Injectable zwitterionic conductive hydrogel as well as preparation method and application thereof

Abstract

The invention discloses injectable zwitterionic conductive hydrogel as well as a preparation method and application thereof. The hydrogel is formed by self-assembly of a zwitterionic copolymer and a PEDOT / PSS conductive polymer, the zwitterionic copolymer is sulfobetaine metacrylic acid ester or a copolymer formed by polymerizing sulfobetaine metacrylic acid ester and other comonomers. According to the invention, the zwitterionic copolymer containing sulfobetaine methacrylate and the PEDOT / PSS conductive polymer solution can be quickly self-assembled to form the hydrogel through simple mechanical mixing, and the hydrogel is endowed with excellent injectability and self-repairing performance based on the dynamic crosslinking characteristic of electrostatic interaction. The hydrogel has excellent immunocompatibility and conductivity, and can play a role in various application scenes of bioelectronics.

Description

Technical Field

[0001] The present invention relates to the technical field of medical materials, and in particular to an injectable zwitterionic conductive hydrogel and a preparation method and application thereof. Background Art

[0002] In recent years, the rapid development of wearable electronic devices, implantable neural interfaces, and intelligent medical systems has placed higher demands on soft conductive materials with excellent biocompatibility, conductivity, and injectability. Conductive hydrogels, which combine the softness of hydrogels with the electron transport capabilities of conductive materials, have become a hot topic in the research of next-generation bioelectronic materials. However, traditional conductive hydrogels often rely on strong covalent crosslinks or large structures, often lacking good injectability, making them unsuitable for minimally invasive surgical procedures.

[0003] Although PEDOT / PSS is regarded as the most promising conductive polymer in the biomedical field, CN119950824A discloses a conductive thermosensitive hydrogel and its preparation method and application, which specifically includes PEDOT:PSS and PluronicF127. It is a conductive thermosensitive hydrogel based on hydrogen bond interaction and is used as a submucosal filling material. PluronicF127 confers thermosensitivity, can be injected at room temperature, and quickly forms a solid structure at body temperature. After the operation, it can be removed by local cooling. The conductive properties of PEDOT:PSS support the electrocautery process and ensure smooth resection. The good conductivity of PEDOT:PSS helps to evenly distribute the current and avoid damage caused by local overheating. The hydrogen bond interaction between PEDOT:PSS and PluronicF127 significantly improves the mechanical strength of the hydrogel, ensuring that the hydrogel provides a long-lasting and stable mucosal lifting effect during surgery, providing a reliable and efficient support material for minimally invasive surgery.

[0004] CN114504683A discloses a composite conductive hydrogel prepared by physical crosslinking of acellular matrix hydrogel and PEDOT:PSS. The composite hydrogel has extremely high electron transfer efficiency and mechanical strength, can effectively restore the microenvironment of cell growth, has excellent biocompatibility, can promote tissue regeneration, and has good fluidity and injectability. It is not restricted by shape and size during use, and can adhere well to tissues after gelation, without the need for additional surgical suturing. The conductivity of the hydrogel can also be adjusted by adjusting the amount of PEDOT:PSS added, and electrical stimulation can be provided to cells by applying a weak external current, which can effectively promote cell proliferation, migration and functionalization. The composite hydrogel is expected to be prepared into a biomedical conductive material for application.

[0005] However, the stability and immunocompatibility of PEDOT:PSS in complex physiological environments are also limited, and it is easy to induce inflammatory reactions or signal attenuation. Zwitterionic polymers have been widely used in biomaterial design in recent years due to their unique neutral structure, excellent anti-protein adsorption, hydration ability and biocompatibility. However, the effective integration of zwitterionic polymers with PEDOT / PSS to achieve hydrogels with high conductivity, good injectability and biostability still faces challenges such as complex material design and difficulty in structural regulation.

