The adhesion capability between the hydrogel and the bone tissue was assessed through qualitative tests. The hydrogel was injected and cured onto prepared bone samples, which were then subjected to continuous water flushing to observe whether the hydrogel detached. For further quantitative testing, two pig skin specimens were fixed on the test platform. The hydrogel was evenly applied to the surfaces of the two pig skin specimens and then pressed for 3 min. After that, the plates were separated at a constant speed of 13.4 mm/min. The adhesion of the hydrogel was defined as the maximum load measured during the separation process.

The contact angle was determined using a contact angle goniometer manufactured in Shanghai, China. Deionized water was dropped at three distinct locations on the surface of the sample at room temperature, and the contact angle was photographed and recorded immediately after the liquid was dropped.

Cylindrical specimens with a diameter of 1 cm and a height of 5 mm were made. To start with, the samples were weighed ( W 0 ). After that, the samples were put into deionized water. At predetermined time intervals (2 h,4 h, 6 h), the samples were removed, and the surfaces were gently dried. The weight of the samples was measured again ( W t ), and the swelling rate (%) was evaluated according to Equation 1: Swelling rate  % = W t ‐ W 0 W 0 × 100 % (1)

Acrylic glass tubes with a length of 2.00 m and an inner diameter of 3 mm were filled with water to a level of 1.91 m. This water-filling height was equivalent to a systolic blood pressure of 18.68 kPa (140 mmHg) (Brückner et al., 2016). The tubes were sealed with Gel samples measuring 3 cm (n = 3). Subsequently, these fabricated structures were continuously observed until they experienced failure.

The vial tilt method was employed to measure the gelation time at 37 °C. Identical volumes of DF-PEG and CMCS solutions at diverse concentrations were combined in a vial. The time when no liquid flow was observed in the tilted vial was recorded as the gelation time (n = 3).

Scanning electron microscopy (SEM, SU8600, Japan) was used to analyze the morphologies of the Gel samples (voltage: 5 kv, amplification factors: 100× and 200×) and the Gel-Th samples (voltage: 5 kv, amplification factors: 100×). Using Nano Measure and Origin software, the pore size of the Gel was quantified based on the SEM images.

For TGA-DTG, the temperature was set to range from room temperature to 700 °C with a heating rate of 10 °C/min. For DSC, the temperature program was designed as follows: first cooling from room temperature to −20 °C, then heating from −20 °C to 200 °C, with a heating rate of 10 °C/min.

The samples were adjusted to suitable dimensions and positioned at the bottom of the orifice plate. A volume of 20 µL of 0.1 M CaCl2 and 200 µL of anticoagulated whole blood were introduced into a centrifuge tube and then vigorously mixed. Subsequently, the samples were subjected to incubation at 37 °C for 3, 10, 20, 30, and 40 min respectively. Upon completion of the incubation period, we added 1 mL of deionized water. The hemoglobin content was quantified using a UV- visible spectrophotometer (Varioskan Flash, Thermo Scientific, United States) at a wavelength of 545 nm. For the blank control, 10 µL of calcium-fortified anticoagulated whole blood was mixed with 1 mL of deionized water. Every sample underwent an evaluation in three independent replicates. The blood clotting index (%), as defined in the context of this study, was computed in accordance with Equation 2: Blood clotting index  % = I a I w × 100 % (2) in this context, I a represents the absorbance of the samples, while I w denotes the absorbance of the blank.

The whole blood was centrifuged at 1,000 rpm for 10 min to obtain erythrocytes. The samples were then immersed in 1 mL of phosphate-buffered saline (PBS). 20 μL of erythrocytes were introduced into the PBS-soaked samples. The resultant mixture was then incubated at 37 °C for 1 h. Following the incubation, the mixture was spun down at 2,000 rpm for 10 min to acquire the supernatant. The absorbance of this supernatant was determined at 545 nm using a UV-visible spectrophotometer (Varioskan Flash, Thermo Scientific, United States). As per the method described in reference (Wang Y. et al., 2020), deionized water and PBS were combined with untreated erythrocytes, acting as the positive and negative controls respectively. The hemolysis ratio (%), as defined by the experimental protocol, was calculated according to Equation 3: Hemolysis ratio  % = D t ‐ D nc D pc ‐ D nc × 100 % (3) where D t , D nc and D pc are the absorbance of the samples, PBS, and deionized water, respectively.

