Injectable mineralized collagen bone repair composite material loaded with stem cell recruitment factor and osteogenic induction factor and preparation method of injectable mineralized collagen bone repair composite material

Abstract

The invention belongs to the field of biomedical materials and tissue engineering, and particularly relates to an injectable mineralized collagen bone repair composite material loaded with stem cell recruitment factors and osteogenic induction factors and a preparation method of the injectable mineralized collagen bone repair composite material. The composite material is composed of 50%-53% of mineralized collagen matrix particles, 15%-20% of SDF-1alpha loaded heparinized gelatin microspheres, 15%-20% of BMP-2 loaded PDA coated mesoporous silica particles, 8%-10% of an NHS-PEG-NHS cross-linking agent and 4%-6% of glycerin. The composite material can realize time-space sequential regulation and control of'collection first and then differentiation '; a bionic bone matrix microenvironment is provided through the mineralized collagen matrix; the problems of two-factor release conflict and stability are solved; in-vitro experiments show that the material can significantly promote stem cell migration and osteogenic differentiation, and animal experiments prove that the material can recover 82-88% of mechanical properties of bone defect parts. The invention solves the technical problems of insufficient stem cell collection, low osteogenic differentiation efficiency and uncontrollable factor release of the existing bone repair material.

Description

Technical Field

[0001] This invention belongs to the field of biomedical materials and tissue engineering, specifically relating to an injectable mineralized collagen bone repair composite material loaded with stem cell recruitment factors and osteogenic inducing factors, and its preparation method. Background Technology

[0002] Repairing large bone defects and irregular bone cavities is a major challenge in orthopedics. These types of bone injuries are commonly seen in comminuted fractures with bone loss caused by severe trauma such as high-energy traffic accidents and falls from heights. Due to the specific nature of the injury site, traditional fixation and repair methods often fail to meet clinical needs, necessitating the use of injectable materials to fill these complex bone defect areas and restore skeletal integrity and function. Injectable materials offer advantages such as ease of use and adaptability to irregular bone cavity morphologies, allowing for better conformation to the defect site and providing initial support and structural foundation for bone repair.

[0003] Currently, commonly used bone repair materials in clinical practice mainly include PMMA bone cement, calcium phosphate cement, and collagen sponge. However, these materials have varying degrees of defects and cannot meet the needs of repairing complex bone defects.

[0004] PMMA bone cement is a widely used orthopedic filling material, but it has significant limitations. First, PMMA bone cement lacks bioactivity and cannot form good biointegration with surrounding bone tissue, hindering long-term bone repair and regeneration. Second, during the polymerization process, PMMA bone cement releases a large amount of heat; this exothermic reaction can damage surrounding cells, affecting cell survival and function, thus impeding the bone repair process. Furthermore, PMMA bone cement is non-degradable, posing a risk of foreign body reactions when present in the body long-term, and its fixed shape may fail to adapt to new physiological needs as bone grows and remodels.

[0005] Calcium phosphate materials possess certain bioactivity and osteoconductivity, but their slow degradation rate cannot match the rate of new bone formation, thus limiting new bone growth within the material and affecting the quality and efficiency of bone repair. Furthermore, calcium phosphate materials are brittle and have poor mechanical properties, making them prone to fracture under physiological loads. They cannot provide sufficient mechanical support to bone defects, and lack effective cell recruitment capabilities, making it difficult to attract enough stem cells to the damaged area to participate in the bone repair process.

[0006] While single collagen / HA (hydroxyapatite) complexes mimic the composition of natural bone tissue to some extent, their weak mechanical properties prevent them from withstanding significant mechanical stress, limiting their application in repairing weight-bearing bone defects. More importantly, these complexes lack the ability to directionally regulate stem cell behavior and cannot effectively guide stem cells to differentiate into osteoblasts, resulting in unsatisfactory bone repair outcomes.

[0007] Bone repair is a complex and ordered biological process, mainly involving two key processes: the efficient recruitment of stem cells and the directed differentiation of stem cells into osteoblasts. However, existing bone repair technologies have significant limitations in both of these aspects.

[0008] Stem cells play a crucial role in bone repair, migrating to the injury site and differentiating into osteoblasts to participate in new bone formation. However, insufficient stem cell homing at the injury site is a common problem. Existing single-factor loading systems, such as collagen scaffolds loaded only with BMP-2 (bone morphogenetic protein-2), while BMP-2 has some osteogenic induction effects, cannot overcome the bottleneck of stem cell recruitment. Due to the lack of effective stem cell recruitment factors, the number of stem cells reaching the injury site is insufficient, limiting the initiation of bone repair and hindering the formation of enough new bone tissue.

[0009] Even if some stem cells manage to reach the injury site, guiding them to efficiently differentiate into osteoblasts remains a key challenge. Currently, single-factor loading systems are relatively inefficient in promoting osteogenic differentiation of stem cells and are prone to uncontrolled differentiation. For example, the burst release of BMP-2 may induce heterotopic ossification, i.e., the formation of bone tissue in non-bone areas, which can not only affect the normal function of surrounding organs and tissues but may also lead to serious complications. Furthermore, the effects of BMP-2 have dose-dependent side effects; excessively high doses may trigger adverse reactions such as inflammatory responses and nerve damage.

