Bone tissue repair material and preparation method thereof

Abstract

The invention relates to the technical field of biological materials, and discloses a bone tissue repair material and a preparation method thereof.According to the material, hydroxyapatite, octacalcium phosphate and silicon-doped beta-tricalcium phosphate are adopted for constructing three-phase ceramic microspheres, a two-stage pore structure is formed through a pore forming agent method, chitosan coating and gradient biomimetic mineralization treatment are sequentially conducted on the surface, and the bone tissue repair material is obtained. And compounding with a strontium-barium double-ion cross-linked sodium alginate-collagen-oxidized chitosan ternary matrix, and adding hemostatic functional microspheres. According to the invention, multiple technical breakthroughs such as in-vivo dynamic mineralization of octacalcium phosphate, strontium-barium double-ion anti-displacement cross-linking, oxidized chitosan covalent cross-linking, gradient mineralization layer interface enhancement, rapid hemostasis of hemostatic microspheres and the like are adopted; the comprehensive technical effects that the mechanical strength of the material is improved, the degradation rate is adjustable, the antibacterial performance is enhanced, the in-vivo stability is remarkably improved, hemostasis is rapid, and the bioactive coating is firm and does not fall off are achieved.

Description

Technical Field

[0001] This invention relates to the field of biomaterials technology, and more specifically, to a bone tissue repair material and its preparation method. Background Technology

[0002] Bone repair materials are used to repair bone defects caused by various reasons, including trauma, tumor resection, and bone defects remaining after infection and debridement. Existing bone repair materials typically combine bone morphogenetic proteins (BMPs) with natural bone matrix (such as collagen), utilizing the osteoinductive activity of BMPs and the scaffolding effect of collagen to promote bone regeneration. However, pure collagen matrix materials have low mechanical strength, which cannot meet the requirements of applications in load-bearing areas. To improve mechanical properties, researchers have added ceramic particles such as hydroxyapatite (HA) or β-tricalcium phosphate (β-TCP) to the collagen matrix to form composite materials.

[0003] Despite improvements in the mechanical properties of ceramic-collagen composites, the following technical challenges remain to be addressed:

[0004] First, it is difficult to precisely match the material degradation rate with the bone regeneration rate. Hydroxyapatite degrades slowly, while β-tricalcium phosphate degrades relatively quickly. Although biphasic ceramics can regulate the degradation rate to some extent, it is still difficult to achieve dynamic synchronization with the osteogenic process. Premature material degradation leads to insufficient structural support, while excessively slow degradation hinders bone tissue ingrowth and remodeling.

[0005] Second, materials containing hydrogel components (such as sodium alginate) have poor stability in vivo. Sodium alginate is usually formed by physical cross-linking with calcium ions to form a gel, but after implantation in the human body, the high concentration of sodium and magnesium ions in the tissue fluid rapidly replaces the calcium ions at the binding sites, causing the cross-linking network to disintegrate and the mechanical properties of the material to drop significantly within a few days.

[0006] Third, bone defects are often accompanied by bleeding, and materials in the blood environment are prone to expansion, disintegration or displacement, affecting the early stability of the implant and lacking hemostatic function, thus prolonging the operation time.

[0007] Fourth, the bioactive coating formed on the material surface by biomimetic mineralization using simulated body fluid (SBF) is prone to peeling or cracking during plastic manipulation, and the bonding strength between the coating and the substrate is insufficient, making it easy to fall off after implantation.

[0008] Fifth, existing materials lack antibacterial properties, resulting in a high risk of infection after implantation, especially in cases of traumatic fractures or open bone defects. Summary of the Invention

[0009] To address the aforementioned technical problems, this invention provides a bone tissue repair material and its preparation method.

[0010] A bone tissue repair material comprising a ceramic phase and an organic matrix phase;

[0011] The ceramic phase is a porous microsphere, composed of hydroxyapatite, octacalcium phosphate, and silicon-doped β-tricalcium phosphate in a mass ratio of 3:2:5, wherein the silicon content of the silicon-doped β-tricalcium phosphate is 1.5-3.5 wt%.

[0012] The porous microspheres have a dual-level pore structure, with macropores having a diameter of 200-500 μm, micropores having a diameter of 5-50 μm, and a porosity of 60-80%. The surface of the porous microspheres is sequentially coated with a chitosan layer and a gradient mineralization layer. The chitosan layer has a thickness of 10-30 μm, and the gradient mineralization layer includes an inner mineralization layer and an outer mineralization layer. The inner mineralization layer has a thickness of 2-5 μm, and the total mineralization layer thickness is 8-15 μm.

[0013] The organic matrix phase includes calf bone collagen, sodium alginate, oxidized chitosan, and hemostatic microspheres;

[0014] The sodium alginate is cross-linked with strontium ions, barium ions, and calcium ions, with a molar ratio of Sr:Ba:Ca of 4:3:3; the oxidized chitosan has an oxidation degree of 15-25% and is used in amounts of 10-20% of the sodium alginate mass; the hemostatic microspheres are sodium alginate-gelatin composite microspheres with a diameter of 50-150 μm, encapsulating thrombin and tranexamic acid inside;

[0015] The porous microspheres of the ceramic phase are loaded with bone morphogenetic proteins.

[0016] The volume ratio of the ceramic phase to the organic matrix phase is 4:6.

[0017] Preferably, the porous microspheres are prepared by a pore-forming agent method, wherein the pore-forming agent is polymethyl methacrylate microspheres, including microspheres with a particle size of 200-500 μm and microspheres with a particle size of 5-50 μm, and the mass ratio of the two particle sizes is 7:3; the particle size of the porous microspheres is 0.5-2 mm.

[0018] Preferably, the chitosan has a molecular weight of 100,000-500,000 Da, a degree of deacetylation ≥85%, and is used in an amount of 3-8% of the mass of the ceramic microspheres.

[0019] Preferably, the gradient mineralization layer is formed by two-step biomimetic mineralization: first, it is soaked in a simulated body fluid of 0.5 times concentration at 37°C for 12 hours to form an inner mineralization layer, and then soaked in a simulated body fluid of standard concentration at 37°C for 24 hours to form an outer mineralization layer; the main component of the gradient mineralization layer is hydroxyapatite, and the calcium-to-phosphorus ratio is 1.60-1.67.

[0020] Preferably, the sodium alginate has a molecular weight of 200,000-300,000 Da and a concentration of 2-4%.

[0021] Preferably, the aldehyde groups of the oxidized chitosan form Schiff base covalent crosslinks with the amino groups of collagen and sodium alginate, and form a double crosslinking network with strontium-barium-calcium ions.

[0022] Preferably, the thrombin activity is ≥1000 IU / mg, and the amount used is 1-3% of the gelatin mass; the amount of tranexamic acid used is 5-10% of the gelatin mass; and the amount of hemostatic microspheres used is 3-8% of the total organic matrix mass.

[0023] A method for preparing a bone tissue repair material includes the following steps:

[0024] Step 1: Silicon-doped tricalcium β-phosphate was prepared using the sol-gel method, with a silicon content of 1.5-3.5 wt%.

[0025] Step 2: Silicon-doped β-tricalcium phosphate, hydroxyapatite and octacalcium phosphate are mixed in a mass ratio of 5:3:2, and polymethyl methacrylate microspheres are added as pore-forming agents. After molding, the mixture is sintered at 1100-1200℃ to form ceramic microspheres with a bilevel pore structure.

[0026] Step 3: Immerse the ceramic microspheres in a chitosan solution for coating, with a coating thickness of 10-30 μm;

[0027] Step 4: Perform gradient biomimetic mineralization on the coated ceramic microspheres: first, soak them in a simulated body fluid at 0.5 times the concentration at 37°C for 12 hours to form an inner mineralization layer, and then soak them in a simulated body fluid at a standard concentration at 37°C for 24 hours to form an outer mineralization layer. The total mineralization layer thickness is 8-15 μm.

[0028] Step 5: Prepare sodium alginate-gelatin composite microspheres, with thrombin and tranexamic acid encapsulated within the microspheres;

[0029] Step 6: Mix calf bone collagen, sodium alginate solution, oxidized chitosan and hemostatic microspheres to prepare an organic matrix paste;

[0030] Step 7: Add strontium ions, barium ions, and calcium ions to the organic matrix paste, with a molar ratio of Sr:Ba:Ca of 4:3:3, and perform composite cross-linking;

[0031] Step 8: Load bone morphogenetic proteins into the pores of the treated ceramic microspheres;

[0032] Step 9: Mix ceramic microspheres loaded with bone morphogenetic proteins with cross-linked organic matrix paste at a volume ratio of 4:6 to form a plastic composite material.

[0033] Preferably: In step 7, the mixed ion crosslinking solution is prepared by dissolving strontium chloride, barium chloride, and calcium chloride in deionized water at a molar ratio of 4:3:3 to prepare a crosslinking solution with a total divalent cation concentration of 50-100 mM; the amount of crosslinking solution added is calculated based on the concentration of sodium alginate, and the total molar amount of divalent ions is 0.4-0.6 times the molar amount of sodium alginate monomer.

[0034] Preferably, in step 4, the composition of the 0.5 times concentration simulated body fluid is: NaCl 4.0 g / L, NaHCO3 0.175 g / L, KCl 0.112 g / L, K2HPO4·3H2O 0.114 g / L, MgCl2·6H2O 0.1525 g / L, CaCl2 0.1835 g / L, Na2SO4 0.0355 g / L, pH 7.4;

[0035] The standard concentration of simulated body fluid consists of: NaCl 8.0 g / L, NaHCO3 0.35 g / L, KCl 0.224 g / L, K2HPO4·3H2O 0.228 g / L, MgCl2·6H2O 0.305 g / L, CaCl2 0.367 g / L, Na2SO4 0.071 g / L, and pH 7.4.

[0036] The beneficial effects of this invention are as follows:

[0037] (1) Dynamic mineralization function: Octalcalcium phosphate (OCP) is introduced as the precursor phase of the ceramic phase. Under physiological conditions, OCP is gradually converted into hydroxyapatite, and the conversion process continuously releases calcium and phosphorus ions and provides the driving force for mineralization. Compared with the existing static ceramics (HA / β-TCP), OCP achieves "dynamic mineralization", which enables the degradation of the material to proceed synchronously with the bone mineralization process in vivo, breaking through the technical bottleneck of the non-adjustable degradation rate of existing materials. The differentiated degradation of the three-phase ceramics (HA / OCP / Si-TCP) realizes the time-sequential function of early cell migration, mid-term accelerated osteogenic formation, and late-term structural stability.

[0038] (2) Significant improvement in in vivo stability: The use of a strontium-barium-calcium ternary ion crosslinking (Sr:Ba:Ca = 4:3:3) replaces the traditional single calcium ion crosslinking. Utilizing the high coordination stability constants and large ionic radius effects of barium and strontium ions with sodium alginate, the fundamental problem of rapid disintegration of sodium alginate gel in vivo due to ion exchange is solved from the perspective of coordination chemistry. Combined with Schiff base covalent crosslinking of oxidized chitosan, a dual crosslinking network of "physical crosslinking + chemical crosslinking" is formed, resulting in a qualitative leap in the stability of the material in vivo. This technological breakthrough allows the material to be applied to bone marrow cavities and hypervascular sites with abundant tissue fluid and high risk of ion exchange, expanding its clinical application scope.

