Controlled Degradable Polylactic Acid-Based Composite Material for Circumcision Suture Device and its Preparation Method

Abstract

This invention belongs to the field of medical polymer materials technology, specifically providing a controllable degradable polylactic acid-based composite material for circumcision suture devices and its preparation method. The composite material uses polylactic acid-polycaprolactone block copolymer as the matrix and adds pH-responsive core-shell nano-hydroxyapatite, compatibilizer, and plasticizer. The core-shell nanoparticles are composed of a zinc-silicon co-doped hydroxyapatite core and a poly(dimethylaminoethyl methacrylate) brush shell. In an acidic degradation environment, the brush shell expands to accelerate the release of zinc ions, thereby achieving intelligent antibacterial properties.

Description

Technical Field

[0001] This invention belongs to the field of medical polymer materials technology, specifically relating to a controllable degradable polylactic acid-based composite material for circumcision suture devices and its preparation method, and particularly to a degradable composite material containing pH-responsive polymer brush grafted with zinc-silicon co-doped core-shell structured nano-hydroxyapatite. Background Technology

[0002] Phimosis and paraphimosis are common urological conditions, and circumcision is the primary treatment in clinical practice. In recent years, disposable circumcision suture devices have become widely used in clinical practice due to their ease of operation and aesthetically pleasing results. Currently, commercially available circumcision suture devices are mainly made of non-degradable materials such as stainless steel or polycarbonate, requiring a second removal procedure after wound healing. This second removal not only increases the patient's psychological burden and the number of medical visits, but also carries the risk of bleeding and infection due to wound traction during removal. This fear and pain are particularly pronounced in children.

[0003] To address the aforementioned issues, biodegradable suture materials have become a research hotspot. Polylactic acid (PLA), a biodegradable polymer approved by the U.S. Food and Drug Administration (FDA), possesses good biocompatibility and processability. However, pure PLA materials inherently suffer from high brittleness and insufficient toughness. More importantly, the accumulation of localized acidic products during PLA degradation can trigger non-infectious inflammatory responses and induce acid-induced autocatalytic accelerated degradation, leading to a rapid loss of mechanical properties in the later stages rather than a gradual decline. This makes it difficult to precisely match the degradation cycle with the 8-12 week healing cycle of circumcision wounds.

[0004] In the prior art, Chinese patent CN108641074B discloses biodegradable materials, their preparation methods, and applications, but their degradation cycle is 6 to 24 months, which is much longer than that required for circumcision; Chinese patent CN107376027A discloses a polymer / hydroxyapatite whisker composite porous scaffold for cartilage tissue repair and its preparation method, which is prepared by solution electrospinning and is not suitable for suture components that require a dense structure; Chinese patent CN115463244B discloses a method for preparing a sustained-release antibacterial polylactic acid dressing and its application, but its tensile strength is only 5 to 15 MPa, which is far from meeting the mechanical requirements of 45 MPa or more for sutures.

[0005] Regarding antibacterial modification, existing technologies using zinc-doped nano-hydroxyapatite have been proven to possess broad-spectrum antibacterial activity. However, its zinc ion release behavior follows a passive diffusion mechanism, with the release rate influenced by environmental pH but lacking active response regulation capabilities. In the wound microenvironment after circumcision, lactic acid produced from the degradation of polylactic acid lowers the local pH, coinciding with the period of highest infection risk. Traditional zinc-doped hydroxyapatite cannot adaptively enhance antibacterial activity during this critical period. Furthermore, existing zinc-doped systems exhibit severe initial burst release of zinc ions followed by insufficient release later, making it difficult to maintain an effective antibacterial concentration throughout the entire wound healing cycle. None of the aforementioned existing technologies address the intelligent feedback regulation between antibacterial function and the acidic degradation environment in biodegradable materials specifically designed for circumcision suture devices. Summary of the Invention

[0006] The purpose of this invention is to overcome the shortcomings of existing technologies and provide a controllable degradable polylactic acid-based composite material for circumcision suture devices and its preparation method. This invention constructs a L-type polylactic acid-polycaprolactone block copolymer matrix and introduces pH-responsive polymer-grafted zinc-silicon co-doped core-shell nano-hydroxyapatite as a smart functional filler. This achieves a positive feedback antibacterial regulation mechanism that automatically triggers the accelerated release of zinc ions when the material degrades to create an acidic environment. Simultaneously, it ensures synergistic optimization of mechanical properties, degradation rate, and biosafety, enabling the suture device to self-degrade and be absorbed in vivo without the need for secondary circumcision.

[0007] To achieve the above objectives, the present invention provides a controllable degradable polylactic acid-based composite material for circumcision suture devices, using 100 parts by weight of L-polylactic acid-polycaprolactone block copolymer as the matrix resin. The block copolymer is prepared by sequential ring-opening polymerization of L-lactide and caprolactone, with a molar ratio of L-lactide to caprolactone of 70:30 to 85:15 and a number average molecular weight of 80,000 to 150,000.

[0008] The composite material also contains 3 to 10 parts by weight of pH-responsive core-shell structured nano-hydroxyapatite. This core-shell structured nano-hydroxyapatite consists of a zinc-silicon co-doped nano-hydroxyapatite core and a surface-grafted poly(2-(dimethylamino)ethyl methacrylate) brush shell. In the zinc-silicon co-doped nano-hydroxyapatite core, zinc ions account for 3% to 8% of the molar ratio of calcium ions, and silicate ions account for 8% to 25% of the molar ratio of phosphate ions, with a core particle size of 30 to 60 nm. The poly(2-(dimethylamino)ethyl methacrylate) brush shell is grown from the surface of the hydroxyapatite core via surface-initiated atom transfer radical polymerization, with a number-average molecular weight of 3000 to 8000 and a shell thickness of 8 to 20 nm.

[0009] The composite material also contains 1 to 5 parts by weight of a polylactic acid-polyethylene glycol-polylactic acid triblock compatibilizer and 2 to 8 parts by weight of a tributyl acetylacetate plasticizer.

[0010] The present invention also provides a method for preparing the composite material, comprising: a first step, using stannous octoate as a catalyst and lauryl alcohol as an initiator, firstly ring-opening polymerization of L-lactide at 160 to 170°C for 2 to 4 hours to form active chain ends, then adding caprolactone and continuing polymerization at 140 to 155°C for 3 to 5 hours to obtain a block copolymer; a second step, preparing zinc-silicon co-doped nano-hydroxyapatite cores by co-precipitation, then grafting with a silane coupling agent and 2-bromoisobutyryl bromylating to construct an atom transfer polymer on the surface. The free radical polymerization initiation site is used to initiate the atom transfer free radical polymerization of 2-(dimethylamino)ethyl methacrylate on the surface of nanoparticles using cuprous bromide and pentamethyldiethylenetriamine as catalytic systems to obtain core-shell structured nano-hydroxyapatite; the third step is to melt-blend and granulate the block copolymer, core-shell nanoparticles, triblock compatibilizer and plasticizer in a twin-screw extruder at 165 to 185°C; the fourth step is to injection mold the granules into stitcher parts at 130 to 150°C.

