Traceable hydrogel for cartilage repair and use thereof

Abstract

This invention discloses a traceable hydrogel for cartilage repair and its applications. The hydrogel, composed of chitosan, sericin, and collagen, forms a three-dimensional biomimetic network, within which a magnetic material is uniformly dispersed. The magnetic material is selected from nano-Fe3O4 or γ-Fe2O3. Addressing the shortcomings of existing hydrogels, such as insufficient mechanical strength, inability to perform non-invasive tracking, easy aggregation of magnetic nanoparticles, and safety risks associated with chemical cross-linking, the traceable hydrogel prepared by this invention can improve the compressive modulus of the hydrogel, promote the deposition of cartilage-specific matrix, and simultaneously retain the complete magnetic signal activity while avoiding the toxicity of chemical cross-linking. The hydrogel can rapidly gel at physiological body temperature, making it suitable for minimally invasive clinical injection, providing a new solution for cartilage defect repair that combines functionality and safety.

Description

Technical Field

[0001] This invention relates to the field of medical materials, specifically to a traceable hydrogel for cartilage repair and its applications. Background Technology

[0002] Articular cartilage defects are a common musculoskeletal disease. Because cartilage lacks blood vessels and nerves, its self-repair ability is extremely weak, and effective physiological spontaneous repair is impossible after a defect is found. Currently, commonly used clinical methods for repairing cartilage defects include microfracture surgery and autologous / allogeneic cartilage transplantation. However, these methods all have unavoidable problems. The regenerated tissue obtained is often fibrocartilage with poor mechanical properties. Furthermore, there are many issues such as limited donor availability and immune rejection, making it difficult to meet the needs of long-term clinical treatment.

[0003] The development of tissue engineering technology has provided a new technical approach for the repair of articular cartilage defects. Among them, hydrogel materials, with their high water content, three-dimensional network structure and excellent biocompatibility, have become one of the core research directions in the field of cartilage tissue engineering.

[0004] However, existing hydrogel systems still have significant technical problems in cartilage repair applications: on the one hand, some hydrogels have insufficient mechanical properties and weak overall mechanical strength, making it difficult to meet the long-term mechanical support requirements under the physiological environment of joint weight-bearing; on the other hand, existing technology systems lack effective in vivo monitoring methods for implanted materials and repair processes. After hydrogels are implanted in the body, it is difficult to achieve continuous and non-invasive dynamic observation of their spatial distribution, degradation behavior and cartilage tissue regeneration process.

[0005] Currently, most mainstream methods for evaluating repair effectiveness use endpoint detection modes such as tissue sections, which cannot achieve continuous tracking of the repair process. While imaging examinations such as MRI can achieve non-invasive detection, they have limitations such as high detection costs, inconvenient operation, and difficulty in high-frequency repetition. They cannot provide real-time feedback on the repair progress of the defect site, which is not conducive to timely adjustment of clinical treatment plans and optimization and improvement of hydrogel material performance. Summary of the Invention

[0006] To address the aforementioned technical problems, this invention provides a traceable hydrogel for cartilage repair and its application, thereby solving the technical problems in existing hydrogel systems for cartilage repair applications, such as insufficient mechanical properties, difficulty in meeting the long-term mechanical support requirements under the physiological environment of joint weight-bearing, and inability to continuously and non-invasively observe the spatial distribution, degradation behavior, and cartilage tissue regeneration process of implanted hydrogels.

[0007] To achieve the above-mentioned technical objectives, in a first aspect, the present invention provides a traceable hydrogel for cartilage repair, comprising chitosan, sericin and collagen forming a hydrogel with a three-dimensional biomimetic network, wherein a magnetic material is uniformly dispersed within the hydrogel; wherein the magnetic material is selected from nano Fe3O4 or γ-Fe2O3.

[0008] Preferably, the traceable hydrogel is prepared by the following method:

[0009] Step 1: Mix chitosan and magnetic material, then add to deionized water. Next, add sericin and collagen under continuous stirring until completely dissolved to obtain a composite precursor solution. The concentrations of chitosan, sericin, collagen, and magnetic material in the composite precursor solution are as follows (by mass percentage): 4% chitosan, 1%–2% sericin, 1%–2% collagen, and 0.2%–0.8%.

[0010] Step 2: Slowly add a mixed solution of sodium β-glycerophosphate and NaHCO3 to the composite precursor solution obtained in Step 1, stir evenly, and then inject it into the target location to form the traceable hydrogel; wherein, the mass ratio of sodium β-glycerophosphate to NaHCO3 is 10:1, and the mass ratio of chitosan to sodium β-glycerophosphate is (1:2) to (1:3).

[0011] Preferably, in step 1, chitosan is dissolved in a dilute acetic acid solution or a vitamin C solution; wherein the concentration of the dilute acetic acid or vitamin C solution is 1 wt%.

[0012] Preferably, in step 1, the magnetic material is added to an aqueous solution containing polyethylene glycol, vortexed for 3 to 5 minutes, and then ultrasonically treated for 10 to 30 minutes to obtain a stable Fe3O4 dispersion; wherein the concentration of PEG in the polyethylene glycol aqueous solution is 1 wt% to 5 wt%.

[0013] Preferably, in step 2, the traceable hydrogel is formed by standing at 37°C for 30 to 60 minutes.

[0014] Secondly, an application of a traceable hydrogel, which is used to prepare implants for treating cartilage defects. Specifically, it is delivered to the treatment site via injection.

[0015] Compared with the prior art, the present invention has the following beneficial effects:

[0016] 1. This invention uses a three-dimensional biomimetic network constructed from chitosan, sericin, and collagen as a hydrogel matrix. It possesses high water content and good biocompatibility similar to the natural extracellular matrix of chondrocytes, providing a suitable three-dimensional microenvironment for chondrocyte adhesion, proliferation, and cartilage-specific matrix secretion. On this basis, magnetic nanoparticles within a limited concentration range act as physical cross-linking points by forming hydrogen bonds and electrostatic interactions with the polymer chains in the hydrogel network, significantly improving the compressive modulus of the hydrogel. This enables it to meet the long-term mechanical support requirements of the joint's weight-bearing physiological environment. The improved mechanical properties further ensure the stability of the three-dimensional network structure during the hydrogel's degradation in vivo, providing stable spatial support for the continuous regeneration of cartilage tissue.

[0017] 2. The introduction of magnetic nanoparticles in this invention simultaneously endows the hydrogel with stable optical and magnetic dual-mode signals. This enables quantitative quality control of the material through in vitro absorbance detection, and also allows for non-invasive, continuous, and dynamic tracking of the implanted material's distribution, degradation behavior, and cartilage regeneration process within the body through histological staining and MRI imaging. This solves the technical pain point of existing technologies where the cartilage repair process cannot be monitored in real time. This solution uses a thermosensitive physical crosslinking system composed of sodium β-glycerophosphate and sodium bicarbonate, which not only enables in-situ gelation of the hydrogel at physiological body temperature, giving the material minimally invasive injectability and allowing for precise filling of irregular cartilage defects and improved fit to the defect site, but also avoids the cytotoxicity of residual chemical crosslinking agents and the damage to the crystal structure and signal activity of magnetic nanoparticles caused by the crosslinking reaction, thus simultaneously ensuring the material's biosafety and the stability of its tracking performance.

[0018] 3. This invention clarifies the concentration window for achieving optimal comprehensive performance of magnetic nanoparticles, providing clear quantitative guidance for material quality control and clinical application safety. Within the preferred concentration range, the relative cell proliferation rate of the material meets the requirements for medical biomaterials, the hemolysis rate is less than 5%, and the inflammatory response in vivo after implantation is mild and can subside spontaneously over time, demonstrating excellent biosafety. Attached Figure Description

[0019] Figure 1 The images shown are scanning electron microscope (SEM) images of the hydrogels prepared for the comparative examples and the examples. Among them, the Control group is the comparative example 1, HG-0.25 is the example 1, HG-0.5 is the example 2, and HG-1 is the comparative example 2.

[0020] Figure 2 The swelling ratios of the hydrogels prepared in the comparative examples and the examples are shown below; where HG-0 is comparative example 1, HG-0.25 is example 1, HG-0.5 is example 2, and HG-1 is comparative example 2.

[0021] Figure 3The compression modulus of the hydrogels prepared in the comparative examples and examples is given; where HG-0 is comparative example 1, HG-0.25 is example 1, HG-0.5 is example 2, and HG-1 is comparative example 2.

[0022] Figure 4 The hysteresis loop diagrams are for the hydrogels prepared in the comparative examples and the examples; where HG-0 is comparative example 1, HG-0.25 is example 1, HG-0.5 is example 2, and HG-1 is comparative example 2.

[0023] Figure 5 The images show the Prussian blue staining results of subcutaneous implantation of hydrogels prepared for comparative and example studies; where Control group is comparative example 1, HG-0.25 is example 1, HG-0.5 is example 2, and HG-1 is comparative example 2.

[0024] Figure 6 Prussian blue staining results of hydrogel scaffold materials prepared in comparative examples and examples implanted into cartilage defects in rats; where Blank represents cartilage defects without any implantation, Control group represents implantation comparative example 1, HG-0.25 represents implantation example 1, HG-0.5 represents implantation example 2, and HG-1 represents implantation comparative example 2.

[0025] Figure 7 The relative proliferation rates of L929 cells cultured with the hydrogel extracts prepared in the comparative examples and the examples for 1, 2, and 3 days were calculated (n=3, *p<0.05, **p<0.01). Among them, Blank was the blank group without hydrogel, Control group was the group with comparative example 1, HG-0.25 was the group with example 1, HG-0.5 was the group with example 2, and HG-1 was the group with comparative example 2.

[0026] Figure 8 The hemolysis rate of the hydrogels prepared in the comparative examples and examples is shown below; where HG-0 is comparative example 1, HG-0.25 is example 1, HG-0.5 is example 2, and HG-1 is comparative example 2.

[0027] Figure 9 The H&E staining results are shown for the hydrogels prepared for subcutaneous implantation in the comparative and example examples; where the Control group is comparative example 1, HG-0.25 is example 1, HG-0.5 is example 2, and HG-1 is comparative example 2.

[0028] Figure 10HE staining results of hydrogel scaffold materials prepared in comparative examples and examples implanted into cartilage defects in rats; where Blank represents cartilage defects without any implantation, Control group represents implantation comparative example 1, HG-0.25 represents implantation example 1, HG-0.5 represents implantation example 2, and HG-1 represents implantation comparative example 2.

[0029] Figure 11 Safranin and Fast Green staining results of hydrogel scaffold materials prepared in comparative examples and examples implanted into cartilage defects in rats; where Blank represents cartilage defects without any implantation, Control group represents implantation comparative example 1, HG-0.25 represents implantation example 1, HG-0.5 represents implantation example 2, and HG-1 represents implantation comparative example 2.

