Development of a two-layer 3D equine endometrial tissue model using genipin-crosslinked collagen scaffolds and 3D printing | Scientific Reports

Scientific Reports volume 15, Article number: 19759 (2025) Cite this article Metrics details Advances in endometrial tissue engineering have enabled the combination of modified scaffolding materials

Scientific Reports volume 15, Article number: 19759 (2025) Cite this article

Metrics details

Advances in endometrial tissue engineering have enabled the combination of modified scaffolding materials with modern cell culture technologies. Genipin and three-dimensional (3D) printing have advanced cell-tissue engineering by enabling the precise layering of cell-containing matrices while ensuring low cytotoxicity. This study aimed to advance equine endometrial tissue engineering by designing customized collagen scaffolds using 3D printing technology, while optimizing the genipin concentration to avoid toxicity. Genipin was tested at concentrations of 4, 2, 1, 0.5, 0.25, 0.125, and 0 mM on equine endometrial epithelial cells (eECs) and mesenchymal stromal cells (eMSCs). Its effects on cell morphology and scaffold properties were evaluated in collagen-based conventional equine endometrial tissue (3D-ET) by assessing percentage of cells spreading within each genipin concentration. Additionally, genipin-collagen scaffolds at 2, 1, 0.5, 0.25, and 0 mM were analyzed for viscoelastic properties using rheological testing. Based on these assessments, 0.5 mM genipin was identified as the optimal concentration and was to develop in vitro 3D-ET. Key 3D printing parameters, including extrusion pressure, printing temperature, pre-printing time, and velocity, were optimized. The structural integrity of the advanced 3D-ET was assessed via phase contrast microscopy. Cellular characterization was performed using Pan-cytokeratin and Vimentin staining. For the characterization of printed 3D-ET, mucin production was assessed using Alcian blue staining, while estrogen receptor alpha (ERα) expression was evaluated by immunofluorescence. A study of oxytocin-stimulated prostaglandin synthesis capacity was performed in an advanced 3D-ET for 24 h, and expression of key genes was analyzed quantitatively using real-time PCR. Genipin exhibited dose-dependent toxicity, with 0.5 mM identified as the optimal concentration based on its support of proliferative activity, cell morphology, and viscoelastic properties. Only eMSCs were successfully 3D-printed in a collagen scaffold with 0.5 mM genipin. While the 3D-printed biomaterial failed to support eECs viability; eECs survived and formed glands only when a conventional seeding method was used. Consequently, a dual-layer 3D-ET model was developed in which eMSCs were printed with 0.5 mM genipin-collagen, and eECs were overlain using conventional methods. This model preserved the structural integrity necessary for glandular-like development and maintained the functional characteristics of equine endometrial tissue. Mucin production was observed, while ERα was detected in the cytoplasm and translocated into the nucleus.Notably, after OT challenge prostaglandin-endoperoxide synthase 2 (PTGS2) expression was significantly elevated in the treatment group compared to controls (p < 0.05). This advanced 3D-ET model offers a robust platform for studying tissue-specific functions and could be further developed by incorporating immune or endothelial cells or creating complex structures such as glands or vessels.

Tissue engineering (TE) strategies using scaffold materials, cells, and bioactive factors offer considerable potential for developing more physiological uterine models. Currently, they often fail to adequately replicate the complexity of in vivo structure and function of the endometrium1. Collagen-based hydrogels are promising biomimetic matrices for replicating the endometrial tissue microenvironment2 due to their ability to deliver bioactive molecules and closely mimic the composition and stiffness of the native extracellular matrix (ECM)3. These hydrogels have been shown to effectively support the co-culture of endometrial epithelial organoids with stromal cells in vitro for several mammalian species4,5,6. Our previous study7 demonstrated that collagen hydrogels facilitate de novo three-dimensional (3D) reconstruction of equine endometrial tissue (3D-ET), while maintaining the functional properties of the incorporated endometrial cells. However, challenges such as scaffold instability and difficulties in controlling scaffold shape, and associated limitations in the uniform distribution of the culture medium, have hindered their use in long-term in vitro studies.

To address these issues, genipin, a natural crosslinker derived from Gardenia jasminoides,8 presents a promising alternative. is a promising stabilizing molecule. Genipin enhances the mechanical properties of bioartificial tissues9 by crosslinking collagen through nucleophilic attacks on lysine and arginine residues, leading to more stable tissue matrices with tunable stiffness10 but minimal cytotoxicity. Although genipin has demonstrated its potential for supporting cellular and tissue functions in other models, such as corneal and bone constructs11, its application in endometrial models is underexplored. In this respect, it is essential to balance the genipin concentration to optimize scaffold stiffness for a specific tissue type to maintain proper cellular function.

3D-bioprinting offers a powerful method to bridge the gap between artificial and natural tissues by layering cell-laden materials into 3D structures8. It provides advantages such as geometric freedom, automation, reproducibility, and customizability12, allowing precise cell and biomaterial deposition into pre-designed architectures, unlike conventional tissue engineering methods13. Bioprinting approaches include extrusion, jetting, and vat polymerization, each with distinct advantages and limitations. Extrusion bioprinting deposits bioink layer by layer using mechanical or pneumatic actuation, requiring crosslinking for mechanical stability. While it accommodates a wide range of materials, its limitations include low resolution and shear stress on cells14. Jetting-based bioprinting enables contactless deposition with high precision and reduced shear stress but is restricted by bioink compatibility and shape fidelity, making it suitable for thin structures like skin and retina15. Vat polymerization produces high-resolution, complex structures via light-induced photopolymerization but necessitates specialized bioinks with potentially cytotoxic photoinitiators16. Given these factors, extrusion-based bioprinting was selected for this study due to its material versatility, avoidance of specialized chemicals, and suitability despite its lower resolution.For endometrial tissues, 3D bioprinting has demonstrated potential for fabricating intricate microstructures using bioinks like alginate and gelatin1,17,18,19. Although collagen-based bioinks show promise, particularly in placental20, skin21,22 and lung23,24 models, their use in bioprinting reproductive tissues and organs remains largely unexplored. Further research could leverage the power of 3D-bioprinting to develop complex endometrial tissues that closely mimic their in vivo counterparts, thereby overcoming the limitations of conventional 3D-ET models. By customizing the composition of these tissue constructs, 3D-bioprinting offers the potential to create more accurate models for studying reproductive health and advancing tissue engineering applications. This innovation could provide significant benefits by enabling more precise investigation of the complexities of endometrial structure and function.

