Shape memory bone repair material and preparation method thereof

Abstract

The invention belongs to the technical field of biomedical engineering, and relates to a shape memory bone repair material and a preparation method thereof. According to the bone repair material, a hydrogel material is filled in pores of a bone repair scaffold shape memory polymer material, and a photothermal conversion material is added into the hydrogel material. According to the shape memory bone repair material and the preparation method thereof disclosed by the invention, after the prepared bone repair material is implanted into a bone defect area, shape change can be realized at a transformation temperature of 38-42 DEG C, so that the edge of a bracket is tightly attached to the edge of the bone defect area, and the bracket keeps necessary strength on the premise of not causing body injury; and bone defect repair is promoted.

Description

Technical Field

[0001] This invention belongs to the field of biomedical engineering technology and relates to a shape memory bone repair material and its preparation method. Background Technology

[0002] Currently, bone defects have a high incidence rate and are difficult to regenerate and repair. In the treatment of bone defects, 3D-printed bone repair scaffolds offer advantages such as precise structural control and personalized fabrication based on the defect area. However, because the shape of a 3D-printed scaffold cannot be altered, its shape must be modified when implanted into complex bone defect areas. This not only increases unnecessary trauma but also leads to a misalignment between the scaffold and the defect area, significantly affecting the healing outcome.

[0003] If we can add a fourth dimension of time response to the 3D printed scaffold based on shape memory materials, we can prepare a 4D printed scaffold and give it adjustable shape characteristics. That is, before implantation, the scaffold can be shrunk; after implantation, the scaffold can return to its initial shape under external stimulation. During bone repair, it is not necessary to remove healthy tissue, and the scaffold can also fit closely to the edge of the defect.

[0004] Chinese patent application CN 111905153 A uses shape memory polyurethane composite materials to repair orbital bone defects, but the transition temperature of this material is 40–60°C, while in clinical practice, temperatures above 42°C can cause osteonecrosis. Furthermore, polyurethane contains diisocyanates, resulting in poor biocompatibility. Chinese patent application CN 105944144 A uses shape memory composite materials containing Fe3O4 magnetic nanoparticles, which tend to aggregate in the liver and spleen, also indicating poor biocompatibility. While Chinese patent application CN 101554488 B proposes a method for initially determining the proportions of different components of shape memory polymers within a transition temperature range of 5–37°C, the actual transition temperature of the obtained material differs from the calculated formula. Experimental exploration to obtain the component proportions within the target transition temperature range remains crucial. Chinese patent application CN 109364302 A reports tissue engineering scaffolds composed of different shape memory polymers, but does not test the shape transition temperature of these scaffolds.

[0005] It is important to emphasize that, for in vivo applications, the transition temperature of shape memory polymer materials should be set between 38 and 42°C. This not only allows for shape recovery without causing damage to the body, but also ensures that the scaffold retains a certain strength after the temperature returns to body temperature (37°C). Therefore, preparing shape memory materials with good biocompatibility and suitable transition temperatures for in vivo applications remains a challenging problem in this field.

[0006] Furthermore, the inherent hydrophobic properties of shape memory polymer scaffolds hinder cell adhesion and migration, and they lack the biological functions to promote bone regeneration. Hydrogels, on the other hand, possess excellent biocompatibility, hydrophilicity, and high porosity, with properties similar to the extracellular matrix. Therefore, these two can be combined to achieve complementary advantages. Simultaneously, stimulus-responsive nanoparticles can be incorporated into the hydrogel, raising its temperature upon application of external stimuli and promoting the shape recovery of the 4D-printed scaffold; furthermore, the nanoparticles can carry bioactive components to promote the bone repair process. In this regard, Chinese patent application CN109701084 A reports a shape memory bioactive nanoparticle / biodegradable polyester composite scaffold. It involves adding biodegradable shape memory polyester, hydroxyapatite nanoparticles, and a rheology modifier to an organic solvent, followed by ultrasonic treatment to obtain an oil phase. Water is then added to the oil phase, and emulsification is performed to form a water-in-oil high internal phase emulsion. This high internal phase emulsion is used as ink, and extrusion 3D printing is employed to obtain a three-dimensional porous emulsion scaffold. The solvent is then evaporated to obtain the final biological scaffold. However, because this scaffold lacks the addition of stimulus-responsive nanoparticles, it cannot achieve shape recovery through external stimulation, thus failing to meet the requirements for in vivo applications.

[0007] Therefore, it is essential to develop bone repair scaffolds that combine 4D-printed scaffolds and hydrogels to match the shape of defective bone tissue, meet the needs of minimally invasive implantation, and at the same time have the biological function of promoting bone defect repair. Summary of the Invention

[0008] The primary objective of this invention is to provide a shape memory bone repair material that can change shape at a transition temperature of 38–42°C after implantation into a bone defect area, so that the edge of the scaffold fits tightly with the edge of the bone defect area, maintains the necessary strength of the scaffold without causing damage to the body, and promotes bone defect repair.

[0009] To achieve this objective, in a basic implementation, the present invention provides a shape memory bone repair material, wherein the bone repair material has hydrogel material filled into the pores of a shape memory polymer material of a bone repair scaffold, and the hydrogel material contains a photothermal conversion material.

[0010] In a preferred embodiment, the present invention provides a shape memory bone repair material, wherein:

[0011] The shape memory polymer material is selected from one or more of polylactic acid (PLA), polylactic acid-co-trimethylene carbonate (PLA-TMC), polylactic acid-co-glycolic acid (PLGA), polylactic acid-co-carprolactone (PLA-CL), polylactic acid-co-ethylene glycol (PLA-PEG), and polyurethane (PU), with polylactic acid-trimethylene carbonate being preferred;

[0012] The hydrogel material is selected from one or more of the following: methacrylated gelatin, methacrylated chitosan, methacrylated sodium alginate, methacrylated chondroitin sulfate, hydroxybutyl chitosan, sodium alginate, hyaluronic acid, chondroitin sulfate, polyethylene glycol diacrylate, and polyvinyl alcohol, with methacrylated gelatin being preferred.

[0013] The photothermal conversion material is selected from one or more of polydopamine, black phosphorus, gold nanorods, and carbon nanotubes, with polydopamine being preferred.

