Bone adhesive with hemostatic and osteoinductive functions and use thereof

Abstract

This application relates to the field of bone repair medical materials technology, and more particularly to a bone adhesive with hemostatic and osteogenic functions and its application. The bone adhesive comprises a solid phase component and a liquid phase component; the solid phase component, by mass, comprises 85-90 parts of calcium phosphate matrix, 8-12 parts of α-amino acid compound, and 1-10 parts of silicon-based active glass; the α-amino acid compound contains at least one of phosphorylated hydroxyl groups and free hydroxyl groups; the liquid-to-solid ratio of the liquid phase component to the solid phase component is 0.3-1 mL / g, and after mixing the solid and liquid phase components, the hydration product of the calcium phosphate matrix is ​​permealuminate. The solution provided by this application can simultaneously meet multiple requirements such as hemostasis in a blood environment, wet-state bonding and fixation, controlled degradation, and osteogenic induction.

Description

Technical Field

[0001] This application relates to the field of medical materials for bone repair, and in particular to a bone adhesive with hemostatic and bone-promoting functions and its application. Background Technology

[0002] In clinical settings such as emergency bone trauma treatment and orthopedic orthopedic surgery / tumor resection, three common clinical challenges are currently faced: difficulty in sealing diffuse bleeding in cancellous bone, difficulty in fixing comminuted fracture surfaces, and slow bone regeneration in defect areas. The core requirement for bone trauma and bone defect repair is to achieve the synergistic unity of interface integration, initial fixation, immediate hemostasis, and long-term bone regeneration. However, a single material system is difficult to adapt to the multiple functional requirements of complex wounds simultaneously.

[0003] For example, the invention patent application with application number 202311759659.0 discloses a bone adhesive that relies on an aldehyde cross-linking mechanism to achieve tissue adhesion. However, free proteins in the blood will prematurely consume the aldehyde functional groups, resulting in insufficient wet adhesion stability. Furthermore, it lacks an active coagulation-promoting function design, leading to low hemostasis efficiency. The overall material degradation rate is too fast to provide long-term stable mechanical support for bone regeneration, making it difficult to simultaneously meet the multiple needs of hemostasis, fixation, and osteogenic formation. The invention patent application with application number 202410060864.6 discloses a non-weight-bearing bone adhesive that uses calcium phosphate extract to provide a calcium-rich environment for curing, which can reduce the proportion of tetracalcium phosphate used and reduce the risk of pH fluctuations. However, it can only achieve basic adhesion in clean wounds without significant bleeding. It lacks an active hemostasis function design and a multi-component synergistic osteogenic regulation mechanism. The mechanical support cycle and degradation rate are difficult to match the rhythm of bone regeneration, and it also cannot achieve the synergistic unity of multi-dimensional functions.

[0004] Current clinical protocols cannot simultaneously address the three core requirements of rapid hemostasis in a blood environment, high-strength adhesion and fixation on wet bone surfaces, and biodegradable bone regeneration. Developing novel bone repair materials that combine these properties has become an urgent research direction in the field of orthopedic biomaterials. Summary of the Invention

[0005] To overcome the problems existing in related technologies, this application provides a bone adhesive with hemostatic and osteogenic functions, which can simultaneously meet multiple needs such as hemostasis in a blood environment, wet bonding and fixation, controlled degradation and osteogenic induction.

[0006] The first aspect of this application provides a bone adhesive with hemostatic and bone-promoting functions, comprising a solid phase component and a liquid phase component; The solid phase component, by mass parts, includes 85-90 parts of calcium phosphate matrix, 8-12 parts of α-amino acid compounds, and 1-10 parts of silicon-based active glass. The α-amino acid compounds contain at least one of phosphorylated hydroxyl groups and free hydroxyl groups; The liquid-to-solid ratio of the liquid phase component to the solid phase component is 0.3-1 mL / g, and after the solid phase component and the liquid phase component are mixed, the hydration product of the calcium phosphate matrix is ​​transparent calcium phosphate.

[0007] In some embodiments, the calcium phosphate matrix includes a first matrix and a second matrix; The first matrix is ​​selected from at least one of β-tricalcium phosphate, α-tricalcium phosphate, tetracalcium phosphate, and octacalcium phosphate; The second matrix is ​​selected from at least one of anhydrous calcium dihydrogen phosphate and calcium dihydrogen phosphate dihydrate; The mass ratio of the first matrix to the second matrix is ​​(4-8):(1-5).

[0008] In some embodiments, the mass ratio of the first substrate to the second substrate is 2:1.

[0009] In some embodiments, the α-amino acid compound is selected from at least one of L0-phosphoserine, O-phospho-L-threonine, and L-serine; The silicon-based active glass is selected from at least one of mesoporous bioactive glass, strontium-doped mesoporous bioactive glass, copper-doped mesoporous bioactive glass, and zinc-doped mesoporous bioactive glass.

[0010] In some embodiments, if the silicon-based active glass is a mesoporous bioactive glass, the mesoporous bioactive glass is prepared by the following method: (1) Dissolve hexadecyltrimethylammonium bromide in deionized water, add ethyl acetate at 40°C, and stir at a constant temperature until an emulsion system is formed; (2) Add alkaline solution dropwise to the emulsion system and stir at a constant temperature to form an alkaline catalytic reaction system; (3) Tetraethyl orthosilicate, triethyl phosphate, and calcium nitrate tetrahydrate aqueous solution were added dropwise to the alkaline catalytic reaction system, and the reaction was carried out under constant temperature and stirring to obtain mesoporous bioactive glass gel. (4) After aging the obtained gel, it is washed with deionized water and anhydrous ethanol alternately to remove residual organic components, and then dried. (5) The dried gel was calcined, cooled, ground and sieved to obtain mesoporous bioactive glass powder, which was then sealed and stored for later use.

[0011] In some embodiments, in step (1), the mass ratio of deionized water to hexadecyltrimethylammonium bromide is 1:1; the volume ratio of ethyl acetate to deionized water is 3:10; and the constant temperature stirring time is 30 min. In step (2), the alkaline solution is 2 mol / L ammonia water, the volume ratio of ammonia water to deionized water is 21:100, and the constant temperature stirring time after dropwise addition is 15 min. In step (3), the volume ratio of tetraethyl orthosilicate to deionized water is 7:50, and the mixture is stirred at a constant temperature for 30 min after being added dropwise; the volume ratio of triethyl phosphate to deionized water is 7:500, and the mixture is stirred at a constant temperature for 30 min after being added dropwise; the calcium nitrate tetrahydrate aqueous solution is prepared by dissolving 18 parts by mass of calcium nitrate tetrahydrate in 25 parts by mass of deionized water, and the mixture is stirred at a constant temperature for 3 h after being added dropwise. In step (4), the aging temperature is 37°C, the aging time is 3 days, the washing is 3 times, and the drying temperature is 60°C. In step (5), the calcination temperature is 650℃.