[0006] Therefore, developing a new hydrogel system that combines injectability, zwitterionic properties and conductive properties is of great significance to promoting the development of fields such as bioelectronics, neural regulation and tissue repair. Summary of the Invention

[0007] The present invention addresses the difficulty of achieving injectability and immunocompatibility of conductive hydrogels and provides an injectable zwitterionic conductive hydrogel. The hydrogel is obtained by electrostatic self-assembly of a zwitterionic copolymer and PEDOT / PSS. Its dynamic cross-linking properties give it excellent injectability and self-healing properties. PEDOT / PSS and zwitterionic polymers give it broad application prospects in bioelectronics-related medical fields.

[0008] To achieve the above object, the technical solution adopted by the present invention is: An injectable zwitterionic conductive hydrogel, which is self-assembled from a zwitterionic copolymer and a PEDOT / PSS conductive polymer; The zwitterionic copolymer is a copolymer formed by polymerization of sulfobetaine methacrylate or sulfobetaine methacrylate and other comonomers.

[0009] In this invention, a zwitterionic copolymer containing sulfobetaine methacrylate and a PEDOT / PSS conductive polymer solution rapidly self-assemble to form a hydrogel through simple mechanical mixing. Dynamic crosslinking due to electrostatic interactions imparts excellent injectability and self-healing properties. Composed of zwitterionics with excellent immunocompatibility and conductive PEDOT / PSS polymers, this hydrogel exhibits excellent immunocompatibility and conductivity, making it suitable for a variety of bioelectronic applications.

[0010] The molecular weight of the zwitterionic copolymer is 20,000-2,000,000.

[0011] The weight percentage of sulfobetaine methacrylate in the zwitterionic copolymer is greater than 50%. Preferably, the weight percentage of sulfobetaine methacrylate is greater than 60%, greater than 70%, or greater than 80%. The higher the content, the more conducive it is to gel formation. For specific application scenarios, other comonomers with corresponding functions can be added to obtain hydrogels with corresponding functions.

[0012] The other comonomers include one or more of acrylate monomers and acrylamide monomers.

[0013] The mass ratio of the zwitterionic copolymer to the PEDOT / PSS conductive polymer is 50 to 0.5:1. Preferably, the mass ratio is 30 to 1:1; more preferably, the mass ratio is 20 to 5:1. A higher zwitterionic copolymer content improves the biocompatibility of the hydrogel, while a higher conductive polymer content improves the conductivity of the hydrogel. However, excessive zwitterionic content can reduce injectability and conductivity, while increased conductive polymer content can reduce gel strength.

[0014] The present invention also provides a method for preparing the injectable zwitterionic conductive hydrogel, comprising the steps of: Step 1, copolymerizing raw materials containing sulfobetaine methacrylate to prepare a zwitterionic copolymer; Step 2: mixing the zwitterionic copolymer freeze-dried powder with the PEDOT / PSS conductive polymer solution to obtain the injectable zwitterionic conductive hydrogel through self-assembly.

[0015] The copolymerization in step 1 contains an initiator, the copolymerization temperature is 60-90° C., the reaction is carried out for 6-18 hours, and the product is dialyzed and freeze-dried to obtain a copolymer.

[0016] The initiator includes one or more of sodium persulfate, ammonium persulfate, and potassium persulfate, and the mass fraction of the initiator in the aqueous solution is 0.1-5%.

[0017] The mixing in step 2 is carried out at room temperature and the self-assembly is carried out for more than 5 seconds. The viscosity of the system will continue to increase after mixing, and the self-assembly to form a hydrogel is usually completed in 3-10 minutes.

[0018] The present invention also provides the use of the injectable zwitterionic conductive hydrogel in the preparation of biomedical materials and bioelectronic materials.