This study aimed to enhance 3D in vitro equine endometrial tissue culture by determining the optimal concentration of genipin to enhance the stability of collagen scaffolds in a 3D-ET model without compromising structural and functional properties. This would improve the mechanical strength and biological functionality, making the model more suitable for precise 3D bioprinting of equine endometrial tissue, which could be used to develop advanced models for the study of endometrial disorders and the development of novel therapeutic strategies.

The animal care and use protocol for this study was approved by the Chulalongkorn University Animal Care and Use Program (CU-ACUP, Protocol No. 2131028) which complied with the ARRIVE guidelines. All biosafety procedures were conducted in accordance with the guidelines set by the Institutional Biosafety Committee of the Faculty of Veterinary Science, Chulalongkorn University (IBC Protocol No. 1631022).

eMSCs and eECs were isolated from fertile mares (1.5-7 years old) during early diestrus (days 2–7 post-ovulation) and cultured, as described by Santiviparat et al. (2024). The eMSCs were incubated in low-glucose DMEM supplemented with 10% fetal bovine serum (FBS), 1% L-glutamine, and 1% antibiotic/antimycotic solution, while eECs were cultured in a semi-defined medium (EC medium) containing MSC medium supplemented with 0.01 µg/ml epidermal growth factor (EGF), 2.436 mM hydrocortisone and 10 mM Rho-associated coiled-coil kinase (ROCK) inhibitor (Y-27632). One endometrial cell line that fully characterized for endometrial traits, as reported by Santiviparat (2024), was used in this study. Cells from passage 5 (eMSCs) and passages 6–7 (eECs) were utilized.

(A) Schematic representation of the overall experimental design, divided into 4 main steps. (B) Schematic representation of genipin crosslinking with collagen.

To assess genipin toxicity, eECs and eMSCs were cultured in previously described EC and MSC media with varying concentrations of genipin (4, 2, 1, 0.5, 0.25, 0.125, and 0 mM) in 96-well plates (18,000 cells/well, three replicates per condition). The MTT (3- [4,5-dimethylthiazol-2-yl]-2,5 diphenyl tetrazolium bromide) assay was used to measure cellular metabolic activity. Cells were incubated for up to 7 days, and proliferative activity was assessed daily using the MTT assay. At designated time points (24, 72, 120, 168 h), culture medium was replaced with MTT reagent. After incubation for 3 h, dimethyl sulfoxide (DMSO) was used to solubilize formazan crystals. Optical density (OD) was measured at 570 nm wavelength using a microplate reader. A standard curve was generated by seeding serial dilutions of eECs and eMSCs (24,000 to 375 cells/well) and analyzing OD after 24 h with the MTT assay.

Each genipin-collagen scaffold concentration was tested in triplicate to assess eEC and eMSC morphology.

The sterilized collagen basement membrane (BM) (Nitta Gelatin Inc., Japan, Lot no. 190709) was prepared using a modified protocol from a previous study (Santiviparat et al., 2024). Briefly, the collagen BM was neutralized with 200 mM HEPES buffer and adjusted to a pH of 7.0–7.5 using 10 mM sterile NaOH. The solution was gently stirred with a magnetic stirrer under sterile conditions for 5 min and stored at 4 °C until use. The collagen was then divided into 7 portions and mixed with genipin at concentrations of 4, 2, 1, 0.5, 0.25, 0.125, and 0 mM. The mixtures were vortexed and centrifuged at 4 °C for 2 min before further use.

Rheological properties of collagen scaffolds crosslinked with genipin at concentrations of 0, 0.25, 0.5, 1, and 2 mM were evaluated using a rotational rheometer (Kinexus Ultra Plus, Malvern Instruments) at the National Metal and Materials Technology Center (MTEC), Thailand. A frequency sweep test was performed with 20 mm parallel plates at a 1.00 mm gap under controlled stress conditions. Measurements were conducted at 25 °C over a frequency range of 0.01–100 rad/s. A strain sweep test was conducted to determine the linear viscoelastic region, with shear strain ranging from 0.1 to 1.0%. Based on these results, a shear strain of 1.0% was selected for subsequent measurements.

eECs and eMSCs were combined at a 1:1 ratio to a total of 150,000 cells and suspended in collagen containing varying genipin concentrations (total volume of 500 µL or 0.965 cm³). The mixture was added to a 24-well plate, and 200 µL of phosphate buffered saline (PBS) was used to cover the collagen in each well. The plate was incubated at 37 °C in a humidified atmosphere of 5% CO₂ for 1.5 h to allow collagen polymerization. After gel formation, the PBS was removed, and 500 µL of EC medium was added to each well. The culture was maintained for 7 days in EC medium, with medium changes every 2–3 days. On day 3, a cell count and morphological assessment were performed on randomly selected fields (10 areas at 10x magnification) for each genipin concentration using a phase-contrast microscope (CK X41 Olympus, Japan). On day 7, the final morphological evaluation of the 3D-ET was conducted for each genipin concentration.

Optimization of 3D printing parameters for collagen-based scaffolds with different genipin concentrations.

