An injectable zwitterionic conductive hydrogel and its preparation method and application

Abstract

The application discloses an injectable zwitterionic conductive hydrogel and a preparation method and application thereof, and the hydrogel is self-assembled from a zwitterionic copolymer and a PEDOT / PSS conductive polymer; the zwitterionic copolymer is a sulfobetaine methacrylate or a copolymer polymerized from the sulfobetaine methacrylate and other comonomers. In the application, the zwitterionic copolymer containing the sulfobetaine methacrylate and the PEDOT / PSS conductive polymer solution can be rapidly self-assembled into the hydrogel through simple mechanical mixing, and the dynamic cross-linking characteristics based on electrostatic interaction endow the hydrogel with excellent injectability and self-repairing performance. The hydrogel has excellent immunocompatibility and conductive performance, and can play a role in various bioelectronic application scenarios.

Description

TECHNICAL FIELD

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

[0002] In recent years, with the rapid development of wearable electronic devices, implantable neural interfaces and intelligent medical systems, there is a higher demand for soft conductive materials with good biocompatibility, conductivity and injectability. Conductive hydrogel has become a hot spot in the research of new generation of bioelectronic materials due to its softness and electronic transmission capacity. However, traditional conductive hydrogel often lacks good injectability due to its dependence on strong covalent cross-linking or large size structure, which is not conducive to minimally invasive surgery.

[0003] Although PEDOT / PSS is considered as the most promising conductive polymer in the field of biomedical applications, CN119950824A discloses a conductive temperature-sensitive hydrogel and its preparation method and application, which specifically includes PEDOT:PSS and Pluronic F127, and is a conductive temperature-sensitive hydrogel based on hydrogen bond interaction, which is used as a submucosal filling material. Pluronic F127 endows temperature sensitivity, which can be injected at room temperature and rapidly forms a solid structure at body temperature. After the operation, it can be removed by local cooling. The conductive performance of PEDOT:PSS supports the electrical cutting process and ensures the smooth progress of the cutting. The good conductivity of PEDOT:PSS helps to uniformly distribute the current and avoid local overheating damage. The hydrogen bond interaction between PEDOT:PSS and Pluronic F127 significantly improves the mechanical strength of the hydrogel, ensuring that the hydrogel provides a durable and stable mucosal lifting effect during the operation, and provides a reliable and efficient support material for minimally invasive surgery.

[0004] CN114504683A discloses a composite conductive hydrogel prepared by physical cross-linking of acellular matrix hydrogel and PEDOT:PSS, which has extremely high electronic transmission efficiency and mechanical strength, can well restore the microenvironment for cell growth, has excellent biocompatibility, can promote tissue regeneration, and has good fluidity and injectability, and is not limited by shape and size during use. After gelation, it can be well adhered to the tissue and does not need additional surgical suture; the conductivity of the hydrogel can be adjusted by adjusting the amount of PEDOT:PSS added, and the cells can be provided with electrical stimulation by applying a weak electric current, which can effectively promote cell proliferation, migration and functionalization, and is expected to be prepared into a biomedical conductive material for application.

[0005] However, the stability and immunocompatibility of PEDOT:PSS in complex physiological environment also have certain limitations, which can easily cause inflammatory response or signal attenuation. Zwitterionic polymers have excellent protein adsorption resistance, hydration ability and biocompatibility due to their unique electrically neutral structure, and have been widely used in biomaterial design in recent years. However, it is still a challenge to effectively integrate zwitterionic polymers with PEDOT / PSS to realize a hydrogel with high conductivity, good injectability and biological stability, which faces the challenges of complex material design and difficult structure regulation.

[0006] Therefore, it is of great significance to develop a new hydrogel system with injectability, zwitterionic characteristics and conductivity to promote the development of the fields of bioelectronics, neural regulation and tissue repair. SUMMARY

[0007] The present application aims at the difficulty of realizing injectability and immunocompatibility of conductive hydrogel, and provides an injectable zwitterionic conductive hydrogel, which is obtained by electrostatic self-assembly of zwitterionic copolymer and PEDOT / PSS. The dynamic cross-linking characteristics of the hydrogel endow it with excellent injectability and self-repairing performance. PEDOT / PSS and zwitterionic polymer endow the hydrogel with wide application prospects in the field of bioelectronic related medical field.

