Conductive polymer hydrogel as well as preparation method and application thereof in biological tissues

Abstract

The invention provides conductive polymer hydrogel, a preparation method thereof and application of the conductive polymer hydrogel in biological tissues. The conductive polymer hydrogel comprises poly (3, 4-ethylenedioxythiophene): polystyrene sulfonate and water, or the conductive polymer hydrogel comprises poly (3, 4-ethylenedioxythiophene): polystyrene sulfonate, an additive component and water, the additive component comprises one or more of a non-conductive polymer and an ionic compound, and the additive component and / or poly (3, 4-ethylenedioxythiophene): polystyrene sulfonate and / or poly (3, 4-ethylenedioxythiophene): polystyrene sulfonate and / or poly (3, 4-ethylenedioxythiophene): polystyrene sulfonate and / or poly (3, 4-ethylenedioxythiophene): polystyrene sulfonate and / or poly (3, 4-ethylenedioxythiophene): polystyrene sulfonate. The polystyrenesulfonate is of a net structure, and the polystyrenesulfonate is of a polystyrenesulfonate structure. The conductive polymer hydrogel has a certain network structure and has better flexibility and conductivity. The conductive polymer hydrogel can meet the wide mechanical-electrical property requirements required by bioelectronic application, and can be effectively used as a biological electrode for epidermal and implantable electrophysiological signal recording by synergistically balancing mechanical compliance and conductivity, and stable and reliable high-signal-to-noise-ratio signals are provided.

Description

Technical Field

[0001] This invention relates to the field of hydrogel technology, specifically to conductive polymer hydrogels, their preparation methods, and their applications in biological tissues. Background Art

[0002] Advanced bioelectronics holds revolutionary potential in biomedical applications, offering superior diagnostic and therapeutic capabilities, such as electrical stimulation of biological tissues and precise recording of biosignals. Unfortunately, traditional bioelectronic devices are made of rigid materials, resulting in significant mechanical mismatches with organs and hindering seamless tissue-electronic interfaces. Previous research has shown that minimizing stiffness mismatch can reduce damage to underlying tissues. Recently, numerous bioelectronic devices, including commercially available silicon probes, epidermal electrodes, and neural interfaces, have emerged, bridging the gap between electronic systems and the human body. However, most approaches focus primarily on the structural design of electronic components, such as employing ultrathin components and large-pore mesh structures to reduce bending stiffness.

[0003] Conductive hydrogels possess tissue-like mechanical properties, making them promising electronic interface materials for biological tissues. Biological tissues are typically soft, have a low elastic modulus (1 kPa–1 MPa), and contain a large amount of water, usually exceeding 70%. In contrast, most inorganic materials and dry polymers (such as metals and polycarbonates) used in bioelectronic devices have much higher elastic moduli (100 MPa–10 GPa) and contain almost no water. Conductive hydrogels are an ideal alternative in engineering materials due to their unique water-rich composition, excellent biocompatibility, intrinsic mechanical compliance, and conductivity. Conductive polymer hydrogels uniquely combine high conductivity, physiological stability, and excellent biocompatibility—advantages not found in systems based solely on ionic salts, metals, or carbon nanomaterials. However, current hydrogels still suffer from some drawbacks, such as relatively high modulus and poor conductivity. Summary of the Invention

[0004] This invention aims to at least partially solve one of the technical problems in related technologies. Therefore, one object of this invention is to provide a conductive polymer hydrogel with a low elastic modulus and suitable conductivity, which can be well applied to different biological tissues and achieve good compatibility with biological tissues.

[0005] In one aspect, the present invention provides a conductive polymer hydrogel. According to embodiments of the present invention, the conductive polymer hydrogel comprises poly(3,4-ethylenedioxythiophene):polystyrene sulfonate and water, or the conductive polymer hydrogel comprises the poly(3,4-ethylenedioxythiophene):polystyrene sulfonate, additives, and water, wherein the additives comprise one or more of non-conductive polymers and ionic compounds, wherein the additives and / or the poly(3,4-ethylenedioxythiophene):polystyrene sulfonate form a network structure. Thus, the above-mentioned conductive polymer hydrogel has a certain network structure, a low elastic modulus, resulting in good flexibility, and also good conductivity. The above-mentioned characteristics of the conductive polymer hydrogel of the present invention allow it to meet the wide range of mechatronic performance requirements for bioelectronic applications. By synergistically balancing mechanical compliance and conductivity, the conductive polymer hydrogel can be effectively used as a bioelectrode for epidermal and implantable electrophysiological signal recording, providing stable and reliable high signal-to-noise ratio signals.

[0006] According to an embodiment of the present invention, the added component includes the non-conductive polymer, which includes one or more of polyacrylamide, polyacrylamide-carrageenan, polyacrylic acid, and polyvinyl alcohol.

[0007] According to an embodiment of the present invention, based on the total mass of the conductive polymer hydrogel, the conductive polymer hydrogel comprises, by mass percentage, 0.03 wt.% to 0.1 wt.% of the poly(3,4-ethylenedioxythiophene):polystyrene sulfonate and 12 wt.% to 20 wt.% of the non-conductive polymer.

[0008] According to embodiments of the present invention, the added component includes the ionic compound, which includes one or more of the following: 1-ethyl-3-methylimidazolium ethyl sulfate; 4-(3-butyl-1-imidazolium)-1-butane sulfonate triester; 1-butyl-3-methylimidazolium tetrafluoroborate; [2-(methacryloyloxy)ethyl]dimethyl-(3-sulfopropyl)ammonium hydroxide; and 3-[dimethyl-[3-(2-methylprop-2-enoylamino)propyl]ammonium]propane-1-sulfonate.

[0009] According to an embodiment of the present invention, based on the total mass of the conductive polymer hydrogel, the conductive polymer hydrogel comprises, by mass percentage, 0.8 wt.% to 1.60 wt.% of the poly(3,4-ethylenedioxythiophene):polystyrene sulfonate and 20 wt.% to 30 wt.% of the ionic compound.

[0010] According to an embodiment of the present invention, the conductive polymer hydrogel comprises poly(3,4-ethylenedioxythiophene):polystyrene sulfonate and water, and the conductive polymer hydrogel comprises, by mass percentage, 0.8 wt.% to 2 wt.% of the poly(3,4-ethylenedioxythiophene):polystyrene sulfonate.

[0011] According to an embodiment of the present invention, the conductive polymer hydrogel has an elastic modulus of less than 130 kPa and / or an electrical conductivity of 0.15 to 6 S / m.

[0012] In another aspect, the present invention provides a method for preparing the aforementioned conductive polymer hydrogel. According to embodiments of the present invention, the method for preparing the conductive polymer hydrogel includes one of the following methods:

[0013] Method 1:

[0014] Adding dopamine hydrochloride or polyvinyl alcohol to an alkaline solution yields a first mixed solution; adding freeze-dried poly(3,4-ethylenedioxythiophene):polystyrene sulfonate to the first mixed solution yields a second mixed solution; adding a polymerizing monomer, a crosslinking agent, and an initiator to the second mixed solution, and reacting, yields a first reaction solution; the first reaction solution self-cures to obtain the conductive polymer hydrogel.

[0015] Method 2:

[0016] An ionic compound, the crosslinking agent, and an initiator are added to an aqueous solution of poly(3,4-ethylenedioxythiophene):polystyrene sulfonate, and the mixture is stirred to obtain a third mixed solution. The third mixed solution is then subjected to a polymerization reaction at a predetermined temperature for a predetermined time to obtain the conductive polymer hydrogel.

[0017] Method 3:

[0018] A polar solvent is added to the aqueous solution of poly(3,4-ethylenedioxythiophene):polystyrene sulfonate and stirred to obtain a fourth mixed solution; the fourth mixed solution is then subjected to drying, annealing and hydration treatments in sequence to obtain the conductive polymer hydrogel.