An in-house developed 3D bioprinter was used in this study. A cartesian robot (Snapmaker 1.0) was equipped with a custom-made syringe pump printhead. The printhead utilized a DC servo motor coupled with a ball screw to drive a disposable syringe (Nipro 10 cc Luer-Slip syringe) and extrude the contained material through tapered plastic dispensing tips (18 gauge). During extrusion, the printhead controlled the volumetric flow rate of material in a positive displacement manner by using closed loop control of the motor’s rotation. As demonstrated in a previous study, this printhead allowed the extrusion pressure to be estimated and logged in real-time25. During extrusion, the custom-made embedded printhead controller measured the electrical current supplied to the motor, continuously. Utilizing a mathematical model of the printhead action together with the measured current, the extrusion pressure was then calculated. The bioprinter was equipped with a temperature control system consisting of two parts: (a) a syringe heater, custom-designed to cover the syringe, which could regulate the temperature of the printing material inside the syringe within the 25˚C to 70˚C range. (b) a print bed heater, custom-designed to interface with the underneath of a 60 mm petri dish, which could regulate its temperature between 25˚C and 70˚C.

Collagen-based scaffolds with genipin cross-linking were tested in triplicate.

for each genipin concentration to optimize and validate 3D printing parameters.

To achieve favorable print quality, the optimal extrusion parameters within the cell viability range must be obtained. The extrusion parameters that this study focused on consisted of (a) the concentration of genipin in the collagen, (b) the temperature of the printhead, and (c) the time delay between the preparation of the collagen mixture and the printing process. Characterization began with measuring extrusion pressure for each genipin concentration (13.26, 4, and 0 mM) and temperature (25, 31 and 37˚C), which covered the genipin concentrations tested with cells previously. This established pressure reference points for 3D printing at different genipin concentrations. Next, the gelation level of the extruded collagen in each group was observed visually and categorized into 3 levels, under-gelation, proper-gelation, and over-gelation. After that, the genipin concentration and temperature which resulted in the most suitable gelation level was selected. Finally, the collagen was extruded with the selected parameters at different delay periods from genipin addition. For each time delay, the extrusion pressure was recorded to aid selection of the most suitable delay period.

A broad range of genipin concentrations (0.5, 0.25, and 0.125 mM) were selected for an experiment to assess the effect of genipin concentration on the extrusion pressure. For each group, 500 mm3 of collagen mixture was extruded from the syringe at a rate of 10 mm3/s at 3 different printhead temperatures, namely 25, 31, and 37 oC. During the extrusion, the pressures were logged continuously at a frequency of 100 times per sec. The recorded pressures were then time-averaged to represent each collagen group and corresponding temperature. Note that to ensure that the averaged pressures were obtained at a steady state of extrusion, only the data in the middle 40 s of the extrusion process were used. According to Fedorovich et al. (2011)26, printing pressures of up to 200 kPa do not significantly compromise the viability of multipotent stromal cells. Their study employed a similar needle size (0.21 mm vs. our 0.18 mm) but used a more viscous material than collagen, thus generating higher shear rates at the same pressure. Therefore, only set-ups where pressure did not exceed 200 kPa in this experiment were chosen for the next analysis.

The concentrations of genipin that were acceptable for cell viability were 0.125, 0.25, and 0.5 mM. Collagen was mixed with genipin at one of the usable concentrations and then loaded into the syringe. The mixtures were then extruded at different temperatures and flow rates in the same manner as in the previous experiment. A camera was used to capture the flow behavior of the collagen for analysis. The footage was then used to visually evaluate the process and chose the mixing ratio and printhead temperature that led to the best gelation state.

The mixtures with the most suitable combination of genipin concentration and temperature were then extruded from the printhead at the same volume and flow rate as in the previous experiment. The extrusion was performed at 3 different delay times (30, 60, and 90 min) from the moment that genipin was added to the collagen. During each extrusion, the printhead recorded the pressure data over time, as in the previous experiment. Next, the mean and standard deviation of the pressure were calculated for each extrusion delay time.

Having selected the most appropriate genipin concentration, printing temperature, and time delay period, the preferred volumetric flow rate needs to be chosen. To characterize this parameter, the collagen was printed into a scaffold shape, as would be used for the 3D cell culture, at different volumetric flow rates of 0.5, 2.5, and 4.5 mm3/s. The scaffold design consisted of two layers, each made up of six parallel lines. The two layers were orientated at 90o to each other so that the lines formed a grid-like structure. Each line was designed to be 1 mm in diameter and 20 mm in length, such that the overall surface area of the scaffold was 20 mm x 20 mm. Each layer was printed at 0.5 mm layer height. The scaffolds were printed directly onto a Corning® 60 mm diameter x 15 mm deep tissue culture-treated nonpyrogenic polystyrene petri dish. The dish was placed on the heated print bed which was set to 50˚C for at least 5 min to stabilize the temperature of the dish before every print. The scaffolds at each flow rate were printed at three different delay times, namely 30, 60, and 90 min from the moment that genipin was added. The pictures of the scaffolds were taken from above. Based on the number of successfully printed pores in the scaffold together with the visual analysis, the optimal print speed and the delay time before printing were chosen for the experiment with the cells.

Validation of optimal genipin concentration and 3D printing protocol for equine endometrial tissue constructs.

In this experiment, the optimal genipin concentration (0.5 mM), and optimal 3D printing protocol were tested with equine endometrial cells in triplicate.

A 0.5 mM genipin solution was gently mixed with collagen using a magnetic stirrer, and air bubbles were removed by centrifugation at 1000 rpm for 1 min. eECs and eMSCs were prepared at a 1:1 ratio to a total cell density of 300,000 cells/mL added to the genipin-crosslinked collagen matrix. The cell-laden collagen solution was loaded into a 10 cc sterile syringe (Nipro, Thailand) and incubated at 37 °C with 5% CO₂ for 60 min to allow crosslinking. A 3D bioprinter was used for printing, and aprotocol for reconstructing equine endometrial tissue was established based on printability, including the selected printing velocity.

The 3D-ET constructs were cultured in EC medium for 7 days with medium changes every 2–3 days. Morphological changes were monitored daily using a phase-contrast microscope (CK X41 Olympus, Japan).