[0030] Figure 12 The results of type II collagen immunohistochemistry after implantation of the hydrogel scaffold materials prepared in the comparative examples and examples into cartilage defects in rats are shown. Among them, Blank is the cartilage defect without any material implanted, Control group is implantation comparative example 1, HG-0.25 is implantation example 1, HG-0.5 is implantation example 2, and HG-1 is implantation comparative example 2. Detailed Implementation

[0031] The technical solutions of the present invention will be clearly and completely described in conjunction with the accompanying drawings of the embodiments of the present invention. Obviously, the described embodiments are only some embodiments of the present invention, and not all embodiments. All other embodiments obtained by those skilled in the art based on the present invention are within the scope of protection of the present invention.

[0032] Unless otherwise specified in the specific circumstances, the numerical ranges listed herein include upper and lower limits, as well as all integers and fractions within that range, but are not limited to the specific values ​​listed when the range is defined.

[0033] I. A traceable hydrogel for cartilage repair

[0034] This invention addresses the technical problems of existing hydrogel systems for cartilage repair, such as the difficulty in matching the mechanical strength to the long-term support requirements of the physiological environment of joint weight-bearing, and the inability to achieve non-invasive, continuous, and dynamic monitoring of the in vivo behavior of the material and the cartilage repair process after implantation. It also addresses the technical challenges of magnetic nanoparticles easily agglomerating in the hydrogel matrix, and the potential for conventional cross-linking systems to damage the crystalline structure and signal activity of magnetic materials, posing biosafety risks. This invention constructs a three-dimensional biomimetic network matrix for the hydrogel using chitosan, sericin, and collagen, which are highly homologous to the extracellular matrix components of cartilage cells. Magnetic nanomaterials are then introduced into this matrix system. To address the dispersibility issue of magnetic nanoparticles, polyethylene glycol is used as a dispersant for pretreatment. To mitigate the impact of the cross-linking system on the activity and biosafety of the magnetic material, a thermosensitive physical cross-linking system composed of sodium β-glycerophosphate and sodium bicarbonate is used to achieve in-situ gelation of the hydrogel. Furthermore, by limiting the concentration ratio of each component in the composite precursor solution, a hydrogel system with both cartilage repair function and in vivo tracking performance is constructed. Therefore, the traceable hydrogel of the present invention is a hydrogel with a three-dimensional biomimetic network formed by chitosan, sericin and collagen, and a magnetic material is uniformly dispersed in the hydrogel; wherein the magnetic material is selected from nano Fe3O4 or γ-Fe2O3.

[0035] The traceable hydrogel described in this invention has achieved many unexpected technical effects after clinical application. In conventional understanding within this field, magnetic nanoparticles in hydrogel systems are merely used as contrast agent components to achieve imaging tracking. However, this invention is the first to discover that Fe3O4 at specific concentrations produces a triple synergistic effect on the hydrogel system: "mechanical enhancement, repair promotion, and process tracking." The traceable hydrogel described in this invention not only possesses excellent imaging tracking capabilities but also acts as a physical cross-linking point through hydrogen bonding and electrostatic interactions with the polymer chains in the three-dimensional network of the hydrogel, thereby increasing the compressive modulus of the hydrogel. Simultaneously, magnetic nanoparticles within this concentration range promote the specific deposition of glycosaminoglycans and type II collagen in the regenerated tissue of cartilage defect areas within the in vivo implantation environment, making the structure and composition of the regenerated tissue closer to natural hyaline cartilage. Furthermore, the aforementioned mechanical regulation, cartilage regeneration regulation, and tracking imaging effects are strictly concentration-dependent. The synergistic effect of the magnetic nanoparticles can only be achieved within a limited concentration window. Beyond this range, the nanoparticles aggregate, leading to a decrease in the mechanical properties and biocompatibility of the hydrogel. Pretreatment of the magnetic nanoparticles with polyethylene glycol not only achieves uniform dispersion of the nanoparticles in the hydrogel matrix through steric hindrance but also forms a flexible interface layer between the nanoparticles and the polymer network, improving the stress transfer efficiency between the nanoparticles and the matrix. This transforms the inherent defect of nanoparticle agglomeration into a favorable factor for enhancing the mechanical properties of the hydrogel. Furthermore, the physical crosslinking system of sodium β-glycerophosphate and sodium bicarbonate achieves in-situ gelation at physiological body temperature without damaging the spinel crystal structure of the magnetic nanoparticles, fully preserving their magnetic and optical signal activity. At the same time, it avoids the cytotoxic risks caused by residual chemical crosslinking agents, achieving simultaneous optimization of magnetic particle dispersibility, signal activity, material mechanical properties, and biocompatibility.

[0036] In some embodiments of the present invention, the traceable hydrogel is prepared by the following method:

[0037] Step 1: Mix chitosan and magnetic material, then add to deionized water. Next, add sericin and collagen under continuous stirring until completely dissolved to obtain a composite precursor solution. The concentrations of chitosan, sericin, collagen, and magnetic material in the composite precursor solution are as follows (by mass percentage): 4% chitosan, 1%–2% sericin, 1%–2% collagen, and 0.2%–0.8%.

[0038] Step 2: Slowly add a mixed solution of sodium β-glycerophosphate and NaHCO3 to the composite precursor solution obtained in Step 1. After stirring evenly, inject it into the target location to form the traceable hydrogel; wherein the mass ratio of sodium β-glycerophosphate to NaHCO3 is 10:1, and the mass ratio of chitosan to sodium β-glycerophosphate is (1:2) to (1:3). In specific implementation, the order of adding chitosan and magnetic materials first, followed by adding sericin and collagen, allows the magnetic materials to be initially dispersed in the cationic polymer chain environment of chitosan, avoiding the steric hindrance of protein molecules that hinders the binding of magnetic materials and chitosan molecular chains. This ensures that the magnetic materials can be uniformly anchored in the subsequently formed three-dimensional network framework, achieving a uniform distribution of tracer signals and mechanical enhancement sites. The chitosan concentration is fixed at 4%. As the main framework material of the three-dimensional hydrogel network, this concentration is the optimal basic concentration for constructing a hydrogel with basic structural stability and injectability. If the concentration is too low, the hydrogel framework structure will be loose and unable to form a continuous three-dimensional interpenetrating network, making it difficult to meet the basic mechanical support requirements for cartilage defect repair. If the concentration is too high, the viscosity of the precursor solution will be too high, which will not only be detrimental to the uniform dispersion of magnetic materials and protein components, but will also significantly reduce the injectability of the material, making it unsuitable for the application requirements of minimally invasive injection filling in clinical practice. The concentrations of sericin and collagen were each limited to 1%–2%. Both are natural macromolecules with high homology to the extracellular matrix of chondrocytes, which can significantly enhance the cell affinity of the hydrogel and the adhesion and proliferation ability of chondrocytes. When combined with chitosan, they can form an interpenetrating biomimetic three-dimensional network. This concentration range can form a matching interpenetrating structure with 4% chitosan backbone. If the concentrations of both are too low, the modification and enhancement of the chitosan network cannot be achieved, resulting in insufficient bioactivity of the hydrogel and difficulty in effectively promoting the secretion of cartilage-specific matrix. If the concentrations of both are too high, it will lead to precursor dissolution. Excessive aggregation of protein molecules in the liquid disrupts the pore uniformity of the hydrogel network and causes uncontrollable gelation time. The excessively rapid in vivo degradation rate cannot match the regeneration cycle of cartilage tissue, resulting in a lack of mechanical support during regeneration. The synergistic combination of the two within this concentration range can simultaneously ensure the structural stability, cell affinity, and degradation rate matching of the hydrogel network. The resulting interpenetrating protein network can form abundant hydrogen bonds with the chitosan backbone, further enhancing the structural integrity of the three-dimensional network and compensating for the deficiencies in mechanical properties and bioactivity of a single chitosan network.The concentration of magnetic materials is limited to 0.2% to 0.8%. As the tracer and mechanical enhancement component of the hydrogel, this concentration range is the optimal window for achieving synergy among tracer performance, mechanical enhancement, and biosafety. If the concentration is below 0.2%, the hydrogel cannot obtain sufficient magnetic and optical signal intensity, resulting in insufficient signal-to-noise ratio for in vitro quantitative detection, and inability to obtain clear and distinguishable positive signals for in vivo histological staining and MRI imaging. At the same time, it is difficult to form sufficient physical cross-linking points, making it difficult to achieve the mechanical enhancement effect. However, if the concentration is above 0.8%, the nanoparticles will aggregate due to high surface energy, which not only destroys the uniformity of the three-dimensional network of the hydrogel, leading to a decrease in mechanical properties, but also reduces the cell compatibility of the material and exacerbates the inflammatory response in vivo. In addition, the aggregated nanoparticles will cause uneven distribution of magnetic signals, affecting the accuracy of tracer imaging. A thermosensitive physical crosslinking system composed of sodium β-glycerophosphate and NaHCO3, with a mass ratio of 10:1, can synergistically regulate the pH and gelation temperature of the system, enabling rapid in-situ gelation of the hydrogel at physiological body temperature. Sodium β-glycerophosphate, as the core thermosensitive crosslinking agent, achieves temperature-responsive sol-gel transition through hydrogen bonding and electrostatic interactions with the amino groups of chitosan molecules. NaHCO3, as a pH buffer, neutralizes the acidity of the chitosan solution, adjusting the system to a near-physiological neutral range, thus preventing local inflammatory reactions caused by implantation of an acidic system. If the mass ratio deviates from 10:1, an excessively high proportion of NaHCO3 will lead to an excessively high pH, ​​causing chitosan precipitation and decreased stability of the precursor solution. Conversely, an excessively low proportion of NaHCO3 will fail to effectively neutralize the pH and increase the gelation temperature, preventing in-situ gelation at physiological body temperature. The mass ratio of chitosan to sodium β-glycerophosphate is limited to (1:2) to (1:3). A chitosan concentration of 4% achieves optimal cross-linking efficiency and gelation performance. If the ratio is lower than 1:2, insufficient sodium β-glycerophosphate will result in inadequate cross-linking, preventing the formation of a structurally stable hydrogel, significantly prolonging gelation time, or even preventing gel formation altogether. If the ratio is higher than 1:3, excessive sodium β-glycerophosphate will lead to over-cross-linking, resulting in a hard, brittle hydrogel with reduced mechanical compatibility with natural articular cartilage. Furthermore, excessively high cross-linking density reduces the porosity of the hydrogel, hindering cell infiltration and nutrient transport, thus impeding cartilage regeneration. In addition, the cross-linking system ratio exhibits synergistic compatibility with the concentrations of each component in the precursor solution. The matched degree of cross-linking allows the magnetic material to be uniformly locked within a three-dimensional network, preventing leakage in vivo. Simultaneously, the stable network structure ensures the slow release of protein components, continuously promoting cartilage regeneration.

