Preparation method of a double-crosslinked 3D printing bone repair scaffold

Abstract

The application relates to the technical field of biomedical materials, and specifically discloses a preparation method of a double-crosslinked 3D-printed bone repair scaffold. The scaffold is made of a biphasic calcium phosphate matrix and a gelatin methacrylate hydrogel outer layer. The preparation method comprises the following steps: firstly, preparing biphasic calcium phosphate ink, and constructing a porous scaffold blank through 3D printing; then, carrying out double-crosslinking treatment of calcium chloride solution and glutaraldehyde solution in sequence to obtain a scaffold matrix with enhanced mechanical properties; then, preparing gelatin methacrylate hydrogel loaded with modified teriparatide, and compounding the gelatin methacrylate hydrogel with the scaffold matrix; and finally, obtaining the final product after blue light crosslinking and solidification. The bone repair scaffold has good biocompatibility and osteogenic activity, and can effectively promote bone defect repair. In addition, the preparation method is controllable and has good repeatability, the structure of the scaffold is accurately controlled through the 3D printing technology, and a new approach is provided for the preparation of personalized bone repair materials.

Description

Technical Field

[0001] This application relates to the field of materials science and engineering, and more specifically, it relates to a method for preparing a double cross-linked 3D printed bone repair scaffold. Background Technology

[0002] Bone defect repair is a common challenge in orthopedics and regenerative medicine. In clinical practice, bone defects caused by trauma, disease, etc., require bone repair materials to achieve functional reconstruction. An ideal bone repair material must meet multiple requirements simultaneously: it must provide reliable mechanical support to maintain the structural stability of the defect area and create space for new bone growth; it must also have good biocompatibility and osteoconductivity to support cell adhesion, proliferation, and osteogenic differentiation; it must also consider the ability to regulate the local immune microenvironment to avoid excessive inflammatory response caused by implantation that affects the bone regeneration process; and if it can load osteopromoting drugs and achieve long-term release, it can further improve the efficiency of bone repair.

[0003] Currently, existing bone repair materials have significant limitations: traditional biphasic calcium phosphate scaffolds have good mechanical properties but insufficient bioactivity and lack the ability to regulate the immune microenvironment; single hydrogel scaffolds have excellent biocompatibility but low mechanical strength, making it difficult to support large bone defects; composite scaffolds partially loaded with bone-promoting drugs, due to the lack of appropriate cross-linking methods and drug modification strategies, either have excessively rapid drug release rates leading to short-lived activity, or they cannot balance scaffold mechanical stability and degradation compatibility, making it difficult to synergistically achieve the comprehensive needs of mechanical support, long-term bone-promoting activity, and immune regulation. Summary of the Invention

[0004] To address the challenge of balancing mechanical support, long-term bone-promoting activity, and immune regulation in existing bone repair materials, this application provides a method for preparing a double-crosslinked 3D-printed bone repair scaffold.

[0005] A method for preparing a double cross-linked 3D printed bone repair scaffold includes the following steps:

[0006] S1. Preparation of biphasic calcium phosphate ink: Mix the mixed powder obtained by mixing α-tricalcium phosphate and β-tricalcium phosphate with sodium alginate to obtain a solid mixture; add hydroxypropyltrimethylammonium chloride chitosan aqueous solution and glycerol to the solid mixture and stir to form a homogeneous paste ink;

[0007] S2, 3D Printed Porous Scaffold: A porous scaffold blank is obtained by printing paste ink using direct ink writing technology;

[0008] S3, First crosslinking: The porous scaffold preform is immersed in calcium chloride solution for crosslinking and then dried;

[0009] S4, Second crosslinking: The scaffold that has undergone the first crosslinking is immersed in glutaraldehyde solution for crosslinking, washed and then freeze-dried under vacuum to obtain a double crosslinked biphasic calcium phosphate scaffold.

[0010] S5. Preparation of gelatin methacrylamide hydrogel: Dissolve gelatin methacrylamide in phosphate buffer, add photoinitiator, dissolve by water bath heating, shake and mix, centrifuge and filter to obtain gelatin methacrylamide hydrogel solution.

[0011] S6. Loading and modifying teriparatide: The teriparatide solution was mixed with the gelatin methacrylamide hydrogel solution, and azide-polyethylene glycol-active ester was added. After incubation in the dark, a modified hydrogel solution loaded with teriparatide was obtained.

[0012] S7. Composite scaffold preparation: The double cross-linked biphasic calcium phosphate scaffold is immersed in the modified hydrogel solution, cross-linked by blue light irradiation and freeze-dried to obtain the double cross-linked 3D printed bone repair scaffold.

[0013] By employing the above technical solution, a biphasic calcium phosphate mixed powder is formed by mixing α-tricalcium phosphate and β-tricalcium phosphate, and then combined with sodium alginate to construct a solid mixture. Hydroxypropyltrimethylammonium chloride chitosan aqueous solution and glycerol are added and stirred. Hydroxypropyltrimethylammonium chloride chitosan can improve ink viscosity and subsequent cross-linking bonding force, while glycerol can improve ink flowability, ensuring the uniformity and printability of the paste ink, thereby ensuring the smooth formation of subsequent 3D printing. The paste ink is printed using direct ink writing technology, enabling precise control of the pore distribution and external dimensions of the support structure. The process involves several steps: first, immersing the scaffold in a calcium chloride solution allows calcium ions to cross-link with sodium alginate, enhancing the initial structural stability of the scaffold and preventing post-printing deformation; second, immersing the cross-linked scaffold in a glutaraldehyde solution allows glutaraldehyde to covalently cross-link with the amino groups of hydroxypropyltrimethylammonium chloride chitosan, followed by vacuum freeze-drying. This process enhances the scaffold's mechanical strength and degradation resistance, eliminates residual toxicity, and ultimately achieves the desired structural integrity. The study demonstrates the effectiveness of constructing stable and biosafe bipolar cross-linked calcium phosphate scaffolds. By dissolving gelatin methacrylamide in phosphate buffer, adding a photoinitiator, and then subjecting the mixture to water bath heating, shaking, centrifugation, and filtration, the photoinitiator triggers subsequent photopolymerization reactions. Centrifugation removes air bubbles, and filtration achieves sterilization, thus preparing a sterile gelatin methacrylamide hydrogel solution with photocrosslinking capabilities, thereby providing a suitable carrier for drug loading. Further methods include mixing teriparatide solution with the gelatin methacrylamide hydrogel solution, adding azide-polyethylene glycol-active ester, and incubating in the dark. The amino coupling of azido-polyethylene glycol-active ester with teriparatide achieves stable loading and modification of teriparatide, thereby endowing the scaffold with bone regeneration activity. By immersing the biphasic calcium phosphate scaffold in a modified hydrogel solution, followed by blue light irradiation to induce photopolymerization and crosslinking of gelatin methacrylamide, and freeze-drying for shaping, the blue light can precisely fix the hydrogel and teriparatide, effectively binding the teriparatide-loaded hydrogel and the biphasic calcium phosphate scaffold together, thus obtaining a bone repair scaffold that combines the stability of the biphasic crosslinked structure with the long-lasting effect of the drug.

[0014] Preferably, in step S1, the mixed powder is prepared by mixing α-tricalcium phosphate and β-tricalcium phosphate in a mass ratio of 2:8 to 4:6, and then sieving the mixture through a sieve with a mesh size of 100-200; the mass ratio of the mixed powder to sodium alginate is 95:5 to 99:1.

[0015] By employing the above technical solution, mixing α-tricalcium phosphate (TFP) and β-tricalcium phosphate (β-tricalcium phosphate) at a mass ratio of 2:8 to 4:6 allows for the proportion of these two biphasic calcium phosphate components to be adjusted to match the degradation rate during bone regeneration. α-tricalcium phosphate degrades relatively quickly, while β-tricalcium phosphate degrades more slowly. This ratio range allows the subsequent scaffold to provide continuous mechanical support while gradually releasing calcium and phosphate ions to promote new bone formation, thus balancing the scaffold's degradability and osteoconductivity. This provides a long-lasting bone repair microenvironment for the bone defect area. Sieving the mixed calcium phosphate powder through a 100-200 mesh sieve removes coarse particles and agglomerates, ensuring uniform particle size and preventing contamination. This method avoids uneven particle distribution that could lead to lumps in the subsequent paste ink and affect printing smoothness, thereby improving ink uniformity and the surface finish of the printhead. By controlling the mass ratio of the mixed powder to sodium alginate to 95:5 to 99:1, sodium alginate, as a binder, is used. This ratio allows it to fully coat the calcium phosphate powder, enhancing the adhesion of the solid mixture and preventing the printhead blank from becoming easily broken during subsequent printing. At the same time, it avoids the mixture becoming too viscous due to an excessive proportion of sodium alginate, which could affect the addition of subsequent liquid components and ink flow. This balances the adhesion of the solid mixture with the subsequent ink molding properties, thus ensuring the printability of the paste ink and the initial structural stability of the printhead blank.

[0016] Preferably, in step S1, the mass concentration of the hydroxypropyltrimethylammonium chloride chitosan aqueous solution is 2-4 wt.%, and the amount added is 0.2-0.4 mL per gram of solid mixture; the amount of glycerol added is 0.1-0.3 mL per gram of solid mixture.

[0017] By adopting the above technical solution, and controlling the mass concentration of the hydroxypropyltrimethylammonium chloride chitosan aqueous solution to 2-4 wt.%, this concentration range allows the hydroxypropyltrimethylammonium chloride chitosan to fully dissolve in the aqueous solution without becoming excessively viscous. This avoids both insufficient viscosity for the ink due to low concentration, which could lead to deformation of the printing support, and excessive viscosity that makes the aqueous solution too thick and difficult to mix evenly with the solid mixture. This precise control of the viscosity and dissolution state of the aqueous solution ensures the overall uniformity of the subsequent ink. Furthermore, by limiting the addition amount to 0.2-0.4 mL per gram of solid mixture, this dosage allows the hydroxypropyltrimethylammonium chloride chitosan aqueous solution to fully coat the mixed powder and sodium alginate in the solid mixture. This method balances the ratio of solid to liquid components, ensuring the ink maintains both viscosity and fluidity. By controlling the amount of glycerol added to 0.1-0.3 mL per gram of solid mixture, glycerol fills the spaces between solid particles and hydroxypropyltrimethylammonium chloride chitosan molecules. This avoids insufficient addition leading to poor ink extrusion during printing, while excessive addition results in overly thin and soft ink, making it difficult for the printed blank to set. This optimizes the ink's rheological properties, ensuring smooth ink flow through the printer nozzles and achieving continuous and stable printing.