[0014] In a preferred embodiment, the present invention provides a shape memory bone repair material, wherein the shape memory polymer material is a polylactic acid-trimethylene carbonate copolymer, which is obtained by polymerizing trimethylene carbonate and L-lactic acid in a molar ratio of 1:(1.5-9.0), preferably by polymerizing trimethylene carbonate and L-lactic acid in a molar ratio of 1:4.

[0015] In a preferred embodiment, the present invention provides a shape memory bone repair material, wherein the bone repair material further contains hydroxyapatite in the shape memory polymer material and / or the hydrogel material.

[0016] In a preferred embodiment, the present invention provides a shape memory bone repair material, wherein the mass ratio of the shape memory polymer material to the added hydroxyapatite is (1-4):1, preferably 4:1; and the mass ratio of the hydrogel material to the added hydroxyapatite is 10:(1-10).

[0017] In a preferred embodiment, the present invention provides a shape memory bone repair material, wherein the bone repair material further contains bone-promoting drugs and / or bone-promoting bioactive ingredients added to the hydrogel material.

[0018] In a preferred embodiment, the present invention provides a shape memory bone repair material, wherein the osteogenic drug or the osteogenic bioactive ingredient is selected from one or more of pargeline, simvastatin, strontium ranelate, dexamethasone, teriparatide, bone morphogenetic protein 2 (BMP-2), lomoxozimilab, and osteogenic peptides (e.g., OGP, W9, CGRP, BFP-1, FGF-18, and GHK).

[0019] In a preferred embodiment, the present invention provides a shape memory bone repair material, wherein the mass ratio of the hydrogel material to the osteogenic drug and / or osteogenic bioactive ingredient added thereto is (10-100000000):1.

[0020] In a preferred embodiment, the present invention provides a shape memory bone repair material, wherein the mass ratio of the shape memory polymer material to the hydrogel material is 30:(1-2), and the mass ratio of the hydrogel material to the photothermal conversion material is 1:(0.005-0.5).

[0021] The second objective of this invention is to provide a method for preparing the above-mentioned bone repair material, so as to better prepare the above-mentioned bone repair material. The prepared bone repair material can change shape at a transition temperature of 38 to 42°C after being implanted into the bone defect area, so that the edge of the scaffold can fit tightly with the edge of the bone defect area, maintain the necessary strength of the scaffold without causing damage to the body, and promote the repair of bone defects.

[0022] To achieve this objective, in a basic embodiment, the present invention provides a method for preparing the above-mentioned bone repair material, the method comprising the following steps:

[0023] (1) Prepare the shape memory polymer material described above;

[0024] (2) Prepare a precursor solution of the hydrogel material containing the photothermal conversion material;

[0025] (3) Immerse the shape memory polymer material in the precursor solution for a certain period of time;

[0026] (4) Irradiate the soaked shape memory polymer material to form the hydrogel material from the precursor solution.

[0027] In a preferred embodiment, the present invention provides a method for preparing the above-mentioned bone repair material, wherein in step (1), the shape memory polymer material of the desired shape and size is prepared by using 3D printing technology.

[0028] In a preferred embodiment, the present invention provides a method for preparing the above-mentioned bone repair material, wherein the 3D printing parameters are: barrel temperature 45-50℃, nozzle size 0.4-0.8mm, layer height 0.2-0.6mm, nozzle moving speed 5-10mm / s, line spacing 0.4-0.8mm; extrusion speed 1-2mm. 3 / s.

[0029] In a preferred embodiment, the present invention provides a method for preparing the above-mentioned bone repair material, wherein in step (2), hydroxyapatite encapsulated with photothermal conversion material (preferably polydopamine) is added to the precursor solution, and preferably, osteopromoting drugs and / or osteopromoting bioactive ingredients, such as pargelin, are also added to the precursor solution.

[0030] In a preferred embodiment, the present invention provides a method for preparing the above-mentioned bone repair material, wherein in step (2), the precursor solution is formed by dissolving the precursor of the hydrogel material (preferably methacrylated gelatin) in a photoinitiator solution.

[0031] In a preferred embodiment, the present invention provides a method for preparing the above-mentioned bone repair material, wherein the photoinitiator is selected from one or more of lithium phenyl (2,4,6-trimethylbenzoyl) phosphate (LAP), 2-hydroxy-4'-(2-hydroxyethoxy)-2-methylphenylacetone (Irgacure 2959), and phenyl bis(2,4,6-trimethylbenzoyl)phosphine oxide (BAPO).

[0032] In a preferred embodiment, the present invention provides a method for preparing the above-mentioned bone repair material, wherein in step (3), the soaking time is 10 to 60 minutes.

[0033] In a preferred embodiment, the present invention provides a method for preparing the above-mentioned bone repair material, wherein in step (4), the irradiation is performed by irradiating with a light source with a wavelength of 360-410 nm for 3-10 minutes.

[0034] The beneficial effect of the present invention is that, by utilizing the shape memory bone repair material and its preparation method of the present invention, the prepared bone repair material can change its shape at a transition temperature of 38-42°C after being implanted into the bone defect area, so that the edge of the scaffold can fit tightly with the edge of the bone defect area, maintain the necessary strength of the scaffold without causing damage to the body, and promote the repair of bone defects.

[0035] Existing 3D-printed bone repair scaffolds, due to their immutable shape, suffer from difficulty in fitting the scaffold edges to the defect edges, significantly impacting healing outcomes. Regarding shape memory materials, existing technologies utilize shape memory alloys with poor biocompatibility, are non-degradable, and have a 6% allergy rate to nickel. Furthermore, the polymers used in current technologies lack suitable transition temperatures for in vivo application; temperatures above 42°C can lead to osteonecrosis, while temperatures below body temperature result in lower strength within the body, failing to meet the mechanical requirements for bone defect repair. Therefore, shape memory polymer materials with appropriate transition temperatures are needed. Additionally, the inherent hydrophobic properties of existing polymer bone repair scaffolds hinder cell adhesion and migration, and they lack the biological functions to promote bone defect repair.

[0036] This invention, by adjusting the proportions of different polymers in the shape memory polymer and the proportion of inorganic materials incorporated, not only obtains a bone scaffold material with a transition temperature suitable for clinical applications, but also solves the shape adaptation problem of traditional 3D printed bone repair scaffolds. Furthermore, by coating the outer layer of the shape memory scaffold material with hydrogel, it solves the problem that polymer scaffolds are hydrophobic and lack the function of promoting bone regeneration. Attached Figure Description

[0037] Figure 1 This is a flowchart illustrating the digital design, preparation, and clinical application of the bone repair material in Example 1.