[0039] In some embodiments of the present invention, in step 1, chitosan is dissolved in a dilute acetic acid solution; wherein the concentration of the dilute acetic acid is 1 wt%. Using a 1 wt% dilute acetic acid solution to dissolve chitosan, as a cationic polysaccharide, allows the molecular chains of chitosan to expand and dissolve uniformly in an acidic aqueous solution, achieving a thorough and uniform dissolution of chitosan and forming a chitosan solution with stable molecular chains. This lays a uniform matrix foundation for the subsequent construction of the three-dimensional network. If the concentration of dilute acetic acid is below 1 wt%, the system is insufficiently acidic, and chitosan cannot be completely dissolved, resulting in undissolved flocculent precipitates. This leads to decreased homogeneity of the precursor solution, preventing the formation of a continuous and uniform three-dimensional network structure. It also affects the uniform dispersion of magnetic materials and protein components in the system. If the concentration of dilute acetic acid is above 1 wt%, the system is too acidic, requiring the addition of more NaHCO3 for pH neutralization. This not only disrupts the original cross-linking system's ratio balance and affects the stability of gelation performance but also results in excessively high ionic strength in the final system, impacting the dispersion stability of magnetic nanoparticles. Furthermore, an excessively acidic environment can cause molecular chain denaturation of sericin and collagen, damaging their bioactivity and reducing the cell affinity of the hydrogel. In addition, chitosan dissolved in a 1 wt% dilute acetic acid solution exhibits optimal molecular chain extension, enabling it to fully form intermolecular hydrogen bonds and electrostatic interactions with magnetic materials and protein components, ensuring the uniform formation of the subsequent interpenetrating network structure.

[0040] In some embodiments of the present invention, in step 1, the magnetic material is added to an aqueous solution containing polyethylene glycol, vortexed for 3 min to 5 min, and then ultrasonically treated for 10 min to 30 min to obtain a stable Fe3O4 dispersion; wherein the concentration of PEG in the polyethylene glycol aqueous solution is 1 wt% to 5 wt%. This invention uses a polyethylene glycol (PEG) aqueous solution to pretreat magnetic materials, and combines vortex oscillation and ultrasonic treatment processes to prepare a stable Fe3O4 dispersion. Addressing the issue of Fe3O4 nanoparticles' high surface energy easily agglomerating in aqueous systems, PEG is physically adsorbed onto the surface of Fe3O4 nanoparticles through hydrophobic segments. The hydrophilic long chains extend into the aqueous solution, effectively preventing nanoparticle aggregation and sedimentation through steric hindrance, thus achieving long-term stable dispersion in the aqueous system. The PEG concentration is limited to 1wt%~5wt%. This concentration range allows for the formation of a complete and continuous adsorption coating layer on the surface of Fe3O4 nanoparticles. If the PEG concentration is below 1wt%, it cannot adequately coat the nanoparticles, resulting in insufficient steric hindrance and difficulty in effectively preventing nanoparticle aggregation, thus failing to obtain a stable monodisperse. If the PEG concentration is above 5wt%, it leads to excessive system viscosity and excessive entanglement between PEG molecular chains, which in turn induces nanoparticle aggregation. Furthermore, excessive PEG hinders the intermolecular interactions between chitosan, protein components, and Fe3O4 nanoparticles, weakening their mechanical reinforcing effect as physical cross-linking points. Meanwhile, the dispersion process employed in this invention, combining 3-5 minutes of vortex oscillation with 10-30 minutes of ultrasonic treatment, first achieves preliminary uniform mixing of the PEG aqueous solution and Fe3O4 nanoparticles through 3-5 minutes of vortex oscillation, breaking up the macroscopic agglomerates of nanoparticles. If the vortex time is less than 3 minutes, preliminary and sufficient mixing cannot be achieved, and the agglomerates cannot be effectively dispersed, making it difficult to achieve uniform dispersion at the nanoscale during subsequent ultrasonic treatment. If the vortex time is greater than 5 minutes, it cannot further improve the dispersion effect; instead, it will lead to an increase in system temperature, affecting the adsorption stability of PEG on the surface of nanoparticles. The ultrasonic treatment lasts for 10 to 30 minutes. Ultrasonic cavitation breaks down the micro-agglomeration between nanoparticles, allowing PEG molecular chains to be fully adsorbed onto the fresh surface of the nanoparticles, forming a monodisperse nanoparticle dispersion. If the ultrasonic treatment time is less than 10 minutes, the micro-agglomeration of nanoparticles cannot be completely broken, resulting in insufficient dispersion stability and easy sedimentation during subsequent gelation. If the ultrasonic treatment time is more than 30 minutes, the system temperature will continue to rise, which will not only destroy the molecular chain structure of PEG, but also cause oxidation of the Fe3O4 nanoparticle surface, destroying its spinel crystal structure and leading to a decrease in magnetic signal activity.Moreover, the PEG concentration and dispersion process have a synergistic effect, which can obtain a long-term stable Fe3O4 monodisperse, ensuring that it can be uniformly dispersed in the subsequent precursor solution. This leads to the uniform distribution of mechanical enhancement sites and tracer signals in the hydrogel, avoiding the problems of decreased mechanical properties and uneven tracer signals caused by particle agglomeration. At the same time, the flexible interface layer formed by PEG on the surface of nanoparticles achieves good stress transfer between nanoparticles and hydrogel polymer network, further enhancing the mechanical properties of hydrogel.

[0041] In some embodiments of the present invention, in step 2, the traceable hydrogel is formed by standing at 37°C for 30 to 60 minutes. 37°C is the physiological body temperature, which perfectly matches the temperature of the implantation environment, ensuring consistency between the gel-forming properties of the hydrogel prepared in vitro and those formed in situ in vivo. This provides predictable and repeatable gelation behavior for clinical applications. The 30 to 60 minute standing gelation time is the optimal gelation period for matching the aforementioned crosslinking system ratio and precursor component concentration. This time range ensures that the hydrogel completes sufficient physical crosslinking reactions, forming a stable and uniform three-dimensional network structure. If the gelation time is less than 30 minutes, the cross-linking reaction is not fully completed, the three-dimensional network structure of the hydrogel is not fully formed, and the mechanical strength is insufficient. After implantation, it is prone to collapse and cannot achieve stable filling and support of the defect site. At the same time, the insufficient cross-linking system will lead to the excessive leakage of magnetic materials and protein components, affecting the stability of the tracking performance and the sustainability of the cartilage repair effect. If the gelation time is more than 60 minutes, the cross-linking reaction is fully completed. Continuing to let it stand will not further improve the structural stability of the hydrogel. Instead, it will increase the time cost of in vitro operation. At the same time, in clinical application scenarios, the excessively long gelation time corresponds to the slow gelation rate in vivo, which will cause the injected material to fail to solidify quickly and easily flow out from the cartilage defect site, making it impossible to achieve precise defect filling. The gelation process at 37℃ for 30-60 minutes is a mild physical cross-linking process that will not damage the spinel crystal structure of the magnetic material or the bioactivity of the protein components. It simultaneously ensures the stable performance of the hydrogel's tracer properties, mechanical properties, and bioactivity. Furthermore, this gelation condition is synergistically compatible with the aforementioned cross-linking system ratio, enabling precise control of the gelation rate. This ensures a sufficient time window for clinical injection and allows for rapid in-situ solidification after implantation, fully meeting the application requirements of minimally invasive clinical injection.

[0042] II. An application of a traceable hydrogel

[0043] The traceable hydrogel described in this invention is used to prepare implants for treating cartilage defects. Specifically, it is delivered to the treatment site via injection.

[0044] III. Examples and Comparative Examples

[0045] Example 1

[0046] Step 1: Dissolve chitosan powder in a solution containing 1% glacial acetic acid to prepare a 4% (w / v) chitosan solution;

[0047] Step 2: Dissolve or disperse NaHCO3 in distilled water to prepare a NaHCO3 solution with a concentration of 0.6 mol / L;

[0048] Step 3: Dissolve or disperse sodium β-glycerophosphate in the above sodium bicarbonate solution to prepare a 50% (w / v) sodium β-glycerophosphate / sodium bicarbonate mixed solution;

[0049] Step 4: Stir the chitosan solution under ice bath conditions, and gradually add the above β-glycerophosphate sodium / sodium bicarbonate mixed solution (the volume ratio of chitosan solution to β-glycerophosphate sodium / sodium bicarbonate mixed solution is 30:100), and stir for 15 min;

[0050] Step 5: After mixing the above ingredients evenly, continue stirring, then pour in sericin powder (the mass ratio of sericin to chitosan is 25:100) and collagen powder (the mass ratio of collagen to chitosan is 25:100), and continue stirring for 10 minutes to obtain the thermosensitive sol-gel conversion material.

[0051] The operating procedures for Comparative Examples 1 and 2 were the same as those for Examples 1 and 2, except for the content of Fe3O4, as shown in Table 1. Meanwhile, the sample prepared in Comparative Example 1 was labeled HG-0, the sample in Example 1 was labeled HG-0.25, the sample in Example 2 was labeled HG-0.5, and the sample in Comparative Example 2 was labeled HG-1.

[0052] Table 1

[0053] Group Chitosan (wt%) sericin egg (wt%) Collagen (wt%) <![CDATA[Fe3O4(wt%)]]> Comparative Example 1 4 1.5 2.5 0 Example 1 4 1.5 2.5 0.25 Example 2 4 1.5 2.5 0.5 Comparative Example 2 4 1.5 2.5 1

[0054] IV. Performance Analysis

[0055] 1. Microscopic morphology

[0056] Figure 1 The images show scanning electron microscope (SEM) images of the hydrogels prepared in Comparative Examples 1-3 and Example 1. From... Figure 1As can be seen from the results: Comparative Example 1 (Control group) shows dense but small pores (mostly 10-30 μm) at low magnification, with relatively uniform distribution but thin pore walls; at high magnification, the pore wall surface is smooth, the three-dimensional network structure is relatively loose, and the pore connectivity is generally poor. In Example 1, the pore size is significantly increased (30-50 μm) at low magnification, the pore wall is significantly thickened, and the three-dimensional network is enhanced; at high magnification, uniform fine texture appears on the pore wall surface, indicating that Fe3O4 nanoparticles have been uniformly dispersed in the polymer matrix. Example 2 presents the optimal three-dimensional porous structure: at low magnification, the pore size is further increased to 50-80 μm, the pore shape is regular and nearly circular, the pore walls are continuous and uniform in thickness, the pore connectivity is optimal, and a well-developed three-dimensional interpenetrating network is formed; at high magnification, the pore walls are dense and thick, the surface texture is uniform, and there is no obvious particle agglomeration. Comparative Example 2 shows significant degradation of the three-dimensional network structure: at low magnification, the pore size is abnormally increased (exceeding 100 μm) and the distribution is extremely uneven, with some areas showing pore wall fusion and structural collapse; at high magnification, the pore wall surface is rough, with obvious nanoparticle aggregates visible, and the network structure becomes disordered and broken.

