Polylactic acid-caprolactone copolymer PLCL degradable coronary stent with hybrid interpenetrating network structure and preparation method of polylactic acid-caprolactone copolymer PLCL degradable coronary stent

Abstract

The invention belongs to the technical field of medical apparatus and instruments, and relates to a polylactic acid-caprolactone copolymer PLCL degradable coronary stent with a hybrid interpenetrating network structure and a preparation method of the polylactic acid-caprolactone copolymer PLCL degradable coronary stent. The scaffold is composed of linear polylactic acid (PLLA) and a star-shaped polylactic acid-caprolactone copolymer (PLCL) macromonomer with a photoreactive crosslinking end group, a chemical crosslinking network is formed through photocrosslinking, a physical crystallization network is formed through controlled annealing and programmed crystallization, the two are interpenetrated to form a hybrid interpenetrating network structure, and the cross-linked network and the physical crystallization network are interpenetrated to form a cross-linked network. Therefore, the contradiction between the radial support maintenance and the degradation controllability of the degradable stent is improved. The stent has good radial supporting performance, controllable degradation behavior and good biocompatibility, and is suitable for interventional therapy of coronary artery diseases.

Description

Technical Field

[0001] This invention belongs to the field of biomedical polymer materials and interventional medical devices, specifically relating to a biodegradable coronary stent composed of linear polylactic acid (PLLA) and star-shaped polylactic acid-caprolactone copolymer (PLCL) with photoreactive crosslinking end groups, having a hybrid interpenetrating network structure in which chemical crosslinking networks and physical crystalline networks interpenetrate each other, as well as a method for preparing the stent. Background Technology

[0002] Coronary stents, as the core medical implant for treating coronary atherosclerotic heart disease, have become a focus of research and development in the medical device field in recent years due to their biodegradable form. This is because biodegradable stents can gradually degrade through hydrolysis after implantation and be metabolized and absorbed by the body, effectively avoiding long-term complications such as restenosis and endothelial dysfunction caused by long-term retention of traditional metal stents. However, existing biodegradable coronary stents still face several key technical bottlenecks: First, insufficient mechanical support performance. Most products rely on the physical processing and molding of a single polymer substrate, lacking targeted structural reinforcement design. The radial support force is insufficient to withstand the continuous mechanical load on the blood vessel wall over a long period, making them prone to deformation or collapse during the critical stage of vascular regeneration. Second, imbalanced degradation rate regulation. An excessively rapid degradation rate in the early stages leads to premature loss of mechanical support capacity during the critical period of vascular healing, while excessive degradation products can easily trigger local tissue irritation. Third, poor biocompatibility. Most products lack functional modifications that target and regulate cell adhesion and proliferation, easily inducing excessive macrophage infiltration and persistent inflammatory responses after implantation, thereby hindering the vascular endothelialization process and affecting clinical treatment outcomes. The aforementioned problems severely restrict the clinical translation and application potential of biodegradable coronary stents, which is also the core improvement direction of this invention. Summary of the Invention

[0003] In view of this, the purpose of this application is to provide a method for preparing a polylactic acid-caprolactone copolymer (PLCL) coronary composite stent that combines mechanical support and a controllable degradation rate. The stent prepared by this invention has excellent mechanical support performance, degradation rate and good biocompatibility.

[0004] In a first aspect, this application discloses a biodegradable coronary stent of polylactic acid-caprolactone copolymer (PLCL) with a hybrid interpenetrating network structure, the stent comprising a stent body formed of a polymer composite material; the polymer composite material comprises: linear polylactic acid (PLLA) and star-shaped polylactic acid-caprolactone copolymer (PLCL) macromonomers with photocrosslinked end groups; The support body also has: (1) A chemical crosslinking network formed by the photocrosslinking end groups under light irradiation; (2) A physical crystallization network formed by the hierarchical crystallization of the PLLA and the PLCL; Furthermore, the chemical cross-linked network and the physical crystallization network interpenetrate each other within the scaffold body to form a hybrid interpenetrating network structure.

[0005] Preferably, the photocrosslinking end group is cinnamoyl.

[0006] Preferably, the star-shaped polylactic acid-caprolactone copolymer (PLCL) macromonomer has a 6-arm star structure.

[0007] Preferably, the mass ratio of L-lactide to ε-caprolactone in the synthesis of the six-armed star-shaped PLCL macromonomer is 2.8-3.1:1, and more preferably 121:41 (i.e. 2.95:1).

[0008] Preferably, the mass fraction of PLLA in the polymer composite material is 30% to 80%; the purification and collection during the synthesis of the six-armed star-shaped PLCL macromonomer is performed using supercritical CO2 extraction. More preferably, the mass fraction of PLLA in the polymer composite material is 50% to 60%.

[0009] Preferably, the scaffold is a tubular structure, the tube of which is composed of composite nanofiber tubes; the composite nanofiber tubes are prepared by uniaxial electrospinning and gradient drying of the blended spinning solution to form a green embryo, which is then subjected to thermal densification, photocrosslinking and graded crystallization treatment; wherein, the electrospinning process utilizes a high-voltage electrostatic field to stretch the polymer at a high ratio, so that the polymer molecular chains are highly oriented along the fiber axis, providing an excellent axial mechanical basis for the scaffold; The thermal densification process (i.e., the annealing process before photocrosslinking) enables the highly oriented nanofibers to undergo controlled melting and fusion, eliminating inter-fiber voids and forming a dense and uniform tube wall. This endows the scaffold with excellent radial support and fluid sealing properties, while retaining the microscopic surface texture unique to electrospinning to promote cell adhesion. The temperature (120°C) of the thermal densification process (i.e., the annealing process before photocrosslinking) is precisely controlled between the cold crystallization temperature and melting temperature of PLLA. Within this temperature range, the polymer molecular chains gain sufficient mobility for rearrangement and fusion, while the orientation structure induced by electrospinning is preserved. This unique "orientation retention-dense fusion" mechanism allows the final tube wall material to not only possess the density of traditional extruded tubes but also retain a much higher degree of molecular chain orientation, which is key to achieving a balance between high radial support and excellent axial flexibility.

[0010] The blending spinning solution uses hexafluoroisopropanol as a solvent and contains the following components in a specific ratio: 3.5-4.5g of six-armed star-shaped polylactic acid-caprolactone copolymer macromonomer, 5.5-6.5g of poly-L-lactic acid, 0.8-1.0g of silk fibroin, 45-55mg of sodium heparin, and 6-8mL of PEGylated gold nanorod solution with a concentration of 0.1mg / mL.

[0011] Preferably, the main wavelength of ultraviolet light used in the photocrosslinking process of the scaffold body is 365nm.

[0012] Preferably, the surface of the support body is formed by laser cutting, and the three-dimensional model data of the laser cutting are: diameter 2.5-3.0mm, length 12-15mm, and thickness 60-80μm.

[0013] Preferably, the surface of the stent body further comprises a bioactive modification layer, the bioactive modification layer comprising RGD peptide, chitosan, heparin sodium, vascular endothelial growth factor and stromal cell-derived factor-1α.

[0014] Secondly, this application discloses a method for preparing a biodegradable coronary stent of polylactic acid-caprolactone copolymer (PLCL) with a hybrid interpenetrating network structure, comprising the following steps: Step 1: Synthesis of the six-armed star-shaped polylactic acid-caprolactone copolymer (PLCL) macromonomer (1) Feeding and reaction: Under nitrogen protection, 0.21 g of dipentaerythritol, 60.5 g of L-lactide, 20.5 g of ε-caprolactone and 81 mg of stannous octoate were added to a 500 mL polymerization reactor. Under continuous nitrogen protection, the reaction system was placed in an oil bath at 130 °C and mechanically stirred at 200-250 rpm for 24 h. In this process, dipentaerythritol was used as a six-armed star core. Through the ring-opening polymerization reaction catalyzed by stannous octoate, a PLCL copolymer with a specific chain segment structure was synthesized, which provides a basic polymer skeleton for subsequent functionalization.