[0011] Compared with the prior art, the present invention has the following beneficial effects:

[0012] First, this invention designs a pH-responsive core-shell structured nano-hydroxyapatite filler and constructs a closed-loop positive feedback regulation mechanism of degradation-acid production-polymer brush expansion-accelerated zinc ion release-sterilization and acid neutralization-polymer brush collapse. When the polylactic acid matrix degrades to produce lactic acid, causing the local pH value to drop to 6.0 to 6.5, the tertiary amine groups in the shell poly(2-(dimethylamino)ethyl methacrylate) brush undergo protonation, leading to polymer chain swelling and exposing the zinc-silicon co-doped hydroxyapatite surface of the core. This increases the zinc ion release rate to 3 to 5 times that under pH 7.4 conditions. While the released zinc ions kill bacteria, the dissolution of calcium phosphate in the core neutralizes the local acidic environment. As the pH value returns to neutral, the polymer brush deprotonates and collapses, and the zinc ion release rate automatically decreases, avoiding cytotoxicity caused by excessive release. This intelligent feedback mechanism enables the material to automatically enhance its antibacterial activity during the acidic degradation phase, which carries the highest risk of infection, while maintaining a safe mode of low-level release during the normal healing phase. This achieves intelligent antibacterial control, releasing more when needed and less when not needed.

[0013] Second, this invention employs a zinc-silicon co-doping system instead of the traditional single zinc doping system. When silicate ions replace phosphate ions in the hydroxyapatite lattice, charge-compensating hydroxyl vacancies are generated. These vacancies constitute rapid diffusion channels for zinc ions to dissolve from the deep layers of the lattice to the surface under acidic conditions. Simultaneously, the release of silicon ions promotes fibroblast proliferation and collagen synthesis, exhibiting bioactivity that promotes wound healing. The synergistic effect of zinc-silicon co-doping results in approximately twice the zinc ion release rate in an acidic environment compared to a single zinc doping system, while the baseline release rate in a neutral environment shows no significant difference.

[0014] Third, this invention uses a L-type polylactic acid-polycaprolactone block copolymer to replace the physical blending system. The block structure chemically bonds rigid polylactic acid segments to flexible polycaprolactone segments at the molecular level, eliminating the phase separation and interfacial defects unavoidable in physical blending. The composite material achieves a tensile strength of 48 to 62 MPa and an elongation at break of 18% to 35%, meeting the mechanical performance requirements of the suture device. By adjusting the block ratio and the core-shell nanoparticle content, the complete degradation and absorption period of the material in vivo can be precisely controlled within the range of 8 to 12 weeks.

[0015] Fourth, all degradation products of this invention are substances that can be metabolized by the human body, including lactic acid metabolized through the tricarboxylic acid cycle, caprolactam decomposed through the fatty acid oxidation pathway, hydroxyapatite absorbed and utilized by the skeletal system, zinc and silicon ions participating in normal metabolism as essential trace elements for the human body, and poly(2-(dimethylamino)ethyl methacrylate) degraded into small molecule amines and excreted by the kidneys. The biosafety has been verified to meet the qualified standards through cytotoxicity tests and subcutaneous implantation tests. Attached Figure Description

[0016] Figure 1 Transmission electron microscope image of the pH-responsive core-shell structured nano-hydroxyapatite obtained in Example 2 of this invention.

[0017] Figure 2 This is a schematic diagram of the pH response positive feedback intelligent antibacterial regulation mechanism of the present invention.

[0018] Figure 3 This is a scanning electron microscope image of the tensile cross-section of the composite material obtained in Example 1 of the present invention.

[0019] Figure 4 The cumulative zinc ion release curves of the core-shell nanoparticles obtained in Example 2 of this invention under different pH conditions are shown.

[0020] Figure 5 The curves show the changes in mass loss rate during the in vitro simulated degradation process of Examples 1 to 3 and Comparative Example 1 of the present invention.

[0021] Figure 6 The curves show the changes in antibacterial rate at different time points during in vitro degradation of Example 1 and Comparative Example 2 of the present invention. Detailed Implementation

[0022] The technical solution of the present invention will be described in detail below with reference to specific embodiments. All reagents and instruments used in the following embodiments are commercially available. Specifically, L-lactide with a purity greater than 99.5% was purchased from Zhejiang Hisun Biomaterials Co., Ltd.; caprolactone with a purity greater than 99% was purchased from Sinopharm Chemical Reagent Co., Ltd.; stannous octoate with a purity greater than 95% was purchased from Sigma-Aldrich; 2-(dimethylamino)ethyl methacrylate with a purity greater than 98% was purchased from Shanghai Aladdin Biochemical Technology Co., Ltd., and the polymerization inhibitor was removed by a neutral alumina column before use; 2-bromoisobutyryl bromide with a purity greater than 98% was purchased from Sigma-Aldrich; and cuprous bromide with a purity greater than 99%... The following were purchased from Sigma-Aldrich: pentamethyldiethylenetriamine (greater than 99% purity) from Tokyo Chemical Industry Co., Ltd.; gamma-aminopropyltriethoxysilane (greater than 98% purity) from Nanjing Shuguang Chemical Group Co., Ltd.; triethylamine (greater than 99% purity) from Sinopharm Chemical Reagent Co., Ltd.; calcium nitrate, zinc nitrate, tetraethyl orthosilicate, and diammonium hydrogen phosphate (all analytical grade) from Sinopharm Chemical Reagent Co., Ltd.; polyethylene glycol (number average molecular weight 4000) from Sinopharm Chemical Reagent Co., Ltd.; and tributyl acetylacetate (greater than 99% purity) from Jinan Henghua Technology Co., Ltd. The twin-screw extruder was a Nanjing Keya SHJ-20 model, and the injection molding machine was a Haitian MA900 model.

[0023] Preparation Example 1: Synthesis of L-polylactic acid-polycaprolactone block copolymer:

[0024] 100g of L-lactide and the reactor were placed in a vacuum drying oven and dried at 80°C for 12 hours to remove residual moisture. The dried L-lactide was then transferred to a 500mL stainless steel reactor that had been purged with nitrogen three times. 0.05g of stannous octoate and 0.3g of lauryl alcohol were added as catalyst and initiator, respectively. Under nitrogen protection, the reactor was heated to 165°C, and the stirring speed was maintained at 60 rpm for 3 hours for polymerization. At this point, a sample was taken for testing, and the L-lactide conversion rate reached over 95%, forming a poly(L-lactic acid) prepolymer with active hydroxyl end groups.

[0025] Under conditions of uninterrupted reaction, 42.8 g of caprolactone was injected into the reactor to achieve a molar ratio of L-lactide to caprolactone of 80:20. The reaction temperature was lowered to 150°C, and polymerization was continued with stirring for 4 h, allowing caprolactone to initiate ring-opening polymerization at the active chain ends of poly-L-lactic acid, forming a poly-L-lactic acid-polycaprolactone diblock copolymer. After the reaction, the product was dissolved in dichloromethane, purified three times by methanol precipitation, and dried in a vacuum drying oven at 50°C for 24 h. Gel permeation chromatography showed that the number-average molecular weight of the obtained block copolymer was 112,000, and the molecular weight distribution index was 1.45. Differential scanning calorimetry showed two distinct glass transition temperatures of -55.3°C and 56.8°C, corresponding to the polycaprolactone block and the polylactic acid block, respectively, confirming the successful construction of the block structure.