Cell viability was assessed on days 1, 3, and 7 using 2 µM Calcein-AM (203700-1MG, Merck, Germany) and 2 µM Ethidium homodimer (46043-1MG-F, Sigma). The 3D-ET constructs were incubated with the stains in the dark for 15 min and then evaluated under a fluorescence microscope; live cells were indicated by green Calcein-AM labeling, and dead cells by red Ethidium homodimer labeling.

In this experiment, a structured system was designed to develop each cell-type layer of the endometrial 3D model separately, based on the observation that eMSCs thrived under 3D bioprinting conditions, whereas eECs showed better survival in a conventional 3D-ET system.

Passage 5 eMSCs were digested with 0.25% trypsin-EDTA, centrifuged, and stored at 4 °C. Collagen BM was neutralized and mixed with 0.5 mM genipin. The eMSCs were suspended at a density of 300,000 cells/mL in this crosslinked collagen matrix. The mixture was gently stirred, loaded into a sterile syringe, and incubated at 37 °C with 5% CO₂ for 60 min to allow crosslinking. The 3D bioprinter was used to fabricate endometrial tissue under optimized conditions, followed by a 60 min incubation for polymerization before adding MSC medium. The eMSCs printing layer was tested in five replicates.

After 3 days of eMSC layer development, an eEC layer was prepared by mixing with collagen BM (without genipin) using the conventional 3D-ET method. The MSC medium was removed, and the eMSC layer was washed with 1x PBS before the eEC layer was added on top of the eMSC layer, followed by polymerization at 37 °C with 5% CO₂ for 60 min. EC medium was then added, with medium changes every 2–3 days. The eECs layer was printed in five replicates following the eMSC layer.

Morphological changes in eMSCs and eECs were observed over a 10 day period using a phase-contrast microscope (CK X41 Olympus, Japan). For structural characterization, cryostat sections were used. Briefly, collagen scaffolds were collected on days 9–10, frozen at -20 °C for at least 24 h and embedded in Optimal Cutting Temperature (OCT) compound (Sakura Finetek, Torrance, CA) for rapid cryofixation. The frozen blocks were sectioned at -20 °C using a bench microtome (RM 2035, Leica Biosystems Nussloch GmbH, Nussloch, Germany), producing 30 μm sections that were mounted on coated slides and then stored at -20 °C. eECs were characterized using Pan-cytokeratin (ab234297, Abcam), and eMSCs with Vimentin (V5255, Sigma Aldrich: Table 1), following a modification of the protocol described by Santiviparat et al. (2024).

On day 10 of the 3D-ET printing system culture (7 days after eECs were added), the systems were divided into two groups, a control group and a group treated with 250 nM oxytocin for 24 h, as described by Penrod et al. (2009)27. After treatment, the culture medium was removed, and the scaffold were immediately collected and stored at -80 °C for further analysis.

Mucin characterization was performed on cryostat sections of 3D-ET printed tissue to evaluate mucin production. Histological visualization of sulfated and carboxylated acid mucopolysaccharides, as well as sulfated and carboxylated mucins, was achieved using the Alcian Blue (pH 2.5) stain kit (Cat. No. H-3501, Vector Laboratories, Inc., United States). Staining was conducted under acidic conditions following the manufacturer’s protocol. Nuclei appeared red, while acidic sulfated mucosubstances, hyaluronic acid, and sialomucins were identified by blue staining.

Cryostat sections of the 3D-ETprinted tissue were used to characterize OXTR expression. Sections were washed twice for 3 min with 0.5% Tween 20 in 1x PBS, then blocked with a solution containing 0.2% bovine serum albumin (BSA), 0.1 M glycine, 0.05% Tween 20, and 0.1% Triton X. The primary OXTR antibody (bs-1314R, Bios: Table 1) and ERα (Clone 33, invitrogen: Table 1) was diluted 1:50 in a solution of 1x PBS, 10 mM glycine, 0.05% Tween 20, 0.1% Triton X, and 0.1% H2O2, and added to the sections seperately for overnight incubation at 4 °C. The secondary antibody was diluted in 0.1% Tween 20 in PBS and used to incubate the sections for 30 min at room temperature. Nuclei were counter-stained with DAPI, and sections were mounted with anti-fade reagent under a cover slip. Immunofluorescence microscopy was used for 3D structural assessment.

These steps were adapted from the protocol outlined by Santiviparat et al. (2024). Briefly, primer sequences were generated using Primer3 based on high-quality nucleotide sequences from the National Center for Biotechnology Information (NCBI; http://www.ncbi.nlm.nih.gov/) and previous studies28,29 as detailed in Table 2. Primer specificity and compatibility were verified using NCBI Primer-BLAST and OligoAnalyzer software (http://eu.idtdna.com/analyzer/Applications/OligoAnalyzer/) before synthesis by Eurofins MWG Operon (Eurofins MWG Operon). Glyceraldehyde 3-phosphate dehydrogenase (GAPDH) was used as a reference gene, and OXTR, PTGS2, and prostaglandin F synthase (PGFS) were the target genes. Total RNA was extracted using the RNeasy Minikit (Qiagen, Hilden, Germany) and assessed for quality with a NanoDrop™ 2000 (Wilmington, DE, USA), ensuring an A260/A280 ratio of 1.9–2.1. Residual DNA was removed using DNaseI (RQI Rnase-Free Dnase, Promega WI, USA), and cDNA synthesis was carried out using the ImprompT II Reverse Transcription System (Promega, Madison, WI, USA) using 1000 ng of RNA in accordance with the manufacturer’s instructions. PCR was performed with goTaq® Green Master Mix (Promega, Madison, WI, USA), including initial denaturation, 30–40 amplification cycles, and a final extension. PCR products were separated on 1% agarose gels with 1x Tris/Borate/EDTA buffer (TBE buffer, pH 8.0, 2 mM EDTA, 90 mM Tris, 90 mM boric acid) and visualized with Redsafe (iNtRON Biotechnology, Gyeonggi-do, Korea) and UV light.