[0015] (2) End-group functionalization: After the reaction was completed, the product was dissolved in 400 mL of tetrahydrofuran; under the conditions of 0℃ ice bath and 400-450 rpm stirring, 50 mL of tetrahydrofuran mixture containing cinnamyl chloride and triethylamine was slowly added dropwise through a constant pressure dropping funnel. The addition time was controlled at 1 h. After the addition was completed, the ice bath was removed and the reaction was continued at 25℃ for 5 h. During this process, the photosensitive cinnamyl group was precisely introduced into the end of the star-shaped PLCL polymer through the acylation reaction, giving it the ability to crosslink and cure under ultraviolet light irradiation.

[0016] (3) Purification and collection: The reaction solution was slowly poured into 4 times the volume of methanol that had been pre-frozen to 0°C to precipitate. After filtration through a filter cloth, the white fibrous precipitate was collected. The crude product was then transferred to a supercritical CO2 extraction device for dynamic extraction for 4 hours. The purified product was placed in a vacuum oven at 50°C and dried for 48-50 hours to finally obtain a white flocculent six-armed star-shaped PLCL macromonomer. This purification process removed most of the impurities through low-temperature precipitation and then used supercritical CO2 fluid to efficiently extract the residual solvent and small molecules, ultimately obtaining a high-purity macromonomer with complete photosensitive groups.

[0017] Step 2: Preparation of composite nanofiber tube embryos Weigh out 3.5-4.5g of the six-armed star-shaped PLCL macromonomer prepared in the first step, 5.5-6.5g of poly-L-lactic acid (PLLA), 0.8-1.0g of silk fibroin, 45-55mg of heparin sodium, and 6-8mL of [unclear - possibly a specific ingredient or solution]. PEGylated gold nanorod solution (carboxyl-terminated, catalog number AuB14) was mixed with 90-110 mL of hexafluoroisopropanol, sealed, and placed on a magnetic stirrer. The mixture was stirred in the dark at 25°C and 400-500 rpm for 22-24 hours to obtain a blended spinning solution. A uniaxial electrospinning apparatus was constructed, using a 3.0 mm diameter metal mandrel connected to a speed-controlled motor as the receiving device. The distance between the needle tip and the mandrel surface was adjusted to 15 cm. The spinning solution was transferred to a syringe and attached to a syringe pump. The apparatus was turned on, and spinning was performed under stable temperature and humidity to obtain fiber tubes. The fiber tubes, along with the mandrel, were placed in a fume hood at room temperature for 1 hour to initially evaporate the solvent. They were then transferred to a vacuum drying oven and dried at 40°C and -0.1 MPa for 48 hours to completely remove the hexafluoroisopropanol, ultimately obtaining a composite nanofiber tube preform. During this process, the high tensile stress generated by electrospinning induced a high degree of orientation of the polymer molecular chains along the fiber axis. This orientation structure is the key basis for the subsequent formation of a high-strength physical crystalline network. The added PEGylated gold nanorods act as highly efficient heterogeneous nucleating agents in the system. Their uniform dispersion reduces the nucleation energy barrier of polymer crystals, inducing PLLA and PLCL to form finer and more uniform spherulitic structures during subsequent heat treatment, thereby optimizing the perfection of the physical crystallization network. Simultaneously, the carboxyl groups on the surface of the gold nanorods may promote interactions between polymer segments, facilitating the synergistic penetration of the chemical crosslinking network and the physical crystallization network at the microscale. Furthermore, the added PEGylated gold nanorods not only act as heterogeneous nucleating agents to induce the formation of fine and uniform spherulitic structures, but their surface carboxyl end groups can also form weak hydrogen bonds with the terminal hydroxyl groups of polymer segments. This interaction, during photocrosslinking, can assist in connecting different polymer segments, promoting the uniform distribution of the chemical crosslinking network within the physical crystallization micro-intervals, thus truly achieving hybrid interpenetration at the microscale. Silk fibroin enhances biocompatibility and flexibility, promoting cell adhesion; sodium heparin is uniformly dispersed in the fibrous matrix, endowing the scaffold with immediate anticoagulant activity in the initial stage, avoiding platelet adhesion and thrombosis initiation caused by material exposure during early implantation.

[0018] Step 3: Thermal densification, photocrosslinking, and hierarchical crystallization The composite nanofiber tube preform obtained in the second step is transferred into a processing chamber equipped with an ultraviolet light source and precise temperature control. After nitrogen is introduced to replace the air, the temperature is raised to 120°C at 10°C / min under nitrogen protection and held for 10 min for annealing relaxation and thermal densification. In this key step, the characteristic of the temperature being close to the melting point of the polymer is utilized to cause controlled melting and fusion of the surface layer of the electrospun fibers, eliminating the pore defects between the fibers and forming a dense and airtight tube wall structure to meet the stringent requirements of coronary stents for radial support force. Simultaneously, this process eliminates the internal stress during spinning, allowing the molecular chains to relax and rearrange, providing the optimal thermodynamic environment for subsequent crystallization and crosslinking. Maintaining a temperature of 120℃, the tube is uniformly irradiated for 5 minutes using a 365nm UV LED surface light source (irradiance 10 mW / cm²) to achieve photocrosslinking. Immediately after irradiation, a programmed cooling process is executed: first, the temperature is reduced from 120℃ to 90℃ at a rate of 2℃ / min, then from 90℃ to 60℃ at a rate of 0.5℃ / min, and then annealed at this temperature for 30 minutes. Finally, the temperature is slowly cooled to room temperature at a rate of 1℃ / min, all under a nitrogen atmosphere. Nitrogen protection throughout this process prevents material oxidation, and the graded temperature control and annealing relaxation work together to improve crystallization uniformity. The UV photocrosslinking is highly efficient and uniform, ultimately enhancing the mechanical stability and structural integrity of the scaffold.

[0019] Step 4: Laser cutting and shaping followed by cleaning and sterilization The cross-linked crystallized tubing prepared in the third step was fixed on a rotating chuck, and the three-dimensional model data of the stent was imported. A complete stent mesh structure was cut out using a precision fiber laser. After completion, the stent was immersed in sterile ethanol and ultrasonically cleaned at 40°C (100 W power) for 5 minutes. After removal, it was dried in a clean bench to obtain a laser-cut, structurally complete coronary stent embryo. The coronary stent embryo was placed in a medical packaging bag and sealed, and then sterilized by double-sided electron beam irradiation with an absorbed dose of 25-35 kGy. Among these steps, the use of a precision fiber laser can accurately replicate the three-dimensional model of the stent, forming a well-structured and uniformly sized mesh structure. This ensures effective radial support for the blood vessel after stent implantation and provides a structural basis for the uniform loading of the subsequent surface functional coating.

[0020] Step 5: Surface Biofunctionalization and Terminal Treatment (1) The coronary stent embryo prepared in step 4 was placed in an oxygen plasma cleaner and treated at 100 Pa pressure and 100 W power for 60 s. Then it was immediately immersed in a PBS solution of RGD peptide (arginyl-glycyl-aspartic acid) and slowly shaken at 4 °C for 12 h. This treatment introduces active groups on the surface of the stent to enhance wettability through plasma treatment and uses its activation site to achieve covalent grafting of RGD peptide, thereby specifically enhancing the adhesion and spread of vascular endothelial cells and promoting rapid endothelialization.