[0008] To achieve the above purpose, the technical scheme adopted by the present application is:

[0009] An injectable zwitterionic conductive hydrogel is self-assembled by zwitterionic copolymer and PEDOT / PSS conductive polymer.

[0010] The zwitterionic copolymer is sulfobetaine methacrylate or a copolymer polymerized by sulfobetaine methacrylate and other comonomers.

[0011] In the present application, the zwitterionic copolymer containing sulfobetaine methacrylate and the PEDOT / PSS conductive polymer solution can be rapidly self-assembled into a hydrogel by simple mechanical mixing. The dynamic cross-linking characteristics based on electrostatic interaction endow the hydrogel with excellent injectability and self-repairing performance. The hydrogel is composed of zwitterions with excellent immunocompatibility and PEDOT / PSS polymers with conductivity, so it has excellent immunocompatibility and conductivity, and can play a role in various bioelectronic application scenarios.

[0012] The molecular weight of the zwitterionic copolymer is 20000-2000000.

[0013] The mass percentage of the sulfobetaine methacrylate in the zwitterionic copolymer is more than 50%. Preferably, the mass percentage of the sulfobetaine methacrylate is more than 60%, more than 70% or more than 80%. The higher the content is, the more conducive to gelation is, and other co-monomers with corresponding functions can be added to obtain a hydrogel with corresponding functions according to specific application scenarios.

[0014] The other co-monomers include one or more of oleic acid ester monomers and acrylamide monomers.

[0015] The mass ratio of the zwitterionic copolymer to the PEDOT / PSS conductive polymer is 50-0.5:1. Preferably, the mass ratio of the zwitterionic copolymer to the PEDOT / PSS conductive polymer is 30-1:1; preferably, the mass ratio of the zwitterionic copolymer to the PEDOT / PSS conductive polymer is 20-5:1. The higher the content of the zwitterionic copolymer is, the better the biocompatibility of the hydrogel is, and the higher the content of the conductive polymer is, the better the conductivity of the hydrogel is. However, too high zwitterionic content will lead to a decrease in injectability and conductivity, and an increase in the content of the conductive polymer may lead to a decrease in gel strength.

[0016] The application also provides a preparation method of the injectable zwitterionic conductive hydrogel, comprising the following steps:

[0017] Step 1: copolymerization of raw materials containing sulfobetaine methacrylate to obtain a zwitterionic copolymer;

[0018] Step 2: mixing the zwitterionic copolymer lyophilized powder with a PEDOT / PSS conductive polymer solution to self-assemble to obtain the injectable zwitterionic conductive hydrogel.

[0019] The copolymerization in step 1 contains an initiator, the temperature of the copolymerization is 60-90℃, the reaction time is 6-18h, and the product is dialyzed and lyophilized to obtain the copolymer.

[0020] 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%.

[0021] The mixing in step 2 is carried out at room temperature, and the self-assembly time is more than 5s. After mixing, the viscosity of the system will continuously increase, and the self-assembly of the hydrogel can be achieved usually within 3-10min.

[0022] The application also provides the use of the injectable zwitterionic conductive hydrogel in the preparation of biomedical materials and bioelectronic materials.

[0023] For example, for the instant filling of tissue injury: the injectable zwitterionic conductive hydrogel combines excellent injectability, electrical conductivity and immunocompatibility endowed by zwitterionic polymer, and has unique advantages in minimally invasive repair after muscle injury. Unlike other tissues, muscle tissue has electrophysiological activity and can generate action potentials in response to nerve stimulation to maintain normal motor function. After acute muscle injury, the electrical signal transmission path of the damaged site may be interrupted, affecting tissue function recovery. Therefore, constructing a scaffold material with conductivity is crucial for reestablishing local electrical signal continuity. The injectable zwitterionic conductive hydrogel proposed in the present application can quickly fill the injury 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 function reconstruction.

[0024] For example, for improving the long-term immunocompatibility of implanted electrodes. After the implantable electrode is implanted in the human body, the body will produce a serious immune response to this type of foreign object, and after a period of time, a foreign body reaction will develop, and after the body cannot remove these foreign bodies, the collagen secreted by the tissue will isolate the electrode from the human body, eventually leading to a decline in electrode performance, and even failure. And this reaction is long-term and continuous, which will cause long-term pain to the patient.