[0026] Preparation Example 2: Preparation of pH-responsive core-shell structured nano-hydroxyapatite:

[0027] The first stage involved the preparation of zinc-silicon co-doped nano-hydroxyapatite cores. 14.16 g of calcium nitrate and 0.595 g of zinc nitrate were dissolved in 200 mL of deionized water to form a mixed cation solution, with zinc ions comprising 5% of the molar amount of calcium ions. 3.17 g of diammonium hydrogen phosphate and 0.52 g of tetraethyl orthosilicate were dissolved in 100 mL and 50 mL of deionized water, respectively, and the pH of both solutions was adjusted to 10.5. Silicate ions comprised 15% of the molar amount of phosphate ions. Under constant temperature water bath conditions of 60°C and magnetic stirring, the phosphate and silicate solutions were simultaneously added dropwise to the cation solution at a rate of 2 mL / min. During the addition, ammonia was continuously used to maintain the pH of the reaction system within the range of 10.0 to 10.5. After the addition was complete, the reaction was stirred for another 2 hours. The precipitate was then aged at room temperature for 24 hours, centrifuged, washed three times with deionized water and once with anhydrous ethanol, vacuum dried at 60°C for 12 hours, and then calcined at 400°C for 2 hours to obtain zinc-silicon co-doped nano-hydroxyapatite cores. X-ray diffraction analysis confirmed that the product was a pure-phase hydroxyapatite structure, with a slight shift in interplanar spacing compared to standard hydroxyapatite, confirming that the dopant ions successfully entered the crystal lattice. Transmission electron microscopy revealed rod-shaped particles with a diameter of 35 to 55 nm, such as... Figure 1 As shown. Inductively coupled plasma atomic emission spectrometry (ICP-AES) analysis revealed that the zinc doping concentration was 4.8 mol% and the silicon doping concentration was 14.2 mol%.

[0028] The second stage involved the construction of surface initiation sites. 10 g of zinc-silicon co-doped nano-hydroxyapatite powder was dispersed in 100 mL of anhydrous ethanol and ultrasonically dispersed for 30 min. 0.5 g of gamma-aminopropyltriethoxysilane was dissolved in 20 mL of an ethanol-water solution, and the pH was adjusted to 4.0 for pre-hydrolysis for 1 h. The pre-hydrolyzed solution was added to the nanoparticle dispersion, and the mixture was refluxed and stirred at 70°C for 4 h. After the reaction, the nanoparticles were centrifuged, washed, and dried to obtain amino-functionalized nano-hydroxyapatite. Subsequently, the amino-functionalized nanoparticles were redispersed in 50 mL of anhydrous dichloromethane, and 1.4 mL of triethylamine was added as an acid-binding agent. Under ice bath and nitrogen protection, 0.92 g of 2-bromoisobutyryl bromide was slowly added dropwise. After the addition was complete, the reaction was continued at room temperature for 12 h. The reaction product was centrifuged, washed three times each with dichloromethane and methanol, and then vacuum dried to obtain nano-hydroxyapatite with bromoisobutyryl atom transfer radical polymerization initiation sites on its surface. Fourier transform infrared spectroscopy detected an absorption peak of the amide carbonyl stretching vibration at 1740 cm⁻¹ and a characteristic peak of the carbon-bromine bond at 1260 cm⁻¹, confirming the successful construction of the initiation site. X-ray photoelectron spectroscopy detected a 3d characteristic peak of bromine at 70.2 eV, with an initiation site density of approximately 0.35 mmol / g.

[0029] The third stage involved surface-initiated atom transfer radical polymerization to grow poly(2-(dimethylamino)ethyl methacrylate) brush shells. In a deoxygenated 100 mL Schlenk tube, 5 g of nano-hydroxyapatite containing initiation sites and 40 mL of a mixed solvent of methanol and deionized water (volume ratio 1:1) were added. 7.85 g of 2-(dimethylamino)ethyl methacrylate monomer was added, along with 0.072 g of cuprous bromide and 0.173 g of pentamethyldiethylenetriamine as catalyst and ligand, respectively, under nitrogen protection. The molar ratio of monomer to initiation sites was 30:1. The reaction system was stirred at 30°C for 6 h. After the reaction, tetrahydrofuran was added to terminate the reaction. After centrifugation, washing three times each with methanol and deionized water, and vacuum drying at 40°C for 24 h, core-shell structured nano-hydroxyapatite was obtained. Thermogravimetric analysis showed that the organic shell accounted for 18.5% of the total mass, and the estimated number-average molecular weight of the polymer brush was approximately 5200. Dynamic light scattering tests showed that the hydrodynamic diameter of the core-shell particles in deionized water was 82 nm at pH 7.4, expanding to 128 nm at pH 5.5, with a particle size increase ratio of 1.56, confirming the successful construction of pH responsiveness. The zeta potential was -8.5 mV at pH 7.4, reversing to +22.3 mV at pH 5.5, further confirming the protonation behavior of the tertiary amine group.

[0030] To verify the pH-responsive zinc ion release behavior of the core-shell structure, core-shell nanoparticles were dispersed in phosphate buffer solutions at pH 7.4 and pH 5.5, respectively, and cultured with shaking at 37°C. The supernatant was periodically collected, and the zinc ion concentration was determined by inductively coupled plasma atomic emission spectrometry (ICP-AES). At pH 7.4, the cumulative zinc ion release over 24 hours was 0.85 μg / mL; at pH 5.5, the cumulative zinc ion release over 24 hours was 3.62 μg / mL, which was 4.3 times that under neutral conditions. As a control, the cumulative zinc ion release over 24 hours for zinc-silicon co-doped nano-hydroxyapatite without polymer grafts was 1.12 μg / mL and 2.15 μg / mL at pH 7.4 and pH 5.5, respectively, with an acid / neutral release ratio of only 1.9 times. These results indicate that the core-shell structured pH-responsive polymer brush increases the acceleration factor of zinc ion release under acidic conditions from 1.9 times to 4.3 times, achieving a significantly enhanced pH-triggered intelligent release effect. Meanwhile, the zinc release of the core-shell structure under neutral conditions (0.85 ug / mL) was lower than that of the uncoated system (1.12 ug / mL), indicating that the polymer brush acted as a release barrier in the collapsed state, effectively inhibiting ineffective release under neutral conditions.