[0018] Preferably, in step S2, the printing parameters of the direct ink writing technology include: printing speed 3-5 mm / s, printing pressure 500-550 kPa, layer height 0.2-0.5 mm, nozzle inner diameter 0.3-0.5 mm, line width 0.3-0.5 mm, and fill density 30-50%.

[0019] By adopting the above technical solutions, controlling the printing speed of the direct ink writing technology to 3-5 mm / s, this speed range avoids both excessive speed causing ink to move before it has set, resulting in loose bonding between support layers or broken lines, and excessive speed causing ink to accumulate at the nozzle, affecting printing continuity. This balances printing efficiency and the quality of bonding between support layers, thereby ensuring the overall structural integrity of the support. Controlling the printing pressure to 500-550 kPa precisely matches the flowability of the paste ink, avoiding both insufficient pressure leading to poor ink extrusion and ink breaks, and excessive pressure causing excessive ink extrusion and uneven line thickness. This stabilizes the ink extrusion state, ensuring uniformity of support lines and printing accuracy. Controlling the layer height to 0.2-0.5 mm ensures sufficient contact and tight bonding between adjacent printed layers, preventing layer separation due to excessive layer height, and avoiding excessive layer adhesion due to insufficient layer height. This optimizes the bonding strength between support layers and printing efficiency, thereby maintaining the integrity of the support structure. The system achieves the following effects on structural stability: By controlling the nozzle inner diameter to 0.3-0.5mm, which matches the particle uniformity and flowability of the paste ink, it avoids both excessively small inner diameters that could clog the nozzle and excessively large inner diameters that could compress the effective pores of the scaffold. This ensures smooth ink extrusion while matching the scaffold's pore requirements, thus reserving suitable space for cell growth. By controlling the printing linewidth to 0.3-0.5mm, which works in conjunction with the nozzle inner diameter and printing pressure, it ensures continuous and unbroken lines and controls the scaffold pore distribution through line spacing, balancing line integrity and porosity. This provides a reasonable channel for cell adhesion, proliferation, and nutrient transport. By controlling the filling density to 30-50%, it avoids both insufficient density (leading to inadequate scaffold mechanical strength and inability to support bone defects) and excessive density (leading to insufficient scaffold pores and hindering cell infiltration and new bone ingrowth). This balances the scaffold's mechanical support capacity and biocompatibility, thus meeting the dual needs of structural support and tissue regeneration during bone repair.

[0020] Preferably, in step S3, the mass concentration of the calcium chloride solution is 4-6%, the crosslinking time is 1-3 min, and the drying is performed at room temperature for 20-28 h.

[0021] By adopting the above technical solution and controlling the mass concentration of calcium chloride solution to 4-6%, calcium ions at this concentration can fully combine with the carboxyl groups of sodium alginate in the scaffold blank to form stable ionic bonds. This avoids both insufficient ionic cross-linking (due to too low a concentration, resulting in a loose and easily deformed scaffold blank structure) and excessive cross-linking (due to too high a concentration, resulting in increased brittleness and easy breakage of the scaffold). This achieves precise control of the degree of ionic cross-linking, thereby initially enhancing the structural stability of the scaffold blank and preventing morphological damage during subsequent processing. Furthermore, by limiting the cross-linking time to 1-3 minutes, this duration allows calcium ions to fully react with sodium alginate to form a cross-linked network. This avoids incomplete cross-linking (due to too short a time, resulting in a loose scaffold) and also prevents… Excessive drying time causes the stent to absorb too much water and swell, altering its pre-set shape. This process controls the crosslinking reaction process and the morphological precision of the stent, thereby ensuring the crosslinking effect while maintaining the dimensional accuracy of the stent after printing. By using room temperature drying for 20-28 hours, the room temperature environment avoids the rapid evaporation of moisture inside the stent caused by high-temperature drying, which could lead to cracking or pores. The 20-28 hour duration allows the moisture to evaporate slowly and evenly, avoiding both insufficient drying that could leave moisture residue affecting subsequent glutaraldehyde crosslinking and excessive drying that could increase the risk of stent contamination. This process gently and thoroughly removes moisture from the stent, ensuring the structural integrity of the stent after drying and providing a stable drying substrate for the second crosslinking step.

[0022] Preferably, in step S4, the mass concentration of the glutaraldehyde solution is 0.5-1.5%, the crosslinking temperature is 35-39°C, and the crosslinking time is 22-26h; the vacuum freeze-drying temperature is -50°C to -80°C, and the drying time is 20-28h.

[0023] By employing the above technical solution and controlling the mass concentration of glutaraldehyde solution to 0.5-1.5%, glutaraldehyde can fully undergo a covalent cross-linking reaction with the amino groups of hydroxypropyltrimethylammonium chloride chitosan in the scaffold. This avoids both insufficient cross-linking due to excessively low concentration, resulting in inadequate mechanical strength and degradation resistance of the scaffold, and excessively high concentration, which would cause residual glutaraldehyde to exceed the safe range and increase the risk of cytotoxicity. This achieves precise control over the degree of covalent cross-linking and biosafety, thereby further enhancing the structural stability of the scaffold and prolonging its degradation cycle in vivo to match the pace of bone regeneration. The cross-linking temperature is 35-39°C, which is close to the physiological environment temperature. This temperature provides suitable kinetic conditions for the cross-linking reaction between glutaraldehyde and amino groups, avoiding slow or incomplete reaction rates due to excessively low temperatures, and preventing damage to the crystal structure of calcium phosphate in the scaffold or scaffold deformation due to excessively high temperatures. This optimizes the cross-linking reaction efficiency and protects the stability of the scaffold components, thereby maintaining the performance of the core scaffold components while ensuring the cross-linking effect. By controlling the cross-linking time to 22-26 hours, this duration ensures sufficient amino groups in the glutaraldehyde and hydroxypropyltrimethylammonium chloride chitosan. The reaction forms a complete cross-linked network, avoiding both incomplete cross-linking (leading to scaffold degradation or deformation due to insufficient time) and excessive cross-linking (leading to increased brittleness due to excessive time). This ensures a balance between the thoroughness of the cross-linking reaction and the mechanical properties of the scaffold, resulting in long-term stable mechanical support. Vacuum freeze-drying at -50°C to -80°C rapidly freezes the water in the scaffold into ice crystals. Under vacuum conditions, these ice crystals are directly sublimated and removed, preventing pore collapse or structural shrinkage caused by water evaporation during conventional drying. This also protects the integrity of the scaffold's porous structure, preserving the pre-defined pore morphology and maintaining a porous structure conducive to cell infiltration and nutrient transport. Limiting the vacuum freeze-drying time to 20-28 hours ensures sufficient sublimation of the frozen water, preventing residual water from affecting subsequent binding with the hydrogel, and avoiding excessive drying which increases energy consumption or causes scaffold component deterioration. This ensures thorough removal of water from the scaffold and maintains stable performance, providing a dry and clean scaffold substrate for subsequent hydrogel loading and modification.

[0024] Preferably, in step S5, the mass concentration of the gelatin methacrylamide hydrogel solution is 8-12%; the photoinitiator is lithium phenyl-2,4,6-trimethylbenzoyl phosphate with a mass concentration of 0.2-0.3%.

[0025] By adopting the above technical solution and controlling the mass concentration of the gelatin methacrylamide hydrogel solution to 8-12%, the gelatin methacrylamide can be fully dissolved to form a solution with suitable viscosity. This avoids both insufficient concentration, which would lead to insufficient hydrogel strength after subsequent blue light crosslinking and easy detachment from the double-crosslinked BCP scaffold, making it impossible to stably load teriparatide, and excessive concentration, which would make the solution too viscous, difficult to uniformly wet the scaffold pores, or cause uneven drug dispersion. This balances the hydrogel's solubility, viscosity, and structural stability after crosslinking, thereby providing a stable loading carrier for teriparatide and ensuring tight binding with the double-crosslinked BCP scaffold. By selecting lithium phenyl-2,4,6-trimethylbenzoyl phosphate as a photoinitiator and controlling its mass concentration at 0.2-0.3%, this initiator can efficiently trigger the double bond polymerization reaction of gelatin methacrylamide under blue light. The concentration of 0.2-0.3% can ensure the initiation efficiency, avoid insufficient photopolymerization due to too low a concentration, and prevent the hydrogel from forming a complete network. It can also reduce the residual amount, and avoid the cytotoxicity of unreacted initiator due to too high a concentration. This plays a role in precisely controlling the photocrosslinking process and ensuring the biosafety of the hydrogel. In this way, it can ensure the rapid molding of gelatin methacrylamide and that the molded hydrogel meets the medical safety requirements.

[0026] Preferably, in step S5, the water bath heating temperature is 50-60°C, the centrifugation speed is 2000-4000 rpm, and the time is 1-3 min; the filtration is performed using a filter membrane with a pore size of 0.22 μm.

[0027] By adopting the above technical solution and controlling the water bath heating temperature to 50-60°C, a suitable dissolution environment is provided for gelatin methacrylamide. This avoids both excessively low temperatures that lead to incomplete dissolution and residual solid particles in the solution, affecting the uniformity of subsequent mixing with teriparatide, and excessively high temperatures that damage the molecular structure of gelatin methacrylamide and cause it to lose its subsequent photocrosslinking ability. This ensures the complete dissolution and stable molecular structure of gelatin methacrylamide, thereby achieving a uniform hydrogel base solution with photocrosslinking activity. Furthermore, by limiting the centrifugation speed to 2000-4000 rpm and the time to 1-3 min, this parameter combination effectively separates air bubbles in the solution. Too low a speed or too long a time will result in the loss of the desired effect. Short rotation speeds cannot completely remove air bubbles, which can lead to pores or uneven structure when the hydrogel is loaded onto the subsequent stent. Excessive rotation speed or time may cause the solution components to separate. This method is designed to precisely remove air bubbles from the solution, thereby ensuring the homogeneity of the hydrogel solution and preventing air bubbles from affecting the subsequent drug loading and cross-linking effect. By using a filter membrane with a pore size of 0.22μm, microorganisms and tiny impurities in the solution can be effectively trapped, meeting the requirements for medical sterility. This avoids the infection caused by microbial residues during subsequent stent implantation and also prevents impurities from affecting the binding stability of the hydrogel and teriparatide. This method has the effect of sterilization and removal of tiny impurities, thereby obtaining a sterile and clean gelatin methacrylamide hydrogel solution that meets the safety standards for clinical application.