[0010] To overcome the shortcomings of single-factor loading systems, researchers have attempted to develop simple dual-factor hybrid carriers, such as blending collagen and growth factors. However, the timing of factor release is uncontrollable. During bone repair, stem cell recruitment factors (such as SDF-1α and stromal cell-derived factor-1α) need to be released rapidly to quickly attract stem cells to the injury site, while osteogenic inducing factors (such as BMP-2) need to be released slowly to continuously promote the differentiation of stem cells into osteoblasts. Simple dual-factor hybrid carriers cannot achieve this precise control of release timing, leading to a mismatch between factor release and the physiological needs of bone repair. Furthermore, growth factors are easily inactivated. Enzymes such as collagenase degrade collagen, and the instability of the solution environment also affects growth factor activity, significantly shortening the effective duration of growth factors and preventing them from fully realizing their biological functions. In addition, existing simple dual-factor hybrid carriers lack mechanical support and degradation matching design, failing to provide sufficient mechanical support while gradually degrading as new bone forms, thus affecting the overall effectiveness of bone repair.

[0011] In summary, existing bone repair materials have many shortcomings in addressing large bone defects and irregular bone cavities, failing to meet clinical demands for efficient and safe bone repair. There is an urgent need to develop an injectable mineralized collagen bone repair composite material capable of loading stem cell recruitment factors and osteogenic inducing factors, while possessing good mechanical properties, controllable degradation properties, and precise release timing. Summary of the Invention

[0012] The present invention provides an injectable material that achieves spatiotemporal sequential regulation of "recruitment followed by differentiation"; provides a biomimetic bone matrix microenvironment through a mineralized collagen matrix; solves the problems of dual-factor release conflict and stability; and enables the material to be injected at room temperature, shaped during surgery, and has in-situ self-enhancing mechanical properties.

[0013] On one hand, the present invention provides an injectable mineralized collagen bone repair composite material loaded with stem cell recruitment factors and osteogenic inducing factors, comprising the following components: 50%-53% mineralized collagen matrix microparticles, wherein the mineralized collagen matrix microparticles are formed by electrochemical co-deposition mineralization of type I collagen, and the particle size is 200-400μm; 15%-20% heparinized gelatin microspheres loaded with SDF-1α; 15%-20% BMP-2 loaded PDA-coated mesoporous silica microparticles; 8%-10% NHS-PEG-NHS crosslinking agent; 4%-6% glycerin.

[0014] Furthermore, the injectable mineralized collagen bone repair composite material is composed of the following components: 53% mineralized collagen matrix microparticles; 15% heparinized gelatin microspheres loaded with SDF-1α; 20% BMP-2 loaded PDA-coated mesoporous silica microparticles; 8% NHS-PEG-NHS crosslinking agent; 4% glycerin.

[0015] Furthermore, in the aforementioned injectable mineralized collagen bone repair composite material, the mineralized collagen matrix microparticles are prepared by the following method: a) Dissolve type I collagen in 0.5M acetic acid to a final concentration of 15-30 mg / mL; b) Mix the collagen solution with 5 times the concentration of simulated body fluid at a volume ratio of 4:1, and add sodium β-glycerophosphate to a final concentration of 0.1M; c) Electrochemical co-deposition was performed at a constant potential of 2.0–4.0 V at a temperature of 37 °C. The deposition was terminated when the charge transfer resistance increased by 25%. d) After gradient dehydration, freeze-dry and mechanically pulverize and sieve to obtain 200-400μm particles.

[0016] Furthermore, in the injectable mineralized collagen bone repair composite material, the heparinized gelatin microspheres loaded with SDF-1α are prepared by the following method: S1. Gelatin and sodium heparin are reacted at a mass ratio of 5:1 under EDC / NHS catalysis to prepare heparinized gelatin. S2. Using the oil-in-water emulsification method, the heparinized gelatin solution and paraffin oil are mixed at a volume ratio of 1:3-10 and emulsified by stirring at 300-3000 rpm. S3, add 0.1% glutaraldehyde for cross-linking and curing, wash and load with 50 μg / mL SDF-1α.

[0017] Furthermore, in the injectable mineralized collagen bone repair composite material, the BMP-2-loaded PDA-coated mesoporous silica microparticles are prepared by the following method: Step 1: Prepare mesoporous silica using the CTAB template method; Step 2: Disperse mesoporous silica in a 0.6-0.8 mg / mL BMP-2 solution and allow it to adsorb by shaking at 4°C; Step 3: Add 12-24 mg of dopamine hydrochloride to Tris-HCl buffer and shake at room temperature to form a 10-20 nm PDA coating.

