant raw material availability, straightforward fabrication processes, and low production costs [30]. As a natural-origin bioink, GG-MA thus provides novel perspectives and technical foundations for 3D bioprinting bioinks and biomedical applications.

This study developed a photo-crosslinkable methacrylated guar gum (GG-MA) bioink for 3D bioprinting, demonstrating its capacity to fabricate hydrogel scaffolds with tunable mechanical properties and degradation kinetics for bone marrow mesenchymal stem cell (BMSC) encapsulation (Fig. 1). The photopolymerization mechanism enabled rapid gelation under UV exposure while maintaining precise control over scaffold architecture through digital light processing. Notably, GG-MA scaffolds exhibited adjustable compressive modulus ranging from 0.069 to 0.184 MPa via methacrylation degree, maintained high cell viability post-encapsulation, and promoted chondrocyte differentiation. The demonstrated biocompatibility and functional tunability position GG-MA as a promising candidate for 3D bioprinting bioink, cartilage tissue engineering and regenerative medicine applications.

Fig. 1.

Schematic diagram of the synthesis of photo-crosslinkable GG-MA for 3D bioprinting and cartilage tissue engineering. a Synthetic route of methacrylated guar gum (GG-MA) modified with methacrylic anhydride and used for 3D bioprinting. b 3D bioprinting GG-MA hydrogel with bone marrow mesenchymal stem cells (BMSCs) for cartilage tissue engineering

Materials and methods

Materials

Guar gum (J&K Scientific, Beijing, China), Methacrylic anhydride (Sigma Aldrich, USA), 2-hydroxy-1-(4-(hydroxyethoxy) phenyl)-2-methyl-1-propanone (Irgacure 2959, Sigma Aldrich, USA), cellulose dialysis membranes [molecular weight cutoff (MWCO), 3000 Da, Solarbio, Beijing, China]. The remaining reagents are commercially available and used as instructed.

Synthesis and characterization of GG-MA

Methacrylated guar gum (GG-MA) was synthesized using a previously described protocol with minor modifications [31]. Briefly, 1 g of guar gum (GG) powder was dissolved in 100 mL of deionized water. The solution was stirred at 50 °C until completely dissolved. Methyl acrylate (MAA) was then slowly added at molar ratios of 1:0.5 and 1:2. The pH of the solution was adjusted to 7.5 using 5 M NaOH and reacted for 6 h at 50 °C while continuously stirring. The solution was then dialyzed in deionized water for 5 days, with the water changed every 8 hours, followed by lyophilization to obtain GG-MA. The resultant hydrogels with low and high substitution rates, corresponding to molar ratios of 1:0.5 and 1:2, were designated as GG-MA1 and GG-MA2, respectively.

GG-MA were characterized by Fourier-transform infrared (FTIR) spectroscopy, proton Nuclear Magnetic Resonance (¹H NMR) spectroscopy and rheological tests to verify methacrylation and evaluate GG-MA bioink printability. FTIR analysis was performed on a PerkinElmer 100 spectrometer (Bruker, Switzerland) over 4000–400 cm⁻¹. methacrylation was confirmed by the methacryloyl group characteristic peak at 1705 cm⁻¹ (C=O stretching). ¹H NMR spectroscopy was conducted on a 600 MHz spectrometer (Agilent, USA) at 25 °C using D₂O as solvent and TMS as internal standard. The GG-MA hydrogels with a radius of 10 mm and a height of 1 mm were evaluated using a rheometer (ARES-G2, TA Instruments, USA) for dynamic frequency, oscillatory strain, and temperature sweep scans for rheological evaluation.

