Bone guided regeneration membrane photocurable in situ and its preparation method and application

Abstract

This invention discloses a bone-guided regeneration membrane capable of intraoperative in-situ photocuring, its preparation method, and its applications. The membrane comprises, sequentially from the bone side to the soft tissue side, PDA-Ca²⁺. + The membrane consists of a wetting and adhesion inner layer, a ternary interpenetrating network hydrogel body layer, and a non-crosslinked physical anti-adhesion HA outer layer. The body layer is loaded with VEGF-PLGA fast-release microspheres and BMP-2-PLGA sustained-release microspheres, and contains a LAP photoinitiator. Before photocuring, the membrane is soft and malleable, allowing for intraoperative cutting and shaping to adapt to irregular bone surfaces at nonunion sites. After irradiation with 405 nm visible light for 30–120 seconds, it transforms in situ into a semi-rigid support state, locking the membrane's shape and providing mechanical support. PDA-Ca² + Adhesive coatings enable screwless and sutureless bioadhesive fixation to moistened bone surfaces; PLGA dual microspheres achieve time-controlled release of VEGF and BMP-2 with peak release intervals ≥35 days; Sr-nHA provides sustained release of Sr²⁺. + It synergistically promotes osteogenic effects with BMP-2. This invention is specifically designed to assist in revision surgery of autologous bone grafts for nonunion.

Description

Technical Field

[0001] This invention relates to the field of orthopedic biomaterials, and in particular to a bone-guided regeneration membrane that can be photocured in situ during surgery, its preparation method, and its application. Background Technology

[0002] Nonunion is one of the most serious complications after fracture treatment. It is defined as the failure to achieve bony union after more than 9 months of adequate treatment, with no progress in healing for 3 consecutive months. Treatment for nonunion primarily involves revision surgery. The core strategies include removing sclerotic bone ends and fibrous tissue, implanting autologous cancellous bone, and providing a barrier environment to guide bone regeneration. While autologous bone grafting is the gold standard, its use alone faces challenges such as soft tissue invasion at the graft site, insufficient bone volume maintenance, and graft displacement. Therefore, a functional bone-guided regeneration membrane that can be used directly during surgery is urgently needed to assist in its application.

[0003] Guided bone regeneration membrane (GBR) technology has been widely used in oral and maxillofacial surgery and orthopedics. Existing GBR membranes are mainly divided into two categories: non-degradable and biodegradable membranes. Non-degradable membranes, represented by expanded polytetrafluoroethylene (ePTFE) membranes and titanium mesh, offer good mechanical support and space maintenance, but require secondary surgery for removal, increasing surgical trauma and infection risks, and have a high membrane exposure rate. Biodegradable membranes, represented by collagen membranes, have good biocompatibility and do not require secondary surgery, but suffer from limitations such as insufficient mechanical strength, a degradation rate mismatch with the bone regeneration cycle, and the inability to actively promote osteogenesis. Furthermore, existing commercially available GBR membranes are all pre-shaped products, unable to adaptively reshape according to the irregular bone surface morphology at the nonunion site during surgery, resulting in poor clinical adaptability.

[0004] In recent years, photocurable hydrogels have attracted attention in the field of orthopedic biomaterials due to their adjustable mechanical properties and good biocompatibility. Currently, there is a lack of bone-guided regeneration membranes that simultaneously possess the following characteristics: (1) They can be photocured in situ during surgery to achieve a mechanical phase transition from a soft state to a semi-rigid state to adapt to irregular bone surfaces; (2) They have the ability to self-adhere to moisten bone surfaces and do not require additional fixation devices; (3) They have an outer barrier function that resists soft tissue adhesion; (4) They are equipped with a multi-growth factor time-controlled release system to actively promote vascularization and bone regeneration; (5) Their degradation cycle matches the bone nonunion healing cycle. Summary of the Invention

[0005] To address the shortcomings of the existing technology, the present invention aims to provide a bone-guided regeneration membrane capable of intraoperative in-situ photocuring, its preparation method, and its application. This bone-guided regeneration membrane, through a three-layer functionalized structure design and a visible light in-situ phase transition mechanism, achieves comprehensive functions including intraoperative adaptive shaping, adhesion to a moistened bone surface without additional fixation, and time-controlled release of multiple growth factors to promote bone regeneration. It is specifically designed to assist in revision surgery of autologous bone grafting for nonunion. A bone-guided regeneration membrane capable of intraoperative in-situ photocuring comprises, from the bone surface side to the soft tissue side: The inner layer is a coordination coating consisting of a polymer containing catechol groups and metal ions. The main body layer comprises an interpenetrating network hydrogel with methacrylamide gelatin and methacrylamide silk fibroin as co-crosslinking components. The interpenetrating network hydrogel contains strontium-doped nano-hydroxyapatite, a visible light initiator, and polylactic acid-glycolic acid copolymer microspheres encapsulating bone growth factor. The bone growth factor is selected from at least one of vascular endothelial growth factor and bone morphogenetic protein-2. The outer layer is a non-crosslinked coating containing polysaccharide substances; The compressive modulus of the main layer without exposure to visible light at a wavelength of 405 nm is 1–20 kPa; the compressive modulus after exposure to visible light at a wavelength of 405 nm and complete curing is 30–200 kPa, and the ratio of the compressive modulus after complete curing to the compressive modulus before curing is 7–25.

[0006] The present invention also provides a method for preparing a bone-guided regeneration membrane that can be photocured in situ during surgery, comprising the following steps: S1. Using a double emulsion-solvent evaporation method, polylactic acid-glycolic acid copolymer is used as the wall material to encapsulate vascular endothelial growth factor and bone morphogenetic protein-2, respectively, to prepare a first microsphere encapsulating vascular endothelial growth factor and a second microsphere encapsulating bone morphogenetic protein-2. S2. The methacrylamide gelatin is dissolved in a buffer solution containing the visible light initiator at a temperature of 40–60°C. After cooling to room temperature, the methacrylamide silk fibroin is added and dissolved at room temperature. Then, the strontium-doped nano-hydroxyapatite, the first microsphere, and the second microsphere are added and mixed evenly to obtain the precursor solution. S3. The precursor liquid is injected into the mold to form a film, and visible light with a wavelength of 405 nm and an intensity of 5–10 mW / cm² is applied for irradiation for 10–30 seconds to obtain a hydrogel pre-crosslinked film. S4. The surface of the pre-crosslinked hydrogel membrane corresponding to the bone side is immersed in an alkaline buffer solution containing monomers with catechol groups and metal ions to form the inner layer. S5. A solution of polysaccharide is coated onto the surface of the pre-crosslinked hydrogel membrane corresponding to the soft tissue side and dried and cured to form the outer layer.

[0007] Compared with the prior art, the beneficial effects of the present invention are as follows: The compressive modulus of the ternary interpenetrating network hydrogel main layer of the present invention after photocuring can reach 96.5±12.7 kPa, which is 4.3 times that of the pure GelMA control group and 11.2 times that of the prior art CN113713179B, thus meeting the clinical needs of bone-guided regeneration membranes for mechanical support.

[0008] The PLGA dual microsphere system of this invention achieves temporal separation of VEGF and BMP-2 release peaks with an interval of at least 35 days, with a peak VEGF concentration of 58.3±5.8 ng / mL and a peak BMP-2 concentration of 18.5±1.8 ng / mL, which conforms to the biological sequence of vascularization followed by osteogenic formation.