[0019] For example, it can be used for immediate filling of tissue damage: injectable zwitterionic conductive hydrogels combine excellent injectability, electrical conductivity and immune compatibility imparted by zwitterionic polymers, and have unique advantages in minimally invasive repair after muscle injury. Unlike other tissues, muscle tissue has electrophysiological activity and can respond to nerve stimulation to generate action potentials and maintain normal motor function. After acute muscle injury, the electrical signal conduction path in the damaged area may be interrupted, affecting the recovery of tissue function. Therefore, constructing a scaffold material with conductivity is crucial to rebuilding the continuity of local electrical signals. The injectable zwitterionic conductive hydrogel proposed in the present invention can quickly fill the damaged area and form a continuous conductive network, which not only helps to restore the electrophysiological function of muscle tissue, but also provides structural support and physiological environment regulation in the early repair stage, thereby promoting muscle regeneration and functional reconstruction.

[0020] Another example is improving the long-term immune compatibility of implantable electrodes. Once implanted, these electrodes trigger a severe immune response to these foreign substances, which over time develops into a foreign body reaction. Unable to eliminate these foreign substances, collagen secreted by the tissue isolates the electrodes from the body, ultimately leading to decreased electrode performance and even failure. This reaction is long-lasting and persistent, causing long-term pain for the patient.

[0021] In some embodiments, the hydrogel of the present invention can be injected into the electrode implantation site, and then the electrode is implanted into the hydrogel. The hydrogel can serve as a bridge between the tissue and the electrode. Since the hydrogel provided by the present invention has excellent immune compatibility, it avoids the occurrence of acute inflammation and foreign body reaction, and effectively improves the long-term performance of the electrode.

[0022] Compared with the prior art, the present invention has the following beneficial effects: (1) The injectable zwitterionic conductive hydrogel in the present invention does not require sophisticated structural design. Only two polymers that are already available on the market or can be easily polymerized are used to form a hydrogel that is both conductive and injectable. This avoids the cost and biocompatibility issues caused by the introduction of complex components. The process is simple and easy to promote.

[0023] (2) The injectable zwitterionic conductive hydrogel of the present invention has excellent injectability and immunocompatibility, and shows broad application prospects in biomedical fields such as implantable bioelectronics and tissue repair. BRIEF DESCRIPTION OF THE DRAWINGS

[0024] In order to more clearly illustrate the technical solutions of the embodiments of the present application, the following is a brief introduction to the drawings required for use in the embodiments of the present application. It should be understood that the following drawings only show certain embodiments of the present application and therefore should not be regarded as limiting the scope. For ordinary technicians in this field, other relevant drawings can be obtained based on these drawings without creative work.

[0025] Figure 1 Strain scan of the injectable zwitterionic conductive hydrogel SPPH2 prepared in Example 1 ( Figure 1 A) and angular frequency sweep test ( Figure 1 Middle B) Results.

[0026] Figure 2 Statistical graphs of G' and tan δ of the injectable zwitterionic conductive hydrogels with different component ratios prepared in Example 1.

[0027] Figure 3 The injectable zwitterionic conductive hydrogel SPPH2 prepared in Example 1 was cycled under large and small strains to confirm the self-healing properties of the hydrogel.

[0028] Figure 4 The injection force of the injectable zwitterionic conductive hydrogel SPPH2 prepared in Example 1 ( Figure 4 A) and injection filling capabilities for different interfaces ( Figure 4 Middle B).

[0029] Figure 5 Raman spectra of the injectable zwitterionic conductive hydrogel SPPH2 and pure PEDOT / PSS prepared in Example 1.

[0030] Figure 6 Visualization results of molecular simulation electrostatic potential (ESP) of the injectable zwitterionic conductive hydrogel described in the present invention.

[0031] Figure 7 The ionic conductivity and electronic conductivity of the injectable zwitterionic conductive hydrogel prepared in Example 1.

[0032] Figure 8 This is the electrochemical impedance spectroscopy of the injectable zwitterionic conductive hydrogel prepared in Example 1.

[0033] Figure 9 The response frequency of the rat hind limbs after applying 0.1 V electrical stimulation signals of different frequencies (5 Hz, 10 Hz, 20 Hz, 50 Hz) to the sciatic nerve in different groups as described in Example 1.