[0025] In some embodiments, the hydrogel of the present application can be injected into the electrode implantation site, and then the electrode is implanted into the hydrogel. The hydrogel can act as a bridge between the tissue and the electrode. Since the hydrogel provided by the present application has excellent immunocompatibility, acute inflammation and foreign body reaction are avoided, and the long-term performance of the electrode is effectively improved.

[0026] Compared with the prior art, the present application has the following beneficial effects:

[0027] (1) The injectable zwitterionic conductive hydrogel in the present application does not require sophisticated structural design, but only two polymers available on the market or simply polymerized to form a hydrogel with conductivity and injectability, avoiding the cost and biocompatibility problems caused by the introduction of complex components, and the process is simple and easy to promote.

[0028] (2) The injectable zwitterionic conductive hydrogel in the present application has excellent injectability and immunocompatibility, and has broad application prospects in the fields of implantable bioelectronics and tissue repair. BRIEF DESCRIPTION OF DRAWINGS

[0029] To more clearly illustrate the technical solutions of the embodiments of this application, the accompanying drawings used in the embodiments of this application will be briefly introduced below. It should be understood that the following drawings only show some embodiments of this application and should not be regarded as a limitation of the scope. For those skilled in the art, other related drawings can be obtained based on these drawings without creative effort.

[0030] Figure 1 Strain scan of SPPH2, an injectable zwitterionic conductive hydrogel prepared in Example 1 ( Figure 1 (A) and angular frequency scan test ( Figure 1 (B) Result.

[0031] Figure 2 The graph shows the statistical values ​​of G' and tan δ for injectable zwitterionic conductive hydrogels with different component ratios prepared in Example 1.

[0032] 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.

[0033] 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)

[0034] Figure 5 The Raman spectra of the injectable zwitterionic conductive hydrogel SPPH2 prepared in Example 1 and pure PEDOT / PSS are shown.

[0035] Figure 6 This is a visualization result of the molecular simulated electrostatic potential (ESP) of the injectable zwitterionic conductive hydrogel described in this invention.

[0036] Figure 7 The ionic conductivity and electronic conductivity are those of the injectable zwitterionic conductive hydrogel prepared in Example 1.

[0037] Figure 8 The electrochemical impedance spectroscopy of the injectable zwitterionic conductive hydrogel prepared in Example 1 is shown.

[0038] Figure 9 The hindlimb response frequencies of rats after applying 0.1 V electrical stimulation signals (5 Hz, 10 Hz, 20 Hz, 50 Hz) to the sciatic nerve in different groups as described in Example 1.

[0039] Figure 10The different groups of rats described in Application Example 1 were subjected to electrical stimulation of the hind limbs at 1 Hz and different voltages (0.1 V, 0.3 V, 0.5 V, 0.7 V, 0.9 V), and the amplitude of the swing of the hind limbs under stimulation was recorded.

[0040] Figure 11 The physical diagram of the acute tibialis anterior muscle injury model in Application Example 1 was disconnected and filled with SPPH2 hydrogel.

[0041] Figure 12 After the acute tibialis anterior muscle injury model in Application Example 1 was disconnected and filled with SPPH2 hydrogel, the amplitude statistics of the muscle transmission of the transmitted electrophysiology model were recorded.

[0042] Figure 13 The expression statistics of representative inflammatory factors after DSPH and SPPH2 were implanted subcutaneously in mice for 14 days in Application Example 2.

[0043] Figure 14 After DSPH and SPPH2 were implanted subcutaneously in mice for 28 days in Application Example 2, the Masson trichrome section, F4 / 80 fluorescent staining and CD3 fluorescent staining were recorded.

[0044] Figure 15 The signal-to-noise ratio of the electromyographic signals recorded by the surface electrode, the directly implanted electrode and the SPPH2 treated electrode in Application Example 2.

[0045] Figure 16 The Masson trichrome section of the implantation site of the directly implanted electrode and the SPPH2 treated electrode in Application Example 2. DETAILED DESCRIPTION

[0046] In order to make the purpose, technical scheme and advantages of the present application more clear and obvious, the present application will be further described in detail below in combination with embodiments. It should be understood that the specific embodiments described herein are only used to explain the present application and not to limit the present application. Those skilled in the art can modify or replace equivalently without departing from the spirit and scope of the present application, which should be covered within the protection scope of the present application.

