New application of PLLA-TMC-GA polymer, PHA-loaded PLLA-TMC-GA microsphere, microsphere support and preparation method thereof

Abstract

The invention belongs to the field of medical materials, and particularly relates to a novel application of a PLLA-TMC-GA polymer, a PHA-loaded PLLA-TMC-GA microsphere, a microsphere support and a preparation method of the PHA-loaded PLLA-TMC-GA microsphere. The technical problem to be solved by the invention is to provide a bone defect repair material with osteogenic property and mechanical property and a preparation method thereof. The preparation method comprises the following steps: firstly, based on new application of a polylactic acid-trimethylene carbonate-polyglycolic acid ternary polymer (PLLA-TMC-GA) material, preparing a microsphere, a microsphere bracket, a hydrogel composite bracket and a drug-loaded hydrogel composite bracket on the basis of the ternary polymer and polydopamine modified nano-hydroxyapatite; the products have different degrees of conversion effects from artificial bones to biological bones, and a brand-new effective technical scheme is provided for irregular bone defect repair.

Description

Technical Field

[0001] This invention belongs to the field of medical materials, specifically relating to a new application of PLLA-TMC-GA polymer, PHA-loaded PLLA-TMC-GA microspheres and their microsphere scaffolds and preparation methods. Background Technology

[0002] Bone loss due to debridement caused by bone exposure, trauma, or infection affects more than 20 million patients worldwide each year. Alarmingly, this number is projected to increase with the aging population. The global orthopedic implant market was valued at $33.5 billion in 2022 and is expected to grow at an annual rate of 6.3% from 2023 to 2030. Human bone tissue is a complex, multifaceted structure that is constantly undergoing renewal. However, when bone defects exceed a critical size (2 cm) or are accompanied by severe infection or inflammation, the body's self-repair ability is impaired; such defects are often termed large bone defects. Furthermore, in cases of severe bone infection, clinical intervention requires repeated debridement, leading to irregular shapes of the bone defects and making implant placement extremely difficult. Currently, common bone defect repair techniques in clinical practice include vascularized fibular grafting, the Ilizarov technique, and the Masquelet technique. Unfortunately, these techniques also have drawbacks, such as limited autologous bone sources, low repair rates, and a high risk of recurrent infections. Therefore, selecting a human bone replacement scaffold that can effectively fill large bone defects has become a key issue in orthopedic implantation.

[0003] In recent years, artificial bone scaffolds made from biomaterials such as synthetic polymers, natural polymers, and inorganic materials have been extensively studied. In previous work, the inventors of this invention utilized a polylactic acid-trimethylene carbonate copolymer (PLLA-TMC) and prepared a series of composite bone scaffold materials by blending inorganic particles, natural polymers, and synthetic polymers. High molecular weight polymers, such as polytrimethylene carbonate (PTMC) and polylactic acid (PLLA), have been used in the preparation of bone repair scaffolds due to their good biocompatibility and biodegradability. Among these materials, PLLA has high mechanical properties, but the acidic substances produced during degradation can easily lead to a local inflammatory environment. On the other hand, PTMC has surface erosion degradation characteristics, and its degradation products are non-toxic and non-acidic.

[0004] Cancellous bone, as the bone tissue with the fastest metabolic rate and largest volume ratio in the human body, is a primary target for artificial bone biomimicry. In recent years, biomimicry of the macroporous structure of cancellous bone has been achieved through 3D printing and freeze-drying, but its mechanical strength often fails to match that of cancellous bone. Microsphere scaffolds prepared by thermal sintering can effectively improve the mechanical properties of the scaffolds and achieve controllable porosity. Simultaneously, the addition of inorganic nanoparticles is an effective method to improve the performance of composite scaffold materials. Nano-hydroxyapatite (nHA) has a highly similar composition to inorganic nanoparticles in bone, exhibiting good biocompatibility and accelerating bone regeneration. However, the poor dispersibility of nHA in polymer solutions limits the osteogenic and mechanical properties of composite materials. The function of "Mussel Adhesive Protein (MAP)" has inspired the application of dopamine (PDA) self-polymerization functionalization, demonstrating its effectiveness in improving the dispersibility of inorganic particles. Studies have shown that PDA coating can enhance the interfacial compatibility between nHA and the polymer matrix, thereby improving the biocompatibility and mechanical durability of the composite material.

[0005] In this application and research context, the inventors of this invention aim to provide a material for bone defect repair with superior osteogenic and mechanical properties. Summary of the Invention

[0006] The technical problem solved by this invention is to provide a material for repairing bone defects with osteogenic and mechanical properties, and a method for preparing the same.

[0007] The material for bone defect repair in this invention is based on a new application of a polymer, namely the use of polylactic acid-trimethylene carbonate-polyhydroxyacetic acid ternary polymer (PLLA-TMC-GA) in the preparation of materials for bone defect repair.

[0008] The polylactic acid-trimethylene carbonate-polyhydroxyacetic acid ternary polymer (PLLA-TMC-GA) of this invention is prepared by ring-opening polymerization of three monomers—lactic acid (LA), trimethylene carbonate (TMC), and glycolide (GA)—in the presence of a catalyst. This ternary polymer can be used to prepare biomedical implants such as fully degradable vascular stents and porous tissue engineering scaffolds. These implants gradually degrade and are absorbed by the body, eliminating the need for secondary surgery and reducing patient suffering and medical burden. Furthermore, this material can be used in drug delivery systems to achieve precise treatment by controlling the release rate and location of drugs. It is also applied to 3D printing technology, allowing the ternary polymer to be processed into implants and tissue engineering scaffolds of various complex structures and sizes. This processing method not only improves production efficiency and precision.

[0009] The polylactic acid-trimethylene carbonate-polyhydroxyacetic acid terpolymer (PLLA-TMC-GA) used in this invention is prepared by the following method:

[0010] (1) Mix LA, TMC and GA monomers in the following mass ratio: 50-80 parts of LA, 10-30 parts of TMC and 10-30 parts of GA;

[0011] (2) Add stannous octoate to the mixture obtained in step (1) and dry it in a reaction vessel;

[0012] (3) Polymerize at 130-150℃ for 1-2 hours to obtain the product;

[0013] (4) The product obtained in step (3) is dissolved in dichloromethane, and then anhydrous ethanol is added dropwise to precipitate the material uniformly. Then it is dried to obtain polylactic acid-trimethylene carbonate-polyhydroxyacetic acid terpolymer (PLLA-TMC-GA).

[0014] In the above preparation method:

[0015] Step (1) Mix the LA, TMC and GA monomers in the following weight ratio: 72 parts LA, 18 parts TMC and 10 parts GA.

[0016] Step (2) Add 0.01-0.05% stannous octoate according to the weight percentage of the mixture.

[0017] Preferably, in step (2), 0.05% stannous octoate is added as a percentage by weight of the mixture.

[0018] The drying in step (2) is to use a vacuum water pump, a vacuum oil pump and a diffusion pump in sequence for 1 hour.

[0019] Step (3) is preferably carried out at 140°C for 48 hours.

[0020] The drying process described in step (4) is carried out at 30-40°C under vacuum.

[0021] Preferably, the drying in step (4) is performed at 37°C under vacuum.

[0022] The application of the above-mentioned polylactic acid-trimethylene carbonate-polyhydroxyacetic acid ternary polymer (PLLA-TMC-GA) to bone defect repair specifically involves preparing PHA-loaded PLLA-TMC-GA microspheres by combining the polylactic acid-trimethylene carbonate-polyhydroxyacetic acid ternary polymer (PLLA-TMC-GA) with polydopamine-modified nano-hydroxyapatite (also known as PDA-nHA, or PHA for short).

[0023] The polylactic acid-trimethylene carbonate-polyhydroxyacetic acid ternary polymer microspheres (hereinafter referred to as PHA-loaded PLLA-TMC-GA microspheres) of the present invention are prepared by the following method, including the following steps:

[0024] A. The polylactic acid-trimethylene carbonate-polyhydroxyacetic acid terpolymer is dissolved in a solvent to form an oil phase, and then polydopamine-modified nano-hydroxyapatite is added to form a mixed phase;

[0025] B. Prepare an aqueous solution of polyvinyl alcohol (PVA) as the aqueous phase;

[0026] C. Add the mixed phase obtained in step A dropwise to the aqueous phase obtained in step B, and stir to obtain an emulsion;

[0027] D. Add water to the emulsion obtained in step C and stir to disperse until solidified microspheres are obtained;

[0028] E. Filter, wash and dry the solidified microspheres obtained in step D with water to obtain polylactic acid-trimethylene carbonate-polyhydroxyacetic acid ternary polymer microspheres loaded with polydopamine-modified nano-hydroxyapatite (abbreviated as PHA-loaded PLLA-TMC-GA microspheres).

[0029] In the above technical solution:

[0030] Step A: Mix according to the following ratio: dissolve 1g of polylactic acid-trimethylene carbonate-polyhydroxyacetic acid terpolymer in 5-20ml of solvent.

[0031] Preferably, step A is prepared according to the following ratio: 1g of polylactic acid-trimethylene carbonate-polyhydroxyacetic acid terpolymer is dissolved in 15ml of solvent.

[0032] The solvents used in step A are dichloromethane, dioxane (i.e., 1,4-dioxane), dimethyl sulfoxide, and acetone.

[0033] Preferably, the solvent in step A is dichloromethane.

[0034] In step A, the amount of polydopamine-modified nano-hydroxyapatite added is 5-20 wt% of the oil phase described in step A.

[0035] Preferably, in step A, the amount of polydopamine-modified nano-hydroxyapatite added is 20 wt% of the oil phase described in step A.

[0036] In step A, the polylactic acid-trimethylene carbonate-polyhydroxyacetic acid terpolymer is dissolved in dichloromethane and stirred magnetically until completely dissolved.