Evaluation of printability of GG-MA hydrogel

To evaluate the printability of the hydrogel, the 3D-Bioplotter system (Regenovo, Zhejiang, China) was employed with a pressure-controlled cartridge for 3D printing. 3D bioprinting of GG-MA hydrogel bioinks was conducted by firstly preparing a 0.1% w/v aqueous solution of photoinitiator Irgacure 2959 (I2959). GG-MA powder was weighed and dissolved in the photoinitiator solution to prepare bioinks with concentrations of 2%, 3%, 4%, and 5% w/v. Each bioink was loaded into a syringe, installed into a 3D bioprinter, and the air valve-syringe interface was fastened. 3D scaffolds were printed via computer software control of the air compressor and bioprinter, using a pre-designed model (layer height: 200 μm; infill density: 50%) and parameters: 0.05–0.20 MPa air pressure, 5–10 mm/s printing speed, and 260 μm diameter needle. The optimal bioink concentration was selected by observing printing air pressure, bioink extrusion, and scaffold formation. Two groups of hydrogel scaffolds (GG-MA1 and GG-MA2) with different methacrylation degree were fabricated for subsequent experiments. The as-printed scaffolds were irradiated with a 365 nm UV light source for a duration of 1 minute to facilitate crosslinking and solidification.

Characterization of 3D printing GG-MA hydrogel scaffolds

After the hydrogel scaffolds were completed through 3D printing, they were immediately irradiated with an ultraviolet (UV) light lamp at a wavelength of 365 nm for 1 min to crosslink the scaffolds. For microscopic morphology, the hydrogels were fabricated into cylindrical shapes with a diameter of 10 mm and a height of 5 mm, lyophilized, and then gold-coated before being observed using scanning electron microscopy (SEM, SU-8020, Hitachi, Japan). The GG-MA hydrogels were prepared as cylindrical specimens with a diameter of 10 mm and a height of 10 mm and tested in a universal testing machine (Hengyi, HY-0230, Shanghai, China) at a compression rate of 1 mm/min. The stress-strain curve was recorded, and the compressive Young’s modulus was determined from the initial 10% linear region.

For swelling evaluation, the hydrogel scaffolds of the same size were lyophilized, weighed, and recorded as dry weight (Wd). Then, they were soaked in Phosphate-Buffered Saline (PBS) solution and placed in a 37 °C shaking incubator. Scaffold samples were collected at 2 h, 4 h, 8 h, 12 h, 24 h, 36 h, and 48 h, weighed, and recorded as wet weight (Ws). The swelling ratios were calculated according to the following equation:

$$\mathit{Swelling}(\%)=\frac{\mathrm{W}s-\mathrm{W}d}{\mathrm{W}d}\times 100\%$$ 1

In degradation assessment, the hydrogel scaffolds were lyophilized and weighed (W0). Subsequently, the samples were immersed in PBS and incubated in a shaking incubator at 37 °C. After specified time points, the samples were removed, lyophilized, and weighed (W_t). The degradation ratios were calculated according to the following equation:

$$\mathit{Degradation}\mathit{ratio}\left(\%\right)=\frac{\mathrm{W}0-\mathrm{W}t}{\mathrm{W}0}\times 100\%$$ 2

In vitro cell culture

The bone marrow mesenchymal stem cells (BMSCs) were harvested from the femoral marrow of New Zealand white rabbits (aged 3–7 days) according to our previous study. The ethical guidelines of the Animal Ethics Committee of Guangxi Medical University conducted all procedures (Approval No. 202409915). The isolated cells were cultured in α-MEM medium (Gibco, USA) supplemented with 10% fetal bovine serum (Zhejiang Tianhang Biotechnology Ltd., China) and 1% penicillin/streptomycin (Solarbio, China) at 37 °C, 5% CO₂ in an incubator. BMSCs at passages 3–5 were utilized in subsequent experiments.

Cytotoxicity assay

The BMSCs at a density of 1 × 10⁷ cells/mL were resuspended in a GG-MA solution containing 1% Irgacure 2959 to prepare a bioink, which was then loaded into the cartridge of a 3D printer. The printing was performed using a nozzle with a diameter of 0.26 mm. The printed scaffolds were immediately subjected to UV light irradiation (365 nm, 6 mW/cm²) for 1 min to achieve crosslinking, resulting in stable cell-laden 3D scaffolds. Subsequently, the cell-laden 3D scaffolds were placed in a culture medium and incubated at 37 °C, 5% CO₂, with the medium replaced every 3 days. After 7 and 14 days of incubation, cell viability was assessed using a Live/Dead staining kit (Life Technologies, USA). Images were acquired using a laser confocal microscope (Leica, Germany).