[0047] The raw materials used in the following specific embodiments were purchased from the market.

[0048] Example 1

[0049] Step 1: 10 wt% of sulfobetaine (SBMA) monomer was dissolved in deionized water, then 1 wt% of potassium persulfate was added as initiator. After bubbling deoxygenation for 20 minutes under nitrogen atmosphere, the reaction system was transferred to a 70 °C oil bath, and the polymerization reaction was stirred at constant temperature for 12 hours. After the reaction was completed, the product was dialyzed to remove small molecule impurities, and then freeze-dried to obtain poly sulfobetaine (PSBMA) polymer powder with an average molecular weight of about 200,000.

[0050] Step 2: At room temperature, 50 mg, 100 mg, 150 mg, and 200 mg of PSBMA freeze-dried powder were added to 1 mL of PEDOT / PSS solution (1.1 wt%), respectively, i.e. the mass ratio was 4.5:1, 9:1, 1.36:1, and 18:1, respectively. Then fast stirring until the powder was completely dissolved in the PEDOT / PSS solution, and as the powder was continuously dissolved, the viscosity of the system continuously increased, and after about 5 min, an injectable zwitterionic hydrogel was formed, which was recorded as SPPH1, SPPH2, SPPH3, and SPPH4, respectively.

[0051] Example 2

[0052] Step 1: 20 wt% of sulfobetaine (SBMA) monomer was dissolved in deionized water, then 1 wt% of potassium persulfate was added as initiator. After bubbling deoxygenation for 20 minutes under nitrogen atmosphere, the reaction system was transferred to a 70 °C oil bath, and the polymerization reaction was stirred at constant temperature for 12 hours. After the reaction was completed, the product was dialyzed to remove small molecule impurities, and then freeze-dried to obtain poly sulfobetaine (PSBMA) polymer powder with an average molecular weight of about 600,000.

[0053] Step 2: At room temperature, 200 mg of PSBMA freeze-dried powder was added to the PEDOT / PSS solution (1.1 wt%), and the mass ratio of PSBMA to PEDOT / PSS was 18:1. Then fast stirring until the powder was completely dissolved in the PEDOT / PSS solution, and as the powder was continuously dissolved, the viscosity of the system continuously increased, and finally an injectable zwitterionic hydrogel was formed.

[0054] Example 3

[0055] Step 1: 15 wt% of sulfobetaine (SBMA) monomer and 5 wt% of acrylic acid (AA) were dissolved in deionized water, followed by the addition of 1 wt% of potassium persulfate based on the total mass of monomers as an initiator. After deoxygenation by bubbling nitrogen for 20 minutes, the reaction system was transferred to a 70 °C oil bath and stirred at constant temperature for 12 hours of polymerization reaction. After the reaction was completed, the product was dialyzed to remove small molecular impurities, and then freeze-dried to obtain a poly sulfobetaine-poly acrylic acid copolymer (PSBMA-PAA) polymer powder with an average molecular weight of about 553,000.

[0056] Step 2: At room temperature, 200 mg of PSBMA-PAA freeze-dried powder was added to the PEDOT / PSS solution (1.1 wt%), and the mass ratio of PSBMA to PEDOT / PSS was 18:1. Then it was stirred quickly until the powder was completely dissolved in the PEDOT / PSS solution. As the powder continued to dissolve, the viscosity of the system continued to rise, and finally an injectable zwitterionic hydrogel was formed.

[0057] Comparative Example 1

[0058] Step 1: 40 μL of sodium dodecyl sulfate (DBSA) was added to 1 mL of PEDOT / PSS aqueous solution (1.1 wt%), and the mass ratio of DBSA to PEDOT / PSS was 3.63:1. The solution was mixed uniformly by vortexing, and then vacuumed for 5 min in a vacuum environment. After standing at room temperature for 2 h, the PEDOT / PSS solution formed a gel, which was dialyzed in ultrapure water to remove excess DBSA. The hydrogel was broken into small particles with a diameter of 50 μm using a cell disruptor, and an injectable conductive hydrogel DSPH was obtained.