[0057] Therefore, the addition of an appropriate amount of Fe3O4 nanoparticles (0.25%~0.5%) significantly thickens the pore walls of the hydrogel and makes the three-dimensional network more developed and continuous. This is because the Fe3O4 nanoparticles bind to the molecular chains of chitosan, sericin, and collagen through hydrogen bonds and electrostatic interactions, acting as additional physical cross-linking points and promoting the entanglement and interpenetrating network formation of polymer chains. This optimization of the microstructure directly corresponds to the improvement of macroscopic mechanical properties, verifying the core finding of this invention that "Fe3O4 nanoparticles are not only a contrast agent but also a mechanically enhancing component." When the concentration of Fe3O4 nanoparticles reaches 1% (exceeding the optimal window of 0.2%-0.8% specified in this invention), the nanoparticles exhibit significant aggregation, leading to severe deterioration of the three-dimensional network structure of the hydrogel. This also proves that when the concentration exceeds the specified range, the aggregation effect of nanoparticles will destroy the uniformity of the hydrogel network, resulting in decreased mechanical properties, reduced biocompatibility, and uneven tracer signals, demonstrating that limiting the concentration window is a necessary condition for achieving synergistic effects of multiple functions. The uniform texture and absence of obvious agglomerates on the pore wall surface in the high-magnification electron microscopy images of Examples 1 and 2 demonstrate that the preparation method described in this invention can effectively prevent the aggregation of Fe3O4 nanoparticles and achieve their uniform dispersion in the hydrogel matrix. This also verifies the technical effect of this invention in "transforming the inherent defect of easy aggregation of nanoparticles into a favorable factor for mechanical enhancement." In particular, Example 2 presents an ideal microstructure for cartilage tissue regeneration, characterized by a pore size of 50-80 μm, continuous pore walls, and good connectivity: this pore size range is conducive to the infiltration, adhesion, and proliferation of chondrocytes, while ensuring the transport of nutrients and metabolic waste; simultaneously, the thick and uniform pore walls provide long-term stable mechanical support, matching the physiological load-bearing requirements of articular cartilage. This provides direct microstructural evidence for the in vivo experimental results showing that appropriate amounts of Fe3O4 promote cartilage-specific matrix deposition and accelerate cartilage regeneration. All examples maintained the complete three-dimensional porous structure, without obvious structural damage or decrease in particle crystallinity caused by chemical crosslinking. This indirectly proves that the β-glycerophosphate sodium / sodium bicarbonate thermosensitive physical crosslinking system used in this invention does not destroy the spinel crystal structure of Fe3O4, can completely preserve its magnetic signal activity, and avoids the cytotoxicity risk of chemical crosslinking agents.

[0058] 2. Swelling properties

[0059] The swelling properties of the four groups of hydrogel scaffold materials obtained from the comparative example and the embodiment were tested using a gravimetric method. The results are as follows: Figure 2 As shown. The lyophilized scaffold samples were immersed in PBS buffer solution at 37°C. Samples were removed and weighed at set time intervals, and the weight was calculated according to the formula SR=(W t The swelling ratio is calculated as (−W0) / W0.

[0060] As shown in the figure, both the examples and comparative examples exhibit a typical two-stage swelling behavior of "rapid swelling-plateau equilibrium". The rapid swelling period is from 0 to 60 minutes, during which water molecules rapidly enter the three-dimensional network of the hydrogel, causing a sharp increase in the swelling rate. The slow swelling period is from 60 to 120 minutes, during which the swelling rate gradually slows down. After approximately 120 minutes, all groups reach swelling equilibrium, and the equilibrium swelling rate remains relatively stable without a significant upward trend. However, with the increase of Fe3O4 content, the equilibrium swelling rate of the hydrogel shows a significant gradient decrease: Comparative Example 1 has the highest equilibrium swelling rate, approximately 2.6; Example 1 decreases to approximately 2.0; Example 2 further decreases to approximately 1.5; and Comparative Example 2 has the lowest, approximately 1.25. Simultaneously, Comparative Example 1 has the largest error bar, indicating the worst repeatability and batch stability of its swelling behavior. As the Fe3O4 content increases to 0.5%, the error bar decreases significantly, and the swelling stability reaches its optimal level. However, when the Fe3O4 content further increases to 1% (Comparative Example 2), the error bar slightly increases, and the swelling stability decreases.

[0061] Therefore, the equilibrium swelling ratio of the hydrogel is negatively correlated with the cross-linking density of the three-dimensional network: the higher the cross-linking density, the stronger the restriction of the polymer chains, the weaker the ability to absorb water and swell, and the lower the equilibrium swelling ratio. With the addition of appropriate amounts of Fe3O4 (0.25%, 0.5%), the equilibrium swelling ratio decreased significantly in a gradient, proving that Fe3O4 nanoparticles, through hydrogen bonding and electrostatic interactions, bind to the molecular chains of chitosan, sericin, and collagen, acting as additional physical cross-linking points and effectively increasing the cross-linking density of the hydrogel's three-dimensional network. This result confirms at the molecular level that Fe3O4 is not only a contrast agent but also a core finding of this invention as a mechanically reinforcing component, perfectly consistent with the microstructural changes of thickened pore walls and a denser network observed in scanning electron microscopy images. Within the concentration window of 0.2%-0.8% (0.25%, 0.5%) defined in this invention, as the Fe3O4 concentration increases, the swelling ratio decreases uniformly and the error bars gradually decrease, indicating a uniform increase in physical cross-linking points and a more uniform and stable three-dimensional network structure. When the concentration exceeds the upper limit by 1%, although the swelling rate continues to decrease, the swelling stability decreases significantly (error bar increases). This is because at high concentrations, Fe3O4 nanoparticles aggregate, resulting in extremely uneven distribution of crosslinking points: excessive crosslinking in aggregated areas and insufficient crosslinking in non-aggregated areas. This verifies the conclusion of this invention: exceeding the limited concentration range, the aggregation effect of nanoparticles will destroy the uniformity of the hydrogel network, leading to performance degradation, proving that the 0.2%-0.8% concentration window is a necessary condition for achieving synergistic effects of multiple functions. Meanwhile, an excessively high swelling rate (e.g., 2.6 in Comparative Example 1) will cause excessive water absorption and swelling after implantation, putting pressure on surrounding normal cartilage tissue. Excessive swelling will also loosen the network structure, rapidly reducing mechanical strength and failing to provide long-term stable mechanical support. An excessively low swelling rate (e.g., 1.25 in Comparative Example 2) indicates excessive crosslinking, an overly dense network, and excessively low porosity, which will severely hinder the infiltration and proliferation of chondrocytes, as well as the transport of nutrients and metabolic waste. The equilibrium swelling ratio of Example 2 is approximately 1.5, which is within the optimal range for cartilage repair. It also exhibits optimal swelling stability, indicating that its three-dimensional network structure provides sufficient cross-linking density for long-term mechanical support while maintaining suitable porosity and water absorption capacity, thus providing an ideal growth microenvironment for chondrocytes. The small error bars in Examples 1 and 2 indicate a high degree of consistency in their swelling behavior, suggesting that the Fe3O4 nanoparticles are uniformly dispersed in the hydrogel matrix with a uniform distribution of physical cross-linking points. This verifies that the preparation method used in this invention can effectively prevent nanoparticle aggregation. Furthermore, the thermosensitive physical cross-linking system of sodium β-glycerophosphate / sodium bicarbonate can stably lock the uniformly dispersed Fe3O4 within the three-dimensional network, avoiding structural damage and uneven cross-linking problems that may occur with chemical cross-linking systems.

[0062] 3. Mechanical properties

[0063] Figure 3 The compression modulus of hydrogels with different Fe3O4 contents is shown. As can be seen from the figure, the compression modulus exhibits a typical concentration-dependent trend of "increasing first and then decreasing": as the Fe3O4 content increases from 0% to 0.5%, the compression modulus of the hydrogel increases significantly in a stepwise manner; when the Fe3O4 content further increases to 1% (exceeding the upper limit of the concentration window), the compression modulus decreases sharply.

[0064] Therefore, within the concentration window of 0.2%-0.8% defined in this invention, Fe3O4 nanoparticles are not only a contrast agent for MRI tracking, but also a key functional component that significantly improves the mechanical properties of hydrogels. The enhancement mechanism is completely consistent with previous results regarding microstructure and swelling rate: Fe3O4 nanoparticles bind to the molecular chains of chitosan, sericin, and collagen through hydrogen bonds and electrostatic interactions, acting as additional physical cross-linking points. This effectively increases the cross-linking density and structural integrity of the three-dimensional network, thereby achieving a significant increase in compressive modulus. This result verifies the core finding of this invention: "Fe3O4 has a triple synergistic effect of mechanical enhancement, repair promotion, and process tracking." When the Fe3O4 concentration is below 0.2%, sufficient physical cross-linking points cannot be formed, resulting in limited mechanical enhancement. When the concentration is within the 0.2%-0.8% window, the mechanical properties steadily improve with increasing concentration, reaching an optimal level at 0.5%. When the concentration exceeds the upper limit and reaches 1%, the compressive modulus drops sharply, only slightly higher than the blank control group. This is because at high concentrations, Fe3O4 nanoparticles undergo severe aggregation, failing to function as effective physical cross-linking points and instead disrupting the continuity and uniformity of the hydrogel's three-dimensional network, forming stress concentration points and leading to a sharp deterioration in overall mechanical properties. This result, from a mechanical performance perspective, further validates the scientific validity and necessity of limiting the Fe3O4 concentration to 0.2%-0.8% in this invention. The compressive modulus of natural human articular cartilage is typically in the range of 500-1000 kPa, and the compressive modulus of Examples 1 and 2 both fall within this range, precisely within the mechanical performance range of natural articular cartilage. This means that after implantation, the hydrogel perfectly matches the physiological load-bearing requirements of the joint, preventing premature collapse or disintegration due to insufficient mechanical strength, and avoiding wear on surrounding normal cartilage tissue due to excessive hardness, thus providing a long-term stable mechanical support environment for cartilage tissue regeneration. The steady increase in compressive modulus in Examples 1 and 2 indicates that the Fe3O4 nanoparticles are uniformly dispersed in the hydrogel matrix, fully utilizing their role as physical cross-linking points. This indirectly verifies that the preparation method used in this invention can effectively prevent the aggregation of nanoparticles. At the same time, the thermosensitive physical crosslinking system of sodium β-glycerophosphate / sodium bicarbonate can stably lock Fe3O4 in a three-dimensional network without destroying its crystal structure and dispersion state, thus achieving simultaneous optimization of mechanical properties, dispersibility and biosafety.

