Non-absorbable suture with excellent biocompatibility and preparation method thereof
By using a three-layer composite suture design, combining polytetrafluoroethylene fiber, polyetheretherketone mesh, and phosphocholine zwitterionic polymer, the biocompatibility and mechanical stability issues of non-absorbable sutures are solved, achieving high biocompatibility and long-term mechanical stability.
Patent Information
- Authority / Receiving Office
- CN · China
- Patent Type
- Applications(China)
- Current Assignee / Owner
- BIOPAG (CHONGQING) BIOTECHNOLOGY CO LTD
- Filing Date
- 2026-01-27
- Publication Date
- 2026-05-01
AI Technical Summary
Existing non-absorbable sutures have shortcomings in terms of biocompatibility and long-term mechanical stability, and are prone to causing inflammatory reactions and mechanical property degradation.
The suture employs a three-layer composite structure, with a core layer of polytetrafluoroethylene fiber, a transition layer of polyetheretherketone mesh, and a surface functional layer of covalently grafted phosphocholine zwitterionic polymer. Each layer is fixed by electrospinning interlacing and vapor deposition polymerization, thereby improving biocompatibility and structural stability.
It significantly improves the biocompatibility of sutures, reduces the risk of inflammatory response, and maintains long-term mechanical stability, with a tensile strength retention rate as high as 90.3-97.8% and a CD68 positivity rate as low as 4.3-8.2%.
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Abstract
Description
Technical Field
[0001] This application relates to the field of medical materials technology, specifically to a non-absorbable suture with excellent biocompatibility and its preparation method. Background Technology
[0002] Medical suture materials are biomaterials used in surgery to close wounds and connect tissues. They are mainly divided into absorbable sutures and non-absorbable sutures. Absorbable sutures gradually degrade and are absorbed by tissues through enzymatic or hydrolytic processes, eliminating the need for secondary surgery for removal. They are suitable for short-term wound closure scenarios. Non-absorbable sutures, on the other hand, maintain their structural integrity in the body for a long time without degradation or absorption. They provide durable mechanical support for wound healing and are currently widely used in orthopedic surgery, suturing of internal organs, and suturing of areas with high skin tension, where long-term stable fixation is required.
[0003] However, with the advancement of medical technology, the performance requirements for non-absorbable sutures in clinical practice are becoming increasingly stringent. They not only need excellent long-term mechanical stability but also good biocompatibility to reduce adverse reactions such as inflammation and tissue rejection after implantation. Currently, the mainstream non-absorbable sutures on the market mainly use synthetic polymer materials such as polypropylene and polyester. These materials have advantages such as high strength and good degradation resistance, but their surface chemical properties have poor compatibility with human tissue cells, easily causing protein adsorption, platelet activation, and macrophage aggregation, leading to local inflammatory reactions, which may affect wound healing in severe cases. Furthermore, after long-term implantation, some materials may experience mechanical property degradation due to interfacial interactions, failing to provide sustained stable support.
[0004] To improve the biocompatibility of non-absorbable sutures, various improvement measures have been disclosed in the prior art, such as coating the suture surface with a biocompatible coating or plasma-modifying the suture surface. However, coating methods have problems such as weak adhesion between the coating and the substrate and easy peeling off. After long-term implantation, the biocompatibility will decrease due to coating failure. While simple plasma modification can improve surface hydrophilicity in the short term, it lacks durable and stable functional groups and cannot fundamentally inhibit protein adsorption and inflammatory response.
[0005] Therefore, developing a non-absorbable suture that combines high biocompatibility with excellent long-term mechanical stability has become an urgent problem to be solved in this field. Summary of the Invention
[0006] To overcome the problems of poor biocompatibility and poor long-term mechanical stability of existing non-absorbable sutures, this application provides a non-absorbable suture with excellent biocompatibility and its preparation method.
[0007] In a first aspect, this application provides a non-absorbable suture with excellent biocompatibility, employing the following technical solution: A non-absorbable suture with excellent biocompatibility includes a core layer, a transition layer, and a surface functional layer; the core layer and the transition layer are fixed by electrospinning and interweaving, and the surface functional layer is fixed to the surface of the transition layer by vapor deposition polymerization; The core layer is polytetrafluoroethylene fiber with a diameter of 20-50 μm and a porosity of <5%; the transition layer is polyetheretherketone with a mesh pore size of 10-15 μm and a thickness of 5-10 μm; and the surface functional layer is a zwitterionic polymer covalently grafted with phosphocholine with a thickness of 0.5-1.2 μm.