[0009] The PDA-Ca² of the present invention + The adhesion strength of the adhesive coating on the moist bone surface reached 11.5±2.1 kPa, far exceeding the clinical minimum threshold of 5 kPa, and it also has reversible-irreversible adhesion transformation characteristics, taking into account both intraoperative positioning flexibility and postoperative fixation stability.

[0010] The degradation cycle of the main body layer of this invention is about 20 weeks, and the degradation cycle of the HA outer layer is about 5.2±0.8 weeks. The pH of the degradation solution is maintained at 7.08±0.09 throughout the degradation process, thus avoiding aseptic inflammatory reactions caused by acidic degradation products.

[0011] The Sr-nHA of the present invention can be distributed in a gradient in the main layer, with the concentration on the bone surface side being 1.5–5 times that on the outer layer side, simulating the mineralization gradient of natural bone tissue, and further guiding osteoblasts to migrate and differentiate towards the bone surface side. Detailed Implementation

[0012] The terms first, second, third, fourth, etc. (if present) in the specification and claims of this invention are used to distinguish similar objects and are not necessarily used to describe a particular order or sequence. It should be understood that such data can be interchanged where appropriate so that the embodiments described herein can be implemented in orders other than those described herein. Furthermore, the terms include or have, and any variations thereof, are intended to cover a non-exclusive inclusion; for example, a process, method, system, product, or apparatus that includes a series of steps or units is not necessarily limited to those steps or units explicitly listed, but may include other steps or units not explicitly listed or inherent to such processes, methods, products, or apparatus.

[0013] This invention uses GelMA as the matrix and SilMA as the co-crosslinking component to construct a dual-network structure that synergistically integrates a covalent crosslinking network (free radical polymerization between GelMA methacryloyl groups and between GelMA and SilMA methacryloyl groups) and a physical crosslinking network (SilMA molecular chain segments formed through β-sheet conformation). Sr-nHA is dispersed in the main layer as both an inorganic reinforcing phase and an osteogenic active phase. Compared with the pure GelMA control group, the standard embodiment of the ternary system achieved a compressive modulus of 96.5 kPa after photocuring, which is 4.3 times that of the control group; compared with the optimal compressive modulus of approximately 8.6 kPa reported in CN113713179B, this invention achieves 11.2 times that, fundamentally solving the technical bottleneck that the mechanical strength of existing photocured hydrogels is insufficient to support bone-guided regeneration.

[0014] This invention designs two types of PLGA microspheres with different degradation rates: VEGF fast-release microspheres with an LA:GA molar ratio of 50:50 and BMP-2 sustained-release microspheres with an LA:GA molar ratio of 75:25. The VEGF release peak occurs on days 5–10, promoting early vascular network reconstruction at nonunion sites; the BMP-2 release peak occurs on days 42–56, initiating bone regeneration signals after vascularization is complete. The release peaks of the two growth factors are spaced at least 35 days apart, achieving a time-controlled release strategy of vascularization followed by osteogenic formation, consistent with the biological principles of bone regeneration.

[0015] The bone-guided regeneration membrane of this invention is soft and malleable before photocuring, allowing surgeons to cut and manually shape it to adapt to irregular bone surface morphology at nonunion sites. After irradiation with 405 nm blue-violet light for 30–120 seconds, the membrane transforms in situ into a semi-rigid support state, with the compressive modulus increasing by 7–25 times before and after photocuring. This flexible-to-rigid phase transition mechanism enables the same product to possess both intraoperative malleability and postoperative mechanical support functions, eliminating the need for pre-shaping and making it suitable for nonunion sites of any shape. The 405 nm visible light source is safer than ultraviolet light sources, avoiding phototoxicity to surrounding tissues.

[0016] This invention provides a dopamine polymer and Ca²⁺ mixture on the bone-side surface of the membrane. + The formed PDA-Ca² + The ligand coating achieves an adhesion strength of ≥5 kPa in a moist bone environment. This coating exhibits a unique reversible-irreversible adhesion transition: it is reversibly adherent within 30 seconds of application, allowing the surgeon to reposition and adjust; after 5–10 minutes, it transitions to irreversible adhesion, providing stable fixation. No screws, sutures, or any additional mechanical fixation devices are required throughout the procedure, simplifying the surgical procedure and reducing complications associated with fixation devices.

[0017] This invention incorporates a non-crosslinked hyaluronic acid (HA) physical coating on the soft tissue side of the membrane, effectively preventing soft tissue cells from invading the bone regeneration area in the early postoperative period. This HA coating degrades naturally within 4–8 weeks, a degradation time window that matches the early callus formation time window at the site of nonunion. The non-crosslinked physical coating design avoids the potential cytotoxicity of chemical crosslinking agents.

[0018] This invention is specifically designed for clinical use in revision surgery for nonunion, and is fully compatible with the operational requirements of autologous cancellous bone graft implantation. Sr-nHA continuously releases Sr² during the degradation of the host layer. + Sr²⁻, with a cumulative release of 100–130 μg / mL over 84 days, forms a synergistic osteogenic effect with BMP-2: + By activating the Wnt / β-catenin signaling pathway to upregulate osteogenic gene expression, BMP-2 directly drives osteogenic differentiation through the Smad1 / 5 / 8 signaling pathway, while VEGF promotes angiogenesis in the early stages to provide nutritional support for subsequent bone regeneration. The three form a complementary sequential synergistic mechanism in the time dimension.

[0019] A bone-guided regeneration membrane capable of intraoperative in-situ light curing comprises, from the bone surface side to the soft tissue side, the following components: The inner layer is a coordination coating consisting of a polymer containing catechol groups and metal ions. The main body layer comprises an interpenetrating network hydrogel with methacrylamide gelatin and methacrylamide silk fibroin as co-crosslinking components. The interpenetrating network hydrogel contains strontium-doped nano-hydroxyapatite, a visible light initiator, and polylactic acid-glycolic acid copolymer microspheres encapsulating bone growth factor. The bone growth factor is selected from at least one of vascular endothelial growth factor and bone morphogenetic protein-2. The outer layer is a non-crosslinked coating containing polysaccharide substances; The compressive modulus of the main layer without exposure to visible light at a wavelength of 405 nm is 1–20 kPa; the compressive modulus after exposure to visible light at a wavelength of 405 nm and complete curing is 30–200 kPa, and the ratio of the compressive modulus after complete curing to the compressive modulus before curing is 7–25.

[0020] The inner layer is a coordination coating formed by dopamine polymer and calcium ions, with a thickness of 1–10 μm; The outer layer is a non-crosslinked coating of hyaluronic acid with a thickness of 5–100 μm; In the interpenetrating network hydrogel, the concentration of methacrylamide gelatin is 5%–15% by mass-volume ratio, the concentration of methacrylamide silk fibroin is 2%–10%, and the mass ratio of methacrylamide silk fibroin to methacrylamide gelatin is (0.5–2):1.