For real-time PCR, gene expression was assessed using the KAPA SYBR® FAST qPCR kit, with primers optimized by conventional PCR. Standards for real-time PCR were prepared from purified PCR products and quantified, ensuring a standard curve with R2 > 0.995 and near 100% efficiency. Relative gene expression was calculated using the Livak and Schmittgen method30, normalized to GAPDH, and analyzed by one-way ANOVA.

Statistical analysis was analyzed by SPSS version 29.0.0.0, while all graphs in this study were prepared using GraphPad Prism 9.5.1. Continuous data were tested for normality and equivalence of variance, and reported as mean ± standard error of the mean (SEM). ANOVA was used to analyze treatment differences in all experiments, except for Experiment 4, where the quantitative data derived from real-time PCR was analyzed using an independent t-test. While viability of the endometrial cells within experiment 3.2 was a descriptive analysis.

Varying the genipin concentration significantly affected cell viability. For eMSCs, lower concentrations of genipin (0.25-0 mM) resulted in a higher cell count than higher concentrations (4, 2, 1 mM) (p < 0.05). There was no significant difference in cell number between the 0.5 mM genipin group and the groups with either higher or lower concentrations (Fig. 2A). High concentrations of genipin (4, 2, 1 mM) were also toxic to eECs as indicated by significantly lower cell counts compared to lower concentrations (0.5, 0.25, 0.125, and 0 mM) (p < 0.05) (Fig. 2B).

Genipin concentration significantly influenced the morphology of endometrial cells. At higher concentrations (4, 2, 1 mM), both eECs and eMSCs predominantly exhibited a rounded morphology with minimal spreading during the 7 day culture. In the 1 mM genipin-collagen scaffold, a few cells assumed a spindle-like shape (Fig. 2C). At lower genipin concentrations (0.5–0 mM), however, eMSCs formed spindle-like structures by day 1, and eECs developed into gland-like structures by day 3, which further expanded into larger, branched formations by day 7 (Fig. 2D). The eMSCs yielded a significantly higher percentage of spreading cells after 3 days of culture at lower (0.25–0 mM) compared to higher genipin concentrations (4, 2, 1 mM) (p < 0.05), with no significant difference to the 0.5 mM group. By contrast, eECs yielded significantly fewer cells with spreading morphology at higher (4, 2, 1 mM) compared to lower genipin concentrations (0.5, 0.25, 0.125, and 0 mM) (p < 0.05) (Fig. 2E).

The frequency sweep test for samples with different levels of genipin was performed at 1% shear strain within the linear viscoelastic region. For all samples, higher values of elastic modulus (G’) compared to viscous modulus (G”) were observed from the angular frequency of 0.1 to 100 rad/s. This suggests that all tested collagen genipin mixtures displayed predominantly elastic behavior. The logarithmic average of phase angle (tan-1(G”/G’)) was then calculated for each sample. The values were comparable for different concentrations of genipin (5.05° at 0.25 mM, 5.16° at 0.5 mM, 4.93° at 1 mM, 4.88° at 2 mM). However, the phase angle of the collagen without genipin was significantly higher at 10.43°. The lower phase angles demonstrated that the addition of genipin at different concentrations promoted similar solid-like elastic behavior of the bioink compared to pure collagen. Lastly, the logarithmically averaged elastic modulus was compared among samples as shown in Fig. 3A. The value exhibited a positive correlation with genipin concentration, though the rate of increase slowed at higher concentrations. This concluded that the higher concentration of genipin escalated the rigidity of the viscoelastic bioink.

Results of a 7 days study examining the toxicity of various genipin concentrations (ranging from 0 to 4 mM) on eMSCs and eECs, using the MTT assay. (A) Lower genipin concentrations (0.25-0 mM) resulted in significantly higher cell counts compared to higher concentrations (4, 2, 1 mM). There was no significant difference in cell counts between the 0.5 mM group and either the higher or lower concentration groups (p < 0.05). (B) Cell counts were significantly higher in the lower (0.5-0 mM) than the higher genipin concentrations (4, 2, 1 mM) (p < 0.05). (C & D) Morphological assessment of eECs and eMSCs in conventional 3D-ET using a collagen-based scaffold with varying genipin concentrations over 7 days. (C) At higher genipin concentrations (4, 2, 1 mM), the cells primarily exhibited rounded morphology with limited spreading during the 7 days culture period. In the 1 mM genipin-collagen scaffold, occasional cells displayed a spindle-like morphology (indicated by a yellow arrow; high magnification view in the lower right corner of the image). (D) In scaffolds with lower genipin concentrations (0.5-0 mM), eMSCs formed spindle-like structures on day 1. By day 3, the eECs had developed into gland-like structures, which by day 7 had evolved into more extensive branching formations. (E) A bar graph showing the percentage of spreading cells (eECs and eMSCs) after 3 days of culture in 10 randomly selected areas (10x magnification). The 0.5-0 mM genipin-collagen scaffolds exhibited significantly enhanced cell spreading compared to higher concentrations (p < 0.05). Scale bar 50 μm.

The time-averaged pressure measured during the extrusion of the collagen mixture with varying genipin concentrations and temperatures is shown in Fig. 3A. Across all genipin groups, pressure exhibited a positive correlation with temperature. Similarly, at any given temperature, higher genipin concentrations resulted in increased pressure. The highest recorded time-averaged pressure was 180.17 kPa for the 13.26 mM solution at 37 °C, which was still below the 200 kPa safety threshold for eMSCs26. Therefore, the 0.125, 0.25, and 0.5 mM genipin concentrations used for viability-testing were deemed physically safe for eMSCs and used in subsequent experiments. The extrusion of collagen with these three genipin concentrations was subsequently performed at room temperature (25 °C); frames captured by the camera are displayed in Fig. 3B. The footage indicated that all tested concentrations and temperatures remained in the under-gelation phase, where the collagen exited the nozzle as droplets. Consequently, the highest genipin concentration of 0.5 mM at 37 °C was selected for further investigation of time delay effects.