[0039] (3) Strong bonding of the bioactive coating: An innovative gradient biomimetic mineralization method (first 0.5×SBF, then 1×SBF) is adopted to form a mineralized coating with a gradient transition between the inner and outer layers, eliminating the interface abruptness and stress concentration problems of traditional one-step mineralization. Combined with the flexible buffering effect of the chitosan coating layer, a composite bonding of "physical anchoring + chemical bonding" between the mineralized layer and the ceramic matrix is ​​achieved, which significantly improves the crack resistance and spalling resistance of the coating during plastic operation. This technological breakthrough enables the material to have a stable bioactive surface while maintaining plasticity.

[0040] (4) Integration of rapid hemostasis: Sodium alginate-gelatin microspheres encapsulating thrombin and tranexamic acid are introduced into the material, achieving the dual function of "rapid hemostasis + continuous hemostasis". The microspheres rapidly release hemostatic components in the blood environment, effectively solving the problems of material expansion, disintegration and displacement caused by bleeding at bone defect sites, and improving the early stability of the implant. This is a function that is generally lacking in existing bone repair materials. This embodiment achieves the integration of hemostasis function through innovation in material composition and structure, without the need for additional hemostatic materials.

[0041] (5) Synergistic antibacterial function: Silicon ions and the chitosan coating layer work together to provide antibacterial effects. Chitosan exerts its natural antibacterial activity by disrupting bacterial cell walls, while silicon ions exert their antibacterial effect by interfering with bacterial metabolism. Together, they reduce the risk of infection after implantation. This function is of great value in clinical situations with high infection risk, such as traumatic fractures and open bone defects.

[0042] (6) Synergistic effect of multiple elements: Silicon ions promote heterogeneous nucleation of hydroxyapatite and accelerate in vivo mineralization; chitosan enhances the bonding of ceramic-organic matrix interface; bilevel pores realize the functional partitioning of cell migration channels and bioactive interfaces; BMP achieves dual-phase release in bilevel pores; and multiple technical elements generate a synergistic enhancement effect, making the comprehensive performance of the material better than the simple sum of each individual technical element.

[0043] (7) Balancing plasticity and high strength: By optimizing the volume ratio of ceramic phase to organic matrix phase in a ratio of 4:6 and controlling the gelation rate of strontium-barium-calcium crosslinking, the material maintains a plastic operation time of 25-35 minutes while achieving a compressive strength of 18-28 MPa after curing, thus achieving a balance between plasticity and mechanical strength, and simultaneously meeting the two application conditions of intraoperative shaping and post-implantation load bearing. Attached Figure Description

[0044] Figure 1 This is a bar chart comparing the shear strength of the mineralized layer interface according to the present invention;

[0045] Figure 2 This is a bar chart comparing the spalling rate of the mineralized layer after the plastic operation of this invention;

[0046] Figure 3 This is a bar chart comparing in vitro coagulation times according to the present invention;

[0047] Figure 4 This is a bar chart comparing the hemostasis time in animal models according to the present invention;

[0048] Figure 5 This is the release kinetic curve of the hemostatic active ingredient of the present invention. Detailed Implementation

[0049] The subject matter described herein will now be discussed with reference to exemplary embodiments. It should be understood that these embodiments are discussed only to enable those skilled in the art to better understand and implement the subject matter described herein, and changes may be made to the function and arrangement of the elements discussed without departing from the scope of this specification. Various processes or components may be omitted, substituted, or added as needed in the examples. Furthermore, some features described in the examples may be combined in other examples.

[0050] Example 1

[0051] This embodiment proposes a bone tissue repair material, comprising a ceramic phase and an organic matrix phase;

[0052] The ceramic phase is a porous microsphere composed of hydroxyapatite, octacalcium phosphate, and silicon-doped β-tricalcium phosphate in a mass ratio of 3:2:5, wherein the silicon content of the silicon-doped β-tricalcium phosphate is 2.5 wt%.

[0053] The porous microspheres have a dual-level pore structure with a macropore diameter of 350 μm, a micropore diameter of 30 μm, and a porosity of 70%. The surface of the porous microspheres is sequentially coated with a chitosan layer and a gradient mineralization layer. The chitosan layer has a thickness of 20 μm, and the gradient mineralization layer includes an inner mineralization layer and an outer mineralization layer. The inner mineralization layer has a thickness of 3 μm, and the total mineralization layer thickness is 11 μm.

[0054] The organic matrix phase includes calf bone collagen, sodium alginate, oxidized chitosan, and hemostatic microspheres;

[0055] The sodium alginate is cross-linked with strontium ions, barium ions and calcium ions, and the molar ratio of Sr:Ba:Ca is 4:3:3; the oxidation degree of the oxidized chitosan is 20%, and the amount used is 15% of the mass of sodium alginate; the hemostatic microspheres are sodium alginate-gelatin composite microspheres with a diameter of 100μm, and encapsulate thrombin and tranexamic acid inside.

[0056] The porous microspheres of the ceramic phase are loaded with bone morphogenetic proteins.

[0057] The volume ratio of the ceramic phase to the organic matrix phase is 4:6.

[0058] The porous microspheres are prepared by a pore-forming agent method, wherein the pore-forming agent is polymethyl methacrylate microspheres, including microspheres with a particle size of 200-500 μm and microspheres with a particle size of 25 μm, and the mass ratio of the two particle sizes is 7:3; the particle size of the porous microspheres is 1 mm.

[0059] The chitosan has a molecular weight of 300,000 Da, a degree of deacetylation ≥85%, and is used at 5% of the mass of the ceramic microspheres.

[0060] The gradient mineralization layer is formed through a two-step biomimetic mineralization process: first, it is immersed in a simulated body fluid at 0.5 times the concentration at 37°C for 12 hours to form an inner mineralization layer, and then it is immersed in a simulated body fluid at a standard concentration at 37°C for 24 hours to form an outer mineralization layer; the main component of the gradient mineralization layer is hydroxyapatite, and the calcium-to-phosphorus ratio is 1.63.

[0061] The sodium alginate has a molecular weight of 250,000 Da and a concentration of 3%.

[0062] The aldehyde groups of the oxidized chitosan form Schiff base covalent crosslinks with the amino groups of collagen and sodium alginate, and form a double crosslinked network with strontium-barium-calcium ions.

[0063] The thrombin activity is ≥1000 IU / mg, and the dosage is 2% of the gelatin mass; the tranexamic acid dosage is 8% of the gelatin mass; and the hemostatic microspheres dosage is 5% of the total organic matrix mass.

[0064] Example 2

[0065] The difference between this embodiment and Embodiment 1 is that:

[0066] The silicon content of the silicon-doped β-tricalcium phosphate is 1.5 wt%.

[0067] The porous microspheres have a dual-level pore structure, with macropores having a diameter of 200 μm, micropores having a diameter of 5 μm, and a porosity of 60%.

[0068] The chitosan layer has a thickness of 10 μm; the inner mineralized layer has a thickness of 2 μm, and the total mineralized layer thickness is 8 μm;

[0069] The degree of oxidation of the oxidized chitosan is 15%, and the amount used is 10% of the mass of sodium alginate; the diameter of the hemostatic microspheres is 50 μm;

[0070] The pore-forming agent is polymethyl methacrylate microspheres, including microspheres with a particle size of 200 μm and microspheres with a particle size of 5 μm, with a mass ratio of the two particle sizes of 7:3; the particle size of the porous microspheres is 0.5 mm.

[0071] The chitosan has a molecular weight of 100,000 Da, a degree of deacetylation ≥85%, and is used at 3% of the mass of the ceramic microspheres.

[0072] The main component of the gradient mineralization layer is hydroxyapatite, with a calcium-to-phosphorus ratio of 1.60.

[0073] The sodium alginate has a molecular weight of 200,000 Da and a concentration of 2%.

[0074] The thrombin activity is ≥1000 IU / mg, and the dosage is 1% of the gelatin mass; the tranexamic acid dosage is 5% of the gelatin mass; and the hemostatic microspheres dosage is 3% of the total organic matrix mass.

[0075] Example 3

[0076] The difference between this embodiment and Embodiment 1 is that:

[0077] The silicon content of the silicon-doped β-tricalcium phosphate is 3.5 wt%.

[0078] The porous microspheres have a dual-level pore structure, with macropores having a diameter of 500 μm, micropores having a diameter of 50 μm, and a porosity of 80%.

[0079] The chitosan layer has a thickness of 30 μm; the inner mineralized layer has a thickness of 5 μm; and the total mineralized layer thickness is 15 μm.

[0080] The degree of oxidation of the oxidized chitosan is 25%, and the amount used is 20% of the mass of sodium alginate; the diameter of the hemostatic microspheres is 150 μm;

[0081] The pore-forming agent is polymethyl methacrylate microspheres, including microspheres with a particle size of 500 μm and microspheres with a particle size of 50 μm, with a mass ratio of the two particle sizes of microspheres of 7:3; the particle size of the porous microspheres is 2 mm.

[0082] The chitosan has a molecular weight of 500,000 Da, a degree of deacetylation ≥85%, and is used at 8% of the mass of the ceramic microspheres.

[0083] The main component of the gradient mineralization layer is hydroxyapatite, with a calcium-to-phosphorus ratio of 1.67.

[0084] The sodium alginate has a molecular weight of 300,000 Da and a concentration of 4%.

[0085] The thrombin activity is ≥1000 IU / mg, and the dosage is 3% of the gelatin mass; the tranexamic acid dosage is 10% of the gelatin mass; and the hemostatic microspheres dosage is 8% of the total organic matrix mass.

[0086] Example 4

[0087] This embodiment proposes a method for preparing a bone tissue repair material, including the following steps:

[0088] Step 1: Silicon-doped tricalcium β-phosphate was prepared using the sol-gel method, with a silicon content of 1.5-3.5 wt%.

[0089] Step 2: Silicon-doped β-tricalcium phosphate, hydroxyapatite and octacalcium phosphate are mixed in a mass ratio of 5:3:2, and polymethyl methacrylate microspheres are added as pore-forming agents. After molding, the mixture is sintered at 1100-1200℃ to form ceramic microspheres with a bilevel pore structure.

[0090] Step 3: Immerse the ceramic microspheres in a chitosan solution for coating, with a coating thickness of 10-30 μm;

[0091] Step 4: Perform gradient biomimetic mineralization on the coated ceramic microspheres: first, soak them in a simulated body fluid at 0.5 times the concentration at 37°C for 12 hours to form an inner mineralization layer, and then soak them in a simulated body fluid at standard concentration at 37°C for 24 hours to form an outer mineralization layer. The total mineralization layer thickness is 8-15 μm.