[0037] In step A, the stirring speed of the magnetic stirrer a is 300-350 rpm.

[0038] Preferably, in step A, the stirring speed of the magnetic stirrer a is 300 rpm.

[0039] In step A, polydopamine-modified nano-hydroxyapatite is added to form a mixed phase, which is then magnetically stirred until homogeneous.

[0040] In step A, the stirring speed of the even-numbered magnetic stirrer is 400-500 rpm.

[0041] Preferably, in step A, the stirring speed of the magnetic stirrer b is 500 rpm.

[0042] The concentration of the polyvinyl alcohol aqueous solution in step B is 0.1% to 1% w / v, that is, 10 ml to 1 ml is used to dissolve 1 g of polyvinyl alcohol.

[0043] Preferably, the concentration of the polyvinyl alcohol aqueous solution in step B is 0.1% w / v, that is, 1g of polyvinyl alcohol is dissolved in 10ml.

[0044] In step C, the mixed phase obtained in step A is added dropwise to the aqueous phase obtained in step B by adding 100-150 ml of aqueous phase for every 1 g of polylactic acid-trimethylene carbonate-polyhydroxyacetic acid terpolymer in the mixed phase.

[0045] Preferably, the mixed phase obtained in step A is added dropwise to the aqueous phase obtained in step B by adding 120 ml of aqueous phase for every 1 g of polylactic acid-trimethylene carbonate-polyhydroxyacetic acid terpolymer in the mixed phase.

[0046] In step C, the stirring is performed using magnetic stirring.

[0047] In step C, the stirring speed of the magnetic stirrer is 300-400 rpm.

[0048] Preferably, in step C, the stirring speed of the magnetic stirrer is 350 rpm.

[0049] In step D, the water is deionized water or phosphate buffered saline (PBS).

[0050] Preferably, in step D, the water is deionized water.

[0051] In step D, add 2-3 times the amount of water as the emulsion from step C.

[0052] Preferably, in step D, 2.6 times the amount of water as the emulsion in step C is added.

[0053] In step D, the stirring is performed using magnetic stirring.

[0054] In step D, the stirring conditions are as follows: first, reduce the stirring speed by 50-100 rpm compared to step C, and then continue stirring for 10-14 hours. The purpose of reducing the stirring speed is to promote the formation of dispersed, stable droplets and accelerate the volatilization of dichloromethane. During this prolonged stirring process, microspheres gradually form and solidify.

[0055] Preferably, in step D, the stirring conditions are as follows: first, the stirring speed is reduced by 50 rpm based on step C, and then the stirring time is maintained for 12 hours.

[0056] In step E, the filtration is performed using a material with a particle size of 300μm-500μm to obtain solidified microspheres.

[0057] In step E, the washing involves washing with water 2 to 4 times; preferably, the washing involves washing with water 3 times.

[0058] In step E, the water is deionized water or phosphate buffered saline (PBS).

[0059] Preferably, in step E, the water is deionized water.

[0060] In step E, the drying is carried out in a vacuum drying environment at a temperature of 35-40℃.

[0061] Preferably, in step E, the vacuum drying environment is a vacuum oven.

[0062] Preferably, the drying temperature in step E is 37°C.

[0063] Further applications of the microspheres of this invention include assembling them into scaffolds through sintering. The microsphere scaffold of this invention is made of polylactic acid-trimethylene carbonate-polyhydroxyacetic acid ternary polymer microspheres loaded with polydopamine-modified nano-hydroxyapatite (hereinafter referred to as PHA-loaded PLLA-TMC-GA microspheres), which are pressurized and compressed in a mold and then subjected to gradient sintering to obtain the microsphere scaffold of this invention.

[0064] In the above-mentioned method for preparing microsphere scaffolds,

[0065] The mold can be an 8mm x 4mm cylindrical mold, a 4mm x 4mm square mold, or a 4mm x 4mm x 4mm triangular mold; among them, the cylindrical mold is the most effective.

[0066] The pressurization pressure is 50-150 MPa.

[0067] Preferably, the pressurization pressure is 100 MPa.

[0068] The gradient sintering refers to sintering at temperatures of 110℃, 120℃, 130℃, and 140℃ sequentially.

[0069] The gradient sintering involves sintering at each temperature for 3-5 hours.

[0070] Preferably, the gradient sintering is performed by sintering at each temperature for 4 hours.

[0071] The microsphere scaffold of this invention possesses mechanical strength comparable to cancellous bone and exhibits controllable shape memory function. In vivo experiments have also confirmed that the microsphere scaffold of this invention can transform artificial bone into biological bone.

[0072] One application of the microsphere scaffold of this invention is the addition of sodium alginate hydrogel to the microsphere scaffold, allowing the active ingredients to be loaded through the cross-linked "egg-box" structure of sodium alginate. Specifically, the microsphere scaffold of cross-linked sodium alginate hydrogel of this invention is prepared using the following method:

[0073] (a) Sodium alginate was used to make sodium alginate hydrogel;

[0074] (b) The microsphere scaffold is soaked in the hydrogel obtained in step (a), crosslinked, and dried to form a hydrogel-microsphere scaffold.

[0075] In step (a), the sodium alginate hydrogel contains 2-4 wt% sodium alginate.

[0076] Preferably, the sodium alginate hydrogel in step (a) contains 3 wt% sodium alginate.

[0077] The crosslinking conditions described in step (b) are crosslinking in a calcium ion solution with a content of 0.1-1% for 1-6 hours.

[0078] Preferably, the crosslinking conditions in step (b) are crosslinking in a calcium ion solution with a content of 0.6% for 4 hours.

[0079] In step (b), the drying process is freeze drying.

[0080] The microsphere scaffold of the present invention can also be loaded with hydrogel containing drug-active ingredients to form a drug-loaded hydrogel-microsphere composite scaffold, which is prepared by the following method:

[0081] (a1) A drug-loaded hydrogel made from active ingredient sodium alginate;

[0082] (b1) The microsphere scaffold is soaked in the drug-loaded hydrogel obtained in step (a1), crosslinked, and dried to form a drug-loaded hydrogel-microsphere composite scaffold.

[0083] The sodium alginate hydrogel described in step (a1) contains 2-4 wt% sodium alginate.

[0084] Preferably, the sodium alginate hydrogel in step (a1) contains 3 wt% sodium alginate.

[0085] The active ingredient in step (a1) is selected from icariin.

[0086] In step (a1), the concentration of icariin liquid is 5-6 mg / mL.

[0087] Preferably, the concentration of icariin liquid in step (a1) is 5.4 mg / mL.

[0088] The hydrogel preparation method described in step (a1) is as follows: The hydrogel is prepared according to the following ratio:

[0089] Dissolve 0.15g to 0.18g of icariin in 30mL of deionized water, stir well, and then add 0.6g to 1.2g of sodium alginate to completely dissolve it.

[0090] Preferably, 0.162g of icariin is dissolved in 30mL of deionized water, stirred evenly, and then 0.9g of sodium alginate is added and completely dissolved.

[0091] The crosslinking conditions described in step (b1) are crosslinking in a calcium ion solution with a content of 0.1-1% for 1-12 hours.

[0092] Preferably, the crosslinking conditions in step (b1) are crosslinking in a calcium ion solution with a content of 0.6% for 4 hours.

[0093] In step (b1), the drying process is freeze drying.

[0094] Furthermore, the freeze-drying in step (b1) involves pre-freezing at -20 to 80°C for 10 to 100 minutes and freeze-drying for 24 to 48 hours.

[0095] Preferably, the freeze-drying in step (b1) involves pre-freezing at -80°C for 20 minutes and freeze-drying for 24 hours.

[0096] The icariin-loaded hydrogel-microsphere composite scaffold obtained by the aforementioned method is abbreviated as PLLA-TMC-GA / PHA / Ica, or PTG / PHA / Ica. In vivo experiments have confirmed the transformation of this scaffold from artificial bone to biological bone.

[0097] In summary, this invention is the first to utilize a polylactic acid-trimethylene carbonate-polyhydroxyacetic acid ternary polymer (PLLA-TMC-GA) in the field of bone defect repair. This ternary polymer and polydopamine-modified nano-hydroxyapatite (PHA) are used to prepare microspheres via a single-emulsion method, and the microspheres are further sintered and assembled into a scaffold. The scaffold is then cross-linked with sodium alginate hydrogel to form a composite scaffold, creating a hierarchical porous biocomposite structure. This scaffold possesses mechanical strength comparable to cancellous bone and exhibits controllable shape memory function. This composite scaffold can also load active ingredients with bone metabolism benefits. The cross-linked "egg-box" structure of sodium alginate facilitates the sustained release of these active ingredients, thereby enabling long-term management of bone repair. In vivo experiments have confirmed that the microspheres, microsphere scaffolds, hydrogel composite scaffolds, and drug-loaded hydrogel composite scaffolds of this invention all exhibit varying degrees of transformation from artificial bone to biological bone, providing a novel and effective technical solution for the repair of irregular bone defects. Attached Figure Description

[0098] Figure 1 Synthetic route diagram of terpolymer PLLA-TMC-GA.

[0099] Figure 2 NMR image of PLLA-TMC-GA.

[0100] Figure 3 Characterization diagram of PDA-nHA.

[0101] in Figure 3 A-3F respectively represent:

[0102] Figure 3 A: Schematic diagram of PDA self-polymerization to synthesize PDA-nHA;

[0103] Figure 3 B: Macroscopic and SEM images of n-HA and PDA-nHA;

[0104] Figure 3 C: Dispersion of n-HA and PDA-nHA in PLLA-TMC-GA polymer solution;

[0105] Figure 3 D: FTIR analysis of PDA-nHA;

[0106] Figure 3 E: XRD pattern of PDA-nHA;

[0107] Figure 3 F: TEM images of nHA and PHA, with white arrows indicating PDA coatings.