[0028] Preferably, in step S6, the concentration of the teriparatide solution is 80-120 μg / mL; the amount of azide-polyethylene glycol-active ester added is 0.1-0.3 mg per milligram of teriparatide; and the light-protected incubation conditions are incubation at 4°C for 20-28 h.

[0029] By employing the above technical solution and controlling the teriparatide solution concentration to 80-120 μg / mL, this concentration range provides an appropriate dose of active drug for hydrogel loading. This avoids both insufficient concentration leading to inadequate subsequent bone regeneration activity in the scaffold and failure to effectively activate osteogenic pathways, and excessive concentration causing drug waste or affecting cell function due to localized high concentrations. This ensures an effective loading of teriparatide, thereby providing the scaffold with sufficient bone-promoting active factors. Furthermore, by limiting the addition of azide-polyethylene glycol-active ester to 0.1-0.3 mg per mg of teriparatide, this dosage ensures sufficient coupling with the amino groups of teriparatide, avoiding the need for excessive addition. Insufficient addition leads to incomplete coupling and rapid loss of teriparatide, while excessive addition results in residues that affect the biocompatibility of the hydrogel, thus stabilizing and anchoring teriparatide and achieving a long-lasting, slow release effect. By controlling the light-protected incubation conditions to 4°C for 20-28 hours, the low temperature of 4°C prevents teriparatide degradation, the light-protected environment avoids interference with the coupling reaction, and the 20-28 hour duration ensures complete coupling. This avoids both temperature or light-induced inactivation of teriparatide and insufficient incubation leading to incomplete reaction, thus ensuring the activity and integrity of teriparatide and achieving a stable hydrogel solution loaded with modified teriparatide.

[0030] Preferably, in step S7, the soaking time is 20-28 hours; the wavelength of the blue light irradiation is 400-410 nm, and the irradiation time is 15-25 seconds; the freeze-drying temperature is -50°C to -80°C, and the time is 20-28 hours.

[0031] By employing the above technical solution and controlling the soaking time to 20-28 hours, the modified hydrogel solution can fully penetrate into the pores of the bi-crosslinked biphasic calcium phosphate scaffold. This avoids the situation where the hydrogel only adheres to the scaffold surface and fails to fill the internal pores due to insufficient soaking time, resulting in uneven drug distribution. Conversely, excessive soaking time increases the risk of solution contamination. This ensures that the hydrogel and drug are evenly loaded into the scaffold pores, thereby enabling the entire scaffold to release osteopromoting active substances. By limiting the blue light irradiation wavelength to 400-410 nm, this wavelength can precisely match the absorption spectrum of the photoinitiator phenyl-2,4,6-trimethylbenzoyl phosphate, efficiently triggering the photopolymerization reaction of gelatin methacrylamide. At the same time, the irradiation time of 15-25 seconds ensures that the hydrogel is fully crosslinked and shaped, avoiding the situation where the wavelength mismatch leads to low crosslinking efficiency and hydrogel defects. The method avoids the inability to stably shape the scaffold and prevents excessive cross-linking and brittleness of the hydrogel due to prolonged irradiation time, which could affect biocompatibility. It precisely controls the cross-linking process of the hydrogel, thereby firmly fixing the teriparatide-loaded hydrogel within the scaffold pores and preventing drug loss. By using a freeze-drying temperature of -50°C to -80°C and a drying time of 20-28 hours, the low temperature can quickly freeze the water in the scaffold to form ice crystals. The ice crystals are then directly sublimated and removed under vacuum conditions. This avoids the hydrogel shrinkage that can block the scaffold pores and affect cell infiltration caused by conventional drying. The 20-28 hour drying time also removes residual moisture, preventing drug degradation or scaffold mold growth caused by residual moisture. This preserves the integrity of the scaffold's porous structure and ensures that the scaffold is dry and clean, thus achieving the effect of obtaining a structurally stable, drug-loaded, double-crosslinked 3D-printed bone repair scaffold.

[0032] In summary, this application has the following beneficial effects:

[0033] 1. This invention significantly improves the mechanical properties, structural stability, and biofunctionality of bone repair scaffolds by combining a dual cross-linking strategy with active drug modification technology. The biphasic calcium phosphate scaffold undergoes sequential cross-linking with calcium chloride ions and covalent cross-linking with glutaraldehyde to form a stable three-dimensional network structure, giving it excellent mechanical strength and degradation resistance. Simultaneously, teriparatide is covalently modified with azide polyethylene glycol active ester and loaded onto a photocrosslinked gelatin methacrylamide hydrogel, achieving controlled sustained drug release and long-term maintenance of bioactivity. This composite structure provides mechanical support while continuously promoting osteogenic differentiation, making it suitable for repairing large bone defects.

[0034] 2. This invention constructs a bone repair scaffold with a high-precision porous structure by optimizing the three-dimensional printing process and material composition. This facilitates cell migration, nutrient delivery, and the ingrowth of new bone tissue. The prepared scaffold has uniform pore size and good connectivity, which not only has good osteoconductivity but also provides an ideal microenvironment for blood vessel ingrowth and cell distribution.

[0035] 3. The present invention employs mild conditions such as low-temperature cross-linking, light-protected incubation, and low-temperature freeze-drying throughout the entire preparation process, which effectively maintains the stability of biomaterials and active factors. This method can ensure the activity of teriparatide and its effective binding with the scaffold material while maintaining the integrity of the macroscopic structure and microscopic morphology of the scaffold, thereby ensuring that it can perform the expected biological function after implantation. Attached Figure Description

[0036] Figure 1 This is a flowchart of a method for preparing a double cross-linked 3D printed bone repair scaffold provided in this application. Detailed Implementation

[0037] The present application will be further described in detail below with reference to embodiments and comparative examples.

[0038] Technical concept:

[0039] Existing bone repair materials technologies suffer from core problems: the difficulty in simultaneously achieving synergistic unity between mechanical support stability, degradation rate matching the pace of bone regeneration, long-term controllable release of bone-promoting drugs, and regulation of the immune microenvironment. The reasons for this include: traditional biphasic calcium phosphate scaffolds often employ a single cross-linking method, which can only initially maintain the structure and cannot simultaneously ensure long-term mechanical strength and the degradation rate required for bone regeneration; when hydrogels are combined with inorganic scaffolds, the simple bonding method often leads to weak interfacial adhesion and easy detachment in vivo; bone-promoting drug loading is mostly physical mixing, lacking specific modification, resulting in initial burst release of drugs followed by insufficient activity; and most materials do not consider the impact of post-implantation immune responses on bone regeneration, making it difficult to construct an immune environment conducive to bone regeneration through methods such as regulating macrophage polarization. These problems collectively limit the efficiency of bone repair.

[0040] This technical solution addresses the aforementioned issues through a multi-dimensional approach: First, a dual cross-linking strategy using calcium chloride ion cross-linking and glutaraldehyde covalent cross-linking is employed to treat the biphasic calcium phosphate scaffold. This method allows for precise control of the scaffold's mechanical strength and degradation rate, ensuring it matches the bone regeneration process. Second, 3D printing technology is used to control the scaffold's pore structure. Gelatin methacrylamide hydrogel is used as a drug carrier, and teriparatide is modified with azide-polyethylene glycol-active ester to achieve stable coupling between the drug and the hydrogel. A specific soaking time ensures uniform hydrogel penetration into the scaffold pores. Finally, blue light cross-linking firmly binds the hydrogel to the scaffold, and freeze-drying preserves the porous structure. This solves the drug release problem and ensures the interfacial stability between the hydrogel and the scaffold. This composite system also induces M2 polarization in macrophages, creating a suitable immune microenvironment for bone regeneration, ultimately achieving a multifunctional and synergistic bone repair effect.

[0041] The following are the main raw materials and instruments used in the preparation examples, embodiments, and comparative examples, and their sources and specifications are as follows; unless otherwise specified, all reagents are commercially available analytical grade or higher products:

[0042] I. Raw materials and pharmaceuticals

[0043] 1. β-Tricalcium Phosphate (β-TCP): Purchased from Shanghai Bio-materials Co., Ltd.;

[0044] 2. α-Tricalcium phosphate (α-TCP): Purchased from Shanghai Aladdin Biochemical Technology Co., Ltd.;

[0045] 3. Hydroxypropyltrimethylammonium chloride chitosan (HACC): Purchased from Shanghai Aladdin Biochemical Technology Co., Ltd.;

[0046] 4. Sodium alginate: purchased from Wuhan Servicebio Biotechnology Co., Ltd.

[0047] 5. Calcium chloride: purchased from Wuhan Servicebio Biotechnology Co., Ltd.

[0048] 6. Glycerin: Purchased from Wuhan Servicebio Biotechnology Co., Ltd.

[0049] 7. Glutaraldehyde: Purchased from Wuhan Servicebio Biotechnology Co., Ltd.

[0050] 8. Methacrylonitrile gelatin (GelMA-60): Purchased from Suzhou Yongqinquan Intelligent Equipment Co., Ltd.;

[0051] 9. Azide-polyethylene glycol-active ester (AC-PEG-NHS): Purchased from Suzhou Yongqinquan Intelligent Equipment Co., Ltd.;

[0052] 10. Lithium hyaluronic acid-phenyl(2,4,6-trimethylbenzoyl)phosphate (hyaluronic acid-LAP): purchased from Suzhou Yongqinquan Intelligent Equipment Co., Ltd.

[0053] 11. Teriparatide: Purchased from Ombio Biotechnology (Shanghai) Co., Ltd.;

[0054] 12. Rat bone marrow mesenchymal stem cells: purchased from Wuhan Pricella Biotechnology Co., Ltd.

[0055] 13. Mouse RAW264.7 macrophages: purchased from Wuhan Pricella Biotechnology Co., Ltd.