[0008] This application provides a medical non-absorbable suture with a three-layer composite structure comprising a core layer, a transition layer, and a surface functional layer. The core layer and the transition layer are fixed by an electrospinning interlacing process, which effectively utilizes the partial compatibility of polytetrafluoroethylene and polyetheretherketone in the molten state to form a tight interlacing structure and avoid interlayer delamination. The surface functional layer is fixed by a vapor phase deposition polymerization (glow discharge polymerization) process, which covalently grafts phosphocholine groups onto the active sites on the surface of the transition layer, further improving the structural stability. In the three-layer composite structure of this application, polytetrafluoroethylene fiber is used as the core layer, which has excellent resistance to biodegradation, mechanical strength and chemical stability, ensuring that the suture does not degrade after long-term implantation and continuously provides stable tensile support. Polyetheretherketone (PEEK) is a biocompatible polymer material whose grid structure can provide space for tissue cell growth, playing a role in stress buffering, interface connection and promoting the integration of the suture with the surrounding tissue interface. After argon plasma treatment of the PEEK grid, the surface energy of the transition layer can be increased to 60-65 mN / m, which can enhance its interweaving and bonding force with the core layer on the one hand, and introduce active groups on the other hand, providing reaction sites for covalent grafting of the surface functional layer and ensuring the stability of the functional layer. The surface functional layer is a zwitterionic polymer formed by glow discharge polymerization of phosphorylcholine monomers on the surface of the transition layer. Phosphorylcholine is a core component of human cell membranes and possesses excellent biocompatibility. It can mimic the hydrophilic-hydrophobic balance structure of the cell membrane surface, significantly inhibiting protein adsorption and platelet activation, and reducing the risk of inflammatory responses. The zwitterionic polymer can be covalently grafted onto the surface of the transition layer, resulting in a strong bond that is not easily detached, ensuring functional stability after long-term implantation. In summary, the non-absorbable suture provided in this application has advantages such as good biocompatibility and excellent long-term mechanical stability. After implantation, it can reduce tissue rejection and provide continuous and stable mechanical support for suture sites requiring long-term healing.
[0009] Optionally, the zwitterionic polymer covalently grafted with phosphoric acid choline is selected from polymethacryloyloxyethyl phosphoric acid choline, poly(methacryloyloxyethyl phosphoric acid choline-hydroxyethyl methacrylate), and poly(methacryloyloxyethyl phosphoric acid choline-dimethyl itaconic acid betaine).
[0010] Optionally, the zwitterionic polymer covalently grafted with phosphoric acid choline is poly(methacryloyloxyethyl phosphoric acid choline-itaconic acid dimethyl ester betaine).
[0011] Optionally, the tensile strength of the polytetrafluoroethylene fiber is ≥4N.
[0012] Secondly, this application provides a method for preparing a non-absorbable suture, comprising the following steps: A polyether ether ketone solution was electrospinned onto the surface of polytetrafluoroethylene fibers to form a transition layer with a thickness of 5-10 μm and a mesh pore size of 10-15 μm on the surface of the core fiber. The core fiber covering the transition layer is subjected to plasma treatment to activate the surface of the transition layer, and the surface energy reaches 60-65mN / m. The surface-activated composite fibers are placed in an atmosphere containing phosphorylcholine monomer for glow discharge polymerization, thereby covalently grafting a surface functional layer with a thickness of 0.5-1.2 μm onto the transition layer surface, ultimately obtaining a non-absorbable suture.
[0013] Optionally, the polyetheretherketone solution is prepared using N-methylpyrrolidone with a concentration of 12-18 wt%.
[0014] Optionally, the electrospinning voltage is 15-20kV and the distance is 10-15cm.
[0015] Optionally, the plasma treatment power is 80-120W and the time is 3-8min.
[0016] Optionally, the glow discharge polymerization temperature is 70-90℃ and the power is 100-150W.