[0019] According to embodiments of the present invention, the conductive polymer hydrogel prepared by the above method has a certain network structure and a low elastic modulus, giving it good flexibility and good conductivity. These properties of the conductive polymer hydrogel of the present invention allow it to meet the wide range of mechatronic performance requirements for bioelectronic applications. By synergistically balancing mechanical compliance and conductivity, the conductive polymer hydrogel can be effectively used as a bioelectrode for epidermal and implantable electrophysiological signal recording, providing stable and reliable high signal-to-noise ratio signals.

[0020] According to an embodiment of the present invention, in method one, at least one of the following conditions is met: the concentration of the alkali solution is 0.2–0.8 g / mL; the mass ratio of dopamine hydrochloride to the volume of the alkali solution is 1 mg:(0.01–0.05) ml, or the mass ratio of polyvinyl alcohol to the volume of the alkali solution is 1 mg:(0.0001–0.0005) ml; the mass ratio of dopamine hydrochloride to the mass of poly(3,4-ethylenedioxythiophene):polystyrene sulfonate is (1.5–5): 1. Alternatively, the mass ratio of the polyvinyl alcohol to the poly(3,4-ethylenedioxythiophene):polystyrene sulfonate is (200-300):1; the mass ratio of the polymeric monomer to the poly(3,4-ethylenedioxythiophene):polystyrene sulfonate is (200-300:1), or (150-230):1; the mass fraction of the crosslinking agent is 0.03 wt.% to 0.09 wt.%; and the mass fraction of the initiator is 0.1 wt.% to 0.2 wt.%.

[0021] According to an embodiment of the present invention, in method one, the alkaline solution further includes carrageenan, and the mass ratio of the carrageenan to the mass of the poly(3,4-ethylenedioxythiophene):polystyrene sulfonate is (12-20):1.

[0022] According to an embodiment of the present invention, in method two, at least one of the following conditions is met: the mass ratio of the ionic compound to the poly(3,4-ethylenedioxythiophene):polystyrene sulfonate is (15-25):1; the mass fraction of the crosslinking agent is 0.03 wt.% to 0.09 wt.%; the mass fraction of the initiator is 0.1 wt.% to 0.2 wt.%; the predetermined temperature is 75-110°C; and the predetermined time is 30 min to 60 min.

[0023] Alternatively, in method three, at least one of the following conditions must be met: in the fourth mixed solution, the volume fraction of the polar solvent is 20 vol% to 30 vol%; in the aqueous solution of poly(3,4-ethylenedioxythiophene):polystyrene sulfonate, the concentration of poly(3,4-ethylenedioxythiophene):polystyrene sulfonate is (8 to 16) mg / ml; the annealing treatment includes heating the fourth mixed solution after the drying treatment to the annealing temperature, and then naturally cooling it at room temperature, repeating this process 2 to 5 times; the annealing temperature is 80 to 200°C; the natural cooling time is 10 to 120 min.

[0024] In another aspect, the present invention provides an application of the conductive polymer hydrogel described above, or a conductive polymer hydrogel prepared by the method described above, in biological tissues. Thus, suitable conductive polymer hydrogels can be selected for implantation into different biological tissues based on their elastic modulus and conductivity, thereby enhancing the flexibility and extensibility of the conductive polymer hydrogels used as epidermal or implantable bioelectrodes.

[0025] Additional aspects and advantages of the invention will be set forth in part in the description which follows, and in part will be obvious from the description, or may be learned by practice of the invention. Attached Figure Description

[0026] The above and / or additional aspects and advantages of the present invention will become apparent and readily understood from the description of the embodiments taken in conjunction with the following drawings, in which:

[0027] Figure 1 The diagram shows a schematic diagram, scanning electron microscope image, Fourier transform infrared spectrum, and ultraviolet-visible spectrum of the crosslinking mechanism of the hydrogel IPNCH@CAPAM in the embodiments of the present invention.

[0028] Figure 2 The diagram shows the crosslinking mechanism of the hydrogel PSCH@SBMA in the embodiments of the present invention, Fourier transform infrared spectrum, Raman spectrum and conductive atomic force microscope;

[0029] Figure 3 The diagram shows the crosslinking mechanism of the hydrogel PSCH@SBMA in the embodiments of the present invention, the Fourier transform infrared spectrum, the XRD spectrum and the conductive atomic force microscope.

[0030] Figure 4 The mechanical property test graphs and conductivity test graphs of the conductive polymer hydrogels in Examples 1-14 are shown.

[0031] Figure 5 The effects of raw material dosage on conductivity, impedance, and elastic modulus during the preparation of IPNCH@CAPAM, PSCH@SBMA, and PCH@EG in the examples are shown.

[0032] Figure 6 The relevant mechanical property test graphs of IPNCH@CAPAM, PSCH@SBMA and PCH@EG in the embodiments are shown;

[0033] Figure 7-1 The relevant electrical performance test graphs for IPNCH@CAPAM, PSCH@SBMA, and PCH@EG in the embodiments are shown.

[0034] Figure 7-2The relevant electrical performance test graphs for IPNCH@CAPAM, PSCH@SBMA, and PCH@EG in the embodiments are shown.

[0035] Figure 8-1 The following diagrams show the epidermal electrophysiological signal recordings of IPNCH@CAPAM, PSCH@SBMA, and PCH@EG in the embodiments.

[0036] Figure 8-2 The following diagrams show the epidermal electrophysiological signal recordings of IPNCH@CAPAM, PSCH@SBMA, and PCH@EG in the embodiments.

[0037] Figure 8-3 The following diagrams show the epidermal electrophysiological signal recordings of IPNCH@CAPAM, PSCH@SBMA, and PCH@EG in the embodiments.

[0038] Figure 9-1 The following are relevant test graphs of the implantable electrophysiological signal recorders IPNCH@CAPAM and PSCH@SBMA in the embodiments;

[0039] Figure 9-2 The following are test diagrams of the implantable electrophysiological signal recorders IPNCH@CAPAM, PSCH@SBMA, and PCH@EG in the embodiments. Detailed Implementation

[0040] The present invention will be explained below with reference to embodiments. Those skilled in the art will understand that the following embodiments are for illustrative purposes only and should not be considered as limiting the scope of the invention. Where specific techniques or conditions are not specified in the embodiments, they are performed according to the techniques or conditions described in the literature in the field or according to the product instructions. Reagents or instruments whose manufacturers are not specified are all conventional products that can be obtained commercially.

[0041] The present invention will now be described with reference to specific embodiments. It should be noted that these embodiments are merely descriptive and do not limit the present invention in any way.

[0042] In one aspect, the present invention provides a conductive polymer hydrogel. According to embodiments of the present invention, the conductive polymer hydrogel comprises poly(3,4-ethylenedioxythiophene):polystyrene sulfonate (PEDOT:PSS) and water, or the conductive polymer hydrogel comprises the poly(3,4-ethylenedioxythiophene):polystyrene sulfonate (PEDOT:PSS), additives, and water, wherein the additives include one or more of non-conductive polymers and ionic compounds, wherein the additives and / or the poly(3,4-ethylenedioxythiophene):polystyrene sulfonate form a network structure. Thus, the above-mentioned conductive polymer hydrogel has a certain network structure, a low elastic modulus, resulting in good flexibility, and also good conductivity. The above-mentioned characteristics of the conductive polymer hydrogel of the present invention allow it to meet the wide range of mechatronic performance requirements for bioelectronic applications. By synergistically balancing mechanical compliance and conductivity, the conductive polymer hydrogel can be effectively used as a bioelectrode for epidermal and implantable electrophysiological signal recording, providing stable and reliable high signal-to-noise ratio signals.