[0034] Figure 10The rat hind limbs were electrically stimulated at different voltages (0.1 V, 0.3 V, 0.5 V, 0.7 V, and 0.9 V) at 1 Hz in different groups as described in Example 1, and the statistical graph of the rat hind limb swing amplitude under stimulation was recorded.

[0035] Figure 11 This is a photo of the acute tibialis anterior muscle injury model being severed and filled with SPPH2 hydrogel in Application Example 1.

[0036] Figure 12 The amplitude statistics of the electrophysiological model transmitted by muscle after the acute tibialis anterior muscle injury model in Application Example 1 was severed and filled with SPPH2 hydrogel.

[0037] Figure 13 This is a statistical chart showing the expression of representative inflammatory factors 14 days after DSPH and SPPH2 were implanted subcutaneously in mice in Application Example 2.

[0038] Figure 14 These are Masson trichrome sections, F4 / 80 fluorescence staining, and CD3 fluorescence staining images 28 days after DSPH and SPPH2 were implanted subcutaneously in mice in Application Example 2.

[0039] Figure 15 This is the signal-to-noise ratio of the electromyographic signals recorded by the surface electrodes, directly implanted electrodes, and SPPH2-processed electrodes in Application Example 2.

[0040] Figure 16 Masson tri-color slices of the implantation sites of directly implanted electrodes and SPPH2-treated electrodes in Application Example 2. DETAILED DESCRIPTION

[0041] In order to make the purpose, technical solutions and advantages of the present invention clearer, the present invention is further described in detail below in conjunction with embodiment. It should be understood that the specific embodiments described herein are only used to explain the present invention and are not intended to limit the present invention. Those skilled in the art will make modifications or equivalent substitutions based on understanding the technical solution of the present invention, without departing from the spirit and scope of the technical solution of the present invention, and all should be encompassed within the protection scope of the present invention.

[0042] The raw materials used in the following specific embodiments are all purchased from the market.

[0043] Example 1 Step 1: 10 wt% of sulfobetaine (SBMA) monomer was dissolved in deionized water, followed by the addition of potassium persulfate (1 wt% of the total monomer weight) as an initiator. After deoxygenation for 20 minutes under a nitrogen atmosphere, the reaction system was transferred to a 70°C oil bath and subjected to constant stirring for 12 hours. Following the reaction, the product was dialyzed to remove small molecule impurities and freeze-dried to obtain a polysulfobetaine (PSBMA) polymer powder with an average molecular weight of approximately 200,000.

[0044] Step 2: At room temperature, 50 mg, 100 mg, 150 mg, and 200 mg of freeze-dried PSBMA powder were added to a 1 mL PEDOT / PSS solution (1.1 wt%), yielding weight ratios of 4.5:1, 9:1, 1.36:1, and 18:1, respectively. Rapid stirring was then performed until the powder was completely dissolved in the PEDOT / PSS solution. As the powder dissolved, the viscosity of the system increased, forming injectable zwitterionic hydrogels after approximately 5 minutes. These hydrogels were designated SPPH1, SPPH2, SPPH3, and SPPH4, respectively.

[0045] Example 2 Step 1: 20 wt% of sulfobetaine (SBMA) monomer was dissolved in deionized water, followed by the addition of potassium persulfate (1 wt% of the total monomer weight) as an initiator. After deoxygenation for 20 minutes under a nitrogen atmosphere, the reaction system was transferred to a 70°C oil bath and subjected to constant stirring for 12 hours. Following the reaction, the product was dialyzed to remove small molecule impurities and freeze-dried to obtain a polysulfobetaine (PSBMA) polymer powder with an average molecular weight of approximately 600,000.

[0046] Step 2: At room temperature, add 200 mg of lyophilized PSBMA powder to a 1.1 wt% PEDOT / PSS solution (with a PSBMA:PEDOT / PSS mass ratio of 18:1). Rapid stirring is then performed until the powder is completely dissolved in the PEDOT / PSS solution. As the powder dissolves, the viscosity of the system increases, ultimately forming an injectable zwitterionic hydrogel.