[0092] Step 5: Prepare sodium alginate-gelatin composite microspheres, with thrombin and tranexamic acid encapsulated within the microspheres;

[0093] Step 6: Mix calf bone collagen, sodium alginate solution, oxidized chitosan and hemostatic microspheres to prepare an organic matrix paste;

[0094] Step 7: Add strontium ions, barium ions, and calcium ions to the organic matrix paste, with a molar ratio of Sr:Ba:Ca of 4:3:3, and perform composite cross-linking;

[0095] Step 8: Load bone morphogenetic proteins into the pores of the treated ceramic microspheres;

[0096] Step 9: Mix ceramic microspheres loaded with bone morphogenetic proteins with cross-linked organic matrix paste at a volume ratio of 4:6 to form a plastic composite material.

[0097] In step 7, a mixed ion crosslinking solution is prepared by dissolving strontium chloride, barium chloride, and calcium chloride in deionized water at a molar ratio of 4:3:3 to prepare a crosslinking solution with a total divalent cation concentration of 50-100 mM. The amount of crosslinking solution added is calculated based on the concentration of sodium alginate, and the total molar amount of divalent ions is 0.4-0.6 times the molar amount of sodium alginate monomer.

[0098] In step 4, the composition of the 0.5-fold concentration simulated body fluid is as follows: NaCl 4.0 g / L, NaHCO3 0.175 g / L, KCl 0.112 g / L, K2HPO4·3H2O 0.114 g / L, MgCl2·6H2O 0.1525 g / L, CaCl2 0.1835 g / L, Na2SO4 0.0355 g / L, pH 7.4;

[0099] The standard concentration of simulated body fluid consists of: NaCl 8.0 g / L, NaHCO3 0.35 g / L, KCl 0.224 g / L, K2HPO4·3H2O 0.228 g / L, MgCl2·6H2O 0.305 g / L, CaCl2 0.367 g / L, Na2SO4 0.071 g / L, and pH 7.4.

[0100] Example 5

[0101] This embodiment proposes a method for preparing a bone tissue repair material, including the following specific implementation steps:

[0102] Step 1: Preparation of silicon-doped β-tricalcium phosphate

[0103] This step uses the sol-gel method to prepare silicon-doped β-tricalcium phosphate (Si-TCP) with a silicon content of 1.5-3.5 wt%.

[0104] Silicon-doped β-tricalcium phosphate (TCPP) is the focus of this embodiment. Compared with undoped β-TCP, the introduction of silicon ions has the following technological breakthroughs: (1) Silicon ions are slowly released after implantation, promoting osteoblast proliferation and differentiation; (2) Silicon ions promote heterogeneous nucleation of hydroxyapatite on the ceramic surface, accelerating the in vivo mineralization process; (3) Silicon doping makes the degradation rate of β-TCP adjustable, forming a differentiated three-phase degradation combination with octacalcium phosphate and hydroxyapatite. The silicon content is controlled at 1.5-3.5 wt% based on the following considerations: when it is below 1.5 wt%, the enhancement effect on biological activity is not obvious, and when it is above 3.5 wt%, it will affect the stability of the β-TCP crystal phase.

[0105] Preparation method: Calcium hydrogen phosphate (DHT) ) and calcium carbonate ( Dissolve calcium carbonate in nitric acid solution (10-15% concentration) at a calcium-to-phosphorus molar ratio of 1.5:1. Calcium carbonate reacts with nitric acid to produce carbon dioxide gas; therefore, this process must be performed in a fume hood. Add tetraethyl orthosilicate ( Using silicon as the silicon source, the silicon content is controlled at 1.5-3.5 wt%, preferably 2.5 wt%. The solution is stirred and hydrolyzed at 60-80℃ (preferably 70℃) for 3-5 hours to form a silicon-containing precursor sol. After drying the sol in a drying oven at 120℃ for 24 hours, it is sintered at 800-1000℃ (preferably 900℃) for 2 hours to obtain silicon-doped β-tricalcium phosphate ceramics. X-ray diffraction (XRD) analysis confirms that the product is a β-TCP crystalline phase, and energy dispersive spectroscopy (EDS) verifies that the silicon element is uniformly distributed.

[0106] Preferred option: When the silicon content is 2.5wt%, the bioactivity and crystal phase stability of the material reach the best balance, and the osteoblast proliferation rate is most significantly improved.

[0107] Step 2: Preparation of three-phase ceramic composite powder and fabrication of bi-level porous microspheres

[0108] In this step, Si-TCP, hydroxyapatite (HA), and octacalcium phosphate (OCP) are mixed in a mass ratio of 5:3:2, and polymethyl methacrylate (PMMA) microspheres are added as pore-forming agents. After molding, the mixture is sintered at high temperature to form ceramic microspheres with a bilevel pore structure.

[0109] (1) Composition of the three-phase ceramic: Existing technologies typically employ HA / β-TCP biphase ceramics. This embodiment innovatively introduces octacalcium phosphate (OCP) to form a three-phase ceramic. OCP is a thermodynamically metastable phase that gradually transforms into hydroxyapatite under physiological conditions. This phase transformation process continuously releases calcium and phosphorus ions and provides the driving force for mineralization, achieving the function of "dynamic mineralization". The selection of the three-phase mass ratio of 5:3:2 (Si-TCP:HA:OCP) is based on the precise control of the degradation rate: OCP is transformed into HA within 3-8 weeks, Si-TCP degrades moderately within 4-12 weeks, and HA is maintained for a long time (3-6 months). The three phases form a time-sequential degradation from early to middle to late stages, corresponding to the functional requirements of different stages of bone regeneration. Compared with existing biphase ceramics, this adds a degradation regulation function in the time dimension.

[0110] (2) Bi-level pore structure: In this embodiment, two PMMA microspheres with different particle sizes (200-500 μm and 5-50 μm, respectively) were selected as pore-forming agents. During sintering, the PMMA decomposed and volatilized to form a bi-level pore structure—macropores with a diameter of 200-500 μm and micropores with a diameter of 5-50 μm, resulting in a total porosity of 60-80%. The bi-level pore structure has functional partitions: macropores provide channels for cell migration and blood vessel ingrowth, shortening the time for cells to reach the interior of the material; micropores provide a high specific surface area for protein adsorption, ion exchange, and growth factor loading. Compared with the single-pore size of existing technologies, this hierarchical pore structure achieves synergistic optimization of cell migration speed and bioactive interface area.

[0111] Preparation method: Si-TCP powder, HA powder, and OCP powder are mixed at a mass ratio of 5:3:2 and ball-milled for 24 hours to obtain a uniform composite powder. Two types of PMMA microspheres (70% with a particle size of 200-500 μm and 30% with a particle size of 5-50 μm) are added to the composite powder, with the total PMMA content being 40-60% of the ceramic powder mass, preferably 50%. Polyvinyl alcohol (PVA) binder (5% aqueous solution, 3-5% of the ceramic powder mass) and deionized water are added, mixed evenly, and granulated to form spherical particles with a diameter of 0.5-2 mm. The particles are pre-fired at 600℃ for 2 hours to remove organic matter, with a heating rate of 5℃ / min; then sintered at 1100-1200℃ (preferably 1150℃) for 3-4 hours (preferably 3.5 hours), with a heating rate of 10℃ / min. After cooling, three-phase ceramic microspheres with a bilevel porous structure are obtained. Scanning electron microscopy (SEM) confirmed that a hierarchical pore network with both macropores and micropores was formed inside the microspheres.

[0112] Preferred scheme: When the amount of PMMA is 50% of the mass of ceramic powder, the sintering temperature is 1150℃ and the sintering time is 3.5 hours, the porosity of the obtained ceramic microspheres is stable at 70%, the pore connectivity is optimal, and the mechanical strength and biological activity are optimally balanced.

[0113] Step 3: Chitosan coating of ceramic microspheres

[0114] In this step, ceramic microspheres are immersed in a chitosan solution for impregnation and coating, forming a chitosan thin layer with a thickness of 10-30 μm. The amount of chitosan used is 3-8% of the mass of the ceramic microspheres.

[0115] The introduction of the chitosan coating layer has a triple functional breakthrough: (1) Chitosan has natural antibacterial activity, exerts its antibacterial effect by destroying the bacterial cell wall, and forms a synergistic antibacterial effect with the subsequently introduced silicon ions; (2) The amino groups of chitosan ( ) and phosphate groups on the ceramic surface ( (2) Hydrogen bonds and electrostatic interactions are formed, and hydrogen bond networks are also formed with the collagen matrix in subsequent steps, which significantly enhances the bonding strength of the ceramic-organic matrix interface and prevents the agglomeration and detachment of ceramic microspheres inside the material; (3) Chitosan, as a flexible organic layer, plays a buffering role in the subsequent gradient mineralization steps, protects the ceramic matrix and provides flexible support for the mineralization layer, and improves the crack resistance of the mineralization layer during plastic operation. The coating thickness is controlled at 10-30μm: when it is less than 10μm, the coating is uneven and the function is insufficient, and when it is greater than 30μm, it affects the material exchange between the ceramic phase and the matrix.

[0116] The chitosan has a molecular weight of 100,000-500,000 Da and a degree of deacetylation ≥85%. The molecular weight of 100,000-500,000 Da ensures that the chitosan has appropriate viscosity and film-forming properties, while the degree of deacetylation ≥85% ensures sufficient amino content to exert antibacterial and adhesive functions.

[0117] Preparation method: Chitosan is dissolved in a 1-2% (preferably 1.5%) acetic acid solution to prepare a 2-3% (preferably 2.5%) chitosan solution. Ceramic microspheres are immersed in the chitosan solution under vacuum for 30-60 minutes (preferably 45 minutes) at a vacuum degree of -0.08 MPa, allowing the solution to fully penetrate the pores and coat the surface. The microspheres are then removed, rinsed three times with deionized water to remove excess chitosan, and dried at room temperature for 12-24 hours (preferably 18 hours) until constant weight. A uniform chitosan thin layer forms on the surface of the coated microspheres, with the chitosan content controlled at 3-8% of the ceramic microsphere mass. The characteristic peak of chitosan (1650 nm) is confirmed by Fourier transform infrared spectroscopy (FTIR). Amide I band, 1590 (Amino bending vibration) proves that chitosan was successfully encapsulated.

[0118] The preferred method is to use an acetic acid concentration of 1.5%, a chitosan concentration of 2.5%, an impregnation time of 45 minutes, and a drying time of 18 hours to achieve the best coating thickness uniformity, with an average thickness of 20 μm, and the strongest bond with the ceramic substrate.

[0119] Step 4: Gradient biomimetic mineralization treatment of ceramic microspheres

[0120] This step involves gradient biomimetic mineralization treatment of the chitosan-coated ceramic microspheres: first, they are immersed in a simulated body fluid (0.5×SBF) at 37°C for 12 hours to form an inner mineralization layer of 2-5 μm thickness; then, they are immersed in a simulated body fluid (1×SBF) at a standard concentration at 37°C for 24 hours to form an outer mineralization layer, with a total mineralization layer thickness of 8-15 μm.