[0043] According to embodiments of the present invention, conductive polymer hydrogels with different compositions have different elastic moduli and conductivity. Therefore, in the present invention, conductive polymer hydrogels with more suitable elastic moduli and conductivity can be selected according to the different biological tissues to be implanted.

[0044] According to some embodiments of the present invention, the added components include a non-conductive polymer, which includes one or more of polyacrylamide (PAM), polyacrylamide-carrageenan (CAPAM), polyacrylic acid (PAAc), and polyvinyl alcohol (PVA). Thus, by using the conductive polymer PEDOT:PSS as a conductive dopant in the non-conductive hydrogel template, an interpenetrating polymer network (IPNCH) is formed. This involves mixing or in-situ polymerizing the conductive polymer (PEDOT:PSS) within the non-conductive hydrogel template (i.e., the hydrogel formed by the non-conductive polymer), forming a typical interpenetrating polymer network (IPN). PEDOT:PSS possesses dual conductivity and biocompatibility. The non-conductive hydrogel template (polyacrylamide, polyacrylamide-carrageenan, polyacrylic acid, and polyvinyl alcohol) is used as a supporting substrate. This non-conductive hydrogel template is porous and has a high water content, thus giving it inherent softness and moisture retention, which is beneficial for conformal contact with biological tissues. This conductive polymer hydrogel combines the conductivity of the conductive polymer with the mechanical properties of the non-conductive substrate.

[0045] In some specific embodiments, the conductive polymer hydrogel comprises PEDOT:PSS, polyacrylamide-carrageenan (CAPAM), and water, and this interpenetrating network conductive polymer hydrogel may be simply referred to as IPNCH@CAPAM; in some specific embodiments, the conductive polymer hydrogel comprises PEDOT:PSS, polyacrylamide (PAM), and water, and this interpenetrating network conductive polymer hydrogel may be simply referred to as IPNCH@PAM; in some specific embodiments, the conductive polymer hydrogel comprises PEDOT:PSS, polyacrylic acid (PAAc), and water, and this interpenetrating network conductive polymer hydrogel may be simply referred to as IPNCH@PAAc-PVA.

[0046] According to some embodiments of the present invention, the conductive polymer hydrogel comprises PEDOT:PSS, a non-conductive polymer, and water. Based on the total mass of the conductive polymer hydrogel, by mass percentage, the conductive polymer hydrogel comprises 0.03 wt.% to 0.1 wt.% (e.g., 0.03 wt.%, 0.04 wt.%, 0.05 wt.%, 0.06 wt.%, 0.07 wt.%, 0.08 wt.%, 0.09 wt.%, 0.1 wt.%, etc.) of poly(3,4-ethylenedioxythiophene):polystyrene sulfonate (PEDOT:PSS) and 12 wt.% to 20 wt.% (e.g., 12 wt.%, 13 wt.%, 14 wt.%, 15 wt.%, 16 wt.%, 17 wt.%, 18 wt.%, 19 wt.%, 20 wt.%, etc.) of non-conductive polymer. Thus, the conductive polymer hydrogel with the above proportions has a lower elastic modulus and lower conductivity.

[0047] According to some embodiments of the present invention, the added components include ionic compounds, including one or more selected from 1-ethyl-3-methylimidazolium ethyl sulfate (EMIM: ES), 4-(3-butyl-1-imidazolium)-1-butanesulfonic acid triacid (BIM: BSA3), 1-butyl-3-methylimidazolium tetrafluoroborate (Bmim: BF4), [2-(methacryloyloxy)ethyl]dimethyl-(3-sulfonylpropyl)ammonium hydroxide (SBMA), and 3-[dimethyl-[3-(2-methylprop-2-enoylamino)propyl]ammonium]propane-1-sulfonate (SBAA). Thus, molecules having sulfonate or sulfonylimide anions are incorporated into the conductive polymer PEDOT:PSS as effective dopant to regulate phase separation and promote the aggregation of the conductive polymer PEDOT:PSS, forming a phase-separated conductive polymer hydrogel (PSCH). The aforementioned ionic compounds exhibit excellent solubility in both water and conductive polymer matrices, thus effectively softening the PSS domain in PEDOT:PSS. The highly acidic sulfonate or sulfonylimide groups interact strongly with the positively charged PEDOT segments and polar groups on the polymer chains through electrostatic and hydrogen bonding interactions. This dual interaction not only stabilizes the doped state but also induces controlled, uniform phase separation and conductive polymer aggregation, thereby forming an interconnected conductive network and achieving a synergistic effect of conductivity and stretchability.

[0048] In some specific embodiments, the conductive polymer hydrogel comprises PEDOT:PSS, 1-ethyl-3-methylimidazolium ethyl sulfate (EMIM:ES), and water, and this phase-separated conductive polymer hydrogel may be simply referred to as PSCH@IL1; in some specific embodiments, the conductive polymer hydrogel comprises PEDOT:PSS, 4-(3-butyl-1-imidazolium)-1-butyrylate triacid (BIM:BSA3), and water, and this phase-separated conductive polymer hydrogel may be simply referred to as PSCH@IL2; in some specific embodiments, the conductive polymer hydrogel comprises PEDOT:PSS, 1-butyl-3-methylimidazolium tetrafluoroborate (Bmim:B... F4) and water, the phase-separated conductive polymer hydrogel can be simply referred to as PSCH@IL3; in some specific embodiments, the conductive polymer hydrogel includes PEDOT:PSS, [2-(methacryloyloxy)ethyl]dimethyl-(3-sulfopropyl)ammonium hydroxide (SBMA) and water, the phase-separated conductive polymer hydrogel can be simply referred to as PSCH@SBMA; in some specific embodiments, the conductive polymer hydrogel includes PEDOT:PSS, 3-[dimethyl-[3-(2-methylprop-2-enoylamino)propyl]ammonium]propane-1-sulfonate (SBAA) and water, the phase-separated conductive polymer hydrogel can be simply referred to as PSCH@SBAA.

[0049] According to some embodiments of the present invention, the conductive polymer hydrogel comprises PEDOT:PSS, an ionic compound, and water. Based on the total mass of the conductive polymer hydrogel, the conductive polymer hydrogel comprises, by mass percentage, 0.8 wt.% to 1.60 wt.% (e.g., 0.8 wt.%, 0.9 wt.%, 1.0 wt.%, 1.1 wt.%, 1.2 wt.%, 1.3 wt.%, 1.4 wt.%, 1.5 wt.%, 1.6 wt.%, etc.) of poly(3,4-ethylenedioxythiophene):polystyrene sulfonate (PEDOT:PSS) and 20 wt.% to 30 wt.% (e.g., 20 wt.%, 21 wt.%, 22 wt.%, 23 wt.%, 24 wt.%, 25 wt.%, 26 wt.%, 27 wt.%, 28 wt.%, 29 wt.%, 30 wt.%, etc.) of the ionic compound. Therefore, the conductive polymer hydrogels in the above proportions have low elastic modulus and low conductivity.

[0050] According to some embodiments of the present invention, the conductive polymer hydrogel comprises poly(3,4-ethylenedioxythiophene):polystyrene sulfonate (PEDOT:PSS) and water. In the preparation method, a polar solvent is introduced into the conductive polymer PEDOT:PSS, followed by controlled dry calcination to generate a pure conductive polymer hydrogel (PCH). In another preparation method, a polar solvent (e.g., dimethyl sulfoxide (DMSO), ethylene glycol (EG), glycerol, N,N-dimethylacetamide (DMF), and tetrahydrofuran (THF)) is added to the conductive polymer PEDOT:PSS system. Through heating, drying annealing, and rehydration processes, these polar solvents can be partially removed, resulting in a nearly pure conductive polymer hydrogel. The addition of the polar solvent improves the crystallinity and orderliness of the conductive polymer nanocrystals, thereby enhancing the electrical properties. Furthermore, the pure conductive polymer hydrogel, free of other components, avoids problems such as impaired conductivity, uneven mechanical and electrical properties, and cytotoxicity caused by other fillers in the hydrogel network.