[0012] In some embodiments, the solid phase component, by mass, includes 90 parts of calcium phosphate matrix, 10 parts of α-amino acid compound, and 2 parts of silicon-based active glass. The liquid-to-solid ratio is 0.6 mL / g.

[0013] In some embodiments, the solid phase component further includes 0.5-2 parts of an additive, the additive being selected from at least one of antibiotics, anti-inflammatory drugs, and antimicrobial peptides; The liquid phase component is selected from any one of physiological saline, phosphate buffer, and biomimetic body fluid.

[0014] In some embodiments, the bone adhesive is selected from any one of injectable paste, pre-cured block, or porous sponge.

[0015] The second aspect of this application provides the application of the above-mentioned bone adhesive with hemostatic and osteogenic functions in the preparation of hemostatic materials for bone wounds, adhesive materials for non-load-bearing bone fractures, hemostatic sealing materials for medullary cavities, and maxillofacial bone repair materials.

[0016] Compared with the prior art, the bone adhesive with hemostatic and bone-promoting functions provided in this application has the following beneficial effects: (1) This bone adhesive can achieve multi-pathway synergistic hemostasis. The α-amino acid compounds containing phosphorylated hydroxyl or free hydroxyl groups can specifically bind to phosphatidylserine receptors on the surface of blood cells, actively adhering to and enriching blood cells. The high specific surface area of ​​the silicon-based active glass can physically adsorb plasma proteins, increasing the concentration of local coagulation factors at the wound site. At the same time, the released calcium and silicon active ions can activate the intrinsic coagulation pathway. The calcium ions released by the calcium phosphate matrix can participate in the coagulation cascade reaction and accelerate the formation of fibrin network. The three components work synergistically to achieve multi-pathway coagulation. Related experimental verification shows that the hemolysis rate of this bone adhesive is less than 5%, which meets the blood compatibility safety standards for medical implant materials. The red blood cell adhesion rate is 33.34%, the platelet adhesion rate is 69.09%, the whole blood coagulation index is about 11%, and the activation time of partial thromboplastin is about 28s. At the same time, there is no risk of zeolite-type materials causing exothermic burns to surrounding tissues during hydration. It can meet the hemostasis needs of complex bone wounds.

[0017] (2) The physicochemical and mechanical properties of this bone adhesive are suitable for clinical needs. The α-amino acid in it acts as an organic crosslinking agent, forming calcium bridge coordination bonds with calcium ions on the surface of calcium phosphate particles through phosphate or hydroxyl groups, and forming intermolecular hydrogen bonds through its own amino and carboxyl groups. This dual action can effectively improve the interfacial bonding force of inorganic particles and optimize the bonding strength and mechanical properties after curing. The silicon-based active glass can fill the pores of hydration products, further optimizing the compressive strength of the cured body. The combination of the three components can regulate the hydration reaction rate, ultimately generating a permeable calcium phosphate product with a good match between the degradation rate and new bone growth. Experimental results show that the initial setting time of this material is about 5 minutes and the final setting time is about 10 minutes, which is suitable for the rhythm of clinical operation, leaving sufficient window for intraoperative shaping and application. The curing process is not prone to collapse and has no obvious exothermic effect. The cured body, after α-amino acid crosslinking and pore filling with silicon-based active glass, exhibits a compressive strength of approximately 30 MPa and a wet-state bone interface adhesion strength of approximately 1.2 MPa. It possesses the mechanical support strength suitable for non-load-bearing bone repair and good wet-state bone interface adhesion, providing immediate mechanical fixation for non-load-bearing bone sections or bone-implant interfaces. It meets the fixation needs of most non-load-bearing bone repair scenarios without requiring additional auxiliary fixation devices. Regarding degradation rate, the permealuminate product generated by the hydration reaction degrades to approximately 69.92% of its original mass 28 days after implantation, showing a high degree of matching with the natural new bone growth rate. It can simultaneously provide support space for new bone growth, minimizing the risk of premature degradation and loss of mechanical support, and reducing the likelihood of foreign body residue affecting the bone integration process, thus meeting the operational and prognostic requirements of clinical surgery.

[0018] (3) The bone adhesive exhibits good biocompatibility. The calcium phosphate matrix is ​​a natural inorganic component of bone, α-amino acids are endogenous substances metabolized by the human body, and silicon-based active glass is a commonly used bone repair material in clinical practice. It has virtually no toxic monomer residues and low immunogenicity. At the same time, α-amino acids can provide cell adhesion sites, which is beneficial for cell adhesion and proliferation. Experiments show that it has no significant cytotoxicity to bone marrow mesenchymal stem cells and human umbilical vein endothelial cells. The cell scratch test shows a healing rate of over 75% after 2 days. It is not prone to significant chronic inflammatory reactions after implantation, which can provide a favorable cellular microenvironment for bone repair and help reduce the risk of complications such as foreign body inflammation and bone cement implantation syndrome that are easily caused by traditional bone wax and polymethyl methacrylate bone cement.

[0019] (4) This bone adhesive can induce osteogenic differentiation in multiple dimensions, with a prominent bone integration effect. The silicon ions continuously released by the silicon-based active glass can upregulate the expression of osteogenic genes and induce bone marrow mesenchymal stem cells to differentiate into osteogenic genes. The calcium and phosphorus ions released by the degradation of the calcium phosphate matrix can directly provide raw materials for new bone mineralization. Serine and threonine in α-amino acids are components of bone matrix collagen and can participate in bone matrix synthesis. With the synergistic effect of the three components, a relatively stable osteogenic induction effect can be achieved without the need for additional loading of exogenous osteogenic factors. At the cellular level, this composite system can increase alkaline phosphatase activity and the amount of mineralized nodules formed. At the molecular level, it can upregulate the expression of osteogenic genes such as RUNX2, COL-1, OCN, and OPN throughout the bone repair cycle, inducing osteogenic differentiation. Experiments in a rat femoral defect model showed that basic repair of the defect could be achieved within 12 weeks of implantation. The newly formed bone tissue was dense and closely integrated with the host bone. The overall bone integration efficiency was superior to that of traditional inactive calcium phosphate bone cement, bone wax, and other materials, which helps to shorten the bone repair cycle and improve the prognosis. Attached Figure Description

[0020] The above and other objects, features and advantages of this application will become more apparent from the more detailed description of exemplary embodiments thereof in conjunction with the accompanying drawings, wherein the same reference numerals generally represent the same components in the exemplary embodiments thereof.