Cell migration assays

To validate the impact of GG-MA on the migration of BMSCs, a transwell migration assay was conducted following a previously reported protocol [32]. In brief, the BMSCs were digested with 0.25% trypsin and resuspended in a complete medium (Corning, USA) supplemented with 10% fetal bovine serum. Subsequently, 200 μL of serum-free medium, containing 1 × 10⁵ cells, was added to the upper chamber (Corning Incorporated, USA; pore size: 8 μm), while the lower chamber was filled with 600 μL of the differently treated medium. Sterilized hydrogel samples were incubated in a complete medium with 10% fetal bovine serum for 24 h to prepare the various media for the lower chamber. After incubation, cells in the upper chamber were removed, and the lower surface of the transwell was fixed with 4% paraformaldehyde (PFA, Sigma, USA). Post-staining with 0.1% crystal violet (Sigma, USA), the stained cells on the bottom surface of the transwell were observed under a light microscope. Three random fields of view were selected, and the number of stained cells was quantified using Image J software.

Chondrogenic differentiation

BMSCs-scaffold complexes of the AS, GG-MA1, and GG-MA2 were cultured in chondrogenic induction medium for 7 days, and then were washed with PBS, fixed with 4% paraformaldehyde for 15 min, 0.1% Triton X-100 permeabilized for 10 min, and closed with 5% Bovine Serum Albumin (BSA, Solarbio, China) for 1 h. Rabbit anti-COL2A1 primary antibody (1:200, Abcam, ab34712) was added and incubated overnight at 4 °C, followed by incubation with Alexa Fluor 488-labeled goat anti-rabbit secondary antibody (1:500, Invitrogen, A11008) for 1 h at room temperature away from light. Cell nuclei were restained by 4’,6-Diamidino-2-Phenylindole (DAPI, 1:1000, Sigma) for 5 min. Images were acquired using a laser confocal microscope (Leica TCS SP8), and fluorescence intensity was quantified by ImageJ software (n = 3).

The quantification of DNA and glycosaminoglycan (GAG) content were tested and standard curves were established using calf thymus DNA and chondroitin sulfate, respectively. For DNA content, a series of DNA standards were prepared and mixed with Hoechst 33258 dye, followed by fluorescence measurements to generate a standard curve. BMSCs were co-cultured with hydrogel samples in a chondrogenic induction medium for 7 days. The samples were collected at 7 days, and were processed with protease K. The GAG content was measured with 1,9-dimethylmethylene blue (DMMB) (Sigma-Aldrich, USA) using the chondroitin sulfate as standard curve.

To assess chondrogenic genes expression, the samples were harvested after 7 days of co-culture with hydrogels in a chondrogenic induction medium. Real-time quantitative polymerase chain reaction (qRT-PCR) was used to assess the genes expression of ACAN, COL1A1, COL2A1, and SOX9. Total RNA was extracted and purified and reverse transcribed to complementary DNA (cDNA) using PrimeScript RT reagent Kit (TaKaRa, Japan). Amplification was performed using the LightCycler^® 96 system to assess the expression levels of ACAN, SOX9, COL2A1, and COL1A1. The specificity of the PCR reaction was confirmed by melting curve analysis. The relative expression of each target gene was calculated using the 2^−ΔΔCt method with Glyceraldehyde-3-Phosphate Dehydrogenase (GAPDH) as an internal reference gene. Each PCR experiment was replicated three times. The forward and reverse primer sequences for each gene are shown in Table 1.

Table 1.