[0121] (1) Gradient mineralization: Existing technologies typically use a single concentration of SBF for one-time mineralization, resulting in a significant interface abruptness between the formed mineralized layer and the matrix, leading to interface stress concentration. Under shear forces and bending deformations during plastic operation, interface cracking and spalling are prone to occur. This embodiment innovatively adopts a two-step gradient mineralization strategy: The first step is carried out in a low concentration of SBF (0.5×), where the ion supersaturation is low and the mineralization rate is slow. The resulting inner mineralized layer (2-5μm) has fine grains and low density, forming a good penetration transition with the chitosan coating layer, similar to "rooting" in the chitosan layer; The second step is carried out in a standard concentration of SBF (1×), where the ion supersaturation is increased and the mineralization rate is accelerated. An outer mineralized layer grows on the inner layer, with larger grains and higher density, providing the main biological activity. This gradient mineralization makes the mineralized layer exhibit a gradient transition from "low density to high density" and "small grains to large grains" from the inside to the outside, eliminating interface abruptness and significantly improving the bonding strength between the mineralized layer and the matrix.

[0122] (2) Differences from existing technologies: Existing one-step mineralization is carried out in standard concentration SBF, resulting in a fast mineralization rate. The mineralized layer is directly deposited on the substrate surface, and the interfacial bonding mainly relies on physical adsorption, leading to weak bonding force. In this embodiment, gradient mineralization allows the inner mineralized layer to penetrate the pores and surface irregularities of the chitosan layer, forming a composite bonding method of "physical anchoring + chemical bonding." The outer mineralized layer grows continuously on the inner layer, forming a continuation of crystal orientation, and the overall mineralized layer forms an integrated structure with the substrate. Experimental results show that the interfacial bonding strength of the gradient mineralization layer is more than 3 times higher than that of one-step mineralization, and the mineralization layer peeling rate after plastic operation is reduced from 30-40% to below 5%.

[0123] (3) Control of mineralized layer thickness: The total mineralized layer thickness is controlled at 8-15 μm, with an inner layer of 2-5 μm and an outer layer of 6-10 μm. The inner layer thickness of 2-5 μm ensures sufficient penetration depth to form a strong anchor, while the outer layer thickness of 6-10 μm provides a sufficient bioactive interface. The total thickness does not exceed 15 μm to avoid cracking due to excessive rigidity when the material is bent and deformed, while maintaining the connectivity of pores and preventing them from being completely blocked by the mineralized layer.

[0124] Preparation method:

[0125] Step 1: Prepare a 0.5x concentration of simulated body fluid (0.5×SBF), with an ion concentration half that of standard SBF. The specific composition is as follows: 4.0 g / L 0.175 g / L 0.112 g / L 0.114 g / L 0.1525 g / L 0.1835 g / L 0.0355 g / L, pH adjusted to 7.4 with hydrochloric acid and tris(hydroxymethyl)aminomethane. Chitosan-coated ceramic microspheres were immersed in 0.5×SBF and placed in a 37℃ constant temperature water bath for 12 hours. During immersion, calcium phosphate crystals gradually deposited on the ceramic surface and chitosan layer, forming an inner mineralization layer 2-5 μm thick. Scanning electron microscopy revealed that the inner crystals were needle-like or plate-like, small in size (50-200 nm in length), and loosely distributed.

[0126] Step 2: Remove the microspheres, gently rinse with deionized water, and transfer them to a standard concentration simulated body fluid (1×SBF). The ion concentration of the standard SBF is: 8.0 g / L 0.35 g / L, 0.224 g / L 0.228 g / L 0.305 g / L 0.367 g / L 0.071 g / L, pH 7.4. Continue soaking in a constant temperature water bath at 37℃ for 24 hours. At the standard concentration, the mineralization rate accelerates, depositing an outer mineralized layer on top of the inner layer. The outer crystals are larger (200-500 nm in length) and more densely packed. Fresh SBF was replaced every 12 hours during the soaking process to maintain ion supersaturation.

[0127] Step 3: Remove the microspheres, rinse with deionized water to remove surface salts, and dry at room temperature. Scanning electron microscopy (SEM) reveals a continuous gradient transition between the inner and outer layers. X-ray diffraction (XRD) confirms that the main component of the mineralized layer is hydroxyapatite (HA) crystal phase. Energy dispersive spectroscopy (EDS) analysis shows that the calcium-to-phosphorus ratio (Ca / P) of the mineralized layer is 1.60-1.67, close to the theoretical value of stoichiometric hydroxyapatite (1.67). A double-layered mineralized coating with a thickness of 8-15 μm forms on the surface of the gradient-mineralized ceramic microspheres, which forms a strong bond with the chitosan coating layer and the ceramic matrix.

[0128] Step 5: Preparation of hemostatic microspheres

[0129] This step prepares sodium alginate-gelatin composite microspheres with a diameter of 50-150 μm, which encapsulate thrombin and tranexamic acid.

[0130] The introduction of hemostatic microspheres is a key innovation in solving bleeding problems at bone defect sites. Bone defects caused by bone surgery or trauma are often accompanied by intramedullary hemorrhage or periosteal vascular hemorrhage. The blood environment can cause materials to absorb water, swell, disintegrate, and shift, affecting early stability. This embodiment innovatively incorporates microspheres encapsulating hemostatic active ingredients into the material, achieving the following technological breakthroughs:

[0131] (1) Synergistic hemostasis of dual active ingredients: The microspheres simultaneously encapsulate two hemostatic components, thrombin and tranexamic acid. Thrombin catalyzes the conversion of fibrinogen into fibrin, promoting the rapid formation of blood clots and exerting an immediate hemostatic effect; tranexamic acid is a plasmin inhibitor, preventing the formed blood clot from being degraded by the fibrinolytic system and maintaining the hemostatic effect. The synergistic effect of the two achieves the dual function of "rapid hemostasis + continuous hemostasis".

[0132] (2) Sustained-release function of microsphere encapsulation: Thrombin and tranexamic acid, when directly added to the material, will be rapidly inactivated or lost due to diffusion, and may also have adverse interactions with other components. Sodium alginate-gelatin microspheres are used for encapsulation, utilizing the sustained-release function of the microspheres to allow the hemostatic components to be released gradually after implantation: Sodium alginate swells due to calcium ion cross-linking when it comes into contact with calcium-containing body fluids, gelatin dissolves at body temperature, and the microspheres slowly disintegrate and release their contents within 5-30 minutes, with the release rate matching the hemostatic time window.

[0133] (3) Selection of microsphere size: The diameter of the microspheres is controlled between 50-150 μm. This size range is based on the following considerations: a diameter greater than 150 μm will affect the plasticity and uniformity of the material, while a diameter less than 50 μm is difficult to prepare and has a low encapsulation rate. Microspheres of 50-150 μm are uniformly dispersed in the organic matrix, which does not affect the overall structure and mechanical properties of the material. After implantation, they can quickly release hemostatic components in the blood environment.

[0134] Preparation method: Hemostatic microspheres were prepared using an emulsification cross-linking method. Sodium alginate (molecular weight 100,000-200,000 Da, preferably 150,000 Da) was dissolved in deionized water to prepare a 2% solution. Gelatin (isoelectric point type, dissolution temperature 40℃, Bloom value 200-250) was added to prepare a sodium alginate-gelatin mixed solution (mass ratio 1:1). The solution temperature was maintained at 40℃ to keep the gelatin in a liquid state. Thrombin (activity ≥1000 IU / mg) and tranexamic acid (pharmaceutical grade, purity ≥99%) were added to the mixed solution. The amount of thrombin was 1-3% (preferably 2%) of the gelatin mass, and the amount of tranexamic acid was 5-10% (preferably 7.5%) of the gelatin mass. The mixture was then thoroughly mixed. The above mixed solution was added dropwise to vegetable oil (soybean oil) containing calcium chloride (concentration 2-5%, preferably 3.5%), and emulsified under high-speed stirring (1000-2000 rpm, preferably 1500 rpm). The droplets were sheared into microspheres in the oil phase. After stirring for 30-60 minutes (preferably 45 minutes), sodium alginate and calcium ions crosslinked and solidified, and the gelatin cooled and solidified, forming solid microspheres encapsulating the hemostatic component. The microspheres were collected, washed three times repeatedly with ethanol and deionized water to remove the oil phase, and freeze-dried for 48 hours to obtain dried hemostatic microsphere powder. The particle size distribution of the microspheres was measured by optical microscopy or laser particle size analyzer to confirm a diameter of 50-150 μm. The encapsulation efficiency of tranexamic acid in the microspheres was determined by high-performance liquid chromatography (HPLC), and the thrombin activity retention rate was determined by a thrombin activity assay kit to ensure the activity of the hemostatic component.

[0135] The preferred method is as follows: when the molecular weight of sodium alginate is 150,000 Da, the calcium chloride concentration is 3.5%, the stirring speed is 1500 rpm, and the stirring time is 45 minutes, the resulting microspheres have the most uniform particle size distribution (average diameter 100 μm, particle size distribution index <0.3), the thrombin encapsulation rate is over 85%, the tranexamic acid encapsulation rate is over 90%, and the activity retention rate is >80%.

[0136] Step 6: Preparation of organic matrix paste

[0137] This step prepares an organic matrix paste comprising calf bone collagen, sodium alginate solution, oxidized chitosan, and hemostatic microspheres.

[0138] (1) Introduction of oxidized chitosan: Oxidized chitosan introduces aldehyde groups into the chitosan molecular chain through an oxidation reaction. Modified chitosan with an oxidation degree of 15-25% was used at a dosage of 10-20% of the mass of sodium alginate. The aldehyde groups of oxidized chitosan react with the amino groups on the collagen molecules (…). ) and the amino group on the sodium alginate molecule react to form a Schiff base ( The covalent cross-linking reaction introduces covalent bonds into the sodium alginate-collagen dual network, forming a dual cross-linked network of "physical cross-linking (ionic cross-linking) + chemical cross-linking (covalent bonds)". This innovation significantly enhances the stability of the organic matrix in vivo: the simple physical cross-linking of sodium alginate calcium ions disintegrates rapidly in vivo due to ion replacement, while the covalent cross-linking formed after adding oxidized chitosan is not affected by the ionic environment. Even if the physical cross-linking partially disintegrates, the covalent cross-linking can still maintain the network structure, making the mechanical properties of the material stable over a long period of time.

[0139] The reason for controlling the oxidation degree at 15-25% is that when the oxidation degree is below 15%, the aldehyde content is insufficient, the covalent cross-linking density is low, and the reinforcing effect is not obvious; when the oxidation degree is above 25%, too many aldehyde groups will lead to over-cross-linking, making the material brittle and losing its plasticity, and high aldehyde content may cause cytotoxicity. The dosage of 10-20% of the sodium alginate mass is based on the balance between cross-linking density and material properties.

[0140] (2) Dispersion of hemostatic microspheres: The hemostatic microspheres prepared in step 5 are dispersed in an organic matrix paste, with the amount of microspheres being 3-8% of the total mass of the matrix. The microspheres are uniformly dispersed in the matrix, and after implantation, they rapidly swell and release hemostatic components upon contact with the blood environment, thereby achieving the immediate hemostatic function of the material.

[0141] The sodium alginate has a molecular weight of 200,000-300,000 Da and a concentration of 2-4%. The molecular weight of 200,000-300,000 Da ensures that the sodium alginate has appropriate viscosity and cross-linking properties, while the concentration of 2-4% gives the matrix paste suitable consistency and plasticity.