[0018] On the other hand, a method for preparing the injectable mineralized collagen bone repair composite material of the present invention is also provided, characterized by comprising the following steps: 1) Mineralized collagen matrix microparticles, heparinized gelatin microspheres loaded with SDF-1α, and PDA-coated mesoporous silica microparticles loaded with BMP-2 were mixed in proportion. 2) Add PBS containing 0.1% TWEEN80 and mix until a paste is formed; 3) Add NHS-PEG-NHS crosslinking agent and glycerin before injection and load them into the injection device.

[0019] Furthermore, the injectable mineralized collagen bone repair composite material has a viscosity of 350-400 Pa·s at 25°C and a gelation time of 4.6-6.0 minutes.

[0020] Furthermore, the wet compressive modulus of the injectable mineralized collagen bone repair composite material is 0.57-0.67 MPa.

[0021] Furthermore, in the injectable mineralized collagen bone repair composite material, SDF-1α is released cumulatively at 41%-45% within 24 hours, and BMP-2 is released cumulatively at 76%-85% within 14 days.

[0022] Finally, the invention also provides the application of the injectable mineralized collagen bone repair composite material in the preparation of bone defect repair materials.

[0023] Compared with the prior art, the present invention has at least the following advantages or beneficial effects: Time-controlled release of two factors: SDF-1α is rapidly released through heparinized gelatin microspheres (41%-45% release in 24 hours) to achieve efficient recruitment of stem cells, while PDA-coated mesoporous silica slowly releases BMP-2 (76%-85% release in 14 days) to continuously promote osteogenic differentiation, matching the biological process requirements of bone repair.

[0024] Synergistic effect: The dual-factor composite material increased stem cell migration rate by 230%-245%, ALP activity by 300%-325%, and osteocalcin expression by 180%-201%, which was significantly better than the single-factor system (p<0.01), confirming that BMP-2 and SDF-1α have a synergistic effect in promoting osteoogenesis.

[0025] Excellent mechanical properties: The synergistic effect of mineralized collagen matrix and crosslinking agent enables the composite material to achieve a wet compressive modulus of 0.57-0.67 MPa and a degradation rate of 58%-60% in 8 weeks. The mechanical properties and degradation rate are perfectly matched with the formation of new bone.

[0026] Strong clinical applicability: The material has a viscosity of 350-400 Pa·s at 25℃ and a gelation time of 4.6-6.0 minutes. It can be precisely implanted into irregular bone defect sites through a syringe and solidified in situ during surgery to form mechanical support.

[0027] Biomimetic microenvironment construction: The mineralized collagen matrix (Ca / P=1.67) prepared by electrochemical co-deposition highly mimics the components of natural bone matrix, providing cells with an ideal osteogenic microenvironment.

[0028] Safety assurance: The release rate of BMP-2 is precisely controlled by the PDA coating (10-20nm coating thickness), avoiding side effects such as ectopic ossification caused by the sudden release of growth factors. Animal experiments show that no inflammation or abnormal ossification occurred in the material group. Attached Figure Description

[0029] Figure 1 The image shows the results of alkaline phosphatase activity assay.

[0030] Figure 2 The image shows the results of osteopontin concentration measurement.

[0031] Figure 3 This is a graph showing the results of the relative value determination of mineralized area.

[0032] Figure 4 The graph shows the results of the relative expression level measurement of Runx2.

[0033] Figure 5 The graph shows the results of the relative expression level determination of Osterix. Detailed Implementation

[0034] The technical solution of the present invention will be described below with reference to the embodiments. However, the present invention is not limited to the following embodiments.

[0035] To enable those skilled in the art to better understand and implement the technical solutions of the present invention, the present invention will be further described below in conjunction with specific embodiments and accompanying drawings. However, the embodiments described are not intended to limit the present invention.

[0036] Unless otherwise specified, the experimental and detection methods described in the following embodiments are conventional methods; unless otherwise specified, the reagents and materials are commercially available.

[0037] Example 1

[0038] This embodiment describes the preparation of mineralized collagen matrix (Col-HA) microparticles.

[0039] 1.1 Raw material preparation Type I collagen: derived from beef tendon, molecular weight (MW) 300 kDa, with good biocompatibility and similar composition to human bone collagen.

[0040] Simulated body fluid (SBF×5): 5 times the concentration of standard SBF, containing Ca²⁺ and PO₄³⁻ ions, with a Ca / P molar ratio of 1.67 (consistent with natural bone tissue).

[0041] 0.5M acetic acid: pH 3.0, used to dissolve collagen and disrupt the interactions between collagen peptide chains.

[0042] Sodium β-glycerophosphate (β-GP): final concentration 0.1M, used as a mineralization promoter.

[0043] 1.2 Preparation steps 1.2.1 Collagen Pretreatment Dissolution: The type I collagen lyophilized powder was slowly added to 0.5M acetic acid and stirred at 4°C for 24 hours to achieve a final concentration of type I collagen of 15 mg / mL.