Next, collagen with 0.5 mM genipin was extruded at 37 °C at 3 different waiting times: 30, 60, and 90 min after the addition of genipin. The extrusion pressures were measured and are presented in Fig. 3C. There was a positive correlation between extrusion pressure and waiting time, suggesting progressive gelation of the solution over time. None of the extrusion pressures exceeded 125 kPa, indicating that all waiting times were within safe limits and suitable for the printability experiment.

Various delay times and volumetric flow rates were tested using a 0.5 mM genipin-collagen solution at a printhead temperature of 37 °C. The 2-layer printed scaffold is displayed in Fig. 3D, and the number of successfully printed pores is detailed in Table 3. Results indicated that the scaffold printed after 90 min of waiting time exhibited over-gelation, characterized by a visible separation between the gel and liquid phases, leading to an incomplete scaffold structure. Additionally, the number of successfully printed pores was lowest at 90 min across all flow rates. The largest number of successfully printed pores was observed at the highest flow rate of 4.5 mm³/s for all waiting times. Only minor differences in pore number were noted between the 30 min and 60 min groups. Consequently, the optimal printing conditions for eMSCs applications were determined to be a 0.5 mM genipin-collagen solution at 37 °C, with a flow rate of 4.5 mm³/s and a waiting time of 30–60 min after the addition of genipin. The summary of the 3D printability assessment in this study was presented in Table 4.

(A) Elastic Modulus (G’) of collagen-genipin mixtures at varying genipin concentrations (0 mM, 0.25 mM, 0.5 mM, 1 mM, 2 mM). For each concentration, rheometer measurements were logarithmically averaged over an angular frequency range of 0.1 to 100 rad/s. A higher elastic modulus indicates increased rigidity with higher genipin concentration. Determining the 3D printing conditions for an equine endometrial tissue printing model. (B) Time-averaged extrusion pressure of collagen mixtures with varying genipin concentrations and temperatures (25 °C, 30 °C, 37 °C). The data illustrate a positive relationship between temperature and pressure for all genipin concentrations, with higher genipin concentrations resulting in increased pressure at each temperature. (C) Frames captured during the extrusion of collagen with 0.125, 0.25, and 0.5 mM genipin at room temperature (25 °C). The images indicate that all tested conditions remained in the under-gelation region, where collagen flowed out of the nozzle in droplet form. (D) Extrusion pressure profiles of 0.5 mM genipin-collagen solutions at 37 °C after waiting times of 30, 60, and 90 min. The increasing pressures indicated ongoing gelation, with all pressures remaining below the 125 kPa safety limit for eMSCs. The dotted line shows a second-degree polynomial line of best fit. (E) The 2-layer scaffold printed with a 0.5 mM genipin-collagen solution at a printhead temperature of 37 °C, demonstrating variations in structure at different waiting times and flow rates.

Following the 3D printing of a collagen matrix containing eECs and eMSCs with 0.5 mM genipin, all cells initially displayed a rounded morphology. Calcein-Ethidium staining demonstrated a predominance of viable cells, indicated by green fluorescence. By day 3, fibroblast-like structures began to form among the rounded cells. By day 7, eMSCs had differentiated into more extensive fibroblast-like structures. However, eECs did not develop gland-like formations, which was unfavorable, and different to conventional 3D-ET systems. Viability staining revealed green fluorescence in the fibroblast-like cells on days 3 and 7, whereas the majority of rounded cells exhibited red fluorescence, identifying them as non-viable cells (Fig. 4).

Based on previous findings, this study demonstrated that only eMSCs could grow reliably under 3D printing conditions, whereas eECs showed better survival in a conventional 3D-ET system. Therefore, we designed an endometrial 3D model in which each cell layer was developed separately: the eMSC layer served as the base and was produced through 3D printing, with the eECs subsequently added on top using a conventional method (Experiment 4). This model aimed to replicate endometrial tissue structure, with the hope of recapitulating in vivo conditions, and was evaluated for structural and functional mimicry.

(A) Displays the scaffold within a 60 mm culture plate after 7 days of incubation. (B) Left Panel: Cell morphology from day 1 to day 7. Initially, cells showed a round morphology (day1) with no spreading. By day 3, fibroblast-like structures appeared, surrounded by round cells. By day 7, these structures had expanded significantly, but without any gland-like formations among the eECs (Scale bar 100 μm) Right Panel: Demonstrates cell viability over the same period. Initially, most cells exhibited green fluorescence indicating high viability (day1). By day 3 and day 7, green fluorescence was mainly seen in fibroblast-like and some round cells, denoting ongoing viability. Conversely, most round cells showed red fluorescence (ethidium bromide staining), identifying them as dead. Both eECs and eMSCs were viable in the 0.5 mM genipin-collagen scaffold initially, but only eMSCs consistently tolerated and survived the conditions of 3D printing (Scale bar 100 μm).

Within 24 h of printing, eMSCs exhibited a fibroblast-like morphology, and continued to proliferate until day 3. Following the addition of the eEC layer, eECs adopted a polygonal shape and formed spheroid-like structures by day 2 of their culture (day 5 of eMSC culture). By day 7, the eMSC layer had fully developed fibroblast-like structures, while the eEC layer formed gland-like structures. On day 9 (day 6 of eEC culture), eECs had progressed to yield horizontally orientated tubular structures (Fig. 5A). Immunostaining demonstrated Pan-cytokeratin expression in the eECs, confirming epithelial characteristics (Fig. 5B), and Vimentin expression in eMSCs, confirming their mesenchymal identity within the collagen-genipin scaffold(Fig. 5C).