[0142] Preparation method: Calf bone collagen (type I collagen extracted from calf bones, purity ≥95%, isoelectric point pH 7-9) is dispersed in phosphate buffer (0.01 M) at pH 7.2-7.4 to prepare a collagen suspension with a concentration of 3-5% (preferably 4%). The suspension is stirred at room temperature for 2 hours to ensure complete hydration and dispersion of the collagen. Sodium alginate (molecular weight 200,000-300,000 Da, preferably 250,000 Da, mannuronic acid / guluronic acid ratio M / G = 1.5-2.0) is dissolved in deionized water to prepare a sodium alginate solution with a concentration of 2-4% (preferably 3%). Oxidized chitosan (oxidation degree 15-25%, preferably 20%, prepared by oxidizing chitosan with sodium periodate, molecular weight 8-150,000 Da) is dissolved in a weakly acidic buffer solution (acetic acid-sodium acetate buffer) with pH 5-6 (preferably pH 5.5) to prepare an oxidized chitosan solution with a concentration of 1-2% (preferably 1.5%).

[0143] Collagen suspension, sodium alginate solution, and oxidized chitosan solution are mixed at a volume ratio of 5:3:1, with slow stirring (100-200 rpm) to avoid introducing excessive air bubbles. The mixing time is 10-20 minutes (preferably 15 minutes) to ensure thorough mixing of the three components. The hemostatic microsphere powder prepared in step 5 is added to the mixture, at a rate of 3-8% (preferably 5.5%) of the total matrix mass. The mixture is gently stirred (50-100 rpm) for 5-10 minutes to ensure uniform dispersion of the microspheres, avoiding breakage. The pH is adjusted to 7.2-7.4 to promote the reaction between the aldehyde groups of oxidized chitosan and the amino groups of collagen and sodium alginate. pH adjustment is performed using a 0.1 M sodium hydroxide solution. The resulting organic matrix paste is a uniform, viscous paste with good flowability and plasticity, and an apparent viscosity of 5000-8000 mPa·s. Store the paste in a refrigerator at 4°C for up to 24 hours. Before use, allow it to return to room temperature (20-25°C).

[0144] The preferred formulation is as follows: when the collagen concentration is 4%, the sodium alginate (molecular weight 250,000 Da) concentration is 3%, the oxidized chitosan (oxidation degree 20%) concentration is 1.5%, the hemostatic microsphere dosage is 5.5%, and the mixing time is 15 minutes, the resulting organic matrix paste has the best viscosity, flowability and plasticity, the Schiff base crosslinking degree reaches 60-65%, and the material properties are stable within 24 hours.

[0145] Step 7: Strontium-Barium-Calcium Ternary Ion Crosslinking

[0146] In this step, strontium ions, barium ions, and calcium ions are added to the organic matrix paste for composite cross-linking. The molar ratio of the three ions is Sr:Ba:Ca = 4:3:3.

[0147] (1) Technical background of ion replacement problem: Sodium alginate is a linear polysaccharide composed of α-L-guluronic acid (G unit) and β-D-mannuronic acid (M unit), in which the continuous GG sequence is replaced by divalent cations (such as... Bridges form an "egg-box" structure, achieving physical cross-linking. However, human tissue fluid contains high concentrations of monovalent cations (…). 135-145 mM 3.5-5.0 mM) and other divalent cations ( (0.8-1.2 mM), these ions compete with the carboxyl groups of sodium alginate for binding, gradually replacing the cross-linking sites. This leads to the disintegration of the "egg-box" structure and the collapse of the cross-linked network. Studies have shown that simple calcium-crosslinked sodium alginate gel implanted in the body experiences a 60-80% decrease in mechanical strength within 3 days, which is a key bottleneck limiting the clinical application of sodium alginate-based materials.

[0148] (2) Innovative principle of strontium-barium dual-ion anti-displacement: This implementation method innovatively uses strontium ions ( ) and barium ions ( It partially replaces calcium ions for cross-linking, utilizing the coordination chemical differences between different divalent cations and sodium alginate to achieve anti-displacement function. The specific principle is as follows:

[0149] Ionic radius effect: The ionic radius of barium ions (135 pm) is larger than that of calcium ions (100 pm) and strontium ions (118 pm). The coordination bond formed between barium ions and the guluronic acid carboxyl group of sodium alginate is tighter, resulting in greater steric hindrance, making it difficult for sodium and magnesium ions to approach and displace barium ions. The ionic radius of strontium ions is between that of calcium and barium, and its binding strength with sodium alginate is also higher than that of calcium ions.

[0150] Coordination stability constants: Coordination stability constants of strontium ions and barium ions with sodium alginate ( The levels of ions were significantly higher than those of calcium ions. Literature data shows that... The higher the stability constant, the more stable the binding between the ion and the ligand, and the less likely it is to be replaced.

[0151] The synergistic effect of ternary ions: While barium ions alone offer the highest stability for crosslinking, their rapid gelation can lead to premature gelation during mixing, resulting in uneven material composition. Furthermore, pure barium crosslinked gels are relatively brittle. While strontium ions alone provide better stability than calcium ions, the improvement is limited. This embodiment employs a ternary ion ratio of Sr:Ba:Ca = 4:3:3, achieving an optimal balance between stability, gelation rate, and mechanical properties: Barium ions (30%) provide highly stable crosslinking nodes as "anchor points," strontium ions (40%) provide moderately stable crosslinking and impart osteogenic activity (strontium ions themselves have biological functions of promoting osteoblast differentiation and inhibiting osteoclast activity), and calcium ions (30%) regulate the gelation rate and provide some rapid crosslinking. The synergistic effect of these three ions forms a hierarchical crosslinking network of "strong crosslinking + moderate crosslinking + rapid crosslinking," significantly extending the material's stability time in vivo.

[0152] (3) Essential difference from existing technologies: Existing technologies generally use single calcium ion crosslinking, or add strontium ions (Sr-Ca dual ions) on the basis of calcium crosslinking to achieve bone-promoting function, but have never used a strontium-barium-calcium ternary ion ratio, nor have they recognized the key role of barium ions in resisting ion replacement. This embodiment proposes and verifies for the first time the technical solution of strontium-barium-calcium ternary ion composite crosslinking, fundamentally solving the in vivo stability problem of sodium alginate gel through the principle of coordination chemistry, increasing the mechanical strength retention rate of the material after 3 days in vivo from 20-40% to 75-85%, and extending the in vivo stability time from 3 days to 14-21 days, achieving a qualitative breakthrough.

[0153] (4) Basis for selecting the molar ratio 4:3:3: The molar ratio of Sr:Ba:Ca = 4:3:3 was obtained through systematic experimental optimization. The proportion of barium ions should not be too high (>40%), otherwise gelation will be too fast and the material will become brittle; it should not be too low (<20%), otherwise the anti-replacement effect will be insufficient. The proportion of strontium ions should be kept at a relatively high level (40%) to exert its osteogenic activity; the proportion of calcium ions should be kept at a certain level (30%) to regulate the gelation rate and reduce costs. This ratio achieves the optimal balance between in vivo stability, mechanical properties, plasticity and bioactivity.

[0154] Preparation method: Prepare a mixed ion crosslinking solution, and add strontium chloride (SrCl) Analytical grade), barium chloride ( (analytical grade) and calcium chloride ( Analytical grade sodium alginate was dissolved in deionized water at a molar ratio of 4:3:3 to prepare a crosslinking solution with a total divalent cation concentration of 50-100 mM (preferably 75 mM). The mixed ionic crosslinking solution was slowly added dropwise to the organic matrix paste prepared in step 6 at a dropping rate of 1-2 mL / min, while gently stirring (50-100 rpm) to allow strontium, barium, and calcium ions to diffuse evenly into the matrix and combine with the carboxyl groups of sodium alginate. The amount of crosslinking solution added was calculated based on the concentration of sodium alginate, and the total molar amount of divalent ions was 0.4-0.6 times (preferably 0.5 times) the molar amount of sodium alginate monomer (considering that only guluronic acid participates in crosslinking in sodium alginate, and the average molecular weight of sodium alginate monomer is calculated as 198 g / mol).

[0155] Specific calculation method: Assuming 100 g of sodium alginate solution (3% concentration) is used, the mass of sodium alginate is 3 g, the number of monomer moles is 3 ÷ 198 = 0.0152 mol, and the total number of divalent ions required is 0.0152 × 0.5 = 0.0076 mol. If a 75 mM crosslinking solution is used, the volume of crosslinking solution to be added is 0.0076 ÷ 0.075 = 101 mL.

[0156] After the addition of the crosslinking solution, the matrix gradually gelled, and the viscosity increased to 15000-20000 mPa·s. The entire crosslinking process was carried out at room temperature (20-25℃) for 15-30 minutes (preferably 20 minutes) to obtain a crosslinked matrix paste with suitable viscosity and plasticity. The storage modulus of the matrix was confirmed by rheometer testing. ) and loss modulus ( ), This indicates that the material has gelling properties, with a typical value of = 800-1200 Pa, = 200-400 Pa (measured at a frequency of 1 Hz and a strain of 1%). The contents of strontium, barium, and calcium in the matrix were determined by inductively coupled plasma mass spectrometry (ICP-MS), verifying that the molar ratio of the three ions was close to 4:3:3 (with an allowable deviation of ±5%).

[0157] The preferred scheme is as follows: when the total divalent cation concentration is 75 mM, the total molar amount of divalent ions is 0.5 times the molar amount of sodium alginate monomer, and the crosslinking time is 20 minutes, the initial gelation rate, final gel strength, plasticity time and in vivo stability of the material reach the optimal balance, and the strength retention rate reaches 80% after 3 days in vivo.

[0158] Step 8: Loading of bone morphogenetic proteins

[0159] This step involves loading bone morphogenetic protein (BMP) into the pores of the treated ceramic microspheres.

[0160] BMP is a growth factor with osteogenic induction activity, capable of inducing mesenchymal stem cells to differentiate into osteoblasts. By loading BMP into the bipolar pores of ceramic microspheres, the immobilization and sustained release of BMP are achieved through the physical adsorption of the pores and the electrostatic interaction of the surface mineralization layer.

[0161] Loading method: The gradient mineralized ceramic microspheres prepared in step 4 were immersed in a BMP solution (BMP dissolved in phosphate buffer at pH 7.4, concentration 0.1-0.5 mg / mL, preferably 0.3 mg / mL, using recombinant human bone morphogenetic protein-2, rhBMP-2) under vacuum at 4°C for 2-4 hours (preferably 3 hours), with a vacuum degree of -0.08 MPa, to allow the BMP solution to fully penetrate into the bilevel pores of the microspheres. Macropores (200-500 μm) provide rapid penetration channels for BMP, while micropores (5-50 μm) provide a high specific surface area for BMP adsorption. After immersion, the microspheres were removed, and excess liquid on the surface was gently absorbed with filter paper. They were then frozen at -80°C for 4-6 hours and freeze-dried (vacuum degree <10 Pa, temperature -50°C, time 48 hours) to remove moisture while maintaining the activity of BMP. After freeze-drying, BMP in the microspheres was fixed in the form of dry powder on the pore surface and mineralization layer. The BMP loading is calculated based on the final amount of material used, typically 50-200 μg BMP per gram of ceramic microspheres (preferably 100 μg / g), corresponding to a BMP content of 20-80 μg / g of material in the final product. The BMP loading and loading efficiency in the microspheres are determined by BCA protein quantification or enzyme-linked immunosorbent assay (ELISA).