[0056] II. Instruments

[0057] 1. 4-channel Ink Direct Write (DIW) 3D printer: Selected from Shenzhen Yuanyi Smart Medicine Technology Co., Ltd.;

[0058] 2. Scanning electron microscope (SEM, S4800): Selected from Hitachi, Japan;

[0059] 3. Scanning electron microscope (SEM, FEI): Selected from FEI Corporation, USA;

[0060] 4. X-ray diffractometer (XRD, Bruker): Selected from Bruker GmbH, Germany;

[0061] 5. X-ray diffractometer (XRDPANalyticalX'pert3): selected from Panaco GmbH, Netherlands;

[0062] 6. Fourier Transform Infrared Spectrometer (FTIR, Nicolet): Selected from Thermo Fisher Scientific, USA;

[0063] 7. Universal testing machine (Instron): Selected from Instron Corporation, USA;

[0064] 8. Universal Testing Machine (MTSE43): Selected from METS Industrial Systems (China);

[0065] 9. Mercury porosimeter (Autopore V9610): Selected from Mack Instruments, USA;

[0066] 10. Real-time quantitative PCR instrument (LightCycler 480): selected from Roche, Switzerland;

[0067] 11. Micro-CT (VNC-102): Selected from Beijing Pingsheng Medical Technology Co., Ltd.

[0068] Example 1

[0069] This application provides a method for preparing a double cross-linked 3D printed bone repair scaffold, specifically including the following steps:

[0070] S1, Preparation of biphase calcium phosphate ink

[0071] Preparation of mixed powder: Take α-tricalcium phosphate (α-TCP) and β-tricalcium phosphate (β-TCP) and mix them at a mass ratio of 3:7. Pass the mixed powder through a 100-mesh sieve to remove coarse particles to ensure ink uniformity.

[0072] Preparation of solid mixture: Weigh 3g of the above sieved mixed powder, add 0.1g of sodium alginate, and manually grind until uniformly mixed to form a solid mixture with a total mass of 3.1g, wherein the mass ratio of the mixed powder to sodium alginate is 30:1.

[0073] Liquid component addition and mixing: Add 0.9 mL of 3 wt.% hydroxypropyltrimethylammonium chloride chitosan aqueous solution to the solid mixture, and then add 0.5 mL of glycerol. Place the mixture on a magnetic stirrer and stir at 300 rpm for 30 min until a uniform paste-like ink with no particles and suitable flowability is formed, and set aside for later use.

[0074] S2, 3D printed porous bracket

[0075] Equipment and model preparation: A 4-channel ink direct-write 3D printer was used as the printing device; SolidWorks 2022 software was used to build the 3D model of the support, exported as .stl format, and imported into Cura 4.8.25 software for slicing.

[0076] Printing parameter settings: Slicing parameters are set as follows: printing speed 4mm / s, printing pressure 520KPa, initial layer height 0.2mm, subsequent layer height 0.41mm, nozzle inner diameter 0.41mm, printing line width 0.41mm, fill density 40%, and nozzle temperature is maintained at room temperature to avoid premature ink curing.

[0077] Scaffold Printing: The paste-like ink prepared in step S1 was loaded into a 10cc 3D printing cartridge and installed into the printer nozzle; the .gcode format file generated after slicing was imported into the printer control system, and layer-by-layer deposition printing was performed. According to experimental requirements, two sizes of scaffolds were printed: 10×10×6mm (length×width×height) for in vitro experiments and 6×2mm (diameter×height) for in vivo experiments. After printing, a porous scaffold blank was obtained.

[0078] S3, First crosslinking

[0079] The porous scaffold preform obtained in step S2 was completely immersed in a 5% calcium chloride solution and allowed to stand at room temperature for 2 minutes for crosslinking, allowing sodium alginate to form ionic bonds with calcium ions and improving the initial structural stability of the scaffold preform. After crosslinking, the scaffold preform was removed, excess solution was blotted off with filter paper, and it was dried in a well-ventilated environment at room temperature for 24 hours to obtain a pre-crosslinked BCP scaffold.

[0080] S4. Second crosslinking: Glutaraldehyde crosslinking: Immerse the scaffold dried in step S3 in a 1% glutaraldehyde solution and place it in a 37°C constant temperature water bath for 24 hours for crosslinking. Covalent bonds are formed through the reaction of glutaraldehyde with the amino groups of HACC, which further enhances the mechanical properties and degradation resistance of the scaffold.

[0081] Washing and freeze-drying: After cross-linking, the scaffold was removed and washed repeatedly with deionized water three times for 10 minutes each time to remove residual glutaraldehyde solution and avoid cytotoxicity. The washed scaffold was then placed in a vacuum freeze dryer at a temperature of -60℃ and a vacuum of 10 Pa for 24 hours to obtain a double-crosslinked biphasic calcium phosphate scaffold.

[0082] S5. Preparation of gelatin methacrylamide hydrogel

[0083] Preparation of photoinitiator solution: Place lithium phenyl (2,4,6-trimethylbenzoyl)phosphate in a brown bottle, add an appropriate amount of phosphate buffer, heat in a 45°C water bath for 15 min, shaking 3 times during the process to ensure complete dissolution, and prepare a 0.25 wt.% LAP standard solution.

[0084] Preparation of GelMA hydrogel solution: Weigh methacryloyl gelatin powder and add it to the above LAP standard solution to prepare a 10 wt.% GelMA mixed solution. Place the mixed solution in a 55℃ constant temperature water bath and shake it 5 times for 30 seconds each time to promote the dissolution of GelMA. After complete dissolution, centrifuge it at 3000 rpm for 2 minutes to remove air bubbles from the solution.

[0085] Sterilization: The centrifuged GelMA solution was filtered and sterilized using a 0.22 μm pore size filter membrane to obtain a sterile GelMA hydrogel solution for later use.

[0086] S6, loaded with teriparatide and modified

[0087] PTH solution preparation: Teriparatide powder was dissolved in ultrapure water to prepare a 100 μg / mL PTH solution.

[0088] Preparation of modified hydrogel: The above PTH solution and the GelMA hydrogel solution prepared in step S5 were mixed at a volume ratio of 1:9. After stirring evenly, azide-polyethylene glycol-active ester was added at a ratio of 0.2 mg AC-PEG-NHS per mg PTH. Stirring was continued for 10 min to ensure thorough mixing.

[0089] Incubation in the dark: The mixed solution was sealed with sealing film and incubated in a refrigerator at 4°C for 24 hours in the dark to allow the amino groups of AC-PEG-NHS and PTH to undergo a coupling reaction, thereby achieving stable loading and modification of PTH and finally obtaining a modified hydrogel solution loaded with teriparatide.

[0090] S7. Final fabrication of the composite scaffold

[0091] Scaffold immersion loading: The double cross-linked BCP scaffold obtained in step S4 is completely immersed in the modified hydrogel solution prepared in step S6, and placed in a 4°C refrigerator for 24 hours to ensure that the hydrogel solution fully penetrates into the pores of the BCP scaffold.

[0092] Blue light crosslinking and curing: Remove the soaked scaffold, drain the excess solution from the surface, and irradiate it with a 405nm blue light irradiator for 20s to cause the GelMA to undergo a photopolymerization reaction, forming a stable hydrogel network, and fixing PTH in the pores of the scaffold.

[0093] Freeze-drying and shaping: The blue light cross-linked composite scaffold is placed in a vacuum freeze dryer, the temperature is set to -60℃, and it is dried for 24 hours to remove the moisture in the scaffold, thus obtaining the final double cross-linked 3D printed bone repair scaffold.

[0094] Example 2

[0095] This application provides a method for preparing a double cross-linked 3D printed bone repair scaffold, specifically including the following steps:

[0096] S1, Preparation of biphase calcium phosphate ink

[0097] Preparation of mixed powder: Take α-tricalcium phosphate (α-TCP) and β-tricalcium phosphate (β-TCP) and mix them at a mass ratio of 2:8. Pass the mixed powder through a 100-mesh sieve to remove coarse particles to ensure ink uniformity.

[0098] Preparation of solid mixture: Weigh 9.9g of the above sieved mixed powder, add 0.1g of sodium alginate, and manually grind until uniformly mixed to form a solid mixture with a total mass of 10g, wherein the mass ratio of the mixed powder to sodium alginate is 99:1.

[0099] Liquid component addition and mixing: Add 2 mL of 2 wt.% hydroxypropyltrimethylammonium chloride chitosan (HACC) aqueous solution to the solid mixture, and then add 1 mL of glycerol. Place the mixture on a magnetic stirrer and stir at 300 rpm for 30 min until a uniform paste ink with no particles and suitable flowability is formed, and set aside for later use.

[0100] S2, 3D printed porous bracket

[0101] Equipment and model preparation: A 4-channel ink direct-write 3D printer was used as the printing device; SolidWorks 2022 software was used to build the 3D model of the support, exported as .stl format, and imported into Cura 4.8.25 software for slicing.

[0102] Printing parameter settings: Slicing parameters are set as follows: printing speed 3mm / s, printing pressure 500kPa, initial layer height 0.2mm, subsequent layer height 0.2mm, nozzle inner diameter 0.3mm, printing line width 0.3mm, fill density 30%, and nozzle temperature is maintained at room temperature to avoid premature ink curing.

[0103] Scaffold Printing: The paste-like ink prepared in step S1 was loaded into a 10cc 3D printing cartridge and installed into the printer nozzle; the .gcode format file generated after slicing was imported into the printer control system, and layer-by-layer deposition printing was performed. According to experimental requirements, two sizes of scaffolds were printed: 10×10×6mm (length×width×height) for in vitro experiments and 6×2mm (diameter×height) for in vivo experiments. After printing, a porous scaffold blank was obtained.

[0104] S3, First crosslinking

[0105] The porous scaffold preform obtained in step S2 was completely immersed in a 4% calcium chloride solution and allowed to stand at room temperature for 1 minute for crosslinking, allowing sodium alginate to form ionic bonds with calcium ions and improving the initial structural stability of the scaffold preform. After crosslinking, the scaffold preform was removed, excess solution was blotted off with filter paper, and it was dried in a well-ventilated environment at room temperature for 20 hours to obtain a pre-crosslinked BCP scaffold.

[0106] S4, Second crosslinking

[0107] Glutaraldehyde crosslinking: The scaffold dried in step S3 is immersed in a 0.5% glutaraldehyde solution and placed in a 35°C constant temperature water bath for 22 hours for crosslinking. Covalent bonds are formed through the reaction of glutaraldehyde with the amino groups of HACC, which further enhances the mechanical properties and degradation resistance of the scaffold.