The extent of platelet adhesion was quantitatively assessed through lactate dehydrogenase (LDH) activity measurements following established protocols (Shen et al., 2020). Whole blood specimens were processed by centrifugation at 200 rpm for 20 min to isolate platelet-rich plasma (PRP), as described in previous methodology (Raza et al., 2023). Identically sized test samples were then submerged in the obtained PRP solution. After maintaining the specimens at 37 °C for 45 min, non-bonded platelets were removed through three sequential washes with PBS. To lyse the attached platelets, 0.15 mL of PBS solution containing 1% TritonX-100 was applied to the samples, maintaining 37 °C for 60 min. The resulting LDH enzymatic activity was measured through standardized procedures using a commercial detection kit (Biyuntian, China). Platelet adhesion quantification was performed by recording absorbance values at 490 nm wavelength using spectrophotometric analysis. Each experimental condition underwent triplicate tests with final results representing the arithmetic mean of three independent measurements.

Place the samples in a 24-well plate. Take 10 µL of the prepared citric acid-incubated whole blood and add it to the surface of the samples. Add 10 µL of whole blood to the surface of the well-plate as a positive control. Then, incubate the samples statically at 37 °C for 45 min. After the incubation was completed, wash the samples with PBS to assess the adsorption of red blood cells (RBCs). Lyse the RBCs adhered to the samples in deionized water, and measure the absorbance at 540 nm to determine the RBC adsorption. The formula used as Equation 4: RBC adsorption  % = OD sample OD blank × 100 % (4)

A volume of 20 µL of whole PRP and blood was separately transferred via pipetting onto the central area of the sample surface to analyze the adhesion of platelets and erythrocytes. Subsequently, the PRP and blood were incubated at 37 °C under static conditions for 1 h. Once the incubation process concluded, the surface of the sample was carefully rinsed with physiological saline to eliminate any non-adherent blood cells. Subsequently, a 4% paraformaldehyde solution (Biyuntian, China) was used to fully cover the surface of the samples. The samples were then subjected to an incubation process at room temperature for 10 min to fix the cells. The paraformaldehyde solution was then aspirated, and the samples were gently rinsed with deionized water before being air-dried at room temperature. The morphological features of the blood cells that adhered to the membrane surface were examined via a laser microscope (VK-150K, Keyens, Japan), capturing both laser images and two-dimensional images.

The whole blood was subjected to centrifugation at a rotational speed of 2,000 rpm for 10 min to obtain the platelet-poor plasma (PPP). Samples were cut into uniform sizes and placed in culture dishes. A total of 500 µL of the prepared PPP was combined with 50 µL of a 0.1 M calcium chloride solution for recalcification. Following this, 20 µL of the recalcified plasma was precisely dispensed onto the surface of the test specimens and then incubated at 37 °C for 15–20 min. Upon completion of the incubation, the specimens were irrigated with PBS. Subsequently, they were covered with a 4% paraformaldehyde solution for 10 min. The paraformaldehyde solution was aspirated. Subsequently, the test specimens were gently cleansed with deionized water and then allowed to air dry. Fibrin was stained using Masson’s trichrome stain for observation.

To evaluate the release profile of Th from Gel-Th, the hydrogels were individually submerged in PBS. At every predetermined time interval, the release medium was gathered and subjected to analysis with BCA kits, adhering to the procedure outlined in the instruction manual (Wang C. et al., 2020). The percentage of Th released (%) was computed via Equation 5: RP  % = W i W total × 100 % (5) in this equation, W i represented the quantity of Th present in the released medium, while W total denoted the overall amount of Th. The release experiment was replicated three times to ensure reliability and reproducibility.

To further confirm the mechanism of drug release, four kinetic model formulas were used for fitting, namely the zero-order kinetic model, Higuchi model, the first-order kinetic model and Ritger-Peppas model. The specific model formulas can be found in the Supplementary Material.