[0031] The effects of different pH gradients on the zinc release behavior of core-shell nanoparticles were further investigated, such as... Figure 4 As shown in the figure, the zinc release over 24 hours was measured by dispersing core-shell particles in buffer solutions at pH 7.4, pH 6.8, pH 6.2, pH 5.5, and pH 5.0, respectively, and the values ​​were 0.85, 1.25, 2.18, 3.62, and 4.85 μg / mL. Plotting zinc release against pH revealed a sharp increase in release between pH 6.5 and 6.0, which corresponds precisely to the apparent pKa values ​​of poly(2-(dimethylamino)ethyl methacrylate) from 6.3 to 6.8, confirming that the conformational change caused by protonation of the polymer brush is the molecular switch triggering the accelerated release of zinc ions. At pH 5.0, the release reached 4.85 μg / mL, but pre-cytotoxicity assessment showed that the survival rate of L929 cells at this concentration was still 88.2%, lower than the direct damaging effect on cells from the standard pH 4.0 phosphate buffer solution, indicating that even under extremely acidic conditions, the concentration of zinc ions released from the core-shell structure remains within a safe range.

[0032] To evaluate whether the pH response behavior of the core-shell structure in the composite matrix is ​​consistent with that in the free state, the composite sheet obtained in Example 1 was immersed in buffer solutions with pH 5.5 and pH 7.4, and the zinc ion concentration in the solution was measured after 24 hours. The zinc ion concentration released by the composite material at pH 5.5 was 0.52 μg / mL, and at pH 7.4 it was 0.15 μg / mL, with an acid / neutral release ratio of 3.5 times. Although the absolute release amount was lower than that in the free state due to the core-shell particles being encapsulated by the matrix, the pH response ratio (3.5 times) was basically consistent with that in the free state (4.3 times), indicating that the core-shell nanoparticles retained their complete pH-responsive intelligent release function after being embedded in the composite matrix.

[0033] Preparation Example 3: Synthesis of Triblock Compatibilizers:

[0034] 20 g of polyethylene glycol (PEG) with a number average molecular weight of 4000 was dried under vacuum at 120°C for 6 h. The dried PEG was transferred to a 250 mL three-necked flask purged with nitrogen three times, and 80 g of L-lactide and 0.02 g of stannous octoate were added. The mixture was heated to 155°C under nitrogen protection and stirred at 60 rpm for 5 h to polymerize. The hydroxyl groups at both ends of the PEG acted as bifunctional initiators to initiate the ring-opening polymerization of L-lactide, forming a L-polylactic acid-polyethylene glycol-L-polylactic acid triblock copolymer. The reaction product was dissolved in dichloromethane, purified by methanol precipitation, and then dried under vacuum. Gel permeation chromatography showed a number average molecular weight of 23500 and a molecular weight distribution index of 1.38.

[0035] Example 1

[0036] 100 parts of the L-polylactic acid-polycaprolactone block copolymer obtained in Preparation Example 1, 7 parts of the pH-responsive core-shell structured nano-hydroxyapatite obtained in Preparation Example 2, 3 parts of the triblock compatibilizer obtained in Preparation Example 3, and 5 parts of tributyl acetylacetate were weighed and premixed in a high-speed mixer at 1000 r / min for 5 min. The premix was then added to a twin-screw extruder for melt blending and granulation. The temperatures of each section of the extruder were set as follows: feeding section 155°C, compression section 170°C, metering section 175°C, and die head 180°C. The screw speed was 120 r / min. The extruded molten material was water-cooled, stretched, air-dried, and then pelletized. After vacuum drying the granules at 80°C for 4 hours, they were injection molded. The barrel temperature was 160°C in zone 1, 170°C in zone 2, 175°C in zone 3, and 175°C in the nozzle. The mold temperature was 40°C, the injection pressure was 80MPa, the holding pressure was 50MPa, the holding time was 8s, and the cooling time was 20s.

[0037] The mechanical properties of the obtained composite material were tested. According to GB / T 1040.2-2006 standard, the tensile strength of the dumbbell-shaped specimen at a tensile speed of 5 mm / min was 55.3 MPa, the elongation at break was 25.6%, and the elastic modulus was 2.85 GPa. The notched impact strength of the simply supported beam was tested according to GB / T 1043.1-2008 standard, and the measured value was 11.8 kJ / m². The flexural properties were tested according to GB / T 9341-2008 standard, and the flexural strength was 78.2 MPa, and the flexural modulus was 3.12 GPa. All the above mechanical properties meet the requirements for the material used in the circumcision stapler.

[0038] Differential scanning calorimetry (DSC) was performed on the composite material of Example 1 at a heating rate of 10°C / min. The results showed that the material possesses two glass transition temperatures, located at -48.5°C and 52.3°C, corresponding to the polycaprolactone (PVC) and polylactic acid (PLA) blocks, respectively. However, both transition temperatures are shifted towards each other, indicating that the block structure leads to partial interpenetration between the microphase-separated regions. The crystallization melting peak appeared at 160.8°C, with a crystallinity of 28.5%. Thermogravimetric analysis showed an initial decomposition temperature of 285°C, significantly higher than the processing temperature range, indicating good thermal stability and a suitable processing window.

[0039] The tensile and impact fracture morphologies of the composite material in Example 1 were observed using a scanning electron microscope, as shown below. Figure 3 As shown, the tensile fracture surface exhibits obvious ductile fracture characteristics, with numerous silver streaks and microfibrillated structures. The polycaprolactone blocks are stretched during fracture to form slender microfibers, and the orientation and fracture of these microfibers consume a significant amount of impact energy. On the impact fracture surface, core-shell nanoparticles are uniformly dispersed in the matrix, and no obvious particle debonding voids were observed, indicating a good interfacial bond between the inorganic nanoparticles and the organic matrix after compatibilization with the triblock compatibilizer. Energy-dispersive X-ray spectroscopy surface scanning confirmed that zinc, silicon, calcium, and phosphorus elements are uniformly distributed on the fracture surface, with no obvious enrichment regions.

[0040] Transmission electron microscopy (TEM) was used to observe the ultrathin sections. At low magnification, the core-shell particles were uniformly dispersed with an interparticle spacing of 100 to 300 nm, and no obvious chain-like or clustered aggregates were observed. At high magnification, the core-shell structure of individual core-shell particles could be resolved. The dark inorganic hydroxyapatite core was surrounded by a light-colored organic polymer brush shell, with a shell thickness of approximately 12 to 18 nm, consistent with the results of dynamic light scattering measurements in Preparation Example 2. A gradual contrast transition was observed between the polymer brush shell and the surrounding matrix in some particles, indicating that the polymer chains on the outer layer of the brush shell interpenetrated with the matrix, forming a diffusion interface layer. This interface structure facilitates the effective transfer of stress from the matrix to the nanoparticles. Small-angle X-ray scattering (SAXS) further confirmed the uniform dispersion of the core-shell particles in the matrix. The scattering curve showed a single scattering peak in the low-angle region, corresponding to an average interparticle spacing of approximately 180 nm, consistent with the TEM observations.

[0041] Example 2

[0042] The difference between this embodiment and Example 1 is that the molar ratio of L-lactide to caprolactone in the block copolymer is 75:25, the amount of core-shell nanoparticles added is 5 parts, the amount of triblock compatibilizer is 2 parts, and the amount of plasticizer is 6 parts. Other preparation processes are the same as in Example 1. The resulting composite material has a tensile strength of 50.1 MPa, an elongation at break of 31.2%, and a notched impact strength of 13.5 kJ / m². This formulation is suitable for snap-fit ​​connections in suture devices that require high flexibility.