[0017] In summary, this application has the following beneficial effects: 1. This application provides a medical non-absorbable suture with a three-layer composite structure comprising a core layer, a transition layer, and a surface functional layer. The core layer of the suture is made of high-purity polytetrafluoroethylene fiber, which can provide stable mechanical support. The polyetheretherketone (PEEK) mesh of the transition layer can buffer radial stress, avoid the mechanical property degradation of the core layer due to local stress concentration, and also provide reaction sites for covalent grafting of the surface functional layer, ensuring the stability of the functional layer. The surface functional layer is made of a phosphoric acid choline zwitterionic polymer with a cell membrane-like structure, which is fixed by covalent grafting, which can significantly improve the biocompatibility of the suture and reduce the risk of inflammatory response.
[0018] 2. In the medical non-absorbable suture provided in this application, the core layer and the transition layer are fixed by electrospinning and interlacing, and the transition layer and the surface functional layer are covalently grafted together. The interfaces of each layer are tightly bonded, and there is no risk of interlayer delamination. The entire suture structure is not prone to structural damage or component loss during long-term implantation in the body, and has high biocompatibility.
[0019] 3. The non-absorbable sutures provided in this application showed a CD68 positivity rate of only 4.3-8.2% and a tensile strength retention rate of 90.3-97.8% after 180 days of implantation. Detailed Implementation
[0020] This application provides a non-absorbable suture with excellent biocompatibility, comprising a core layer, a transition layer, and a surface functional layer; the core layer is polytetrafluoroethylene fiber with a diameter of 20-50 μm and a porosity of <5%; the transition layer is polyetheretherketone with a mesh pore size of 10-15 μm and a thickness of 5-10 μm; the surface functional layer is an amphoteric polymer covalently grafted with phosphocholine with a thickness of 0.5-1.2 μm; the core layer and the transition layer are fixed by electrospinning and interweaving, and the surface functional layer is fixed to the surface of the transition layer by vapor deposition polymerization.
[0021] The zwitterionic polymer covalently grafted with phosphoric acid choline is selected from polymethacryloyloxyethyl phosphoric acid choline, poly(methacryloyloxyethyl phosphoric acid choline-hydroxyethyl methacrylate), and poly(methacryloyloxyethyl phosphoric acid choline-itaconic acid dimethyl betaine).
[0022] This application provides a method for preparing a non-absorbable suture with excellent biocompatibility, comprising the following steps: (1) Dissolve polyetheretherketone powder in N-methylpyrrolidone to prepare a 12-18wt% polyetheretherketone solution. Use polytetrafluoroethylene core fiber as the receiving substrate and use electrospinning equipment to deposit a transition layer with a thickness of 5-10μm and a mesh pore size of 10-15μm on the surface of the core fiber under the conditions of 15-20kV voltage and 10-15cm receiving distance. (2) Place the core fiber covering the transition layer in a plasma treatment device, introduce argon gas (purity ≥99.99%), and perform surface activation under the conditions of 80-120W power and 3-8min treatment time, so that the surface energy of the transition layer can reach 60-65mN / m; (3) The activated composite fiber is transferred to the glow discharge polymerization reactor, and after being evacuated to 10-50 Pa, an atmosphere containing phosphorylcholine monomer is introduced. The glow discharge polymerization reaction is carried out at 70-90℃ and 100-150W power for 1.5-3 hours, thereby forming a covalently grafted surface functional layer on the surface of the transition layer, and finally obtaining a non-absorbing suture with a three-dimensional composite structure.
[0023] The raw materials, reagents, solvents, etc. used in this application are all commercially available.
[0024] The following describes this application in further detail with reference to preparation examples, embodiments, and performance testing. Preparation Example 1
[0025] Preparation Example 1 provides a methacryloyloxyethyl phosphorylcholine monomer.
[0026] The preparation method of the above-mentioned methacryloyloxyethyl phosphorylcholine monomer is as follows: 100 mL of anhydrous ethanol is added to a dry 500 mL three-necked flask under a nitrogen atmosphere at 25 °C, and nitrogen is purged three times; then 0.1 mol of phosphoric acid choline and 0.12 mol of triethylamine are added sequentially and stirred until completely dissolved; then 0.1 mol of 2-chloroethyl methacrylate is slowly added dropwise at a rate of 1 mL / min; after the addition is complete, the temperature is raised to 50 °C and stirred at a constant temperature for 36 h; finally, the reaction solution is cooled to room temperature, filtered, distilled under reduced pressure, recrystallized from ethyl acetate, washed, and dried to obtain white needle-like crystals of methacryloyloxyethyl phosphorylcholine monomer. Preparation Example 2
[0027] Preparation Example 2 provides a dimethyl itaconic acid betaine monomer.