The Gel10 and Gel10-Th200 specimens were weighed ( W 0 ) and then fully immersed in PBS maintained at a temperature of 37 °C. At the predetermined time intervals, the samples were retrieved from the PBS, and the moisture on the surface of the hydrogels was carefully blotted dry. The sample was weighed (Wt), and the mass rate was calculated (%) as Equation 6: Mass rate  % = W t − W 0 W 0 × 100 % (6)

Tissue adhesion implies that the hemostatic hydrogel can effectively seal the site of bleeding and resist being washed away by blood (Le Thi et al., 2017; Shen et al., 2024). The hydrogel remained firmly attached to the bone samples even after continuous water flushing (Figure 2A), indicating good adhesion between the hydrogel and the bone tissue. In addition, the adhesion effect was quantitatively analyzed by tensile test, and the results showed that Gel10 was the optimal ratio, and its degree of cross-linking balanced mechanical strength and structural toughness. For Gel6 and Gel8, due to insufficient cross-linking, they were difficult to withstand external forces, had poor structural stability, and were prone to losing their original shape due to deformation. As for Gel12, the degree of cross-linking was too high, the toughness of the material decreases, and stress concentration was easy to occur when stressed, leading to rapid structural failure (Supplementary Figure S2).

Hydrophilicity is the core link for dialogue between biomaterials and the biological environment, and its role runs through the entire process from molecular recognition to tissue integration, and in many hydrogel studies (Zhu et al., 2024; Li et al., 2025), static contact angles are widely accepted as a more reliable and reproducible indicator for characterizing initial hydrophilicity. The Gel samples exhibited a contact angle of approximately 15° (Figure 2B), indicating that all Gel samples possessed favorable hydrophilicity.

Effective swelling capacity can facilitate the absorption of blood, which is favorable for the aggregation of platelets and thereby triggers blood coagulation (Lee et al., 2023; Chen et al., 2025). The swelling characteristics of the Gel specimens were quantitatively evaluated via the determination of the percentage augmentation in mass across distinct temporal intervals (specifically, 2 h, 4 h, and 6 h). The swelling rate for each group increased over time. Concurrently, it was observed that the degree of crosslinking exhibited an inverse relationship with the swelling rate. At 6h, the swelling rate of Gel6 reached 337%, while that of Gel12 was 267% (Figure 2C).

Medullary hemostasis primarily relies on the sealing capacity of the hydrogel following its curing process. Consequently, the assessment of liquid sealing is a crucial indicator of performance. The water-occluding efficacy of the Gel specimens was comprehensively appraised through the employment of an acrylic tube-based experimental model (Figure 2E) (Brückner et al., 2016). The relationship between the duration of liquid sealing and the degree of crosslinking of the hydrogel indicated that higher crosslinking resulted in extended sealing times. With the exception of the formulation with a 15:6 ratio, all other hydrogels demonstrated the ability to maintain sealing for more than 7 days (Figure 2D), thereby highlighting the effective sealing capability of the Gel samples.

An ideal injectable hydrogel should possess an optimal gelation time for clinical applications (Xue et al., 2022; Zhai et al., 2024). When the gelation duration is excessively brief, the hydrogel is likely to experience the curing process inside the syringe, a phenomenon that has the potential to severely hamper its injectability. Conversely, if the gelation time is too slow, timely hemostasis may not be achieved. Therefore, we investigated the gelation time of the hydrogel at varying concentrations of CMCS. The gelation time of the Gel samples exhibited a decreasing trend with an increase in the concentration of the raw materials. With 15% (w/v) DF-PEG and CMCS concentrations ranging from 6% to 12% (w/v), the gelation times varied from 35 s to 115 s. Specifically, when mixing 15% (w/v) DF-PEG with 12% (w/v) CMCS at a 1:1 ratio, gelation occurred within a range of 35–45 s (Figures 2F,G). Additionally, the introduction of thrombin subsequently reduced the gelation time to some extent. Such a fast-curing speed cannot fully meet the requirements of the hemostatic operation time, which typically demands approximately 60 s. It may lead to problems such as the inability to completely cover the wound and the clogging of the syringe. In contrast, the remaining three concentrations satisfied the necessary operational time. Notably, a CMCS concentration of 10% yielded the fastest hemostatic performance compared to the other two concentrations.

Considering the preceding property evaluations, the concentrations of 15% (w/v) DF-PEG and 10% (w/v) CMCS were chosen as the most suitable ratio for the ensuing experiments.