[0043] Example 3

[0044] The difference between this embodiment and Example 1 is that the molar ratio of L-lactide to caprolactone in the block copolymer is 85:15, the amount of core-shell nanoparticles added is 10 parts, the amount of triblock compatibilizer is 4 parts, and the amount of plasticizer is 3 parts. The resulting composite material has a tensile strength of 61.5 MPa, an elongation at break of 18.3%, and a notched impact strength of 8.7 kJ / m². This formulation is suitable for the main structural parts of the suture ring.

[0045] Example 4

[0046] The difference between this embodiment and Example 1 is that the zinc doping content in the core-shell nanoparticles is 3 mol%, the silicon doping content is 8 mol%, the molar ratio of monomer to initiation site during the surface atom transfer radical polymerization reaction is 20:1, the number-average molecular weight of the polymer brush is approximately 3200, and the shell thickness is approximately 8 nm. The amount of core-shell nanoparticles added is 3 parts. The resulting composite material has a tensile strength of 52.8 MPa and an elongation at break of 22.5%. The zinc ion release rate at pH 5.5 is 3.2 times higher than that at pH 7.4.

[0047] Example 5

[0048] The difference between this embodiment and Example 1 is that the zinc doping content in the core-shell nanoparticles is 8 mol%, the silicon doping content is 25 mol%, the molar ratio of monomer to initiation site is 50:1, the number-average molecular weight of the polymer brush is approximately 7800, and the shell thickness is approximately 20 nm. The amount of core-shell nanoparticles added is 8 parts. The resulting composite material has a tensile strength of 48.2 MPa and an elongation at break of 33.8%. The zinc ion release rate at pH 5.5 is 5.1 times higher than that at pH 7.4, indicating the most significant pH response.

[0049] Comparative Example 1

[0050] Physical blending was used instead of block copolymerization. 80 parts of L-polylactic acid homopolymer and 20 parts of polycaprolactone homopolymer were weighed and melt-blended with 7 parts of core-shell nanoparticles obtained in Preparation Example 2 and 5 parts of polyethylene glycol monomethyl ether in a twin-screw extruder. The resulting physical blend had a tensile strength of 38.6 MPa, an elongation at break of 9.8%, and a notched impact strength of 5.3 kJ / m². Compared with Example 1, the tensile strength decreased by 30.2% and the elongation at break decreased by 61.7%, confirming that the interfacial compatibility of the block copolymer system was significantly better than that of physical blending.

[0051] Comparative Example 2: Zinc-doped nano-hydroxyapatite in ungrafted polymer brushes

[0052] Zinc-doped (without silicon co-doping) and ungrafted polymer brush nanoparticles were used to replace the core-shell structured nanoparticles. 100 parts of block copolymer, 7 parts of silane coupling agent-modified zinc-doped nano-hydroxyapatite (5 mol% zinc doping, no silicon doping, no polymer brush), 3 parts of triblock compatibilizer, and 5 parts of plasticizer were weighed, following the same process as in Example 1. The resulting composite material had a tensile strength of 54.8 MPa, with mechanical properties essentially equivalent to Example 1. However, the zinc ion release rate at pH 5.5 was only 1.9 times higher than that at pH 7.4, significantly lower than the 4.3 times higher rate in Example 1. In in vitro degradation antibacterial tests from week 4 to week 8, the inhibition rate of Comparative Example 2 decreased sharply from 92.5% at week 4 to 52.3% at week 8, while the inhibition rate of Example 1 only decreased slowly from 95.2% to 82.8%, confirming the crucial role of the pH-responsive intelligent release mechanism in maintaining long-lasting antibacterial activity.

[0053] Comparative Example 3: Zinc-silicon co-doped but ungrafted polymer brush

[0054] Nano-hydroxyapatite co-doped with zinc and silicon but modified only by a silane coupling agent and without grafted polymer brushes was used. The resulting composite material had a tensile strength of 53.5 MPa and an elongation at break of 23.8%, with mechanical properties similar to those of Example 1. The zinc ion release rate ratio at pH 5.5 and pH 7.4 was 2.8 times higher than that of Comparative Example 2 (1.9 times), indicating that the lattice vacancies generated by silicon doping did indeed accelerate zinc dissolution under acidic conditions; however, it was still significantly lower than that of Example 1 (4.3 times), confirming that the pH-responsive expansion and collapse behavior of the polymer brush shell is an indispensable element for achieving high-rate intelligent release. Meanwhile, the zinc ion release amount of this comparative example under neutral conditions was 1.08 ug / mL, higher than the 0.85 ug / mL of the core-shell structure, indicating that the lack of a polymer brush barrier effect leads to increased ineffective zinc loss under neutral conditions. At week 8 of in vitro degradation, the antibacterial rate of this comparative example was 65.8%, which was higher than that of Comparative Example 2 (52.3%) but still significantly lower than that of Example 1 (82.8%). The data from the three sets of comparative examples formed a complete progressive comparison chain, clearly revealing the independent and irreplaceable functional contributions of the zinc-silicon co-doped core and the polymer brush shell in the core-shell structure.

[0055] The materials obtained in Examples 1 to 5 and Comparative Examples 1 to 3 were injection molded into 10 mm × 10 mm × 1 mm sheets, sterilized, and then immersed in a pH 7.4 phosphate buffer solution at 37°C, with the solution being replaced weekly. Samples were taken at weeks 1, 2, 4, 6, 8, 10, and 12 to determine the mass loss rate, molecular weight retention rate, tensile strength retention rate, and pH value of the degradation solution.

[0056] The degradation results of Example 1 are as follows Figure 5 As shown: Week 1: Mass loss rate 0.8%, molecular weight retention 95.2%, strength retention 97.5%; Week 2: Mass loss rate 2.1%, molecular weight retention 88.6%, strength retention 93.2%; Week 4: Mass loss rate 8.5%, molecular weight retention 72.3%, strength retention 82.8%; Week 6: Mass loss rate 22.3%, molecular weight retention 48.5%, strength retention 58.6%; Week 8: Mass loss rate 45.8%, molecular weight retention 25.2%, strength retention 32.5%; Week 10: Mass loss rate 78.5%, molecular weight retention 8.6%; Week 12: Mass loss rate 96.8%, the sample is basically completely degraded. The pH value of the degradation solution remained within the range of 6.52 to 7.38 throughout the cycle, reflecting the buffering effect of the alkaline degradation products of zinc-silicon co-doped hydroxyapatite in the core-shell nanoparticles on the acidic environment.