[0108] Figure 4 Particle size distribution of microspheres.

[0109] in: Figure 4 A represents the microsphere size at different PVA concentrations; Figure 4 B represents the microsphere size at different polymer concentrations.

[0110] Figure 5 Characterization diagram of microspheres.

[0111] in Figure 5 A- Figure 5 E represents:

[0112] Figure 5 A represents a schematic diagram of microsphere preparation using the single emulsion method;

[0113] Figure 5 B indicates: SEM image of microspheres;

[0114] Figure 5 C indicates: microsphere XRD;

[0115] Figure 5 D indicates: FTIR analysis of microspheres;

[0116] Figure 5 E represents the contact angle of the microsphere material.

[0117] Compared with the control group, the P values ​​were <0.05 (*), <0.01 (**), <0.001 (***), and <0.0001 (****). Figure 6 Characterization diagram of sintering temperature.

[0118] in: Figure 6 A- Figure 6 E represents:

[0119] Figure 6 A: Schematic diagram of sintered microsphere scaffold;

[0120] Figure 6 B: Permeability of the sintered support;

[0121] Figure 6 C: SEM image of the sintered support;

[0122] Figure 6 D: Porosity of the microsphere scaffold;

[0123] Figure 6 E: Compression performance of the microsphere scaffold.

[0124] Compared with the control group, the P values ​​were <0.05 (*), <0.01 (**), <0.001 (***), and <0.0001 (****). Figure 7 Characterization of microsphere scaffolds.

[0125] in Figure 7 A- Figure 7 F represents:

[0126] Figure 7 A: Macroscopic images of microsphere scaffolds, including: (a) PTG scaffold; (b) PTG / 20PHA scaffold;

[0127] Figure 7 B: SEM image of the microsphere scaffold;

[0128] Figure 7 C: Microsphere size distribution;

[0129] Figure 7 D: Compression performance of the microsphere scaffold;

[0130] Figure 7 E: Porosity of the microsphere scaffold;

[0131] Figure 7 F: Shape memory function, in which Figure 7 F also includes four smaller figures, a and b, which respectively represent: (a) scaffold filling irregular bone defects; (b) PTG shape memory function; (c) PTG / PHA scaffold shape memory function; and (d) scaffolds at different temperatures having shape memory function.

[0132] Compared with the control group, the P values ​​were <0.05 (*), <0.01 (**), <0.001 (***), and <0.0001 (****). Figure 8 Figure showing the in vitro degradation characteristics of the microsphere scaffold.

[0133] in, Figure 8 A- Figure 8 E represents:

[0134] Figure 8 A: Figure 8 A contains Figure 8 A1- Figure 8 A3 represents: (A1) SEM images of the microsphere scaffold after 8 and 12 weeks of in vitro degradation; (A2) erosion and degradation pores on the surface of the microspheres; (A3) dissolution of the microsphere scaffold after 16 weeks of degradation.

[0135] Figure 8 B: Comparison of the compressive modulus of the microsphere scaffold before and after degradation;

[0136] Figure 8 C: Mass loss during stent degradation;

[0137] Figure 8 D: Changes in porosity during scaffold degradation;

[0138] Figure 8E: Changes in hydrophilicity during scaffold degradation.

[0139] Compared with the control group, the P values ​​were <0.05 (*), <0.01 (**), <0.001 (***), and <0.0001 (****). Figure 9 Biocompatibility diagram of microsphere scaffolds.

[0140] in Figure 9 A- Figure 9 Representative I:

[0141] Figure 9 A: Fluorescence detection of MC3T3-E1 cells cultured on PTG / PHA / Ica microsphere scaffolds at 1, 3, and 7 days;

[0142] Figure 9 B: Percentage of PTG / PHA / Ica stent release;

[0143] Figure 9 C: Fluorescence intensity of MC3T3-E1 cells cultured on PTG / PHA / Ica microsphere scaffolds at 1, 3, and 7 days;

[0144] Figure 9 D: CCK-8 results of MC3T3-E1 cells on different scaffolds; (E) Cell viability analysis;

[0145] Figure 9 F: ALP activity in MC3T3-E1 cells; (G) Cytoskeleton staining of MC3T3-E1 cells on different scaffolds;

[0146] Figure 9 H: Staining of MC3T3-E1 cells cultured on PTG / PHA / Ica microsphere scaffolds for 3 and 7 days under an ALP microscope;

[0147] Figure 9 I: ALP staining to evaluate the osteogenic effect of different scaffolds.

[0148] Compared with the control group, the P values ​​were <0.05 (*), <0.01 (**), <0.001 (***), and <0.0001 (****). Figure 10 Image showing the in vivo healing effect of PTG / PHA / Ica stents.

[0149] in Figure 10 A- Figure 10 F represents:

[0150] Figure 10 A: Flowchart for constructing a skull defect model;

[0151] Figure 10B: Macroscopic images of surrounding tissues at different postoperative time points;

[0152] Figure 10 C: Micro-CT image of skull defect in rat;

[0153] Figure 10 D: Micro-CT image of a rat skull defect without a scaffold (red represents newly formed bone tissue);

[0154] Figure 10 E: Quantitative analysis of Micro-CT data: BV / TV (%) at the stent filling site;

[0155] Figure 10 F: Quantitative analysis of Micro-CT data: BMD at the stent filling site (mg HA / ccm).

[0156] Compared with the control group, the P values ​​were <0.05 (*), <0.01 (**), <0.001 (***), and <0.0001 (****). Figure 11 Histological analysis of the skull was performed at 6 and 12 weeks postoperatively.

[0157] in, Figure 11 A represents the hematologic staining results at 6 and 12 weeks after stent implantation, where B represents host bone, NB represents newly formed bone, and S represents the stent.

[0158] Figure 11 B indicates the result of Masson trichrome staining. Detailed Implementation

[0159] The present invention will be explained and described below with reference to embodiments. Those skilled in the art will understand that the following embodiments are for illustrative purposes only and should not be considered as limiting the scope of the invention. Where specific techniques or conditions are not specified in the embodiments, they are performed according to the techniques or conditions described in the literature in the field or according to the product instructions. Reagents or instruments whose manufacturers are not specified are all conventional products that can be obtained commercially.

[0160] The following are the definitions of the English abbreviations and explanations of the main terms used in this invention:

[0161] Polylactic acid-trimethylene carbonate-polyhydroxyacetic acid ternary polymer is abbreviated as PLLA-TMC-GA; lactic acid is abbreviated as LA, and the LA monomer used in this invention is L-lactic acid monomer; trimethylene carbonate is abbreviated as TMC; glycolide is abbreviated as GA; dopamine is abbreviated as PDA; nano-hydroxyapatite is abbreviated as nHA or n-HA; polydopamine-modified nano-hydroxyapatite can also be called PDA-nHA, abbreviated as PHA; polylactic acid-trimethylene carbonate-polyhydroxyacetic acid ternary polymer microspheres loaded with polydopamine-modified nano-hydroxyapatite are abbreviated as PHA-loaded PLLA-TMC-GA microspheres; polyvinyl alcohol Alcohol, abbreviated as PVA; polylactic acid-trimethylene carbonate-polyhydroxyacetic acid ternary polymer microspheres loaded with polydopamine-modified nano-hydroxyapatite, abbreviated as PLLA-TMC-GA microspheres or PLLA-TMC-GA / PHA microspheres; icariin, abbreviated as Ica; hydrogel-microsphere composite scaffold loaded with icariin, abbreviated as PLLA-TMC-GA / PHA / Ica or PTG / PHA / Ica; potassium bromide, abbreviated as KBr; Tris solution, Tris is tris(hydroxymethyl)aminomethane, generally abbreviated as Tris, Tris solution is Tris (60.57g, 1N) Approximately 420 ml of HCl; add double-distilled water to a final volume of 1000 ml. Preparation method: First, dissolve Tris in a small amount of double-distilled water (300-500 ml). After adding HCl, adjust the pH to 7.6 with HCl (1N) or NaOH (1N). Finally, add double-distilled water to a final volume of 1000 ml. This solution is a stock solution and should be stored at 4°C.

[0162] The following embodiments and experiments demonstrate the beneficial effects of the technical solution of the present invention.

[0163] 1. Synthesis and Characterization of PLLA-TMC-GA

[0164] The synthesis of PLLA-TMC-GA was achieved through ring-opening polymerization of cyclic monomers: LA, TMC, and GA monomers were added to a reaction flask in a mass ratio of 72:18:10 along with stannous octoate (0.05 wt%). The mixture was vacuum dried for 1 hour using a vacuum water pump, a vacuum oil pump, and a diffusion pump, respectively. The product was then polymerized at 140°C for 48 hours to obtain the final product. The product was removed from the flask and completely dissolved in dichloromethane in a beaker. Anhydrous ethanol was then added dropwise using a separatory funnel to allow the material to settle uniformly. Finally, the ternary polymer was collected and dried to constant weight in a vacuum oven at 37°C. It was then weighed and recovered using a balance. The synthetic route is shown below. Figure 1As shown. The PLLA-TMC-GA ternary polymer was analyzed by 1H NMR spectroscopy. The specific steps were as follows: 2-10 mg of sample was weighed and carefully placed into a sample tube. 0.5 ml of CDCL3 (deuterated chloroform) was injected into the sample tube using a syringe to ensure complete dissolution. Subsequently, the sample was tested using the instrument, and the positions of individual hydrogen atoms on the terminal residues were determined to ascertain the content. The data were analyzed and plotted using Origin software. The NMR spectrum is shown below. Figure 2 Through preparation methods and testing, the inventors confirmed that the ternary polymers prepared using the following methods are all suitable for the standards of preparing bone defect repair materials.