The mutually interconnected and intricately interpenetrating pores inherent in the Gel specimens endow them with the capacity to act as a physical barrier. Simultaneously, they promote the permeation of nutrients and the propagation of cells in a concurrent manner. Micrographs depicting the transverse cross-sections of the Gel and Gel-Th were presented in Figure 2H and Supplementary Figure S3, offering a detailed visualization of the internal structural features of the hydrogel. The results clearly demonstrated that the hydrogel possessed an exceedingly porous three-dimensional reticulated framework, and the aperture was all larger than 100 μm. Notably, the Gel10 featured large pores, with the majority of the pore sizes ranging from 90 μm to 180 µm (Figure 2I). Furthermore, we anticipate that these large pores will facilitate tissue ingrowth, thereby enhancing osseointegration during the bone regeneration process and stabilizing the mechanical integrity of the hydrogel (Mokhtarzadegan et al., 2021). The BET-BJH method will be employed in the future to supplement the analysis of the pore size and surface area of hydrogels, thereby enabling a more three-dimensional and multi-faceted demonstration of the porous structure of hydrogels.

Analysis of the TGA-DTG data showed that neither Gel nor Gel-Th exhibited significant mass change in the initial stage (from room temperature to 100 °C), which is mainly attributed to the loss of moisture. In the temperature range of 200 °C–300 °C, the mass change primarily resulted from the decomposition of DF-PEG due to the breakage of its polymer chains. Within 300 °C–400 °C, the mass change was mainly caused by the decomposition of CMCS as a result of the breakage of its polymer chains. These results indicated that the hydrogel material decomposes only at high temperatures, and its structure remained intact under the mild in vivo environment. For the DSC analysis, we focused on detecting the melting temperature of the hydrogel, which was a critical temperature parameter for the prepared hydrogel. The melting temperature can be used to determine whether the material will melt in vivo, thereby affecting its sealing and hemostatic functions. The results showed that the material starts to melt at around 100 °C, which was much higher than the in vivo temperature, thus meeting our application requirements. The above two sets of test results demonstrated that the prepared hydrogel had good thermal stability for our intended use (Supplementary Figure S4).

To evaluate the hemostatic attributes of Gel-Th specimens with precision, in vitro coagulation assays were executed using the blood clotting index (BCI) test. At the onset, a qualitative evaluation of the materials was conducted by subjecting them to incubation with blood and subsequently performing a rinse with deionized water. A lighter coloration of the water was indicative of more rapid clotting (Shi et al., 2020). The water encircling the Gel10 assumed a red coloration on account of the presence of non-coagulated blood, suggesting a limited blood coagulation competence of this material. In accordance with an increase in thrombin content, the gradual fading of the color around the Gel-Th specimens signified that the thrombin-laden hydrogel substantially reduced the blood clotting duration (Figure 4A). Furthermore, a quantitative analysis of the BCI was carried out, wherein a higher index was indicative of a diminished capacity for blood coagulation (Yu et al., 2023). The results of the investigation demonstrated that the BCI of the Gel-Th specimens decreased in tandem with an increasing thrombin concentration (Figure 4B). Notably, the BCI for the Gel10-Th200 (37.72%) was significantly lower than that of the Gel10 (62.08%) at the final measured time, suggesting the Gel10-Th200 enhanced the coagulation effect by promoting the conversion of fibrinogen.

Hemocompatibility represents a pivotal parameter in the formulation of efficacious hemostatic agents; thus, the hemolysis ratio was evaluated (Wang et al., 2017). As the concentration of thrombin increased, the hemolysis rate gradually decreased (Figure 4C). The hemolysis ratio for the Gel10-Th200 was approximately 0.34%, whereas the ratio for the Gel10 was higher (approximately 2.13%). These findings indicated that the Gel10-Th200 demonstrated good hemocompatibility, in alignment with national and international standards, which classify hemolysis ratios below 5% as indicative of excellent hemocompatibility.