[0051] In some specific embodiments, dimethyl sulfoxide (DMSO) is added as a polar solvent during the preparation of the conductive polymer hydrogel, and the resulting pure conductive polymer hydrogel can be abbreviated as PCH@DMSO; in some specific embodiments, ethylene glycol (EG) is added as a polar solvent during the preparation of the conductive polymer hydrogel, and the resulting pure conductive polymer hydrogel can be abbreviated as PCH@EG; in some specific embodiments, glycerol is added as a polar solvent during the preparation of the conductive polymer hydrogel, and the resulting pure conductive polymer hydrogel can be abbreviated as PCH@Glycerol; in some specific embodiments, N,N-dimethylacetamide (DMF) is added as a polar solvent during the preparation of the conductive polymer hydrogel, and the resulting pure conductive polymer hydrogel can be abbreviated as PCH@DMF; in some specific embodiments, tetrahydrofuran (THF) is added as a polar solvent during the preparation of the conductive polymer hydrogel, and the resulting pure conductive polymer hydrogel can be abbreviated as PCH@THF.

[0052] According to some embodiments of the present invention, the conductive polymer hydrogel comprises PEDOT:PSS and water, and based on the total mass of the conductive polymer hydrogel, by mass percentage, the conductive polymer hydrogel comprises 0.8 wt.% to 2 wt.% (e.g., 0.8 wt.%, 0.9 wt.%, 1.0 wt.%, 1.1 wt.%, 1.2 wt.%, 1.3 wt.%, 1.4 wt.%, 1.5 wt.%, 1.6 wt.%, 1.7 wt.%, 1.8 wt.%, 1.9 wt.%, 2.0 wt.%, etc.) of poly(3,4-ethylenedioxythiophene):polystyrene sulfonate (PEDOT:PSS). Thus, the conductive polymer hydrogel with the above proportions has a lower elastic modulus and lower conductivity.

[0053] According to some embodiments of the present invention, in addition to the components mentioned above, the conductive polymer hydrogel may also contain some raw materials that have not fully reacted during the preparation process and / or components involved in side reactions.

[0054] According to some embodiments of the present invention, the elastic modulus of the conductive polymer hydrogel is less than 130 kPa (e.g., 130 kPa, 100 kPa, 80 kPa, 70 kPa, 50 kPa, 40 kPa, 30 kPa, 20 kPa, 10 kPa, 8 kPa, 5 kPa, 3 kPa, 1 kPa, 0.8 kPa, 0.5 kPa, 0.2 kPa, 0.1 kPa, etc.).

[0055] According to some embodiments of the present invention, the conductivity of the conductive polymer hydrogel is 0.15 to 6 S / m (e.g., 0.15 S / m, 0.2 S / m, 0.8 S / m, 1 S / m, 1.2 S / m, 1.5 S / m, 1.8 S / m, 2 S / m, 2.3 S / m, 2.5 S / m, 3 S / m, 3.5 S / m, 4 S / m, 4.5 S / m, 5 S / m, 5.5 S / m, 6 S / m, etc.).

[0056] In another aspect, the present invention provides a method for preparing the aforementioned conductive polymer hydrogel. According to embodiments of the present invention, the method for preparing the conductive polymer hydrogel includes one of the following methods:

[0057] According to some embodiments of the present invention, a method for preparing an interpenetrating network conductive polymer hydrogel (IPNCH) includes: adding dopamine hydrochloride or polyvinyl alcohol to an alkaline solution to obtain a first mixed solution; adding freeze-dried poly(3,4-ethylenedioxythiophene):polystyrene sulfonate (PEDOT:PSS) to the first mixed solution to obtain a second mixed solution; adding a polymerizing monomer, a crosslinking agent, and an initiator to the second mixed solution, and reacting to obtain a first reaction solution; and self-curing the first reaction solution to obtain the conductive polymer hydrogel.

[0058] According to some embodiments of the present invention, in the above-described method, the alkaline solution also includes carrageenan (CA).

[0059] The following describes in detail, according to some specific embodiments of the present invention, several methods for preparing conductive polymer hydrogels (IPNCH):

[0060] In some embodiments of the present invention, the preparation method of interpenetrating network conductive polymer hydrogel (IPNCH@CAPAM) includes: firstly, completely dissolving carrageenan (CA) in deionized water, then adding NaOH to the carrageenan solution to form an alkaline environment to obtain an alkaline solution; then adding dopamine hydrochloride (DAHCl) to the alkaline solution, and reacting for a certain time to obtain a first mixed solution; dispersing freeze-dried PEDOT:PSS into the first mixed solution to obtain a second mixed solution; then completely dissolving the monomer acrylamide (AAm), crosslinking agent, and initiator in the above-mentioned second mixed solution to obtain a first reaction solution; finally, pouring the first reaction solution into a corresponding mold and self-curing to obtain conductive polymer hydrogel (IPNCH@CAPAM).

[0061] In some embodiments, in the preparation method of the above-described conductive polymer hydrogel (IPNCH@CAPAM), the mass ratio of carrageenan to poly(3,4-ethylenedioxythiophene):polystyrene sulfonate is (12-20):1, for example (12:1, 13:1, 14:1, 15:1, 15.6:1, 17:1, 18:1, 19:1, 20:1, etc.); the alkali concentration is 0.2-0.8 g / mL (for example, 0.2 g / mL, 0.3 g / mL). / mL, 0.4g / mL, 0.5g / mL, 0.6g / mL, 0.7g / mL, 0.8g / mL, etc.); the mass ratio of dopamine hydrochloride to alkaline solution volume is 1mg:(0.01~0.05)ml; the mass ratio of dopamine hydrochloride to poly(3,4-ethylenedioxythiophene):polystyrene sulfonate is (1.5~5):1; the mass ratio of the polymer monomer acrylamide (AAm) to poly(3,4-ethylenedioxythiophene):polystyrene sulfonate is (1.5~5):1. The ratio is (200-300):1, for example, 200:1, 210:1, 220:1, 230:1, 240:1, 250:1, 260:1, 270:1, 280:1, 290:1, 300:1, etc.; based on the total mass of the conductive polymer hydrogel, the mass of the crosslinking agent added is 0.03wt.% to 0.09wt.% (for example, 0.03wt.%, 0.04wt.%, 0.05wt.%, 0.06wt.%, 0.09wt.%). 0.7 wt.%, 0.08 wt.%, 0.09 wt.%, etc.); based on the total mass of the conductive polymer hydrogel, the added mass of the initiator is 0.1 wt.% to 0.2 wt.% (e.g., 0.1 wt.%, 0.11 wt.%, 0.12 wt.%, 0.13 wt.%, 0.14 wt.%, 0.15 wt.%, 0.16 wt.%, 0.17 wt.%, 0.18 wt.%, 0.19 wt.%, 0.2 wt.%, etc.).

[0062] In other embodiments of the present invention, the preparation method of interpenetrating network conductive polymer hydrogel (IPNCH@PAM) includes: firstly, adding NaOH to deionized water to form an alkaline environment to obtain an alkaline solution; then adding dopamine hydrochloride (DAHCl) to the alkaline solution, reacting for a certain time to obtain a first mixed solution; dispersing freeze-dried PEDOT:PSS into the first mixed solution to obtain a second mixed solution; then completely dissolving the monomer acrylamide (AAm), crosslinking agent and initiator in the above-mentioned second mixed solution to obtain a first reaction solution; finally pouring the first reaction solution into a corresponding mold and self-curing to obtain conductive polymer hydrogel (IPNCH@PAM).