[0044] Purification: The solution was centrifuged at 8000 rpm for 30 min to remove undissolved impurities. The supernatant was dialyzed against deionized water using a 10 kDa dialysis bag for 48 h (water changed every 4 h) to obtain a purified collagen solution.

[0045] 1.2.2 Electrochemical Co-deposition Electrolyte preparation: Mix SBF×5 with collagen solution at a volume ratio of 4:1, add β-GP to a final concentration of 0.1M, and stir until homogeneous.

[0046] Deposition parameters: Voltage mode: constant potential 2.0V; Temperature: 37℃.

[0047] Electrode configuration: Anode: titanium alloy mesh (pore size 200μm); Cathode: platinum sheet, electrode spacing 10mm.

[0048] The charge transfer resistance (Rct) was monitored in real time using electrochemical impedance spectroscopy (EIS), and the reaction was terminated when Rct suddenly increased by 25%.

[0049] 1.2.3 Post-processing technology Gradient dehydration: Replace with 30%, 50%, and 70% ethanol for 15 min each time; finally replace with tert-butanol twice (20 min each time) to reduce ice crystal damage.

[0050] Freeze-drying: Pre-freeze at -80℃ for 2 hours to fix the porous structure; maintain at -50℃ and 5Pa vacuum for 12 hours, then raise the temperature to 25℃ at 1℃ / min and reduce the vacuum to 0.1Pa to remove bound water, thus obtaining No. 1 mineralized collagen matrix (Col-HA), which has a porous scaffold structure.

[0051] 1.2.4 Particle Preparation The No. 1 mineralized collagen matrix (Col-HA) was mechanically pulverized, and 200μm particles were sieved and labeled as No. 1 Col-HA particles for subsequent applications.

[0052] Example 2

[0053] This embodiment describes the preparation of mineralized collagen matrix (Col-HA) microparticles.

[0054] In this example, all steps were identical except that the collagen pretreatment concentration was 30 mg / mL; the electrochemical co-deposition potential was 3.0 V; and 300 μm particles were sieved during particle preparation. These particles were labeled 2# Col-HA for subsequent applications.

[0055] Example 3

[0056] This embodiment describes the preparation of mineralized collagen matrix (Col-HA) microparticles.

[0057] In this example, the steps were identical except that the collagen pretreatment concentration was 20 mg / mL; the electrochemical co-deposition potential was 4.0 V; and 400 μm particles were sieved during particle preparation. These particles were labeled 3# Col-HA for subsequent applications.

[0058] Example 4

[0059] This embodiment describes the preparation of microspheres loaded with SDF-1α.

[0060] Synthesis of heparinized gelatin: 10 g of gelatin was dissolved in 100 mL of PBS (pH 6.0) and stirred in a water bath at 50 °C until completely dissolved. Sodium heparin (mass ratio to gelatin 1:5) was weighed and dissolved in PBS (pH 6.0). EDC (5 times the molar amount of heparin) and NHS (1:1 molar amount of EDC) were added, and the mixture was stirred at room temperature for 30 min. The activated heparin solution was added dropwise to the gelatin solution, and the reaction was carried out at 50 °C for 2 h (pH maintained at 6.0). Glycine (0.1 M) was added to quench the remaining EDC, and the mixture was dialyzed at 4 °C (MWCO 12 kDa) for 48 h. The gelatin was then lyophilized to obtain heparinized gelatin.

[0061] Preparation of microspheres by oil-in-water emulsion method: Oil phase preparation: Paraffin oil was mixed with Span 80 (2% v / v) and preheated to 50°C; Aqueous phase preparation: Heparinized gelatin was dissolved in PBS (5% w / v) and kept at 50°C; Emulsification: The aqueous phase was slowly added dropwise to the oil phase (volume ratio 1:10), and stirred at 3000 rpm for 15 min to form an emulsion; Crosslinking and curing: Glutaraldehyde solution was added dropwise to a final concentration of 0.1%, and stirring was continued for 2 h; Microsphere collection: The microspheres were centrifuged at 4000 rpm, washed three times with isopropanol, activated, and then SDF-1α (final concentration 50 μg / mL) was added. The mixture was shaken at 4°C for 24 h, and 5 wt% mannitol was added to obtain 1#Hep-Gel@SDF-1α.

[0062] Example 5

[0063] This embodiment describes the preparation of microspheres loaded with SDF-1α.

[0064] The difference between this example and Example 4 is that the emulsification conditions are as follows: the aqueous phase is slowly added dropwise to the oil phase (volume ratio 1:5), and the mixture is stirred at 1500 rpm for 10 min to form an emulsion. All other conditions are the same, and 2#Hep-Gel@SDF-1α is prepared.

[0065] Example 6

[0066] This embodiment describes the preparation of microspheres loaded with SDF-1α.

[0067] The difference between this example and Example 4 is that the emulsification conditions are as follows: the aqueous phase is slowly added dropwise to the oil phase (volume ratio 1:3), and the mixture is stirred at 800 rpm for 8 minutes to form an emulsion. All other conditions are the same, and 3#Hep-Gel@SDF-1α is prepared.