[0162] The preferred scheme is as follows: when the BMP concentration is 0.3 mg / mL, the impregnation time is 3 hours, and the loading amount is 100 μg / g ceramic microspheres, the loading efficiency reaches 75-80%, the distribution uniformity of BMP in the microspheres is optimal, the activity retention rate is >90%, and the release curve exhibits ideal dual-phase characteristics.

[0163] Step 9: Mixing ceramic microspheres with organic matrix and material molding

[0164] This step involves mixing BMP-loaded ceramic microspheres with the cross-linked organic matrix paste prepared in step 7 to form the final plastic bone tissue repair material.

[0165] Mixing method: BMP-loaded ceramic microspheres were mixed with a cross-linked organic matrix paste at a volume ratio of 4:6. This ratio was based on the following considerations: the ceramic phase ratio (40%) provides sufficient mechanical support and osteoconductive function, while the organic matrix phase ratio (60%) ensures the material's plasticity and bioactivity. An excessively high ceramic ratio (>50%) will make the material too hard and lose its plasticity, while an excessively low ceramic ratio (<30%) will result in insufficient mechanical strength.

[0166] During mixing, use gentle manual or mechanical mixing methods (for mechanical mixing, use a planetary mixer at 30-50 rpm, preferably 40 rpm) to avoid excessive shearing that could damage the chitosan coating and mineralization layer on the surface of the ceramic microspheres. Control the mixing time to 5-10 minutes (preferably 7 minutes) to ensure uniform dispersion of the ceramic microspheres in the organic matrix. Confirm the mixing uniformity by visual and tactile inspection: there should be no obvious agglomeration of ceramic microspheres or separation of the organic matrix in the material. The mixed material should be a uniform grayish-white paste with good plasticity and adhesion, and can be extruded into shape using a syringe (5-10 mm inner diameter).

[0167] After the ceramic microspheres are mixed with the organic matrix, the octacalcium phosphate and hydroxyapatite in the ceramic microspheres slowly dissolve and release in the aqueous environment. and The released calcium ions participate in the continuous cross-linking of sodium alginate, forming a dual cross-linking mode of "exogenous cross-linking + endogenous cross-linking" with the exogenous strontium-barium-calcium ion cross-linking in step 7, thereby enhancing the cross-linking density and long-term stability. The calcium ion release rate of the ceramic phase is approximately 5-10 μg / (g·h), and the total amount of calcium ions released cumulatively within 48 hours reaches 3-5% of the theoretical solubility, which is sufficient to maintain the stability of the sodium alginate cross-linking network.

[0168] Plasticity verification: Take an appropriate amount of the mixed material (10-20 g), and knead, flatten, and shape it into different shapes (such as cylinders, strips, and blocks) by hand. The material deforms under external force but does not crack or disintegrate. Let the shaped material stand for 15-30 minutes (preferably 20 minutes). Due to the continuous ionic crosslinking and covalent crosslinking of oxidized chitosan, the material gradually solidifies, eventually forming a solid material with a certain strength. The solidified material retains the shape it was shaped and does not undergo significant shrinkage or expansion deformation (volume change rate <5%).

[0169] The compressive strength of the cured material was tested using a universal testing machine (compression rate 1 mm / min, sample size cylindrical, diameter 10 mm, height 15 mm), confirming it reached 18-28 MPa (preferably 23 MPa). Scanning electron microscopy revealed that the ceramic microspheres were uniformly dispersed in the organic matrix, with good interfacial bonding and no obvious gaps. A BMP release test (the material was immersed in simulated body fluid, kept at 37°C with constant temperature oscillation, and samples were taken periodically to determine the released BMP concentration using ELISA) confirmed that BMP is released in a dual-phase manner within the bipolar pore structure—30% is released initially (1-3 days) to rapidly initiate osteogenic formation, and the remaining 70% is released sustainably in the later stage (4-28 days) to maintain the osteogenic effect, conforming to a first-order kinetic release model.

[0170] The preferred scheme is as follows: when the volume ratio of ceramic phase to organic matrix phase is 4:6, the mixing speed is 40 rpm, the mixing time is 7 minutes, and the curing time is 20 minutes, the plasticity, post-curing strength, uniform dispersion of ceramic microspheres and BMP release curve of the material are all optimal, the compressive strength is 23±2 MPa, and the plastic operation time window is 25-30 minutes.

[0171] Thus, a bone tissue repair material with a three-phase ceramic composition, a bilevel porous structure, chitosan coating, a gradient mineralization layer, strontium-barium dual-ion crosslinking, oxidized chitosan covalent crosslinking, and hemostatic function has been prepared. This material can be used directly or packaged and refrigerated (3 months at 4℃, 12 months at -20℃).

[0172] To verify the technical effect of this embodiment, the following three experiments were conducted to compare and verify the advantages of the material of this invention over the prior art.

[0173] Experiment 1: In vivo stability test (anti-displacement effect of strontium-barium-calcium ternary ion crosslinking)

[0174] 1. Experimental Objective

[0175] To verify the stability improvement effect of strontium-barium-calcium ternary ion crosslinking (Sr:Ba:Ca = 4:3:3) compared with traditional calcium ion crosslinking in a simulated in vivo environment, the mechanical strength retention rate of the material in simulated body fluid was tested.

[0176] 2. Preparation of experimental samples

[0177] Four groups of sodium alginate-collagen composite material samples were prepared (excluding ceramic phases to highlight the influence of ionic crosslinking):

[0178] Sample A: Traditional calcium ion crosslinking (control group), using 100 mM CaCl2 crosslinking solution

[0179] Sample B: Sr-calcium biionic crosslinked (Sr:Ca = 5:5), total ion concentration 75 mM

[0180] Sample C: Barium-calcium dual ion crosslinked (Ba:Ca = 3:7), total ion concentration 75 mM

[0181] Sample D: Strontium-Barium-Calcium ternary crosslinked (Sr:Ba:Ca = 4:3:3, this invention), total ion concentration 75 mM

[0182] The concentration of sodium alginate (molecular weight 250,000 Da) in each sample group was 3%, the concentration of collagen was 4%, and the amount of oxidized chitosan (oxidation degree 20%) was 15% of the mass of sodium alginate. All samples were molded into cylindrical specimens (diameter 10 mm, height 15 mm), and the initial compressive strength was tested after crosslinking at room temperature for 30 minutes.

[0183] 3. Experimental conditions

[0184] The cross-linked sample was immersed in simulated body fluid (SBF, containing Na) + 142 mM, Mg 2+ The samples were placed in a 1.0 mM ion exchange medium (simulating a high ion exchange environment in vivo) and then placed in a 37°C constant temperature shaking incubator (60 rpm). The samples were taken out at 0, 1, 3, 7, 14 and 21 days, the surface moisture was blotted with filter paper and the compressive strength was tested immediately.

[0185] 4. Experimental Procedure

[0186] (1) Use a universal testing machine to test the compressive strength of the samples at each time point, with a compression rate of 1 mm / min, and record the yield stress value.

[0187] (2) Calculate the mechanical strength retention rate: Strength retention rate (%) = (compressive strength on day n / initial compressive strength) × 100%.

[0188] (3) Five parallel samples were tested for each group of samples, the average value was taken, and the standard deviation was calculated.

[0189] (4) The residual contents of Sr, Ba and Ca in the sample were determined by ICP-MS to analyze the ion loss.

[0190] (5) The storage modulus G' of the samples at different time points was tested by rheometer to evaluate the stability of the gel network.

[0191] 5. Experimental or test results

[0192] Table 1. Compressive strength and strength retention rate of samples with different crosslinking methods in simulated body fluid.

[0193]

[0194] Note: Data are expressed as mean ± standard deviation (n=5).

[0195] Table 2. Ion residue content in samples at different time points (ICP-MS analysis)

[0196]

[0197] Note: Ion residual rate = (Ion content on day n / Initial ion content) × 100%; "-" indicates that the sample does not contain this ion.

[0198] 6. Analysis and Summary

[0199] (1) Comparison of mechanical strength retention rate: The experimental results show that after immersion in simulated body fluid for 3 days, the mechanical strength retention rate of sample A (traditional calcium ion crosslinking) dropped sharply from 100% to 33.6%, and after 21 days it was only 6.1%, almost completely losing its mechanical properties. The retention rate of sample B (Sr-Ca dual ion crosslinking) was 51.3% after 3 days, which was an improvement over sample A but still not ideal. The retention rate of sample C (Ba-Ca dual ion crosslinking) was 65.5% after 3 days, which was better than samples A and B. The mechanical strength retention rate of sample D (Sr-Ba-Ca ternary ion crosslinking of the present invention, molar ratio 4:3:3) reached 81.4% after 3 days and was still 57.8% after 21 days, which was 48 percentage points higher than that of traditional calcium crosslinking (from 33.6% to 81.4%), verifying the significant advantage of the technical solution of the present invention in terms of in vivo stability.

[0200] (2) Ion Residual Analysis: ICP-MS data showed that calcium ions in sample A lost more than 60% within 3 days (residual rate 38.5%), confirming the rapid replacement of calcium ions by high concentrations of sodium and magnesium ions. The residual rates of strontium and barium ions in sample D after 3 days were 78.5% and 84.7%, respectively, significantly higher than the calcium ion residual rate in sample A (38.5%), demonstrating that the binding of strontium and barium ions with sodium alginate is more stable and has stronger resistance to replacement. In particular, the residual rate of barium ions was the highest (84.7%), verifying the anti-replacement effect of the large ionic radius and high coordination stability constant of barium ions. The synergistic use of ternary ions resulted in a total ion residual rate of 78.9%, providing an ionic basis for the long-term stability of the material.

[0201] (3) Verification of coordination chemistry principle: The experimental results are consistent with the predictions of coordination chemistry theory. The coordination bond formed between barium ions (ionic radius 135 pm) and sodium alginate has the highest stability, followed by strontium ions (118 pm), and calcium ions (100 pm) have the lowest stability. This invention uses a Sr:Ba:Ca ratio of 4:3:3, achieving an optimal balance between stability (contribution of barium ions), bioactivity (strontium ions promote bone formation), and gelation rate (regulation by calcium ions). Sample D achieved a strength retention rate of 95.6% after 1 day, higher than the 84.5% to 85.1% of other groups, indicating that the initial curing effect of ternary ion crosslinking is better and the gel network is denser.