[0059] Performance Test

[0060] 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 flat plate fixture with a diameter of 20 mm and a gap of 1500 μm. The test temperature was 37 °C. The sample was oscillated for 1 minute before testing, and then subjected to an oscillation strain sweep (strain range: 0.1-2000%, fixed frequency of 10 rad / s) and an oscillation frequency sweep (frequency range: 0.1-100 rad / s, fixed strain of 1%). The results are shown in Figure 1 Figure 1 The strain sweep (A) and angular frequency sweep test (B) of the injectable zwitterionic conductive hydrogel SPPH2 prepared in Example 1 are shown in Figure 1 Figure 1 ​​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 indicating that the material exhibits solid-like behavior.

[0061] Figure 2 The strength of the hydrogel formed by different mass ratios of zwitterionic PSBMA and PEDOT / PSS components was explored. Figure 2 The G' and tan δ statistical diagram of the injectable zwitterionic conductive hydrogel prepared in Example 1 with different component ratios can be seen that with the increase of PSBMA content, the mechanical strength of the obtained injectable conductive hydrogel is rising.

[0062] Figure 3 The injectable zwitterionic conductive hydrogel SPPH2 prepared in Example 1 was cycled at large and small strains to confirm the self-healing performance of the hydrogel, and the self-healing ability of the SPPH2 hydrogel was explored. In the rotational rheometer test, the hydrogel was subjected to 1000% and 1% strain, respectively. Large strain can destroy the network of the hydrogel, and low strain condition can provide self-healing for the hydrogel. After 3 cycles of strain, it can be seen that the mechanical properties of the hydrogel do not change significantly, proving its excellent self-healing performance.

[0063] Figure 4 The injection force of the injectable zwitterionic conductive hydrogel SPPH2 prepared in Example 1 Figure 4 and the injection filling ability for different interfaces Figure 4 and B). The injection force of SPPH1, SPPH2, SPPH3, SPPH4 hydrogels with different compositions was explored. The above four kinds of hydrogels were injected using a 23G injection needle, and the injection force was determined using a universal testing machine. The results show that the four kinds of hydrogels all have excellent injectability and can complete injection under small injection force. In addition, as shown in Figure 4 B, in some complex interfaces and gelatin-simulated tissues, this injectable hydrogel can effectively fill the tissue interface.

[0064] Figure 5 Raman spectrum of the injectable zwitterionic conductive hydrogel SPPH2 prepared in Example 1 and pure PEDOT / PSS. The interaction principle between PSBMA and PEDOT / PSS was explored. From the results of the Raman spectrum, it can be seen that the oxidation peak of PEDOT / PSS appears a significant blue shift, indicating its transition to the reduced state, which is a manifestation of the weakening of the interaction between PSS and PEDOT. It can be understood that PSBMA and PSS have electrostatic interaction, thereby promoting the aggregation of PEDOT, and finally forming an injectable conductive hydrogel.

[0065] Figure 6 The electrostatic potential map of the interaction between PSBMA and PEDOT / PSS polymer molecules is shown. It can be seen that there is a clear infiltration between the electron cloud of PSBMA and PEDOT / PSS, further proving that there is an electrostatic interaction between PSBMA and PEDOT / PSS.

[0066] Figure 7 The ionic conductivity (Ri) and electronic conductivity (Re) of the prepared SPPH1, SPPH2, SPPH3, SPPH4 hydrogels are shown. It can be seen that as the content of PSBMA increases, its conductivity decreases, which is inevitable because the corresponding conductive components will decrease significantly.

[0067] Figure 8 The electrochemical impedance spectra of the SPPH1, SPPH2, SPPH3, SPPH4 hydrogels prepared in Example 1 measured by an electrochemical workstation are shown. It can be seen that as the content of PSBMA increases, its electrochemical impedance rises.

[0068] Application Example 1

[0069] Representative injectable zwitterionic conductive hydrogel for nerve protection and acute muscle injury repair

[0070] 1. Animal surgery procedure

[0071] In this application example, the performance of the injectable zwitterionic conductive hydrogel in protecting nerves during stimulation of the nerve by a stimulation electrode and repairing the muscle electrophysiological signal transmission function in acute muscle injury is described. First, after the SD rat is anesthetized, the right hind limb sciatic nerve is exposed, and then the positive and negative electrodes of the stimulation electrode are in contact with the sciatic nerve or the conductive hydrogel SPPH2 of the application is first injected on the nerve, and then the stimulation electrode is inserted into the SPPH2 hydrogel. The sciatic nerve is subjected to 0.1V different frequency (5Hz, 10Hz, 20Hz, 50Hz) electrical stimulation signals by the stimulation electrode. The response frequency of the rat hind limb is recorded. In addition, the rat hind limb is subjected to 1Hz different voltage (0.1V, 0.3V, 0.5V, 0.7V, 0.9V) electrical stimulation, and the amplitude of the rat hind limb swing under stimulation is recorded.