[0021] The polylactic acid-glycolic acid copolymer microspheres encapsulating bone growth factors comprise a first microsphere and a second microsphere: The first microspheres are loaded with vascular endothelial growth factor, and the molar ratio of lactic acid to glycolic acid in the polylactic acid-glycolic acid copolymer of the first microspheres is 50:50. The particle size of the first microspheres is 5–30 μm. The second microsphere encapsulates bone morphogenetic protein-2. The molar ratio of lactic acid to glycolic acid in the polylactic acid-glycolic acid copolymer of the second microsphere is 75:25. The particle size of the second microsphere is 10–50 μm.

[0022] The degree of methacrylation substitution of the methacrylamide gelatin is 40%–80%; The degree of methacrylation substitution of the methacrylated silk fibroin is 15%–40%; The interpenetrating network hydrogel includes a covalent cross-linked network formed by free radical polymerization between methacryloyl groups of the methacryloyl gelatin and between methacryloyl groups of the methacryloyl gelatin and the methacryloyl silk fibroin, and a physical cross-linked network formed by molecular chain segments of the methacryloyl silk fibroin through β-sheet conformation.

[0023] The visible light initiator is lithium phenyl-2,4,6-trimethylbenzoylphosphite, and the concentration of the visible light initiator in the interpenetrating network hydrogel is 0.1%–0.5% by mass-volume ratio. The strontium-doped nano-hydroxyapatite accounts for 0.5%–5% of the dry weight of the main layer, wherein the strontium doping molar ratio is 5%–10%, and the particle size of the strontium-doped nano-hydroxyapatite is 20–100 nm.

[0024] The strontium-doped nano-hydroxyapatite is distributed in a concentration gradient along the thickness direction within the main body layer; The concentration of the strontium-doped nano-hydroxyapatite increases from the side closer to the outer layer to the side closer to the inner layer, and the concentration of the strontium-doped nano-hydroxyapatite on the side closest to the inner layer is 1.5–5 times that on the side closest to the outer layer.

[0025] In embodiments of the present invention, a method for preparing a bone-guided regeneration membrane capable of intraoperative in-situ photocuring is included, comprising the following steps: S1. Using a double emulsion-solvent evaporation method, polylactic acid-glycolic acid copolymer is used as the wall material to encapsulate vascular endothelial growth factor and bone morphogenetic protein-2, respectively, to prepare a first microsphere encapsulating vascular endothelial growth factor and a second microsphere encapsulating bone morphogenetic protein-2. S2. The methacrylamide gelatin is dissolved in a buffer solution containing the visible light initiator at a temperature of 40–60°C. After cooling to room temperature, the methacrylamide silk fibroin is added and dissolved at room temperature. Then, the strontium-doped nano-hydroxyapatite, the first microsphere, and the second microsphere are added and mixed evenly to obtain the precursor solution. S3. The precursor liquid is injected into the mold to form a film, and visible light with a wavelength of 405 nm and an intensity of 5–10 mW / cm² is applied for irradiation for 10–30 seconds to obtain a hydrogel pre-crosslinked film. S4. The surface of the pre-crosslinked hydrogel membrane corresponding to the bone side is immersed in an alkaline buffer solution containing monomers with catechol groups and metal ions to form the inner layer. S5. A solution of polysaccharide is coated onto the surface of the pre-crosslinked hydrogel membrane corresponding to the soft tissue side and dried and cured to form the outer layer.

[0026] In step S1, the molar ratio of lactic acid to glycolic acid in the polylactic acid-glycolic acid copolymer wall material used in preparing the first microsphere is 50:50; the molar ratio of lactic acid to glycolic acid in the polylactic acid-glycolic acid copolymer wall material used in preparing the second microsphere is 75:25. In step S3, the specific process of injecting the precursor liquid into the mold for casting includes: Multiple precursor solutions with progressively increasing concentrations of strontium-doped nano-hydroxyapatite were prepared. According to the order of increasing concentration of strontium-doped nano-hydroxyapatite in the precursor liquid, each portion of the precursor liquid is injected layer by layer into the mold for casting. After each of the precursor liquid sub-liquids has completed the casting operation, the entire film is irradiated with visible light.

[0027] In step S4, the pH of the alkaline buffer solution containing the monomer with catechol groups and metal ions is 8.0–8.5, the concentration of dopamine hydrochloride in the buffer solution is 1–5 mg / mL, the concentration of calcium chloride is 5–20 mM, and the immersion reaction time is 12–24 hours.

[0028] The present invention relates to the application of the intraoperatively photocurable bone-guided regeneration membrane in the preparation of medical devices for autologous bone grafting surgery, wherein the autologous bone grafting surgery is used to treat nonunion or delayed fracture healing, including tibial nonunion, femoral nonunion, humeral nonunion, delayed fracture healing or segmental bone defects.

[0029] The present invention will be further described in detail below with reference to specific embodiments. These embodiments are for illustrative purposes only and are not intended to limit the scope of the invention.

[0030] Materials and reagents Methacrylamide gelatin (GelMA, 60% methacrylation degree, approximately 50 kDa) was purchased from Suzhou Yongqinquan Intelligent Equipment Co., Ltd. (EFL).

[0031] Methacrylated silk fibroin (SilMA, 25% or 30% methacrylation substitution) was purchased from Suzhou Yongqinquan Intelligent Equipment Co., Ltd. or prepared according to the methods described in the literature.

[0032] Strontium-doped nano-hydroxyapatite (Sr-nHA, 8% or 10% strontium doping molar ratio, particle size 40–60 nm) was prepared by wet chemical co-precipitation.

[0033] Polylactic acid-glycolic acid copolymer (PLGA, molecular weight 50,000 Da) was selected from two types with LA:GA molar ratios of 50:50 and 75:25, and was purchased from Shandong Daigang Biotechnology Co., Ltd.

[0034] Recombinant human vascular endothelial growth factor (rhVEGF) and recombinant human bone morphogenetic protein-2 (rhBMP-2) were purchased from PeproTech, Inc., USA.

[0035] The phenyl-2,4,6-trimethylbenzoyl lithium hypophosphite (LAP) was purchased from Suzhou Yongqinquan Intelligent Equipment Co., Ltd.

[0036] Dopamine hydrochloride was purchased from Sigma-Aldrich, USA. Calcium chloride (CaCl2, analytical grade) was purchased from Sinopharm Chemical Reagent Co., Ltd.

[0037] Hyaluronic acid (HA, molecular weight 100–300 kDa) was purchased from Shandong Huaxi Haiyu Biotechnology Co., Ltd.

[0038] Polyvinyl alcohol (PVA, molecular weight 30,000–70,000 Da) was purchased from Sigma-Aldrich, USA. Tris-HCl buffer (pH 8.0–8.5) PBS buffer (pH 7.4) was prepared according to standard methods.

[0039] The control group used pure GelMA hydrogel membranes, free of SilMA, Sr-nHA, and PLGA microspheres. GelMA was dissolved in PBS buffer containing 0.25% (w / v) LAP in a 50°C water bath, resulting in a final GelMA concentration of 10% (w / v). After cooling to room temperature, the membrane was cast into a mold to a thickness of 0.8 mm. The bone side of the membrane was coated with PDA-Ca² as per step S5. + The adhesion layer (dopamine hydrochloride concentration 2 mg / mL, CaCl2 concentration 10 mM, reaction time 18 hours) was not coated with the HA outer layer. The parameters of the control group are shown in Table 1.