[0028] The preparation method of the above-mentioned itaconic acid dimethyl ester betaine monomer is as follows: (1) At 25°C and under a nitrogen atmosphere, 100 mL of acetonitrile was added to a dry 250 mL three-necked flask, and nitrogen was purged three times. Then, 0.1 mol of dimethyl itaconic acid was added and stirred until completely dissolved. Then, 0.105 mol of N,N-dimethylaminopropylamine was slowly added dropwise while stirring continuously. After the addition was complete, the temperature was raised to 45°C and the reaction was stirred at a constant temperature for 14 h while maintaining nitrogen protection during the reaction. After the reaction was completed, the reaction solution was cooled to room temperature, concentrated to 1 / 3 of the original volume by vacuum distillation, filtered, and the filtrate (containing the tertiary amine intermediate) was collected.
[0029] (2) Add 0.11 mol sodium chloroacetate to the filtrate containing the above tertiary amine intermediate and stir to disperse evenly; then adjust the pH of the system to 7.8 with 0.1 mol / L sodium hydroxide solution; then raise the temperature to 60℃ and stir the reaction at a constant temperature for 20 h; after the reaction is completed, cool the reaction solution to room temperature, add anhydrous ethanol to dilute, filter, distill under reduced pressure and recrystallize to obtain a white powdery itaconic acid dimethyl betaine monomer. Example 1
[0030] Example 1 provides a non-absorbable suture with excellent biocompatibility.
[0031] The method for preparing the above-mentioned biocompatible non-absorbable sutures includes the following steps: (1) Dissolve polyetheretherketone powder in N-methylpyrrolidone to prepare a 15wt% polyetheretherketone solution. Use polytetrafluoroethylene fiber (diameter 30±2μm, porosity 3%, tensile strength 5.0N) as the receiving substrate. Use electrospinning equipment to deposit a transition layer with a thickness of 8μm and a mesh pore size of 10-15μm on the surface of the core fiber under the conditions of 20kV voltage and 15cm receiving distance. (2) The core fiber covering the transition layer is placed in a plasma processing device, and argon gas with a purity of ≥99.99% is introduced. Surface activation is carried out under the conditions of 100W power and 5min processing time, so that the surface energy of the transition layer can reach 65mN / m. (3) The activated composite fiber is transferred to the glow discharge polymerization reactor, and after being evacuated to 30 Pa, an atmosphere of methacryloyloxyethyl phosphorylcholine monomer is introduced. The glow discharge polymerization reaction is carried out at 80 °C and 120 W power for 2 h, thereby forming a surface functional layer with a covalent graft thickness of 1 μm on the surface of the transition layer, and finally obtaining a non-absorbable suture with a three-dimensional composite structure. Example 2
[0032] Example 2 provides a non-absorbable suture with excellent biocompatibility.
[0033] The difference between the above embodiment and Embodiment 1 is that the thickness of the transition layer is 5 μm and the thickness of the surface functional layer is 0.5 μm. Example 3
[0034] Example 3 provides a non-absorbable suture with excellent biocompatibility.
[0035] The difference between the above embodiment and Embodiment 1 is that the thickness of the transition layer is 10 μm and the thickness of the surface functional layer is 1.2 μm. Example 4
[0036] Example 4 provides a non-absorbable suture with excellent biocompatibility.
[0037] The difference between the above embodiments and Embodiment 1 is that the methacryloyloxyethyl phosphorylcholine monomer is replaced with methacryloyloxyethyl phosphorylcholine monomer and hydroxyethyl methacrylate monomer in a weight ratio of 1:1. Example 5
[0038] Example 5 provides a non-absorbable suture with excellent biocompatibility.
[0039] The difference between the above embodiments and Embodiment 1 is that the methacryloyloxyethyl phosphorylcholine monomer is replaced with a methacryloyloxyethyl phosphorylcholine monomer and a dimethyl itaconic acid betaine monomer in a weight ratio of 1:1. Example 6
[0040] Example 6 provides a non-absorbable suture with excellent biocompatibility.
[0041] The difference between the above embodiment and Example 1 is that the weight ratio of methacryloyloxyethyl phosphorylcholine monomer and itaconic acid dimethyl betaine monomer is 1:0.8. Example 7
[0042] Example 7 provides a non-absorbable suture with excellent biocompatibility.