[0021] Figure 1 These are graphs showing the compressive strength variations in Examples 1-15 of this application; Figure 2 This is a comparison diagram of the wet bonding strength of Embodiment 5 and Embodiment 11 of this application; Figure 3 This is a graph showing the setting time of the bone adhesive in different liquid phase environments according to Example 11 of this application; Figure 4 These are the XRD patterns of Embodiments 5 and 11 of this application; Figure 5This is the microstructure of the cured bone adhesive of Example 11 of this application; Figure 6 This is a SEM image of the interface between the bone adhesive and bovine bone in Embodiment 11 of this application; Figure 7 These are the in vitro degradation curves of Examples 5 and 11 of this application; Figure 8 These are the blood cell adhesion rates and SEM images of Examples 10-15 of this application; Figure 9 These are the hemolysis rate test results of Example 11 of this application at different concentrations; Figure 10 It is the whole blood coagulation index of Examples 10-15 of this application; Figure 11 These are activation partial thromboplastin time graphs for different amounts of BG added in Examples 10-15 of this application; Figure 12 These are the proliferation activity test results of different cells from Examples 5 and 11 of this application; Figure 13 These are the cell scratch test results from Examples 5 and 11 of this application; Figure 14 These are the ALP activity and mineralized nodule test results from Example 11 of this application; Figure 15 This refers to the expression levels of osteogenic-related genes at different culture times in Example 11 of this application; Figure 16 This is the process of constructing and filling a rat femoral defect model as shown in the embodiments of this application; Figure 17 These are two-dimensional sagittal section images taken 12 weeks post-operation for Examples 5 and 11; Figure 18 These are three-dimensional reconstruction images of Examples 5 and 11 12 weeks post-surgery; Figure 19 These are the results of quantitative analysis of bone morphometrics in Examples 5 and 11. Detailed Implementation

[0022] Preferred embodiments of the present application will now be described in more detail with reference to the accompanying drawings. While preferred embodiments of the present application are shown in the drawings, it should be understood that the present application may be implemented in various forms and should not be limited to the embodiments set forth herein. Rather, these embodiments are provided to make the present application more thorough and complete, and to fully convey the scope of the present application to those skilled in the art.

[0023] The terminology used in this application is for the purpose of describing particular embodiments only and is not intended to be limiting of the application. The singular forms “a,” “the,” and “the” used in this application and the appended claims are also intended to include the plural forms unless the context clearly indicates otherwise. It should also be understood that the term “and / or” as used herein refers to and includes any or all possible combinations of one or more of the associated listed items.

[0024] It should be understood that although the terms "first," "second," "third," etc., may be used in this application to describe various information, this information should not be limited to these terms. These terms are only used to distinguish information of the same type from one another. For example, without departing from the scope of this application, first information may also be referred to as second information, and similarly, second information may also be referred to as first information. Thus, a feature defined as "first" or "second" may explicitly or implicitly include one or more of that feature. In the description of this application, "multiple" means two or more, unless otherwise explicitly specified.

[0025] In clinical settings such as emergency bone trauma treatment, orthopedic correction, and bone tumor resection, three common clinical challenges currently facing the field of bone repair are: difficulty in sealing diffuse bleeding from cancellous bone leading to unclear surgical fields and prolonged operation time; poor adaptability of irregular fracture surfaces in comminuted fractures making immediate stable fixation difficult; and disruption of the osteogenic microenvironment in the defect area resulting in slow bone regeneration. The core requirement for bone trauma and bone defect repair lies in achieving a synergistic unity of interface integration, initial fixation, immediate hemostasis, and long-term bone regeneration. However, existing single-material or simple compound systems are insufficient to simultaneously meet the multiple functional requirements of complex bleeding wounds.

[0026] This application provides a bone adhesive with hemostatic and osteogenic functions. Through the synergistic compounding of calcium phosphate matrix, α-amino acid compounds and silicon-based active glass, it can stably hydrate in a blood environment to generate a permeable calcium phosphate solidification system. It has multiple advantages such as immediate hemostasis, high-strength adhesion on wet surfaces, biodegradability and synergistic osteogenic function, which can solve the technical pain point that existing bone repair materials cannot meet the triple clinical needs at the same time.

[0027] The technical solutions of the embodiments of this application are described in detail below with reference to the accompanying drawings.

[0028] This application provides a bone adhesive with hemostatic and osteogenic functions, comprising a solid phase component and a liquid phase component; The solid phase component, by mass parts, includes 85-90 parts of calcium phosphate matrix, 8-12 parts of α-amino acid compounds, and 1-10 parts of silicon-based active glass. The α-amino acid compounds contain at least one of phosphorylated hydroxyl groups and free hydroxyl groups; The liquid-to-solid ratio of the liquid phase component to the solid phase component is 0.3-1 mL / g, and after the solid phase component and the liquid phase component are mixed, the hydration product of the calcium phosphate matrix is ​​transparent calcium phosphate.

[0029] In this application, the bone adhesive consists of a solid phase component and a liquid phase component. Under normal conditions, the solid and liquid phase components are separately packaged, and only mixing is required during use. The optimal liquid-to-solid ratio of the liquid phase component to the solid phase component is 0.3-1 mL / g. Within this range, the hemolysis rate of the material is less than 5%, which meets the safety requirements for medical implant materials.

[0030] In this application, the calcium phosphate matrix is ​​divided into a first matrix and a second matrix. The core purpose is to directionally control the hydration product to form a permeapatite phase, avoiding the formation of a hydroxyapatite phase. The first matrix is ​​the matrix framework phase, primarily providing osteoconductivity and long-term mechanical support, and can be at least one of beta-tricalcium phosphate (β-TCP), α-tricalcium phosphate, tetracalcium phosphate, and octacalcium phosphate. The second matrix is ​​the hydration initiating phase, and can be at least one of monobasic calcium phosphate (MCPA) and calcium dihydrogen phosphate dihydrate. After hydrolysis, the second matrix forms a weakly acidic environment, promoting the dissolution of calcium and phosphate ions and their hydration reaction with the first matrix, ultimately generating a solidified, dense permeapatite product. The mass ratio of the first matrix to the second matrix is ​​(4-8):(1-5). In a preferred embodiment, the mass ratio of the first matrix to the second matrix is ​​2:1.