[0040] Example 1 (Standard Formulation) The material composition parameters of Example 1 are shown in Table 1: GelMA concentration 8% (w / v), SilMA concentration 4% (w / v), SilMA to GelMA mass ratio 0.5:1, Sr-nHA accounts for 2% (w / w) of the dry weight of the main layer, LAP concentration 0.25% (w / v), and film thickness 0.8 mm.

[0041] Step S1: Preparation of VEGF fast-release microspheres. VEGF-PLGA microspheres were prepared using a W1 / O / W2 double emulsion-solvent evaporation method. Inner aqueous phase W1: rhVEGF... 165 The growth factor was dissolved in 0.1% (w / v) BSA-PBS solution at a concentration of 100 μg / mL. Oil phase O: PLGA (molecular weight 50,000 Da) with a LA:GA ratio of 50:50 was dissolved in dichloromethane at a concentration of 5% (w / v). W1 was injected into phase O and homogenized at 10,000 rpm for 30 seconds to form a primary emulsion W1 / O. Aqueous phase W2: 1% (w / v) PVA aqueous solution. The W1 / O primary emulsion was poured into W2 and mechanically stirred at 3,000 rpm for 2 minutes to form a W1 / O / W2 biemulsion. The dichloromethane was evaporated by continuous stirring at room temperature for 4 hours. The microspheres were collected by centrifugation, washed three times with deionized water, and freeze-dried. The resulting VEGF-PLGA microspheres had a particle size of 5–30 μm and an encapsulation efficiency of 63.2 ± 2.8%.

[0042] Step S2: Preparation of BMP-2 sustained-release microspheres. The method is the same as in S1, but PLGA is replaced with LA:GA (75:25 ratio, molecular weight 50,000 Da), and the concentration of rhBMP-2 in the aqueous phase is 200 μg / mL. The resulting BMP-2-PLGA microspheres have a particle size of 10–50 μm and an encapsulation efficiency of 57.6 ± 2.4%. Adjusting the LA:GA ratio from 50:50 to 75:25 increases the crystallinity of PLGA and reduces its degradation rate, thereby achieving delayed release of BMP-2.

[0043] Step S3: Preparation of the precursor solution. Dissolve GelMA in PBS buffer containing 0.25% (w / v) LAP in a 50°C water bath to a final GelMA concentration of 8% (w / v), and stir until completely dissolved. After cooling to room temperature (25°C), add SilMA powder (final concentration 4% (w / v)) and stir to dissolve at room temperature. Room temperature dissolution of SilMA, rather than heating, is used here to avoid damaging the β-sheet conformation of the silk fibroin molecular chain at high temperatures. Add Sr-nHA nanoparticles (2% (w / w) dry weight, 8% strontium doping molar ratio, particle size 40–60 nm), and sonicate for 10 minutes (100 W power, ice bath conditions) to ensure uniform suspension of the nanoparticles. Add the VEGF-PLGA microspheres (10 mg / mL) prepared in step S1 and the BMP-2-PLGA microspheres (10 mg / mL) prepared in step S2, and gently stir to mix thoroughly to obtain the precursor solution.

[0044] Step S4: Casting. The precursor solution is injected into a polytetrafluoroethylene (PTFE) flat mold for casting, with the film thickness controlled to 0.8 mm using a doctor blade. Under light-protected conditions, selective low-intensity pre-illumination (405 nm, 5–10 mW / cm², 10–30 seconds) can be applied to partially pre-crosslink the film, giving it a certain initial operational strength (able to be picked up with tweezers without breaking) while maintaining a soft and malleable state. The compressive modulus of the pre-illuminated film is controlled within the range of 1–20 kPa.

[0045] Step S5: PDA adhesion layer coating. Prepare Tris-HCl buffer (pH 8.5), add dopamine hydrochloride (final concentration 2 mg / mL) and CaCl2 (final concentration 10 mM). Immerse the bone-side down membrane obtained in Step S4 into the above solution and react at room temperature in the dark for 18 hours. Dopamine self-polymerizes under alkaline conditions to form polydopamine (PDA) and deposits on the membrane surface, while Ca²⁺… + It forms a coordination bond with the catechin group of PDA to form PDA-Ca² + Adhesive coating. After the reaction, the coating was gently rinsed three times with deionized water to remove unreacted dopamine and loosely attached PDA particles. The resulting PDA coating thickness was 4.2 ± 0.7 μm, and the wet adhesion strength was 11.5 ± 2.1 kPa. The coating provided reversible adhesion within 30 seconds of application, allowing for intraoperative repositioning; after 5–10 minutes, due to PDA-Ca² + Further cross-linking of the coordination network and interfacial dehydration effects lead to irreversible adhesion.

[0046] Step S6: HA Anti-adhesion Layer Coating. Hyaluronic acid was dissolved in deionized water to prepare a 1.5% (w / v) HA solution. The HA solution was uniformly coated onto the soft tissue side of the membrane using a coating stick and allowed to dry and cure at room temperature, forming a 28.5 ± 4.2 μm thick HA non-crosslinked physical coating. This coating adheres to the surface of the host layer through physical adsorption, without chemical crosslinking, thus avoiding the potential cytotoxicity of crosslinking agents.

[0047] Step S7: Sterilization and Cutting. Cut the coated film to the required size and sterilize it with ethylene oxide (EtO). After sterilization, the retention rates of VEGF and BMP-2 bioactivity are detected by ELISA. The film is released only if the retention rates of both growth factors are ≥80%. After aseptic packaging, store at 4°C protected from light.

[0048] Example 2 (High SilMA Formulation) Example 2 enhances the contribution of the β-sheet physical crosslinking network by increasing the proportion of SilMA and decreasing the proportion of GelMA. Specific material parameters are shown in Table 1: GelMA concentration 6% (w / v), SilMA concentration 8% (w / v), SilMA to GelMA mass ratio 1.33:1, SilMA substitution degree 25%, Sr-nHA 2% (w / w) of dry weight, strontium doping molar ratio 8%, and other parameters (LAP, microsphere addition amount, PDA coating conditions, HA coating conditions) are the same as in Example 1. Preparation steps S1–S7 are the same as in Example 1.

[0049] Example 2 was designed to verify the enhancing effect of a high SilMA content on the mechanical properties after photocuring. Since SilMA molecular segments spontaneously form β-sheet physical crosslinks at room temperature, high SilMA formulations are expected to achieve higher compressive modulus and slower degradation rate after photocuring.

[0050] Example 3 (High Sr-nHA formulation) Example 3 enhances the inorganic enhancement effect and Sr² by increasing the Sr-nHA content and strontium doping ratio. + Release amount. Specific material parameters are shown in Table 1: GelMA concentration 8% (w / v), SilMA concentration 4% (w / v), SilMA substitution degree 30%, Sr-nHA 5% (w / w) of dry weight, Strontium doping molar ratio 10%, film thickness 1.0 mm, and other parameters are the same as in Example 1. Preparation steps S1–S7 are the same as in Example 1.

[0051] The purpose of Example 3 is to verify the effect of high inorganic phase content on mechanical properties and Sr². +Impact on sustained release capability. The degree of SilMA substitution was increased from 25% to 30% to increase the covalent crosslinking density between SilMA and GelMA. The film thickness was increased from 0.8 mm to 1.0 mm, making it suitable for clinical scenarios with large bone defect volumes.