[0021] (2) Anticoagulation coating: The stent was immersed in 1 mg / mL chitosan solution (dissolved in 1% v / v acetic acid solution) and 2 mg / mL heparin sodium solution (dissolved in deionized water) for 5 min each time, and gently rinsed three times with acetic acid-sodium acetate buffer solution at pH=5 in between. This process was repeated 5 times. This process adopts electrostatic layer-by-layer self-assembly technology, through the alternating adsorption of chitosan (positively charged) and heparin sodium (negatively charged) on the surface of the stent to build a stable multilayer film, which aims to achieve controllable loading and long-term release of heparin molecules, thereby significantly improving the anticoagulation performance and blood compatibility of the stent.

[0022] (3) Growth factor coating: The scaffold was immersed in a PBS solution containing 10 μg / mL vascular endothelial growth factor and 10 μg / mL stromal cell-derived factor-1α and adsorbed at 4°C for 2 h. This step fixes VEGF and SDF-1α on the scaffold surface by physical adsorption to synergistically promote the directional migration, proliferation and homing of vascular endothelial cells and stem cells, thereby actively accelerating the regeneration and repair of vascular tissue after implantation.

[0023] (4) Packaging: Place the functionalized stent in a medical packaging bag and seal it.

[0024] Preferably, the stent prepared by the present invention has excellent radial support force, slow degradation rate and good biocompatibility, and is suitable for interventional treatment of coronary artery disease.

[0025] Preferably, the 50 mL tetrahydrofuran mixture in the first step of the end-group functionalization of the present invention contains 1.11 g of cinnamyl chloride and 0.61 g of triethylamine.

[0026] Preferably, the temperature during the extraction process in the first step of purification and collection described in this invention is 40°C and the pressure is 15 MPa.

[0027] Preferably, the first step of purification and collection in this invention uses a 200-mesh filter cloth to collect the white fibrous precipitate.

[0028] Preferably, the concentration of the PEGylated gold nanorod solution in the two steps of the present invention is 0.1 mg / mL.

[0029] Preferably, the two-step electrospinning of the present invention is set with a propulsion flow rate of 1.0-1.5 mL / h, an applied voltage of 12-18 kV, and a spindle speed of 800-1000 rpm.

[0030] Preferably, the composite nanofiber tube obtained in the second step of the present invention has a thickness of 120±10μm.

[0031] Preferably, the three-dimensional model data of the stent in the fourth step of the present invention are: diameter 2.5-3.0mm, length 12-15mm, and thickness 60-80μm.

[0032] Preferably, the precision fiber laser used in the fourth step of the present invention has a power of 1.2W and a scanning speed of 5mm / s.

[0033] Preferably, the concentration of RGD peptide in the fifth step of the present invention is 1 mg / mL.

[0034] The present invention has the following beneficial effects: Compared with the prior art, the present invention has at least the following beneficial effects: 1) By interpenetrating and synergistically combining chemical cross-linking networks and physical crystallization networks, the structural stability and radial support retention capacity of the scaffold body are improved while maintaining the material's processability, thereby reducing the risk of early collapse. 2) Controlled annealing and programmed crystallization control can regulate crystal morphology and chain segment movement, thereby making degradation behavior more predictable and process repeatable; 3) By processing the surface microstructure and modifying its bioactivity, the space for endothelialization guidance and blood compatibility improvement can be expanded without sacrificing the manufacturability of the scaffold itself; 4) This invention innovatively employs a composite molding process of "electrospinning orientation-thermal densification fusion". Unlike the traditional approach of electrospinning which only retains the porous structure, this invention utilizes the molecular chain orientation advantage of electrospinning to enhance the material's potential. Subsequently, a thermal densification process is used to eliminate pores to obtain a dense scaffold body, solving the problem of insufficient support in porous materials. At the same time, the special surface morphology formed by fiber fusion still maintains good biocompatibility. 5) The gold nanorods introduced in this invention are not used as mechanical fillers, but as crystallization regulators and network synergists. Trace amounts of gold nanorods refine grains through heterogeneous nucleation, promoting uniform interpenetration of the chemical and physical dual networks, achieving a significant leap in performance with extremely low addition amounts. Attached Figure Description

[0035] Figure 1 This is a comparison of the number of lymphocytes in the stent implantation area of ​​miniature pigs in Examples 1-3 and Comparative Examples 7-10.

[0036] Figure 2 This is a comparison chart of inflammation scores in miniature pigs between Examples 1-3 and Comparative Examples 7-10.

[0037] Figure 3 This is a comparison chart of the percentage narrowing of the stent implantation area in miniature pigs in Examples 1-3 and Comparative Examples 7-10. Detailed Implementation

[0038] To make the advantages of this application clearer, the following detailed description is provided in conjunction with embodiments. It should be understood that the specific embodiments described herein are merely illustrative and not intended to limit the scope of this application.

[0039] In this embodiment, poly-L-lactic acid, ε-caprolactone, dipentaerythritol, and chitosan were purchased from Hubei Langbowan Biopharmaceutical Co., Ltd.; vascular endothelial growth factor was purchased from Shenzhen Zike Biotechnology Co., Ltd.; stromal cell-derived factor-1α was purchased from Shanghai Yaji Biotechnology Co., Ltd.; and PEGylated gold nanorods (carboxyl-terminated, catalog number AuB14) were purchased from Nanjing Dongna Biotechnology Co., Ltd.