[0165] (1) Mix LA, TMC and GA monomers in the following mass ratio: PLLA 50-80 parts, TMC 10-30 parts, GA 10-30 parts;

[0166] (2) Add stannous octoate to the mixture obtained in step (1) and dry it in a reaction vessel;

[0167] (3) Polymerize at 130-150℃ for 1-2 hours to obtain the product;

[0168] (4) The product obtained in step (3) is dissolved in dichloromethane, and then anhydrous ethanol is added dropwise to precipitate the material uniformly. Then it is dried to obtain polylactic acid-trimethylene carbonate-polyhydroxyacetic acid terpolymer (PLLA-TMC-GA).

[0169] In the above preparation method:

[0170] Step (1) Mix the LA, TMC and GA monomers in the following weight ratio: 72 parts PLLA, 18 parts TMC and 10 parts GA.

[0171] Step (2) Add 0.01-0.05% stannous octoate according to the weight percentage of the mixture.

[0172] Preferably, in step (2), 0.05% stannous octoate is added as a percentage by weight of the mixture.

[0173] The drying in step (2) involves vacuum drying for 1 hour sequentially using a vacuum water pump, a vacuum oil pump, and a diffusion pump.

[0174] Step (3) is preferably carried out at 140°C for 48 hours.

[0175] The drying process described in step (4) is carried out at 30-40°C under vacuum.

[0176] Preferably, the drying in step (4) is performed at 37°C under vacuum.

[0177] 2. Synthesis and characterization of polydopamine-modified n-HA (PHA)

[0178] To synthesize PHA, 1 g of n-HA powder was dispersed in 250 mL of Tris solution at pH 8.5 and stirred on a magnetic stirrer (500 rpm). Then, 0.5 g of dopamine hydrochloride (2 mg / mL) was weighed and added to the n-HA suspension. The reaction was carried out at room temperature in the dark with continuous magnetic stirring for 24 h. After the reaction was completed, the product was separated by filtration, washed three times with deionized water, and freeze-dried to obtain a brownish-black powder.

[0179] Characterization methods: Fourier transform infrared spectroscopy was performed on PHA using an infrared spectrometer. Samples were prepared by potassium bromide pelleting, and samples were collected from 500 to 4000 cm⁻¹. -1 The infrared characteristic peaks of both substances were detected and observed within the specified range. X-ray diffraction (XRD) was used to characterize n-HA and PHA, analyzing their crystal forms. Specifically, appropriate amounts of n-HA and PHA powder were evenly spread in the sample cell of a glass slide. The accelerating voltage was set to 40 kV, the tube current to 30 mA, CuKα radiation was used, and the scanning angle range was 10-90°. Origin software was used to analyze and plot the data. The dispersibility of n-HA and PHA in the ternary polymer solution was detected. Specifically, 1 g of n-HA and PHA powder were weighed into glass sample vials, and each vial was then filled with a completely dissolved ternary polymer solution (w / v = 1:15). The vials were sealed with adhesive and the caps tightened to prevent dichloromethane evaporation. The vials were stirred vigorously for 1 min. The deposition of n-HA and PHA was observed at 5 min, 10 min, and 30 min. The microscopic surface morphology of n-HA and PHA was analyzed using scanning electron microscopy. The procedure was as follows: n-HA and PHA powders were adhered to the stage with conductive adhesive and sputtered with gold. The sputtered samples were then placed in the machine, and the voltage was set to 10KV for scanning. Images were taken at different magnifications as needed.

[0180] Experimental results: In this invention, the inventors synthesized PDA-nHA using the self-polymerization principle of dopamine under weakly alkaline conditions. Figure 3 A). PDA-modified nHA appeared as a brownish-black powder, indicating that PDA was successfully deposited on the n-HA surface. Scanning electron microscopy revealed that PHA agglomeration was eliminated, thereby improving dispersibility. Figure 3 B). Furthermore, the inventors found that the particle size of the modified n-HA particles was significantly reduced. Interestingly, although initially uniformly suspended, the n-HA group exhibited a noticeable stratification effect from day one, while the PHA group showed no significant deposition. Figure 3C). These phenomena indicate that there are considerable van der Waals forces and hydrogen bonds between nHA particles, leading to aggregation. Conversely, PDA functionalization tends to reduce the surface energy of nHA, thereby improving the interfacial compatibility between nHA and the polymer. This enhanced interfacial compatibility allows PHA to be uniformly dispersed in solution.

[0181] FT-IR characterization analysis of PDA, HA and PHA structures Figure 3 D) indicates that in the HA sample, 3447.4 cm -1 The peak at 564.3 cm⁻¹ is a characteristic peak of hydroxyl groups. -1 and 603.4cm -1 The characteristic peak at 1039.7 cm⁻¹ is due to the PO bending of the phosphate group. -1 and 1096.1cm -1 The characteristic peak at this location is due to the stretching band of OPO phosphate ions. Furthermore, the effect of PDA modification on the properties of nHA single crystals was investigated by XRD. Figure 3 E). XRD patterns showed similar Bragg peaks between nHA and PHA, with corresponding 2θ values ​​(25.8°, 31.7°, 46.7°, 49.5°, and 53.1°) matching the standard HA reflectance. These values ​​are consistent with the crystal planes (002), (211), (300), (222), and (213), indicating that PDA functionalization does not affect the structure or crystallinity of nHA. Finally, a PDA coating was observed on the surface of nHA in TEM images. Figure 3 F).

[0182] 3. Preparation and characterization of PLLA-TMC-GA microspheres

[0183] (1) Investigation of the effect of different PVA concentrations on the particle size of PLLA-TMC-GA microspheres: PLLA-TMC-GA microspheres were prepared using a single emulsion method. The specific steps were as follows: 1 g of PLLA-TMC-GA ternary polymer material was dissolved in 20 mL of dichloromethane solution (w / v = 1:20, 300 rpm) at room temperature as the oil phase O. 120 mL of PVA solutions of different concentrations (w / v = 0.1%, 0.25%, 0.5%, 1%) were prepared as the aqueous phase W. The polymer solution was then slowly added dropwise to the PVA solution, and the mixture was stirred at 350 rpm to form an emulsion O / W. 350 mL of deionized water was poured into the emulsion, and the stirring speed was reduced to 300 rpm during this period to promote the formation of dispersed and stable droplets. The stirring process was then continued for 12 h to promote the formation and solidification of microspheres. Afterwards, the formed microspheres were filtered and rinsed three times with deionized water, and then dried and collected in a vacuum oven at 37 °C. The microstructure of the prepared microspheres with different components was captured by scanning electron microscopy (SEM). The particle size distribution of the microspheres was then analyzed using Nano Measurer software. The selected microspheres were randomly distributed, with a total of 100 microspheres measured for each component. The specific steps for using Nano Measurer software were as follows: first, import the JPG format SEM image; select a reference scale bar of the same length as the SEM image and set the same unit (μm); then, measure the diameter of each microsphere individually to obtain the final particle size data for all microspheres. The results are shown below. Figure 4 A.

[0184] (2) Investigation of the effect of different polymer concentrations on the particle size of PLLA-TMC-GA microspheres: PLLA-TMC-GA microspheres were prepared using a single emulsion method. The specific steps were as follows: 1 g of PLLA-TMC-GA ternary polymer material was dissolved in different volumes of dichloromethane solution (w / v = 1:5, 1:10, 1:15, 1:20, 300 rpm) at room temperature as the oil phase (O). 120 mL of 0.25% (w / v) PVA solution was prepared as the aqueous phase (W). The polymer solution was then slowly added dropwise to the PVA solution, and the mixture was stirred at 350 rpm to form an emulsion (O / W). 350 mL of deionized water was poured into the emulsion, and the stirring speed was reduced to 300 rpm during this period to promote the formation of dispersed and stable droplets. The stirring process was then continued for 12 h to promote the formation and solidification of microspheres. Afterward, the formed microspheres were filtered and rinsed three times with deionized water, then dried and collected in a vacuum oven at 37℃. The microsphere particle size data were measured according to method 3(1), and the results are shown in [see figure]. Figure 4 B.

[0185] 4. Preparation and characterization of PLLA-TMC-GA / PHA composite microspheres

[0186] Preparation of composite microspheres: PLLA-TMC-GA / PHA composite microspheres were prepared using a single emulsion method. The specific steps were as follows: 1 g of PLLA-TMC-GA ternary polymer was dissolved in 15 mL of dichloromethane solution at room temperature. After the polymer was completely dissolved, different amounts of PHA powder (5 wt%, 10 wt%, 15 wt%, 20 wt%) were added to form a mixed S / O phase, and stirred at 500 rpm until homogeneous. 120 mL of a 0.1% (w / v) PVA solution was prepared as the aqueous phase (W). The S / O phase mixture was slowly added dropwise to the PVA solution, and stirred at 350 rpm to form an emulsion S / O / W. 350 mL of deionized water was poured into the emulsion, and the stirring speed was reduced to 300 rpm during this period to promote the formation of dispersed, stable droplets and accelerate the volatilization of dichloromethane. The stirring process continued for 12 h to promote the formation and solidification of microspheres. Afterwards, the formed microspheres were filtered and rinsed three times with deionized water, then dried and collected in a vacuum oven at 37℃. The same method was used to prepare a control group with 10 wt% unmodified HA. For simplicity, the initial PLLA-TMC-GA microspheres were uniformly referred to as PTG, and the microspheres prepared by blending PLLA-TMC-GA with 5 wt% PHA, 10 wt% PHA, 15 wt% PHA, 20 wt% PHA and 10 wt% HA were respectively referred to as PTG / 5PHA, PTG / 10PHA, PTG / 15PHA, PTG / 20PHA and PTG / 10HA.