[0063] In some embodiments, in the preparation method of the above-mentioned conductive polymer hydrogel (IPNCH@PAM), the concentration of the alkali solution is 0.2-0.8 g / mL (e.g., 0.2 g / mL, 0.3 g / mL, 0.4 g / mL, 0.5 g / mL, 0.6 g / mL, 0.7 g / mL, 0.8 g / mL, etc.); the mass ratio of dopamine hydrochloride to the volume of the alkali solution is 1 mg:(0.01-0.05) ml; the mass ratio of dopamine hydrochloride to poly(3,4-ethylenedioxythiophene):polystyrene sulfonate is (1.5-5):1; the mass ratio of the polymeric monomer acrylamide (AAm) to poly(3,4-ethylenedioxythiophene):polystyrene sulfonate is (200-300):1, for example, 200:1, 210:1, 220:1, 230:1, 240:1, 250:1. 1. 260:1, 270:1, 280:1, 290:1, 300:1, etc.; Based on the total mass of the conductive polymer hydrogel, the mass fraction of the crosslinking agent is 0.03wt.% to 0.09wt.% (e.g., 0.03wt.%, 0.04wt.%, 0.05wt.%, 0.06wt.%, 0.07wt.%, 0.08wt.%, 0.09wt.%, etc.); Based on the total mass of the conductive polymer hydrogel, the mass fraction of the initiator is 0.1wt.% to 0.2wt.% (e.g., 0.1wt.%, 0.11wt.%, 0.12wt.%, 0.13wt.%, 0.14wt.%, 0.15wt.%, 0.16wt.%, 0.17wt.%, 0.18wt.%, 0.19wt.%, 0.2wt.%, etc.).

[0064] In some embodiments of the present invention, the preparation method of interpenetrating network conductive polymer hydrogel (IPNCH@PAAc) includes: firstly, adding NaOH to deionized water to form an alkaline environment to obtain an alkaline solution; then adding dopamine hydrochloride (DAHCl) to the alkaline solution, reacting for a certain time to obtain a first mixed solution; dispersing freeze-dried PEDOT:PSS into the first mixed solution to obtain a second mixed solution; then completely dissolving the monomer acrylic acid (AAc), crosslinking agent and initiator in the above-mentioned second mixed solution to obtain a first reaction solution; finally, pouring the first reaction solution into a corresponding mold and self-curing to obtain conductive polymer hydrogel (IPNCH@PAAc).

[0065] In some embodiments, in the preparation method of the above-mentioned interpenetrating network conductive polymer hydrogel (IPNCH@PAAc), the concentration of the alkali solution is 0.2-0.8 g / mL (e.g., 0.2 g / mL, 0.3 g / mL, 0.4 g / mL, 0.5 g / mL, 0.6 g / mL, 0.7 g / mL, 0.8 g / mL, etc.); the mass ratio of dopamine hydrochloride to the volume of the alkali solution is 1 mg:(0.01-0.05) ml; the mass ratio of dopamine hydrochloride to poly(3,4-ethylenedioxythiophene):polystyrene sulfonate is (1.5-5):1; the mass ratio of the polymeric monomer acrylic acid (AAc) to poly(3,4-ethylenedioxythiophene):polystyrene sulfonate is (200-300):1, for example, 200:1, 210:1, 220:1, 230:1, 240:1, 25... The crosslinking agent has a mass fraction of 0.03 wt.% to 0.09 wt.% (e.g., 0.03 wt.%, 0.04 wt.%, 0.05 wt.%, 0.06 wt.%, 0.07 wt.%, 0.08 wt.%, 0.09 wt.%) based on the total mass of the conductive polymer hydrogel; the initiator has a mass fraction of 0.1 wt.% to 0.2 wt.% (e.g., 0.1 wt.%, 0.11 wt.%, 0.12 wt.%, 0.13 wt.%, 0.14 wt.%, 0.15 wt.%, 0.16 wt.%, 0.17 wt.%, 0.18 wt.%, 0.19 wt.%, 0.2 wt.%) based on the total mass of the conductive polymer hydrogel.

[0066] In some embodiments of the present invention, the preparation method of interpenetrating network conductive polymer hydrogel (IPNCH@PAAc-PVA) includes: firstly, completely dissolving polyvinyl alcohol (PVA) in deionized water, then adding NaOH to the polyvinyl alcohol solution to form an alkaline environment to obtain an alkaline solution, and reacting for a certain period of time to obtain a first mixed solution; dispersing freeze-dried PEDOT:PSS into the first mixed solution to obtain a second mixed solution; then completely dissolving the monomer acrylic acid (AAc), crosslinking agent and initiator in the above-mentioned second mixed solution to obtain a first reaction solution; finally, pouring the first reaction solution into a corresponding mold and self-curing to obtain interpenetrating network conductive polymer hydrogel (IPNCH@PAAc-PVA).

[0067] In some embodiments, in the preparation method of the above-mentioned interpenetrating network conductive polymer hydrogel (IPNCH@PAAc-PVA), the concentration of the alkali solution is 0.2-0.8 g / mL (e.g., 0.2 g / mL, 0.3 g / mL, 0.4 g / mL, 0.5 g / mL, 0.6 g / mL, 0.7 g / mL, 0.8 g / mL, etc.); the mass ratio of polyvinyl alcohol to the volume of the alkali solution is 1 mg: (0.0001-0.0005) ml; polyethylene... The mass ratio of enol to poly(3,4-ethylenedioxythiophene):polystyrene sulfonate is (200-300):1, for example, 200:1, 210:1, 220:1, 230:1, 240:1, 250:1, 260:1, 270:1, 280:1, 290:1, 300:1; the mass ratio of the polymeric monomer acrylic acid (AAc) to the poly(3,4-ethylenedioxythiophene):polystyrene sulfonate is (150-... The crosslinking agent ratio is 230:1, for example, 150:1, 160:1, 170:1, 180:1, 190:1, 198:1, 200:1, 210:1, 220:1, 230:1, etc.; based on the total mass of the conductive polymer hydrogel, the mass fraction of the crosslinking agent is 0.03wt.% to 0.09wt.% (for example, 0.03wt.%, 0.04wt.%, 0.05wt.%, 0.06wt.%, 0.07wt.%). 0.08 wt.%, 0.09 wt.%, etc.); based on the total mass of the conductive polymer hydrogel, the mass fraction of the initiator added is 0.1 wt.% to 0.2 wt.% (e.g., 0.1 wt.%, 0.11 wt.%, 0.12 wt.%, 0.13 wt.%, 0.14 wt.%, 0.15 wt.%, 0.16 wt.%, 0.17 wt.%, 0.18 wt.%, 0.19 wt.%, 0.2 wt.%, etc.).

[0068] In some embodiments, the crosslinking agent described above includes, but is not limited to, NN-methylenebisacrylamide (BIS), and the initiator includes, but is not limited to, sodium persulfate (SPS), ammonium persulfate (APS), and other initiators.

[0069] According to some embodiments of the present invention, a method for preparing conductive polymer hydrogel (PSCH) includes: adding an ionic compound, a crosslinking agent, and an initiator to an aqueous solution of poly(3,4-ethylenedioxythiophene):polystyrene sulfonate (PEDOT:PSS), and stirring to obtain a third mixed solution; and polymerizing the third mixed solution at a predetermined temperature for a predetermined time to obtain conductive polymer hydrogel (PSCH).