[0038] Figure 2 The figures show the mechanical property test results of P(DLLA-TMC) and P(LLA-TMC). Figures a and b show the elastic modulus and tensile strength of P(DLLA-TMC) and P(LLA-TMC), respectively.

[0039] Figure 3 The results show the effects of the ratio of L-lactic acid (LLA) and trimethylene carbonate (TMC) on the P(LLA-TMC) transition temperature and the effects of the ratio of nano-hydroxyapatite (nHA) to P(LLA-TMC) on the P(LLA-TMC) / nHA transition temperature in Example 2.

[0040] Figure 4 The results show the effect of the ratio of nHA to P(LLA-TMC) on the mechanical properties of P(LLA-TMC) / nHA. Figure a shows the elastic modulus, and Figure b shows the tensile strength.

[0041] Figure 5 The image shows the scanning electron microscope results of the P(LLA-TMC) / nHA scaffold.

[0042] Figure 6 The graph shows the mechanical performance test results for different supports.

[0043] Figure 7 The images show the transmission electron microscopy (TEM) results of different nanoparticles.

[0044] Figure 8 Figures show the characterization results of different materials, where: Figure a shows the SEM detection results of different hydrogel materials; Figure b shows the SEM detection results of cross-sections of different scaffolds; and Figure c shows the water contact angle detection results of different scaffolds.

[0045] Figure 9 The temperature response shape memory effect diagrams of P(LLA-TMC) / nHA and P(LLA-TMC) / nHA+GelMA / nHA@PDA-PGL stents are shown, where: Figure a shows the shape recovery process of the two stents at 42℃; Figure b shows the shape fixation rate of the two stents at 4℃ and the shape recovery rate at 42℃.

[0046] Figure 10 Figures show the photothermal response detection results of P(LLA-TMC) / nHA, P(LLA-TMC) / nHA+GelMA, P(LLA-TMC) / nHA+GelMA / nHA@PDA, and P(LLA-TMC) / nHA+GelMA / nHA@PDA-PGL. Figure a shows the temperature change curves of different scaffolds under NIR light irradiation over time; Figure b shows the temperature change curves of P(LLA-TMC) / nHA+GelMA / nHA@PDA-PGL under NIR light irradiation of different intensities over time; Figure c shows the temperature change curves of P(LLA-TMC) / nHA+GelMA / nHA@PDA-PGL under NIR light irradiation of 1.2 W / cm². 2 The photothermal stability test results of P(LLA-TMC) / nHA+GelMA / nHA@PDA-PGL under NIR light irradiation are shown in the figure.

[0047] Figure 11 The image shows the results of the near-infrared light response shape recovery rate test for different supports.

[0048] Figure 12 Figure 1 shows the results of detecting the osteogenic differentiation of hBMSCs in vitro using different scaffolds. Figure 2 shows the expression of genes related to osteogenic differentiation of hBMSCs. Figure 3 shows the alkaline phosphatase activity of hBMSCs.

[0049] Figure 13 The images show the detection results of different scaffolds promoting the repair of critical bone defects in the rat skull. Figure a shows a three-dimensional reconstruction image of the rat skull and materials obtained by micro-CT scanning; Figure b shows the bone volume / tissue volume (BV / TV) ratio based on bone histomorphometric analysis. Detailed Implementation

[0050] The following examples further illustrate specific embodiments of the present invention.

[0051] Example 1: Digital design, preparation, and efficacy evaluation process of bone repair materials

[0052] The operation process of this embodiment is as follows: Figure 1 As shown.

[0053] (1) Obtain three-dimensional data of the bone defect area through CT scan.

[0054] Three-dimensional imaging data of craniofacial bone defects were obtained through spiral CT scans. Using Mimics software, three-dimensional reconstruction of the craniofacial region was performed based on the DICOM data from the CT scans, and the data was exported as an STL file. The STL file was then imported into the 3D reverse engineering software Geomagic Studio for further design. In Geomagic Studio, a triangular facet mesh was fixed, and noise from the DICOM data was cleaned.

[0055] (2) Design and preparation of personalized composite 4D printed bone repair materials

[0056] Based on CT scan data, personalized 3D scaffolds were generated using the modeling software nTopology, with porosity, pore size, and filament diameter set according to different needs. The personalized 3D scaffolds were exported as STL files for subsequent printing. 4D printing ink was prepared, and personalized 4D-printed bone repair scaffolds were printed using an extrusion-type bio-3D printer. Because solvent evaporation would cause the printed scaffold volume to shrink, multiple cubic scaffolds were first printed to determine the volume shrinkage rate. Based on this, the STL model was enlarged so that the edge of the final printed scaffold was 1 mm larger than the bone defect boundary. The scaffolds were dried in a vacuum drying oven for 72 hours to fully remove the solvent, obtaining a preliminary personalized 4D-printed scaffold, and then sterilized by gamma ray (25 kGy) radiation. Then, under sterile conditions, a hydrogel precursor solution was prepared, and the 4D-printed scaffold was immersed in the hydrogel precursor solution and incubated on a constant-temperature horizontal shaker for 10 minutes to allow the hydrogel precursor solution to fully wet the scaffold pores. The 4D-printed scaffold was then removed, and ultraviolet light irradiation was applied to cross-link the hydrogel, resulting in a personalized composite 4D-printed bone repair material with a hydrogel-coated surface.

[0057] (3) Evaluate the ability of personalized 4D printed bone repair materials to adapt to complex bone defects and promote bone defect healing.

[0058] Personalized 4D printed bone repair materials at temperatures above T transAfter being shrunk in the environment, the scaffold was implanted into the bone defect site. Subsequently, the NIR irradiation group underwent NIR irradiation immediately after scaffold implantation until the scaffold shape was restored. CT scans were used to reconstruct three-dimensional images, and the degree of fit between the scaffold and the bone defect edge was calculated based on the contact area between the scaffold and the bone defect edge. Simultaneously, bone regeneration was evaluated through bone morphometric analysis.