[0187] Characterization of the composite microspheres: The microscopic surface morphology of the microspheres was analyzed using scanning electron microscopy (SEM). The procedure was as follows: the microspheres were adhered to the stage with conductive adhesive and then sputter-coated with gold. The sputter-coated samples were placed in the microscope, and scanning was performed at a voltage of 10 kV. Images were taken at different magnifications as needed. Fourier transform infrared (FTIR) spectroscopy was used to analyze the PLLA-TMC-GA / PHA microspheres. The specific procedure was as follows: 1-3 mg of PLLA-TMC-GA / PHA microspheres were weighed and ground together with potassium bromide, pressed into tablets, and samples were collected from 500 to 4000 cm⁻¹. -1 Data within the specified range were used to detect and observe infrared characteristic peaks and plot infrared images.

[0188] XRD characterization of PLLA-TMC-GA / PHA microspheres was performed using an X-ray diffractometer to analyze the crystal form of each substance. The specific steps were as follows: appropriate amounts of PLLA-TMC-GA / PHA microspheres with different components were evenly spread in the sample cell of a glass slide, the accelerating voltage was set to 40KV, the tube current to 30mA, CuKα rays were used, and the scanning angle range was 10-90°.

[0189] Experimental results: Six microsphere compositions (PTG, PTG / 5PHA, PTG / 10PHA, PTG / 15PHA, PTG / 20PHA, PTG / 10HA) were prepared using a single emulsion method. Figure 5 A). To ensure that the porosity of the scaffold is greater than 40% and the pore size is greater than 250 μm, based on previous reports, a microsphere size of approximately 350-500 μm was selected.

[0190] Use SEM to observe the microspheres, such as Figure 5 As shown in Figure B, the surface of the PTG microspheres is smooth. However, with the increase of PHA content (5–20 wt%), the surface of the microspheres gradually becomes rough and small pores appear. Notably, a large number of micropores were also observed on the surface of the PTG / 10HA microspheres. These micropores are considered to be a result of interfacial incompatibility between nHA and the polymer, as nHA is not completely encapsulated in the microspheres. Furthermore, the encapsulation efficiency of the PHA microspheres prepared under various formulation conditions was determined (as shown in Table 1). As the PHA content increased from 5% to 20%, the actual encapsulation efficiency of the PHA content increased from 1.1% to 5.4%. Therefore, it is not necessary to further increase the gradient of PTG encapsulation for PHA.

[0191] Encapsulation efficiency of 1PLLA-TMC-GA / PHA composite microspheres

[0192]

[0193] FTIR spectroscopy further confirmed the presence of nHA and PHA in the composite microspheres, such as Figure 5 As shown in D. In the spectrum, 1736.3 cm⁻¹ -1 The peak at [location] corresponds to the C=O stretching vibration peak on the PLLA molecular chain. Additionally, a peak at 1403.3 cm⁻¹... -1 The peak represents the characteristic -CH2 peak in the GA molecular chain. (At 2945.8 cm⁻¹) -1 The nearby absorption peaks are caused by the CH stretching vibrations induced by the -CH and CH3 groups of LA and the -CH2 group of TMC. (1184.8 cm⁻¹) -1 The peak at 1038.2 cm⁻¹ is a characteristic peak resulting from the splitting of the LA and TMC chain segments after the COC stretching vibration. -1 603.1cm -1 and 564.5cm -1 The peaks are typical characteristic peaks of HA in (PO4)3-. These characteristic peaks are also retained in the PTG / PHA microspheres, and their positions do not change with increasing PHA content. Further XRD analysis of the phase composition of the composite microspheres (…) Figure 5C). The diffraction peaks at 25.8°, 31.7°, 46.7°, 49.5°, and 53.1° are characteristic peaks of HA, and no other crystalline peaks were observed. These diffraction peaks become more pronounced with increasing PHA content, indicating that the incorporation of PHA does not change the crystallinity of the PTG polymer. Furthermore, PTG / HA and PTG / PHA microspheres exhibit similar Bragg peaks, indicating that the addition of PHA to the microspheres does not affect the crystal structure of nHA.

[0194] PLLA-TMC-GA / PHA Hydrophilicity Angle Test: To test the hydrophilicity of the composite polymer, 1g of PLLA-TMC-GA was added to 15mL of dichloromethane solution and magnetically stirred until the ternary polymer material was completely dissolved (note that dichloromethane is volatile; when the polymer is completely dissolved, monitor the amount of dichloromethane that has evaporated and add more solution as needed). Then, PHA powder was added to the polymer solution in different proportions (5wt%, 10wt%, 15wt%, 20wt%) and magnetically stirred until homogeneous at room temperature. The homogeneously mixed PLLA-TMC-GA / PHA solution was then poured onto a glass plate. The glass plate was placed in a well-ventilated outdoor area to allow the dichloromethane to fully evaporate, resulting in a PLLA-TMC-GA / PHA film. When demolding, alcohol could be dropped onto the glass plate, and the film was slowly demolded to ensure its integrity. The film was then placed in a 37℃ vacuum oven and dried until constant weight was achieved. PLLA-TMC-GA films and unmodified n-HA blended PLLA-TMC-GA films were prepared using the same method described above.

[0195] The hydrophilicity of material surfaces plays a crucial role in controlling cell adhesion and proliferation. To assess the hydrophilicity of PTG / PHA surfaces, a static droplet method was used. Figure 5 E) Measurement of water contact angle (WCA). Considering that PLLA and PTMC are inherently hydrophobic polymers, PTG exhibits a relatively large WCA (86.4 ± 2.0°). However, the hydrophilicity of the material can be improved by increasing the PHA content in the composite (ranging from 5-20 wt%). When 20 wt% PHA was added to the polymer film, the WCA decreased to 62 ± 2°. This shift towards hydrophilicity can be attributed to the strong hydrophilicity of HA and the hydrophilic groups on the PDA surface, such as carboxyl, amino, imino, and phenolic groups. These groups collectively contribute to improving the hydrophobic properties of the polymer.

[0196] 5. Preparation of sintered microsphere scaffolds

[0197] Microsphere scaffolds were prepared using thermal sintering technology. The microspheres were compressed at 100 MPa in an 8 mm x 4 mm cylindrical mold. The scaffolds were sintered for 4 hours at four different temperature gradients (110℃, 120℃, 130℃, and 140℃). After removing the mold, the original microsphere scaffolds were obtained. Icariin was immersed in sodium alginate hydrogel and injected into the internal pores of the microsphere scaffolds to form a hydrogel-microsphere composite scaffold. The microsphere scaffolds were immersed in a specially prepared 3 wt% sodium alginate solution, then crosslinked in a calcium ion solution (0.6%) for 4 hours, followed by freeze-drying. The crosslinked scaffolds were then treated in an icariin solution (5.4 mg / mL) at 100 rpm for 12 hours to obtain the PLLA-TMC-GA / PHA / Ica microsphere scaffold (referred to as PTG / PHA / Ica).

[0198] (1) Characterization of sintering behavior of PTG / PHA microsphere scaffold

[0199] Considering that bone strength varies with age and individual patient conditions, developing scaffolds with controllable strength is imperative. Typically, two key factors for biomimetic scaffolds are mechanical properties and permeability. PTG scaffolds are prepared by sintering at 110℃, 120℃, 130℃, and 140℃ to facilitate evaluation of the correlation between sintering temperature and permeability and mechanical properties. Figure 6 A). For example Figure 6 As shown in Figure B, within 3 seconds of ink droplet application, the ink completely penetrated into the microsphere scaffold sintered at 110℃, indicating sufficient porosity between the microspheres. However, as the sintering temperature increased, the ink's penetration into the scaffold significantly decreased, until the microsphere scaffold sintered at 130℃ ultimately hindered ink entry. Scanning electron microscopy and porosity measurements further evaluated the pore size between the microspheres. Figure 6 C and Figure 6 D). At 110℃, the microspheres were not fully bonded, with a porosity of 40.75±1.45%, and significant pore sizes existed between them. When the sintering temperature increased to 140℃, the bonding surface between the microspheres significantly increased, and the porosity was only 11.87±0.27%. Correspondingly, the compression test results showed that the compressive modulus of the scaffold reached a peak of 143.32 MPa at 140℃, which is nearly 10 times that of the scaffold sintered at 110℃. Figure 6(E). Furthermore, sintering at 120℃ and 130℃ also improved the mechanical properties of the scaffold, with compressive moduli of (14.74±3.40) MPa and (29.46±10.28) MPa, respectively, but the porosity decreased more significantly. In summary, the pore size and mechanical properties of the scaffold can be adjusted by controlling the sintering temperature. While ensuring that the mechanical properties match those of human cancellous bone (compressive modulus of 10–2000 MPa), the micropore size of the scaffold is typically 250–500 μm to accommodate the transport of nutrients or metabolites within the bone tissue. Therefore, 110℃ was ultimately selected as the optimal sintering temperature for the scaffold.

[0200] (2) Physicochemical characterization of PTG / PHA microsphere scaffold

[0201] Figure 7 Image A shows the composite microsphere scaffold sintered at 110℃. The microsphere scaffolds with added PHA are all brownish-black in color. The surface morphology of the microsphere scaffolds was observed using scanning electron microscopy, as shown below. Figure 7 As shown in Figure B, PHA is uniformly dispersed on the surface and inside the microspheres, confirming successful encapsulation. The inventors also conducted statistical analysis on the particle size of the microspheres. Figure 7 C) All microspheres are larger than 350 μm, meeting the expected requirements. Hydrogel was injected into the pore structure of the scaffold to design a biomimetic scaffold with a gradient structure.