[0072] To evaluate the application prospect of the present application in acute muscle injury repair, an acute tibialis anterior muscle injury model was established. After anesthesia of the rats, the tibialis anterior muscle of the rats was exposed, and then the muscle was transected to cause acute muscle injury. An electrical stimulation electrode was inserted into the upper part of the transected muscle, and a muscle electrical signal recording electrode was inserted into the lower part of the transected muscle. Then, an electrical stimulation of 5 V and 1 Hz was applied, and the electrical signal was recorded. Subsequently, the transected part was continuously filled with the injectable conductive hydrogel, and the recovery of the muscle physiological signal transmission ability was judged according to the amplitude of the muscle electrical signal.

[0073] 2. Result analysis

[0074] After applying electrical stimulation of different frequencies to the sciatic nerve of the rats, the statistical results of the response frequency of the muscle to the electrical stimulation are shown in Figure 9 , and the results show that the presence of the SPPH2 hydrogel does not cause attenuation of the electrical stimulation signal frequency. Further, as shown in Figure 10 , when different voltages are applied to the sciatic nerve, there is almost no difference in the swing amplitude of the hind limbs between the group containing the hydrogel and the group not containing the hydrogel. It is further proved that the presence of the conductive hydrogel does not cause loss of the electrical stimulation signal. In addition, when the electrical stimulation electrode directly contacts the nerve tissue, frequent and strong short-time 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 nerve. The introduced SPPH2 hydrogel can act as a bridge between the tissue and the electrode, avoiding direct contact between the electrode and the nerve, and thus protecting the nerve.

[0075] Figure 11 Figure 1 is a photograph of the transection and filling of the SPPH2 hydrogel in the acute tibialis anterior muscle injury model in application example 1. Figure 12 Figure 2 is a statistical diagram of the amplitude of the electrical physiological signal transmitted by the muscle after the transection and filling of the SPPH2 hydrogel in the acute tibialis anterior muscle injury model in application example 1.

[0076] As shown in Figure 11 and Figure 12 , in the acute tibialis anterior muscle injury model, an electrical physiological signal amplitude of about 16 mV can be recorded under 5 V electrical stimulation when the muscle is not transected; and when the muscle is completely transected, it is almost impossible to detect an effective electrical signal. After injection of 50 μL of the SPPH2 hydrogel into the transected part, the electrical physiological signal is partially recovered, and the detected amplitude is about 10 mV. After further injection of the SPPH2 hydrogel to completely fill the transected area, the electrical signal amplitude is increased to about 15 mV, which is almost close to the pre-injury level. The above results show that the hydrogel can effectively reconstruct the electrical signal conduction pathway in acute muscle injury, and has a broad application prospect in the instant repair and functional recovery of muscle tissue.

[0077] Application Example 2 Representative injectable zwitterionic conductive hydrogel for improving long-term performance of implanted electrodes

[0078] 1. Experimental procedure

[0079] This application example illustrates the fibrotic response of the injectable conductive hydrogel in a mouse subcutaneous implantation model, and its application in improving the long-term performance of electromyographic electrodes. The specific operation is as follows:

[0080] First, the SPPH2 hydrogel prepared as in Example 1 and the DSPH hydrogel prepared in Comparative Example 1 were injected subcutaneously into the back of female mice, with an injection amount of 100 μL per mouse, and the inflammatory condition (representative cytokine immunohistochemistry) was evaluated on day 14 and the fibrotic response (MASSON trichrome, F4 / 80, CD3 immunofluorescence staining) was evaluated on day 28.

[0081] Subsequently, to evaluate the performance of the hydrogel in improving the long-term performance of implanted electromyographic signal electrodes, three groups of experiments were set up: one group used a commercial electromyographic electrode to measure the electromyographic signal of rats on the skin, denoted as (Surface group), one group implanted a bare platinum electrode directly into the tibialis anterior muscle of rats (denoted as Bare Pt group); the other group first injected SPPH2 hydrogel at the implantation site, and then implanted the electrode into the hydrogel (denoted as SPPH2 / Pt group). During the experiment, the rats were first anesthetized, and then an electrical stimulus of 5V 1Hz was used to stimulate the sciatic nerve of the rats to obtain the electromyographic signal. The electromyographic signal was recorded for 4 weeks during the experiment, and after 4 weeks, the muscle tissue was subjected to Masson staining and H&E staining to confirm the development of inflammation and fibrosis.