[0040] Example 1 S1. Synthesis of the six-armed star-shaped polylactic acid-caprolactone copolymer (PLCL) macromonomer: In a 500 mL polymerization reactor that had undergone three vacuum-nitrogen-purging cycles, 0.21 g of dipentaerythritol, 60.5 g of L-lactide, and 20.5 g of [unspecified ingredient] were precisely added. ε-caprolactone and 81 mg stannous octoate were reacted at 130 °C in an oil bath with mechanical stirring at 200 rpm for 24 h under continuous nitrogen protection. After the reaction, the product was dissolved in 400 mL of tetrahydrofuran. A mixture of 1.11 g cinnamyl chloride and 0.61 g triethylamine was added dropwise through a constant pressure dropping funnel over 1 h in an ice bath at 0 °C with stirring at 400 rpm. After removing the ice bath, the reaction was continued at 25 °C for 5 h. The reaction solution was slowly poured into 4 times its volume of pre-cooled methanol at 0 °C to precipitate the product. The white fibrous precipitate was collected using a 200-mesh filter cloth. The crude product was transferred to a supercritical CO2 extraction device and dynamically extracted for 4 h at 40 °C and 15 MPa. After purification, the product was dried in a vacuum oven at 50 °C for 48 h to finally obtain a white flocculent six-armed star-shaped PLCL macromonomer. S2. Preparation of composite nanofiber tube embryos: Weigh 3.5g of the six-armed star-shaped PLCL macromonomer prepared in step S1, 5.5g of poly-L-lactic acid (PLLA), 0.8g of silk fibroin, 45mg of heparin sodium, and 6mL of... PEGylated gold nanorod solution (carboxyl-terminated, concentration 0.1 mg / mL, catalog number AuB14) was mixed with 90 mL of hexafluoroisopropanol, sealed, and placed on a magnetic stirrer. The mixture was stirred in the dark at 25°C and 400 rpm for 22 h to obtain a blended spinning solution. A uniaxial electrospinning apparatus was constructed, using a 3.0 mm diameter metal mandrel connected to a speed-controlled motor as the receiving device. The distance between the needle tip and the mandrel surface was adjusted to 15 cm. The spinning solution was transferred into a syringe and installed on a syringe pump. The flow rate was set to 1.0 mL / h, the applied voltage to 12 kV, and the mandrel speed to 800 rpm. The apparatus was turned on and spun under stable temperature and humidity until a 110 μm thick fiber tube was formed on the mandrel surface. The fiber tube, along with the mandrel, was placed in a fume hood at room temperature for 1 h to initially evaporate the solvent. Then, it was transferred to a vacuum drying oven and dried at 40°C and -0.1 MPa for 48 h to completely remove the hexafluoroisopropanol, finally obtaining a composite nanofiber tube preform. S3. Thermal densification, photocrosslinking, and graded crystallization: The composite nanofiber tube preform obtained in step S2 is transferred into a processing chamber equipped with an ultraviolet light source and precise temperature control. After replacing the air with nitrogen, the temperature is raised to 120°C at 10°C / min and held for 10 min under nitrogen protection for annealing relaxation and thermal densification. The temperature is maintained at 120°C, and the tube is irradiated uniformly in all directions for 5 min using an ultraviolet LED surface light source with a main wavelength of 365nm (irradiance of 10mW / cm²) to complete photocrosslinking. After irradiation, the tube is immediately subjected to programmed cooling: first, the temperature is reduced from 120°C to 90°C at 2°C / min, then from 90°C to 60°C at 0.5°C / min, and annealed at this temperature for 30 min. Finally, the temperature is slowly cooled to room temperature at a rate of 1°C / min. The entire process is carried out under a nitrogen atmosphere. S4. Laser Cutting and Cleaning: The cross-linked crystallized tubing prepared in step S3 is fixed on a rotating chuck. The three-dimensional model data of the stent (diameter 2.5 mm, length 12 mm, thickness 60 μm) is imported, and a complete stent mesh structure is directly cut out using a precision fiber laser with a power of 1.2 W and a scanning speed of 5 mm / s. Subsequently, the stent is immersed in sterile ethanol and cleaned at 40 °C and 100 W power for 5 min. After removal, it is dried in an ultra-clean workbench to obtain a laser-cut and structurally complete coronary stent embryo. The coronary stent embryo is placed in a medical packaging bag and sealed, and then sterilized by double-sided electron beam irradiation with an absorbed dose of 25 kGy. S5. Surface Biofunctionalization and Terminal Treatment: The coronary stent embryos prepared in step S4 were placed in an oxygen plasma cleaner and treated at 100 Pa and 100 W for 60 s. Immediately after treatment, they were immersed in a 1 mg / mL RGD peptide PBS solution and incubated with slow shaking at 4 °C for 12 h. Subsequently, they were immersed sequentially in a 1 mg / mL chitosan solution dissolved in 1% v / v acetic acid and a 2 mg / mL heparin sodium solution dissolved in deionized water, each for 5 min. The stents were gently rinsed three times with an acetate-sodium acetate buffer solution at pH 5. This process was repeated 5 times to complete the anticoagulant coating. Next, the stents were immersed in a PBS solution containing 10 μg / mL vascular endothelial growth factor and 10 μg / mL stromal cell-derived factor-1α and coated with growth factors at 4 °C for 2 h. Finally, the functionalized stents were sealed in a medical packaging bag.

[0041] Example 2 S1. Synthesis of the six-armed star-shaped polylactic acid-caprolactone copolymer (PLCL) macromonomer: In a 500 mL polymerization reactor that had undergone three vacuum-nitrogen-purging cycles, 0.21 g of dipentaerythritol, 60.5 g of L-lactide, and 20.5 g of [unspecified ingredient] were precisely added. ε-caprolactone and 81 mg stannous octoate were reacted at 130 °C in an oil bath with mechanical stirring at 230 rpm for 24 h under continuous nitrogen protection. After the reaction, the product was dissolved in 400 mL of tetrahydrofuran. A mixture of 1.11 g cinnamyl chloride and 0.61 g triethylamine was added dropwise through a constant pressure dropping funnel over 1 h in an ice bath at 0 °C with stirring at 430 rpm. After removing the ice bath, the reaction was continued at 25 °C for 5 h. The reaction solution was slowly poured into 4 times its volume of pre-cooled methanol at 0 °C to precipitate the product. The white fibrous precipitate was collected using a 200-mesh filter cloth. The crude product was transferred to a supercritical CO2 extraction device and dynamically extracted for 4 h at 40 °C and 15 MPa. After purification, the product was dried in a vacuum oven at 50 °C for 49 h to finally obtain a white flocculent six-armed star-shaped PLCL macromonomer. S2. Preparation of composite nanofiber tube embryos: Weigh 4g of the six-armed star-shaped PLCL macromonomer prepared in step S1, 6g of poly-L-lactic acid (PLLA), 0.9g of silk fibroin, 50mg of heparin sodium, and 7mL of... PEGylated gold nanorod solution (carboxyl-terminated, concentration 0.1 mg / mL, catalog number AuB14) was mixed with 100 mL of hexafluoroisopropanol, sealed, and placed on a magnetic stirrer. The mixture was stirred in the dark at 25°C and 450 rpm for 23 h to obtain a blended spinning solution. A uniaxial electrospinning apparatus was constructed, using a 3.0 mm diameter metal mandrel connected to a speed-controlled motor as the receiving device. The distance between the needle tip and the mandrel surface was adjusted to 15 cm. The spinning solution was transferred into a syringe and installed on a syringe pump. The flow rate was set to 1.3 mL / h, the applied voltage to 15 kV, and the mandrel speed to 900 rpm. The apparatus was turned on and spun under stable temperature and humidity until a 120 μm thick fiber tube was formed on the mandrel surface. The fiber tube, along with the mandrel, was placed in a fume hood at room temperature for 1 h to initially evaporate the solvent. Then, it was transferred to a vacuum drying oven and dried at 40°C and -0.1 MPa for 48 h to completely remove the hexafluoroisopropanol, finally obtaining a composite nanofiber tube preform. S3. Thermal densification, photocrosslinking, and graded crystallization: The composite nanofiber tube preform obtained in step S2 is transferred into a processing chamber equipped with an ultraviolet light source and precise temperature control. After replacing the air with nitrogen, the temperature is raised to 120°C at 10°C / min and held for 10 min under nitrogen protection for annealing relaxation and thermal densification. The temperature is maintained at 120°C, and the tube is irradiated uniformly in all directions for 5 min using an ultraviolet LED surface light source with a main wavelength of 365nm (irradiance of 10mW / cm²) to complete photocrosslinking. After irradiation, the tube is immediately subjected to programmed cooling: first, the temperature is reduced from 120°C to 90°C at 2°C / min, then from 90°C to 60°C at 0.5°C / min, and annealed at this temperature for 30 min. Finally, the temperature is slowly cooled to room temperature at a rate of 1°C / min. The entire process is carried out under a nitrogen atmosphere. S4. Laser Cutting and Cleaning: The cross-linked crystallized tubing prepared in step S3 is fixed on a rotating chuck. The three-dimensional model data of the stent (diameter 2.7 mm, length 14 mm, thickness 70 μm) is imported, and a complete stent mesh structure is directly cut using a precision fiber laser with a power of 1.2 W and a scanning speed of 5 mm / s. Subsequently, the stent is immersed in sterile ethanol and cleaned at 40 °C and 100 W power for 5 min. After removal, it is dried in a clean bench to obtain a laser-cut and structurally complete coronary stent embryo. The coronary stent embryo is placed in a medical packaging bag and sealed, and then sterilized by double-sided electron beam irradiation with an absorbed dose of 30 kGy. S5. Surface Biofunctionalization and Terminal Treatment: The coronary stent embryos prepared in step S4 were placed in an oxygen plasma cleaner and treated at 100 Pa and 100 W for 60 s. Immediately after treatment, they were immersed in a 1 mg / mL RGD peptide PBS solution and incubated with slow shaking at 4 °C for 12 h. Subsequently, they were immersed sequentially in a 1 mg / mL chitosan solution dissolved in 1% v / v acetic acid and a 2 mg / mL heparin sodium solution dissolved in deionized water, each for 5 min. The stents were gently rinsed three times with an acetate-sodium acetate buffer solution at pH 5. This process was repeated 5 times to complete the anticoagulant coating. Next, the stents were immersed in a PBS solution containing 10 μg / mL vascular endothelial growth factor and 10 μg / mL stromal cell-derived factor-1α and coated with growth factors at 4 °C for 2 h. Finally, the functionalized stents were sealed in a medical packaging bag.