[0202] Bone defects typically require implant sites with superior mechanical properties. The compressive modulus of the microsphere scaffold is calculated using stress-strain curves, such as... Figure 7 As shown in Figure D, the increase in inorganic salt PHA (0–20 wt%) significantly improved the compressive modulus of the microsphere scaffold. For example, the compressive modulus of the PTG / 20PHA scaffold reached 124 ± 7.73 MPa, far exceeding the 21.8 ± 1.98 MPa of the original PTG scaffold, meeting the requirement of the average compressive modulus of cortical bone (100–230 MPa). Furthermore, the inventors found that incorporating hydrogel into the PTG / PHA / Ica scaffold further increased the compressive modulus of the PTG / 20PHA scaffold by approximately 20 MPa. This is attributed to the biphasic structure formed by the network between the microspheres and the hydrogel filling them. The abundant amino, carboxyl, and hydroxyl groups in the hydrogel form stable covalent and non-covalent interactions with the catechol surface groups of PHA. The "egg-box" structure formed by sodium alginate through cationic crosslinking supports the mechanical properties of the scaffold.

[0203] After incorporation of hydrogel, an increase in the porosity of the microsphere scaffold was observed. Figure 7E). The porosity of the PTG / 20PHA / Ica scaffold was 47.92±0.32%, significantly higher than that of the other components. Notably, the addition of 10% HA did not significantly increase the compressive modulus (42.46±11.71 MPa) and porosity (44.24±1.54%) of the scaffold, which also proves that the unmodified HA was not completely encapsulated in the microspheres, consistent with previous dispersion experiments and observations of the microsphere microstructure.

[0204] Previous research by He et al. has shown that PTMC-TCP microsphere scaffolds (porosity <45%, maximum compressive modulus 18.68±1.64 MPa) can also be prepared by thermal sintering. The PTG / 20PHA / Ica scaffold has a compressive modulus approximately 100 MPa higher than the PTMC-TCP scaffold, possesses the same inorganic salt dopant, and maintains sufficient porosity. The larger surface area of ​​the modified nHA leads to the densification of this scaffold, indicating that the incorporation of GA enhances the mechanical strength lacking in the PTMC scaffold during implantation.

[0205] The shape memory function of polymers fixes polymer segments into a temporary shape through vitrification, making bone defect filling mechanically programmable. This polymer then recovers its original shape under the influence of external factors (such as heat). Figure 7 Fa illustrates the process by which shape memory scaffolds fill irregular bone defects. For example... Figure 7 As shown in Figure Fb, PTG polymer was molded into a palm shape, subjected to a high temperature of 42°C to achieve a curled appearance, and then fixed at 0°C. Subsequently, the sample was placed at 37°C, simulating human body temperature. The inventors observed that the material quickly recovered the shape of the palm. Furthermore, the inventors also tested the shape memory function of the PTG / PHA microsphere scaffold. Figure 7 Fc) revealed that the star-shaped microsphere scaffold could rapidly recover its shape. Control experiments were conducted at 25℃ and 45℃ (…). Figure 7 (Fd). In contrast, at the normal surgical environment temperature of 25°C, the stent unfolds much more slowly, taking approximately 320 seconds to fully unfold, more than three times that at 37°C. At 45°C, the glass transition temperature of the synthetic polymer, the stent fully folds into a pentagon within 60 seconds, exhibiting three upward folding angles. This indicates that the stent cannot fully expand when the temperature approaches the material's glass transition temperature. However, as the temperature approaches the material's glass transition temperature, the recovery rate increases, and the shape memory function becomes more pronounced. Therefore, during surgery, the stent can be cryogenically fixed, and after implantation, it can recover its permanent shape upon stimulation by body temperature, ensuring close contact between the stent and the surrounding bone defect, thereby enhancing its support function.

[0206] (3) Degradation behavior of PTG / PHA microsphere scaffold

[0207] The inventors evaluated the compressive modulus of the scaffold before and after degradation. Figure 8 B), and combined Figure 8 SEM images of the microsphere scaffold after 8 and 12 weeks of in vitro degradation are shown in Figure A. After 12 weeks of degradation, the compressive modulus of the scaffold inevitably decreased. Although the compressive modulus of the PTG / 20PHA / Ica scaffold decreased from (165.23±8.64) MPa to (105.67±7.04) MPa, it still maintained the highest compressive modulus. The compressive moduli of the PTG / 20PHA group and the PTG / 15PHA group were (43±6.58) MPa and (40.6±3.55) MPa, respectively, with a statistically significant difference (p<0.05). Notably, compared with PTG / 20PHA, the decrease in compressive modulus of the PTG / 20PHA / Ica scaffold after 12 weeks of degradation did not exceed 50%. This may be because the degradation of PTG led to the large release of PHA particles, while the cross-linked sodium alginate encapsulation structure delayed the rapid release of PHA. Simultaneously, the hydrogel and microsphere scaffold form a continuous network, with a stable PTG / PHA matrix supporting the hydrogel and preventing structural collapse. This indicates that the scaffold maintains mechanical properties similar to the replaced bone tissue as much as possible during degradation, thus avoiding the consequences of stiffness differences and rapid decrease in bone density. In the mass loss test, PTG degraded the slowest, with a mass loss increase of only 7.14±1.12% after 12 weeks. The mass loss of PTG / 20PHA / Ica was 15.23±0.79%, followed by the PTG / 20PHA scaffold, with a mass loss of 13.39±0.71%. Figure 8 C). Furthermore, compared to the inventors' previous work, the PLLA-TMC scaffold showed almost no change in mass loss during the 3-month degradation process, indicating that the introduction of the GA unit accelerated the degradation rate of the polymer, while the addition of PHA further improved the degradation performance of the microsphere scaffold.

[0208] At weeks 2 and 4, the loss rates of the PTG / 20PHA / Ica scaffolds were 9.03±0.62% and 11.33±0.89%, respectively, significantly higher than those of the PTG / 20PHA scaffolds (4.81±1.98% and 8.73±1.38%). This was primarily due to the preferential breakage of glycosidic bonds in the outer sodium alginate hydrogel during the early stages of degradation, followed by the degradation of the inner polymer under the influence of water molecules. Ideally, the hydrogel possesses good biocompatibility and a porous structure, providing a favorable platform for cell adhesion. As the hydrogel degrades, the inner polymer microspheres begin to dissolve and degrade, further inducing tissue cell proliferation and differentiation. Therefore, this gradient degradation pattern is more suitable for the growth of new bone tissue.

[0209] After the second week, the pH of the scaffold degradation solution began to decrease due to the hydrolysis of ester bonds on the ternary polymer molecular chains under the influence of water molecules. The decrease in PTG was the most significant, from 7.35±0.08 to 6.97±0.04, while the PTG / 20PHA / Ica ratio only decreased from 7.33±0.33 to 7.29±0.04. This alkaline degradation environment indicates that the PTMC degradation products are not acidic and, together with PHA, promote the acidic degradation of PLA. In this weakly alkaline environment, the presence of hydroxyl groups strongly attacks the electrophilic groups of the PLA chains, accelerating PLA degradation. This further alleviates the problem of low polymer degradation rate. After TMC surface degradation and PHA particle release, the scaffold exhibited good porosity (…). Figure 8 D). The porosity of the scaffolds increased to varying degrees after 12 weeks of degradation. In the water absorption test ( Figure 8 E) The PTG scaffold consistently maintained the lowest water absorption rate (45.13±5.58%~54.25±4.26%), while the water absorption rates of PTG / 20PHA and PTG / 20PHA / Ica increased to 78.77±4.41% and 84.65±2.05%, respectively. This indicates that PHA effectively increased the water absorption rate of the scaffold, and the hydrophilic hydroxyl and carboxyl groups distributed on the main chain determine the strong hydrophilicity of sodium alginate.

[0210] (4) Evaluation of in vitro biocompatibility and osteogenic behavior of microsphere scaffolds

[0211] Icariin (Ica) is a cyclic flavonoid compound extracted from traditional Chinese herbal medicine. It possesses the ability to enhance osteoblast differentiation and proliferation, promote calcification, and promote periosteal formation. Ica standard concentration curves were constructed by measuring the absorbance values ​​of different concentrations of Ica standard solutions at 273 nm. The fitted linear curve equation was y = 0.9760x, R² = 0.9953 > 0.99. Figure 9In Sc B, the amount of drug released gradually increased with increasing release time. During the first 4 days, the overall drug release showed a linear and stable trend, but the microsphere scaffold released a large amount of icariin (45±2.12%), which may be related to the rapid diffusion of the drug-loaded drug on the microsphere surface. After day 4, the drug release rate gradually slowed down, and after 10 days, the increase in the drug release rate became very slow with prolonged time. Finally, when the drug was released at day 25, the release rate remained at 75.12±1.27%. The slowing of the drug release rate during the prolonged release period may be due to the drug being present not only on the scaffold surface but also within the scaffold's microporous structure. Notably, the numerous microsphere gaps and the network formed by sodium alginate cross-linking provided storage areas for the drug. After the rapid surface release of drug molecules, these molecules continued to be slowly and orderly discharged from the interior of the scaffold. Therefore, the characteristics of early burst release and late sustained release of the drug demonstrate the excellent encapsulation effect of the hydrogel microsphere scaffold, enabling icariin to promote the early differentiation of mesenchymal stem cells into osteoblasts. Meanwhile, the sustained release in the mid-to-late stages can continuously affect the proliferation and differentiation of osteoblasts, while inhibiting the development of osteoclasts.