[0047] Example 3 Step 1: 15 wt% sulfobetaine (SBMA) monomer and 5 wt% acrylic acid (AA) were dissolved in deionized water. Potassium persulfate (1 wt% of the total monomer weight) was then added as an initiator. After deoxygenation under nitrogen bubbling for 20 minutes, the reaction system was transferred to a 70°C oil bath and stirred for 12 hours. After completion of the reaction, the product was dialyzed to remove small molecule impurities and freeze-dried to obtain a polysulfobetaine-polyacrylic acid copolymer (PSBMA-PAA) powder with an average molecular weight of approximately 553,000.

[0048] Step 2: At room temperature, add 200 mg of freeze-dried PSBMA-PAA powder to a 1.1 wt% PEDOT / PSS solution (with a PSBMA:PEDOT / PSS mass ratio of 18:1). Rapid stirring is then performed until the powder is completely dissolved in the PEDOT / PSS solution. As the powder dissolves, the viscosity of the system increases, ultimately forming an injectable zwitterionic hydrogel.

[0049] Comparative Example 1 Step 1: Add 40 μL of sodium dodecyl sulfate (DBSA) to 1 mL of a 1.1 wt% PEDOT / PSS aqueous solution (with a mass ratio of DBSA to PEDOT / PSS of 3.63:1). Vortex the solution to mix thoroughly, then evacuate the solution under vacuum for 5 minutes. After standing at room temperature for 2 hours, the PEDOT / PSS solution forms a gel. The resulting gel is dialyzed in ultrapure water to remove excess DBSA. A cell disruptor is used to fragment the hydrogel into small particles with a diameter of 50 μm, yielding the injectable conductive hydrogel DSPH.

[0050] Performance Testing The injectable zwitterionic conductive hydrogel prepared in Example 1 was subjected to rheological characterization. All rheological tests were performed on a rotational rheometer (TA Instruments, Discovery HR-20) using a 20 mm diameter, 1500 μm gap plate fixture at 37°C. The sample was allowed to rest for 1 minute before testing, and then subjected to an oscillatory strain sweep (strain range: 0.1–2000%, fixed frequency: 10 rad / s) and an oscillatory frequency sweep (frequency range: 0.1–100 rad / s, fixed strain: 1%). The results are shown in Table 1. Figure 1 As shown, Figure 1 Strain scan of the injectable zwitterionic conductive hydrogel SPPH2 prepared in Example 1 ( Figure 1 A) and angular frequency sweep test ( Figure 1(B) Results. These curves show that the elastic and viscous properties of the material change with strain and angular frequency, with the elastic properties dominating, further demonstrating that the material exhibits solid-like behavior.

[0051] Figure 2 The strength of the hydrogels formed by different mass ratios of zwitterionic PSBMA and PEDOT / PSS components was investigated. Figure 2 The G' and tan δ statistical graphs of the injectable zwitterionic conductive hydrogels with different component ratios prepared in Example 1 show that the mechanical strength of the injectable conductive hydrogels increases continuously with the increase of PSBMA content.

[0052] Figure 3 The injectable zwitterionic conductive hydrogel SPPH2 prepared in Example 1 was cycled under high and low strains to confirm its self-healing properties. The self-healing ability of the SPPH2 hydrogel was explored. In rotational rheometer testing, strains of 1000% and 1% were applied to the hydrogel, respectively. High strains disrupt the hydrogel network, while low strains allow for self-repair. After three strain cycles, the hydrogel's mechanical properties remained unchanged, demonstrating its excellent self-healing properties.

[0053] Figure 4 The injection force of the injectable zwitterionic conductive hydrogel SPPH2 prepared in Example 1 ( Figure 4 A) and injection filling capabilities for different interfaces ( Figure 4 (B) The injection force of SPPH1, SPPH2, SPPH3, and SPPH4 hydrogels with different compositions was investigated. A 23G injection needle was used to inject the four hydrogels, and the injection force was measured using a universal testing machine. The results showed that all four hydrogel compositions had excellent injectability and could be injected with relatively low injection force. Figure 4 As shown in Figure B, this injectable hydrogel can effectively fill the tissue interface in some complex interfaces and gelatin-simulated tissues.