Primer sequence list

Gene	Forward primer (5’-3’)	Reverse primer (5’-3’)
GAPDH	GGCTCTCTGCTCCTCCCTGT	CTGTGCCGTTGAACTTGCCG
ACAN	TTCGGCAAGCTACCTTCTGG	TTCTCCTGAACCACTGACGC
COL2A1	GTCCTACAATGTCAGGGCCA	ACCCCTCTCTCCCTTGTCAC
COL1A1	GATCCTGCCGATGTCGCTAT	GGGACTTCTTGAGGTTGCCA
SOX9	TCCAGCAAGAACAAGCCACA	CGAAGGGTCTCTTCTCGCTC

Statistical analysis

Data were analyzed using SPSS 20.0 and are presented as mean ± standard deviation (SD). A t-test was conducted for single-sample comparisons, while one-way ANOVA was used for two-by-two comparisons between groups. A p-value of <0.05 was considered statistically significant. * denotes statistical differences between this experimental group compared with the blank group, *p < 0.05, **p < 0.01, ***p < 0.001 and ****p < 0.0001; # denotes statistical differences between comparisons experimental groups, #p < 0.05, ##p < 0.01, ###p < 0.001 and ####p < 0.0001, n = 3.

Results

Characterization of GG-MA and rheological properties of GG-MA hydrogels

GG-MA was successfully synthesized via the protocol outlined in Fig. 1, with its structure validated by Fourier-transform infrared spectroscopy (FTIR) and proton nuclear magnetic resonance (¹H NMR). As shown in the FTIR spectrum (Fig. 2a), GG-MA exhibited a characteristic stretching vibration band of the methacrylate C=O bond at 1705 cm⁻¹ [33]. In the ¹H NMR spectrum (Fig. 2b), two distinct new peaks emerged at 5.7 ppm and 6.1 ppm, assigned to the vinyl protons of the methacrylate C=C bond, alongside a peak at 1.9 ppm corresponding to the methyl group. Using the ¹H NMR data, the degree of methacrylate (MA) substitution (DS) on the GG backbone was quantified [34]. The DS values for GG-MA1 and GG-MA2 were determined to be 6% and 13%, respectively.

Fig. 2.

Characterization of GG-MA and rheological properties of GG-MA hydrogels. a ¹H NMR and b FTIR spectra of the GG and GG-MA hydrogels: Peaks at 5.6 and 6.1 ppm confirm the methacrylation of GG, while the peak at 1.9 ppm indicates the presence of methyl protons on the GG backbone. The arrow represents a new peak at 1705 cm⁻¹ that appears due to the stretching band of the C=O bond. Rheological properties of GG-MA: c The viscosity and shear-thinning curves, d temperature scanning curves, and e oscillatory frequency sweep curves of GG-MA1 and GG-MA2

The rheological properties of hydrogels are crucial parameters for evaluating the printability of bioinks [35]. The rheological behavior of GG-MA hydrogels was systematically characterized through three orthogonal testing protocols: oscillatory strain sweep (0.1–100% strain, 1 Hz), temperature ramping (20–40 °C, 1 °C/min), and oscillatory frequency sweep (0.1–100 Hz, 1% strain). As shown in Fig. 2c, both GG-MA1 and GG-MA2 hydrogels exhibited pronounced shear-thinning rheological behavior, and their viscosities decreased as the shear rate increased. Across the entire shear rate range, GG-MA2 had a notably higher storage modulus (G′) than GG-MA1. More importantly, for both formulations, G′ was far greater than the loss modulus (G′′) at all shear rates. This G′′ dominated relationship not only verified their elastic nature but also indicated excellent injectability, which is a key requirement for extrusion-based 3D bioprinting. The temperature scanning mode provides insights into the viscoelastic behavior of the hydrogels at different temperatures. As shown in Fig. 2d, between 20 °C and 40 °C, the G′ and G′′ of both GG-MA concentrations remained almost unchanged, with G′ consistently higher than G′′. This stability shows that the GG-MA hydrogels have good linear viscoelastic stability when exposed to physiological temperature variations. During the oscillatory-frequency test (Fig. 2e), as the angular frequency increased from 0 to 100 rad/s, the G′ and G′′ of both GG-MA hydrogels showed minimal fluctuations. Moreover, G′ was always significantly larger than G′′. This result demonstrates that the GG-MA hydrogels possess excellent mechanical stability, as their elastic-dominated behavior is maintained even under dynamic frequency conditions. In summary, regardless of the grafting rate, the GG-MA hydrogels exhibit stable rheological properties, such as shear-thinning for easy injection, thermal viscoelastic stability, and frequency-independent mechanical robustness. These features make them highly suitable as bioinks for 3D bioprinting.