[0065] 4. Magnetic property detection

[0066] The magnetic properties of the four groups of hydrogel scaffold materials obtained in the comparative examples were tested using a vibrating sample magnetometer (VSM, LakeShore 7400). The results are as follows: Figure 4 As shown.

[0067] As can be seen from the figure, the hysteresis loops of the Fe3O4-containing samples in Examples 1-2 and Comparative Example 2 are all smooth "S"-shaped curves passing through the origin, with no obvious remanence (residual magnetization) or coercivity (the reverse magnetic field required to reduce the magnetization to zero). This indicates that the Fe3O4 nanoparticles in the hydrogel maintain a single-domain superparamagnetic state, exhibiting strong magnetism in the presence of an external magnetic field, and the magnetism disappears immediately after the external magnetic field is removed, without any hysteresis. As the mass fraction of Fe3O4 increases from 0.25% to 1%, the saturation magnetization of the hydrogel (the plateau value of magnetization approaching under high field) shows an approximately linear gradient increase: the specific magnetization of Comparative Example 1 is almost 0, proving that the chitosan-sericulture-collagen matrix itself is non-magnetic, and all magnetic signals come from the Fe3O4 nanoparticles; the saturation magnetization of Example 1 is approximately 0.8 emu / g, the saturation magnetization of Example 2 is approximately 1.4 emu / g, and the saturation magnetization of Comparative Example 2 is approximately 1.6 emu / g. It is worth noting that even when the Fe3O4 content reached 1% (exceeding the concentration window of 0.2%-0.8% specified in this invention), the hysteresis loop of Comparative Example 2 still did not show ferromagnetic characteristics (no remanence and coercivity), indicating that the crystalline structure of Fe3O4 nanoparticles was not destroyed, and the particle size remained below the superparamagnetic critical size (Fe3O4 is about 20nm).

[0068] Therefore, within the concentration window of 0.2%-0.8% defined in this invention, the hydrogel exhibits clear and controllable magnetic properties. The saturation magnetization increases linearly with the Fe3O4 concentration, indicating that ideal MRI imaging contrast can be obtained by precisely adjusting the amount of Fe3O4 added. The saturation magnetization in Examples 1 and 2 is within the effective signal range for MRI imaging, capable of generating sufficiently clear and distinguishable in vivo tracer signals, meeting the clinical needs for non-invasive continuous dynamic monitoring of hydrogel implantation location, degradation process, and cartilage repair status. In this experiment, all Fe3O4-containing hydrogels maintained complete superparamagnetic characteristics, and the saturation magnetization was linearly positively correlated with the Fe3O4 content, directly proving that the β-glycerophosphate / sodium bicarbonate thermosensitive physical crosslinking system used in this invention, while achieving in-situ gelation at physiological body temperature, does not destroy the spinel crystal structure of Fe3O4 nanoparticles, thus fully preserving their magnetic signal activity. Meanwhile, superparamagnetism is a unique property of single-domain nanoparticles. Once the nanoparticles undergo severe aggregation and the particle size exceeds the superparamagnetic critical size, they will transform into ferromagnetism, exhibiting significant remanence and coercivity. In this invention, even with an Fe3O4 content of 1%, the hydrogel still maintains superparamagnetism, indicating that the preparation method described in this invention can effectively prevent the aggregation of Fe3O4 nanoparticles through steric hindrance, maintaining the monodisperse state of the particles even at high concentrations. The flexible coating layer formed by PEG on the surface of the nanoparticles not only improves dispersibility but also avoids direct interaction between the nanoparticles and the crosslinking agent, further protecting their crystalline structure and magnetic properties. More importantly, the maintenance of superparamagnetism is crucial for the biosafety of implantable materials. Because superparamagnetic nanoparticles are non-magnetic in the absence of an external magnetic field, they will not attract each other and aggregate to form large particles, thus avoiding the risk of vascular blockage or local inflammatory reactions. The magnetism disappears immediately after the external magnetic field is removed, without causing long-term magnetic interference to surrounding normal tissues. This result proves that the traceable hydrogel of this invention has good in vivo biosafety and meets the safety requirements for clinical implantable materials. Although Comparative Example 2 exhibited the highest saturation magnetization and theoretically the strongest tracer signal, combined with previous scanning electron microscopy, swelling performance, and compressive modulus results, it was found that a 1% Fe3O4 concentration already led to severe degradation of the hydrogel's three-dimensional network structure and a precipitous drop in mechanical properties. This further demonstrates the scientific validity of limiting the Fe3O4 concentration to 0.2%-0.8% in this invention: within this concentration window, the hydrogel can simultaneously achieve excellent mechanical support properties, good cartilage regeneration promotion, and clear MRI tracer function, achieving an optimal balance of triple synergistic effects.

[0069] 5. Biosafety

[0070] (1) Subcutaneous implantation

[0071] Figure 5The image shows the Prussian blue staining results of the hydrogel subcutaneous implants prepared in the comparative and example cases. Using an SD rat subcutaneous implantation model, the hydrogel scaffold materials of each group were implanted subcutaneously into the back of the rats. Skin tissue surrounding the implantation area was removed at 1, 2, 3, and 4 weeks post-surgery, and after fixation, embedding, sectioning, and Prussian blue staining, the tissue was stained.

[0072] As shown in the figure, no blue signal was observed in Comparative Example 1 at any time point, proving that the chitosan-sericin-collagen hydrogel matrix itself does not contain iron. All detected blue signals originated from the implanted Fe3O4 nanoparticles, excluding interference from endogenous iron. In the first week, all examples showed obvious blue areas, indicating successful material implantation and distribution of magnetic particles within them, with Comparative Example 2 showing the deepest and most concentrated blue. From the second to the fourth week, the blue signal in Example 1 gradually faded over time, and by the fourth week, the blue area was quite blurred, indicating that the magnetic particles may have been lost relatively quickly due to hydrogel degradation or diffusion. In the fourth week, Example 2 still maintained a relatively clear and stable blue signal area. This indicates that the magnetic particles at this concentration are uniformly dispersed and well retained in the hydrogel, the hydrogel structure is stable, and it can maintain its tracking ability for a relatively long time. Although Comparative Example 2 had the strongest initial signal, the blue area appeared more diffuse or aggregated in the later stages (third and fourth weeks), and the surrounding tissue reaction may have been more severe (morphological changes in the pink area). High concentrations of particles may cause local aggregation or inflammatory reactions, affecting the uniform degradation of the material and the accuracy of tracking.

[0073] These Prussian blue staining images demonstrate the in vivo retention and degradation behavior of the hydrogel in SD rats over a period of up to 4 weeks after subcutaneous implantation. The blue regions specifically indicate the distribution and retention of Fe3O4 magnetic nanoparticles. This result visually demonstrates the excellent in vivo tracking function of the hydrogel of this invention, successfully achieving in vivo visual monitoring of the material by introducing magnetic particles. Specifically, Examples 1, 2, and Comparative Example 2 all exhibited significant blue signals at each time point, and the signal intensity was positively correlated with particle concentration, verifying the effectiveness of the tracer. In particular, Example 2 maintained a clear and relatively concentrated blue region even in the fourth week, indicating that the magnetic particles at this concentration were not only uniformly dispersed but also that the hydrogel matrix structure was stable, enabling long-term and stable tracking, effectively matching the cartilage repair cycle. In contrast, the signal in Example 1 decayed rapidly over time, while Comparative Example 2, although initially showing a strong signal, may have become diffused later due to particle aggregation or local inflammation. This further corroborates the necessity of limiting the magnetic material concentration window (0.2%~0.8%) in this invention to achieve the optimal balance between tracking performance and biocompatibility. Meanwhile, the critical repair period for cartilage tissue regeneration is typically 4-8 weeks post-surgery. In Example 2 of this invention, the tracer signal remained stable for at least 4 weeks, remaining clearly identifiable and uniformly distributed throughout the entire period, providing continuous dynamic monitoring information on the implant material's location, degradation process, and cartilage repair status for clinical use. In contrast, the signal in Example 1 disappeared too early, failing to cover the complete critical repair period; while the signals in Comparative Example 2 lasted longer, they posed serious safety risks and could not meet the actual needs of clinical tracer monitoring. The uniform distribution of the early blue signals in Examples 1 and 2 directly demonstrates that the preparation method described in this invention can effectively prevent the secondary aggregation of Fe3O4 nanoparticles in vivo; the stable signal anchoring proves that the three-dimensional network structure formed by the thermosensitive physical crosslinking system is uniform and stable, enabling the gradual release of nanoparticles during in vivo degradation, avoiding the risk of nanoparticle burst release caused by uneven local crosslinking and sudden network collapse that may occur with chemical crosslinking systems.

[0074] (2) Implantation at cartilage defect sites

[0075] Figure 6 Prussian blue staining results of hydrogel scaffold materials prepared in comparative and example studies implanted into cartilage defect sites in rats. An in situ cartilage defect model of the knee joint was used in SD rats. Cylindrical bone defects approximately 2 mm in diameter and 1.5 mm deep were created in the femoral condyle, and hydrogel scaffold materials from each group were implanted. Animals were sacrificed 4 weeks post-surgery, and tissue from the defect area was removed, fixed, decalcified, embedded, sectioned, and then stained with Prussian blue.

[0076] The Prussian blue staining results of the in situ model of knee cartilage defect show that no blue signal was observed in the Blank blank defect group and the Control blank hydrogel group at weeks 2 and 4. This proves that the normal cartilage, bone tissue and chitosan-sericin-collagen hydrogel matrix itself do not contain iron elements that can be stained. All detected blue signals came from the implanted Fe3O4 nanoparticles, excluding the interference of endogenous iron and the material matrix. The anchoring effect of Fe3O4 under dynamic joint loading showed a strict concentration dependence: In the second week, both Example 1 and Comparative Example 2 showed obvious blue areas, indicating that the material was successfully implanted and the magnetic particles were distributed within it, with Comparative Example 2 showing the deepest and most concentrated blue. In the fourth week, the blue signal of Example 1 gradually faded over time, and by the fourth week, the blue area had become quite blurred, indicating that the magnetic particles may have been lost relatively quickly due to hydrogel degradation or diffusion. In the fourth week, Example 2 still maintained a relatively clear and stable blue signal area, indicating that the magnetic particles at this concentration were uniformly dispersed and well preserved in the hydrogel, and the hydrogel structure was stable, maintaining a long-term tracking ability. Although Comparative Example 2 had the strongest initial signal, the blue area appeared more diffuse or aggregated in the later period (fourth week), and the surrounding tissue reaction may have been more severe (morphological changes in the pink area), suggesting that high concentrations of particles may cause local aggregation or inflammatory reactions, affecting the uniform degradation of the material and the accuracy of tracking.