[0059] Example 2: Preparation of bone repair materials and study on preparation conditions

[0060] (1) Study on the influence of lactic acid crystal morphology

[0061] Good mechanical strength is an important requirement for bone repair materials. The crystallinity of lactic acid (LA) used in polylactic acid-trimethylene carbonate (PLA-TMC) can affect the mechanical properties of the prepared PLA-TMC copolymer. Therefore, P(DLLA-TMC) and P(LLA-TMC) copolymers (Jinan Daigang Biotechnology Co., Ltd., China) prepared from amorphous racemic lactic acid (DLLA) and semi-crystalline L-lactic acid (LLA) were used to prepare samples for mechanical property testing via solvent casting. Their mechanical properties were compared to select a more ideal PLA-TMC material for subsequent experiments. The specific method is as follows:

[0062] Using a universal mechanical testing machine (UTM-5105), In China, tensile strength of P(DLLA-TMC) and P(LLA-TMC) samples was gradually increased at a rate of 1.0 mm / min until fracture at room temperature, and stress-strain curves were generated by recording force and displacement data. The elastic modulus of the polylactic acid-trimethylene carbonate copolymer was calculated by measuring the slope of the stress-strain curve in the linearly changing region (~10%), and the tensile strength was determined by the end stress of the linearly changing region. The results showed that the elastic modulus and strength of P(LLA-TMC) were 38.1±5.8 MPa and 10.7±2.9 MPa, respectively, both significantly higher than those of P(DLLA-TMC) (1±0.4 MPa and 0.5±0.1 MPa). Figure 2 Therefore, P(LLA-TMC) was chosen for subsequent research.

[0063] (2) Study on the effect of LLA to TMC ratio on P(LLA-TMC) transition temperature

[0064] The ratio of LLA to TMC affects the transformation temperature (T0) of the polymerized P(LLA-TMC). transTherefore, P(LLA-TMC) (abbreviated as 91P(LLA-TMC), 82P(LLA-TMC), and 73P(LLA-TMC), respectively, purchased from Jinan Daigang Biotechnology Co., Ltd.) with different LLA to TMC ratios (9:1, 8:2, and 7:3, w / w) were dissolved in 1,4-dioxane (1:4, w / v) and stirred at 300 rpm for 2 hours at room temperature. Then, samples for differential scanning calorimetry (DSC) testing were prepared in a polytetrafluoroethylene mold using a solvent casting method. The samples for DSC testing were then dried in a vacuum drying oven for 72 hours, and test samples were obtained using a punch according to the sample requirements of the experimental instrument. Under nitrogen protection, thermal analysis was performed on samples 91P (LLA-TMC), 82P (LLA-TMC), and 73P (LLA-TMC) using a differential scanning calorimeter (DSC214, NETZSCH, Germany). The specific procedure was as follows: First, the samples were heated from 0°C to 110°C in the first heating cycle; then cooled to 0°C at a rate of 10°C / min to eliminate their thermal history; the thermal history (T0) was determined from the DSC curves of the second heating cycle (from 0°C to 100°C, with a heating rate of 10°C / min). trans Values. The results showed that the T values ​​for 91P (LLA-TMC), 82P (LLA-TMC), and 73P (LLA-TMC) were... trans The temperatures were 40.6±1.0℃, 31.6±1.1℃, and 23.3±1.7℃, respectively. Figure 3 a) 82P (LLA-TMC) was selected for subsequent research.

[0065] (3) Study on the ratio of nHA to P(LLA-TMC) in 4D printing ink preparation

[0066] Nano-hydroxyapatite can make the scaffold more similar to the chemical composition of natural bone, increasing the osteogenic activity of the scaffold. Therefore, 82P (LLA-TMC) was blended with nHA to prepare 4D printing ink. Nano-hydroxyapatite (nHA) (… The mass ratios of nHA (nHA) and 82P(LLA-TMC) were set to 1:4, 1:2, and 1:1.33, respectively. Then, different masses of nHA were dispersed in 1,4-dioxane according to these ratios and ultrasonically agitated (300W) for 5 minutes to ensure uniform dispersion. Next, the corresponding mass of 82P(LLA-TMC):1,4-dioxane was dissolved in the above solution to achieve a 1:5 (w / v) ratio. The mixture was then manually stirred for 5 minutes, followed by stirring at 300 rpm for 12 hours at room temperature. Samples for DSC analysis were then formed in a PTFE mold using solvent casting. DSC results showed that as the nHA content in the P(LLA-TMC) / nHA composite increased, the Tg of the P(LLA-TMC) / nHA composite increased. trans The temperature increased from 33.8 ± 0.8 °C (nHA: 82P(LLA-TMC) = 1:4) to 36.0 ± 0.7 °C (nHA: 82P(LLA-TMC) = 1:2) and 38.8 ± 0.6 °C (nHA: 82P(LLA-TMC) = 1:1.33); however, when nHA: 82P(LLA-TMC) increased to 1:1, T trans The temperature dropped to 34.9 ± 1.5℃. Figure 3 b).

[0067] Next, the effect of nHA content on the mechanical properties of P(LLA-TMC) / nHA composites was further investigated using a universal testing machine. The results showed that with the increase of nHA content in the P(LLA-TMC) / nHA composites, the elastic modulus and tensile strength of the composites gradually increased. Both the elastic modulus and tensile strength reached their maximum when nHA:82P(LLA-TMC) = 1:1.33, and decreased when nHA:82P(LLA-TMC) = 1:1. Figure 4 Therefore, a P(LLA-TMC) / nHA composite material with a mass ratio of nHA:82P(LLA-TMC) of 1:1.33 was ultimately chosen for 4D printing ink preparation because of its T0 at ~40℃. trans It not only avoids damage to tissues / cells, but also has good mechanical properties.

[0068] (4) Preparation of 4D printed scaffold P(LLA-TMC) / nHA

[0069] A 4D printing scaffold was printed using an extrusion-type bio-3D printer (DHD CellPrint, Beijing Double Helix 3D Biotechnology Co., Ltd., China) with P(LLA-TMC) / nHA 4D printing ink prepared at a mass ratio of nHA:82P(LLA-TMC) of 1:4. The printing parameters were set as follows: barrel temperature 45–50℃, nozzle size 0.4–0.8 mm, layer height 0.2–0.6 mm, nozzle movement speed 5–10 mm / s, line spacing 0.4–0.8 mm; extrusion speed 1–2 mm / s. 3 / s. After printing, dry in a vacuum drying oven for 72 hours to completely remove the solvent, obtaining the 4D printed scaffold.