[0057] The molecular weight changes of the material in Example 1 during degradation were tracked using gel permeation chromatography. The number-average molecular weight was 108,500 before degradation, decreasing to 92,300 in week 2, 71,200 in week 4, 38,500 in week 6, 15,200 in week 8, and 4,800 in week 10. Plotting molecular weight against time showed an approximately linear decay curve for the block copolymer, with a linear fit coefficient of determination (R²) of 0.987. Example 2, due to its higher polycaprolactone block ratio, experienced a slightly faster degradation rate, reaching 85.2% by week 10. Example 3, with its highest polylactic acid block ratio and largest core-shell nanoparticle content, exhibited the slowest degradation rate, with a mass loss of 88.5% by week 12 and near-complete degradation by week 14.

[0058] The physical blend material in Comparative Example 1 exhibited typical collapse-type degradation during the degradation process. The mass loss was slow, only 5.2% before week 6, but rapidly decreased to 62.3% from week 6 to week 8. The pH of the degradation solution dropped to 4.85 in week 6, triggering an acid-induced autocatalytic effect. The degradation behavior of the block copolymer system showed an approximately linear characteristic; a linear fit of the molecular weight decay data yielded a coefficient of determination (R²) of 0.987. The degradation behavior of the material in Comparative Example 2 was basically consistent with that in Example 1, with a mass loss rate of 95.5% in week 12. However, the pH of the degradation solution dropped to 5.85 to 6.12 from week 4 to week 8, significantly lower than the 6.52 to 6.85 in the same period of Example 1. This indicates the lack of a core-shell structure with a pH-responsive polymer brush shell. Although its zinc-silicon co-doped hydroxyapatite also has acid-base buffering properties, its buffering efficiency is lower than that of the core-shell structure. The pH value of the degradation solution of Comparative Example 3 was 6.15 to 6.55 from week 4 to week 8, falling between that of Example 1 and Comparative Example 2. This indicates that silicon doping increases the reactivity of the hydroxyapatite surface and the release rate of alkaline ions, but it still cannot replace the pH-responsive release mechanism of the polymer brush shell. The systematic comparison of the above degradation behavior fully confirms the indispensable synergistic necessity of both the zinc-silicon co-doped core and the pH-responsive polymer brush shell in the core-shell structure.

[0059] Further subcutaneous implantation in rats was used for in vivo degradation assays to verify the results. In Example 1, the in vivo mass loss rates at weeks 4, 8, and 12 were 12.5%, 52.3%, and 98.2%, respectively, with the in vivo degradation rate being slightly faster than in vitro. Complete degradation and absorption were achieved within 10 to 12 weeks.

[0060] According to GB / T 31402-2015 standard, the antibacterial properties were tested using Staphylococcus aureus ATCC 6538 and Escherichia coli ATCC 25922 as test strains. Example 1 showed inhibition rates of 95.2% and 91.8% against Staphylococcus aureus and Escherichia coli, respectively. Example 5, due to its maximum zinc doping content and polymer brush thickness, achieved inhibition rates of 98.2% and 95.6%, respectively. In Comparative Example 2, the initial inhibition rates of the ungrafted polymer brush were 92.5% and 88.3%, but these rates dropped sharply to 52.3% and 45.6% at week 8 of in vitro degradation, while Example 1 maintained effective antibacterial levels of 82.8% and 76.5% at the same time point.

[0061] The mechanism of this difference lies in the following: In Comparative Example 2, zinc ions are released passively through diffusion. Initially, the release is severe due to the large concentration gradient, but later, the release weakens due to the formation of a zinc-poor surface layer. In Example 1, the core-shell structure experiences a local pH decrease in the later stages of degradation due to acid production from the polylactic acid matrix degradation. This triggers the expansion of the polymer brush, exposing more zinc-silicon co-doped hydroxyapatite surface, creating a positive feedback effect where more ions are released when antibacterial properties are needed. Figure 6 As shown, this method maintains an effective release concentration of zinc ions throughout the 8-12 week degradation period. Zinc ion release kinetics tests revealed that in Example 1, the average daily zinc release was 0.12 to 0.18 ug / mL in the first four weeks. From weeks 4 to 8, due to acid production from matrix degradation triggering a pH response, the average daily zinc release increased to 0.20 to 0.32 ug / mL. From weeks 8 to 12, it decreased to 0.08 to 0.12 ug / mL, exhibiting a unique mid-term enhanced release curve. In Comparative Example 2, the average daily zinc release continuously decreased from 0.35 ug / mL in week 1 to 0.03 ug / mL in week 8, falling below the minimum inhibitory concentration (MIC) of 0.05 ug / mL for Staphylococcus aureus, thus losing its effective antibacterial ability.

[0062] The antibacterial effect on the material surface was observed using a double staining method with live and dead bacteria. Samples from Example 1 and Comparative Example 2 were taken out at week 6 of in vitro degradation and co-cultured with Staphylococcus aureus for 6 hours before observation under a fluorescence microscope. The proportion of dead bacteria on the surface of the material in Example 1 exceeded 88%, while the proportion of dead bacteria on the surface of the material in Comparative Example 2 was only 35%. Further testing of anti-biofilm formation ability showed that the amount of biofilm formed on the surface of the material in Example 1 at week 6 of degradation was only 12.5% ​​of that in Comparative Example 2. This result has important clinical significance for circumcision suture devices, because once a biofilm forms on the surface of the suture device, it will significantly reduce the penetration effect of local antibiotics and lead to prolonged postoperative infection. The positive feedback mechanism of the core-shell structured nanoparticles of this invention, which automatically enhances zinc ion release in the middle of degradation, precisely compensates for the key deficiency of traditional passive release systems in the later stages of degradation, resulting in weak antibacterial activity.

[0063] To evaluate the long-term antibacterial stability of core-shell nanoparticles in composite materials, the material obtained in Example 1 was stored for 6 months under accelerated aging conditions of 60°C and 75% relative humidity. After aging, the antibacterial performance was tested again. The inhibition rates against Staphylococcus aureus and Escherichia coli after accelerated aging were 93.8% and 90.2%, respectively, representing only a decrease of 1.4% and 1.6% compared to the pre-aging levels of 95.2% and 91.8%. These results indicate that the polymer brush shell layer in the core-shell structure, which coats the zinc-silicon co-doped hydroxyapatite core, not only achieves pH-responsive intelligent release but also acts as a protective barrier during storage, effectively preventing unnecessary loss of zinc ions due to environmental humidity and ensuring the stability of the antibacterial performance throughout the product's shelf life.

[0064] Following the GB / T 16886.5-2017 standard, L929 mouse fibroblasts were used as target cells, and cytotoxicity was tested using the extraction method. At 24 h and 72 h of culture, the cell viability of the extract group from Example 1 was 98.5% and 96.2%, respectively, with a cytotoxicity rating of 0 to 1. Further degradation solutions from the material of Example 1 at weeks 4, 8, and 12 were collected for cytotoxicity testing. At each time point, the cell viability of the degradation solution group was 95.8%, 92.5%, and 90.3%, respectively, all at grade 1, confirming the biocompatibility of the degradation products throughout the entire degradation cycle.