[0031] In this application, α-amino acid compounds possess dual chemical activity as both protonic acids and Lewis bases. The amino, carboxyl, and phosphate groups contained in the molecule can act as ligands and interact with Ca. 2+ When metal ions form stable chelates, a phosphate α-amino acid salt coordination network is ultimately constructed, simultaneously enhancing the cross-linking between calcium phosphate particles and achieving chemical adhesion between the material and bone tissue. If prioritizing improved wet adhesion strength is required, O-phosphoserine (OPLS) can be used, as its side-chain phosphorylated hydroxyl groups have higher activity and stronger coordination ability, resulting in higher wet adhesion strength than L-serine at the same addition amount. To reduce raw material costs, L-serine can be used, relying on its free side-chain hydroxyl groups for coordination, making it suitable for cost-sensitive primary healthcare settings. To enhance osteogenic activity, O-phospho-L-threonine can be used, as it is a natural monomer of bone matrix collagen, participating in cross-linking while directly providing raw materials for new bone matrix synthesis and increasing the expression of osteogenic-related genes.

[0032] In this application, silicon-based active glass and α-amino acids can synergistically enhance hemostatic and osteogenic properties through non-specific physical adsorption and specific receptor binding. Specifically, the silicon-based active glass, with its high specific surface area, non-specifically adsorbs plasma proteins, continuously releasing Si during the adsorption process. 4+ With Ca 2+ This activates the intrinsic coagulation pathway and enhances blood cell affinity; while the α-amino acid side chain functional group can specifically bind to phosphatidylserine receptors on platelet and erythrocyte membranes, and simultaneously through PO4 3- The high-density negative charge further enriches blood cells.

[0033] In this application, mesoporous bioactive glass (BG) is sufficient for routine clinical scenarios; if antibacterial effects on wounds are also required, copper / zinc doped mesoporous bioactive glass can be used, as copper / zinc ions can be continuously released to achieve broad-spectrum antibacterial effects; if further improvement in osteogenic efficiency is required, strontium doped mesoporous bioactive glass can be used, as strontium ions can synergistically upregulate the expression of osteogenic-related genes with silicon ions.

[0034] To facilitate a further understanding of this application, the solutions described below are further described in conjunction with embodiments. Those skilled in the art will understand that the examples described in this application are only a portion of the examples, and any other suitable specific examples are within the scope of this application.

[0035] I. Preparation of Silicon-based Active Glass The silicon-based active glass described in this embodiment is selected from any one of mesoporous bioactive glass, strontium-doped mesoporous bioactive glass, copper-doped mesoporous bioactive glass, and zinc-doped mesoporous bioactive glass. The specific preparation steps of the mesoporous bioactive glass are as follows: (1) Weigh 200g of cetyltrimethylammonium bromide and dissolve it in 200mL of deionized water. Heat the solution in a water bath to 40°C. After it is completely dissolved, add 60mL of ethyl acetate and stir at a constant temperature for 30min to form a stable emulsion system. (2) Add 42 mL of 2 mol / L ammonia water dropwise to the above emulsion system. After the addition is complete, continue stirring at a constant temperature for 15 min to form an alkaline catalytic reaction system. (3) Add 28 mL of tetraethyl orthosilicate dropwise to the above alkaline catalytic reaction system. After the addition is complete, stir at a constant temperature for 30 min. Then add 2.8 mL of triethyl phosphate dropwise. After the addition is complete, stir at a constant temperature for 30 min. Then add calcium nitrate tetrahydrate aqueous solution (prepared by dissolving 18 g of calcium nitrate tetrahydrate in 25 mL of deionized water). After the addition is complete, stir at a constant temperature for 3 h. The reaction yields mesoporous bioactive glass gel. (4) The obtained gel was aged at 37°C for 3 days. After aging, it was washed three times with deionized water and anhydrous ethanol to remove residual organic template agent. After washing, the gel was dried in an oven at 60°C until constant weight. (5) Place the dried gel in a muffle furnace and calcine it at 650°C to remove residual organic matter. After naturally cooling to room temperature, grind it through a 200-mesh sieve to obtain mesoporous bioactive glass powder. Seal and dry it for later use.

[0036] Correspondingly, the preparation method of strontium, copper and zinc doped mesoporous bioactive glass is basically the same as the above steps. The only difference is that in step (3), while adding calcium nitrate tetrahydrate, the doped ions replace 2%-10% of the total calcium ions. Equivalent amounts of strontium nitrate tetrahydrate, copper nitrate trihydrate or zinc nitrate hexahydrate are added to obtain strontium-doped, copper-doped or zinc-doped mesoporous bioactive glass respectively.

[0037] The bone adhesive described in this application consists of separately packaged solid and liquid components, which are mixed in proportion before use. The specific preparation method is as follows: 2.1 Preparation of solid components Weigh out the following raw materials according to the specified proportions: 85-90 parts of calcium phosphate matrix, 8-12 parts of α-amino acid compounds, and 1-10 parts of silicon-based active glass. Optionally, add 0.5-2 parts of functional additives.

[0038] The calcium phosphate matrix is ​​composed of a first matrix and a second matrix, with a mass ratio of (4-8):(1-5) between the first matrix and the second matrix. The first matrix is ​​selected from at least one of β-tricalcium phosphate, α-tricalcium phosphate, tetracalcium phosphate, and octacalcium phosphate, and the second matrix is ​​selected from at least one of anhydrous calcium dihydrogen phosphate and calcium dihydrogen phosphate dihydrate. The α-amino acid compound is selected from at least one of L-phosphoserine, O-phospho-L-threonine, and L-serine; The silicon-based active glass is selected from at least one of the mesoporous bioactive glass prepared in step one above and strontium / copper / zinc doped mesoporous bioactive glass. The functional additive is selected from at least one of antibiotics, anti-inflammatory drugs, and antimicrobial peptides.

[0039] When preparing the solid component, the weighed raw materials are added to a planetary ball mill, and anhydrous ethanol is used as the ball milling medium. The milling is carried out at a power of 30kW for 4 hours to make the powder uniformly mixed. After ball milling, the material is placed in a 120℃ oven to dry for 24 hours. After cooling, it is passed through a 200-mesh sieve to obtain the solid component, which is then sealed and stored for later use.