[0054] Figure 5 Raman spectra of the injectable zwitterionic conductive hydrogel SPPH2 prepared in Example 1 and pure PEDOT / PSS. The interaction between PSBMA and PEDOT / PSS was investigated. Raman spectroscopy reveals a significant blue shift in the oxidation peak of PEDOT / PSS, indicating a transition to a reduced state. This weakening of the interaction between PSS and PEDOT suggests that electrostatic interaction between PSBMA and PSS promotes PEDOT aggregation, ultimately forming the injectable conductive hydrogel.

[0055] Figure 6 The electrostatic potential diagram of the interaction between PSBMA and PEDOT / PSS polymer molecules is shown. It can be seen that there is obvious infiltration between the electron clouds of PSBMA and PEDOT / PSS, further demonstrating the existence of electrostatic interaction between PSBMA and PEDOT / PSS.

[0056] Figure 7 The ionic conductivity (Ri) and electronic conductivity (Re) of the prepared SPPH1, SPPH2, SPPH3, and SPPH4 hydrogels are shown. It can be seen that with the increase of PSBMA content, the conductivity decreases, which is inevitable because the corresponding conductive component will show a significant decrease.

[0057] Figure 8 The electrochemical impedance spectra of the SPPH1, SPPH2, SPPH3, and SPPH4 hydrogels prepared in Example 1 measured by an electrochemical workstation are shown. It can be seen that the electrochemical impedance increases with the increase of the PSBMA content.

[0058] Application Example 1 Representative injectable zwitterionic conductive hydrogels for neuroprotection and acute muscle injury repair 1. Animal surgery process This application example describes the ability of an injectable zwitterionic conductive hydrogel to protect nerves during nerve stimulation with a stimulating electrode and to restore muscle electrophysiological signal transmission in acute muscle injury. First, after anesthetizing SD rats, the sciatic nerve of the right hind limb was exposed. The positive and negative electrodes of an electrical stimulation electrode were then placed in contact with the sciatic nerve, or the conductive hydrogel SPPH2 of the present invention was injected into the nerve before the stimulation electrode was inserted into the SPPH2 hydrogel. A 0.1V electrical stimulation signal at different frequencies (5Hz, 10Hz, 20Hz, and 50Hz) was applied to the sciatic nerve via the electrical stimulation electrode. The response frequency of the rat's hind limb was recorded. Furthermore, the rat's hind limb was stimulated at different voltages (0.1V, 0.3V, 0.5V, 0.7V, and 0.9V) at 1Hz, and the amplitude of the rat's hind limb swing during stimulation was recorded.

[0059] In order to evaluate the application prospects of the present invention in the repair of acute muscle damage, an acute tibialis anterior muscle injury model was established. After anesthetizing the rat, the tibialis anterior muscle of the rat was exposed, and then the muscle was transected to cause acute muscle damage. An electrical stimulation electrode was inserted into the upper part of the severed muscle, and an electromyographic signal recording electrode was inserted into the lower part of the severed muscle. Subsequently, an electrical stimulation of 5V, 1Hz was applied, and the electrical signal was recorded. Then, the severed part was continuously filled with injectable conductive hydrogel, and the recovery of the muscle physiological signal transmission ability was judged based on the amplitude of the electromyographic signal.