[0212] As is well known, the implantation of biomaterials should take into account their cell compatibility. Ideal biomaterials should be able to promote the proliferation and differentiation of osteoblasts and their precursor cells, such as... Figure 9 As shown in Figure A, the strong green fluorescence indicates that MC3T3-E1 cells adhere firmly to the scaffold surface, indicating no scaffold cytotoxicity. Combined with semi-quantitative fluorescence intensity analysis (…), Figure 9 C) It was observed that on day 1, only a small number of MC3T3-E1 cells adhered to PTG and PTG / 10HA scaffolds, while cell adhesion was more concentrated on PTG / 20PHA and PTG / 20PHA / Ica scaffolds. By day 7, MC3T3-E1 cells on the PTG / PHA / Ica scaffold showed the highest fluorescence intensity, which is consistent with the absorbance measurements. Figure 9 D). The OD values ​​of the PTG scaffold group were significantly lower than those of the PHA scaffold group on days 1, 3, and 7. This may be due to the smooth surface of the PTG microspheres, which makes cell adhesion difficult. The addition of PHA not only increased the roughness of the microsphere surface but also increased cell adhesion due to the "barnacle effect" of PHA on the microsphere surface. In addition, the high similarity of HA to the inorganic components of the bone matrix and the Ca2+ released by HA can promote osteocyte migration by stimulating (PO4)3-. The inventors also determined the cell viability ( Figure 9 E) After treatment with PTG, PTG / 10HA, PTG / 20PHA, and PTG / 20PHA / Ica, the cells showed good proliferation. Interestingly, as... Figure 9As shown in Figure B, cells first grow at the junctions of the microspheres, then gradually diffuse and occupy most of the surface of the microspheres, eventually encapsulating them. Simultaneously, cytoskeleton staining ( Figure 9 G) shows that MC3T3-E1 cells have the highest density at the microsphere junctions, indicating that this stacking structure between microspheres provides a good platform for cell proliferation and adhesion.

[0213] ALP is an early marker of osteogenic alkaline phosphatase (ALP). Staining and activity analysis were used to assess the bone-strengthening effect of PLLA-TMC-GA / PHA / Ica. For example... Figure 9 As shown in Figure H, significant differences in ALP staining were observed under a microscope at 3 and 7 days after osteogenic induction between the PTG group and the PTG / 20PHA and PTG / 20PHA / Ica groups. Furthermore, macroscopic images show that the PTG / 20PHA / Ica group exhibited the highest degree of ALP staining. Figure 9 I). Day 7 marks a crucial stage in osteoblast differentiation. ALP activity analysis ( Figure 9 F) showed that the differences in ALP protein expression gradually widened over time, with the mean ALP expression level in the PTG / 20PHA / Ica group significantly higher than that in other groups on day 14. The mean ALP expression level in the PTG / 20PHA group significantly increased with increasing PHA content. This may be related to the early release of Ca from PHA. 2+ It is associated with the induction of osteoblast differentiation. At the same time, PDA has significant adhesion ability, which can recruit more signaling proteins and affect MC3T3-E1 cells.

[0214] (5) Evaluation of the in vivo osteogenic effect of microsphere scaffolds

[0215] To determine the ability of the PTG / 20PHA / Ica microsphere scaffold to repair bone defects, the inventors established a rat model of skull defects. Figure 10 A). It was divided into 5 groups: blank group, PTG group, PTG / 10HA group, PTG / 20PHA group, and PTG / 20PHA / Ica group. From a macroscopic perspective ( Figure 10 B), at 6 weeks post-operation, the unfilled group still had unabsorbed hematoma and calcified tissue, delaying bone defect repair. Until 12 weeks post-operation, no significant inflammatory response was observed around the scaffold-filled defect. The scaffold remained at the defect site and fused well with the surrounding tissue, particularly with a repair membrane covering it, indicating good biocompatibility and biodegradability.

[0216] Skull bones were collected at 6 and 12 weeks, and bone formation was assessed using MICRO-CT. Figure 10C shows the presence of the microsphere scaffold in the defect area, and the scaffold components exhibited different degrees of degradation during new tissue growth. Notably, at 12 weeks, the PTG / 20PHA and PTG / 20PHA / Ica groups showed significant scaffold degradation, with a small number of microsphere particles visible, consistent with previous in vitro degradation experiments. After adjusting the threshold, the newly formed bone (indicated in red) was marked. Figure 10 (D) During the repair process, new bone development proceeds from the periphery towards the center. At 6 weeks, compared with other groups, the PTG / 20PHA and PTG / 20PHA / Ica groups showed significant bone formation at the defect periphery. With increasing time points (6 and 12 weeks), the control group showed the least amount of new bone formation around the bone defect, while the degree of new bone growth in all other groups was significantly higher than that in the control group. Among them, the PTG / 20PHA / Ica group showed the fastest new bone growth, with complete bone bridge formation observed, confirming that PTG / 20PHA / Ica has good bone induction capabilities.

[0217] Quantitative micro-CT analysis of bone tissue was performed, including estimation of bone volume fraction (BV / TV) and bone mineral density (BMD). Figure 10 E, Figure 10 F). At week 6, the BV / TV of the PTG / PHA / Ica scaffold reached 18.96% ± 1.83%, 4.4 times that of the control group. After 12 weeks, the PTG / PHA / Ica scaffold had the highest percentage of new bone formation, with a BV / TV of 25.74% ± 2.96%, exceeding that of other components, indicating significant new bone formation in the defect area. Figure 10 E).

[0218] The dynamic degradation behavior of the scaffolds plays a crucial role in this specific bone regeneration process. At 12 weeks, the in vitro mass loss rates of PTG, PTG / 10HA, PTG / 20PHA, and PTG / PHA / Ica scaffolds were all lower than the in vivo bone volume / total volume (BV / TV) percentages of newly formed bone (7.14±1.12% vs. 13.46±2.85%, 11.86±0.85% vs. 18.18±0.94%, 12.72±1.65% vs. 20.43±0.76%, and 15.23±0.79% vs. 31.0±3.9%, respectively). PTMC is a fatty polyester that is primarily degraded by enzymatic hydrolysis and aqueous solutions via lipases, thus exhibiting a faster degradation rate in vivo. Consequently, it increases the space for new bone formation until tissue regeneration is successfully completed.

[0219] Similarly, bone mineral density (BMD) is also an important indicator for assessing bone strength. Higher BMD indicates a larger volume of new bone and better support. At week 6, the BMD of the defect area in the unfilled scaffold group was only (74.25±5.62) mg HA / ccm, increasing to (185.48±12.14) mg HA / ccm at week 12, but still significantly lower than other groups. Figure 10 F). In the PTG / PHA / Ica scaffold group, the area with the highest bone mineral density at week 6 was (260.68±9.16) mg HA / ccm, and at week 12 it was (315.93±16.93) mg HA / ccm. This indicates that the filling scaffold can effectively replace bone trabeculae, forming a three-dimensional network structure that facilitates nutrient transport within the medullary cavity. Intrascaffold Ca 2+ The continuous release of Ica helps form a natural matrix structure around the tissue, enabling newly formed bone tissue to withstand mechanical loads in the defect area.

[0220] Histological analysis was performed at 6 and 12 weeks postoperatively to assess bone formation. Notably, in the PTG / PHA / Ica group, bone regeneration was already evident, and more importantly, fibrous tissue and new bone growth could be observed in the gaps between the scaffold microspheres. Figure 11 A). The porous microsphere scaffold exhibits good osteoconductive properties. At week 12, except for a small amount of new bone formation within the PTG microsphere scaffold, the other two groups showed significant induction of new bone ingrowth. The generated bone tissue had a certain interfacial bridging with the original bone. Masson staining, a technique commonly used to stain collagen fibers, confirmed the large amount of bone matrix formation induced by the PTG / PHA / Ica scaffold, thus highlighting the scaffold's ability to promote bone regeneration and repair. Figure 11 B). The degree of scaffold degradation varied among different microsphere scaffolds within the tissue. At week 6, undegraded microsphere components were still visible in the PTG and PTG / PHA scaffolds, while no scaffold component residue was observed in the PTG / PHA / Ica scaffold. This difference can be attributed to Ica promoting the circulation of metabolites within the neovascular microcirculation, thereby inducing premature hydrogel decomposition and accelerating the degradation of the PTG / PHA / Ica scaffold. Furthermore, this method also incorporates the degradation of polymers by lipases, consistent with patterns observed in previous in vivo and in vitro experiments.

[0221] This invention provides a simple and easy-to-operate method for preparing PLLA-TMC-GA / PHA microspheres and their scaffolds, and enhances the mechanical strength of the scaffolds by incorporating dopamine-modified nano-hydroxyapatite. The three-dimensional network structure constructed from the hydrogel and microspheres facilitates a gradient degradation pattern. Furthermore, co-culturing MC3T3-E1 cells with the scaffold demonstrates good biocompatibility, and the hydrogel loaded with active ingredients such as icariin on the microsphere scaffold can effectively catalyze osteogenic differentiation of osteoblasts. In vivo experiments confirm that the scaffold promotes new bone ingrowth and bone bridge formation; this multi-layered composite scaffold exhibits excellent repair effects on irregular bone defects.

Claims

1. Use of a polylactic-trimethylene carbonate-polyglycolic acid terpolymer in the preparation of a material for repairing bone defects.

2. Use according to claim 1, characterized in that: The preparation method of the polylactic-trimethylene carbonate-polyglycolic acid terpolymer comprises the following steps: (1) mixing lactic acid LA, trimethylene carbonate TMC and glycolide GA three monomers according to the following mass ratio: LA 50-80 parts, TMC 10-30 parts, and GA 10-30 parts; (2) adding stannous octoate to the mixture obtained in step (1) and drying in a reaction container; (3) polymerizing at 130-150℃ for 1-2h to obtain a product; (4) dissolving the product obtained in step (3) in dichloromethane, then adding anhydrous ethanol dropwise to uniformly precipitate the material, and then drying to obtain the polylactic-trimethylene carbonate-polyglycolic acid terpolymer; Preferably, at least one of the following is met in the above preparation method: In step (1), the lactic acid LA, trimethylene carbonate TMC and glycolide GA three monomers are mixed according to the following weight ratio: LA 72 parts, TMC 18 parts, and GA 10 parts; Or, in step (2), stannous octoate is added at 0.01-0.05% by weight of the mixture; Or, preferably, in step (2), stannous octoate is added at 0.05% by weight of the mixture; Or, in step (2), the drying is vacuum drying for 1h using a vacuum water pump, a vacuum oil pump and a diffusion pump in sequence; Or, in step (3), the polymerization is preferably carried out at 140℃ for 48h; Or, in step (4), the drying is oven drying at 30-40℃ in a vacuum environment; Or, preferably, in step (4), the drying is oven drying at 37℃ in a vacuum environment.