[0077] Unlike the static environment of subcutaneous implantation, under the dynamic shear force of repeated joint loading, the signal decay rate of each group showed a pattern highly consistent with the mechanical performance: Example 2, with the best mechanical performance, had the longest signal duration (≥4 weeks); Example 1, with the second best mechanical performance, completely disappeared at 4 weeks; while the comparative examples 2, with a sharp drop in mechanical performance, showed a significant signal decay at 2 weeks and completely disappeared at 4 weeks, much faster than the metabolic rate under the subcutaneous implantation environment. Throughout the observation period, the blue signals of all Fe3O4-containing groups appeared only in the cartilage defect filling area. No scattered blue particles were detected in the surrounding normal cartilage tissue, subchondral bone, synovial tissue, or joint cavity, proving that Fe3O4 nanoparticles at the specified concentration would not migrate ectopically due to synovial fluid erosion or kinetic shear force. The results of this in-situ cartilage defect model provide the most clinically valuable validation of the technical efficacy of this invention. Example 2, even up to 4 weeks post-surgery, still provided a clearly identifiable tracer signal in the cartilage defect area, and the signal remained strictly confined to the implantation site. This demonstrates that the hydrogel system constructed in this invention can maintain the integrity of its three-dimensional network structure under the complex physiological environment of repeated joint flexion and extension and weight-bearing, effectively locking Fe3O4 nanoparticles and preventing early leakage due to synovial fluid erosion and shear forces. This provides a solid experimental foundation for non-invasive, continuous, and dynamic monitoring of the cartilage repair process in clinical practice, solving the technical problems of signal loss and inaccurate tracking in existing traceable hydrogels under dynamic joint conditions. Comparative Example 2 showed a significant signal attenuation in week 2 and complete disappearance in week 4, profoundly revealing the dangers of exceeding the concentration window from an in-situ perspective: high concentrations of Fe3O4 cause nanoparticle aggregation, disrupting the continuity and mechanical strength of the hydrogel's three-dimensional network. Under dynamic joint loading, the network structure prematurely disintegrates, failing to effectively lock in the nanoparticles, leading to their rapid loss and metabolism. This not only fails to achieve long-term tracking but may also result in the loss of mechanical support for the defect site due to premature material disintegration. This result is more convincing than in vitro mechanical experiments and subcutaneous implantation experiments, demonstrating that a concentration window of 0.2%-0.8% is a necessary condition for the hydrogel of this invention to simultaneously achieve the triple functions of mechanical support, cartilage repair, and stable tracking in a joint environment. The critical period for cartilage defect repair in rats is 4-6 weeks post-surgery. The tracking signal of Example 2 in this invention can be stably sustained for at least 4 weeks, precisely covering the most critical stage of cartilage regeneration, providing clinicians with real-time information on the location, degradation process, and defect filling status of the implanted material. In Example 1, the signal disappeared too early, making it impossible to monitor material behavior during the repair process; in Comparative Example 2, the signal was unstable and lost too early, neither of which could meet the actual tracing requirements for clinical cartilage repair. Meanwhile, the fact that Example 2 maintained a stable tracing signal within 4 weeks indirectly proves that its three-dimensional network structure can withstand the physiological load and repeated shear forces of the joint, without early collapse or disintegration.This is completely consistent with the previous compression modulus test results, further verifying the effectiveness of the design concept of improving the mechanical properties of hydrogels by using Fe3O4 as a physical cross-linking point in the real in vivo environment.

[0078] (3) In vitro cell proliferation

[0079] The four groups of hydrogel scaffold materials obtained in the comparative and example cases were subjected to in vitro cell proliferation assays, and the results are as follows: Figure 7 As shown. The scaffold material was added to complete culture medium at a mass / volume ratio of 0.1 g / mL for extraction (37℃, 24 h) to prepare 100% and 50% concentration extracts. The absorbance at ODλ = 490 nm-1 of L-929 cells cultured in wells containing different concentrations of extract for 1, 2, and 3 days was measured using the MTT assay. The relative cell proliferation rate (RGR = ODexperimental group / ODcontrol group) was calculated as 100%. Figure 7 The relative proliferation rate and toxicity score of L-929 cells were measured over 3 days in well plates containing different concentrations of scaffold material extract.

[0080] As shown in the figure, the relative cell proliferation rate of Comparative Example 1 was not statistically different from that of the Blank control group at all time points (Day 1: 102%, Day 2: 99%, Day 3: 109%), demonstrating that the chitosan-sericin-collagen three-dimensional network matrix and the thermosensitive physical cross-linking system of sodium β-glycerophosphate / sodium bicarbonate were completely non-cytotoxic. In Example 1, the relative cell proliferation rate remained stable between 93% and 98% within 3 days (according to ISO 10993-5 Medical Device Biological Evaluation Standard, toxicity score 0, no toxicity). In Example 2, the relative cell proliferation rate remained stable between 88% and 90% within 3 days, both exceeding 80% (toxicity score 1, very slight toxicity, no clinical significance). The slight differences between these two groups and the Blank group essentially disappeared by Day 3, indicating that the cells had good adaptability to low concentrations of Fe3O4. In contrast, the relative cell proliferation rate of Comparative Example 2 was significantly lower than that of the other groups at all time points (p<0.05), remaining stable between 74% and 78%, with a toxicity score of grade 2 (mild toxicity). This toxicity did not show a significant trend of abating with prolonged culture time. From day 1 to day 3, the relative cell proliferation rate of each group remained stable or slightly increased, without showing a continuous decline over time, demonstrating that the hydrogel extract did not contain any harmful substances that would lead to the accumulation of cytotoxicity.

[0081] Therefore, the cell proliferation rate of Comparative Example 1 is completely consistent with that of the Blank group, proving that the β-glycerophosphate sodium / sodium bicarbonate physical crosslinking system used in this invention completely solves the cytotoxicity problem caused by the residue of traditional chemical crosslinking agents (such as glutaraldehyde and genipin). Physical crosslinking achieves gelation only through hydrogen bonding and electrostatic interactions, without introducing any toxic chemicals or destroying the bioactivity of natural polymers. This is the fundamental reason for the excellent biocompatibility of the hydrogel of this invention. Cytotoxicity exhibits a concentration-dependent effect: within the defined window of 0.2%-0.8%, Fe3O4 shows no significant cytotoxicity; once the window is exceeded to 1%, statistically significant mild toxicity immediately appears. This result explains the potential inflammatory risk that Comparative Example 2 may have in in vivo experiments at the molecular and cellular level, further proving that limiting the Fe3O4 concentration to 0.2%-0.8% in this invention is not only necessary for achieving mechanical enhancement and stable tracking, but also an important condition for ensuring the biosafety of the material. Even at a high concentration of 1%, the toxicity of Comparative Example 2 is only mild, far lower than the cytotoxicity level of Fe3O4 nanoparticles without PEG coating in the prior art. This demonstrates that the preparation method described in this invention not only solves the problem of nanoparticle dispersion, but also that the flexible coating layer formed by PEG on the Fe3O4 surface effectively blocks direct contact between nanoparticles and cell membranes, reducing oxidative stress damage and significantly lowering the inherent cytotoxicity of magnetic nanoparticles. The low toxicity observed in the limited concentration group in in vitro experiments is completely consistent with the results observed in in vivo subcutaneous implantation and in situ joint experiments, which showed no significant inflammatory response, no tissue necrosis, and no ectopic deposition, forming a complete chain of evidence for biosafety from the in vitro cellular level to the in vivo tissue level. This indicates that the hydrogel system of this invention will not cause serious local or systemic toxicity reactions in clinical applications and has extremely high clinical translational safety.

[0082] (4) Blood compatibility

[0083] The blood compatibility of the four groups of hydrogel scaffold materials obtained from the comparative and example cases was evaluated, and the results are as follows: Figure 8 As shown in the figure. Fresh anticoagulated rabbit blood was diluted to 40% (v / v) with physiological saline. 10 mL of physiological saline (experimental group), 10 mL of physiological saline (negative control group), and 10 mL of deionized water (positive control group) were added to each group's hydrogel sample. After incubation at 37℃ for 30 min, 250 μL of diluted blood was added. After incubation at 37℃ for 1 h, 200 μL of supernatant was centrifuged (1000 r / min, 10 min), and the absorbance of the supernatant was measured at 540 nm to calculate the hemolysis rate.

[0084] As shown in the figure, the static hemolysis test using fresh anticoagulated rabbit blood met the ISO 10993-4 standard for biological evaluation of medical devices. *** represents p < 0.001 (extremely significant statistical difference), and * represents p < 0.05 (significant statistical difference). The hemolysis rates of all four hydrogel groups were below 1.5%, far below the 5% safety threshold specified by the ISO standard, demonstrating that the hydrogel system of this invention has excellent blood compatibility and does not cause significant red blood cell rupture or hemolysis. The blank hydrogel (Comparative Example 1) had the highest hemolysis rate, approximately 1.4%. After the addition of Fe3O4, the hemolysis rates of all three composite hydrogel groups decreased significantly (p < 0.001), stabilizing between 0.5% and 0.6%, a reduction of approximately 60% compared to Comparative Example 1. This is completely contrary to the initial prediction of this invention, as it is generally believed in the art that nanoparticles may damage the red blood cell membrane, leading to an increased hemolysis rate. However, the addition of Fe3O4 in this invention significantly improved the blood compatibility of the hydrogel. It is noteworthy that the hemolysis rate remained extremely low throughout the entire range of Fe3O4 concentration increasing from 0.25% to 1%, with only minor statistical differences among the three groups (p<0.05) and no obvious concentration-dependent increasing trend. Even in Comparative Example 2, which was outside the concentration window, the hemolysis rate was essentially the same as that of the two example groups, and no deterioration in blood compatibility was observed.