[0070] In some embodiments of the present invention, a method for preparing a phase-separated conductive polymer hydrogel (PSCH) includes: vigorously stirring a PEDOT:PSS aqueous solution for a certain period of time, then adding an ionic compound, a crosslinking agent, and an initiator to the PEDOT:PSS aqueous solution and stirring to obtain a third mixed solution; pouring the third mixed solution into a mold and covering it with a glass slide, and then carrying out a thermal polymerization reaction on a hot plate at a predetermined temperature for a predetermined time to obtain the conductive polymer hydrogel (PSCH).

[0071] In some embodiments, the crosslinking agent described above includes, but is not limited to, NN-methylenebisacrylamide (BIS), and the initiator includes, but is not limited to, sodium persulfate (SPS), ammonium persulfate (APS), and other initiators.

[0072] In some embodiments, the aforementioned ionic compounds include one or more of 1-ethyl-3-methylimidazolium ethyl sulfate (EMIM: ES), 4-(3-butyl-1-imidazolium)-1-butyrate triacid (BIM: BSA3), 1-butyl-3-methylimidazolium tetrafluoroborate (Bmim: BF4), [2-(methacryloyloxy)ethyl]dimethyl-(3-sulfonylpropyl)ammonium hydroxide (SBMA), and 3-[dimethyl-[3-(2-methylprop-2-enoylamino)propyl]ammonium]propane-1-sulfonate (SBAA). Thus, molecules with sulfonate or sulfonylimide anions are incorporated into the conductive polymer PEDOT:PSS as effective dopant to regulate phase separation and promote the aggregation of the conductive polymer PEDOT:PSS, forming a phase-separated conductive polymer hydrogel (PSCH). The aforementioned ionic compounds exhibit excellent solubility in both water and conductive polymer matrices, thus effectively softening the PSS domain in PEDOT:PSS. The highly acidic sulfonate or sulfonylimide groups interact strongly with the positively charged PEDOT segments and polar groups on the polymer chains through electrostatic and hydrogen bonding interactions. This dual interaction not only stabilizes the doped state but also induces controlled, uniform phase separation and conductive polymer aggregation, thereby forming an interconnected conductive network and achieving a synergistic effect of conductivity and stretchability.

[0073] In some embodiments, in the method for preparing conductive polymer hydrogel (PSCH), the mass ratio of the ionic compound to poly(3,4-ethylenedioxythiophene):polystyrene sulfonate is (15–25):1, for example, 15:1, 16:1, 17:1, 18:1, 19:1, 20:1, 20.8:1, 21:1, 22:1, 23:1, 24:1, 25:1, etc.; based on the total mass of the conductive polymer hydrogel, the mass fraction of the crosslinking agent added is 0.03 wt.% to 0.09 wt.% (for example, 0.03 wt.%, 0.04 wt.%, 0.05 wt.%). The initiator was added at a mass fraction of 0.1 wt.% to 0.2 wt.% (e.g., 0.1 wt.%, 0.11 wt.%, 0.12 wt.%, 0.13 wt.%, 0.14 wt.%, 0.15 wt.%, 0.16 wt.%, 0.17 wt.%, 0.18 wt.%, 0.19 wt.%, 0.2 wt.%, etc.) based on the total mass of the conductive polymer hydrogel; the predetermined temperature was 75–110 °C; and the predetermined time was 30–60 min. Therefore, a conductive polymer hydrogel (PSCH) with a high yield can be prepared under the above conditions, and the conductive polymer hydrogel (SPCH) exhibits good flexibility and high conductivity.

[0074] According to some embodiments of the present invention, a method for preparing a pure conductive polymer hydrogel (PCH) includes: adding a polar solvent to an aqueous solution of poly(3,4-ethylenedioxythiophene):polystyrene sulfonate (PEDOT:PSS) and stirring to obtain a fourth mixed solution; subjecting the fourth mixed solution to drying, annealing and hydration treatments in sequence to obtain a conductive polymer hydrogel (PCH).

[0075] In some embodiments of the present invention, the method for preparing pure conductive polymer hydrogel (PCH) includes: vigorously stirring a PEDOT:PSS aqueous solution for a certain period of time, then adding a polar solvent to the PEDOT:PSS aqueous solution, stirring and reacting at room temperature for a certain period of time to obtain a fourth mixed solution; directly pouring the fourth mixed solution into a template, drying it at 50-75°C for a certain period of time, then annealing it multiple times (2-5 times) at an annealing temperature and rehydrating it to obtain pure PEDOT:PSS conductive polymer hydrogel (PCH).

[0076] In some specific embodiments, the volume fraction of the polar solvent in the fourth mixed solution is 20 vol% to 30 vol% (e.g., 20 vol%, 22 vol%, 25 vol%, 28 vol%, 30 vol%); the concentration of poly(3,4-ethylenedioxythiophene):polystyrene sulfonate in the aqueous solution is (8-16) mg / ml (e.g., 8 mg / ml, 9 mg / ml, 10 mg / ml, 11 mg / ml, 12 mg / ml, 13 mg / ml, 14 mg / ml, 15 mg / ml, 16 mg / ml); the annealing treatment includes heating the dried fourth mixed solution to the annealing temperature, and then allowing it to cool naturally at room temperature, repeating this process 2-5 times; the annealing temperature is 80-200°C; the natural cooling time is 10-120 min. Therefore, under the above conditions, conductive polymer hydrogels (PCH) with high yield can be prepared, and the conductive polymer hydrogels (PCH) have good flexibility and high conductivity.

[0077] According to embodiments of the present invention, in the above-described method for preparing pure conductive polymer hydrogel (PCH), a polar solvent is introduced into the conductive polymer PEDOT:PSS, followed by controlled dry calcination to generate pure conductive polymer hydrogel (PCH). In the above-described preparation method, a polar solvent (e.g., dimethyl sulfoxide (DMSO), ethylene glycol (EG), glycerol, N,N-dimethylacetamide (DMF), and tetrahydrofuran (THF)) is added to the conductive polymer PEDOT:PSS system. Through heating, drying annealing, and rehydration processes, during the annealing process, the PEDOT segments can be further rearranged and reinforced with π-π stacking, and the PSS groups migrate to the outer layer, forming PEDOT-rich conductive channels. These polar solvents can be removed to a certain extent, resulting in a nearly pure conductive polymer hydrogel. Moreover, the removal of these polar solvents is not a simple boiling evaporation (simple boiling evaporation, lacking sufficient time and temperature control, often fails to effectively induce the ordered arrangement of PEDOT segments, leading to low conductivity), but rather results from diffusion and desorption processes, as well as differences in interaction with the matrix. The addition of polar solvents improves the crystallinity and orderliness of the conductive polymer nanocrystals, thereby enhancing electrical properties. Furthermore, a pure conductive polymer hydrogel without other components avoids problems such as impaired conductivity, uneven mechanical and electrical properties, and cytotoxicity caused by other fillers in the hydrogel network.

[0078] In another aspect, the present invention provides an application of the conductive polymer hydrogel described above, or a conductive polymer hydrogel prepared by the method described above, in biological tissues. Thus, suitable conductive polymer hydrogels can be selected for implantation into different biological tissues based on their elastic modulus and conductivity, thereby enhancing the flexibility and extensibility of the conductive polymer hydrogels used as epidermal or implantable bioelectrodes.

[0079] According to embodiments of the present invention, the conductive polymer hydrogel of the present invention can be applied to different biological tissues such as brain tissue, nerve tissue, heart, muscle and skin epidermis, depending on its different elastic modulus (i.e. flexibility) and conductivity.