[0052] Example 4 (Gradient nHA formulation) Example 4 constructed a gradient distribution of Sr-nHA in the host layer to simulate the mineralization gradient of natural bone tissue from cancellous bone to compact bone. Specific material parameters are shown in Table 1: GelMA concentration 8% (w / v), SilMA concentration 4% (w / v), SilMA substitution degree 25%, total Sr-nHA content 3% (w / w) of dry weight, Strontium doping molar ratio 8%, and film thickness 0.8 mm.

[0053] The gradient casting process is as follows: Four precursor solutions with increasing Sr-nHA content were prepared, with Sr-nHA concentrations of 1.2%, 2.3%, 3.6%, and 4.8% (w / w, corresponding to the contents of layers 1 to 4 in Table 5). The PTFE mold was placed horizontally, and the lowest concentration sub-solution (1.2%, corresponding to layer 1 on the outer side) was injected first. The thickness of this layer was controlled by a scraper to be approximately 1 / 4 of the total film thickness (approximately 0.2 mm). After allowing it to stand at room temperature for 2 minutes to allow the surface to stabilize initially, 2.3% (layer 2), 3.6% (layer 3), and 4.8% (layer 4, bone side) sub-solutions were injected sequentially, with each layer having a thickness of approximately 0.2 mm. After the four layers were cast sequentially, a low-intensity pre-lighting (5–10 mW / cm², 20 seconds) was applied to the entire structure to achieve interfacial fusion between the layers through partial crosslinking of GelMA and SilMA. The remaining steps S5–S7 were the same as in Example 1.

[0054] The EDS line scan analysis of Example 4 is shown in Table 5. The Sr-nHA content on the bone surface (layer 4) was 4.8 ± 0.5 wt%, and on the outer layer (layer 1) it was 1.2 ± 0.2 wt%, with a gradient factor of 4.0, which falls within the 1.5–5 times range defined in claim 6. The Ca / P molar ratio at each layer remained stable at 1.57–1.59, close to the theoretical stoichiometry of 1.67, confirming that the crystal structure of Sr-nHA remained intact during the gradient casting process.

[0055] Example 5 (Alternative Adhesive Coating Solution) PDA-Ca² in step S5 + The system was replaced with tannic acid-Fe³ + (TA-Fe³) + Coordination system. Specific procedure: Immerse the bone-side side of the membrane in a Tris-HCl buffer (pH 7.4) containing tannic acid (5 mg / mL) and FeCl3 (2 mM), and react at room temperature for 6 hours to form TA-Fe³⁺.+ Coordination adhesion coating. This alternative also utilizes phenolic hydroxyl-metal ion coordination bonds to achieve wet surface adhesion, with an adhesion strength of 4–8 kPa. The remaining material composition and preparation steps are the same as in Example 1.

[0056] Performance testing methods (1) Compression modulus test. Each group of samples (before and after photocuring) was cut into circular pieces with a diameter of 10 mm and a thickness consistent with the prepared film thickness (n=5). Uniaxial unconstrained compression tests were performed at room temperature using a universal testing machine (Instron 5944 or equivalent) at a compression rate of 1 mm / min. The compression modulus was taken as the slope of the linear region of the stress-strain curve (strain range of 5%–15%). Samples before photocuring were tested directly; samples after photocuring were tested after being irradiated with 405 nm blue-violet light (light intensity 30 mW / cm²) until fully cured. The modulus increase was calculated as follows: Modulus increase = Compression modulus after curing / Compression modulus before curing.

[0057] (2) Determination of photocuring time. A rotational rheometer (Anton Paar MCR 302 or equivalent) was used with a parallel plate fixture (20 mm in diameter). The oscillation mode was set (frequency 1 Hz, strain 1%), and the measurement was performed at a constant temperature of 37°C. The precursor liquid was placed on the lower plate, and 405 nm light (light intensity 30 mW / cm²) was turned on. The change of storage modulus G' over time was recorded simultaneously. The photocuring time was defined as the time required for G' to reach 90% of the plateau value (n=5).

[0058] (3) Determination of moisture content and porosity. Moisture content test: After photocuring, the sample was equilibrated and swollen in PBS at 37°C for 24 hours, and the wet weight Ww was measured. Then, it was freeze-dried to constant weight and the dry weight Wd was measured (n=5). Moisture content = (Ww) / (Wd) Wd) / Ww ×100%. Porosity test: Liquid displacement method was used, with anhydrous ethanol as the displacement solution (n=5). Average pore size test: The cross-section of the freeze-dried sample was observed using a scanning electron microscope (SEM, Hitachi S-4800 or equivalent model), and the average value of more than 50 pore sizes was measured using ImageJ software.

[0059] (4) PDA Wet Adhesion Strength Test. The lap shear method was used. The PDA coated side was attached to the surface of fresh pig bone (wetted with physiological saline), a pre-pressure of 5 N was applied and maintained for 30 seconds, followed by immersion in physiological saline at 37°C for 10 minutes. The lap shear test was performed using a universal testing machine at a rate of 1 mm / min. Adhesion strength = maximum shear force / adhesion area (n=5). The PDA coating thickness was measured using cross-sectional SEM (n=5).

[0060] (5) Cumulative release rate test of VEGF and BMP-2. Membrane samples (approximately 10 mm × 10 mm) from each embodiment were placed in 24-well plates containing 1 mL of PBS (pH 7.4) and incubated at 37°C (n=3). Every 3 days, all supernatant was collected and an equal volume of fresh PBS was added. The release concentration at each time point was quantitatively determined using a human VEGF ELISA kit and a human BMP-2 ELISA kit (PeproTech). Cumulative release rate = Total cumulative release up to that time point / Initial encapsulation volume × 100%. Monitoring continued for 84 days.

[0061] (6) Sr² + Cumulative release test. The sample soaking conditions were the same as in (5), and the supernatant was analyzed by inductively coupled plasma atomic emission spectrometry (ICP-OES, Agilent 5110 or equivalent) to detect Sr². + Concentration. Sr² + The cumulative release amount is the sum of the release amounts at each time point (n=3). Monitoring lasted for 84 days.

[0062] (7) In vitro degradation test. The initial dry weight W0 of each group of samples (approximately 10 mm × 10 mm, after photocuring) was weighed and placed in a sealed centrifuge tube containing 2 mL of degradation solution (PBS, pH 7.4, containing 2 U / mL type I collagenase), and incubated at 37°C (n=3). The degradation solution was changed every 3 days, and the pH value was recorded. Samples were removed at preset time points (weeks 2, 4, 6, 8, 10, 12, 16, 20, and 24), washed three times with deionized water, freeze-dried to constant weight, and the residual dry weight Wt was measured. Residual mass = Wt / W0 × 100%. The complete degradation time of the HA outer layer was defined as the time point at which the HA coating completely disappeared in SEM cross-sectional observation.