[0043] The difference between the above embodiment and Example 1 is that the weight ratio of methacryloyloxyethyl phosphorylcholine monomer and itaconic acid dimethyl betaine monomer is 1:0.5. Example 8
[0044] Example 8 provides a non-absorbable suture with excellent biocompatibility.
[0045] The difference between the above embodiment and Example 1 is that the weight ratio of methacryloyloxyethyl phosphorylcholine monomer and itaconic acid dimethyl betaine monomer is 1:0.3. Example 9
[0046] Example 9 provides a non-absorbable suture with excellent biocompatibility.
[0047] The difference between the above embodiment and Example 1 is that the weight ratio of methacryloyloxyethyl phosphorylcholine monomer and itaconic acid dimethyl betaine monomer is 1:1.5. Comparative Example 1
[0048] Comparative Example 1 provides a non-absorbable suture with excellent biocompatibility.
[0049] The difference between the above comparative example and Example 1 is that the non-absorbable suture does not contain a surface functional layer. Comparative Example 2
[0050] Comparative Example 2 provides a non-absorbable suture with excellent biocompatibility.
[0051] The difference between the above comparative example and Example 1 is that the thickness of the transition layer is 3 μm and the thickness of the surface functional layer is 2 μm. Comparative Example 3
[0052] Comparative Example 3 provides a non-absorbable suture with excellent biocompatibility.
[0053] The difference between the above embodiment and Embodiment 1 is that the polytetrafluoroethylene fiber is replaced with polypropylene fiber. Comparative Example 4
[0054] Comparative Example 4 provides a non-absorbable suture with excellent biocompatibility.
[0055] The difference between the above embodiment and Embodiment 1 is that the polytetrafluoroethylene fiber is replaced with polyester fiber. Performance testing
[0056] The performance of the non-absorbable sutures obtained in Examples 1-9 and Comparative Examples 1-2, as well as commercially available polypropylene and polyester sutures, was tested. The specific process is as follows: (1) 130 SPF-grade SD rats, weighing 200-250g, were selected, with half males and half females; they were randomly divided into 13 groups (experimental groups 1-9 and control groups 1-4). (2) Make a 2cm incision in the back muscle layer of SD rats and implant sterilized sutures into the muscle layer (the sutures of Examples 1-9 were used in experimental groups 1-9, the sutures of Comparative Examples 1-2 were used in control groups 1-2, and commercially available polypropylene fiber sutures and polyester fiber sutures were used in control groups 3-4 respectively). The skin was sutured in layers and the rats were fed normally for 180 days.
[0057] (3) After 180 days, the rats were sacrificed, and the muscle tissue within 1 cm around the suture (including the suture-tissue interface) was completely dissected and immediately fixed in 4% paraformaldehyde fixative for 24 h, and then paraffin embedded and sectioned. (4) The sections were dewaxed and hydrated, antigen retrieval was performed, blocking was performed, primary antibody incubation (adding anti-CD68 monoclonal antibody) was performed, secondary antibody incubation (adding goat anti-mouse IgG) was performed, DAB staining was performed, counterstaining and dehydration were performed. Then, under an optical microscope, 10 non-overlapping suture-tissue interface fields were randomly selected, and the number of CD68 positive cells (brownish-yellow stained cells) and the total number of cells (total number of cell nuclei stained with hematoxylin) in each field were counted. The CD68 positivity rate of each group of mice was calculated and the average value was taken. The results are shown in Table 1 below. The formula for calculating the CD68 positivity rate is as follows: CD68 positivity rate = (average number of CD68-positive cells in 10 fields of view / average total number of cells in 10 fields of view) × 100%.
[0058] Note: CD68 is a specific marker of macrophage activation. Using immunohistochemistry, anti-CD68 specific antibodies are used to bind to positive cells in the tissue surrounding the suture. After the colorimetric reaction, the intensity of the inflammatory response induced after suture implantation is reflected by statistically analyzing the proportion of CD68 positive cells to the total number of cells. The lower the positive rate, the better the biocompatibility.
[0059] (4) The sutures of the experimental group and the control group were completely removed, sterilized with ethylene oxide, and rinsed three times with physiological saline after sterilization. Then, the tensile strength was tested using a universal testing machine. The tensile strength retention rate was calculated based on the initial tensile strength of the sutures. The results are shown in Table 1 below. The formula for calculating the tensile strength retention rate is as follows: Tensile strength retention rate = (remaining tensile strength after 180 days / initial tensile strength) × 100%.