3. Poly (lactic-co-trimethylene carbonate) -polyglycolic acid ternary polymer microspheres loaded with polydopamine-modified nanohydroxyapatite, characterized in that: Prepared from the polylactic-trimethylene carbonate-polyglycolic acid terpolymer and polydopamine-modified nano-hydroxyapatite.

4. The method for preparing the polydopamine-modified nano-hydroxyapatite- loaded polylactic-co-trimethylene carbonate-co-polyglycolic acid ternary polymer microspheres according to claim 3, characterized in that: Comprising the following steps: A. Dissolve the polylactic-trimethylene carbonate-polyglycolic acid terpolymer in a solvent to form an oil phase, then add polydopamine-modified nano-hydroxyapatite to form a mixed phase; B. Prepare a polyvinyl alcohol aqueous solution as an aqueous phase; C. Add the mixed phase obtained in step A to the aqueous phase obtained in step B dropwise and stir to obtain an emulsion; D. Add water to the emulsion obtained in step C and stir to obtain dispersed microspheres; E. Filter, wash and dry the solidified microspheres obtained in step D to obtain polylactic-trimethylene carbonate-polyglycolic acid terpolymer microspheres loaded with polydopamine-modified nano-hydroxyapatite.

5. The method for preparing poly (dopamine modified nano-hydroxyapatite) loaded poly (lactic-co-trimethylene carbonate-co-glycolic acid) microspheres according to claim 3, characterized in that: At least one of the following is met: In step A, dissolve 1g of polylactic-trimethylene carbonate-polyglycolic acid terpolymer in 5-20ml of solvent according to the following ratio; Or, preferably, in step A, dissolve 1g of polylactic-trimethylene carbonate-polyglycolic acid terpolymer in 15ml of solvent according to the following ratio; Or, in step A, the solvent is dichloromethane, dioxane, dimethyl sulfoxide and acetone; Or, preferably, in step A, the solvent is dichloromethane; Or, in step A, the amount of polydopamine-modified nano-hydroxyapatite added is 5-20wt% of the oil phase in step A; Or, preferably, in step A, the amount of polydopamine modified nano-hydroxyapatite added is 20wt% of the oil phase in step A; Or, in step A, the polylactic acid-trimethylene carbonate-polyglycolic acid ternary polymer is dissolved in dichloromethane using magnetic stirring a until completely dissolved; Or, in step A, the stirring speed of the magnetic stirring a is 300-350 rpm; Or, preferably, in step A, the stirring speed of the magnetic stirring a is 300 rpm; Or, in step A, the polydopamine modified nano-hydroxyapatite is added to the mixed phase to form a mixture, which is stirred using magnetic stirring b until uniform; Or, in step A, the stirring speed of the magnetic stirring b is 400-500 rpm; Or, preferably, in step A, the stirring speed of the magnetic stirring b is 500 rpm; Or, in step B, the concentration of the polyvinyl alcohol aqueous solution is 0.1%-1% w / v, i.e. 10ml-1ml of polyvinyl alcohol is dissolved per 1g of polyvinyl alcohol; Or, preferably, in step B, the concentration of the polyvinyl alcohol aqueous solution is 0.1% w / v, i.e. 10ml of polyvinyl alcohol is dissolved per 1g of polyvinyl alcohol; Or, in step C, the mixed phase obtained in step A is added dropwise to the aqueous phase obtained in step B according to 100-150ml of aqueous phase per 1g of polylactic acid-trimethylene carbonate-polyglycolic acid ternary polymer in the mixed phase; Or, preferably, in step C, the mixed phase obtained in step A is added dropwise to the aqueous phase obtained in step B according to 120ml of aqueous phase per 1g of polylactic acid-trimethylene carbonate-polyglycolic acid ternary polymer in the mixed phase; Or, in step C, the stirring is performed using magnetic stirring; Or, in step C, the stirring speed of the magnetic stirring is 300-400 rpm; Or, preferably, in step C, the stirring speed of the magnetic stirring is 350 rpm; Or, in step D, the water is deionized water or phosphate buffered saline solution; Or, preferably, in step D, the water is deionized water; Or, in step D, 2-3 times the amount of water of the emulsion in step C is added; Or, preferably, in step D, 2.6 times the amount of water of the emulsion in step C is added; Or, in step D, the stirring is performed using magnetic stirring; Or, in step D, the stirring conditions are first reduced by 50-100 rpm based on step C, and then the stirring time is continued for 10-14h; Or, preferably, in step D, the stirring conditions are first reduced by 50 rpm based on step C, and then the stirring time is continued for 12h; Or, in step E, the filtration is performed using a material with a particle size of 300μm-500μm to obtain solidified microspheres; Or, in step E, the washing is performed 2-4 times with water; preferably, the washing is performed 3 times with water; Or, in step E, the water is deionized water or phosphate buffered saline solution; Or, preferably, in step E, the water is deionized water; Or, in step E, the drying is performed in a vacuum drying environment, and the drying temperature is 35-40℃; Or, preferably, in step E, the vacuum drying environment uses a vacuum oven; Or, preferably, in step E, the drying temperature is 37℃; Or, preferably, in step A, the solvent is dichloromethane.

6. A microsphere stent characterized by: The polylactic acid-trimethylene carbonate-polyglycolic acid ternary polymer microspheres loaded with the polydopamine modified nano-hydroxyapatite of claim 3 are compressed in a mold under pressure, and the obtained product is sintered by gradient sintering.

7. The method of claim 6, wherein the microscaffold is prepared by: The polylactic acid-trimethylene carbonate-polyglycolic acid ternary polymer microspheres loaded with the polydopamine modified nano-hydroxyapatite of claim 3 are compressed in a mold under pressure, and the obtained product is sintered by gradient sintering. Further, at least one of the following conditions is met: Or, the mold is a cylindrical mold with a size of 8mm x 4mm, a square mold with a size of 4mm x 4mm, or a triangular mold with a size of 4mm x 4mm x 4mm; Or, preferably, the mold is a cylindrical mold with a size of 8mm x 4mm; Or, the pressure for the compression under pressure is 50-150Mpa; Or, preferably, the pressure for the compression under pressure is 100Mpa; Or, the gradient sintering refers to sintering at temperatures of 110℃, 120℃, 130℃, and 140℃, respectively; Or, the gradient sintering is sintering at each temperature for 3-5h, respectively; Or, preferably, the gradient sintering is sintering at each temperature for 4h, respectively.

8. A microspheres scaffold of cross-linked sodium alginate hydrogel, characterized in that: The microsphere scaffold cross-linked sodium alginate hydrogel of claim 6 is prepared by the following method: (a) sodium alginate is prepared into a sodium alginate hydrogel; (b) the microsphere scaffold is soaked into the hydrogel obtained in step (a), cross-linked, and dried to form a hydrogel-microsphere scaffold. Further, at least one of the following conditions is met: The sodium alginate content in the sodium alginate hydrogel of step (a) is 2-4wt%; Or, preferably, the sodium alginate content in the sodium alginate hydrogel of step (a) is 3wt%; Or, the cross-linking condition of step (b) is cross-linking in a calcium ion solution with a content of 0.1-1% for 1-6h; Or, preferably, the cross-linking condition of step (b) is cross-linking in a calcium ion solution with a content of 0.6% for 4h; Or, the drying of step (b) is freeze-drying.

9. Drug-loaded hydrogel-microsphere composite scaffold, characterized in that: The microsphere scaffold loaded with a hydrogel containing a pharmaceutically active ingredient of claim 6 is prepared by the following method: (a1) an active ingredient and sodium alginate are prepared into a drug-loaded hydrogel; (b1) the microsphere scaffold is soaked into the drug-loaded hydrogel obtained in step (a1), cross-linked, and dried to form a drug-loaded hydrogel-microsphere composite scaffold; Further, at least one of the following conditions is met: Or, the sodium alginate content in the sodium alginate hydrogel of step (a1) is 2-4wt%; Or, preferably, the sodium alginate content in the sodium alginate hydrogel of step (a1) is 3wt%; Or, the active ingredient of step (a1) is selected from icariin; Or, the liquid concentration of icariin of step (a1) is 5-6mg / mL; Or, preferably, the liquid concentration of icariin of step (a1) is 5.4mg / mL; Or, the preparation method of the hydrogel of step (a1) is as follows: Or, 0.6-1.2g of sodium alginate is added to the solution after stirring, and then it is completely dissolved; Or, preferably, 0.162 g of icariin is dissolved in 30 mL of deionized water, stirred uniformly, then 0.9 g of sodium alginate is added and fully dissolved; Or, the cross-linking condition in step (b1) is cross-linking in a calcium ion solution with a content of 0.1-1% for 1-12 h; Or, preferably, the cross-linking condition in step (b1) is cross-linking in a calcium ion solution with a content of 0.6% for 4 h; Or, the drying in step (b1) is freeze-drying; Or, further, the freeze-drying in step (b1) is pre-freezing at-20-80 ℃ for 10-100 min, and freeze-drying for 24-48 h; Or, preferably, the freeze-drying in step (b1) is pre-freezing at-80 ℃ for 20 min, and freeze-drying for 24 h.

10. Use of the polydopamine-modified nanohydroxyapatite-loaded polylactic acid-trimethylene carbonate-polyhydroxyacetic acid ternary polymer microspheres in claim 3, the microsphere scaffold in claim 6, the cross-linked sodium alginate hydrogel microsphere scaffold in claim 8, and the drug-loaded hydrogel-microsphere composite scaffold in claim 9 in the preparation of a material for repairing bone defects.

Images