[0060] 2. Results Analysis After applying electrical stimulation of different frequencies to the sciatic nerve of rats, the statistical results of the muscle response frequency to electrical stimulation are as follows: Figure 9 As shown, the results show that the presence of SPPH2 hydrogel does not attenuate the frequency of the electrical stimulation signal. Figure 10 As shown in the figure, when different voltages were applied to the sciatic nerve, there was almost no difference in the swing amplitude of the rat hind limbs between the groups containing hydrogel and those without hydrogel. This further proves that the presence of conductive hydrogel will not cause the loss of electrical stimulation signals. In addition, when the electrical stimulation electrode directly contacts the nerve tissue, frequent and high-intensity short-term stimulation can easily cause local tissue damage. At the same time, the significant mechanical modulus mismatch between the rigid electrode and the soft nerve tissue can also cause irreversible damage to the structurally fragile nerves. The introduced SPPH2 hydrogel can act as a bridge between the tissue and the electrode, avoiding direct contact between the electrode and the nerve, thereby protecting the nerve.

[0061] Figure 11 This is a photo of the acute tibialis anterior muscle injury model being severed and filled with SPPH2 hydrogel in Application Example 1. Figure 12 The amplitude statistics of the electrophysiological model transmitted by muscle after the acute tibialis anterior muscle injury model in Application Example 1 was severed and filled with SPPH2 hydrogel.

[0062] like Figure 11 and Figure 12 As shown in the acute tibialis anterior muscle injury model, when the muscle was not severed, an electrophysiological signal amplitude of approximately 16mV could be recorded under 5V electrical stimulation; however, when the muscle was completely severed, almost no effective electrical signal could be detected. After injecting 50μL of SPPH2 hydrogel into the severed area, the electrophysiological signal was partially restored, with a detected amplitude of approximately 10mV. After further injection of SPPH2 hydrogel to completely fill the severed area, the electrical signal amplitude increased to approximately 15mV, almost approaching the pre-injury level. These results indicate that this hydrogel can effectively reconstruct the electrical signal conduction pathway in acute muscle injury, showing broad application prospects in the immediate repair and functional recovery of muscle tissue.

[0063] Application Example 2: Representative injectable zwitterionic conductive hydrogels for improving the long-term performance of implantable electrodes 1. Experimental steps This application example describes the fibrotic response of the injectable conductive hydrogel in a mouse subcutaneous implantation model and its application in improving the long-term performance of myoelectric electrodes. The specific steps are as follows: First, the SPPH2 hydrogel prepared in Example 1 and the DSPH hydrogel prepared in Comparative Example 1 were injected subcutaneously into the back of female mice, with an injection volume of 100 μL per mouse. The inflammatory condition (representative cytokine immunohistochemistry) was evaluated on the 14th day, and the fibrosis response (MASSON trichrome, F4 / 80, and CD3 immunofluorescence staining) was evaluated on the 28th day.

[0064] Subsequently, to evaluate the hydrogel's ability to improve the long-term performance of implantable myoelectric electrodes, three experiments were conducted: one group measured rats' EMG signals using commercially available EMG electrodes applied epidermally (the Surface group); one group implanted bare platinum electrodes directly into the rat tibialis anterior muscle (the BarePt group); and another group injected SPPH2 hydrogel into the implantation site before implanting the electrodes within the hydrogel (the SPPH2 / Pt group). During the experiments, the rats were anesthetized, and then their sciatic nerves were stimulated with 5V / 1Hz electrical stimulation to obtain EMG signals. EMG signals were recorded for four weeks, and after four weeks, the muscle tissue was stained with Masson's and H&E staining to confirm the development of inflammation and fibrosis.

[0065] 2. Results Analysis First, mouse tissues were recovered on day 14 and immunohistochemical analysis of representative proinflammatory cytokines was performed. Figure 13 As shown, TNF-α, CCR7, IL-, and IL-17 are four typical pro-inflammatory cytokines, and higher expression levels indicate more severe inflammation. Compared with the control group, DSPH, SPPH2 triggered a weaker inflammatory response, with significantly lower release of representative pro-inflammatory cytokines than the control group, indicating that SPPH2 has superior immune compatibility.