[0202] (4) Clinical application significance: The material of this invention retains more than 80% of its mechanical strength after 3 days in a simulated body fluid environment, meeting the structural stability requirements of early bone defect repair (1-3 days after implantation). It can effectively withstand the physiological load of the implantation site and the ion replacement pressure of the tissue fluid environment. Compared with traditional calcium cross-linked materials that lose most of their strength within 3 days, the material of this invention can be stably maintained in vivo for 14-21 days, providing a sufficient time window for the bone regeneration process. It is particularly suitable for clinical situations where the in vivo stability of materials is extremely important, such as the repair of bone defects in the medullary cavity and the repair of bone in highly vascularized areas.

[0203] Experiment 2: Test of interfacial bonding strength of gradient mineralized layer

[0204] 1. Experimental Objective

[0205] The gradient biomimetic mineralization method of the present invention (first 0.5×SBF, then 1×SBF) improves the bonding strength between the mineralized layer and the matrix interface and the plastic operation resistance to spalling compared with the traditional one-step mineralization method.

[0206] 2. Preparation of experimental samples

[0207] Two groups of chitosan-coated ceramic microsphere samples were prepared:

[0208] Sample E (control group, traditional one-step mineralization): Chitosan-coated ceramic microspheres were directly immersed in a standard concentration of simulated body fluid (1×SBF) and soaked at 37°C for 36 hours to form a mineralization layer in one step.

[0209] Sample F (in this invention, gradient mineralization): Chitosan-coated ceramic microspheres were first immersed in 0.5×SBF at 37°C for 12 hours to form an inner mineralization layer, and then transferred to 1×SBF at 37°C for 24 hours to form an outer mineralization layer. The total mineralization time was 36 hours.

[0210] The ceramic microspheres in both groups of samples had the same composition (HA:OCP:Si-TCP = 3:2:5, bilevel porous structure, microsphere size 1.0-1.5 mm), and the chitosan coating thickness was 20 μm in both groups. Thirty microsphere samples were prepared from each group for subsequent testing.

[0211] 3. Experimental conditions

[0212] (1) Interface shear strength test: The mineralized layer was sheared at room temperature using a micromechanical testing system with a loading rate of 0.1 mm / min. The maximum shear stress at the interface peeling was recorded.

[0213] (2) Plasticity operation simulation test: The mineralized ceramic microspheres were mixed with organic matrix paste (sodium alginate-collagen-oxidized chitosan) at a volume ratio of 4:6 and standardized plasticity operation was carried out (kneading 10 times, each time applying pressure of 50 N; bending deformation 5 times, bending angle of 90°; syringe extrusion molding 2 times, inner diameter of 8 mm). After the operation was completed, the peeling of the mineralized layer was observed and statistically analyzed.

[0214] (3) Scanning electron microscopy observation: After selecting typical samples for freeze drying and gold sputtering, the microstructure and interface structure of the mineralized layer were observed using a scanning electron microscope (SEM, accelerating voltage 15 kV).

[0215] 4. Experimental Procedure

[0216] (1) The interfacial shear strength of the mineralized ceramic microspheres was tested using a micromechanical testing system. Fifteen samples were tested in each group, and the average value and standard deviation were taken.

[0217] (2) Mix another 15 microsphere samples with the organic matrix and perform standardized plastic operation. After the operation, immerse the material in deionized water and rinse gently. Collect the detached mineralized layer fragments, dry and weigh them, and calculate the peeling rate = (mass of detached fragments / total mass of initial mineralized layer) × 100%. The mass of the initial mineralized layer is determined by the mass difference of microspheres before and after mineralization.

[0218] (3) Use energy dispersive spectroscopy (EDS) to determine the calcium-to-phosphorus ratio (Ca / P) at different depths of the mineralized layer and evaluate the compositional gradient of the mineralized layer.

[0219] 5. Experimental or test results

[0220] Table 3 Comparison of interfacial bonding strength and anti-stripping performance of different mineralization methods

[0221]

[0222] Note: Data are expressed as mean ± standard deviation (n=15). The density of the mineralized layer was calculated by three-dimensional reconstruction using micro-CT.

[0223] Figure 1 This is a bar chart comparing the shear strength at the mineralized layer interface.

[0224] Figure 2 This is a bar chart comparing the spalling rate of the mineralized layer after plastic manipulation.

[0225] Table 4. EDS energy dispersive spectroscopy analysis: Changes in calcium-to-phosphorus ratio at different depths of the mineralized layer.

[0226]

[0227] Note: The Ca / P molar ratio is calculated using the formula: Ca / P = (Ca at% / P at%). The theoretical stoichiometry for hydroxyapatite is Ca / P = 1.67.

[0228] 6. Analysis and Summary

[0229] (1) Significantly improved interfacial bonding strength: Experimental results show that the gradient mineralization method of the present invention (sample F) improves the interfacial shear strength from 2.8 MPa to 8.5 MPa, an increase of 3.0 times, compared with the traditional one-step mineralization (sample E). This improvement is attributed to the inner mineralization layer formed by gradient mineralization penetrating into the pores of the chitosan layer, forming a composite bonding mode of physical anchoring and chemical bonding, while the interfacial bonding of traditional one-step mineralization mainly relies on physical adsorption, resulting in weak bonding force.

[0230] (2) Significantly improved resistance to spalling during plastic operation: After standardized plastic operation (kneading, bending, extrusion), the spalling rate of the mineralized layer in sample F was only 4.2%, which was 30.5 percentage points lower than that of sample E (34.7%), representing a reduction of 88%. This demonstrates that the gradient mineralized layer has excellent crack resistance and spalling resistance when subjected to shear force and bending deformation during plastic operation, overcoming the technical bottleneck of easy detachment of traditional one-step mineralized coatings during plastic operation.

[0231] (3) Verification of gradient composition structure: EDS energy dispersive spectroscopy analysis showed that the Ca / P ratio of the mineralized layer of sample E remained at 1.63-1.65 from the surface layer to the interface layer, with no obvious gradient and uniform composition. The Ca / P ratio of the mineralized layer of sample F gradually decreased from 1.66 at the surface layer to 1.52 at the interface layer, forming an obvious compositional gradient. The inner mineralized layer with a low Ca / P ratio has higher reactivity and flexibility, and is more likely to form chemical bonds with the chitosan coating layer. At the same time, it acts as a buffer layer to reduce stress concentration between the outer dense mineralized layer and the flexible matrix.

[0232] Experiment 3: Rapid Hemostasis Function Test

[0233] 1. Experimental Objective

[0234] The study aimed to verify the rapid hemostatic effect of the hemostatic microspheres (encapsulating thrombin and tranexamic acid) introduced in the material of this invention on bleeding at bone defect sites, and to evaluate the hemostasis time, bleeding volume, and material stability in the blood environment.

[0235] 2. Preparation of experimental samples

[0236] Two sets of complete bone tissue repair material samples were prepared:

[0237] Sample G (control group, without hemostatic microspheres): ceramic microspheres (coated with chitosan and subjected to gradient mineralization, loaded with BMP) and organic matrix (calf bone collagen, sodium alginate, oxidized chitosan, Sr-Ba-Ca ternary ion crosslinking) were mixed at a volume ratio of 4:6, without the addition of hemostatic microspheres.

[0238] Sample H (of the present invention, containing hemostatic microspheres): ceramic microspheres and organic matrix (containing 5.5% hemostatic microspheres, with thrombin and tranexamic acid encapsulated within the microspheres) were mixed at a volume ratio of 4:6.

[0239] The other components of the two groups of samples were identical. Each group of samples was molded into a cylindrical specimen (10 mm in diameter and 10 mm in height) and weighed approximately 1.0 g.

[0240] 3. Experimental conditions

[0241] (1) In vitro coagulation time test: Fresh anticoagulated rabbit blood (purchased from the experimental animal center, with normal coagulation function) was used and preheated in a constant temperature water bath at 37℃. 0.5 mL of anticoagulated rabbit blood was added to 0.05 mL of 0.2 M calcium chloride solution to initiate coagulation. The material sample (0.5 g) was added immediately, and the time from the addition of calcium chloride to complete blood coagulation (coagulation time) was recorded. Criteria for judging blood coagulation: When the test tube is tilted at 45°, the blood is considered to have stopped flowing when coagulation is considered complete.

[0242] (2) Measurement of bleeding volume (in vitro simulation): An in vitro bleeding model was established. Fresh heparinized rabbit blood was pumped through a simulated bone defect cavity (a cylindrical cavity with an inner diameter of 15 mm and a depth of 20 mm) with a constant flow rate (2 mL / min, simulating the flow rate of bone marrow cavity bleeding) using a peristaltic pump. The outflowing blood was collected, and the total bleeding volume and the time to achieve hemostasis were measured (hemostasis was considered to have occurred when the bleeding volume dropped to below 0.1 mL / min).

[0243] (3) Animal model hemostasis test: A femoral medullary cavity burr hole hemorrhage model was established using New Zealand white rabbits (weight 2.5-3.0 kg, n=12). Under general anesthesia and aseptic conditions, the lateral femur was incised, and an 8 mm diameter drill bit was used to drill a hole in the middle of the femoral shaft to the medullary cavity (depth 15 mm), causing medullary cavity hemorrhage. Immediately, a material sample (weighing approximately 1.5 g) was filled into the defect cavity, and the time from the start of material filling to complete cessation of bleeding (hemostasis time) was recorded. Hemostasis judgment criteria: no fresh blood seepage from the surface of the defect site for 30 seconds was considered hemostasis. At the same time, the total blood loss was recorded (calculated by weighing the gauze that absorbed blood during the operation).

[0244] 4. Experimental Procedure

[0245] (1) In vitro coagulation time test: 10 parallel samples were tested for each group of samples. The coagulation time was accurately recorded using a stopwatch, and the average value and standard deviation were taken. The blank control group (only anticoagulant + calcium chloride was added, without any other materials) was tested simultaneously.

[0246] (2) Measurement of bleeding volume: Each group of samples was tested 6 times, and the time to achieve hemostasis and the total bleeding volume were recorded. The mean and standard deviation were calculated.

[0247] (3) Animal model hemostasis test: Rabbits were randomly divided into two groups of 6 each. Bone defects were filled with sample G and sample H respectively, and the hemostasis time and bleeding volume were recorded. The wound healing and infection rate were assessed after 7 days of postoperative observation.

[0248] (4) Hemostatic microsphere release kinetics test: The organic matrix (0.5 g) containing hemostatic microspheres was immersed in a simulated body fluid containing calcium ions (containing 2.5 mM CaCl₂). The sample was collected in whole blood (simulated blood calcium concentration) and oscillated at 37°C for 1, 3, 5, 10, 15 and 30 minutes. The thrombin activity in the released solution was measured using a thrombin activity assay kit, and the tranexamic acid concentration was measured by HPLC. The release curve was plotted.

[0249] (5) Stability test of the material in the blood environment: The two groups of samples were immersed in fresh rabbit blood (the blood volume was 10 times the material volume), and left to stand at 37°C. The samples were taken out after 10, 30, 60 and 120 minutes, respectively. The surface blood was gently rinsed with physiological saline. The dimensional change rate (expansion rate or shrinkage rate) and shape retention of the samples were measured to evaluate the stability of the material in the blood environment.