[0080] Example

[0081] The relevant performance testing methods involved in the following embodiments include the following tests and testing methods:

[0082] 1. Mechanical properties (elastic modulus and viscous modulus):

[0083] The elastic and viscous moduli of the hydrogel were measured using a rheometer (MCR301, Anton-Paar). In frequency sweep testing, the shear modulus was measured at 1% strain within a frequency range of 0.1–10 Hz. The mechanical properties of the conductive polymer hydrogel were tested using an electric universal material measuring instrument (WDW3020, Changchun Kexin) equipped with a 100 N load sensor. Tensile tests were performed on rectangular samples (25 mm x 8 mm x 2 mm) at a tensile rate of 60 mm / min. -1 The cyclic stretching rate is 60 mm / min. -1 The stretch ratio is 200%. The formula for calculating tensile stress (σ) is:

[0084] Where F is the tensile load and A is the cross-sectional area.

[0085] The elongation (λ) is calculated as the elongation length (l) relative to the original length (l0):

[0086] To test the interfacial toughness of the hydrogel with different organs, the organs were cut after being frozen at -20°C. Hearts were cut into slices measuring 1 cm x 1 cm x 5 mm. Skin was cut into slices 1 cm long and 1 cm wide, retaining its original thickness. Before use, organ specimens were kept moist in phosphate-buffered saline (PBS) at 4°C. The hydrogel was then attached to the heart, skin, and glass. A standard 90° peel test was performed on the prepared samples using a mechanical testing machine (WDW3020, Changchun Kexin Test Instruments Co., Ltd.). All tests were conducted at a constant peel speed of 30 mm / min. Peel strength was calculated by dividing the peel force by the width of the adhered sample.

[0087] 2. Analysis of electrical and electrochemical properties

[0088] The electrical and electrochemical properties of the hydrogel were tested using a CHI660E electrochemical workstation (Shanghai Zhenhua Instruments Co., Ltd.). To measure conductivity, the hydrogel was filled between two parallel copper electrodes connected in an electrical circuit, and the corresponding voltage data was obtained by applying a 5mA DC current. Resistance (R) was calculated using Ohm's law.

[0089]

[0090] Where R is the average resistance, V is the voltage, and I is the current. Then, the conductivity (σ) is determined according to the following formula:

[0091]

[0092] Where σ is the conductivity, ρ is the resistivity, L is the distance between the two copper electrodes, and A is the cross-sectional area of ​​the hydrogel sample (calculated as length × width). Generally, the hydrogel is manufactured using a mold measuring 10*10*1mm. All actual geometric parameters, including material thickness and electrode spacing, are precisely measured using calipers to ensure accuracy.

[0093] The hydrogel electrode is adhered between two reference electrodes (Ra and Rc), ensuring equal spacing (Lab = Lbc), and the interfacial resistance (Rb) is measured. The resistance between each pair of the three points (Rab, Rbc, and Rac) is measured, and the interfacial resistance Rb is calculated using the following formula:

[0094]

[0095] This value reflects the contact resistance of the hydrogel-electrode interface and helps to gain a deeper understanding of the stability and quality of electrical contacts in bioelectronic applications.

[0096] To measure electrical impedance (EIS), a hydrogel was sandwiched between two steel electrodes (10 mm x 10 mm) and equilibrated in PBS before testing. The electrodes were connected to an electrochemical workstation. Impedance measurements were performed between 10⁻¹ and 10⁴ Hz with an amplitude of 5 mV. For volume coefficient (CV) testing, a hydrogel film was sandwiched as an electrode with a coating area of ​​1 cm². CV was measured at a rate of 100 mV / s from open-circuit voltage (OCV) -0.2 V to OCV +0.2 V. The CSC was calculated from the CV data.

[0097]

[0098] Where v is the scan rate, E1 and E2 are the potential windows, i is the current at each potential, and A is the electrode area.

[0099] 3. In vitro biocompatibility

[0100] Hydrogels were prepared and immersed in culture medium to remove unreacted monomers. Simultaneously, a 24-hour UV sterilization treatment was performed. Typically, 10 mL of DMEM medium and 1 g of the above sample were immersed separately at 37°C for 24 hours. Before use, 10% v / v fetal bovine serum and 100 U / mL penicillin-streptomycin-supplemented DMEM were added. Unincubated samples served as controls. Except for the blank group, logarithmic growth phase BHK21 cells were seeded at a density of 5000 cells per well in 96-well plates. After cell adhesion (approximately 4 hours), the experimental groups were replaced with 100 μL of extraction medium for different hydrogels and cultured for 24 and 72 hours. After culture, 100 μL of fresh DMEM medium was added to each well, followed by 10 μL of CCK-8 reagent. After another 2 hours of culture, the optical density (OD, representing cell density) of each well was measured using a microplate reader at 450 nm. The cell viability was calculated using the following formula:

[0101]

[0102] Where Ac is the absorbance of the control group (supplemented with DMEM, without extraction of hydrogel), As is the absorbance of the experimental group, and A b It is the absorbance of the empty autoarray (without cells).

[0103] 4. Biocompatibility in vivo

[0104] Hydrogel samples were cut into 1 cm x 1 cm length and width for implantation; the sham-operated group served as the control group. After aseptic preparation, all samples were immersed in PBS for 24 hours to remove unreacted monomers and sterilized with ultraviolet light before use. Animals were anesthetized with 10% chloral hydrate, and the implants were placed subcutaneously on the dorsal side of the animal, ensuring no overlap between implants. Animals were sacrificed 7 and 14 days after implantation, and the relevant subcutaneous areas were excised and fixed with paraformaldehyde for 24 hours for histological analysis. The fixed tissue samples were then placed in 70% ethanol for histological processing and H&E staining.

[0105] 5. Monitor epidermal electrophysiological signals

[0106] Electromyography (EMG), electrocardiography (ECG), and electroencephalography (EEG) signals from volunteers were detected using ESP32, AD8232, and TGAM (Jiangsu New Future Technology Co., Ltd.). Informed written consent was obtained from all participants prior to the study. Since this experiment did not require the collection of any human samples, approval from national or institutional ethics committees was not necessary. For EMG, hydrogels were adhered to both ends of the muscles being tested, serving as working and reference electrodes. For ECG, hydrogels were adhered to both wrists. For EEG, hydrogels were adhered to both ears and above the left eyebrow. After signal extraction, bandpass filtering was performed using Matlab. The signal-to-noise ratio (SNR) was calculated using the following formula:

[0107]

[0108] 6. Monitoring implantable electrophysiological signals

[0109] PhysioTel TM A fully implantable telemetry system (DSI, USA) combined with hydrogel electrodes is used for implantable electrocardiogram (ECG) and electromyography (EMG) recording. Implantable PhysioTel TM The lead wire and hydrogel electrode were co-molded in a mold, allowing the tough lead wire to be completely covered by hydrogel. Four hydrogel electrodes were implanted on each side of the back muscles and at both ends of the apical line. Surgical details are as follows: Adult BALB / c mice were used solely for evaluating the performance of the hydrogel electrodes. Mice were anesthetized with pentobarbital (1 wt%, 80 mg / kg), and an intermuscular pocket was created between the internal oblique and transversus abdominis muscles through a tendon incision. The transmitter was inserted into the subfascia of the external oblique muscle and fixed with nylon sutures, with the antenna perpendicular to the midline. The hydrogel electrode was then implanted at the designated site, and the skin was sutured. After implantation, the muscles and skin returned to normal. The device was turned on for 30 minutes weekly to record electrocardiogram and electromyogram signals for 3 weeks.

[0110] 7. Material Characterization

[0111] Scanning electron microscopy (SEM): The prepared hydrogel was flash-frozen in liquid nitrogen, becoming brittle and exposing the cross-section. During embrittlement, care was taken to avoid touching or squeezing the fracture surface with tweezers to prevent affecting its pore structure. Then, at 4 × 10⁻⁶ mm, the cross-section was examined. -3 The hydrogel was rapidly transferred to a freeze dryer under vacuum conditions by Torr and dried for 24 hours. Cross-sectional images were obtained using a scanning electron microscope (S-4800, Hitachi).