[0068] Example 7

[0069] This example describes the preparation of a carrier for BMP-2 (PDA coating thickness 10 nm, BMP-2 drug loading 6 wt%).

[0070] The main reagents used in this example include: Cetyltrimethylammonium bromide (CTAB, 99%); tetraethyl orthosilicate (TEOS, 98%); bone morphogenetic protein-2 (BMP-2, recombinant human); dopamine hydrochloride (DA·HCl, 98%); tris(hydroxymethyl)aminomethane (Tris, 99.9%); phosphate-buffered saline (PBS, pH 7.4); mannitol.

[0071] 7.1 Synthesis of mesoporous silica (mSiO2) by CTAB template method CTAB solution preparation: Dissolve 1g CTAB in 50mL deionized water and stir at 65℃ until completely dissolved (concentration 2% w / v).

[0072] 5 mL of TEOS was slowly added dropwise to the CTAB solution, and the mixture was stirred at 500 rpm for 24 h at 80 °C. The reaction solution was allowed to stand at room temperature for 12 h to form a white precipitate. The precipitate was centrifuged at 8000 rpm for 10 min, washed three times with deionized water and twice with ethanol, and then vacuum dried at 60 °C for 12 h. The temperature was then raised to 250 °C and held for 30 min. The temperature was then raised to 350 °C and held for 1 h. The temperature was then raised to 550 °C and held for 4 h. The mixture was then allowed to cool naturally to room temperature to obtain mSiO2 powder.

[0073] 7.2 BMP-2 Adsorption The required mass of BMP-2 is calculated based on a drug loading of 6 wt%. Assuming the mass of mSiO2 is 100 mg, the required mass of BMP-2 is 100 mg × 6% = 6 mg.

[0074] A BMP-2 solution of appropriate concentration was prepared using PBS, such that 10 mL of the solution contained 6 mg of BMP-2, i.e., a concentration of 0.6 mg / mL.

[0075] Disperse 100 mg mSiO2 in 10 mL of the above-prepared BMP-2 solution to ensure that BMP-2 can be fully adsorbed into the mSiO2 channels. Shake at 100 rpm and 4 °C for 12 h. Then centrifuge at 8000 rpm for 10 min and wash three times with PBS to remove unadsorbed protein, thus obtaining mSiO2 adsorbed with BMP-2.

[0076] 7.3 Dopamine (PDA) Coating A product with a PDA coating thickness of 10 nm was prepared by controlling the reaction conditions. A certain amount of DA·HCl was dissolved in 10 mL of Tris-HCl buffer (10 mM, pH 8.5). After preliminary experiments, it was determined that adding 12 mg of DA·HCl could achieve the expected coating thickness.

[0077] Add mSiO2 that adsorbs BMP-2, shake at 100 rpm at room temperature for 6 h, then centrifuge at 8000 rpm for 10 min, wash 3 times with PBS and 2 times with deionized water, add 5 wt% mannitol as a protective agent, and freeze dry to obtain PDA-mSiO2-BMP-2 powder with a PDA coating thickness of 10 nm and a drug loading of 6 wt%, labeled as 1#PDA-mSiO2-BMP-2.

[0078] Example 8

[0079] This example describes the preparation of a carrier for BMP-2 (PDA coating thickness 15 nm, BMP-2 drug loading 7 wt%).

[0080] The main reagents used in this example are the same as those in Example 8.

[0081] 8.1 BMP-2 Adsorption The required mass of BMP-2 is calculated based on a drug loading of 7wt%. Assuming the mass of mSiO2 is 100mg, the required mass of BMP-2 is 100mg × 7% = 7mg.

[0082] A BMP-2 solution with a concentration of 0.7 mg / mL was prepared using PBS. 100 mg of mSiO2 was dispersed in 10 mL of this solution to ensure that BMP-2 was fully adsorbed into the mSiO2 channels. The solution was shaken at 100 rpm and 4 °C for 12 h. Then, the solution was centrifuged at 8000 rpm for 10 min and washed three times with PBS to remove unadsorbed proteins, thus obtaining mSiO2 adsorbed with BMP-2.

[0083] 8.2 Dopamine (PDA) Coating 18 mg of DA·HCl was dissolved in 10 mL of Tris-HCl buffer (10 mM, pH 8.5), and mSiO2 adsorbed with BMP-2 was added. The mixture was shaken at 100 rpm at room temperature for 6 h, followed by centrifugation at 8000 rpm for 10 min. The powder was then washed three times with PBS and twice with deionized water. 5 wt% mannitol was added as a protective agent, and the powder was freeze-dried to obtain PDA-mSiO2-BMP-2 powder with a PDA coating thickness of 15 nm and a drug loading of 7 wt%, which was labeled as 2#PDA-mSiO2-BMP-2.