[0070] P(LLA-TMC) / nHA.

[0071] (5) Preparation of nano-hydroxyapatite@polydopamine-pergillin (nHA@PDA-PGL) nanoparticles

[0072] Poly-dopamine (PDA) possesses good biocompatibility, biodegradability, and excellent near-infrared (NIR) photothermal conversion efficiency. Furthermore, its active groups, such as catechol and amino groups, give it good drug-loading capacity. Therefore, coating nHA with PDA can achieve both photothermal conversion and the loading of bioactive ingredients. A solution containing 2 mg / mL dopamine hydrochloride was prepared using deionized water. Tris-HCl (China) concentration of 10 mM The solution (from China) was adjusted to pH 8.5 using NaOH solution. 100 mg of nHA was weighed and dispersed in 20 mL of the above solution. The mixture was ultrasonically agitated for 10 minutes to ensure uniform dispersion of the nHA particles. The mixture was then stirred at 600 rpm for 12 hours at room temperature. After centrifugation at 12000 rpm for 15 minutes, the particles were washed three times with deionized water (each time at 12000 rpm for 15 minutes) to remove unreacted dopamine. The particles were then freeze-dried for 24 hours to obtain nHA@polydopamine particles (nHA@PDA). Then, a solution containing 100 mM of the osteopromoting small molecule drug pargyline (PGL) was prepared. PBS solution (Beijing Solarbio Science & Technology Co., Ltd., China) was prepared, and nHA@PDA nanoparticles (100 mg / mL) were added. The mixture was sonicated for 10 minutes, reacted in a rotary shaker at room temperature for 12 hours, and then centrifuged at 12,000 rpm for 15 minutes to remove the supernatant. The mixture was then lyophilized to obtain nHA@PDA-PGL nanoparticles and stored at -20°C.

[0073] (6) Preparation of methacrylated gelatin loaded with nHA@PDA-PGL

[0074] First, 60 mg of nHA@PDA and nHA@PDA-PGL nanoparticles were weighed separately and dispersed in 1 mL of PBS. The dispersions were ultrasonically agitated for 10 minutes to ensure uniform dispersion, preparing a 6 wt% nHA@PDA nanoparticle dispersion. Then, a 0.5% (w / v) photoinitiator (lithium phenyl (2,4,6-trimethylbenzoyl)phosphate, LAP) solution was prepared using PBS. Subsequently, methacrylated gelatin (GelMA, EFL-GM-90, Suzhou Yongqinquan Intelligent Equipment Co., Ltd., China) was dissolved in the LAP solution at a ratio of 20% (w / v). The solution was heated in a 65°C water bath for 30 minutes, with frequent shaking to ensure complete dissolution of the GelMA. Finally, the nHA@PDA and nHA@PDA-PGL nanoparticle solutions were mixed with the GelMA solution (1:1, v / v). Subsequently, the solutions were ultrasonically vibrated for 5 minutes to prepare GelMA / nHA@PDA and GelMA / nHA@PDA-PGL hydrogel precursor solutions.

[0075] (7) Preparation of composite 4D printed bone repair materials

[0076] First, the 4D printed scaffold P(LLA-TMC) / nHA was immersed in GelMA solution, GelMA / nHA@PDA precursor solution and GelMA / nHA@PDA-PGL precursor solution, respectively. Then, it was placed on a constant temperature horizontal shaker at 60 rpm and incubated for 10 minutes to ensure that GelMA, GelMA / nHA@PDA precursor solution and GelMA / nHA@PDA-PGL precursor solution were fully immersed in the scaffold pores. Then, the 4D printed scaffold P(LLA-TMC) / nHA was removed and irradiated with ultraviolet light (UV) at a wavelength of 405nm for 200 seconds to crosslink the GelMA / nHA@PDA precursor solution and the GelMA / nHA@PDA-PGL precursor solution, respectively, to prepare composite 4D printed bone repair materials P(LLA-TMC) / nHA+GelMA, P(LLA-TMC) / nHA+GelMA / nHA@PDA and P(LLA-TMC) / nHA+GelMA / nHA@PDA and P(LLA-TMC) / nHA+GelMA / nHA@PDA-PGL with surface coatings of GelMA, GelMA / nHA@PDA and GelMA / nHA@PDA-PGL.

[0077] Therefore, by combining 4D printing of P(LLA-TMC) / nHA with GelMA / nHA@PDA and GelMA / nHA@PDA-PGL, it is possible to first shrink the material and implant it into the body in a minimally invasive manner; then, under NIR irradiation, the scaffold is heated to restore the scaffold to its initial shape and match the bone defect.

[0078] Example 3: Testing of various materials prepared in Example 2

[0079] (1) Surface morphology of 4D printed bracket

[0080] The surface and cross-sectional morphology and internal pore structure of the printed scaffold were observed using a scanning electron microscope (SEM) (JSM-IT200, JEOL, Japan). The results showed that the filament diameter of the 4D printed scaffold P(LLA-TMC) / nHA was 541±32 μm, and the macropore diameter was 647±20 μm. Figure 5 a) High-magnification SEM images show uniformly distributed nHA particles within the P(LLA-TMC) scaffold. Figure 5 b).

[0081] (2) Mechanical performance testing of 4D printed bracket

[0082] The compressive modulus and compressive strength of pure P(LLA-TMC) and P(LLA-TMC) / nHA stents were tested using a universal testing machine. The P(LLA-TMC) / nHA stent was placed on the testing platform and, at room temperature, gradually compressed at a rate of 1.0 mm / min until 80% strain was reached. Force and displacement data were recorded to generate stress-strain curves. The compressive modulus of the stent was calculated by measuring the slope of the stress-strain curve in the linearly changing region (~10%), and the compressive strength was determined by the end stress in the linearly changing region. Results are as follows: Figure 6 The compressive moduli of the P(LLA-TMC) and P(LLA-TMC) / nHA stents were 16.9±4.9 MPa and 73.1±12.0 MPa, respectively. Figure 6 a) The compressive strengths were 1.4±0.4MPa and 6.1±1.3MPa, respectively. Figure 6 (b) This indicates that the addition of nHA significantly improves the mechanical properties of the scaffold. Studies have shown that the compressive modulus of human cancellous bone is 10–157 MPa, and its compressive strength is 1.5–3.8 MPa. Therefore, the P(LLA-TMC) / nHA scaffold possesses mechanical properties that match those of cancellous bone, sufficient to maintain its intact three-dimensional structure and provide the necessary mechanical support for bone regeneration.