[0250] 5. Experimental or test results

[0251] Table 5 Comparison of in vitro coagulation time and bleeding volume

[0252]

[0253] Note: Data are expressed as mean ± standard deviation. Accelerated clotting rate = Clotting time of blank control / Clotting time of sample. Reduction rate of bleeding volume = (Bleeding volume of sample G - Bleeding volume of sample H) / Bleeding volume of sample G × 100%.

[0254] Table 6 Evaluation of hemostasis effect in animal models (New Zealand white rabbit femoral medullary hemorrhage model)

[0255]

[0256] Note: Hemostasis time and bleeding volume data are expressed as mean ± standard deviation. Wound healing score: 0 is the worst (excessive exudation, redness, necrosis), and 10 is the best (dry, no redness, good epithelialization).

[0257] Figure 3 This is a bar chart comparing in vitro clotting times.

[0258] Figure 4 A bar chart comparing hemostasis time in animal models.

[0259] Figure 5 Release kinetics curve of hemostatic active ingredient.

[0260] Table 7. Stability of materials in a blood environment (dimensional changes and shape retention)

[0261]

[0262] Note: A positive rate of change indicates expansion, and a negative rate indicates contraction. Shape retention score: 0 points indicates complete disintegration and loss of shape, and 5 points indicates complete shape retention.

[0263] 6. Analysis and Summary

[0264] (1) Significant rapid hemostasis effect: In vitro experiments showed that sample H containing hemostatic microspheres reduced the clotting time from 285 seconds in the blank control to 52 seconds, accelerating clotting by 5.48 times; compared with sample G without hemostatic microspheres (clotting time 238 seconds), sample H shortened the time by 186 seconds (78%). Animal model experiments further verified that the hemostasis time of sample H was 2.8 minutes, which was 10.9 minutes (79.7%) shorter than the 13.7 minutes of sample G, close to the design goal of "reducing the hemostasis time from 10-15 minutes to 2-3 minutes" expected in the technical plan. In terms of bleeding volume, sample H reduced the bleeding volume in the animal model from 22.3 mL to 4.1 mL, a reduction of 81.6%, exceeding the expected effect of "reducing the bleeding volume by 80%" in the technical plan.

[0265] (2) Validation of the synergistic hemostatic mechanism of the two active ingredients: The release kinetics test of the hemostatic microspheres showed that thrombin and tranexamic acid were rapidly released in the blood environment, with release rates exceeding 70% within 5 minutes and exceeding 90% within 15 minutes. The rapid release of thrombin (28.5% released in 1 minute) enabled immediate hemostasis, while the sustained release of tranexamic acid (97.2% cumulative release in 30 minutes) maintained the hemostatic effect and prevented fibrinolysis. The release curves of both exhibited an ideal dual-phase characteristic of "rapid initial release + sustained mid-term release", validating the sustained-release functional design of the sodium alginate-gelatin microspheres.

[0266] (3) Improved stability of the material in a blood environment: Blood environment stability tests showed that sample G without hemostatic microspheres rapidly absorbed water and swelled in blood, with a volume expansion of 85.3% within 60 minutes, and the shape retention score dropped to 1.3 points, with severe disintegration at the material edges. Sample H containing hemostatic microspheres only expanded by 20.1% in volume within 60 minutes in blood, maintained a shape retention score of 4.2 points, and the material shape remained intact without significant disintegration. This difference is attributed to the rapid hemostasis function of the hemostatic microspheres reducing the contact time between the material and flowing blood, while the double cross-linked network formed by Sr-Ba-Ca ternary ionic cross-linking and oxidized chitosan covalent cross-linking provides structural stability against blood erosion.

[0267] (4) Reduced infection risk and improved wound healing quality: Postoperative observation of animal models showed that 2 rabbits (33.3%) in sample G developed postoperative infection (wound redness and swelling, increased exudation), with a wound healing score of only 6.2 points; no infection occurred in sample H (0%), and the wound healing score was as high as 9.1 points. The rapid hemostasis function shortened the operation time and bleeding exposure time, reduced bacterial contamination and infection risk, and at the same time, the reduction of hematoma improved the local microenvironment and promoted wound healing.

[0268] (5) Clinical value in preventing material displacement: In the animal model, 3 rabbits (50%) in sample G experienced material displacement within the bone defect cavity (assessed by postoperative X-ray imaging), affecting the repair effect; no displacement occurred in sample H (0%). This demonstrates that the rapid hemostasis function effectively solved the problems of material expansion, disintegration, and displacement caused by bleeding at the bone defect site, improved the early stability of the implant, ensured close contact between the material and the bone defect cavity, and provided good initial conditions for subsequent bone regeneration.

[0269] The embodiments of the present invention have been described above. However, the embodiments are not limited to the specific implementation methods described above. The specific implementation methods described above are merely illustrative and not restrictive. Those skilled in the art can make more equivalent embodiments under the guidance of the present embodiments, and all of them are within the protection scope of the present embodiments.

Claims

1. A bone tissue repair material, characterized in that, Including ceramic phase and organic matrix phase; The ceramic phase is a porous microsphere, composed of hydroxyapatite, octacalcium phosphate, and silicon-doped β-tricalcium phosphate in a mass ratio of 3:2:5, wherein the silicon content of the silicon-doped β-tricalcium phosphate is 1.5-3.5 wt%. The porous microspheres have a dual-level pore structure, with macropores having a diameter of 200-500 μm and micropores having a diameter of 5-50 μm, resulting in a porosity of 60-80%. The surface of the porous microspheres is sequentially coated with a chitosan layer and a gradient mineralization layer. The chitosan layer has a thickness of 10-30 μm, and the gradient mineralization layer includes an inner mineralization layer and an outer mineralization layer. The inner mineralization layer has a thickness of 2-5 μm, and the total mineralization layer thickness is 8-15 μm. The organic matrix phase includes calf bone collagen, sodium alginate, oxidized chitosan, and hemostatic microspheres; The sodium alginate is cross-linked with strontium ions, barium ions, and calcium ions, with a molar ratio of Sr:Ba:Ca of 4:3:3; the oxidized chitosan has an oxidation degree of 15-25% and is used in amounts of 10-20% of the sodium alginate mass; the hemostatic microspheres are sodium alginate-gelatin composite microspheres with a diameter of 50-150 μm, encapsulating thrombin and tranexamic acid inside; The porous microspheres of the ceramic phase are loaded with bone morphogenetic proteins. The volume ratio of the ceramic phase to the organic matrix phase is 4:

6.

2. The bone tissue repair material according to claim 1, characterized in that, The porous microspheres are prepared by a pore-forming agent method, wherein the pore-forming agent is polymethyl methacrylate microspheres, including microspheres with a particle size of 200-500 μm and microspheres with a particle size of 5-50 μm, and the mass ratio of the two particle sizes is 7:3; the particle size of the porous microspheres is 0.5-2 mm.

3. The bone tissue repair material according to claim 1, characterized in that, The chitosan has a molecular weight of 100,000-500,000 Da, a degree of deacetylation ≥85%, and is used in an amount of 3-8% of the mass of the ceramic microspheres.

4. The bone tissue repair material according to claim 1, characterized in that, The gradient mineralization layer is formed through a two-step biomimetic mineralization process: first, it is soaked in a simulated body fluid at 0.5 times the concentration at 37°C for 12 hours to form an inner mineralization layer, and then soaked in a simulated body fluid at a standard concentration at 37°C for 24 hours to form an outer mineralization layer; the main component of the gradient mineralization layer is hydroxyapatite, with a calcium-to-phosphorus ratio of 1.60-1.

67.

5. The bone tissue repair material according to claim 1, characterized in that, The sodium alginate has a molecular weight of 200,000-300,000 Da and a concentration of 2-4%.

6. The bone tissue repair material according to claim 1, characterized in that, The aldehyde groups of the oxidized chitosan form Schiff base covalent crosslinks with the amino groups of collagen and sodium alginate, and form a double crosslinked network with strontium-barium-calcium ions.

7. The bone tissue repair material according to claim 1, characterized in that, The thrombin activity is ≥1000 IU / mg, and the amount used is 1-3% of the gelatin mass; the amount of tranexamic acid used is 5-10% of the gelatin mass; and the amount of hemostatic microspheres used is 3-8% of the total organic matrix mass.

8. A method for preparing the bone tissue repair material according to claim 1, characterized in that, Includes the following steps: Step 1: Silicon-doped tricalcium β-phosphate was prepared using the sol-gel method, with a silicon content of 1.5-3.5 wt%. Step 2: Silicon-doped β-tricalcium phosphate, hydroxyapatite and octacalcium phosphate are mixed in a mass ratio of 5:3:2, and polymethyl methacrylate microspheres are added as pore-forming agents. After molding, the mixture is sintered at 1100-1200℃ to form ceramic microspheres with a bilevel pore structure. Step 3: Immerse the ceramic microspheres in a chitosan solution for coating, with a coating thickness of 10-30 μm; Step 4: Perform gradient biomimetic mineralization on the coated ceramic microspheres: first, soak them in a simulated body fluid at 0.5 times the concentration at 37°C for 12 hours to form an inner mineralization layer, and then soak them in a simulated body fluid at standard concentration at 37°C for 24 hours to form an outer mineralization layer. The total mineralization layer thickness is 8-15 μm. Step 5: Prepare sodium alginate-gelatin composite microspheres, with thrombin and tranexamic acid encapsulated within the microspheres; Step 6: Mix calf bone collagen, sodium alginate solution, oxidized chitosan and hemostatic microspheres to prepare an organic matrix paste; Step 7: Add strontium ions, barium ions, and calcium ions to the organic matrix paste, with a molar ratio of Sr:Ba:Ca of 4:3:3, and perform composite cross-linking; Step 8: Load bone morphogenetic proteins into the pores of the treated ceramic microspheres; Step 9: Mix ceramic microspheres loaded with bone morphogenetic proteins with cross-linked organic matrix paste at a volume ratio of 4:6 to form a plastic composite material.

9. The preparation method according to claim 7, characterized in that, In step 7, a mixed ion crosslinking solution is prepared by dissolving strontium chloride, barium chloride, and calcium chloride in deionized water at a molar ratio of 4:3:3 to prepare a crosslinking solution with a total divalent cation concentration of 50-100 mM. The amount of crosslinking solution added is calculated based on the concentration of sodium alginate, and the total molar amount of divalent ions is 0.4-0.6 times the molar amount of sodium alginate monomer.

10. The preparation method according to claim 7, characterized in that, In step 4, the composition of the 0.5-fold concentration simulated body fluid is as follows: NaCl 4.0 g / L, NaHCO3 0.175 g / L, KCl 0.112 g / L, K2HPO4·3H2O 0.114 g / L, MgCl2·6H2O 0.1525 g / L, CaCl2 0.1835 g / L, Na2SO4 0.0355 g / L, pH 7.4; The standard concentration of simulated body fluid consists of: NaCl 8.0 g / L, NaHCO3 0.35 g / L, KCl 0.224 g / L, K2HPO4·3H2O 0.228 g / L, MgCl2·6H2O 0.305 g / L, CaCl2 0.367 g / L, Na2SO4 0.071 g / L, and pH 7.4.

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