[0040] 2.2 Selection of liquid phase components The liquid phase component can be selected from any one of physiological saline, phosphate buffer, or biomimetic body fluid, depending on the clinical application scenario.

[0041] 2.3 Usage Method Before clinical use, the liquid phase component and the solid phase component are mixed and stirred evenly at a liquid-solid ratio of 0.3-1 mL / g to obtain a bone adhesive that can be directly applied. Depending on different clinical needs, the mixed material can be prepared into dosage forms such as injectable paste, pre-cured block, and porous sponge.

[0042] After the solid and liquid components are mixed, the calcium phosphate matrix undergoes a hydration reaction, and the final hydration product is phosphatidylcholine, whose degradation rate matches the natural new bone growth rate well.

[0043] To verify the technical effect of the technical solution of this application, the following embodiments are set up. The liquid phase component in each embodiment is physiological saline, and the liquid-solid ratio is 0.6 mL / g. Except for the different solid phase component ratio, the other preparation steps are the same as the preparation method described above.

[0044] III. Examples 3.1 Examples 1-9 In this series of examples, the α-amino acid compound is fixed as LO-phosphoserine, and the amount added is 10 parts by weight. No silicon-based active glass is added. The mass ratio of β-tricalcium phosphate (first matrix) and anhydrous calcium dihydrogen phosphate (second matrix) in the calcium phosphate matrix is ​​adjusted. The specific formulation is shown in Table 1. Table 1: Solid phase composition ratios of Examples 1-9 (unit: parts by mass)

[0045] Example 5 is referred to as TPCS.

[0046] 3.2 Examples 10-15 In this series of examples, the calcium phosphate matrix is ​​fixed at 60 parts by weight of β-tricalcium phosphate and 30 parts by weight of anhydrous calcium dihydrogen phosphate, with 10 parts by weight of LO-phosphoserine added. The amount of mesoporous bioactive glass added is adjusted. The specific formulation is shown in Table 2. Table 2: Solid phase composition ratios of Examples 10-15 (unit: parts by mass)

[0047] In Table 2, Example 11 is denoted as TPCS-BG.

[0048] IV. Performance Testing and Result Analysis 4.1 Physicochemical Performance Testing 4.1.1 Compressive strength test Test method: The solid and liquid components prepared in each embodiment were mixed uniformly in proportion, and cylindrical bone adhesive samples (Φ6 mm × 12 mm) were prepared using a mold and cured at 37 ℃ and 95% relative humidity for 1 day. Uniaxial compressive strength tests were conducted using a universal testing machine equipped with a 300 kN sensor. The loading rate stress rate was set to 0.5 MPa / s. The test ended when the specimen fractured (the load dropped to 50% of the maximum load or the sample showed obvious fragmentation). Each test was repeated 3 times, and the maximum compressive stress was taken as the compressive strength.

[0049] Test results: like Figure 1 As shown, a represents the change in compressive strength in Examples 1-9. When the mass ratio of β-tricalcium phosphate to anhydrous calcium dihydrogen phosphate is 60:30 (Example 5), the compressive strength reaches the highest value of 38.91±3.93MPa, which is significantly higher than other ratio groups. This indicates that the hydration reaction of the calcium phosphate matrix is ​​the most complete under this ratio, and the density of the cured product is the best.

[0050] like Figure 1 As shown, b represents the change in compressive strength in Examples 10-15. As the amount of mesoporous bioactive glass added increases, the compressive strength of the cured body decreases. When the amount of mesoporous bioactive glass added is 2 parts by mass (Example 11), the compressive strength is 30.01±2.72MPa, which can meet the mechanical support requirements for the repair of non-load-bearing bone defects, while retaining the biological properties of mesoporous bioactive glass.

[0051] 4.1.2 Wet Bond Strength Test Test Method: The bond strength was tested using a bovine bone lap shear test. Two smooth bovine bone pieces, each measuring 100mm × 20mm × 3mm, were bonded together at their ends using the test bone adhesive. The bonded area was 20mm × 20mm. The pieces were then cured at 37℃ and 95% relative humidity for 1 day. Tensile tests were performed using a universal testing machine equipped with a 300kN sensor. The applied stress rate was 0.1MPa / s. The maximum tensile stress at the point of fracture was recorded. Each test was repeated three times, and the average value was taken as the bond strength.

[0052] Test results: such as Figure 2 As shown, the bonding strength of Example 5 (TPCS) was 1.23 ± 0.13 MPa, and the bonding strength of Example 11 (TPCS-BG) was 1.21 ± 0.15 MPa. There was no statistically significant difference between the two groups, indicating that the addition of 2 parts by weight of mesoporous bioactive glass did not significantly affect the core bonding performance of the material, and the shear strength of about 1.2 MPa was sufficient to meet the bonding requirements for bone repair in non-weight-bearing areas.

[0053] 4.1.3 Setting Time Test Test Method: The setting time was determined using a Gilmore apparatus, referring to GB / T 1346-2011 "Standard Consistency Water Requirement, Setting Time and Soundness Test Method for Cement" and ASTM C266-07 standard. Samples with a diameter of 10mm × 5mm were prepared. Deionized water, PBS, sheep whole blood, and physiological saline were used as the liquid phase components, respectively. The initial setting time (the time when the initial setting needle leaves only a slight contact mark on the sample surface) and the final setting time (the time when the final setting needle leaves no obvious indentation on the sample surface) were recorded.

[0054] Test results: such as Figure 3 As shown, Example 11 (TPCS-BG) exhibited stable coagulation behavior in four different liquid phase environments. The initial coagulation time was consistently around 5 minutes (deionized water: 4.75±0.94 min, PBS: 5.83±1.03 min, whole blood: 4.92±1.46 min, normal saline: 5.58±1.56 min), while the final coagulation time was concentrated around 10 minutes (deionized water: 9.58±1.88 min, PBS: 12.25±2.04 min, whole blood: 8.75±1.47 min, normal saline: 11.00±2.17 min). This is well-suited to the rhythm of clinical intraoperative procedures, and there was no significant coagulation delay in the blood environment, which can avoid problems such as postoperative displacement or untimely bleeding closure.

[0055] 4.1.4 Phase Analysis Test method: Samples from Example 5 and Example 11 were cured for 1 day and their phases were characterized by X-ray diffraction.