Evaluation of printability of GG-MA hydrogels

Extrusion-based 3D bioprinting critically depends on achieving optimal balance between extrusion pressure, material flowability, and filament integrity [36]. To evaluate the printability of GG-MA hydrogels, we systematically characterized their rheological behavior and extrusion performance across concentrations (2–5% w/v) without UV light crosslinking. As demonstrated in Fig. 3, 2% and 3% GG-MA exhibited insufficient cohesive strength under extrusion, forming discontinuous droplets rather than stable filaments at needle tips. This behavior correlated with their limited shear-thinning capacity, which failed to maintain continuous material flow under controlled pressure conditions. In contrast, 4% and 5% GG-MA formulations demonstrated superior extrusion characteristics. The 4% formulation achieved stable filament output at remarkably low pressure (0.08 MPa), maintaining stability of the shape over six-layer deposition in 5 min. The 5% formulation required elevated pressure (0.18 MPa) to initiate flow, suggesting increased viscoelastic resistance that could compromise print fidelity through nozzle clogging risks. This systematic evaluation establishes 4% GG-MA as the optimal bioink concentration for high-fidelity bioprinting, balancing rheological requirements with practical extrusion constraints. Subsequent experiments utilized this formulation for complex scaffold fabrication to validate its biomedical application potential.

Fig. 3.

Evaluation of printability of GG-MA hydrogel

Characterization of 3D printing GG-MA hydrogel scaffolds

Scanning electron microscopy (SEM) analysis revealed that 3D-printed GG-MA hydrogel scaffolds exhibited a hierarchical interconnected porous architecture (Fig. 4a). The GG-MA2 formulation (5% w/v) exhibited a degree of methacrylation (DM), which induced increase in crosslinking density compared to GG-MA1. This enhanced crosslinking network resulted in smaller pore diameters (125.65 ± 56.81 μm vs.189.95 ± 61.26 μm) (Fig. 4b) and higher pore density (2.4 × 10⁶ pores/cm³ vs. 1.5 × 10⁶ pores/cm³) (Fig. 4c). The increased crosslinking density consequently improved the Young’s modulus of GG-MA2 scaffolds to 0.184 MPa, demonstrating a 2.7-fold enhancement over GG-MA1 (0.069 MPa) (Fig. 4d). This mechanical reinforcement originated from the denser covalent crosslink network formed through methacrylate group polymerization, which effectively transferred stress across the scaffold structure. These results demonstrate that the mechanical strength of GG-MA hydrogel scaffolds can be controlled by adjusting the substitution rate of methacrylate groups during the modification reaction. Such optimized mechanical properties are critical for maintaining three-dimensional spatial fidelity during cell encapsulation and subsequent tissue regeneration processes. The degradation profiles of GG-MA1 and GG-MA2 hydrogel scaffolds were systematically evaluated through weight loss analysis (Fig. 4e). Over 60 days, GG-MA1 exhibited a degradation rate of 90.71%, while GG-MA2 demonstrated a significantly slower degradation rate of 61.41% (p < 0.01). This controlled degradation behavior of GG-MA2 offers sustained structural integrity critical for tissue engineering applications requiring prolonged mechanical support. The swelling characteristics were quantified through equilibrium swelling tests (Fig. 4f). Both scaffolds achieved swelling equilibrium within 24 h, with GG-MA1 exhibiting a swelling ratio of 1917% and GG-MA2 demonstrating a controlled swelling ratio of 1139%. The observed results revealed that GG-MA possesses three-dimensional pore structures, excellent mechanical properties, and controlled degradation behavior.