[0042] Example 3 S1. Synthesis of the six-armed star-shaped polylactic acid-caprolactone copolymer (PLCL) macromonomer: In a 500 mL polymerization reactor that had undergone three vacuum-nitrogen-purging cycles, 0.21 g of dipentaerythritol, 60.5 g of L-lactide, and 20.5 g of [unspecified ingredient] were precisely added. ε-caprolactone and 81 mg stannous octoate were reacted at 130 °C in an oil bath with mechanical stirring at 250 rpm for 24 h under continuous nitrogen protection. After the reaction, the product was dissolved in 400 mL of tetrahydrofuran. A mixture of 1.11 g cinnamyl chloride and 0.61 g triethylamine was added dropwise through a constant pressure dropping funnel over 1 h in an ice bath at 0 °C with stirring at 450 rpm. After removing the ice bath, the reaction was continued at 25 °C for 5 h. The reaction solution was slowly poured into 4 times its volume of pre-cooled methanol at 0 °C to precipitate the product. The white fibrous precipitate was collected using a 200-mesh filter cloth. The crude product was transferred to a supercritical CO2 extraction device and dynamically extracted for 4 h at 40 °C and 15 MPa. After purification, the product was dried in a vacuum oven at 50 °C for 50 h to finally obtain a white flocculent six-armed star-shaped PLCL macromonomer. S2. Preparation of composite nanofiber tube embryos: Weigh 4.5g of the six-armed star-shaped PLCL macromonomer prepared in step S1, 6.5g of poly-L-lactic acid (PLLA), 1.0g of silk fibroin, 55mg of heparin sodium, and 8mL of... PEGylated gold nanorod solution (carboxyl-terminated, concentration 0.1 mg / mL, catalog number AuB14) was mixed with 110 mL of hexafluoroisopropanol, sealed, and placed on a magnetic stirrer. The mixture was stirred in the dark at 25°C and 500 rpm for 24 h to obtain a blended spinning solution. A uniaxial electrospinning apparatus was constructed, using a 3.0 mm diameter metal mandrel connected to a speed-controlled motor as the receiving device. The distance between the needle tip and the mandrel surface was adjusted to 15 cm. The spinning solution was transferred into a syringe and installed on a syringe pump. The flow rate was set to 1.5 mL / h, the applied voltage to 18 kV, and the mandrel speed to 1000 rpm. The apparatus was turned on and spun under stable temperature and humidity until a 130 μm thick fiber tube was formed on the mandrel surface. The fiber tube, along with the mandrel, was placed in a fume hood at room temperature for 1 h to initially evaporate the solvent. Then, it was transferred to a vacuum drying oven and dried at 40°C and -0.1 MPa for 48 h to completely remove the hexafluoroisopropanol, finally obtaining a composite nanofiber tube preform. S3. Thermal densification, photocrosslinking, and graded crystallization: The composite nanofiber tube preform obtained in step S2 is transferred into a processing chamber equipped with an ultraviolet light source and precise temperature control. After replacing the air with nitrogen, the temperature is raised to 120°C at 10°C / min and held for 10 min under nitrogen protection for annealing relaxation and thermal densification. The temperature is maintained at 120°C, and the tube is irradiated uniformly in all directions for 5 min using an ultraviolet LED surface light source with a main wavelength of 365nm (irradiance of 10mW / cm²) to complete photocrosslinking. After irradiation, the tube is immediately subjected to programmed cooling: first, the temperature is reduced from 120°C to 90°C at 2°C / min, then from 90°C to 60°C at 0.5°C / min, and annealed at this temperature for 30 min. Finally, the temperature is slowly cooled to room temperature at a rate of 1°C / min. The entire process is carried out under a nitrogen atmosphere. S4. Laser Cutting and Cleaning: The cross-linked crystallized tubing prepared in step S3 is fixed on a rotating chuck. The three-dimensional model data of the stent (diameter 3.0 mm, length 15 mm, thickness 80 μm) is imported, and a complete stent mesh structure is directly cut using a precision fiber laser with a power of 1.2 W and a scanning speed of 5 mm / s. Subsequently, the stent is immersed in sterile ethanol and cleaned at 40 °C and 100 W power for 5 min. After removal, it is dried in a clean bench to obtain a laser-cut and structurally complete coronary stent embryo. The coronary stent embryo is placed in a medical packaging bag and sealed, and then sterilized by double-sided electron beam irradiation with an absorbed dose of 35 kGy. S5. Surface Biofunctionalization and Terminal Treatment: The coronary stent embryos prepared in step S4 were placed in an oxygen plasma cleaner and treated at 100 Pa and 100 W for 60 s. Immediately after treatment, they were immersed in a 1 mg / mL RGD peptide PBS solution and incubated with slow shaking at 4 °C for 12 h. Subsequently, they were immersed sequentially in a 1 mg / mL chitosan solution dissolved in 1% v / v acetic acid and a 2 mg / mL heparin sodium solution dissolved in deionized water, each for 5 min. The stents were gently rinsed three times with an acetate-sodium acetate buffer solution at pH 5. This process was repeated 5 times to complete the anticoagulant coating. Next, the stents were immersed in a PBS solution containing 10 μg / mL vascular endothelial growth factor and 10 μg / mL stromal cell-derived factor-1α and coated with growth factors at 4 °C for 2 h. Finally, the functionalized stents were sealed in a medical packaging bag.

[0043] Comparative Example 1: The six-armed star-shaped PLCL macromonomer in step S1 was replaced with ordinary PLCL. The specific preparation method was as follows: In a 500mL polymerization reactor that had undergone three vacuum-nitrogen cycles, 0.05g of 1-dodecyl alcohol (monofunctional initiator), 60.5g of L-lactide, 20.5g of ε-caprolactone, and 81mg of stannous octoate were precisely added. Under continuous nitrogen protection, the reaction was carried out in an oil bath at 130℃ with mechanical stirring at 230rpm for 24h. After the reaction, the product was dissolved in 400mL of tetrahydrofuran, and then slowly poured into 4 times the volume of pre-cooled methanol at 0℃ to precipitate. The white precipitate was collected using a 200-mesh filter cloth. The crude product was transferred to a supercritical CO2 extraction device and dynamically extracted for 4h at 40℃ and 15MPa. After purification, it was dried in a vacuum oven at 50℃ for 49h to finally obtain a linear polylactic acid-caprolactone copolymer (PLCL). The remaining steps were the same as in Example 2.

[0044] Comparative Example 2: No PEGylated gold nanorods were added in step S2, and the remaining steps were the same as in Example 2.

[0045] Comparative Example 3: The six-armed star-shaped PLCL macromonomer in step S1 was replaced with ordinary PLCL (ordinary PLCL is the same as in Comparative Example 1). At the same time, PEGylated gold nanorods were not added in step S2. The remaining steps were the same as in Example 2.