[0065] Subcutaneous implantation experiments were conducted on SD rats according to GB / T 16886.6-2015. One week post-operation, a small number of macrophages and lymphocytes were observed infiltrating around the implant, which is a normal foreign body reaction. Four weeks post-operation, the inflammatory response significantly decreased, with mild erosion of the material surface. Eight weeks post-operation, the implant shrank significantly, inflammation largely subsided, and new capillaries were observed ingrained into the degradation area. Twelve weeks post-operation, the implant was almost completely degraded, and the implantation site was replaced by normal connective tissue. Hematoxylin-eosin staining and Masson's trichrome staining of tissue sections showed no abnormal tissue reaction or chronic inflammation.

[0066] Immunohistochemical analysis was performed on the tissue surrounding the implantation site to detect the expression levels of the inflammatory factors interleukin-6 (IL-6) and tumor necrosis factor-alpha (TNF-alpha). One week post-operation, the expression levels of IL-6 and TNF-alpha in the Example 1 implantation group were 2.3 times and 1.8 times higher than those in the negative control group, respectively, indicating a mild inflammatory response. Four weeks post-operation, the expression levels of these two inflammatory factors decreased to 1.3 times and 1.1 times higher than those in the negative control group, respectively, approaching normal levels. From eight to twelve weeks post-operation, there was no statistically significant difference in inflammatory factor expression levels compared to the negative control group. Compared to the material implantation group in Comparative Example 2, the inflammatory factor expression levels in Example 1 were significantly lower from four to eight weeks post-operation. This may be related to the healing-promoting effect of silicon ions in the core-shell nanoparticles and the reduced infection-mediated inflammatory response due to the antibacterial activity of zinc ions. The zinc content in the tissue surrounding the implantation site was determined using laser ablation inductively coupled plasma mass spectrometry (ICP-MS). Four weeks post-operation, the zinc content was 1.5 times that of normal tissue, and it returned to normal levels after twelve weeks, indicating that the released zinc ions were cleared through normal metabolic pathways without local accumulation.

[0067] The suture staple component was injection molded using the formulation of Example 5, and the main ring structure component was injection molded using the formulation of Example 3. These components were then assembled into a complete simulated suture device. The ring clamping force was 12.5 to 15.8 N, meeting the requirement of over 10 N for foreskin tissue fixation. The suture staple had a puncture force of 3.2 to 4.5 N when piercing fresh pigskin, and no breakage occurred during the puncture and bending process. The simulated suture device was immersed in a phosphate buffer solution at 37°C. At weeks 2, 4, and 6, the ring clamping force retention rates were 95.2%, 82.5%, and 55.8%, respectively, maintaining over 80% clamping force during the first four weeks of the primary healing period. By week 8, the clamping force retention rate had decreased to 22.3%, indicating significant loosening of the suture device, eliminating the need for manual ring removal.

[0068] To simulate the degradation environment in actual use after circumcision, a simulated suture device was placed in simulated bodily fluid containing a certain concentration of artificial sebum and artificial sweat for degradation testing, to more closely resemble the moist and warm microenvironment of the penile foreskin tissue. Results showed that the degradation rate of the suture device under simulated bodily fluid conditions was approximately 15% to 20% faster than in pure phosphate buffer solution, achieving complete degradation by week 10, which highly coincides with the conventional time window of 8 to 10 weeks after clinical circumcision for ring removal. During the degradation process in simulated bodily fluid, the core-shell nanoparticles exhibited a more significant pH-responsive effect. This is because the organic acid components in sebum and sweat further lowered the local pH value, more fully triggering the expansion of the polymer brush and the accelerated release of zinc ions.

[0069] The effects of surface-initiated atom transfer radical polymerization (ATRP) conditions on the pH response of core-shell particles were investigated. When the molar ratio of monomer to initiation site increased from 20:1 to 50:1, the number-average molecular weight of the polymer brush increased from 3200 to 7800, the particle size expansion ratio at pH 5.5 increased from 1.32 to 1.72, and the acid / neutral ratio of zinc ion release increased from 3.2 to 5.1 times. However, when the molar ratio exceeded 60:1, the excessively long polymer brush led to poor dispersibility of the core-shell particles in organic solvents, making them prone to aggregation during extrusion. Therefore, a molar ratio of 25:1 to 50:1 was preferred. Increasing the reaction temperature from 25°C to 40°C accelerated the polymerization rate but broadened the molecular weight distribution; 30°C was the optimal reaction temperature. The optimal molar ratio of the catalyst cuprous bromide to the ligand pentamethyldiethylenetriamine was 1:1.

[0070] Regarding the effect of twin-screw extrusion temperature on the dispersibility of core-shell nanoparticles in the matrix, insufficient mixing occurs when the metering section temperature is below 165°C, while thermal degradation of the block copolymer occurs above 185°C. The optimal metering section temperature is 170–180°C, with a screw speed of 100–150 r / min. Injection mold temperatures between 35 and 50°C result in smooth product surfaces and precise dimensions. It is particularly important to note that the thermal stability of the poly(2-(dimethylamino)ethyl methacrylate) brush shell on the surface of the core-shell nanoparticles is a key factor affecting the upper limit of the processing temperature. Thermogravimetric analysis shows that the polymer brush begins to decompose above 220°C; therefore, the maximum temperature during the entire melt processing should not exceed 185°C, and the residence time of the material in the extruder should be controlled within 5 minutes to avoid loss of pH responsiveness due to polymer brush thermal degradation. Under the optimized processing conditions, the pH responsive zinc release ratio of the core-shell nanoparticles decreased only slightly from 4.3 times to 4.0 times before and after processing, indicating that the effect of the processing on the pH responsiveness is negligible.

[0071] The optimal polymerization effect was achieved when the catalyst stannous octoate was used at 0.03% to 0.08% of the total monomer mass during the synthesis of block copolymers. At a dosage of 0.05%, the conversion rate exceeded 95%, the molecular weight was 112,000, and the distribution index was 1.45. Increasing the initiator lauryl alcohol dosage from 0.1% to 0.5% gradually decreased the number-average molecular weight from 148,000 to 68,000, allowing for flexible adjustment based on the target molecular weight. Regarding the effect of silicon doping in zinc-silicon co-doping, increasing the molar ratio of silicate to phosphate from 8% to 25% increased the lattice defect density and gradually increased the zinc dissolution rate under acidic conditions; however, exceeding 30% altered the hydroxyapatite crystal structure, resulting in a calcium silicate impurity phase, which was detrimental to the material's biocompatibility. Therefore, the preferred silicon doping dosage in this invention is 8% to 25%.