[0084] Example 9

[0085] This example describes the preparation of a carrier for BMP-2 (PDA coating thickness 20 nm, drug loading 8 wt%).

[0086] The main reagents used in this example are the same as those in Example 8.

[0087] 9.1 BMP-2 Adsorption The required mass of BMP-2 is calculated based on a drug loading of 8wt%. Assuming the mass of mSiO2 is 100mg, the required mass of BMP-2 is 100mg × 8% = 8mg.

[0088] A BMP-2 solution with a concentration of 0.8 mg / mL was prepared using PBS. 100 mg of mSiO2 was dispersed in 10 mL of this solution to ensure that BMP-2 was fully adsorbed into the mSiO2 channels. The solution was shaken at 100 rpm and 4 °C for 12 h. Then, the solution was centrifuged at 8000 rpm for 10 min and washed three times with PBS to remove unadsorbed proteins, thus obtaining mSiO2 adsorbed with BMP-2.

[0089] 9.2 Dopamine (PDA) Coating 24 mg of DA·HCl was dissolved in 10 mL of Tris-HCl buffer (10 mM, pH 8.5), and mSiO2 adsorbed with BMP-2 was added. The mixture was shaken at 100 rpm at room temperature for 6 h, followed by centrifugation at 8000 rpm for 10 min. The powder was then washed three times with PBS and twice with deionized water. 5 wt% mannitol was added as a protective agent, and the powder was freeze-dried to obtain PDA-mSiO2-BMP-2 powder with a PDA coating thickness of 20 nm and a drug loading of 8 wt%, which was labeled as 3#PDA-mSiO2-BMP-2.

[0090] Example 10

[0091] This example demonstrates the preparation of composite materials. Composite materials are formulated according to the following component ratios:

[0092]

[0093] Mixing process: After mixing all dry powder components, add PBS containing 0.1% TWEEN80 to adjust to a paste consistency; Before injection, add cross-linking agent NHS-PEG-NHS, add glycerin, and then load it into injection equipment such as pipettes, syringes, and pushers for application.

[0094] Test Example 1 In this test, the shear rate (1s⁻¹) of composite materials 1#-3# was measured using a rheometer; the gelation time of composite materials 1#-3# was measured using the inverted bottle method; the mechanical properties of composite materials 1#-3# were determined by uniaxial compression test according to ISO 5833 "Surgical implants - Bone cement" standard; SDF-1α release was tested using ELSA method (PBS, pH 7.4); BMP-2 release was tested using ELSA method (containing 0.1% BSA); and the degradation cycle of composite materials 1#-3# was tested using simulated body fluid (SBF). The test results are shown in Table 1.

[0095] Table 1. Performance test results of different composite materials

[0096] Test Example 2 This test case is an in vitro test.

[0097] In this example, the single-loaded BMP-2 group sample is 3#PDA-mSiO2-BMP-2 prepared in Example 9, and the single-loaded SDF-1α group sample is 3#Hep-Gel@SDF-1α prepared in Example 6.

[0098] (1) The Transwell migration assay was used to test the chemotactic recruitment ability of the 1#-3# composite material for bone marrow mesenchymal stem cells (BMSCs). Sample grouping: blank collagen group (collagen, negative control); single BMP-2 loading group (single BMP-2 loading, positive control); single SDF-1α loading group (SDF-1α); 1#-3# composite material-dual factor co-loading group (composite material (BMP-2+SDF-1α), experimental group).

[0099] BMSCs (1×10 5Cells / well were seeded in the upper chamber of a Transwell, and the lower chamber was filled with a pre-incubated extract (DMEM diluted to 100 μg / mL) for 24 h. After incubation at 37°C for 12 h, the cells in the lower chamber were stained with crystal violet, and the number of migrating cells (n=5) was counted under a microscope. The rate of change in cell migration was calculated, and the results are shown in Table 2. The formula for calculating the rate of change in cell migration is as follows: Change rate of cell migration count (%) = (mean number of cell migrations in the experimental group - mean number of cell migrations in the negative control group) / mean number of cell migrations in the negative control group × 100 Table 2. Chemotactic recruitment capacity of different materials for bone marrow mesenchymal stem cells

[0100] As shown in Table 2, the dual-factor materials of the present invention (composite materials 1#-3#) exhibit significantly higher chemotactic recruitment capacity for bone marrow mesenchymal stem cells than the blank control collagen group, and significantly higher than the single-load SDF-1α group and the single-load BMP-2 group (BMP-2). This indicates that BMP-2 and SDF-1α work synergistically to enhance the chemotactic recruitment capacity of bone marrow mesenchymal stem cells in the composite materials containing both BMP-2 and SDF-1α.

[0101] (2) The osteogenic differentiation capacity of composite materials 1#-3# was tested using an in vitro induction model to verify the synergistic promoting effect of the two factors on osteogenic differentiation of BMSCs. Sample grouping: Blank collagen group (Collagen, negative control); BMP-2 single loading group (BMP-2 single loading group, positive control); SDF-1α single loading group (SDF-1α); 1#-3# composite material - dual factor co-loading group (composite material (BMP-2+SDF-1α), experimental group).