[0083] (3) Characterization of nanoparticles

[0084] like Figure 7As shown, transmission electron microscopy (TEM) (JEM-1400PLUS, JEOL, Japan) revealed that the nanoparticles were needle-like, but a uniform coating was visible on the surfaces of nHA@PDA and nHA@PDA-PGL. The average diameters of the different nanoparticles were obtained by measuring the diameters of multiple particles. The average diameters of nHA, nHA@PDA, and nHA@PDA-PGL were 103.4 ± 8.3 nm, 129.9 ± 8.6 nm, and 130.2 ± 10.4 nm, respectively.

[0085] (4) Characterization of GelMA hydrogel

[0086] SEM showed that GelMA, GelMA / nHA@PDA, and GelMA / nHA@PDA-PGL all exhibited a uniform and highly interconnected network with a large number of large pores. Figure 8 a). By measuring the diameter of multiple pores, the average pore size of GelMA was 106.5 ± 22.5 μm, while the average pore sizes of GelMA / nHA@PDA and GelMA / nHA@PDA-PGL were slightly increased, at 110.7 ± 16.7 μm and 108.3 ± 14.2 μm, respectively. High-magnification scanning electron microscopy images showed that, compared to GelMA, nHA@PDA and nHA@PDA-PGL nanoparticles exhibited a more uniform distribution on the hydrogel surface. Figure 8 The arrow in the middle (a) indicates a significant increase in surface roughness.

[0087] (5) Characterization of composite 4D printed bone repair materials

[0088] Cross-sectional images of the scaffolds were observed using SEM. Typical porous hydrogel deposits were observed in the pores surrounding the scaffolds in the P(LLA-TMC) / nHA+GelMA, P(LLA-TMC) / nHA+GelMA / nHA@PDA, and P(LLA-TMC) / nHA+GelMA / nHA@PDA-PGL groups, with some hydrogel adhering to the upper scaffold. Figure 8 b) This demonstrates the successful integration of the 4D-printed scaffold with the bioactive hydrogel.

[0089] Water contact angle test results show ( Figure 8 c) Compared with the P(LLA-TMC) / nHA scaffold, the addition of gel significantly reduced the water contact angle (p<0.05) of the P(LLA-TMC) / nHA+GelMA, P(LLA-TMC) / nHA+GelMA / nHA@PDA and P(LLA-TMC) / nHA+GelMA / nHA@PDA-PGL groups, indicating that the hydrogel coating significantly improved the hydrophilicity of the scaffold.

[0090] Example 4: Temperature and photothermal response shape memory properties of various materials prepared in Example 2.

[0091] (1) Detect the shape fixation rate (R) of the scaffold material under temperature stimulation. f ) and shape recovery rate (R r ).

[0092] The shape memory properties of 4D printed scaffolds P(LLA-TMC) / nHA and P(LLA-TMC) / nHA+GelMA / nHA@PDA-PGL were compared. To demonstrate the shape memory properties of the materials, a flower-shaped scaffold (petal diameter 2cm) was first printed. Figure 9 a demonstrates the shape fixation and recovery process of two materials, showing that both materials can maintain their shape above T. trans At the specified temperature, it almost fully recovers to its original shape. Next, a rectangular support (length = 60 mm, width = 5 mm) was printed, and the R-values ​​of the two materials were tested. f and R r The initial angle is denoted as θ0 (180°). The support is folded into a U-shape in a 50°C water bath; the angle at this point is recorded as θ1. After fixing at 4°C for 5 minutes, the angle between the two ends of the U-shape is recorded as θ2. The material is then placed in a 50°C water bath, and the time required for shape recovery is recorded. The angle between the two petals after final shape recovery is measured and recorded as θ3. All angles are measured using a compass. Shape fixation rate (R...) f (shape fixed ratio) and shape recovery rate (R) r The shape recovery ratio (R) is calculated using the following formula: f =(180°-θ2) / (180°-θ1)×100%, R r = (θ3-θ2) / (180°-θ2)×100%. The results show that R... f The percentages were 97.5 ± 0.7% and 97.2 ± 0.6%, respectively; R r The percentages were 90.4 ± 0.9% and 89.3 ± 0.9%, respectively. Figure 9 b). Therefore, the prepared bone repair material has excellent shape memory properties and can almost completely recover to its initial shape at 42°C.

[0093] (2) Evaluation of photothermal responsiveness

[0094] The near-infrared photothermal and shape memory properties of the composite 4D printed bone repair material were characterized using an MDL-III-808 near-infrared laser (Changchun New Industrial Optoelectronic Technology Co., Ltd., Changchun, China) and an 875-1i infrared thermal camera (Testo, Model United Nations, Germany). Different rectangular scaffolds (length = 60 mm, width = 5 mm) were continuously irradiated with the near-infrared laser for 180 s (808 nm, 1.2 W / cm²). 2 The results show () Figure 10 a) The temperatures of the P(LLA-TMC) / nHA+GelMA / nHA@PDA and P(LLA-TMC) / nHA+GelMA / nHA@PDA-PGL groups gradually increased over time, eventually rising by 55℃ (to 80℃); the temperatures of P(LLA-TMC) / nHA and P(LLA-TMC) / nHA+GelMA showed a certain increase in the first 30 seconds (only an increase of 5℃), and then showed almost no significant change in the later stages.

[0095] Then, different irradiation powers (0.3 W / cm²) were applied. 2 0.6W / cm 2 and 1.2W / cm 2 Irradiation of the P(LLA-TMC) / nHA+GelMA / nHA@PDA-PGL scaffold was performed. Results showed that the scaffold temperature gradually increased with increasing irradiation power. Figure 10 b) indicates that it has excellent photothermal conversion effect.