[0046] Comparative Example 4: The photocrosslinking in step S3 was omitted, that is, after annealing and relaxation, the step crystallization process was directly carried out, and the remaining steps were the same as in Example 2.

[0047] Comparative Example 5: Step S3 was skipped, i.e. after photocrosslinking was completed, the sample was directly cooled to room temperature naturally, and the remaining steps were the same as in Example 2.

[0048] Comparative Example 6: The annealing relaxation in step S3 was omitted, that is, the composite nanofiber tube was directly heated to 120°C for photocrosslinking, and the remaining steps were the same as in Example 2.

[0049] Comparative Example 7: The RGD peptide grafting in step S5 was omitted, i.e., the anticoagulant coating was applied directly after plasma treatment, and the remaining steps were the same as in Example 2.

[0050] Comparative Example 8: The anticoagulant coating in step S5 was omitted, that is, after the RGD peptide grafting was completed, the growth factor coating was applied directly, and the remaining steps were the same as in Example 2.

[0051] Comparative Example 9: The growth factor coating in step S5 was omitted, that is, after the anticoagulant coating was completed, the product was directly sealed and packaged, and the remaining steps were the same as in Example 2.

[0052] Comparative Example 10: The simplified surface modification process only involves plasma activation, i.e., after oxygen plasma treatment, without RGD peptide grafting, anticoagulant coating, and growth factor coating, and is directly sealed and packaged. The remaining steps are the same as in Example 2.

[0053] Experiment 1: Mechanical Property Testing Samples: Six stents were taken from each of the following groups: Examples 1-3 and Comparative Examples 1-6.

[0054] Tensile property testing: The tensile strength and elongation at break of the support were determined using a GL028 electronic universal testing machine (purchased from Wuhan Guoliang Instrument Co., Ltd.). The strain rate was 10 mm / min, and the test was repeated 3 times.

[0055] Radial support force testing: Referring to ISO 25539-2:2020 "Cardiovascular implants - Stents - Part 2: Stents", the strain rate was 0.2 mm / min. The radial support force required when the stent was compressed to 50% of its original diameter was measured. Each sample was measured three times. The experimental results are shown in Table 1.

[0056] Elongation at break (%) = (L2 - L1) ÷ L1 × 100% in: L1 is the original length of the gauge length of the specimen (i.e., the effective test section length marked on the specimen before the tensile test). L2 is the length of the gauge length when the specimen breaks (the actual length of the gauge length after being stretched to break).

[0057] Table 1 Table 1 compares the mechanical properties of the stents in Examples 1-3 and Comparative Examples 1-6. The table shows that the tensile strength, elongation at break, and radial compressive strength of the stents in Examples 1-3 are significantly better than those in Comparative Examples 1-6, and the differences between Examples 1-3 are not significant. Specifically, Example 2 is representative, with its tensile strength increasing by 90.26% compared to Comparative Example 1, 19.23% compared to Comparative Example 2, 165.20% compared to Comparative Example 3, 67.06% compared to Comparative Example 4, 42.26% compared to Comparative Example 5, and 54.35% compared to Comparative Example 6; and its elongation at break increasing by 68.1% compared to Comparative Example 1, 16.6% compared to Comparative Example 2, 121.06% compared to Comparative Example 3, 47.37% compared to Comparative Example 4, and [missing data - likely a percentage increase in elongation at break]. The radial compressive strength was 27.11% higher than that of Comparative Example 1, 77.52% higher than that of Comparative Example 2, 142.31% higher than that of Comparative Example 3, 75.81% higher than that of Comparative Example 4, 41.04% higher than that of Comparative Example 5, and 61.54% higher than that of Comparative Example 6. These results fully demonstrate that this invention, by combining a six-armed star-shaped PLCL macromolecular monomer with trace amounts of gold nanorods and utilizing the heterogeneous nucleation (crystallization promotion) effect of gold nanorods, significantly optimizes the physical crystalline network morphology (such as crystallinity and crystal form distribution) of PLLA and PLCL, making the physical network and the chemical network formed by photocrosslinking more tightly integrated. This synergistic effect of the microstructure (rather than the direct load-bearing effect of the gold nanorods themselves), combined with photocrosslinking and hierarchical crystallization processes, overcomes the defects of a single network, significantly improving the mechanical load-bearing capacity, flexibility, and radial support stability of the scaffold.

[0058] It is worth noting that Comparative Example 6, due to the omission of the annealing relaxation step, exhibits significantly inferior overall mechanical properties compared to Examples 1 and 5. Specifically, the elongation at break of Comparative Example 6 (118.45%) decreased by approximately 44% compared to Example 2, indicating that unrelieved internal stress led to severe brittleness of the material. Simultaneously, its radial support force (2.34 N / mm) was also lower than that of Comparative Example 5, demonstrating that annealing relaxation is crucial for eliminating interfiber porosity and achieving thermal densification. This comparison powerfully demonstrates that annealing relaxation is not only for stress relief but also a necessary prerequisite for constructing dense, high-strength tubular structures.

[0059] Experiment 2: Degradation Performance Test Samples: Six stents were taken from each of the following groups: Examples 1-3 and Comparative Examples 5-6.

[0060] Pretreatment: Each support was placed in a vacuum drying oven at 85℃ and dried for 12 hours. After removal, it was accurately weighed using an analytical balance with an accuracy of 0.01 mg and recorded as the initial mass (m1).

[0061] Accelerated degradation experiment: The pretreated scaffold was immersed in a sealed centrifuge tube containing 10 mL of 0.1 mol / L PBS solution. The pH of this solution was 7.4, and it contained 10 mL of PBS solution. 7 U / L of lysozyme; after sealing, place the centrifuge tubes in a constant temperature shaking water bath at 37℃ and a shaking rate of 130r / min for accelerated degradation. To avoid the accumulation of degradation products interfering with the test results, replace the PBS solution in the centrifuge tubes with fresh PBS every 7 days.

[0062] The support was removed on days 14, 21, and 28 of degradation, rinsed thoroughly with double-distilled water, and the surface moisture was absorbed by filter paper. The support was then vacuum dried at 60°C for 12 hours until constant weight, which was recorded as the remaining mass (m2). The degradation rate was then calculated. The experimental results are shown in Table 2.

[0063] Degradation rate (%) = (m1 - m2) / m1 × 100% Table 2 Table 2 compares the degradation rates of Examples 1-3 and Comparative Examples 5-6. As shown in the table, the degradation rates of the stents in Examples 1-3 on days 14, 21, and 28 were significantly lower than those in Comparative Examples 5-6, but the differences were not substantial. This fully demonstrates that the degradation process of Examples 1-3 was slow and stable, exhibiting excellent controllability.

[0064] In contrast, Comparative Example 5 omitted the graded crystallization process and completed the crystallization process only through natural cooling. As a result, a dense and uniform ordered crystalline network could not be formed inside the material. The structure was loose and the grains were arranged randomly, making it easier for water and degradation media to penetrate into the material and thus triggering rapid degradation. Comparative Example 6 skipped the initial annealing and relaxation step. The composite nanofiber tubes contained residual internal stress generated during the spinning process. The molecular chains were arranged randomly and there were local structural defects. These defects became the preferential attack sites for water molecules and enzymes in the early stage of degradation. This not only led to a significant acceleration of the degradation rate, but also caused uneven degradation process and extremely poor stability.

[0065] Experiment 3: Biocompatibility Testing Animals: Forty healthy Tibetan miniature pigs, half male and half female, weighing 25-30 kg, were selected and provided by the Experimental Animal Center of Southern Medical University. They were given free access to food and water. All miniature pigs were acclimatized for 7 days before the experiment.