[0063] (8) Sr-nHA gradient distribution EDS analysis (specifically for Example 4). The freeze-dried sample of Example 4 was cut into sections along the thickness direction, sputter-coated with gold, and then subjected to cross-sectional line scanning analysis using a scanning electron microscope-energy dispersive spectroscopy (SEM-EDS, Hitachi S-4800 / Oxford X-MaxN). The film thickness direction was divided into four layers (layer 1: outer 1 / 4; layer 2: middle 2 / 4; layer 3: middle 3 / 4; layer 4: bone surface 4 / 4). The Sr-nHA content (wt%), Ca content (wt%), P content (wt%), and Sr content (wt%) of each layer were analyzed, and the Ca / P molar ratio and Sr / (Sr+Ca) molar percentage were calculated (n=3).

[0064] Test Results (1) Mechanical and photocuring properties. The test results of the mechanical and photocuring properties of each group of samples are shown in Table 1.

[0065] Table 1. Mechanical and photocuring properties (n=5, mean±SD) As shown in Table 1, the compressive modulus of all examples containing SilMA and Sr-nHA (Examples 1–4) after photocuring was significantly higher than that of the pure GelMA control group. Example 1 (standard formulation) achieved a compressive modulus of 96.5 ± 12.7 kPa after photocuring, which was 4.3 times that of the control group (22.3 ± 3.8 kPa). Example 2 (high SilMA formulation) achieved a compressive modulus of 186.2 ± 21.4 kPa after photocuring, the highest value among all examples, indicating that increasing the SilMA proportion can significantly improve mechanical properties by increasing the β-sheet physical crosslinking density. Example 3 (high Sr-nHA formulation) achieved a compressive modulus of 158.7 ± 18.3 kPa after photocuring, confirming the reinforcing effect of inorganic nanoparticles. The compressive modulus of all embodiments before photocuring was in the range of 1–20 kPa (5.6–8.3 kPa), and the compressive modulus after photocuring was in the range of 30–200 kPa (96.5–186.2 kPa), with a modulus increase of 17.2–22.4 times, all within the range of 7–25 times as defined in claim 1.

[0066] Regarding the photocuring time, Example 2 (35±4 s) had the shortest time, followed by Example 1 (48±5 s), both significantly lower than the 120-second upper limit specified in claim 7. The control group had the longest photocuring time (72±8 s), because the pure GelMA system lacked the β-sheet pre-assembly assistance of SilMA, and the formation of the cross-linked network relied entirely on free radical polymerization. Water content and porosity decreased with increasing SilMA and Sr-nHA content. Example 2 had the lowest water content (73.6±2.1%), while the control group had the highest (88.2±1.5%), reflecting the effect of increased cross-linking density on water retention capacity in the dual-network system. The average pore size of each example was 78–112 μm, which is beneficial for osteoblast migration and proliferation.

[0067] Regarding PDA wet adhesion strength, all groups exceeded the minimum threshold of 5 kPa defined in claim 2, with Example 2 showing the highest (12.8 ± 2.3 kPa). This may be related to the denser surface of the substrate layer and more uniform PDA coating deposition in the high SilMA formulation. The PDA coating thickness ranged from 3.8 to 4.5 μm, all falling within the range of 1–10 μm defined in claim 2. The HA coating thickness ranged from 27.8 to 30.1 μm, falling within the range of 5–100 μm defined in claim 2.

[0068] (2) Drug release kinetics. The cumulative release rate test results of VEGF and BMP-2 are shown in Table 2.

[0069] Table 2. Drug release kinetics—cumulative release rate (n=3, mean±SD, %) As shown in Table 2, VEGF from PLGA fast-release microspheres with LA:GA=50:50 exhibits a typical burst-release-sustained-release curve: 12.3±1.8% is released on day 1 (Example 1), 52.4±3.2% is released cumulatively on day 7, 78.5±3.8% on day 14, and 93.1±2.1% on day 28. The release rate then plateaus, reaching a cumulative release rate of 98.9±0.5% on day 84. BMP-2 from PLGA sustained-release microspheres with LA:GA=75:25 exhibits a delayed-release curve: only 7.8±0.9% is released cumulatively on day 7 (Example 1), 32.7±2.1% on day 28, 58.5±3.4% on day 42 (at which point the cumulative release rate exceeds half, and the release rate is fastest), and 93.2±1.8% on day 84. The cumulative release rate of VEGF exceeded 50% by day 7, while that of BMP-2 exceeded 50% by day 42. The time difference between the two reaching 50% release rate was 35 days, consistent with the release peak time of at least 35 days as defined in claim 3. The encapsulation rates of VEGF in each embodiment were 62.8%–64.1% (all ≥60%, satisfying claim 3), and the encapsulation rates of BMP-2 were 56.9%–58.3% (all ≥55%, satisfying claim 3). The release behavior among the four embodiments was highly consistent, indicating that changes in the composition of the host layer had no significant impact on the release kinetics within the PLGA microspheres.

[0070] (3) ELISA three-factor release concentrations. VEGF, BMP-2, and Sr² in Example 1 + The results of the release concentration test are shown in Table 3.

[0071] Table 3. ELISA three-factor release concentrations – representative data from Example 1 (n=3, mean±SD) As shown in Table 3, the VEGF release concentration peaked at 58.3 ± 5.8 ng / mL on day 7, and then gradually decreased, dropping to 9.5 ± 1.6 ng / mL on day 28 and only 0.3 ± 0.1 ng / mL on day 84. The BMP-2 release concentration peaked at 18.5 ± 1.8 ng / mL on day 42, and then slowly decreased, reaching 16.2 ± 1.5 ng / mL on day 56 and 6.5 ± 0.8 ng / mL on day 84. The VEGF peak occurred on day 7 and the BMP-2 peak occurred on day 42, with a peak interval of 35 days, satisfying the requirements of claim 9 (VEGF peak on days 5–10, measured day 7), BMP-2 peak on days 42–56, measured day 42), and the peak interval of claim 3 (≥35 days, measured day 35). The peak concentration of VEGF, 58.3 ng / mL, falls within the range of 45–70 ng / mL as defined in claim 9; the peak concentration of BMP-2, 18.5 ng / mL, falls within the range of 15–20 ng / mL as defined in claim 9.

[0072] Sr² + The release of Sr² showed a continuous linear cumulative trend, with a cumulative release of 25.8 ± 2.8 μg / mL on day 7, reaching 101.8 ± 6.8 μg / mL on day 42, and 125.3 ± 7.8 μg / mL on day 84, falling within the range of 100–130 μg / mL defined in claim 17. + The sustained release of Sr² overlaps temporally with the delayed release of BMP-2: the peak release period of BMP-2 (Day 28–56) coincides precisely with the release of Sr². + The rapid growth phase of cumulative release (82.5–113.5 μg / mL) allows both osteogenic factors to function simultaneously within the critical time window for nonunion repair.

[0073] (4) Degradation performance. The in vitro degradation test results of each group of samples are shown in Table 4.

[0074] Table 4. Degradation performance (n=3, mean±SD, mass residue %) As shown in Table 4, the control group (pure GelMA) exhibited the fastest degradation rate, with a residual mass of only 5.6 ± 1.5% at week 10, indicating complete degradation around week 10. Example 1 (standard formulation) showed a significantly slower degradation rate, with a residual mass of 40.3 ± 3.5% at week 10, indicating complete degradation around week 20. Examples 2 (high SilMA formulation) and 3 (high Sr-nHA formulation) showed the slowest degradation, with complete degradation around week 24. This difference in degradation rate stems from the dual inhibitory effect of β-sheet physical cross-linking in the ternary interpenetrating network and the Sr-nHA nanoparticles on collagenase hydrolysis: the β-sheet crystalline regions of SilMA exhibit high resistance to enzymatic degradation, while the Sr-nHA nanoparticles filling the network pores delay enzyme molecule penetration.