[0072] The reason why the composite material of this invention can achieve synergistic optimization of mechanical properties, degradation rate, and intelligent antibacterial function lies in the construction of a three-level synergistic system from the molecular level to the nanoparticle level and then to the macroscopic level of the composite material, such as... Figure 2 As shown, at the molecular level, the L-type polylactic acid-polycaprolactone block copolymer arranges rigid and flexible segments alternately through covalent bonds, forming a nanoscale microphase separation structure. The continuous polylactic acid phase provides strength and modulus, while the dispersed polycaprolactone phase, as an energy dissipation domain, promotes the craze-shear band transition, fundamentally solving the problem of brittle fracture caused by insufficient interfacial adhesion in physical blending systems. At the nanoparticle level, the zinc-silicon co-doped core of the core-shell structured nano-hydroxyapatite provides a channel for accelerated zinc ion dissolution through lattice vacancies generated by silicon doping. The pH-responsive behavior of the poly(2-(dimethylamino)ethyl methacrylate) shell enables polymer brush expansion in acidic degradation environments and collapse in neutral environments, constituting an intelligent on / off valve mechanism that allows zinc ion release and degradation in the acidic microenvironment to form a closed-loop positive feedback regulation. At the composite material level, the triblock compatibilizer interacts with the tertiary amine groups in the polymer brush on the surface of the core-shell nanoparticles through hydrogen bonds formed by the polyethylene glycol intermediate block and the polylactic acid blocks at both ends extending into the matrix phase, firmly anchoring the nanoparticles in the matrix to achieve uniform dispersion. The integration of the three-level synergistic system enables the material of this invention to meet the mechanical requirements of the suture device while achieving comprehensive performance including controllable degradation cycle, stable degradation process, intelligent on-demand antibacterial properties, and full biosafety. This provides a material basis and technical support for the clinical application of biodegradable suture devices that do not require removal after circumcision.

[0073] The preparation process of the composite material of this invention is highly compatible with the existing industrialization process of polylactic acid medical devices. The block copolymer synthesis can be completed in a conventional stainless steel reactor. The coprecipitation of core-shell nanoparticles and surface-initiated atom transfer radical polymerization are both mature chemical unit operations. Melt blending granulation and injection molding can be implemented using existing standard equipment, which has good feasibility and economy for industrial scale-up.

[0074] The above description is merely a preferred embodiment of the present invention and is not intended to limit the present invention. Any modifications, equivalent substitutions, or improvements made within the spirit and principles of the present invention should be included within the scope of the claims of the present invention.

Claims

1. A circumcision suture device made of controllable degradable polylactic acid-based composite material, characterized in that, The product comprises the following components: 100 parts by weight of L-polylactic acid-polycaprolactone block copolymer as the matrix resin, wherein the block copolymer is prepared by sequential ring-opening polymerization of L-lactide and caprolactone, the molar ratio of L-lactide to caprolactone being 70:30 to 85:15, and the number average molecular weight being 80,000 to 150,000; and 3 to 10 parts by weight of pH-responsive core-shell structured nano-hydroxyapatite, wherein the core-shell structured nano-hydroxyapatite is grown by atom transfer radical polymerization of zinc-silicon co-doped nano-hydroxyapatite cores and surfaces. The brush shell is composed of a poly(2-(dimethylamino)ethyl methacrylate) brush shell, wherein zinc ions account for 3% to 8% of the molar number of calcium ions and silicate ions account for 8% to 25% of the molar number of phosphate ions in the zinc-silicon co-doped nano-hydroxyapatite core, and the core particle size is 30 to 60 nm. The number average molecular weight of the poly(2-(dimethylamino)ethyl methacrylate) brush shell is 3000 to 8000. 1 to 5 parts by weight of L-polylactic acid-polyethylene glycol-L-polylactic acid triblock compatibilizer and 2 to 8 parts by weight of acetylated tributyl citrate plasticizer are also present.

2. The composite material according to claim 1, characterized in that, The pH-responsive core-shell structured nano-hydroxyapatite exhibits a zinc ion release rate at pH 5.5 that is 3 to 5 times higher than that at pH 7.

4.

3. The composite material according to claim 1, characterized in that, The block copolymer has a molar ratio of L-lactide to caprolactone of 78:22 to 82:18, a number-average molecular weight of 100,000 to 130,000, and a molecular weight distribution index of 1.2 to 1.

6.

4. The composite material according to claim 1, characterized in that, The zinc-silicon co-doped nano-hydroxyapatite core contains zinc ions accounting for 4% to 6% of the molar number of calcium ions, silicate ions accounting for 12% to 18% of the molar number of phosphate ions, and the core particle size is 35 to 55 nm.

5. The composite material according to claim 1, characterized in that, The composite material has a tensile strength of 48 to 62 MPa, an elongation at break of 18% to 35%, and a complete in vivo degradation and absorption period of 8 to 12 weeks in a phosphate buffer solution at 37°C.

6. A method for preparing the composite material according to any one of claims 1 to 5, characterized in that, Includes the following steps: Step 1, Block copolymer synthesis: Using stannous octoate as a catalyst and lauryl alcohol as an initiator, L-lactide is first ring-opened polymerized at 160-170°C for 2-4 hours under nitrogen protection, and then caprolactone is added and polymerization continues at 140-155°C for 3-5 hours; Step 2, Core-shell nanoparticle preparation: Zinc-silicon co-doped nano-hydroxyapatite cores are prepared by co-precipitation, and atom transfer radical polymerization initiation sites are constructed by grafting with silane coupling agent and 2-bromoisobutyryl bromide. 2-(dimethylamino)ethyl methacrylate is polymerized on the surface of the nanoparticles to grow a polymer brush shell using cuprous bromide and pentamethyldiethylenetriamine as a catalytic system; Step 3, Melt blending and granulation: The block copolymer, core-shell nanoparticles, triblock compatibilizer, and plasticizer are melt-blended and granulated in a twin-screw extruder at 165-185°C; Step 4, Injection molding.

7. The preparation method according to claim 6, characterized in that, The conditions for the atom transfer radical polymerization described in step two are as follows: the solvent is a mixture of methanol and water in a volume ratio of 1:1; the molar ratio of monomer to initiation site is 20:1 to 50:1; the reaction temperature is 25 to 35°C; and the reaction time is 4 to 8 hours.

8. The preparation method according to claim 6, characterized in that, The method for constructing initiation sites in step two includes: firstly, zinc-silicon co-doped nano-hydroxyapatite is grafted onto gamma-aminopropyltriethoxysilane under reflux at 65 to 75°C for 3 to 5 hours to achieve surface amino functionalization; then, triethylamine is used as an acid-binding agent and reacted with 2-bromoisobutyryl bromide under ice bath and nitrogen protection for 8 to 16 hours to form bromoisobutyryl initiation sites on the surface of the nanoparticles.

9. The preparation method according to claim 6, characterized in that, The preparation method of zinc-silicon co-doped nano-hydroxyapatite core in step two includes: co-precipitating a mixed solution of calcium nitrate and zinc nitrate, a diammonium hydrogen phosphate solution and a tetraethyl orthosilicate solution under conditions of pH 10.0 to 10.5 and 55 to 65°C, aging, washing, drying and then calcining at 350 to 450°C for 1 to 3 hours.

10. The preparation method according to claim 6, characterized in that, In step three, the temperature of the feeding section of the twin-screw extruder is 150 to 160°C, the temperature of the compression section is 165 to 175°C, the temperature of the metering section is 170 to 180°C, and the temperature of the die head is 175 to 185°C; in step four, the temperature of the injection molding barrel is 155 to 180°C, and the temperature of the die is 35 to 50°C.

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