[0056] Test results: XRD patterns are as follows Figure 4 As shown, in addition to the incompletely reacted β-tricalcium phosphate, the sample exhibited obvious diffraction peaks characteristic of hydroxyapatite, proving that the hydration product of the calcium phosphate matrix is ​​hydroxyapatite. Compared with traditional hydroxyapatite-type bone cement, it has a shorter curing time and a faster degradation rate.

[0057] In Example 5 (TPCS), in addition to the characteristic peak of incompletely reacted β-tricalcium phosphate, the hydration product showed obvious characteristic peaks of phosphatidylcholine, proving that the hydration product of the calcium phosphate matrix is ​​phosphatidylcholine, which has a shorter curing time and a faster degradation rate compared with traditional hydroxyapatite type bone cement.

[0058] 4.1.5 Interface Morphology Observation Test method: Take the sample after bonding the bone adhesive to the bovine bone in Example 11, after curing, dehydrate and sputter gold, and observe the micromorphology of the bonding interface using a scanning electron microscope (SEM).

[0059] Test results: SEM images as follows Figure 5 , Figure 6 As shown, a continuous transition layer is formed between the bone adhesive and the bone tissue. The pores and depressions on the surface of the bone fragments are fully filled by the material. Under high magnification, the sheet-like crystalline structure of the material can be seen extending into the bone tissue, forming a highly interlocked interface structure with no obvious gaps. This indicates that the material has good wettability and fluidity, and can achieve a dual bonding effect of mechanical interlocking and chemical bonding.

[0060] 4.1.6 Degradation performance test Test method: Take the bone adhesive prepared in Example 5 and Example 11, prepare a cured sample with a diameter of 10 mm × 5 mm, weigh it, place it in PBS buffer solution with pH 7.4, and soak it at a constant temperature of 37°C. Take out the sample at 1 day, 7 days, 14 days and 28 days respectively, dry it and weigh it, and calculate the proportion of the remaining mass to the initial mass.

[0061] Test results: Degradation curve as shown Figure 7 As shown, the mass loss rate within the first day of degradation was about 15%, mainly due to the rapid dissolution of soluble components on the material surface; at 28 days, the remaining mass ratios of the Example 5 (TPCS) group and the Example 11 (TPCS-BG) group were 71.46±1.24% and 69.92±3.88%, respectively. The degradation rate matched the natural new bone growth rate well, neither degrading prematurely and losing mechanical support, nor affecting the bone integration process due to foreign matter residue.

[0062] 4.2 Blood compatibility and hemostatic performance test 4.2.1 Blood cell adhesion test Test method: The solidified samples of Examples 10-15 were co-incubated with red blood cell suspension and platelet-rich plasma for 1 h, respectively. After washing with PBS to remove unadhered blood cells, they were fixed with 2.5% glutaraldehyde, dehydrated in a gradient, and the adhesion was observed by SEM. At the same time, the red blood cell adhesion rate and platelet adhesion rate were quantitatively calculated.

[0063] Test results: such as Figure 8 As shown, a) represents the erythrocyte adhesion rate in Examples 10-15; b) represents the platelet adhesion rate in Examples 10-15; c) represents the SEM images of erythrocyte adhesion in Examples 5 and 11; and d) represents the SEM images of platelet adhesion in Examples 5 and 11. Figure 8 It can be seen that as the amount of mesoporous bioactive glass added increases, blood cell adhesion first increases and then decreases; the erythrocyte adhesion rate of Example 11 reached 33.34%, and the platelet adhesion rate reached 69.09%, which were the highest values ​​among all groups, significantly higher than that of Example 10 without BG addition (erythrocyte adhesion rate 25.28%, platelet adhesion rate 31.62%). SEM observation showed that the erythrocytes adhering to the material surface were morphologically intact, and the platelets were fully spread and formed a dense fibrin network, indicating excellent coagulation effect.

[0064] 4.2.2 Hemolysis rate test Test method: The bone adhesive of Example 11 was mixed with fresh rabbit blood at solid-liquid ratios of 0.5 mg / mL, 1 mg / mL, 2 mg / mL and 5 mg / mL, incubated at 37°C for 1 h and then centrifuged. The supernatant was taken and the absorbance was measured at a wavelength of 545 nm to calculate the hemolysis rate.

[0065] Test results: such as Figure 9 As shown, a) are optical images of bone adhesives with liquid-to-solid ratios of 0.5 mg / mL, 1 mg / mL, 2 mg / mL, and 5 mg / mL after interaction with red blood cells; b) are the hemolysis rates of bone adhesives with liquid-to-solid ratios of 0.5 mg / mL, 1 mg / mL, 2 mg / mL, and 5 mg / mL. It can be seen that within the solid-to-liquid ratio range of 0.5-5 mg / mL, the hemolysis rate of all material groups is less than 5%, meeting the blood compatibility safety standards for medical implant materials and posing no risk of hemolysis.

[0066] 4.2.3 Whole Blood Coagulation Index (BCI) Test Test method: The solidified samples of Examples 10-15 were co-incubated with fresh anticoagulated whole blood. Distilled water was added at set time points to dissolve the uncoagulated blood. The absorbance at a wavelength of 545 nm was measured, and the whole blood coagulation index was calculated. The lower the BCI value, the stronger the procoagulant performance.

[0067] Test results: such as Figure 10 As shown, the BCI in Example 11 was 11%, the lowest among all groups, indicating the best procoagulant performance. As the amount of BG added further increased, the BCI gradually rose but remained significantly lower than that of the group without BG, indicating that 2 parts by mass is the optimal addition ratio of BG to balance hemostatic performance.

[0068] 4.2.4 Activated Partial Thromboplastin Time (APTT) Test Test method: Mix the bone adhesive from Examples 10-15 with anemic platelet plasma, add activating reagent and calcium ions, and measure the coagulation time. The lower the APTT value, the higher the activation efficiency of the intrinsic coagulation pathway.

[0069] Test results: such as Figure 11 As shown, the APTT of Example 11 was 28s, which was significantly shorter than that of other groups, indicating that it could efficiently activate the intrinsic coagulation pathway, further verifying the optimal hemostatic effect of adding 2 parts by weight of BG.

[0070] 4.3 Biocompatibility Testing 4.3.1 Cytotoxicity test Test methods: The effects of extracts from the blank control group (control), calcium sulfate bone cement positive control (CSC), Example 5 (TPCS), and Example 11 (TPCS-BG) on the proliferation activity of rat bone marrow mesenchymal stem cells (BMSCs) and human umbilical vein endothelial cells (HUVECs) were tested using the CCK-8 assay. OD values ​​were measured at 1, 3, and 5 days of culture, and live cell staining was performed simultaneously.