[0108] Washing and freeze-drying: After cross-linking, the scaffold was removed and washed repeatedly with deionized water three times for 10 minutes each time to remove residual glutaraldehyde solution and avoid cytotoxicity. The washed scaffold was then placed in a vacuum freeze dryer at -50℃ and 10 Pa for 20 hours to obtain a double-crosslinked biphasic calcium phosphate scaffold.

[0109] S5. Preparation of gelatin methacrylamide hydrogel

[0110] Preparation of photoinitiator solution: Place lithium phenyl (2,4,6-trimethylbenzoyl)phosphate in a brown bottle, add an appropriate amount of phosphate buffer, heat in a 45°C water bath for 15 min, shaking 3 times during the process to ensure complete dissolution, and prepare a 0.2 wt.% LAP standard solution.

[0111] Preparation of GelMA hydrogel solution: Weigh methacryloyl gelatin powder and add it to the above LAP standard solution to prepare an 8 wt.% GelMA mixed solution. Place the mixed solution in a 50℃ constant temperature water bath and shake it 5 times for 30 seconds each time to promote the dissolution of GelMA. After complete dissolution, centrifuge it at 2000 rpm for 1 minute to remove air bubbles from the solution.

[0112] Sterilization: The centrifuged GelMA solution was filtered and sterilized using a 0.22 μm pore size filter membrane to obtain a sterile GelMA hydrogel solution for later use.

[0113] S6, loaded with teriparatide and modified

[0114] PTH solution preparation: Teriparatide (PTH) powder was dissolved in ultrapure water to prepare an 80 μg / mL PTH solution.

[0115] Preparation of modified hydrogel: The above PTH solution and the GelMA hydrogel solution prepared in step S5 are mixed at a volume ratio of 1:9. After stirring evenly, azide-polyethylene glycol-active ester is added at a ratio of 0.1 mg AC-PEG-NHS per mg PTH. Stir for 10 min to ensure thorough mixing.

[0116] Incubation in the dark: The mixed solution was sealed with sealing film and incubated in a refrigerator at 4°C for 20 hours in the dark to allow the amino groups of AC-PEG-NHS and PTH to undergo a coupling reaction, thereby achieving stable loading and modification of PTH and finally obtaining a modified hydrogel solution loaded with teriparatide.

[0117] S7. Final fabrication of the composite scaffold

[0118] Scaffold immersion loading: The double crosslinked BCP scaffold obtained in step S4 is completely immersed in the modified hydrogel solution prepared in step S6, and placed in a refrigerator at 4°C for 20 hours to ensure that the hydrogel solution fully penetrates into the pores of the BCP scaffold.

[0119] Blue light crosslinking and curing: Remove the soaked scaffold, drain the excess solution from the surface, and irradiate it with a 400nm blue light irradiator for 15s to cause the GelMA to undergo a photopolymerization reaction, forming a stable hydrogel network, and fixing PTH in the pores of the scaffold.

[0120] Freeze-drying and shaping: The blue light cross-linked composite scaffold is placed in a vacuum freeze dryer, the temperature is set to -50℃, and it is dried for 20 hours to remove the moisture in the scaffold, thus obtaining the final double cross-linked 3D printed bone repair scaffold.

[0121] Example 3

[0122] This application provides a method for preparing a double cross-linked 3D printed bone repair scaffold, specifically including the following steps:

[0123] S1, Preparation of biphase calcium phosphate ink

[0124] Preparation of mixed powder: Take α-tricalcium phosphate (α-TCP) and β-tricalcium phosphate (β-TCP) and mix them at a mass ratio of 4:6. Pass the mixed powder through a 200-mesh sieve to remove coarse particles to ensure ink uniformity.

[0125] Preparation of solid mixture: Weigh 9.5g of the above sieved mixed powder, add 0.5g of sodium alginate, and manually grind until uniformly mixed to form a solid mixture with a total mass of 10g, wherein the mass ratio of the mixed powder to sodium alginate is 95:5.

[0126] Liquid component addition and mixing: Add 4 mL of 4 wt.% hydroxypropyltrimethylammonium chloride chitosan (HACC) aqueous solution to the solid mixture, and then add 3 mL of glycerol. Place the mixture on a magnetic stirrer and stir at 300 rpm for 30 min until a uniform paste ink with no particles and suitable flowability is formed, and set aside for later use.

[0127] S2, 3D printed porous bracket

[0128] Equipment and model preparation: A 4-channel ink direct-write 3D printer was used as the printing device; SolidWorks 2022 software was used to build the 3D model of the support, exported as .stl format, and imported into Cura 4.8.25 software for slicing.

[0129] Printing parameter settings: Slicing parameters are set as follows: printing speed 5mm / s, printing pressure 550kPa, initial layer height 0.5mm, subsequent layer height 0.5mm, nozzle inner diameter 0.5mm, printing line width 0.5mm, fill density 50%, and nozzle temperature is maintained at room temperature to avoid premature ink curing.

[0130] Scaffold Printing: The paste-like ink prepared in step S1 was loaded into a 10cc 3D printing cartridge and installed into the printer nozzle; the .gcode format file generated after slicing was imported into the printer control system, and layer-by-layer deposition printing was performed. According to experimental requirements, two sizes of scaffolds were printed: 10×10×6mm (length×width×height) for in vitro experiments and 6×2mm (diameter×height) for in vivo experiments. After printing, a porous scaffold blank was obtained.

[0131] S3, First crosslinking

[0132] The porous scaffold preform obtained in step S2 was completely immersed in a 6% calcium chloride solution and allowed to stand at room temperature for 3 minutes for crosslinking, allowing sodium alginate to form ionic bonds with calcium ions and improving the initial structural stability of the scaffold preform. After crosslinking, the scaffold preform was removed, excess solution was blotted off with filter paper, and it was dried in a well-ventilated environment at room temperature for 28 hours to obtain a pre-crosslinked BCP scaffold.

[0133] S4, Second crosslinking

[0134] Glutaraldehyde crosslinking: The scaffold dried in step S3 is immersed in a 1.5% glutaraldehyde solution and placed in a 39°C constant temperature water bath for 26 hours for crosslinking. Covalent bonds are formed through the reaction of glutaraldehyde with the amino groups of HACC, which further enhances the mechanical properties and degradation resistance of the scaffold.

[0135] Washing and freeze-drying: After cross-linking, the scaffold was removed and washed repeatedly with deionized water three times for 10 minutes each time to remove residual glutaraldehyde solution and avoid cytotoxicity. The washed scaffold was then placed in a vacuum freeze dryer at -80℃ and 10 Pa for 28 hours to obtain a double-crosslinked biphasic calcium phosphate scaffold.

[0136] S5. Preparation of gelatin methacrylamide hydrogel

[0137] Preparation of photoinitiator solution: Place lithium phenyl (2,4,6-trimethylbenzoyl)phosphate in a brown bottle, add an appropriate amount of phosphate buffer, heat in a 45°C water bath for 15 min, shaking 3 times during the process to ensure complete dissolution, and prepare a 0.3 wt.% LAP standard solution.

[0138] Preparation of GelMA hydrogel solution: Weigh methacryloyl gelatin powder and add it to the above-mentioned LAP standard solution to prepare a 12 wt.% GelMA mixed solution. Place the mixed solution in a 60℃ constant temperature water bath and shake it 5 times for 30 seconds each time to promote the dissolution of GelMA. After complete dissolution, centrifuge it at 4000 rpm for 3 minutes to remove air bubbles from the solution.

[0139] Sterilization: The centrifuged GelMA solution was filtered and sterilized using a 0.22 μm pore size filter membrane to obtain a sterile GelMA hydrogel solution for later use.

[0140] S6, loaded with teriparatide and modified

[0141] PTH solution preparation: Teriparatide powder was dissolved in ultrapure water to prepare a 120 μg / mL PTH solution.

[0142] Preparation of modified hydrogel: The above PTH solution and the GelMA hydrogel solution prepared in step S5 are mixed at a volume ratio of 1:9. After stirring evenly, azide-polyethylene glycol-active ester is added at a ratio of 0.3 mg AC-PEG-NHS per mg PTH. Stir for 10 min to ensure thorough mixing.

[0143] Incubation in the dark: The mixed solution was sealed with sealing film and incubated in a refrigerator at 4°C for 28 hours in the dark to allow the amino groups of AC-PEG-NHS and PTH to undergo a coupling reaction, thereby achieving stable loading and modification of PTH and finally obtaining a modified hydrogel solution loaded with teriparatide.

[0144] S7. Final fabrication of the composite scaffold

[0145] Scaffold immersion loading: The double cross-linked BCP scaffold obtained in step S4 is completely immersed in the modified hydrogel solution prepared in step S6, and placed in a refrigerator at 4°C for 28 hours to ensure that the hydrogel solution fully penetrates into the pores of the BCP scaffold.

[0146] Blue light crosslinking and curing: Remove the soaked scaffold, drain the excess solution from the surface, and irradiate it with a 410nm blue light irradiator for 25s to cause the GelMA to undergo a photopolymerization reaction, forming a stable hydrogel network, and fixing PTH in the pores of the scaffold.

[0147] Freeze-drying and shaping: The blue light cross-linked composite scaffold is placed in a vacuum freeze dryer, the temperature is set to -80℃, and it is dried for 28 hours to remove the moisture in the scaffold, thus obtaining the final double cross-linked 3D printed bone repair scaffold.

[0148] Comparative Example 1

[0149] The difference between this comparative example and Example 1 is that it only proceeds to step S4, omitting steps S5-S7; the specific steps are as follows:

[0150] The operations of steps S1-S4 are exactly the same as in Example 1. After obtaining the bi-crosslinked biphasic calcium phosphate scaffold, it is directly used as the final product, which is a pure BCP scaffold without any hydrogel loading.

[0151] Comparative Example 2

[0152] The difference between this comparative example and Example 1 is that: in step S5, pure GelMA hydrogel is prepared without teriparatide, and step S6 is omitted, i.e., teriparatide is not loaded, nor is AC-PEG-NHS modification added; the pure GelMA hydrogel is directly loaded onto the BCP scaffold. The specific differences in the steps are as follows:

[0153] Step S5: The procedure for preparing gelatin methacrylamide hydrogel is the same as in Example 1, but only a 10 wt.% GelMA hydrogel solution is prepared, containing 0.25 wt.% LAP and excluding teriparatide.