[0096] Next, the P(LLA-TMC) / nHA+GelMA / nHA@PDA-PGL stent was subjected to a treatment at 1.2 W / cm². 2 Irradiation was performed for 6 minutes (3 minutes on, 3 minutes off), repeated 5 times. The results showed that the heating curve and final temperature remained almost unchanged, demonstrating the excellent photothermal stability of the scaffold. Figure 10 c). The above results indicate that P(LLA-TMC) / nHA@PDA exhibits good photothermal performance.

[0097] (3) Shape recovery rate under NIR light stimulation (R rNIR )

[0098] A rectangular support (length = 60mm, width = 5mm) was used, with the initial angle recorded as θ0 (180°). The support was folded into a U-shape in a 50°C water bath, and the angle at this point was recorded as θ1. After fixing at 4°C for 5 minutes, the angle between the two ends of the U-shape was recorded as θ2. Then, 808nm NIR light (1.2W / cm²) was used. 2 Irradiate the support for 180 seconds, and the angle between the two sides of the final U-shape is recorded as θ. NIR NIR light-triggered shape recovery rate (RrNIR The calculation formula is: R rNIR =(θ rNIR -θ2) / (180°-θ2)×100%. The results show that the R-values ​​of the 4D printed scaffolds P(LLA-TMC) / nHA, P(LLA-TMC) / nHA+GelMA, P(LLA-TMC) / nHA+GelMA / nHA@PDA, and P(LLA-TMC) / nHA+GelMA / nHA@PDA-PGL are... rNIR The percentages were 2.7% ± 0.5%, 3.0% ± 0.7%, 90.0% ± 1.4%, and 89.6% ± 1.7%, respectively. Figure 11 The above results indicate that the addition of nHA@PDA enables the stent to undergo photothermal conversion under NIR light irradiation, thereby raising the overall stent temperature above the shape transition temperature and achieving shape recovery of the P(LLA-TMC) / nHA+GelMA / nHA@PDA and P(LLA-TMC) / nHA+GelMA / nHA@PDA-PGL stents.

[0099] Example 5: In vitro osteoblast differentiation assay of human bone marrow mesenchymal stem cells using materials prepared in Example 2.

[0100] (1) Culture of human bone marrow mesenchymal stem cells

[0101] Human bone marrow mesenchymal stem cells (hBMSCs) were purchased from the ScienCell research laboratory (CA, USA). Cells were incubated in α-MEM (Gibco, Grand Island, USA) basal medium supplemented with 10% (v / v) fetal bovine serum (FBS) (ScienCell), 100 U / mL penicillin G, and 100 mg / mL streptomycin (Gibco, NY, USA) at 37°C with 5% CO2. The medium was changed every 2–3 days for routine expansion. Cells expanded to passages 5–7 were used for subsequent experiments.

[0102] (2) Expression of osteogenic related genes

[0103] Use a 12-well plate Transwell chamber ( In China, hBMSCs were inoculated into the lower chamber of a transwell chamber at a density of 1 × 10⁻⁶. 5 / well, cultured using basal medium (α-MEM). On day 3 post-inoculation, P(LLA-TMC) / nHA+GelMA, P(LLA-TMC) / nHA+GelMA / nHA@PDA, and P(LLA-TMC) / nHA+GelMA / nHA@PDA-PGL were placed in the upper chamber of a Transwell chamber for co-culture. After 7 days of co-culture, the effects of the prepared materials on the expression of osteogenic differentiation-related genes (Runx2, COLI, and OCN) in human bone marrow mesenchymal stem cells were detected by real-time quantitative PCR (RT-qPCR). The results showed that the P(LLA-TMC) / nHA+GelMA / nHA@PDA-PGL scaffold promoted the expression of almost all detected osteogenic-related genes, and the differences were statistically significant (*p<0.05 and **p<0.01). Figure 12 a) indicates that P(LLA-TMC) / nHA+GelMA / nHA@PDA-PGL enhances osteogenic differentiation of human bone marrow mesenchymal stem cells.

[0104] (3) Quantitative detection of alkaline phosphatase activity

[0105] Use a 12-well plate Transwell chamber ( In China, hBMSCs were inoculated into the lower chamber of a transwell chamber at a density of 1 × 10⁻⁶. 5 / well, cultured using basal medium (α-MEM). On day 3 post-inoculation, P(LLA-TMC) / nHA+GelMA, P(LLA-TMC) / nHA+GelMA / nHA@PDA, and P(LLA-TMC) / nHA+GelMA / nHA@PDA-PGL were placed in the upper chamber of a Transwell for co-culture. After 7 days of co-culture, alkaline phosphatase (ALP) activity was measured using a quantitative assay kit (Beyotime, China). The results showed that the average ALP activity of the P(LLA-TMC) / nHA+GelMA / nHA@PDA-PGL scaffold group was significantly higher than that of the other two groups, and the difference was statistically significant (*p<0.05 and **p<0.01). Figure 12 b) indicates that osteogenic differentiation of stem cells in the P(LLA-TMC) / nHA+GelMA / nHA@PDA-PGL group was enhanced.

[0106] Example 6: Experiment on the effect of materials prepared in Example 2 on promoting the repair of critical bone defects in the skull of rats

[0107] Eight-week-old male rats (Sprague Dawley, Vital Rivers, China) were selected and housed in a specific pathogen-free environment at a constant temperature of 25°C with a 12-hour day-night cycle. The experiment was divided into five groups: a control group, a P(LLA-TMC) / nHA+GelMA group, a P(LLA-TMC) / nHA+GelMA / nHA@PDA group, a P(LLA-TMC) / nHA+GelMA / nHA@PDA-PGL group, and a P(LLA-TMC) / nHA+GelMA / nHA@PDA-PGL+NIR group, with five animals in each group. Before the experiment, a scaffold was prepared as a 5.5 mm diameter disc, then the diameter was reduced to 4.5 mm at 50°C and maintained at -20°C before use. After general anesthesia of the rats, a 5 mm diameter critical bone defect in the skull was created, and then the different scaffolds were implanted.

[0108] The P(LLA-TMC) / nHA+GelMA / nHA@PDA-PGL+NIR group received 5 minutes of near-infrared light (808nm) after stent implantation to restore its shape; other groups received no such treatment. Eight weeks later, skull fragments were harvested for micro-CT analysis. Three-dimensional reconstructed images show (…). Figure 13 a) In the blank group, there was only a very small amount of bone regeneration in the skull defect; in the P(LLA-TMC) / nHA+GelMA and P(LLA-TMC) / nHA+GelMA / nHA@PDA groups, a small amount of bone regeneration was observed inside and around the scaffold, but there was a significant gap between the scaffold and the edge of the bone defect.