[0071] Test results: such as Figure 12 As shown, a) is live cell staining of BMSCs; b) is CCK-8 assay result of BMSCs; c) is live cell staining of HUVECs; and d) is CCK-8 assay result of HUVECs. It can be seen that all groups showed good cell proliferation activity and no obvious cytotoxicity. After 5 days of culture, the BMSCs proliferation activity of the Example 11 (TPCS-BG) group was significantly higher than that of other groups, indicating that the material can promote BMSCs proliferation and has excellent biocompatibility.

[0072] 4.3.2 Cell Scratch Assay Test methods: The cell scratch assay was used to verify the effect of materials on the migration ability of BMSCs. A blank control group (control), calcium sulfate bone cement positive control (CSC), Example 5 (TPCS) and Example 11 (TPCS-BG) were used respectively. After BMSCs were cultured to monolayer fusion, scratches were made, and material extract was added for culture. The scratch healing rate within 2 days was observed and calculated.

[0073] Test results: such as Figure 13 As shown, a) is a crystal violet staining image of Transwell cell migration; b) is a semi-quantitative analysis of Transwell staining results; c) is a cell scratch detection image and healing effect; d) is a quantitative analysis of cell scratches. It can be seen that the scratch healing rate of Example 11 exceeded 75% on day 2, significantly higher than other groups, indicating that the material can effectively promote cell migration and recruit endogenous stem cells to participate in bone repair.

[0074] 4.4 In vitro osteogenic performance testing 4.4.1 Alkaline phosphatase (ALP) activity test Test method: BMSCs were co-cultured with the extracts of blank control group (control), calcium sulfate bone cement positive control (CSC), Example 5 (TPCS), and Example 11 (TPCS-BG) for 7 days and 14 days respectively. Intracellular ALP activity was measured using an ALP detection kit, and ALP staining was observed at the same time.

[0075] Test results: such as Figure 14As shown in a and b, where a is an optical image of ALP staining; b is a quantitative analysis of ALP. It can be seen that at 7d and 14d, the ALP activity of Example 11 was significantly higher than that of other groups, and the ALP staining intensity was also significantly higher, indicating that the material can significantly promote early osteoblast differentiation.

[0076] 4.4.2 Staining test for mineralized nodules Test method: BMSCs were co-cultured with the extracts of blank control group (control), calcium sulfate bone cement positive control (CSC), Example 5 (TPCS) and Example 11 (TPCS-BG) for 14 days. The formation of mineralized nodules was observed by Alizarin Red S (ARS) staining, and the OD value was quantitatively measured after dissolving in hydrochloric acid.

[0077] Test results: such as Figure 14 As shown in c and 14d, where c is the optical image of ARS staining; d is the quantitative analysis of ARS. It can be seen that the number and size of mineralized nodules in group 11 of Example 11 were significantly better than those in other groups, and the quantitative OD value of ARS was the highest among all groups, indicating that the material can effectively promote extracellular matrix mineralization.

[0078] 4.4.3 Osteogenesis-related gene expression test Test method: BMSCs were co-cultured with the extracts of blank control group (control), calcium sulfate bone cement positive control (CSC), Example 5 (TPCS), and Example 11 (TPCS-BG) for 7 days and 14 days respectively. The expression levels of osteogenic related genes such as RUNX2, COL-1, ALP, OCN, and OPN were detected by RT-qPCR.

[0079] Test results: such as Figure 15 As shown, after 7 days of culture, the expression of RUNX2 and COL-1 in the Example 11 group was significantly upregulated, initiating the osteogenic differentiation program; after 14 days of culture, the expression of OCN and OPN was further enhanced, promoting the mineralization process, indicating that the material can achieve full-cycle regulation of osteogenic differentiation at the transcriptional level.

[0080] 4.5 In vivo bone repair performance test Test methods: such as Figure 16 As shown, a femoral defect model was constructed in SD rats. Male SD rats were used as experimental animals and were bred and operated on at Guangzhou Shuiyuntian Biotechnology Co., Ltd. Thirty-two rats were randomly divided into four groups of eight each: a blank control group (defect refilled in situ after reconstruction), a CSC group (calcium sulfate bone cement positive control), a TPCS group (experimental group), and a TPCS-BG group (experimental group). Unilateral left femoral surgery was performed. Specific details of the experimental procedure are as follows: Select the left hind limb, disinfect it appropriately, and administer inhalation anesthesia using veterinary isoflurane in a 2.0%–2.5% isoflurane-oxygen mixture. Then maintain anesthesia in a 1.5%–1.8% isoflurane-oxygen mixture, and adjust the gas concentration as needed. After anesthetizing the rats, they were fixed on the animal operating table and covered with surgical drapes, exposing their legs. A longitudinal incision was made on the lateral side of the femur, the skin and subcutaneous tissue were cut, the muscles and periosteum were separated, and the femur was exposed. Use a dental drill and a Φ5 mm drill bit to create a bone defect perpendicular to the femur; Complete bone tissue was removed and the medullary cavity was exposed. The material of each group was filled according to the grouping, and then the removed bone tissue was backfilled in situ. The wound was stitched up layer by layer with sutures and secured with an elastic bandage; Penicillin was administered intramuscularly for three consecutive days after the surgery to prevent infection. Rats were sacrificed and samples were collected at weeks 4 and 12 after material implantation.

[0081] Test results: Imaging results at 12 weeks post-surgery are as follows Figure 17 , 18 As shown, the bone defects in the TPCS-BG group were almost completely repaired, with the newly formed bone tightly fused with the host bone, and the bone structure was dense, closely resembling normal bone morphology. The quantitative analysis results of bone morphometrics are as follows: Figure 19 As shown, the TPCS-BG group had significantly higher bone volume fraction (BV / TV), trabecular thickness (Tb.Th), and trabecular number (Tb.N) than other groups, and the lowest trabecular spacing (Tb.Sp), indicating the best osseointegration effect.