[0066] Furthermore, tissue samples taken 28 days later were analyzed: the degree of fibrosis was assessed by MASSON trichrome staining, macrophage recruitment was confirmed by F4 / 80 staining, and the recruitment of immune cells around the implant was confirmed by CD3 staining. Figure 14 As shown, a dense collagen capsule formed around the DSPH in the control group. Notably, virtually no fibrosis was observed around the SPPH2 hydrogel, demonstrating its superior ability to resist foreign body reactions. Immunofluorescence staining also confirmed that fewer macrophages and immune cells were recruited around SPPH2. In summary, the SPPH2 hydrogel exhibits superior immunocompatibility and resistance to foreign body reactions, demonstrating promising potential for improving the long-term performance of implantable medical devices.

[0067] Based on the above results, the present invention further evaluated the role of the hydrogel in improving the long-term performance of electromyographic electrodes. Figure 15As shown. Compared with the electromyographic signals measured directly on the epidermis by commercial electrodes, both directly implanted electrodes and electrodes treated with SPPH2 showed better signal-to-noise ratio (SNR) in the initial stage; however, as the implantation time prolonged, the signal-to-noise ratio of the directly implanted electrodes decreased significantly, and eventually dropped to a level similar to the signal-to-noise ratio of the electromyographic signals measured on the epidermis. Although the signal-to-noise ratio of the electrodes treated with SPPH2 decreased to a certain extent, they were still able to record electromyographic signals with a higher signal-to-noise ratio on the 28th day, reflecting better long-term performance. Further MASSON slice analysis of the implantation site showed that a more obvious fibrosis reaction occurred at the bare electrode implantation site, while almost no fiber deposition was observed in the SPPH2-treated group ( Figure 16 This result further confirms that SPPH2 hydrogel has the ability to improve the long-term performance of implantable electrodes.

Claims

1. An injectable zwitterionic conductive hydrogel, characterized in that: It is self-assembled from zwitterionic copolymers and PEDOT / PSS conductive polymers; The zwitterionic copolymer is a copolymer formed by polymerization of sulfobetaine methacrylate or sulfobetaine methacrylate and other comonomers.

2. The injectable zwitterionic conductive hydrogel according to claim 1, characterized in that The molecular weight of the zwitterionic copolymer is 20,000-2,000,000.

3. The injectable zwitterionic conductive hydrogel according to claim 1, characterized in that The mass proportion of sulfobetaine methacrylate in the zwitterionic copolymer is greater than 50%.

4. The injectable zwitterionic conductive hydrogel according to claim 1, characterized in that The mass ratio of the zwitterionic copolymer to the PEDOT / PSS conductive polymer is 50-0.5:

1.

5. The injectable zwitterionic conductive hydrogel according to claim 1, characterized in that The other comonomers include one or more of acrylate monomers and acrylamide monomers.

6. The method for preparing the injectable zwitterionic conductive hydrogel according to any one of claims 1 to 5, characterized in that: Including steps: Step 1, copolymerizing raw materials containing sulfobetaine methacrylate to prepare a zwitterionic copolymer; Step 2: mixing the zwitterionic copolymer freeze-dried powder with the PEDOT / PSS conductive polymer solution to obtain the injectable zwitterionic conductive hydrogel through self-assembly.

7. The method for preparing the injectable zwitterionic conductive hydrogel according to claim 6, characterized in that: The copolymerization in step 1 contains an initiator, the copolymerization temperature is 60-90° C., the reaction is carried out for 6-18 hours, and the product is dialyzed and freeze-dried to obtain a copolymer.

8. The method for preparing the injectable zwitterionic conductive hydrogel according to claim 6, characterized in that: The initiator includes one or more of sodium persulfate, ammonium persulfate, and potassium persulfate, and the mass fraction of the initiator in the aqueous solution is 0.1-5%.

9. The method for preparing the injectable zwitterionic conductive hydrogel according to claim 6, characterized in that: The mixing in step 2 is carried out at room temperature and the self-assembly is carried out for more than 5 seconds.

10. Use of the injectable zwitterionic conductive hydrogel according to any one of claims 1 to 5 in the preparation of biomedical materials and bioelectronic materials.

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