[0154] Step S6: This step is omitted, as there is no preparation of teriparatide solution, no preparation of modified hydrogel, and no light-protected incubation.

[0155] Step S7: The double crosslinked BCP scaffold obtained in step S4 is completely immersed in the pure GelMA hydrogel solution prepared in step S5. The subsequent soaking for 24 hours, crosslinking by blue light irradiation, and freeze drying are all the same as in Example 1, to obtain a scaffold loaded only with GelMA hydrogel.

[0156] Comparative Example 3

[0157] The difference between this comparative example and Example 1 is that AC-PEG-NHS is not added in step S6, meaning that teriparatide is not modified by azide-polyethylene glycol-active ester; the specific differences are as follows:

[0158] Step S6: The PTH solution is prepared in the same way as in Example 1;

[0159] Preparation of modified hydrogel: Mix only the PTH solution with the GelMA hydrogel solution prepared in step S5 at a volume ratio of 1:9, without adding AC-PEG-NHS, and stir for 10 min to ensure uniform mixing;

[0160] The same light-protected incubation process as in Example 1 was performed to obtain a GelMA hydrogel solution loaded with unmodified teriparatide.

[0161] The remaining steps are completely consistent with those in Example 1, and the final result is a GelMA hydrogel scaffold loaded with unmodified teriparatide.

[0162] I. Physical and Chemical Properties Characterization Tests of Scaffold Foundation

[0163] The double-crosslinked 3D-printed bone repair scaffolds prepared in Examples 1-3, as well as the pure BCP scaffolds, scaffolds loaded only with GelMA hydrogel, and GelMA hydrogel scaffolds loaded with unmodified teriparatide prepared in Comparative Examples 1-3, were used. Three parallel samples were prepared for each sample, and the following tests were performed: Scanning electron microscopy was used to observe the surface morphology and pore structure of the scaffolds. All samples were freeze-dried, sputter-coated with gold, and then images were taken at 5kV accelerating voltage to analyze pore uniformity and hydrogel coverage. X-ray diffraction was used to analyze the crystal structure of the scaffolds, with a scanning range of 2θ = 10°-80° and a step size of 0.02°, and the changes in the α-TCP / β-TCP characteristic peaks of each group of scaffolds were compared. Fourier transform infrared spectroscopy was used to detect the chemical structure of the scaffolds, with a scanning range of... resolution The coupling of GelMA, PTH, and AC-PEG-NHS was verified. The mechanical properties of the scaffold were tested using a universal mechanical testing machine, with a compression rate of 1 mm / min until fracture. The compressive strength and elastic modulus were recorded. The porosity of the scaffold was determined using a mercury porosimeter under a pressure range of 0-414 MPa, and the total porosity and pore size distribution were calculated. The degradation performance was evaluated using an in vitro degradation test. Each group of scaffolds was immersed in PBS buffer at 37°C, with the PBS being replaced every two days. The scaffolds were removed at 1, 2, and 3 months, freeze-dried, and weighed to calculate the mass loss rate. The drug release performance was detected using high-performance liquid chromatography (HPLC). Scaffolds containing PTH were placed in PBS buffer at 37°C and shaken in the dark. Samples were taken at 1, 4, 8, 12, 24, 48, and 72 hours and at 7, 14, and 21 days to determine the PTH concentration in the supernatant and plot the drug release curve.

[0164] II. In vitro biocompatibility assessment test of scaffold

[0165] The scaffolds from Examples 1-3 and Comparative Examples 1-3 were used to prepare extracts for each scaffold according to the International Organization for Standardization / EN10993-5 standard. Specifically, 1 mL of serum-free DMEM medium was used for every 100 mg of sample. The samples and medium were mixed and incubated at 37°C for 24 h. After incubation, the extracts were filtered through a 0.22 μm filter for sterilization. Finally, 10% fetal bovine serum and 1% penicillin / streptomycin were added to the sterilized extracts. Rat bone marrow mesenchymal stem cells were used as the research object, and the following tests were performed in a 37°C, 5% CO2 saturated humidity incubator: Cell proliferation was detected using the CCK-8 assay, with rBMSCs at a concentration of 5 × 10⁻⁶ cells / mL. 4 Seeds were planted at a density of 1 / mL in 96-well plates, with 100 μL of culture medium containing the extract of each scaffold added to each well. Six replicates were made per group. After culturing for 1, 3, and 7 days, culture medium containing 10% CCK-8 reagent was added, and the plates were incubated at 37°C in the dark for 2 hours. The absorbance at 450 nm was measured using a microplate reader. Cell viability was detected by staining with 5 × 10⁶ cells / mL. 4Cells were seeded per well in 24-well plates and cultured for 24 hours. Afterward, the extracts from each scaffold were added, and the plates were cultured for another 24 hours. The culture medium was discarded, and the cells were washed twice with 1×AssayBuffer. 1 mL of 1×AssayBuffer containing 2 μL Calcein-AM stock solution was added, and the plates were incubated at 37°C in the dark for 25 minutes. Then, 3 μL PPI stock solution was added, and the plates were incubated at room temperature in the dark for 5 minutes. After washing, the distribution of live and dead cells was observed under a fluorescence microscope. Cell migration was detected using the scratch assay. rBMSCs were seeded in 6-well plates until 90% confluence, and scratched with the tip of a pipette before adding the extracts from each scaffold. The culture medium contained 10 μg / mL mitomycin C to inhibit proliferation. The cells were photographed using an optical microscope at 0, 24, and 48 h, and the scratch closure rate was calculated using ImageJ. For cell adhesion and spreading assays, rBMSCs were co-cultured with the extracts of each scaffold for 3 and 5 days, fixed with 4% paraformaldehyde for 1 h, permeabilized with 0.5% Triton X-100 for 30 min, stained with TRITC-phalloidin at room temperature in the dark for 1 h, and stained with DAPI. The cytoskeleton morphology was observed using a laser confocal microscope, and the three-dimensional adhesion of cells on the scaffold surface was observed using SEM.

[0166] 3. In vitro osteogenic performance evaluation test of scaffold

[0167] The scaffold extracts from Examples 1-3 and Comparative Examples 1-3 were prepared using the same method as in the biocompatibility test. Osteogenic induction was induced by co-culturing with rBMSCs. The osteogenic induction medium was high-glucose DMEM containing 10% fetal bovine serum, 1% penicillin / streptomycin, 0.008 μg / mL β-glycerophosphate, 1.76 μg / mL ascorbic acid, and 0.66 μg / mL dexamethasone. Three replicates were set up for each sample. The following tests were performed: ALP staining and activity detection were performed 7 days after osteogenic induction. The medium was discarded, the sample was washed twice with PBS, and alkaline phosphatase chromogenic reagent was added. The sample was incubated at 37°C for 30 min, and the staining was observed under an optical microscope. Simultaneously, the activity was measured using an alkaline phosphatase activity assay kit. The single-sample count was calculated. ALP activity in cells was detected. Alizarin Red staining was performed 14 days after osteogenic induction. Cells were washed twice with PBS, fixed with 4% paraformaldehyde for 30 min, stained with 0.1% Alizarin Red S solution (pH=8.3) at room temperature for 30 min, washed with PBS until no excess staining was found, and the formation of calcium nodules was observed under a stereomicroscope. The staining was eluted with 10% acetic acid, and the absorbance at 405 nm was measured using an ELISA reader to quantify the degree of mineralization. Osteogenesis-related gene expression was detected using qRT-PCR. Cells were collected 14 days after osteogenic induction, total RNA was extracted with TRIzol reagent, reverse transcribed into cDNA, and the mRNA expression levels of ColI, ALP, OCN, BMP2, RUNX2, and OPN genes were detected using a SYBR Green assay kit. The relative expression levels of GAPDH were calculated using the GAPDH internal reference gene. Osteogenesis-related protein expression was detected using Western blot and immunofluorescence assays. Western blot analysis was performed 14 days after osteogenic induction. Cells were collected, and total protein was extracted using RIPA lysis buffer containing protease inhibitors. Protein concentration was determined using the BCA method. 20 μg of protein was loaded onto each sample. After SDS-PAGE electrophoresis, membrane transfer, and blocking with 5% skim milk, primary antibodies (ALP, OCN, RUNX2, and OPN) were added and diluted 1:1000. The membrane was incubated overnight at 4°C. Secondary antibodies were then added and diluted 1:8000, and the membrane was incubated at room temperature for 1 hour. Finally, ECL substrate was used for color development, and ImageJ was used for quantification. Band gray values; Immunofluorescence detection was performed 3 days after osteogenic induction. After cell fixation and permeabilization, BMP2 and ALP primary antibodies were added and diluted 1:200, and incubated overnight at 4°C. Secondary fluorescent antibody was incubated at room temperature for 2 hours. Nuclei were stained with DAPI, and protein localization and expression were observed using LSCM. Transcriptome analysis was performed 7 days after osteogenic induction. Total RNA was extracted from rBMSCs of each sample. RNA purity was detected by Nanodrop, and RNA integrity was assessed using an Agilent 2100 bioanalyzer. Libraries were constructed using the VAHTSU Universal V6 RNA-seq library preparation kit, and transcriptome sequencing was performed by OE Biotech Ltd. to analyze differentially expressed genes and enriched osteogenic-related signaling pathways.

[0168] IV. Evaluation Test of Scaffold In Vitro Immunomodulatory Performance

[0169] The scaffold extracts from Examples 1-3 and Comparative Examples 1-3 were prepared using the same methods as in the biocompatibility test. Mouse RAW264.7 macrophages were used as the research object and cultured in DMEM medium containing 10% fetal bovine serum inactivated at 56°C for 30 min and 1% penicillin / streptomycin at 37°C in a 5% CO2 incubator. The following assays were performed: Macrophage polarization assays were performed using immunofluorescence double staining, qRT-PCR, and flow cytometry. Macrophages were first activated by stimulating them with medium containing 1000 ng / mL LPS for 6 h, followed by culturing with the respective scaffold extracts for another 24 h. For immunofluorescence double staining, after cell smear fixation and permeabilization, primary antibodies against M2 macrophage marker CD206 and M1 macrophage marker iNOS were added at a 1:200 ratio and incubated overnight at 4°C. Secondary fluorescent antibodies were then added and incubated at room temperature for 2 h. Finally, nuclei were stained with DAPI and observed under a fluorescence microscope. qRT-PCR was performed on macrophages. The assays were performed at 6 h and 24 h after culture. Total RNA was extracted from the cells, and the mRNA expression levels of TNF-α, IL-1β (M1-related genes) and TGF-β, IL-10 (M2-related genes) were detected. Flow cytometry was performed after 24 h of culture. Cells were collected, washed with PBS, centrifuged, and then labeled with direct antibodies against CD86 (M1 macrophage marker) and CD206 (M2 macrophage marker). The cells were reacted at room temperature in the dark for 30 min. After washing with PBS, the proportion of positive cells was detected by flow cytometry. ROS scavenging capacity was detected after 6 h of LPS stimulation of macrophages, followed by 6 h and 24 h of culture with the extracts of each scaffold. Cells were then washed with PBS, and serum-free medium containing 10 μM DCFH-DA probe was added. The cells were incubated at 37 °C in the dark for 30 min. After washing with PBS, the fluorescence intensity was observed under an inverted fluorescence microscope, and the fluorescence values ​​at 488 nm excitation and 525 nm emission wavelengths were measured using a microplate reader to assess the intracellular ROS level.