[0082] In summary, Example 11, with its solid-phase composition consisting of 60 parts β-TCP, 30 parts MCPA, 10 parts OPLS, and 2 parts mesoporous bioactive glass, and a liquid-to-solid ratio of 0.6 mL / g, achieves the following technical advantages: 1. Physicochemical properties adapted to clinical needs: initial setting time is about 5 minutes, final setting time is about 10 minutes, compressive strength after curing is about 30 MPa, wet bone interface bonding strength is about 1.2 MPa, residual mass after 28 days of degradation is about 69.92%, and the degradation rate matches the new bone growth rate. 2. Excellent hemostatic performance: the hemolysis rate is less than 5%, the red blood cell adhesion rate is 33.34%, the platelet adhesion rate is 69.09%, the whole blood coagulation index is about 11%, and the activation time of partial thromboplastin is about 28 seconds, which can achieve multi-pathway synergistic hemostasis; 3. Good biocompatibility: No obvious cytotoxicity, can promote the proliferation and migration of BMSCs, and the cell scratch healing rate exceeds 75% after 2 days; 4. Outstanding osteogenic performance: It can regulate the expression of osteogenic-related genes throughout the entire cycle, promote the enhancement of ALP activity and the formation of mineralized nodules, and achieve near-complete repair of femoral defects in rats after 12 weeks, with excellent osseointegration effect.

[0083] The various embodiments of this application have been described above. These descriptions are exemplary and not exhaustive, nor are they limited to the disclosed embodiments. Many modifications and variations will be apparent to those skilled in the art without departing from the scope and spirit of the described embodiments. The terminology used herein is chosen to best explain the principles, practical application, or improvement of the technology in the market, or to enable others skilled in the art to understand the embodiments disclosed herein.

Claims

1. A bone adhesive having hemostatic and osteoinductive functions, characterized by, Includes solid phase components and liquid phase components; The solid phase component, by mass parts, includes 85-90 parts of calcium phosphate matrix, 8-12 parts of α-amino acid compounds, and 1-10 parts of silicon-based active glass. The α-amino acid compounds contain at least one of phosphorylated hydroxyl groups and free hydroxyl groups; The liquid-to-solid ratio of the liquid phase component to the solid phase component is 0.3-1 mL / g, and after the solid phase component and the liquid phase component are mixed, the hydration product of the calcium phosphate matrix is ​​transparent calcium phosphate.

2. The bone adhesive having hemostatic and pro-osteogenic functions according to claim 1, wherein, The calcium phosphate matrix includes a first matrix and a second matrix; The first matrix is ​​selected from at least one of β-tricalcium phosphate, α-tricalcium phosphate, tetracalcium phosphate, and octacalcium phosphate; The second matrix is ​​selected from at least one of anhydrous calcium dihydrogen phosphate and calcium dihydrogen phosphate dihydrate; The mass ratio of the first matrix to the second matrix is ​​(4-8):(1-5).

3. The bone adhesive having hemostatic and pro-osteogenic functions according to claim 2, characterized by, The mass ratio of the first matrix to the second matrix is ​​2:

1.

4. The bone adhesive having hemostatic and pro-osteogenic functions according to claim 1, wherein, The α-amino acid compound is selected from at least one of L-phosphoserine, O-phospho-L-threonine, and L-serine; The silicon-based active glass is selected from at least one of mesoporous bioactive glass, strontium-doped mesoporous bioactive glass, copper-doped mesoporous bioactive glass, and zinc-doped mesoporous bioactive glass.

5. The bone adhesive having hemostatic and pro-osteogenic functions according to claim 4, wherein, If the silicon-based active glass is a mesoporous bioactive glass, the mesoporous bioactive glass is prepared by the following method: (1) Dissolve hexadecyltrimethylammonium bromide in deionized water, add ethyl acetate at 40°C, and stir at a constant temperature until an emulsion system is formed; (2) Add alkaline solution dropwise to the emulsion system and stir at a constant temperature to form an alkaline catalytic reaction system; (3) Tetraethyl orthosilicate, triethyl phosphate, and calcium nitrate tetrahydrate aqueous solution were added dropwise to the alkaline catalytic reaction system, and the reaction was carried out under constant temperature and stirring to obtain mesoporous bioactive glass gel. (4) After aging the obtained gel, it is washed with deionized water and anhydrous ethanol alternately to remove residual organic components, and then dried. (5) The dried gel was calcined, cooled, ground and sieved to obtain mesoporous bioactive glass powder, which was then sealed and stored for later use.

6. The bone adhesive having hemostatic and pro-osteogenic functions according to claim 5, wherein, In step (1), the mass ratio of deionized water to hexadecyltrimethylammonium bromide is 1:1; the volume ratio of ethyl acetate to deionized water is 3:10; and the constant temperature stirring time is 30 min. In step (2), the alkaline solution is 2 mol / L ammonia water, the volume ratio of ammonia water to deionized water is 21:100, and the constant temperature stirring time after dropwise addition is 15 min. In step (3), the volume ratio of tetraethyl orthosilicate to deionized water is 7:50, and the mixture is stirred at a constant temperature for 30 min after being added dropwise; the volume ratio of triethyl phosphate to deionized water is 7:500, and the mixture is stirred at a constant temperature for 30 min after being added dropwise; the calcium nitrate tetrahydrate aqueous solution is prepared by dissolving 18 parts by mass of calcium nitrate tetrahydrate in 25 parts by mass of deionized water, and the mixture is stirred at a constant temperature for 3 h after being added dropwise. In step (4), the aging temperature is 37°C, the aging time is 3 days, the washing is 3 times, and the drying temperature is 60°C. In step (5), the calcination temperature is 650℃.

7. The bone adhesive having hemostatic and pro-osteogenic functions according to claim 1, wherein, The solid phase components, by mass, include 90 parts of calcium phosphate matrix, 10 parts of α-amino acid compounds, and 2 parts of silicon-based active glass; The liquid-to-solid ratio is 0.6 mL / g.

8. The bone adhesive having hemostatic and pro-osteogenic functions according to claim 1, wherein, The solid phase component further includes 0.5-2 parts of additives, wherein the additives are selected from at least one of antibiotics, anti-inflammatory drugs, and antimicrobial peptides; The liquid phase component is selected from any one of physiological saline, phosphate buffer, and biomimetic body fluid.

9. The bone adhesive having hemostatic and pro-osteogenic functions according to claim 1, wherein, The bone adhesive formulation is selected from any one of injectable paste, pre-cured block, or porous sponge.

10. The application of the bone adhesive with hemostatic and osteogenic functions as described in any one of claims 1 to 9 in the preparation of hemostatic materials for bone wounds, adhesive materials for non-load-bearing bone fractures, hemostatic sealing materials for medullary cavities, and maxillofacial bone repair materials.

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