Mucin production was identified by alcian blue staining surrounding the gland-like structures of eECs. The gland-like structures formed by eECs after 3D printing exhibited ERα localization in the cytoplasm, with translocation to the nucleus (Fig. 6B, Left panel). OXTR protein was detected in the cytoplasm of cells within the gland-like structures of the eEC layer (Fig. 6B, Right panel). Oxytocin treatment did not alter OXTR or PGFS gene expression compared to controls but led to the expected upregulation of PTGS2 expression (P < 0.05).

Morphological characterization of endometrial cells in the 3D-ET printed model. (A) This panel illustrates the morphological evolution of endometrial cells in a 0.5 mM genipin-collagen scaffold. Initially, printed eMSCs exhibited a fibroblast-like shape and underwent rapid growth. After over-layering the eECs in a collagen matrix, the eECs began to divide and form spheroid-like structures. By day 7, the lower layer showed eMSC progression to fully developed fibroblast-like structures, whereas the eECs in the upper layer formed gland-like structures. By day 9 of co-culture, eECs had further differentiated into more complex tubular structures orientated horizontally. The orange circle outlined the boundary of the gland-like structure. All images included a scale bar of 100 μm. Immunostaining revealed (B) Pan-cytokeratin in the eECs glandular structures, indicative of epithelial differentiation (Scale bar 50 μm), while (C) Vimentin staining confirmed the mesenchymal origin and characteristics of the eMSCs in the scaffold (Scale bar 50 μm).

(A) Blue staining (orange arrow) indicated the presence of mucin produced by glandular-like structures of eECs in the 3D-ET printed tissue, while red staining marked cell nuclei (Left panel: scale bar 50 μm; Right panel: scale bar 20 μm). (B) Left panel: Green fluorescence showed ERα localized in the cytoplasm and translocated into the nucleus of eECs after 3D printing. Right panel: Green fluorescence indicated OXTR protein expression within the cytoplasm of eEC-derived cells in the 3D-ET printing system (Scale bar: 20 μm). (C) Oxytocin stimulation did not alter OXTR or PGFS gene expression but significantly upregulated PTGS2, a downstream target of the oxytocin-PGFS pathway (p < 0.05), indicating that the 3D-ET printed model retained functional characteristics of eECs.

This study developed a 3D equine endometrium model based on a 3D printing method using eMSCs in matrix made up of collagen, with genipin as a stabilizer. Genipin concentrations of 2–4 mM were found to be cytotoxic to eMSCs and eECs, with moderate toxicity also observed at 1 mM; it was concluded that genipin concentrations in the matrix should not exceed 1 mM to ensure cell viability. This finding aligned with studies on human MSCs, where concentrations above 1 mM were cytotoxic, and 0.5-1 mM were optimal for maintaining viability 31,32. In addition, lower concentrations of genipin (≤ 0.5 mM) better preserved cell spreading and function, possibly as a result of reduced accumulation of reactive oxygen species (ROS) associated with increased ECM stiffness11. The rheology test indicated that collagen gels crosslinked with 0.5 mM genipin exhibited a viscoelastic stiffness of approximately 1.2 kPa. These results align with previous studies showing that MSCs can tolerate stiffness up to 10 kPa. Our results fall within the reported modulus range of 0.0292–3.36 kPa, suggesting that MSCs can survive within this environment33.Furthermore, the observed stiffness closely resembles that of native human endometrial tissue, which ranges from 0.25 kPa in the non-pregnant state to 1.25 kPa during pregnancy34,35. Therefore, a genipin concentration of 0–0.5 mM was selected to preserve endometrial structure and function while mimicking in vivo tissue stiffness. Future studies should assess the mechanical properties of native equine endometrial tissue to refine in vitro models by comparing them with the Young’s modulus of collagen-genipin scaffolds, providing critical data for model development under physiological conditions.The study also established a 3D printing process for creating collagen scaffolds as part of an in vitro equine endometrial model, offering a baseline method for reconstructing complex tissue layers. However, all genipin concentrations recommended for enhancing collagen stability were below the gelation threshold when tested using pressure extrusion before 3D printing. Higher genipin concentrations and temperatures have been shown to accelerate gelation and reduce the pre-printing time (Fig. 3)36,37; in the current study, a genipin concentration of 0.5 mM was identified as optimal for the scaffold. A printing protocol that incorporated a temperature of 37 °C and a 60 min pre-printing delay improved collagen printability without significantly affecting cell viability, similar to previous studies38,39 study. To further enhance scaffold properties, modifications such as photochemical crosslinking or combining collagen with other stabilizing bioinks like alginate or gelatin could be used to improve printability31,40,41,42.

Both conventional and 3D bioprinting techniques were tested for reconstructing equine endometrial tissue in vitro using modified collagen-genipin scaffolds. eMSCs successfully maintained their morphology and viability with both methods, whereas eECs exhibited proper development only in the conventional seeding approach. The viability and functionality of eECs appeared to be influenced by three primary factors: genipin toxicity, material properties, and the bioprinting process itself. In this study, the selected genipin concentration (0.5 mM) was found to be optimal for eECs, as it supported their viability and functionality in both 2D and conventional 3D cultures. However, the sensitivity of eECs became apparent during 3D bioprinting, particularly with extrusion-based bioprinting. This mechanical-driven dispensing system generates high pressure when printing high-viscosity bioinks, which can lead to increased cell death43 In addition, the relationship between shear stress, nozzle size, and bioink viscosity, might cause cell membrane rupture leading to impaired cell viability and function, and improper eEC development44,45. In contrast, eMSCs remained viable, retained their fibroblast-like morphology, and exhibited exponential growth post-printing, suggesting cell-specific tolerance to shear stress, although this study did not quantify shear stress during the 3D printing process. Future research should prioritize reducing shear stress to improve eEC viability and functionality following 3D printing. Shifting from nozzle-based to nozzle-free or hybrid bioprinting methods may minimize cell damage, enhance viability and function and possibly also aid in understanding bioprinting mechanisms to maintain cellular activity and optimize overall printing performance46.