Fig. 4.

Characterization of 3D-printed GG-MA scaffolds. a SEM images of microstructure of GG-MA1 and GG-MA2 hydrogels. Bar = 500 μm. b Pore diameters and c pore density of 3D-printed GG-MA1 and GG-MA2 hydrogels. d The Young’s modulus of GG-MA1 and GG-MA2 hydrogels. e Degradation ratio and f swelling ratio of 3D-printed GG-MA1 and GG-MA2 hydrogels. Experimental values are expressed as mean ± SD (n = 3). ^**p < 0.01, ^***p < 0.001, and ^****p < 0.0001

Cytocompatibility of 3D printing GG-MA hydrogel scaffolds

Cell viability evaluation is critical for validating bioink performance in 3D bioprinting applications [37]. Following a 3D culture of BMSCs loaded onto GG-MA scaffolds, Live/Dead staining was conducted to evaluate cytotoxicity, as illustrated in Fig. 5. On day 7 post-culture, a substantial population of viable cells was observed within GG-MA scaffolds, which was superior to that of widely used bioprinting inks such as alginate sodium (AS). By day 14, with prolonged culture duration, both the quantity and density of cells across all groups progressively increased, indicating successful adhesion and proliferation of cells within the three-dimensional scaffolds demonstrating favorable biocompatibility. Notably, the GG-MA2 group displayed a significantly higher cell proliferation rate, which may be attributed to its enhanced crosslinking density. This improved structural integrity effectively maintained scaffold architecture, preventing pore collapse and thereby promoting cell migration and nutrient diffusion.

Fig. 5.

Live/dead staining of BMSCs encapsulated in alginate sodium (AS) and GG-MA hydrogels after 7 and 14 days of 3-D culture. Calcein-AM stains live cells (green), while PI stains dead cells (red). Bar = 400 μm

Cellular migration of BMSCs in 3D printing GG-MA hydrogel scaffolds

The migratory capacity of BMSCs loaded onto GG-MA scaffolds using a transwell assay as illustrated in Fig. 6a. The results demonstrated that the GG-MA group exhibited a significant enhancement in BMSCs migration compared to the AS group (Fig. 6b). Specifically, the number of migrated cells in the GG-MA2 group was 3.06-fold higher than that in the AS group, indicating that GG-MA2 had a significant promoting effect on cell migration (Fig. 6c). This result suggested that GG-MA2 with superior crosslinking density may optimize scaffold bioactivity through increased surface charge density facilitating BMSC adhesion and sustained release of bioactive moieties that promote cell adhesion and migration. This finding provides an important experimental basis for the application of GG-MA serving as an effective biomaterial to promote the BMSCs migration in tissue engineering and regenerative medicine.

Fig. 6.

GG-MA promotes cell migration. a Schematic diagram of cellular migration assay. b Images of migrated BMSCs, Bar = 500 μm. c Quantitative analysis of migrated cells using Image J software. Experimental values are expressed as mean ± SD (n = 3). ^****p < 0.0001 (vs. AS group); ^##p < 0.01 (intergroup comparisons)

The chondrogenic differentiation of BMSCs in 3D printing GG-MA hydrogel scaffolds