[0085] Therefore, the hemolysis rates in all groups were far below the safety standard of 5%, proving that the hydrogel of this invention does not cause severe hemolytic reactions upon contact with blood. This is particularly important for articular cartilage repair materials, as the material inevitably comes into contact with blood during surgical implantation, and good blood compatibility is a necessary condition to avoid postoperative thrombosis, local inflammatory reactions, and tissue necrosis. The blank hydrogel (Comparative Example 1) had a large number of free chitosan amino groups on its surface. These positively charged amino groups interact electrostatically with the negatively charged erythrocyte membrane, leading to erythrocyte deformation and rupture, thus causing a certain hemolysis rate. However, when Fe3O4 nanoparticles were added, the nanoparticles combined with the chitosan amino groups through electrostatic interactions, acting as physical cross-linking points, significantly reducing the number of free positive charges on the hydrogel surface. Simultaneously, the addition of nanoparticles made the hydrogel surface smoother and the structure more uniform, reducing mechanical damage to erythrocytes—a truly unexpected improvement. Even at a high concentration of 1%, Fe3O4 did not increase the hemolysis rate, directly proving the effectiveness of the preparation method described in this invention. The hydrophilic, flexible coating layer formed by PEG on the surface of Fe3O4 nanoparticles effectively blocks direct contact between the nanoparticles and the erythrocyte membrane, avoiding mechanical puncture damage to the cell membrane caused by the sharp edges of the nanoparticles. It also reduces non-specific protein adsorption, fundamentally eliminating the risk of hemolysis that may be caused by magnetic nanoparticles. In this experiment, the hemolysis rate in all groups was extremely low, indirectly proving that the β-glycerophosphate / sodium bicarbonate physical crosslinking system does not introduce any blood-toxic chemicals. Compared with traditional chemical crosslinking agents (such as glutaraldehyde), which leave toxic monomers and severely damage blood compatibility, the physical crosslinking system of this invention achieves gelation only through hydrogen bonds and electrostatic interactions, completely preserving the good blood compatibility of the natural polymer. These hemolysis test results, together with previous MTT cytotoxicity experiments (toxicity grade 0-1 in the limited concentration group), subcutaneous implantation experiments (no obvious inflammatory response), and in-situ joint experiments (no ectopic deposition), constitute a complete chain of evidence for biosafety from the in vitro cellular level, blood level, to the in vivo tissue level. This comprehensively demonstrates the safety of the hydrogel system of this invention in clinical applications, laying a solid foundation for its further clinical translation.

[0086] (5) H&E staining results after subcutaneous implantation

[0087] Figure 9 The images show the H&E staining results of the hydrogel subcutaneous implants prepared in the comparative and example cases. Using an SD rat subcutaneous implantation model, the hydrogel scaffold materials of each group were implanted subcutaneously into the back of the rats. Skin tissue surrounding the implantation area was removed at weeks 1, 2, 3, and 4 post-surgery. After washing with PBS, the tissue was fixed in 4% paraformaldehyde fixative for 24 hours. Following routine tissue embedding and sectioning, hematoxylin-eosin (H&E) staining was performed, and histopathological observation was conducted under a microscope.

[0088] As shown in the figure, in the first week after surgery, different degrees of inflammatory response were observed in the subcutaneous implantation areas of each group: In the Control group (blank hydrogel), a small number of scattered inflammatory cells (dark purple cell nuclei) were infiltrated around the implantation area, and the interstitial spaces were slightly widened; the degree of inflammatory cell infiltration in the HG-0.25 group was similar to that in the Control group, and a small amount of fibrous tissue hyperplasia was observed locally; the HG-0.5 group had the least inflammatory cell infiltration, the interstitial spaces were basically normal, and only a small number of inflammatory cells were occasionally observed; the HG-1 group showed a large number of dense inflammatory cell aggregations, significantly widened interstitial spaces, and inflammatory exudate was observed locally, with a significantly stronger inflammatory response than other groups. In the second week post-surgery, the inflammatory response showed a differentiated trend: the number of inflammatory cells in the Control group decreased slightly, but scattered infiltration was still visible; the number of inflammatory cells in the HG-0.25 group decreased further, and fibrous tissue hyperplasia began to be obvious; the number of inflammatory cells in the HG-0.5 group basically subsided, and new capillaries and fibroblasts were visible around the implantation area, and the tissue repair process accelerated; the number of inflammatory cells in the HG-1 group was still large, and local tissue necrosis (nuclear pyknosis and dissolution) occurred, fibrous tissue hyperplasia was disordered, and the repair process was hindered. At week 3 post-surgery, the repair status of each group showed significant differences: In the Control group, inflammatory cells continued to decrease and fibrous tissue gradually matured, but the boundary between the implantation area and the surrounding tissue was still relatively clear; In the HG-0.25 group, inflammation basically subsided, the fibrous tissue arrangement tended to be regular, new blood vessels increased, and it integrated well with the surrounding tissue; In the HG-0.5 group, inflammation completely subsided, the implantation area was replaced by mature fibrous tissue, the structure was dense and there was no obvious boundary with the surrounding tissue, and the repair effect was the best; In the HG-1 group, a lot of inflammatory cell infiltration was still visible, the fibrous tissue was excessively proliferated and arranged in a disordered manner, there were still necrotic areas in some places, and the tissue repair was seriously delayed. At the 4th week post-surgery, the final repair status was clear: in the Control group, the fibrous tissue in the implantation area was basically mature, but a small number of inflammatory cells remained; in the HG-0.25 group, the fibrous tissue was highly mature, with a regular structure and close integration with the surrounding tissue; in the HG-0.5 group, the implantation area was completely replaced by normal fibrous tissue, with no inflammatory cells remaining, and the tissue structure was no different from the surrounding normal tissue; in the HG-1 group, inflammatory cells persisted, excessive proliferation of fibrous tissue formed a scar-like structure, the local necrotic area was not completely repaired, and the biocompatibility was the worst.

[0089] Therefore, the hydrogel scaffold of Example 2 (HG-0.5 group) of this invention exhibits optimal in vivo biocompatibility: the inflammatory response is mildest in the early stage (weeks 1-2), enabling rapid initiation of tissue repair; inflammation completely subsides in the middle stage (week 3), fibrous tissue matures and integrates well with surrounding tissues; and there is no residual inflammation in the late stage (week 4), with the implantation area completely replaced by normal tissue. This indicates that the 0.5% Fe3O4 composite hydrogel does not induce persistent or severe immune rejection in vivo, and its safety is significantly superior to that of high-concentration magnetic materials (HG-1 group). Meanwhile, although the HG-0.25 group showed good biocompatibility, its repair speed and maturity were slightly inferior to the HG-0.5 group. The HG-1 group, due to its excessively high Fe3O4 concentration, experienced persistent chronic inflammation and tissue necrosis. This demonstrates that the 0.2%-0.8% Fe3O4 concentration window is the key range for achieving good biocompatibility; exceeding this range severely compromises the in vivo safety of the material, providing important histological safety evidence for its clinical application. This result is completely consistent with the in vitro MTT cytotoxicity assay (toxicity grade 0-1) and hemolysis rate assay (hemolysis rate <0.6%), validating the clinical safety of the material at the in vivo tissue level. The two comparative groups outside the concentration window exhibited persistent chronic inflammatory response, disordered degradation behavior, and extremely poor tissue integration. This is because at high concentrations, Fe3O4 nanoparticles severely aggregated, disrupting the uniformity of the hydrogel's three-dimensional network and leading to irregular material degradation. Furthermore, the aggregated nanoparticles continuously released and stimulated surrounding tissues, triggering a chronic inflammatory response. This result further demonstrates from an in vivo histopathological perspective that a concentration window of 0.2%-0.8% is a necessary condition for achieving the synergistic effect of hydrogel biocompatibility, mechanical stability, and controllable degradation. The regeneration cycle of cartilage tissue is typically 4-8 weeks. Ideally, cartilage repair materials should maintain structural integrity and provide mechanical support in the early stages of repair, gradually degrading and being replaced by newly formed tissue in the later stages. In this invention, the degradation rate of Example 2 precisely meets this requirement: it maintains an intact three-dimensional structure providing mechanical support for the first two weeks, and begins to degrade slowly and uniformly from the third to fourth weeks, while allowing for the orderly ingrowth of new tissue. In contrast, Comparative Example 1 degrades too quickly to provide long-term support, and Comparative Example 2 degrades too slowly, hindering tissue regeneration; neither can meet the needs of cartilage repair. Example 2 still maintains an intact three-dimensional structure in the second week, while Comparative Example 1 shows significant disintegration in the second week. This difference directly proves that an appropriate amount of Fe3O4 as a physical crosslinking point can significantly improve the structural stability of the hydrogel three-dimensional network, slow down the degradation rate of the material, and enable it to provide mechanical support in vivo for a longer period of time. This is completely consistent with the previous results of the compression modulus test (Example 2 has the highest compression modulus) and the swelling rate test (Example 2 has a moderate swelling rate).

[0090] (6) HE staining results after implantation at the cartilage defect site

[0091] Figure 10 HE staining results were obtained for the hydrogel scaffold materials prepared in the comparative and example cases after implantation into the cartilage defect sites of rats. An in situ cartilage defect model of the knee joint of SD rats was used. Hydrogel scaffold materials from each group were implanted. Four weeks post-operation, the animals were sacrificed, and tissue from the defect area was removed. After fixation, decalcification, paraffin embedding, sectioning, and H&E staining, the tissue was stained.

[0092] from Figure 10 As can be seen, the H&E staining results at the second week post-surgery (early repair stage) showed significant differentiation in the degree of cartilage defect filling and the cartilage matrix secretion capacity among the groups: the Blank defect group showed obvious depression in the defect area, with only a thin layer of fibrin clots covering the surface, and the subchondral bone was completely exposed; the Control hydrogel group showed only a small amount of loose fibrous tissue filling in the defect area, with a sharp and clear boundary with the surrounding normal cartilage; the HG-0.25 group showed significant improvement in defect filling and increased thickness of new tissue; the HG-0.5 group showed the best early repair effect, with the defect almost completely filled, the thickness of new tissue approaching that of the surrounding normal cartilage, and the boundary with the surrounding normal cartilage was no longer obvious; the HG-1 group, which exceeded the concentration window, had poor defect filling effect, with thin and unevenly distributed new tissue, and a small amount of scattered inflammatory cell infiltration was visible. H&E staining results at week 4 post-surgery (mid-repair stage) showed a more significant differentiation in cartilage regeneration quality and interface integration among the groups: the Blank group still exhibited a marked depression, filled only by a large amount of messy fibrous scar tissue; the Control group showed increased defect filling, but the newly formed tissue was still mainly fibrous scar tissue, with a clear and sharp boundary with the surrounding normal cartilage; the HG-0.25 group showed further thickening of the newly formed tissue and improved integration with the surrounding cartilage, but some boundaries still existed, and the newly formed cartilage... The bone thickness was slightly thinner than normal cartilage; the HG-0.5 group achieved near-perfect cartilage repair, with the defect completely filled, the surface of the newly formed cartilage smooth and flat, and the thickness completely consistent with the surrounding normal cartilage. Moreover, the newly formed cartilage formed seamless integration with the surrounding normal cartilage and the subchondral bone below, with the boundary almost completely disappearing; the repair effect of the HG-1 group was significantly deteriorated, the defect was still not completely filled, the newly formed tissue was mainly dense fibrous scar, with clear boundaries with the surrounding cartilage and obvious gaps, and continuous chronic inflammatory cell infiltration and undegraded material residue were visible.