[0066] Grouping: 40 miniature pigs were randomly divided into Example 1-3 groups and Comparative Examples 7-10 groups, with 10 pigs in each group.

[0067] After 7 days of acclimatization, all miniature pigs were administered clopidogrel 10 mg / (kg·d) (purchased from Hangzhou Sanofi Sandebao Minsheng Pharmaceutical Co., Ltd.) and aspirin 5 mg / (kg·d) (purchased from Beijing Huaxi United Technology Development Co., Ltd.) by gavage for 5 consecutive days to prevent early postoperative in-stent thrombosis. After gavage, the pigs were fasted for 12 hours and anesthetized with sodium pentobarbital. Following anesthesia, the right femoral artery was dissected layer by layer, and a 6F arterial sheath was inserted according to interventional surgery standards. Heparin 6000U / pig (purchased from Shanghai Maibo Pharmaceutical Technology Co., Ltd.) was injected through the sheath for anticoagulation to prevent intravascular coagulation. A corresponding stent was precisely implanted in the proximal-mid segment of the left anterior descending coronary artery. The incision was sutured layer by layer postoperatively, and all miniature pigs were administered clopidogrel by gavage daily. Clopidogrel and aspirin (dose as described above) were administered intramuscularly daily at 400,000 U / (kg·d) of penicillin (purchased from Beijing Lebo Biotechnology Co., Ltd.) for anti-infection treatment to avoid interference with experimental results from surgical site infection. This was continued for 4 weeks via gavage. After the last administration, miniature pigs were anesthetized and euthanized, and intact heart tissue was immediately removed. The tissue was fixed in 4% neutral formaldehyde solution according to histopathological standards. The coronary artery segment at the stent placement site was separated, and excess myocardial tissue around it was removed. After routine paraffin embedding, sections were prepared, dewaxed with xylene, dehydrated with graded ethanol, and stained with hematoxylin-eosin (hematoxylin staining for 5 min → washing with tap water → eosin staining for 3 min → mounting with neutral resin) to prepare 4 μm thick tissue sections for later use.

[0068] Lymphocyte count determination: Under a 400x optical microscope, five non-overlapping fields of view were randomly selected around the scaffold. The number of lymphocytes in each field of view was counted, and the average value was taken as the lymphocyte count of the sample (cells / 400x field of view). The experimental results are as follows: Figure 1 As shown.

[0069] Inflammation score determination: Based on the standards in Table 3, two pathologists observed tissue sections under a 400x microscope in a double-blind manner, and inflammation scores were assigned to the stent placement sites in miniature pigs according to the standards. If the scores from the two pathologists differed, a consensus was reached through joint review. The experimental results are as follows: Figure 2 As shown.

[0070] Table 3. Standards for Inflammation Score Measurement Determination of neonatal endometrial area and percentage of stenosis: Using a computer image analysis system (MediaCybernetics Image Pro Plus 6.0, USA), hematoxylin-eosin stained sections were examined under a microscope at 40x magnification to observe the area surrounding the external elastic lamina and the residual lumen area. The percentage of stenosis was calculated. The experimental results are as follows: Figure 3 As shown.

[0071] Figure 1 This is a comparison of the number of lymphocytes in the stent implantation area of ​​miniature pigs in Examples 1-3 and Comparative Examples 7-10. Figure 2 This is a comparison chart of inflammation scores in miniature pigs from Examples 1-3 and Comparative Examples 7-10. The results show that compared to Comparative Examples 7-10, the number of lymphocytes and the inflammation score in miniature pigs from Examples 1-3 were significantly reduced, and the data from Examples 1-3 were not significantly different. This indicates that Examples 1-3 can effectively reduce the local inflammatory response after stent implantation and significantly improve its biocompatibility.

[0072] Specifically, Comparative Example 7, lacking RGD peptide grafting, had a lack of cell adhesion sites on the stent surface, resulting in delayed endothelial coverage and prolonged exposure of the material to the vascular environment, leading to persistent inflammation. Comparative Example 8, lacking an anticoagulant coating, significantly increased the risk of thrombosis on the stent surface, with the thrombus organization process directly exacerbating the infiltration and aggregation of inflammatory cells. Comparative Example 9, lacking a growth factor coating, failed to accelerate endothelial repair, delaying the healing process and correspondingly prolonging the subclinical inflammatory phase. Comparative Example 10, having only undergone plasma activation and lacking any bioactive components on the surface, triggered the strongest foreign body reaction and thrombosis, resulting in the worst performance across all inflammation-related indicators.

[0073] Figure 3 This chart compares the percentage of stenosis in the implantation area of ​​the stents in Examples 1-3 with that in Comparative Examples 7-10. The results show that the percentage of stenosis in Examples 1-3 was significantly lower than that in Comparative Examples 7-10, and the data fluctuation between the example groups was minimal. This fully demonstrates that the stent design of the examples can effectively reduce the risk of postoperative luminal stenosis and exhibits superior biocompatibility.

[0074] Based on the comparative examples, Comparative Example 7, due to incomplete endothelialization, lost the natural inhibitory effect of intact endothelium on smooth muscle cell proliferation; in Comparative Example 8, the early formation of microthrombi not only directly increased the volume of the intima, but the large amount of growth factors released by activated platelets also strongly stimulated the migration and proliferation of smooth muscle cells; in Comparative Example 9, due to the slow speed of endothelial repair, the inhibitory effect on proliferation was delayed; and in Comparative Example 10, due to severe inflammatory response and thrombosis, the strongest pro-proliferative microenvironment was formed, ultimately leading to the most severe luminal stenosis.

[0075] This invention addresses the core problems of existing biodegradable coronary stents, including insufficient radial support, uncontrolled degradation rate, poor biocompatibility, and high risk of postoperative inflammation and restenosis. Its core research direction is to achieve "synergistic optimization of mechanical support, controllable degradation, and biological function." It innovatively constructs a hybrid interpenetrating structure where chemical cross-linking networks and physical crystallization networks intersect. Through the synergistic ratio of six-armed star-shaped PLCL macromolecules and linear PLLA, combined with electrospinning orientation, thermal densification fusion, photocrosslinking, and hierarchical crystallization processes, along with heterogeneous nucleation regulation using PEGylated gold nanorods, and layered biofunctionalization modification with RGD peptides, heparin sodium, and growth factors, significant technical achievements are ultimately achieved: the stent's mechanical properties are balanced and stable, radial support and structural integrity are maintained for a long time, effectively avoiding the risk of early collapse; the degradation process is stable and controllable, with the rate precisely matched to the vascular repair process, avoiding local irritation caused by excessive degradation products; excellent biocompatibility significantly reduces post-implantation inflammation, promotes rapid vascular endothelial repair, and reduces the risk of postoperative restenosis. This invention is not a localized optimization of a single performance, but rather achieves synergistic improvement and breakthrough balance in multiple core dimensions such as mechanical support stability, controllable degradation behavior, biocompatibility, and clinical treatment adaptability, providing a more clinically valuable solution for interventional treatment of coronary artery disease.

Claims

1. A biodegradable coronary stent, characterized in that, The scaffold includes a scaffold body formed of a polymer composite material; the polymer composite material comprises: linear polylactic acid PLLA and star-shaped polylactic acid-caprolactone copolymer PLCL macromonomer with photocrosslinked end groups; The support body also has: (1) A chemically cross-linked network formed by the photocrosslinking end groups under light irradiation; and (2) A physical crystallization network formed by the hierarchical crystallization of the PLLA and the PLCL; Furthermore, the chemical cross-linked network and the physical crystallization network interpenetrate each other within the scaffold body to form a hybrid interpenetrating network structure.