To evaluate the chondrogenic differentiation of BMSCs in the GG-MA hydrogels, we conducted immunofluorescent staining and quantitative GAG detection. COL2A1 (type II collagen) is a key marker gene and important effector molecule during the chondrogenic differentiation process of BMSCs. Immunofluorescence analysis (Fig. 7a, b) revealed that the GG-MA2 group exhibited significantly enhanced COL2A1 expression intensity compared to both the AS group and the GG-MA1 group. Quantitatively, the fluorescence intensity in the GG-MA2 group was 2.7-fold higher than that in the AS group and 1.3-fold greater than that in the GG-MA1 group. The quantitative GAG/DNA content (Fig. 7c) revealed that the GG-MA2 group exhibited significantly higher values compared to both the AS and GG-MA1 groups, indicating that the GG-MA2 scaffold markedly enhanced the GAG synthesis efficiency of BMSCs. qRT-PCR analysis (Fig. 7d–g) demonstrated that the GG-MA hydrogels exhibited markedly elevated expression levels of chondrogenic markers ACAN, COL2A1, and SOX9 compared to the AS group. Quantitatively, the genes’ transcriptional level in the GG-MA2 group reached 9.73-fold, 16.83-fold, and 1.37-fold enhancements for ACAN, COL2A1, and SOX9, respectively, when contrasted with the AS group. Conversely, the gene expression of COL1A1, which was a fibrosis marker in the GG-MA2 group, showed a significant 62% reduction compared to the AS group, indicating a pronounced suppression of fibrotic markers during chondrogenesis. These findings indicated that the high cross-linking density of GG-MA hydrogels better promote chondrogenic differentiation of BMSCs, suggesting a promising application potential for GG-MA in cartilage tissue engineering.

Fig. 7.

The chondrogenic differentiation of BMSCs in GG-MA hydrogels. a Immunofluorescence of COL2A1, Bar = 75 μm. b Quantitative analysis of immunofluorescence. c Quantitative GAG/DNA content in AS, GG-MA1and GG-MA2 hydrogels. d–g qRT-PCR assessment of chondrogenesis-related gene expression. Experimental values are expressed as mean ± SD (n = 3). ^**p < 0.01, ^***p < 0.001, and ^****p < 0.0001 (vs. AS group); ^##p < 0.01, and ^###p < 0.001 (intergroup comparisons)

Discussion

3D bioprinting facilitates the customization of tissue engineering scaffold shape and anisotropic properties tailored to specific cartilage regeneration requirements [38]. In this study, we developed a novel photo-crosslinked GG-MA hydrogel, which exhibits considerable potential as a bioink for 3D bioprinting. 2% and 3% GG-MA exhibited insufficient cohesive strength under extrusion, forming discontinuous droplets rather than stable filaments at needle tips. This behavior correlated with their limited shear-thinning capacity, which failed to maintain continuous material flow under controlled pressure conditions [39]. The 4% w/v GG-MA was identified as the optimal concentration. The 4% w/v GG-MA hydrogel not only had superior printability (encompassing favorable extrudability, shape fidelity, and structural stability) but also excellent biological and physicochemical characteristics conducive to cartilage tissue engineering. Specifically, 4% w/v GG-MA hydrogel provides a suitable microenvironment that balances matrix stiffness and cell adhesion, thereby supporting optimal cell viability and promoting chondrogenic differentiation of BMSCs [40]. The gelation kinetics of 4% w/v GG-MA are compatible with 3D bioprinting processes: methacrylated guar gum chains form a crosslinked network at an appropriate rate under UV irradiation, thereby avoiding nozzle blockage or shape collapse and ensuring the coordination of printing efficiency and scaffold stability. Though 5% GG-MA formulations also demonstrated superior extrusion characteristics which required elevated pressure (0.18 MPa) to initiate flow with increased viscoelastic resistance, but elevated pressure might have a negative impact on cell activity. This systematic evaluation establishes 4% GG-MA as the optimal bioink concentration for high-fidelity bioprinting, balancing rheological requirements with practical extrusion constraints. Subsequent experiments utilized this formulation for complex scaffold fabrication to validate its biomedical application potential.