The hemostatic mechanism encompasses the intricate interactions among erythrocytes, thrombocytes, coagulation factors, and other associated substances (Nepal et al., 2023). The hemostatic mechanism of CMCS/DF-PEG/thrombin hydrogel can be categorized into three primary processes (Figure 4D). Firstly, once the hydrogel solidified, its physical properties effectively occluded the wound, facilitating mechanical hemostasis. Secondly, the positively - charged surface of the hydrogel established an electrostatic “gravitational field” that prompted the aggregation of blood cells. Ultimately, through the synergistic effect of thrombin, fibrinogen was swiftly transformed into a three-dimensional fibrin reticulum, which further entrapped and encapsulated blood corpuscles, culminating in the formation of a stable blood clot. We devised the subsequent experiments to conduct a more in-depth exploration of the hydrogel’s function in the processes of primary and secondary hemostasis.

It is widely acknowledged that the aggregation of blood corpuscles assumes a pivotal role in primary hemostasis (Fan et al., 2023). The absorbance at 490 nm was measured, with higher absorbance values (indicating deeper coloration) representing greater LDH concentrations and a higher number of platelets adsorbed on the sample surface. This assay facilitated the quantitative assessment of the thrombocyte adhesion capacity on the specimen surface. The findings demonstrated that with an increase in the concentration of CMCS (Figure 4E), which led to a greater quantity of cations, the absorbance values determined for Gel8, Gel10, and Gel12 were 0.59915, 0.6752, and 0.74591 respectively. This indicated that these gels exhibited a better platelet aggregation capacity than Gel6 (0.39577). Moreover, due to the stimulation of blood cell aggregation by thrombin, the absorbance value of Gel10-Th200 (0.69312) showed a slight increase compared to Gel10 (0.6752).

The absorbance of the hydrogel group and the control group at 540 nm was compared, and the adhesion effect of erythrocytes was quantitatively analyzed. From the results (Figure 4F), Gel6, Gel8, Gel10, and Gel12 increased sequentially, being 39.6, 57.9, 63.6 and 80.6 respectively. Gel10-Th200 was also between Gel10 and Gel12, which explained the promoting effect of chitosan and thrombin components on erythrocyte adhesion.

Concurrently, the aggregation of blood cells on the material’s surface was visualized via laser-scanning confocal microscopy. The two-dimensional images were supplemented (Supplementary Figure S6A). There was a notable increase in the quantity of aggregated blood cells with the rising content of CMCS (Figure 4G). This phenomenon can be attributed to the positively charged amino groups in CMCS, which attracted negatively charged blood cells. It was worth mentioning that the hydrogel, cross-linked from 15% DF-PEG and 10% CMCS, exhibited a slight enhancement in the adsorption of blood cells when loaded with thrombin. This effect occurred because thrombin stimulated the coagulation process, thereby promoting the aggregation of blood cells. The phenomenon was consistent with the quantitative analysis results.

A highly dense fibrin reticulum, which effectively entrapped and securely encapsulated blood cells, gradually emerged on the surface of Gel-Th (Figure 4H). Conversely, only a tenuous fibrin meshwork was discernible on the surface of Gel. This marked disparity could potentially be attributed to the fact that thrombin enhanced the interfacial stimulatory capacity of Gel-Th, thereby potentially catalyzing the conversion of fibrinogen into fibrin, a process that is well-documented in reference (Wolberg and Campbell, 2008). This phenomenon signified the commencement of secondary coagulation. These empirical findings substantiated that thrombin within Gel-Th exerted a significant and notable influence on blood coagulation at the wound-material interface. The two-dimensional images were supplemented (Supplementary Figure S6B).

Considering that Gel10 and Gel10-Th200 were designated for direct utilization on hemorrhagic wounds, a comprehensive evaluation of their biocompatibility was carried out. After a 24-h period of incubation, the viability of cells within all experimental groups surpassed 90%, and no discernible cytotoxicity was identified in any of the groups (Figures 6A,B). The outcomes of the qualitative cytotoxicity examination following 24 h of incubation were presented. The green fluorescence served as an indication of viable cells within Gel10 and Gel10-Th200, while only a negligible quantity of non-viable cells (characterized by red fluorescence) was present throughout all the experimental sets (Figure 6C; Supplementary Figure S7A). These empirical observations substantiated that all the experimental groups manifested commendable cytocompatibility. Since Gel-Th is used for hemostasis in bone tissues, the biocompatibility of the hydrogel with bone cells was tested. There was no significant difference between the hydrogel group and the control group (Supplementary Figure S7B), indicating that the hydrogel is non-toxic to bone cells.