In a functional assessment of 3D-printed endometrial tissue, ERα expression and mucin secretion served as key markers, confirming the functional integrity of the 3D-ET printed model, as they are specific to in vivo endometrial tissue 47,48. Additionally, the functional test included oxytocin treatment to evaluate the tissue’s physiological responsiveness. OXTR gene expression levels were not significantly different between treatment and control. However, immunolabeling demonstrated OXTR protein expression in gland-like structures of the eECs. This finding likely reflected the in vivo scenario where OXTR gene expression is similar between mid and late diestrus and early pregnancy, with changes in oxytocin responsiveness regulated at the protein rather than at the mRNA level49. The significant increase in PTGS2 gene expression in the oxytocin-treated group compared to controls indicates that the 3D-printed eECs retained functional responsiveness to oxytocin stimulation. This supports the hypothesis that eECs maintain their structural and functional integrity within the downstream cascade of the oxytocin-induced PG synthesis pathway. However, no significant difference in PGFS gene expression was observed between the oxytocin-treated and control groups. This suggests that while the 3D-printed structure mimics the architecture and some functions of in vivo endometrial tissue, it may have limitations in fully replicating endometrial function.The lack of PGFS gene activation may be due to differences in dose and exposure time requirements across tissues and species. For example, Penrod et al. (2013)50 reported that a similar oxytocin dose (250 nM) increased PGFS expression after 6 h in equine endometrial explants. While Xiao et al. (1999)51 found that oxytocin-induced PGFS activation in bovine endometrial epithelial cells peaked at 3 h before declining. These findings suggest that the oxytocin exposure duration in this study may not have been optimal for PGFS gene activation. Future studies should determine the optimal oxytocin dose and incubation time for inducing PGFS expression in in vitro models. Additionally, PGF2α secretion in the culture medium could be evaluated as a marker of oxytocin-stimulated pathway activation, providing a more dynamic assessment of oxytocin responsiveness.5051. In conclusion, this study demonstrated that genipin helps stabilize collagen scaffolds, and that its toxicity for endometrial cells was both dose- and cell-type-dependent. A concentration of 0.5 mM was found to be optimal for scaffold stability while maintaining acceptable toxicity levels for eMSCs. Genipin-stabilized collagen enhanced production of the bottom layer of a 3D-ET system when integrated with 3D printing, and analysis of 3D printing characteristics emphasized the importance of temperature and time delay after genipin addition for printability. The resulting 3D-ET model successfully mimicked in vivo endometrial tissue, exhibiting gland-like structures, normal stromal cell appearance, and showed eEC responsiveness to oxytocin. This model holds promise as a more complex in vitro system for endometrial research. Genipin further improved scaffold longevity for long-term studies. The 3D-printing approach allowed precise customization of tissue layers, with potential future advances including incorporation of additional cell types, and integration of tubular or vascular structures to better replicate the complexity of in vivo endometrial tissue for functional and pathological studies.

All data generated or analyzed during this study are included in the published article and the datasets used and/or analysed are available from the corresponding author on reasonable request.

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This research project is supported by the Second Century Fund (C2F), Chulalongkorn University (PN) of 2019 (Batch1/2019), 90th Anniversary Ratchadaphiseksomphot Endowment fund (GCUGR1125662081D, No.1-81) and the National Research Council of Thailand (N41A660173), in addition with, The European Union’s Horizon 2020 Research and Innovation Program under the Marie Skłodowska-Curie action, project “WhyNotDry” (GA-101131087).

Center of Excellence for Veterinary Clinical Stem Cells and Bioengineering, Chulalongkorn University, Bangkok, Thailand

Sawita Santiviparat, Sudchaya Bhanpattanakul & Theerawat Tharasanit

Center of Excellence in Animal Fertility Chulalongkorn University (CU-AF), Chulalongkorn University, Bangkok, Thailand

Sawita Santiviparat & Theerawat Tharasanit

Maxwell H. Gluck Equine Research Center, University of Kentucky, Lexington, KY, USA

Tom A. E. Stout

Department of Anatomy, Faculty of Veterinary, Science Chulalongkorn University, Bangkok, Thailand

Sayamon Srisuwattanasagul

Department of Obstetrics, Gynecology and Reproduction, Faculty of Veterinary Science, Chulalongkorn University, Bangkok, Thailand

Sawita Santiviparat & Theerawat Tharasanit

Department of Pathology, Faculty of Veterinary Science, Chulalongkorn University, Bangkok, Thailand

Sudchaya Bhanpattanakul

Institute for Molecular Systems Engineering and Advanced Materials, Heidelberg University, Im Neuenheimer Feld 225, 69120, Heidelberg, Germany

Setthibhak Suthithanakom & Kai Melde

Micro/Nano Electromechanical Integrated Device Research Unit, Faculty of Engineering, Chulalongkorn University, Bangkok, Thailand

Setthibhak Suthithanakom

Max Plank Institute for Medical Research, Jahnstr.29, 69120, Heidelberg, Germany

Setthibhak Suthithanakom & Kai Melde

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SS collected the samples, conducted the experiments, performed cell culture, carried out and analyzed RT-PCR, real-time PCR, and fluorescent staining, drafted the main manuscript, and prepared Figs. 1, 2 and 4, and 6. SSu managed the 3D printing process, contributed to manuscript drafting and prepared figures 3. SB set up the experiments and performed cell culture. SSr provided the antibody and cryostat and established the protocol for fluorescent staining. TT, TS, KM and SSr conceptualized and supervised the study, reviewed the manuscript, and approved the final version.

Correspondence to Theerawat Tharasanit.

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Santiviparat, S., Suthithanakom, S., Bhanpattanakul, S. et al. Development of a two-layer 3D equine endometrial tissue model using genipin-crosslinked collagen scaffolds and 3D printing. Sci Rep 15, 19759 (2025). https://doi.org/10.1038/s41598-025-04013-4

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