[0102] Prepare a bone induction medium (containing sodium β-glycerophosphate, ascorbic acid, and dexamethasone). Place composite materials #1-#3 and the remaining group samples into 24-well plates, inoculating each well with 3 × 10⁶ samples. 4 MC3T3-E1 osteogenic progenitor cells were cultured, and the cell culture supernatant was collected on day 14 of induction culture. The levels of alkaline phosphatase (ALP) and osteopontin (OPN) were detected by ELISA kit. The results are as follows: Figure 1 , Figure 2 As shown.

[0103] Depend on Figure 1 It can be seen that the ALP activity of the 1#-3# composite material was 300%-325% higher than that of the blank group (p<0.001), and the ALP activity of the 1#-3# composite material group was much higher than that of the single-loaded BMP-2 group and the single-loaded SDF-1α group, indicating that BMP-2 and SDF-1α in the composite material have a synergistic effect.

[0104] Depend on Figure 2 It can be seen that the OCN expression level of composite materials 1#-3# increased by 180%-201% compared with the blank group (p<0.01), which was significantly better than BMP-2 and SDF-1α, indicating that BMP-2 and SDF-1α in composite materials have a synergistic effect in increasing osteocalcin concentration.

[0105] In summary, in vitro induction model tests show that BMP-2 and SDF-1α in composite materials 1#-3# have a synergistic effect on osteogenic differentiation capacity.

[0106] The synergistic effect of BMP-2 and SDF-1α dual factors on extracellular matrix mineralization and the expression of key osteogenic genes in composite materials was tested. Sample grouping: blank collagen group (collagen, negative control); single BMP-2 loading group (single BMP-2 loading, positive control); single SDF-1α loading group (SDF-1α); 1#-3# composite material-dual factor co-loading group (composite material (BMP-2+SDF-1α), experimental group).

[0107] Alizarin Red Staining: The above samples were co-cultured with BMSCs for 14 days. To prevent calcium nodule loss and detachment, half the medium was replaced every 3 days. After osteogenic differentiation, the culture medium was aspirated from the culture plate, and the plate was washed three times with PBS buffer. 2 ml of 4% paraformaldehyde tissue fixative was added to each well for 30 minutes, followed by thorough washing with PBS buffer. Alizarin Red staining solution was then added for 8 minutes. The Alizarin Red staining solution was then aspirated with a pipette, and the plate was washed three more times with PBS buffer. The mineralization area was analyzed using ImageJ, and the data are expressed as a multiple of the control group. Results are as follows: Figure 3 As shown.

[0108] qRT-PCR detection: The above samples were co-cultured with BMSCs cells for 14 days, the culture medium was removed, and the cells were washed twice with sterile PBS. 1 mL of Trizol reagent was added to each well, and the cells were lysed at room temperature for 5 minutes. The lysate was transferred to a centrifuge tube, and total RNA was extracted according to the Trizol reagent instructions. The RNA was then reverse transcribed into cDNA using a reverse transcription kit.

[0109] Using cDNA as a template, qRT-PCR reactions were performed using a qRT-PCR kit and primers Runx2-F, Runx2-R and Osterix-F, Osterix-R. GAPDH was used as an internal reference gene, and the relative gene expression level was calculated using the 2⁻ΔΔCt method. Results are as follows: Figure 4 , Figure 5 As shown.

[0110] Primer sequences: Runx2-F: 5'-CCTCCATCCATTCACTGCAC-3'; Runx2-R: 5'-GGCTCAGGTAGGAGGGGTAA-3'.

[0111] Osterix-F: 5'-CCCAGTCAACAAGGAAGCC-3'; Osterix-R: 5'-AGGCTTGTGAAGGTGGGTGT-3'.

[0112] GAPDH-F: 5'-CAAGAGCACAAGAGGAAGAGAG-3'; GAPDH-R: 5'-CTACATGGCAACTGTGAGGAG-3'.

[0113] Depend on Figure 3 It can be seen that the mineralized area of ​​composite material test groups 1#-3# is 2.1-2.7 times larger than that of the blank group (p<0.001). The mineralized area of ​​composite material test groups 1#-3# is much higher than that of the blank control group, and much higher than that of the single-carrier BMP-2 and single-carrier SDF-1α groups. Depend on Figure 5 It can be seen that the expression levels of Runx2 and Osterix mRNA in the composite material 1#-3# test group were significantly upregulated, significantly higher than those in the single vector group, indicating that the composite material 1#-3# dual factor promotes osteogenic differentiation through multi-pathway synergy.

[0114] (4) The in vivo bone repair performance of composite materials 1#-3# was tested using a rabbit radius defect model.

[0115] Experimental groups: Blank collagen group (Collagen, negative control); BMP-2 single loading group (BMP-2 single loading, positive control); SDF-1α single loading group (SDF-1α); 1#-3# composite material-dual factor co-loading group (composite material (BMP-2+SDF-1α), experimental group).