[0075] The outer layer of HA was completely degraded in all embodiments within 5.0–5.5 weeks. This time window corresponds to the early callus formation stage in nonunion sites (4–8 weeks post-surgery), indicating that the outer layer of HA provides early anti-adhesion protection and then degrades naturally, no longer hindering normal integration of soft tissue with the membrane surface. At week 12 in all embodiments, the pH of the degradation solution remained between 7.08 and 7.15, close to the physiological pH (7.4), significantly better than the control group's 6.82 ± 0.12. This indicates that the degradation products of the ternary system are less acidic, and the alkaline ions (Ca²⁺) released during Sr-nHA degradation are less concentrated. + Sr² + It acts as a pH buffer, which helps to avoid aseptic inflammatory reactions caused by acidic degradation products.

[0076] (5) Sr-nHA gradient distribution EDS analysis. The cross-sectional EDS line scan analysis results of Example 4 are shown in Table 5.

[0077] Table 5. EDS analysis of Sr-nHA gradient distribution – Example 4 (n=3, mean±SD) As shown in Table 5, the Sr-nHA content in Example 4 increased from 1.2 ± 0.2 wt% on the outer layer (layer 1) to 4.8 ± 0.5 wt% on the bone surface (layer 4). The concentration on the bone surface was 4.0 times that on the outer layer, falling within the 1.5–5 times range defined in claim 6. The Ca and P contents showed the same increasing trend, with Ca increasing from 0.42 ± 0.06 wt% to 1.71 ± 0.15 wt% and P increasing from 0.19 ± 0.03 wt% to 0.77 ± 0.07 wt%. The Ca / P molar ratio at each layer remained stable within the range of 1.57–1.59, close to the theoretical Ca / P molar ratio of 1.67 for stoichiometric hydroxyapatite, confirming that the gradient casting process did not damage the crystal structure integrity of Sr-nHA. The Sr / (Sr+Ca) molar percentage remained stable across all layers (7.5%–8.1%), indicating that the Sr doping ratio did not segregate due to gradient casting, and that the Sr-nHA nanoparticles maintained a consistent chemical composition across all layers.

[0078] Results Discussion The GelMA-SilMA-Sr-nHA ternary interpenetrating network hydrogel bone-guided regeneration membrane of the present invention achieves comprehensive performance that cannot be achieved by existing technologies in four dimensions: mechanical properties, drug controlled release, degradation behavior, and bone surface adhesion.

[0079] In terms of mechanical properties, the core advantage of the ternary system lies in the synergistic enhancement mechanism of the covalent-physical dual network. The methacryloyl groups of GelMA undergo free radical polymerization under LAP photoinitiation to form a covalent cross-linked network (label 19), while the SilMA molecular chain segments spontaneously fold to form β-sheet physical cross-linked domains (label 20). The two networks interpenetrate to form an interpenetrating network structure. This dual cross-linking mechanism enables the compressive modulus of Example 1 to reach 96.5 kPa after photocuring, which is 4.3 times that of the pure GelMA control group (22.3 kPa). Compared with the optimal compressive modulus (approximately 8.6 kPa) of the GelMA-HAMA-SilMA system reported in CN113713179B, the present invention achieves 11.2 times that. The reason for such a significant difference is that CN113713179B uses HAMA (hyaluronic acid methacrylamide) as the third component. HAMA has strong hydrophilicity but limited mechanical contribution. In contrast, this invention replaces HAMA with Sr-nHA nano-inorganic phase, which not only improves the matrix stiffness through the physical filling and reinforcement effect of nanoparticles, but also further strengthens the network structure through the interfacial interaction between the nanoparticle surface and GelMA / SilMA.

[0080] Regarding controlled drug release, the PLGA dual-rate microsphere system successfully achieved the time-sequential release of VEGF and BMP-2. As shown in Table 6, the peak concentration of VEGF (58.3 ng / mL) occurred on day 7, and the peak concentration of BMP-2 (18.5 ng / mL) occurred on day 42, with a peak interval of 35 days. The biological basis for this time-sequential design is that the first stage of nonunion repair (1–2 weeks post-surgery) is dominated by vascularization, requiring high concentrations of VEGF to promote the establishment of angiogenesis and provide nutrition and oxygen for subsequent bone regeneration; the second stage (4–8 weeks post-surgery) is dominated by osteogenic differentiation, requiring BMP-2 to drive mesenchymal stem cells to differentiate into osteogenic cells and form new bone matrix. If both factors are released simultaneously, high concentrations of BMP-2 may induce heterotopic ossification or low osteogenic efficiency under conditions of insufficient blood supply. The time-sequential controlled release strategy of this invention ensures the physiological sequence of vascularization first, followed by osteogenic formation.

[0081] Sr² + Regarding the synergistic effect of -BMP-2, as shown in Table 6, Sr² + A cumulative release of 125.3 μg / mL was achieved over 84 days, exhibiting a continuous linear release trend. During the peak BMP-2 release period (Days 28–56), Sr² + The cumulative release increased from 82.5 μg / mL to 113.5 μg / mL, with a high degree of overlap in the time window between the two increases. Sr² + By activating the Wnt / β-catenin signaling pathway to upregulate the expression of osteogenic transcription factors such as Runx2 and Osx, and forming a dual-pathway synergy with the direct osteogenic driving effect of BMP-2 through the Smad1 / 5 / 8 signaling pathway, it is expected to achieve osteogenic effects superior to those of BMP-2 release alone.

[0082] Regarding the PDA adhesive layer, PDA-Ca² + The adhesion strength of the coordination coating on the moist bone surface ranged from 9.2 to 12.8 kPa (Table 2), all significantly exceeding the minimum required threshold of 5 kPa. The reversible-irreversible adhesion transition of this coating has significant clinical value: the surgeon can freely peel it off and reposition it within 30 seconds of application (reversible adhesion), avoiding the operational pressure of traditional one-time application and positioning; after 5–10 minutes, PDA-Ca²... + The coordination network further matures under the action of interfacial dehydration, transforming into an irreversible adhesion that eliminates the need for screws, sutures, or any additional mechanical fixation devices. This characteristic greatly simplifies the entire intraoperative procedure (U1 debridement → U2 bone grafting → U3 adhesion → U4 shaping → U5 light curing), reducing surgical difficulty and time.

[0083] Regarding degradation behavior, the main body layer of Example 1 was completely degraded in approximately 20 weeks, while Examples 2 and 3 were completely degraded in approximately 24 weeks. This degradation period matches the healing cycle of nonunion (typically 12–24 weeks). During degradation, the pH value remained near neutral (7.08–7.15), avoiding the accumulation of acidic degradation products commonly found in polylactic acid materials. This is attributed to the Ca²⁺ released during the degradation of Sr-nHA. + and Sr² + Basic ions have a neutralizing and buffering effect on acidic products.