[0060] Table 1 Performance test results of Examples 1-9, Comparative Examples 1-2, and commercially available sutures
[0061] According to the test results in Table 1, the CD68 positivity rate of the non-absorbable sutures obtained in Examples 1-9 was 4.3-8.2% and the tensile strength retention rate was 90.3-97.8% after 180 days of implantation. In contrast, the CD68 positivity rate of the non-absorbable sutures obtained in Comparative Examples 1-2 was as high as 12.3-14.5% and the tensile strength retention rate was 83.2-87.4% after 180 days of implantation. Commercially available polypropylene and polyester sutures showed a CD68 positivity rate as high as 26.0-27.8% and a tensile strength retention rate of only 66.5-72.3% after 180 days of implantation. Therefore, this application demonstrates that by using polytetrafluoroethylene fiber as the core layer and sequentially setting a transition layer with a thickness of 5-10 μm and a surface functional layer with a thickness of 0.5-1.2 μm on its surface, a non-absorbable suture with good biocompatibility and excellent long-term mechanical stability can be obtained. This suture is suitable for medical suturing scenarios requiring durable support and has excellent application prospects.
[0062] Although the present invention has been described in detail above with general descriptions and specific embodiments, modifications or improvements can be made to it, which will be obvious to those skilled in the art. Therefore, all such modifications or improvements made without departing from the spirit of the present invention fall within the scope of protection claimed by the present invention.
Claims
1. A non-absorbable suture with excellent biocompatibility, characterized in that, It includes a core layer, a transition layer, and a surface functional layer; the core layer and the transition layer are fixed by electrospinning interweaving, and the surface functional layer is fixed to the surface of the transition layer by vapor deposition polymerization. The core layer is polytetrafluoroethylene fiber with a diameter of 20-50 μm and a porosity of <5%; the transition layer is polyetheretherketone with a mesh pore size of 10-15 μm and a thickness of 5-10 μm; and the surface functional layer is a zwitterionic polymer covalently grafted with phosphocholine with a thickness of 0.5-1.2 μm.
2. The non-absorbable suture according to claim 1, characterized in that, The zwitterionic polymer covalently grafted with phosphoric acid choline is selected from polymethacryloyloxyethyl phosphoric acid choline, poly(methacryloyloxyethyl phosphoric acid choline-hydroxyethyl methacrylate), and poly(methacryloyloxyethyl phosphoric acid choline-itaconic acid dimethyl betaine).
3. The non-absorbable suture according to claim 1, characterized in that, The zwitterionic polymer covalently grafted with phosphoric acid choline is poly(methacryloyloxyethyl phosphorylcholine-itaconic acid dimethyl betaine).
4. The non-absorbable suture according to claim 1, characterized in that, The tensile strength of the polytetrafluoroethylene fiber is ≥4N.
5. The method for preparing non-absorbable sutures according to any one of claims 1-4, characterized in that, Includes the following steps: A polyether ether ketone solution was electrospinned onto the surface of polytetrafluoroethylene fibers to form a transition layer with a thickness of 5-10 μm and a mesh pore size of 10-15 μm on the surface of the core fiber. The core fiber covering the transition layer is subjected to plasma treatment to activate the surface of the transition layer, and the surface energy reaches 60-65mN / m. The surface-activated composite fibers are placed in an atmosphere containing phosphorylcholine monomer for glow discharge polymerization, thereby covalently grafting a surface functional layer with a thickness of 0.5-1.2 μm onto the transition layer surface, ultimately obtaining a non-absorbable suture.
6. The method for preparing non-absorbable sutures according to claim 1, characterized in that, The polyetheretherketone solution is prepared using N-methylpyrrolidone with a concentration of 12-18 wt%.
7. The method for preparing non-absorbable sutures according to claim 1, characterized in that, The electrospinning voltage is 15-20kV and the distance is 10-15cm.
8. The method for preparing non-absorbable sutures according to claim 1, characterized in that, The plasma treatment power is 80-120W, and the time is 3-8 minutes.
9. The method for preparing non-absorbable sutures according to claim 1, characterized in that, The glow discharge polymerization temperature is 70-90℃ and the power is 100-150W.