[0093] Therefore, this invention, for the first time, directly demonstrates the practical effectiveness of its triple synergistic effect of "mechanical enhancement, repair promotion, and process tracking" under the real physiological environment of dynamic joint load. Fe3O4 is not only a mechanical enhancer and tracer, but also a bioactive factor that significantly promotes cartilage-specific regeneration, achieving a qualitative leap from "fibrous repair" to "hyaluronic cartilage regeneration." Furthermore, the repair effect of the HG-1 group was even worse than that of the blank hydrogel group, proving that the 0.2%-0.8% concentration window is the only effective range for achieving the triple synergistic effect. Exceeding this range not only fails to achieve the expected results but also causes persistent chronic inflammation due to nanoparticle aggregation, severely inhibiting cartilage regeneration. In addition, this result also proves that mechanical enhancement is a prerequisite for achieving high-quality cartilage repair. The degradation rate of the hydrogel in this invention is perfectly matched with the cartilage regeneration cycle, and it exhibits excellent biocompatibility under the complex physiological environment of the joint, providing comprehensive effectiveness and safety assurance for its clinical translation and application.

[0094] (7) Safranin-Fix-Green staining results after implantation at the cartilage defect site

[0095] Figure 11 This image shows the results of Safranin-Fix-Green staining of hydrogel scaffold materials prepared in comparative and example studies implanted into cartilage defects in rats. An orthotopic cartilage defect model of the knee joint in SD rats was used. Hydrogel scaffold materials from each group were implanted. Four weeks post-surgery, animals were sacrificed, and tissue from the defect area was removed. After fixation, decalcification, paraffin embedding, and sectioning (7 μm thickness), Safranin-Fix-Green staining was performed. Red or purplish-red staining represents the expression of cartilage-specific matrix (glycosaminoglycans, GAGs).

[0096] As shown in the figure, in the second week after surgery (early repair stage), Safranin-Fix-Green staining revealed a differentiation in the secretion capacity of cartilage-specific matrix (glycosaminoglycans, GAGs) in each group: the Blank blank defect group showed no red GAG staining in the defect area, with only a small amount of fibrin; the Control blank hydrogel group also showed no obvious GAG deposition in the defect area, with a clear boundary with the surrounding normal cartilage (red); the HG-0.25 group showed scattered light red GAG staining in the defect area, indicating the initial secretion of cartilage matrix; the HG-0.5 group showed large areas of continuous red GAG staining in the defect area, with a range close to the surrounding normal cartilage, indicating a large amount of cartilage matrix secretion; the HG-1 group showed only a very small amount of scattered red staining in the defect area, with significantly insufficient GAG deposition. By the fourth week post-surgery (mid-repair phase), the differences in GAG expression among the groups became more significant: in the Blank group, the defect area still showed no GAG staining and was filled with fibrous scar tissue; in the Control group, a small amount of red GAG appeared at the edge of the defect area, but the main body was still unstained fibrous tissue; in the HG-0.25 group, the GAG ​​staining area of ​​the defect area expanded and the color deepened, but the thickness was slightly thinner than the surrounding normal cartilage; in the HG-0.5 group, the defect area was completely filled with a uniform dark red GAG matrix, the staining intensity was consistent with the surrounding normal cartilage, and the surface was smooth; in the HG-1 group, the GAG ​​staining of the defect area was still very weak, mainly consisting of unstained dense fibrous tissue, with clear boundaries and gaps with the surrounding normal cartilage.

[0097] Therefore, the hydrogel scaffold of Example 2 (HG-0.5 group) of the present invention can significantly promote the secretion and deposition of cartilage-specific matrix (GAG), achieving matrix regeneration close to that of natural cartilage. Moreover, the GAG ​​distribution in the repaired tissue is uniform and the strength is consistent with that of normal cartilage, indicating that it can effectively induce chondrocytes to secrete functional matrix and promote hyaline cartilage regeneration. At the same time, although the HG-0.25 group has a certain amount of GAG secretion, the repair quality is inferior to that of the HG-0.5 group. In the HG-1 group, GAG secretion is severely suppressed due to the high Fe3O4 concentration, proving that the Fe3O4 concentration window of 0.2%-0.8% is the key range for achieving efficient regeneration of cartilage matrix. Exceeding this range will inhibit matrix synthesis. In addition, the low GAG expression in the Control group and Blank group further highlights the technical advantages of the hydrogel of the present invention (especially HG-0.5) in promoting the regeneration of cartilage-specific matrix, providing functional matrix support for articular cartilage repair.

[0098] (8) Immunological results after implantation at cartilage defect sites

[0099] Figure 12This section presents the immunohistochemical results of type II collagen implanted into cartilage defect sites in rats using hydrogel scaffold materials prepared in Comparative Examples 1-3 and Example 1. An orthotopic cartilage defect model of the knee joint in SD rats was used. Hydrogel scaffold materials from each group were implanted. Four weeks post-operation, animals were sacrificed, and tissue from the defect area was harvested. After fixation, decalcification, paraffin embedding, and sectioning (5 μm thickness), immunohistochemical staining was performed. The main steps included antigen thermal retrieval, endogenous peroxidase blockade, primary antibody incubation, enhancement solution treatment, enzyme-labeled secondary antibody incubation, DAB staining, and counterstaining.

[0100] As shown in the figure, in the second week after surgery (early repair stage), immunohistochemical staining of type II collagen revealed significant differences in cartilage-specific collagen secretion among the groups: the defect areas in the Blank blank defect group and the Control blank hydrogel group showed almost no brownish-yellow type II collagen positive staining, with only the surrounding normal cartilage showing positive staining; the defect area in the HG-0.25 group began to show scattered light brownish-yellow type II collagen staining, indicating the initial synthesis of cartilage collagen; the defect area in the HG-0.5 group showed large areas of continuous brownish-yellow type II collagen staining, with an area close to the surrounding normal cartilage, indicating a large amount of type II collagen secretion; the defect area in the HG-1 group showed only a very small amount of scattered brownish-yellow staining, indicating a significant deficiency in type II collagen deposition. By the fourth week post-surgery (mid-repair phase), the differences in type II collagen expression among the groups became more significant: The Blank and Control groups showed no obvious positive staining for type II collagen in the defect area, with fibrous tissue repair being the primary feature; the HG-0.25 group showed an expanded and darker range of type II collagen staining in the defect area, but the staining intensity and continuity were still weaker than normal cartilage; the HG-0.5 group showed the defect area completely filled with uniform, dark brownish-yellow type II collagen, with staining intensity and distribution continuity highly consistent with the surrounding normal cartilage, indicating that the newly formed tissue was rich in mature type II collagen matrix; the HG-1 group showed extremely weak type II collagen staining in the defect area, with a clear boundary with the surrounding normal cartilage, suggesting that high concentrations of Fe3O4 severely inhibited the synthesis and deposition of type II collagen.

[0101] Therefore, the hydrogel scaffold of Example 2 (HG-0.5 group) of the present invention can significantly promote the synthesis and deposition of cartilage-specific type II collagen, achieving collagen matrix regeneration consistent with natural cartilage. This indicates that it can effectively induce chondrocyte differentiation and functional collagen secretion, which is a key marker of high-quality hyaline cartilage regeneration. Meanwhile, although the Example 1 (HG-0.25) group had some type II collagen secretion, its repair quality was inferior to that of Example 2 (HG-0.5 group). In contrast, the type II collagen synthesis of the Comparative Example 2 group was severely inhibited because the Fe3O4 concentration exceeded the effective window of 0.2%-0.8%, proving that this concentration range is a necessary condition for achieving efficient cartilage collagen regeneration. In addition, the low type II collagen expression in the Control group and Blank group further highlights the technical advantages of the hydrogel of the present invention (especially HG-0.5) in promoting cartilage-specific collagen regeneration, providing structurally complete and functionally mature collagen matrix support for cartilage repair, and verifying the core efficacy of the present invention in promoting repair.

[0102] Finally, it should be noted that the above embodiments are only used to illustrate the technical solutions of the present invention and not to limit the technical solutions. Those skilled in the art should understand that any modifications or equivalent substitutions to the technical solutions of the present invention without departing from the spirit and scope of the present invention should be covered within the scope of the claims of the present invention.

Claims

1. A traceable hydrogel for cartilage repair, characterized in that, A hydrogel with a three-dimensional biomimetic network is formed by chitosan, sericin and collagen, and a magnetic material is uniformly dispersed in the hydrogel; wherein the magnetic material is selected from nano Fe3O4 or γ-Fe2O3.

2. The traceable hydrogel according to claim 1, characterized in that, The traceable hydrogel was prepared by the following method: Step 1: Mix chitosan and magnetic material, then add to deionized water. Next, add sericin and collagen under continuous stirring until completely dissolved to obtain a composite precursor solution. The concentrations of chitosan, sericin, collagen, and magnetic material in the composite precursor solution are as follows (by mass percentage): 4% chitosan, 1%–2% sericin, 1%–2% collagen, and 0.2%–0.8%. Step 2: Slowly add a mixed solution of sodium β-glycerophosphate and NaHCO3 to the composite precursor solution obtained in Step 1, stir evenly, and then inject it into the target location to form the traceable hydrogel; wherein, the mass ratio of sodium β-glycerophosphate to NaHCO3 is 10:1, and the mass ratio of chitosan to sodium β-glycerophosphate is (1:2) to (1:3).

3. The traceable hydrogel according to claim 2, characterized in that, In step 1, chitosan is dissolved in a dilute acetic acid solution or a vitamin C solution; wherein the concentration of the dilute acetic acid or vitamin C solution is 1%.

4. The traceable hydrogel according to claim 1, characterized in that, In step 1, the magnetic material is added to an aqueous solution containing polyethylene glycol, vortexed for 3 to 5 minutes, and then ultrasonically treated for 10 to 30 minutes to obtain a stable Fe3O4 dispersion; wherein the concentration of PEG in the polyethylene glycol aqueous solution is 1 wt% to 5 wt%.

5. The traceable hydrogel according to claim 1, characterized in that, In step 2, the traceable hydrogel is formed by standing at 37°C for 30 to 60 minutes.

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