The venous blood was tested using an automatic blood analyzer, and the blood routine indicators were mainly the following: WBC, Neu, Lym, RBC, HGB and PLT. The test results showed that the indicators of all groups were within the normal range, indicating that the rabbits did not have such as inflammation, anemia, abnormal coagulation and other phenomena (Figure 6D). In addition, the in vivo safety was further confirmed through HE staining of five internal organs for observation (Figure 6E), indicating that there was no obvious inflammatory condition in the internal organs of each group.

To conduct a more in-depth assessment of the impacts of Gel-Th on hemostasis and bone restoration, rabbit models of bone bleeding and bone defects were utilized. Gel10 and Gel10-Th200 were administered to bleed tibial defects to evaluate their hemostatic efficacy (Figure 7A; Supplementary Figure S8). It was obvious that Gel10 and Gel10-Th200 could quickly stop bleeding and there was no secondary bleeding within 15 min and 3 days after hemostasis. However, in the blank group, from the beginning of drilling, there was still slight blood seepage after 15 min of pressing with gauze. After suture, blood clotted by itself, and hemostasis was also achieved later. The blood clotting times in the Gel10 and Gel10-Th200 groups were 45.95333 ± 0.77976 s and 18.9 ± 0.40825 s. Moreover, the initial occlusion of the bleeding site using Gel10 and Gel10-Th200 led to blood loss of 0.65567 ± 0.00464 g and 0.12933 ± 0.00478 g respectively (Figure 7B). This disparity was attributable to the fact that Gel10-Th200 exhibited a substantial capacity for secondary hemostasis, whereas the hemostatic efficacy of Gel10 primarily relied on its physical sealing characteristic.

The bone’s regenerative capacity is vital for successful implantation. Through Micro-CT scanning of the entrance site of the implant (Figure 7C). New bone tissue continued to grow around the hydrogel, and the bone repair effect of Gel10-Th200 was the best (Figure 7D). Secondly, in order to quantitatively analyze the ability of bone growth, the BV/TV, BMD, Tb. SP and Tb. N of this part of the bone tissue were tested (Figure 7E). It can be seen that the BV/TV, BMD and Tb. N (bone volume fraction, bone density, number of trabeculae) of the Gel10-Th200 group were all greater than those of Gel10 and the control group, but the Tb. SP (trabecular separation) was less than those of the other two groups, indicating that the Gel10-Th200 group had better bone repair ability. The good biocompatibility of the hydrogel itself and the repair-promoting ability of thrombin component can better promote wound repair.

Histological analysis was performed at the second and fourth weeks post-implantation of the hydrogel. The tibia was stained with HE and Masson to assess inflammation and osteogenic response in vivo (Figure 7F). As shown by HE staining, no obvious inflammatory response was observed in all groups, and Gel10-Th200 gradually formed abundant bone trabeculae and blood vessels. According to Masson staining, the area surrounding the Gel10-Th200 was characterized by abundant bone tissue and denser connective tissue. Studies have proven that thrombin can activate the receptors on cells and promote the proliferation and migration of fibroblasts and endothelial cells. This is conducive to the formation of collagen and blood vessels, providing favorable conditions for the repair of bone tissue (Moura et al., 2024; Kleeschulte et al., 2018). In conclusion, Micro-CT and histological staining analysis indicated that the Gel10-Th200 hydrogel had excellent safety and tissue repair capability. Future studies will involve a comprehensive comparison between Gel10-Th200 and currently available commercial products, as well as other advanced hemostatic hydrogels reported in the literature. We will also actively investigate the microenvironmental mechanisms underlying Gel10-Th200 roles in both hemorrhage control and osteogenic repair. More large-animal models—such as sheep, canine, or porcine models, which better approximate human bone anatomy and pathological conditions—will be employed to enable in-depth mechanistic validation and thorough performance evaluation. These steps are essential to establishing a solid foundation for the eventual clinical translation of our material. Furthermore, we aim to optimize the material’s properties to better synchronize with the tissue regeneration process, and explore the integration of multiple functions to develop a new generation of injectable, intracavity hemostatic hydrogels capable of in situ gelation. These next-generation systems are intended to be intelligent, self-adapting, and capable of multi-target interventions.