[0109] The P(LLA-TMC) / nHA+GelMA / nHA@PDA-PGL group showed more new bone formation, but gaps still existed between the scaffold and the bone defect margins; the P(LLA-TMC) / nHA+GelMA / nHA@PDA-PGL+NIR group showed abundant new bone formation, and the scaffold margins adhered closely to the bone defect margins. Quantitative analysis further confirmed the results observed in the above images. Figure 13 b). Temperatures exceeding 42°C can potentially lead to osteonecrosis, thereby impairing bone regeneration. However,

[0110] The P(LLA-TMC) / nHA+GelMA / nHA@PDA-PGL+NIR group showed a large amount of new bone formation, indicating that the photothermal process for achieving shape restoration did not lead to osteonecrosis.

[0111] As can be seen from the above examples:

[0112] (1) The composite 4D printed bone repair material P(LLA-TMC) / nHA+GelMA / nHA@PDA-PGL can be shrunk and implanted into complex bone defect areas. After implantation, it restores the original personalized design shape and can achieve precise matching and fit with the shape of the bone defect area. This overcomes the problems of traditional static 3D printed scaffolds requiring the trimming of surrounding healthy tissue and the non-fitting of the implant with the edge of the defect area.

[0113] (2) Composite 4D printed bone repair materials and their components, including P(LLA-TMC), nHA, PDA, PGL and GelMA hydrogel, all have good biocompatibility and biodegradability, which meet the requirements of bone repair.

[0114] (3) The ratio of LLA to TMC in shape memory polymers affects the transition temperature; the addition of nano-hydroxyapatite also affects the transition temperature of the polymer, and the addition of nano-hydroxyapatite makes the scaffold more similar to the composition of natural bone, increasing the osteogenic activity of the scaffold. Therefore, by adjusting the ratio of LLA to TMC and the proportion of nano-hydroxyapatite added, the transition temperature of the 4D printed scaffold can be achieved at 38℃~42℃, which will not cause damage to the body and is suitable for in vivo application; at the same time, the transition temperature of 38℃~42℃ is higher than the body temperature, so that the scaffold still has a certain strength at body temperature.

[0115] (4) Hydrogels can improve the hydrophobic properties of 4D printed shape memory polymer scaffolds and the large pores between fibers, which is beneficial to cell adhesion and proliferation. Secondly, the nanoparticles loaded with polydopamine coating in the hydrogel endow the bone repair scaffold with photothermal effect, enabling the scaffold to heat up under near-infrared light with a wavelength of 808nm and achieve shape restoration. At the same time, the nanoparticles loaded in the hydrogel contain osteogenic small molecule drugs and / or osteogenic bioactive components, which have the biological function of promoting bone defect repair.

[0116] Obviously, those skilled in the art can make various modifications and variations to this invention without departing from its spirit and scope. Therefore, if these modifications and variations fall within the scope of the claims and their equivalents, this invention is also intended to include these modifications and variations. The above embodiments or implementations are merely illustrative examples of this invention, and it can also be implemented in other specific ways or forms without departing from its gist or essential characteristics. Therefore, the described embodiments should be considered illustrative rather than limiting in any respect. The scope of this invention should be defined by the appended claims, and any changes equivalent to the intent and scope of the claims should also be included within the scope of this invention.

Claims

1. A shape memory bone repair material, characterized in that: The bone repair material is made by filling the pores of a shape memory polymer material for bone repair scaffold with hydrogel material, and the hydrogel material contains a photothermal conversion material.

2. The bone repair material according to claim 1, characterized in that: The shape memory polymer material is selected from one or more of polylactic acid, polylactic acid-trimethylene carbonate copolymer, polylactic acid-glycolic acid copolymer, polylactic acid-caprolactone copolymer, polylactic acid-polyethylene glycol copolymer, and polyurethane; The hydrogel material is selected from one or more of the following: methacrylated gelatin, methacrylated chitosan, methacrylated sodium alginate, methacrylated chondroitin sulfate, hydroxybutyl chitosan, sodium alginate, hyaluronic acid, chondroitin sulfate, polyethylene glycol diacrylate, and polyvinyl alcohol. The photothermal conversion material is selected from one or more of polydopamine, black phosphorus, gold nanorods, and carbon nanotubes.

3. The bone repair material according to claim 1, characterized in that: The shape memory polymer material is a polylactic acid-trimethylene carbonate copolymer, which is obtained by polymerizing trimethylene carbonate and L-lactic acid in a molar ratio of 1:(1.5-9.0).

4. The bone repair material according to claim 1, characterized in that: The bone repair material also contains hydroxyapatite in the shape memory polymer material and / or the hydrogel material.

5. The bone repair material according to claim 1, characterized in that: The bone repair material also contains bone-promoting drugs and / or bone-promoting bioactive ingredients added to the hydrogel material.

6. The bone repair material according to claim 1, characterized in that: The mass ratio of the shape memory polymer material to the hydrogel material is 30:(1-2), and the mass ratio of the hydrogel material to the photothermal conversion material is 1:(0.005-0.5).

7. A method for preparing a bone repair material according to any one of claims 1-6, characterized in that, The preparation method includes the following steps: (1) Prepare the shape memory polymer material described above; (2) Prepare a precursor solution of the hydrogel material containing the photothermal conversion material; (3) Immerse the shape memory polymer material in the precursor solution for a certain period of time; (4) Irradiate the soaked shape memory polymer material to form the hydrogel material from the precursor solution.

8. The preparation method according to claim 7, characterized in that: In step (1), the shape memory polymer material of the required shape and size is prepared using 3D printing technology.

9. The preparation method according to claim 8, characterized in that, The 3D printing parameters are as follows: barrel temperature 45–50℃, nozzle size 0.4–0.8 mm, layer height 0.2–0.6 mm, nozzle movement speed 5–10 mm / s, line spacing 0.4–0.8 mm; extrusion speed 1–2 mm / s. 3 / s.

10. The preparation method according to claim 7, characterized in that: In step (3), the soaking time is 10-60 minutes.

Images