[0082] 2. Results analysis

[0083] First, the mouse tissues were recovered on day 14 and subjected to immunohistochemical analysis of representative pro-inflammatory cytokines. The statistical results are shown in Figure 13 TNF-α, CCR7, IL-, and IL-17 are four typical pro-inflammatory cytokines, and the higher the expression, the more severe the inflammation. Compared with the control group DSPH, the inflammatory response induced by SPPH2 was weaker, and the release amount of representative pro-inflammatory cytokines was significantly lower than that of the control group, indicating that SPPH2 has better immunocompatibility.

[0084] Further, the 28-day tissue samples were detected and analyzed: the degree of fibrosis was evaluated by MASSON trichrome staining, the recruitment of macrophages was confirmed by F4 / 80 staining, and the recruitment of immune cells around the implant was confirmed by CD3 staining. The results are shown in Figure 14As shown, the control group DSPH formed a dense collagen envelope around the tissue; notably, almost no fibrotic response was observed around the SPPH2 hydrogel, indicating its excellent ability to resist foreign body response. At the same time, the immunofluorescence staining results also confirmed that the number of macrophages and immune cells recruited around SPPH2 was less. In summary, SPPH2 hydrogel has more excellent immunocompatibility and the ability to resist foreign body response, and shows good potential in improving the long-term performance of implanted medical devices.

[0085] Based on the above results, the application further evaluates the role of the hydrogel in improving the long-term performance of the myoelectric electrode, and the results are as shown in Figure 15 As shown. Compared with the myoelectric signal directly measured by the commercial electrode on the skin, the directly implanted electrode and the electrode treated by SPPH2 both showed more excellent signal-to-noise ratio (SNR) at the initial stage; but as the implantation time was prolonged, the signal-to-noise ratio of the directly implanted electrode decreased significantly and finally decreased to a level similar to the signal-to-noise ratio of the myoelectric signal measured on the skin. Although the signal-to-noise ratio of the electrode treated by SPPH2 decreased to a certain extent, it could still record the myoelectric signal with a higher signal-to-noise ratio at the 28th day, showing better long-term performance. Further MASSON section analysis of the implantation site showed that the bare electrode implantation site had a more obvious fibrotic response, while the SPPH2 treatment group almost did not observe fibrosis deposition Figure 16 ). This result further confirms that the SPPH2 hydrogel has the ability to improve the long-term performance of the implanted electrode.

Claims

1. An injectable zwitterionic conductive hydrogel, characterized in that, It is formed by self-assembly of zwitterionic copolymer and PEDOT / PSS conductive polymer in 3-10 minutes; The zwitterionic copolymer is a copolymer of betaine sulfonate methacrylate or betaine sulfonate methacrylate and other comonomers; the mass percentage of betaine sulfonate methacrylate in the zwitterionic copolymer is more than 50%. The mass ratio of the zwitterionic copolymer to the PEDOT / PSS conductive polymer is 20~5:

1.

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

3. 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.

4. The method for preparing the injectable zwitterionic conductive hydrogel according to any one of claims 1-3, characterized in that, Including the following steps: Step 1: Copolymerize the raw material containing sulfobetaine methacrylate to obtain a zwitterionic copolymer; Step 2: Mix the zwitterionic copolymer lyophilized powder with the PEDOT / PSS conductive polymer solution and self-assemble for 3-10 minutes to obtain the injectable zwitterionic conductive hydrogel.

5. The method for preparing the injectable zwitterionic conductive hydrogel according to claim 4, characterized in that, In step 1, the copolymer contains an initiator, the copolymerization temperature is 60-90℃, the reaction time is 6-18h, and the product is dialyzed and freeze-dried to obtain the copolymer.

6. The method for preparing the injectable zwitterionic conductive hydrogel according to claim 5, 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%.

7. The application of the injectable zwitterionic conductive hydrogel according to any one of claims 1-3 in the preparation of biomedical materials and bioelectronic materials.

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