[0084] The above embodiments are only used to illustrate the technical solutions of the present invention, and are not intended to limit it. Although the present invention has been described in detail with reference to the foregoing embodiments, those skilled in the art should understand that modifications can still be made to the technical solutions described in the foregoing embodiments, or equivalent substitutions can be made to some of the technical features. Such modifications or substitutions do not cause the essence of the corresponding technical solutions to deviate from the spirit and scope of the technical solutions of the embodiments of the present invention.

Claims

1. A bone-guided regeneration membrane capable of intraoperative in-situ light curing, characterized in that, From the bone side to the soft tissue side, the following are included in sequence: The inner layer is a coordination coating consisting of a polymer containing catechol groups and metal ions. The main body layer comprises an interpenetrating network hydrogel with methacrylamide gelatin and methacrylamide silk fibroin as co-crosslinking components. The interpenetrating network hydrogel contains strontium-doped nano-hydroxyapatite, a visible light initiator, and polylactic acid-glycolic acid copolymer microspheres encapsulating bone growth factor. The bone growth factor is selected from at least one of vascular endothelial growth factor and bone morphogenetic protein-2. The outer layer is a non-crosslinked coating containing polysaccharide substances; The compressive modulus of the main layer without exposure to visible light at a wavelength of 405 nm is 1–20 kPa; the compressive modulus after exposure to visible light at a wavelength of 405 nm and complete curing is 30–200 kPa, and the ratio of the compressive modulus after complete curing to the compressive modulus before curing is 7–25.

2. The bone-guided regeneration membrane capable of intraoperative in-situ light curing according to claim 1, characterized in that: The inner layer is a coordination coating formed by dopamine polymer and calcium ions, with a thickness of 1–10 μm; The outer layer is a non-crosslinked coating of hyaluronic acid with a thickness of 5–100 μm; In the interpenetrating network hydrogel, the concentration of methacrylamide gelatin is 5%–15% by mass-volume ratio, the concentration of methacrylamide silk fibroin is 2%–10%, and the mass ratio of methacrylamide silk fibroin to methacrylamide gelatin is (0.5–2):

1.

3. The bone-guided regeneration membrane capable of intraoperative in-situ photocuring according to claim 1 or 2, characterized in that, The polylactic acid-glycolic acid copolymer microspheres encapsulating bone growth factors comprise a first microsphere and a second microsphere: The first microspheres are loaded with vascular endothelial growth factor, and the molar ratio of lactic acid to glycolic acid in the polylactic acid-glycolic acid copolymer of the first microspheres is 50:

50. The particle size of the first microspheres is 5–30 μm. The second microsphere encapsulates bone morphogenetic protein-2. The molar ratio of lactic acid to glycolic acid in the polylactic acid-glycolic acid copolymer of the second microsphere is 75:

25. The particle size of the second microsphere is 10–50 μm.

4. The bone-guided regeneration membrane capable of intraoperative in-situ light curing according to claim 2, characterized in that: The degree of methacrylation substitution of the methacrylamide gelatin is 40%–80%; The degree of methacrylation substitution of the methacrylated silk fibroin is 15%–40%; The interpenetrating network hydrogel includes a covalent cross-linked network formed by free radical polymerization between methacryloyl groups of the methacryloyl gelatin and between methacryloyl groups of the methacryloyl gelatin and the methacryloyl silk fibroin, and a physical cross-linked network formed by molecular chain segments of the methacryloyl silk fibroin through β-sheet conformation.

5. The bone-guided regeneration membrane capable of intraoperative in-situ light curing according to claim 1, characterized in that: The visible light initiator is lithium phenyl-2,4,6-trimethylbenzoylphosphite, and the concentration of the visible light initiator in the interpenetrating network hydrogel is 0.1%–0.5% by mass-volume ratio. The strontium-doped nano-hydroxyapatite accounts for 0.5%–5% of the dry weight of the main layer, wherein the strontium doping molar ratio is 5%–10%, and the particle size of the strontium-doped nano-hydroxyapatite is 20–100 nm.

6. The bone-guided regeneration membrane capable of intraoperative in-situ light curing according to claim 5, characterized in that: The strontium-doped nano-hydroxyapatite is distributed in a concentration gradient along the thickness direction within the main body layer; The concentration of the strontium-doped nano-hydroxyapatite increases from the side closer to the outer layer to the side closer to the inner layer, and the concentration of the strontium-doped nano-hydroxyapatite on the side closest to the inner layer is 1.5–5 times that on the side closest to the outer layer.

7. A method for preparing a bone-guided regeneration membrane capable of intraoperative in-situ photocuring as described in any one of claims 1 to 6, characterized in that, Includes the following steps: S1. Using a double emulsion-solvent evaporation method, polylactic acid-glycolic acid copolymer is used as the wall material to encapsulate vascular endothelial growth factor and bone morphogenetic protein-2, respectively, to prepare a first microsphere encapsulating vascular endothelial growth factor and a second microsphere encapsulating bone morphogenetic protein-2. S2. The methacrylamide gelatin is dissolved in a buffer solution containing the visible light initiator at a temperature of 40–60°C. After cooling to room temperature, the methacrylamide silk fibroin is added and dissolved at room temperature. Then, the strontium-doped nano-hydroxyapatite, the first microsphere, and the second microsphere are added and mixed evenly to obtain the precursor solution. S3. The precursor liquid is injected into the mold to form a film, and visible light with a wavelength of 405 nm and an intensity of 5–10 mW / cm² is applied for irradiation for 10–30 seconds to obtain a hydrogel pre-crosslinked film. S4. The surface of the pre-crosslinked hydrogel membrane corresponding to the bone side is immersed in an alkaline buffer solution containing monomers with catechol groups and metal ions to form the inner layer. S5. A solution of polysaccharide is coated onto the surface of the pre-crosslinked hydrogel membrane corresponding to the soft tissue side and dried and cured to form the outer layer.

8. The method according to claim 7, characterized in that, In step S3, the specific process of injecting the precursor liquid into the mold for casting includes: Multiple portions of precursor liquid with strontium-doped nano-hydroxyapatite concentrations increasing sequentially are prepared; each portion of the precursor liquid is injected layer by layer into a mold for casting according to the order of strontium-doped nano-hydroxyapatite concentration from low to high; after each portion of the precursor liquid has completed the casting operation, visible light is applied to the entire film.

9. The method according to claim 7, characterized in that: In step S1, the molar ratio of lactic acid to glycolic acid in the polylactic acid-glycolic acid copolymer wall material used in preparing the first microsphere is 50:50; the molar ratio of lactic acid to glycolic acid in the polylactic acid-glycolic acid copolymer wall material used in preparing the second microsphere is 75:

25. In step S4, the pH of the alkaline buffer solution containing the monomer with catechol groups and metal ions is 8.0–8.5, the concentration of dopamine hydrochloride in the buffer solution is 1–5 mg / mL, the concentration of calcium chloride is 5–20 mM, and the immersion reaction time is 12–24 hours.

10. The application of a bone-guided regeneration membrane capable of intraoperative in-situ light curing as described in any one of claims 1 to 6 in the preparation of a medical device for autologous bone grafting surgery, characterized in that: The autologous bone grafting procedure is used to treat nonunion or delayed fracture healing, including tibial nonunion, femoral nonunion, humeral nonunion, delayed fracture healing, or segmental bone defects.

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