[0116] Twenty-four New Zealand rabbits (3kg) were used to create 10mm defects in the midshaft of both radii to simulate clinical conditions. The experimental groups were treated with the same materials. After 8 weeks, the biomechanical properties of each test group were quantitatively analyzed using Micro-CT. The test results are shown in Table 3.

[0117] Table 3. In vivo bone repair performance of composite materials #1-#3

[0118] As shown in Table 3, the maximum load recovery rate of composite material group 1#-3# reached 82.72%-88.25%, and the elastic modulus recovery rate reached 82.06%-85.32%, which is close to the level of the healthy side bone (p>0.05). In contrast, the highest recovery rate of the single factor group was only 65.14% and 60.71% (p<0.01), indicating that the dual factors in composite material group 1#-3# synergistically promote functional bone regeneration.

Claims

1. An injectable mineralized collagen bone repair composite material loaded with stem cell recruitment factors and osteogenic inducing factors, characterized in that, It consists of the following components: 50%-53% mineralized collagen matrix microparticles, wherein the mineralized collagen matrix microparticles are formed by electrochemical co-deposition mineralization of type I collagen, and the particle size is 200-400 μm; 15%-20% heparinized gelatin microspheres loaded with SDF-1α; 15%-20% BMP-2 loaded PDA-coated mesoporous silica microparticles; 8%-10% NHS-PEG-NHS crosslinking agent; 4%-6% glycerin.

2. The injectable mineralized collagen bone repair composite material according to claim 1, characterized in that, It consists of the following components: 53% mineralized collagen matrix microparticles; 15% heparinized gelatin microspheres loaded with SDF-1α; 20% BMP-2 loaded PDA-coated mesoporous silica microparticles; 8% NHS-PEG-NHS crosslinking agent; 4% glycerin.

3. The injectable mineralized collagen bone repair composite material according to claim 1, characterized in that, The mineralized collagen matrix microparticles are prepared by the following method: a) Dissolve type I collagen in 0.5M acetic acid to a final concentration of 15-30 mg / mL; b) Mix the collagen solution with 5 times the concentration of simulated body fluid at a volume ratio of 4:1, and add sodium β-glycerophosphate to a final concentration of 0.1M; c) Electrochemical co-deposition was performed at a constant potential of 2.0–4.0 V at a temperature of 37 °C. The deposition was terminated when the charge transfer resistance increased by 25%. d) After gradient dehydration, freeze-dry and mechanically pulverize and sieve to obtain 200-400μm particles.

4. The injectable mineralized collagen bone repair composite material according to claim 1, characterized in that, The heparinized gelatin microspheres loaded with SDF-1α were prepared by the following method: S1. Gelatin and sodium heparin are reacted at a mass ratio of 5:1 under EDC / NHS catalysis to prepare heparinized gelatin. S2. Using the oil-in-water emulsification method, the heparinized gelatin solution and paraffin oil are mixed at a volume ratio of 1:3-10 and emulsified by stirring at 300-3000 rpm. S3, add 0.1% glutaraldehyde for cross-linking and curing, wash and load with 50 μg / mL SDF-1α.

5. The injectable mineralized collagen bone repair composite material according to claim 1, characterized in that, The BMP-2-loaded PDA-coated mesoporous silica microparticles were prepared by the following method: Step 1: Prepare mesoporous silica using the CTAB template method; Step 2: Disperse mesoporous silica in a 0.6-0.8 mg / mL BMP-2 solution and allow it to adsorb by shaking at 4°C; Step 3: Add 12-24 mg of dopamine hydrochloride to Tris-HCl buffer and shake at room temperature to form a 10-20 nm PDA coating.

6. A method for preparing the injectable mineralized collagen bone repair composite material according to any one of claims 1-5, characterized in that, Includes the following steps: 1) Mineralized collagen matrix microparticles, heparinized gelatin microspheres loaded with SDF-1α, and PDA-coated mesoporous silica microparticles loaded with BMP-2 were mixed in proportion. 2) Add PBS containing 0.1% TWEEN80 and mix until a paste is formed; 3) Add NHS-PEG-NHS crosslinking agent and glycerin before injection and load them into the injection device.

7. The injectable mineralized collagen bone repair composite material according to claim 1, characterized in that, The composite material has a viscosity of 350-400 Pa·s at 25°C and a gelation time of 4.6-6.0 minutes.

8. The injectable mineralized collagen bone repair composite material according to claim 1, characterized in that, The wet compressive modulus of the composite material is 0.57-0.67 MPa.

9. The injectable mineralized collagen bone repair composite material according to claim 1, characterized in that, The SDF-1α was released cumulatively at 41%-45% within 24 hours, and the BMP-2 was released cumulatively at 76%-85% within 14 days.

10. The application of the injectable mineralized collagen bone repair composite material according to claim 1 in the preparation of bone defect repair materials.

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