[0170] V. Evaluation Test of Bone Repair Performance and Biocompatibility in the Scaffold

[0171] Fifty male SD rats aged 6-8 weeks were randomly divided into 5 groups of 10 each. Each group received a corresponding scaffold implanted. Comparative groups 1-3 received pure BCP scaffolds, scaffolds loaded only with GelMA hydrogel, and GelMA hydrogel scaffolds loaded with unmodified teriparatide, respectively. Examples 1-3 were combined into a composite scaffold group and implanted with a double-crosslinked 3D-printed bone repair scaffold. The blank control group received no implantation. This method was used to construct a rat model of critical skull defect with a diameter of 5 mm. After anesthesia, a 1.5 cm sagittal incision was made in the scalp, and the skull was exposed through blunt dissection. After preparing the defect area with a trephine, the corresponding scaffold was implanted. For three days post-surgery, rats were injected intramuscularly daily with 200,000 U / kg penicillin to prevent infection. Rats were sacrificed at 6 and 12 weeks post-surgery for the following tests: For Micro-CT imaging analysis, whole skull specimens were taken from rats and fixed with 4% paraformaldehyde for 48 hours. Micro-CT scans were performed using bone window parameters with a slice thickness of 35 μm. Three-dimensional reconstruction was performed using Recon software, and Avatar software was used to quantitatively analyze the new bone volume fraction (BV / TV) and trabecular bone number in the region of interest (ROI) of the defect area. The histological analysis included the amount of bone (Tb.N), trabecular thickness (Tb.Th), and trabecular spacing (Tb.Sp). For histological staining analysis, skull specimens were decalcified with EDTA, embedded in paraffin, and cut into 5μm thick sections. HE staining was performed to observe the morphology of the defect area, and Masson's trichrome staining was performed to observe collagen deposition. Images were taken using an optical microscope, and ImageJ was used to analyze the proportion of new bone area to the defect area. For immunohistochemical and immunofluorescence staining analysis, after dewaxing paraffin sections to water, immunohistochemical staining was used to detect the expression of Runx2, OCN, and ALP. The procedure involved adding primary antibody and incubating overnight at 4°C, followed by adding secondary antibody and incubating at room temperature for 1 hour, and finally developing the color with DAB. Immunofluorescence was used to detect the expression of VEGF and CD31. The procedure involved adding fluorescent secondary antibody and incubating, followed by staining the nucleus with DAPI, and observing the distribution and intensity of positive signals using a fluorescence microscope. Systemic biocompatibility was assessed 12 weeks post-surgery. Heart, liver, spleen, lung, and kidney tissues from rats were collected, fixed in 4% paraformaldehyde, embedded in paraffin, and sectioned. After HE staining, the tissue morphology was observed using an optical microscope to assess the toxic effects of the scaffold on major organs.

[0172] The test results for each embodiment and comparative example are as follows:

[0173] I. Physical and Chemical Properties of the Support Foundation

[0174] Example 1: SEM observation showed a uniform porous structure with a pore size of 400-500 μm. GelMA hydrogel uniformly covered the BCP surface, and PTH was dispersed without agglomeration. XRD retained the characteristic peaks of α-TCP / β-TCP, with corresponding 2θ values ​​of approximately 25.8° and 31.7°. FTIR showed characteristic peaks of GelMA amide bond at 1650 cm⁻¹, PTH amino group at 1240 cm⁻¹, and AC-PEG-NHS azide group at 2100 cm⁻¹, verifying successful coupling of the components.

[0175] Example 2: SEM observation showed that the porous structure was uniform, the GelMA hydrogel was completely covered, and the PTH was uniformly dispersed; XRD retained the characteristic peaks of α-TCP / β-TCP, and FTIR verified that the coupling of GelMA, PTH, and AC-PEG-NHS was successful.

[0176] Example 3: SEM observation showed uniform pore distribution, tight binding between GelMA hydrogel and BCP, and no PTH aggregation; XRD retained α-TCP / β-TCP characteristic peaks, and FTIR showed clear characteristic peaks, indicating stable coupling effect.

[0177] Comparative Example 1: SEM observation showed a porous structure without any hydrogel coverage and a relatively rough surface; XRD showed only α-TCP / β-TCP characteristic peaks, and FTIR showed no GelMA or PTH related infrared peaks.

[0178] Comparative Example 2: SEM observation of the porous structure showed that the GelMA hydrogel was uniformly covered and there was no PTH component; XRD retained the characteristic peaks of α-TCP / β-TCP, and FTIR showed the GelMA infrared peak, with no PTH or AC-PEG-NHS peaks.

[0179] Comparative Example 3: SEM observation showed porous structure with GelMA hydrogel coverage and slight local aggregation of PTH; XRD retained the characteristic peaks of α-TCP / β-TCP, and FTIR showed GelMA and PTH peaks, but no AC-PEG-NHS peak.

[0180] II. In vitro biocompatibility of scaffolds

[0181] Example 1: LSCM observation showed that the cells were spindle-shaped with a regular skeletal arrangement and a long axis length of about 80 μm.

[0182] Example 2: LSCM observation showed that the cells spread well, the cytoskeleton was evenly distributed, and the long axis length was about 75 μm.

[0183] Example 3: LSCM observation showed that the cells were well spread, with a clear cytoskeleton and a long axis length of approximately 85 μm.

[0184] Comparative Example 1: LSCM observation showed that the cells were round, not fully spread, and had a short and disordered cytoskeleton.

[0185] Comparative Example 2: LSCM observation showed that the cells spread out generally, the cytoskeleton was arranged relatively regularly, and the long axis length was about 65 μm.

[0186] Comparative Example 3: LSCM observation showed that the cells spread well, the cytoskeleton was arranged relatively regularly, and the long axis length was about 70 μm.

[0187] Blank control: LSCM observation showed normal cell spread, regular cytoskeleton arrangement, and a long axis length of approximately 72 μm.

[0188] III. In-vivo biosafety

[0189] Examples 1-3, Comparative Examples 1-3, and Blank Control: 12 weeks post-surgery, HE staining of heart, liver, spleen, lung, and kidney tissues in all groups showed no abnormal pathological changes and normal cell morphology, confirming that the scaffold had no systemic toxicity.

[0190] The physical and chemical properties characterization data of the scaffold foundation are shown in Table 1.

[0191] Table 1:

[0192] Group Porosity (%) Compressive strength (MPa) In vitro degradation rate (%) after 12 weeks PTH 14-day cumulative release rate (%) Example 1 60.2±2.3 78.5±3.1 31.8±2.1 84.5±3.2 Example 2 58.6±2.5 72.3±2.8 30.5±1.9 82.1±2.9 Example 3 62.1±2.4 81.7±3.3 33.2±2.4 85.7±3.5 Comparative Example 1 52.3±2.1 95.6±4.2 18.5±1.5 - Comparative Example 2 57.8±2.3 70.2±2.6 28.9±1.8 - Comparative Example 3 59.5±2.2 75.1±2.9 30.1±2.0 98.2±2.7

[0193] The data on the in vitro biocompatibility assessment of the stent are shown in Table 2.

[0194] Table 3:

[0195] Group CCK-8 absorbance (450nm, 7d) viable cell rate (%, 24h) Scratch closure rate (%, 72h) Example 1 1.78±0.09 96.5±1.2 69.2±3.5 Example 2 1.65±0.08 95.3±1.5 65.8±3.2 Example 3 1.82±0.10 97.1±1.1 71.5±3.6 Comparative Example 1 1.21±0.07 90.2±1.8 52.3±2.8 Comparative Example 2 1.43±0.08 93.5±1.4 58.6±3.0 Comparative Example 3 1.52±0.09 94.8±1.3 62.1±3.1

[0196] The data on the in vitro osteogenic performance of the scaffold are shown in Table 3.

[0197] Table 3:

[0198] Group ALP activity (U / g protein, 14d) Alizarin Red Absorbance (405nm, 14d) Relative expression levels of osteogenic genes (14 days, with blank control as 1). OCN protein grayscale values ​​(relative values, 14d) Example 1 118.5±5.2 1.48±0.07 ALP:2.8; OCN:3.1; BMP2:2.9; RUNX2:3.0 1.85±0.12 Example 2 105.3±4.8 1.32±0.06 ALP:2.5; OCN:2.7; BMP2:2.6; RUNX2:2.8 1.62±0.10 Example 3 122.7±5.5 1.53±0.08 ALP:3.0;OCN:3.3;BMP2:3.1;RUNX2:3.2 1.92±0.13 Comparative Example 1 45.2±3.1 0.65±0.04 ALP:1.1;OCN:1.0;BMP2:1.2;RUNX2:1.1 0.82±0.06 Comparative Example 2 68.3±3.5 0.92±0.05 ALP:1.5;OCN:1.4;BMP2:1.6;RUNX2:1.5 1.15±0.08 Comparative Example 3 92.6±4.2 1.18±0.06 ALP:2.2;OCN:2.0;BMP2:2.1;RUNX2:2.3 1.48±0.10 Blank control 42.8±2.9 0.61±0.03 ALP:1.0; OCN:1.0; BMP2:1.0; RUNX2:1.0 0.80±0.05

[0199] The data on the in vitro immunomodulatory performance of the scaffold are shown in Table 4.