Guar gum, a natural polysaccharide characterized by excellent cell compatibility, exhibits substantially enhanced mechanical properties following methacrylation modification, and its rapid photopolymerization capability facilitates the fabrication of high-fidelity cartilage tissue engineering scaffolds [41]. Methacrylation markedly enhances the mechanical strength and stability of guar gum. The increased crosslinking density consequently improved the Young’s modulus of GG-MA2 scaffolds to 0.184 MPa, which was higher than that of methacrylated gellan gum hydrogel (6.4–20 kPa) or methacrylated gelatin hydrogel (1.9–150 kPa). This mechanical reinforcement originated from the denser covalent crosslink network formed through methacrylate group polymerization, which effectively transferred stress across the scaffold structure [42]. Compared to methacrylated hyaluronic acid (MEHA), GG-MA possesses superior mechanical strength and stability, overcoming the inherent limitations of MEHA (i.e., weak gel strength and rapid degradation). The GG-MA exhibits better BMSC adhesion without the need for additional adhesive ligands in comparison to another methacrylated polysaccharide methacrylated gellan gum (MGG). The performance of GG-MA hydrogel scaffolds can be controlled by adjusting the substitution rate of methacrylate groups during the modification reaction, and demonstrates good thermal viscoelastic stability, comparable biocompatibility, superior degradation controllability, and lower production costs, supporting its potential for biomedical applications [43].

The mechanical properties of scaffolds exert a profound influence on proliferation and chondrogenic differentiation of BMSCs [44]. The moderate methacrylation degree ensures the exposure of sufficient cell-adhesive sites (e.g., hydroxyl groups) that can specifically bind to integrin receptors (e.g., α5β1, αvβ3) on the BMSCs surface, triggering the formation of focal adhesions that connect the extracellular matrix (ECM) to the intracellular cytoskeleton [45]. Excessively high methacrylation leads to over-crosslinking of GG-MA chains, reducing the accessibility of cell-adhesive sites and increasing matrix rigidity, thereby inhibiting integrin binding [46]. In contrast, excessively low methacrylation results in insufficient crosslinking, leading to inadequate matrix stiffness and unstable ECM structure, which fails to provide effective mechanical signals for integrin-mediated mechanotransduction [47]. The methacrylation degree of GG-MA might regulate cellular functions through its effects on integrin binding and mechanotransduction [48]. Once activated via binding to GG-MA, integrins transmit extracellular mechanical signals into the cell through downstream signaling pathways (e.g., FAK-RhoA-ROCK), regulating cytoskeleton rearrangement, BMSC behavior, and ECM secretion, which are tightly modulated by the methacrylation degree [49].

While UV crosslinking has demonstrated efficacy in solidifying glycidyl methacrylate-modified GG-MA hydrogels for cartilage engineering applications, its clinical translation faces some limitations that demand systematic evaluation [50]. The residual cytotoxicity of conventional UV initiators (e.g., I2959) may induce oxidative stress and cell damage through reactive oxygen species (ROS) generation [51]. The limited penetration depth (~200 μm) of UV light restricts the fabrication of thick tissue engineering scaffolds, and may create gradient crosslinking densities that weaken mechanical integrity and nutrient diffusion efficiency in thick constructs, which compromises overall mechanical and biological performance [52]. Additionally, prolonged UV irradiation may induce DNA damage in encapsulated MSCs, impairing cell viability and chondrogenic differentiation. These limitations highlight the need for alternative crosslinking approaches (e.g., visible light crosslinking, dual crosslinking systems) to enhance the applicability of GG-MA bioink in cartilage tissue engineering [53].

Conclusions

In this study, we developed a novel photocrosslinked methacrylated guar gum (GG-MA) hydrogel, a promising tunable hydrogel candidate with tailored mechanical properties, favorable biocompatibility, and excellent printability, which advances the development of functional bioinks for 3D bioprinting in cartilage tissue engineering. This work provides a viable strategy for customizable hydrogel scaffold using natural polysaccharide-based bioinks for in vivo cartilage repair validation and individualized clinical application of cartilage regeneration.

Funding

This work was supported by the Guangxi Scientific Research and Technological Development Foundation (Grant No. GuikeAB23026049, AB25069177), the Guangxi Natural Science Foundation (No. 2019GXNSFAA185060), and the National Natural Science Foundation of China (Grant No. 82360426).

Data availability

All data generated or analyzed during this study are included in this published article.

Compliance with ethical standards

Conflict of interest

The authors declare no competing interests.