2. The biodegradable coronary stent according to claim 1, characterized in that, The photocrosslinking end group is cinnamyl.

3. The biodegradable coronary stent according to claim 1, characterized in that, The star-shaped polylactic acid-caprolactone copolymer (PLCL) macromonomer has a 6-arm star structure.

4. The biodegradable coronary stent according to claim 3, characterized in that, The mass ratio of L-lactide to ε-caprolactone in the synthesis of the six-armed star-shaped PLCL macromonomer is 2.8-3.1:

1.

5. The biodegradable coronary stent according to claim 4, characterized in that, The mass fraction of PLLA in the polymer composite material is 30% to 80%; during the synthesis of the six-armed star-shaped PLCL macromonomer, supercritical CO2 extraction is used for purification and collection.

6. The biodegradable coronary stent according to claim 1, characterized in that, The scaffold is a tubular structure, and its body is composed of composite nanofiber tubes. The composite nanofiber tubes are prepared by uniaxial electrospinning and gradient drying of the blended spinning solution to form a green embryo, which is then subjected to thermal densification, photocrosslinking and graded crystallization treatment. The thermal densification is the annealing process before photocrosslinking. The blended spinning solution uses hexafluoroisopropanol as a solvent and contains the following components in a specific ratio: 3.5-4.5g of six-arm star-shaped polylactic acid-caprolactone copolymer macromonomer, 5.5-6.5g of poly-L-lactic acid, 0.8-1.0g of silk fibroin, 45-55mg of sodium heparin and 6-8mL of PEGylated gold nanorod solution with a concentration of 0.1mg / mL.

7. The biodegradable coronary stent according to claim 6, characterized in that, The main wavelength of ultraviolet light used in the photocrosslinking process of the scaffold body is 365nm.

8. The biodegradable coronary stent according to claim 7, characterized in that, The surface of the support body is formed by laser cutting. The three-dimensional model data of the laser cutting are: diameter 2.5-3.0mm, length 12-15mm, and thickness 60-80μm.

9. The biodegradable coronary stent according to claim 8, characterized in that, The surface of the scaffold body further includes a bioactive modification layer, which contains RGD peptide, chitosan, sodium heparin, vascular endothelial growth factor, and stromal cell-derived factor-1α.

10. The method for preparing the biodegradable coronary stent according to any one of claims 1-9, characterized in that, Includes the following steps: Step 1: Synthesis of the six-armed star-shaped polylactic acid-caprolactone copolymer (PLCL) macromonomer (1) Under nitrogen protection, 0.21 g of dipentaerythritol, 60.5 g of L-lactide, 20.5 g of ε-caprolactone and 81 mg of stannous octoate were added to a 500 mL polymerization reactor. Under continuous nitrogen protection, the reaction system was placed in an oil bath at 130 °C and mechanically stirred at 200-250 rpm for 24 h. (2) Terminal functionalization: After the reaction was completed, the product was dissolved in 400 mL of tetrahydrofuran; under the conditions of 0℃ ice bath and 400-450 rpm stirring, 50 mL of tetrahydrofuran mixture containing 1.11 g of cinnamyl chloride and 0.61 g of triethylamine was slowly added dropwise through a constant pressure dropping funnel. The addition time was controlled at 1 h. After the addition was completed, the ice bath was removed and the reaction was continued at 25℃ for 5 h. (3) Purification and collection: The reaction solution was slowly poured into 4 times the volume of methanol that had been pre-frozen to 0°C to precipitate. After filtration through a 200-mesh filter cloth, the white fibrous precipitate was collected. The crude product was transferred to a supercritical CO2 extraction device and dynamically extracted at 40°C and 15 MPa for 4 h. The purified product was placed in a vacuum oven at 50°C and dried for 48-50 h to finally obtain a white flocculent six-armed star-shaped PLCL macromonomer. Step 2: Preparation of composite nanofiber tube embryos Weigh 3.5-4.5 g of the six-armed star-shaped PLCL macromonomer prepared in the first step, 5.5-6.5 g of poly-L-lactic acid, 0.8-1.0 g of silk fibroin, 45-55 mg of heparin sodium, and 6-8 mL of PEGylated gold nanorod solution with a concentration of 0.1 mg / mL. Pour in 90-110 mL of hexafluoroisopropanol, seal, and stir in the dark at 25℃ and 400-500 rpm for 22-24 h to obtain the blended spinning solution. Build a uniaxial electrospinning device with a diameter of 3... A 0mm controllable speed metal mandrel was used as the receiving device. The feed flow rate was set to 1.0-1.5mL / h, the applied voltage to 12-18kV, the mandrel rotation speed to 800-1000rpm, and the distance between the needle tip and the mandrel receiving device to 15cm. Fiber tubes were spun under stable temperature and humidity. The fiber tubes and mandrel were placed in a fume hood at room temperature for 1h for initial dissolution. Then, they were dried under vacuum at 40℃ and -0.1MPa for 48h to completely remove hexafluoroisopropanol, finally obtaining a composite nanofiber tube preform with a thickness of 120±10μm. Step 3: Thermal densification, photocrosslinking, and hierarchical crystallization The composite nanofiber tube preform obtained in the second step was transferred into a processing chamber equipped with an ultraviolet light source and precise temperature control. After replacing the air with nitrogen, the temperature was raised to 120°C at 10°C / min and held for 10 min under nitrogen protection for annealing and relaxation to loosen the molecular chains. The temperature was maintained at 120°C, and the tube was irradiated uniformly in all directions for 5 min using an ultraviolet LED surface light source with a main wavelength of 365nm to complete photocrosslinking. Immediately after irradiation, a programmed cooling process was performed: first, the temperature was reduced from 120°C to 90°C at 2°C / min, then from 90°C to 60°C at 0.5°C / min, and annealed at this temperature for 30 min. Finally, the temperature was slowly cooled to room temperature at a rate of 1°C / min. The entire process was carried out under a nitrogen atmosphere. Step 4: Laser cutting and shaping followed by cleaning and sterilization The cross-linked crystallized tubing prepared in the third step was fixed on a rotating chuck, and the three-dimensional model data of the stent was imported. A complete stent mesh structure was cut out using a precision fiber laser. After completion, the stent was immersed in sterile ethanol and ultrasonically cleaned at 40°C for 5 minutes. After removal, it was dried in a clean bench to obtain a laser-cut and structurally complete coronary stent embryo. The coronary stent embryo was placed in a medical packaging bag and sealed. It was then sterilized by double-sided electron beam irradiation with an absorbed dose of 25-35 kGy. Step 5: Surface Biofunctionalization and Terminal Treatment (1) The coronary stent embryo prepared in the fourth step was placed in an oxygen plasma cleaner and treated at 100 Pa pressure and 100 W power for 60 s. Then it was immediately immersed in RGD peptide PBS solution with a concentration of 1 mg / mL and slowly shaken and incubated at 4℃ for 12 h. (2) Anticoagulant coating: The stent was immersed in 1 mg / mL chitosan and 2 mg / mL heparin sodium solution in sequence for 5 min each time, and gently rinsed three times with acetic acid-sodium acetate buffer solution at pH=5 in between. This process was repeated 5 times. (3) Growth factor coating: The scaffold was immersed in a PBS solution containing 10 μg / mL vascular endothelial growth factor and 10 μg / mL stromal cell-derived factor-1α and adsorbed at 4°C for 2 h. (4) Packaging: Place the functionalized stent in a medical packaging bag and seal it.

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