[0200] Table 4:

[0201] Group M2 macrophage percentage (CD206⁺%, 24h) Relative expression level of M1-related gene TNF-α (24h, with blank control as 1). ROS fluorescence values ​​(488nm excitation / 525nm emission, 24h) Example 1 44.8±3.2 0.32±0.04 82.5±4.1 Example 2 41.5±2.9 0.38±0.05 88.6±4.3 Example 3 46.2±3.5 0.29±0.03 79.8±3.8 Comparative Example 1 22.3±2.1 0.85±0.07 135.2±6.5 Comparative Example 2 28.6±2.5 0.68±0.06 112.8±5.2 Comparative Example 3 35.1±2.8 0.45±0.04 98.5±4.7 Blank control 18.5±1.8 1.00±0.08 152.6±7.1

[0202] The data on bone repair performance in the scaffold are shown in Table 5.

[0203] Table 5:

[0204] Group BV / TV (%) at 12 weeks post-surgery Number of trabeculae (Tb.N) (pieces / mm²) at 12 weeks post-surgery Tb.Th (μm) at 12 weeks post-surgery Percentage of new bone area to defect area at 12 weeks post-surgery (%) Number of VEGF⁺ vessels per field of view at 12 weeks post-surgery Example 1 44.8±3.5 1185±42 85.2±4.1 84.5±4.2 27.8±2.3 Example 2 41.2±3.2 1120±38 80.5±3.8 79.8±3.8 25.3±2.1 Example 3 46.5±3.7 1210±45 88.6±4.3 86.2±4.5 29.1±2.5 Comparative Example 1 22.5±2.8 850±32 62.3±3.2 45.2±3.1 12.8±1.5 Comparative Example 2 28.6±3.0 925±35 68.5±3.5 58.3±3.5 16.5±1.8 Comparative Example 3 35.8±3.3 1050±39 75.2±3.7 72.1±3.9 21.2±2.0 Blank control 15.2±2.1 780±28 55.8±2.9 32.5±2.7 9.6±1.2

[0205] Combining Examples 1-3 and Comparative Example 1 with Tables 1, 2, 3, 4, and 5, it can be seen that the performance of the pure BCP scaffold differs significantly from that of the Example group due to the absence of GelMA hydrogel and modified PTH. From the perspective of basic scaffold characteristics, the lack of GelMA hydrogel encapsulation results in relatively low scaffold porosity and a significantly slower degradation rate, making it difficult to match the "degradation-osteoogenesis" rhythm in bone regeneration. From a biocompatibility perspective, the absence of hydrogel to improve surface hydrophilicity and cell adhesion environment leads to poor cell spreading morphology on the scaffold surface, and limitations in proliferation and migration. From a functional perspective, the lack of PTH's bioactivity regulation significantly restricts osteogenic physiological processes and immune regulation capabilities. In in vivo bone defect repair, the quantity, quality, and vascularization of new bone formation are also far inferior to those in the Example group. This difference confirms the role of GelMA hydrogel as a bioactive carrier. It can not only optimize the surface properties of the scaffold and improve cell compatibility, but also provide a loading and release environment for PTH. As a key factor in promoting bone regeneration, PTH can synergistically activate osteogenic pathways and immune regulation. Without the synergistic effect of these two, the scaffold can only provide basic mechanical support and cannot achieve the function of actively regulating bone regeneration.

[0206] Combining Examples 1-3 and Comparative Example 2 with Tables 1, 3, 4, and 5, it can be seen that although the scaffold loaded with GelMA hydrogel alone shows some performance improvement compared to the pure BCP scaffold, and the hydrophilicity and biocompatibility of GelMA improve the cell spreading and proliferation environment and make the scaffold degradation rate closer to the bone regeneration requirements, its overall function is still inferior to the example group due to the lack of PTH. In terms of osteogenic properties, the alkaline phosphatase activity, extracellular calcium nodule mineralization degree, and osteogenic-related gene expression levels of this scaffold are significantly lower than those of the example group loaded with PTH. In terms of immune regulation, its ability to induce macrophage M2 polarization and its effect on scavenging reactive oxygen species are weak, and it cannot effectively construct an immune microenvironment conducive to bone regeneration. Reflected in in vivo bone defect repair, the proportion of new bone area and the number of angiogenesis are also relatively low. PTH is the core active ingredient that promotes osteogenic differentiation and immune microenvironment remodeling. GelMA hydrogel only plays the role of carrier. Without the activity enhancement of PTH, the scaffold cannot achieve the synergistic bone-promoting effect of carrier and drug, and the bone repair efficiency is naturally limited.

[0207] Based on Examples 1-3 and Comparative Example 3, and in conjunction with Tables 1, 3, 4, and 5, it can be seen that although the scaffold loaded with unmodified PTH achieved initial loading of PTH through GelMA hydrogel, the lack of AC-PEG-NHS coupling modification prevented PTH from forming a stable loading state, resulting in an excessively fast release rate and a short duration of bioactivity. This issue directly affects the long-term performance of the scaffold: in terms of osteogenic promotion, although it can activate osteogenic processes to a certain extent in the short term, it is difficult to maintain the sustained high expression of osteogenic genes and the large accumulation of calcium nodules; in terms of immune regulation, it cannot induce macrophage M2 polarization and scavenge reactive oxygen species in the long term, and the anti-inflammatory effect is not long-lasting; ultimately, in the later stage of bone defect repair in vivo, the increase in new bone volume fraction, bone density and trabecular bone thickness is not as good as in the example group. This modification can achieve local efficient sustained release of PTH and amplification of early osteogenic signals. Through coupling reaction, it stabilizes the structure and release rhythm of PTH, enabling it to continuously participate in the osteogenic and immune regulation process. Without this modification, the function of PTH cannot be exerted in the long term, and the bone repair performance of the scaffold will naturally decline.

[0208] This specific embodiment is merely an explanation of this application and is not intended to limit it. After reading this specification, those skilled in the art can make modifications to this embodiment without contributing any inventive step, but such modifications are protected by patent law as long as they fall within the scope of the claims of this application.

Claims

1. A method for preparing a double cross-linked 3D printed bone repair scaffold, characterized in that: Includes the following steps: S1. Preparation of biphasic calcium phosphate ink: Mix the mixed powder obtained by mixing α-tricalcium phosphate and β-tricalcium phosphate with sodium alginate to obtain a solid mixture; add hydroxypropyltrimethylammonium chloride chitosan aqueous solution and glycerol to the solid mixture and stir to form a homogeneous paste ink; S2, 3D Printed Porous Scaffold: A porous scaffold blank is obtained by printing paste ink using direct ink writing technology; S3, First crosslinking: The porous scaffold preform is immersed in calcium chloride solution for crosslinking and then dried; S4, Second crosslinking: The scaffold that has undergone the first crosslinking is immersed in glutaraldehyde solution for crosslinking, washed and then freeze-dried under vacuum to obtain a double crosslinked biphasic calcium phosphate scaffold. S5. Preparation of gelatin methacrylamide hydrogel: Dissolve gelatin methacrylamide in phosphate buffer, add photoinitiator, dissolve by water bath heating, shake and mix, centrifuge and filter to obtain gelatin methacrylamide hydrogel solution. S6. Loading and modifying teriparatide: The teriparatide solution was mixed with the gelatin methacrylamide hydrogel solution, and azide-polyethylene glycol-active ester was added. After incubation in the dark, a modified hydrogel solution loaded with teriparatide was obtained. S7. Composite scaffold preparation: The double cross-linked biphasic calcium phosphate scaffold is immersed in the modified hydrogel solution, cross-linked by blue light irradiation and freeze-dried to obtain the double cross-linked 3D printed bone repair scaffold; The mixed powder is prepared by mixing α-tricalcium phosphate and β-tricalcium phosphate at a mass ratio of 2:8 to 4:6, and then sieving it through a sieve with a mesh size of 100-200; the mass ratio of the mixed powder to sodium alginate is 95:5 to 99:

1. The mass concentration of the gelatin methacrylamide hydrogel solution is 8-12%; the photoinitiator is lithium phenyl-2,4,6-trimethylbenzoyl phosphate with a mass concentration of 0.2-0.3%.

2. The method for preparing a double cross-linked 3D printed bone repair scaffold according to claim 1, characterized in that: In step S1, the mass concentration of the hydroxypropyltrimethylammonium chloride chitosan aqueous solution is 2-4 wt.%, and the amount added is 0.2-0.4 mL per gram of solid mixture; the amount of glycerol added is 0.1-0.3 mL per gram of solid mixture.

3. The method for preparing a double cross-linked 3D printed bone repair scaffold according to claim 1, characterized in that: In step S2, the printing parameters of the direct ink writing technology include: printing speed 3-5 mm / s, printing pressure 500-550 kPa, layer height 0.2-0.5 mm, nozzle inner diameter 0.3-0.5 mm, line width 0.3-0.5 mm, and fill density 30-50%.

4. The method for preparing a double cross-linked 3D printed bone repair scaffold according to claim 1, characterized in that: In step S3, the mass concentration of the calcium chloride solution is 4-6%, the crosslinking time is 1-3 min, and the drying is performed at room temperature for 20-28 h.

5. The method for preparing a double cross-linked 3D printed bone repair scaffold according to claim 1, characterized in that: In step S4, the mass concentration of the glutaraldehyde solution is 0.5-1.5%, the crosslinking temperature is 35-39°C, and the crosslinking time is 22-26h; the vacuum freeze-drying temperature is -50°C to -80°C, and the drying time is 20-28h.

6. The method for preparing a double cross-linked 3D printed bone repair scaffold according to claim 1, characterized in that: In step S5, the water bath heating temperature is 50-60°C, the centrifugation speed is 2000-4000 rpm, and the time is 1-3 min; the filtration is performed using a filter membrane with a pore size of 0.22 μm.

7. The method for preparing a double cross-linked 3D printed bone repair scaffold according to claim 1, characterized in that: In step S6, the concentration of the teriparatide solution is 80-120 μg / mL; the amount of azide-polyethylene glycol-active ester added is 0.1-0.3 mg per milligram of teriparatide; and the light-protected incubation conditions are incubation at 4°C for 20-28 h.

8. The method for preparing a double cross-linked 3D printed bone repair scaffold according to claim 1, characterized in that: In step S7, the soaking time is 20-28 hours; the wavelength of the blue light irradiation is 400-410 nm, and the irradiation time is 15-25 seconds; the freeze-drying temperature is -50°C to -80°C, and the time is 20-28 hours.

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