[0112] Fourier transform infrared spectroscopy was performed using a Nicolet 6700FTIR spectrometer (Thermo Fisher Scientific (Shanghai) Co., Ltd.) at a range of 400-4000 cm⁻¹. -1 The results were obtained within the scanning range. UV-Vis absorption spectra were recorded using a ULS4096CL-EVO spectrometer (Avantes, China) within the relevant wavelength range. Raman spectra were acquired using a LabRAM HR800 confocal microscope (HORIBAJOBINYVON) equipped with a 633 nm laser, and the spectra were acquired under ambient conditions. Atomic force microscopy (AFM) tests were performed using an MFP-3D-SA (Asylum Research) in tapping mode with a silicon cantilever to evaluate the surface morphology and phase distribution of the samples. X-ray diffraction (XRD) patterns were acquired using Cu-Kα radiation on a Bruker D8 advanced diffractometer at 40.0 kV and 120 mA.

[0113] 8. Statistical Analysis

[0114] Statistical analysis was performed using GraphPadPrism (GraphPad software) and Origin (Origin software). All data are expressed as mean ± standard deviation (SD). Error bars represent standard deviation. Tukey's multiple comparison test was used to assess statistical significance, with thresholds of *P<0.05, **P<0.01, and *8*P<0.001°. Significance was defined as P<0.001.

[0115] Example 1

[0116] The preparation methods of interpenetrating network conductive polymer hydrogels (IPNCH@CAPAM) include:

[0117] First, carrageenan (CA, 0.05 g) was completely dissolved in 5 mL of deionized water at 60 °C. Then, NaOH (300 μL, 0.5 g / mL) was added to the carrageenan solution to create an alkaline environment, resulting in an alkaline solution. Next, 10 mg of dopamine hydrochloride (DAHC1) was added to the alkaline solution, and after reacting for 20 min, a first mixed solution was obtained. 3 mg of freeze-dried PEDOT:PSS was dispersed into the first mixed solution to obtain a second mixed solution. Then, 2.2 mol / L (5 mL) of monomeric acrylamide (AAm) was dispersed into the second mixed solution. Finally, 3 mg of NN-methylenebisacrylamide (BIS) and 8 mg of sodium persulfate (SPS) were completely dissolved in the second mixed solution to obtain a first reaction solution. The first reaction solution was then poured into a corresponding mold and allowed to self-cur to obtain a conductive polymer hydrogel (IPNCH@CAPAM) with a yield of 95.22%.

[0118] Example 2

[0119] The preparation method of interpenetrating network conductive polymer hydrogel (IPNCH@PAM) includes: firstly, adding NaOH (300 μL, 0.5 g / mL) to 5 mL of deionized water to form an alkaline environment to obtain an alkaline solution; then adding 10 mg of dopamine hydrochloride (DAHC1) to the alkaline solution and reacting for 20 min to obtain a first mixed solution; dispersing 3 mg of freeze-dried PEDOT:PSS into the first mixed solution to obtain a second mixed solution; then dispersing 2.2 mol / L (5 mL) of monomeric acrylamide (AAm) into the second mixed solution; then completely dissolving 3 mg of NN-methylenebisacrylamide (BIS) and 8 mg of sodium persulfate (SPS) in the above second mixed solution to obtain a first reaction solution; finally, pouring the first reaction solution into a corresponding mold and self-curing to obtain conductive polymer hydrogel (IPNCH@PAM).

[0120] Example 3

[0121] The preparation method of interpenetrating network conductive polymer hydrogel (IPNCH@PAAc) includes: firstly, adding NaOH (300 μL, 0.5 g / mL) to 5 mL of deionized water to form an alkaline environment to obtain an alkaline solution; then adding 10 mg of dopamine hydrochloride (DAHC1) to the alkaline solution and reacting for 20 min to obtain a first mixed solution; dispersing 3 mg of freeze-dried PEDOT:PSS into the first mixed solution to obtain a second mixed solution; then dispersing 2.2 mol / L (5 mL) of monomer acrylic acid (AAc) into the second mixed solution; then completely dissolving 3 mg of NN-methylenebisacrylamide (BIS) and 8 mg of sodium persulfate (SPS) in the second mixed solution to obtain a first reaction solution; finally, pouring the first reaction solution into a corresponding mold and self-curing to obtain the conductive polymer hydrogel (IPNCH@PAAc).

[0122] Example 4

[0123] The preparation method of interpenetrating network conductive polymer hydrogel (IPNCH@PAAc-PVA) includes: firstly, 750 mg of polyvinyl alcohol (PVA) is dissolved in 5 mL of deionized water and heated to 80-90 °C to completely dissolve it. Then, NaOH (300 μL, 0.5 g / mL) is added to the polyvinyl alcohol solution to form an alkaline environment, resulting in an alkaline solution. After reacting for 20 min, a first mixed solution is obtained. 3 mg of freeze-dried PEDOT:PSS is dispersed into the first mixed solution to obtain a second mixed solution. Then, 1.65 mol / L (5 mL) of monomeric acrylic acid (AAc) is dispersed into the second mixed solution. Then, 3 mg of NN-methylenebisacrylamide (BIS) and 8 mg of sodium persulfate (SPS) are completely dissolved in the above second mixed solution to obtain a first reaction solution. Finally, the first reaction solution is poured into a corresponding mold and self-cured to obtain conductive polymer hydrogel (IPNCH@PAAc-PVA).

[0124] Example 5

[0125] The method for preparing phase-separated conductive polymer hydrogel (PSCH@IL1) includes: vigorously stirring a PEDOT:PSS aqueous solution (5 ml, containing 0.06 g of PEDOT:PSS) for 6 hours; then adding 1.25 g of 1-ethyl-3-methylimidazolium ethyl sulfate (EMIM:ES), 3 mg of NN-methylenebisacrylamide (BIS), and 8 mg of ammonium persulfate (APS) to the PEDOT:PSS aqueous solution and stirring for 24 hours to obtain a third mixed solution; pouring the third mixed solution into a mold and covering it with a glass slide, and then carrying out a thermal polymerization reaction on a hot plate at 90°C for 45 min to obtain the conductive polymer hydrogel (PSCH@IL1).

[0126] Example 6

[0127] The method for preparing the phase-separated conductive polymer hydrogel (PSCH@IL2) is basically the same as that in Example 5, except that the ionic compound is 4-(3-butyl-1-imidazolium)-1-butyric acid triacid salt (BIM: BSA3).

[0128] Example 7

[0129] The method for preparing the phase-separated conductive polymer hydrogel (PSCH@IL3) is basically the same as that in Example 5, except that the ionic compound is 1-butyl-3-methylimidazolium tetrafluoroborate (Bmim: BF4).

[0130] Example 8

[0131] The method for preparing the phase-separated conductive polymer hydrogel (PSCH@SBMA) is basically the same as in Example 5, except that the ionic compound is [2-(methacryloyloxy)ethyl]dimethyl-(3-sulfopropyl)ammonium hydroxide (SBMA). The yield was 89.84%.

[0132] Example 9

[0133] The method for preparing the phase-separated conductive polymer hydrogel (PSCH@SBAA) is basically the same as that in Example 5, except that the ionic compound is 3-[dimethyl-[3-(2-methylprop-2-enoylamino)propyl]ammonium]propane-1-sulfonate (SBAA).

[0134] Example 10

[0135] The method for preparing pure conductive polymer hydrogel (PCH@EG) includes: vigorously stirring a PEDOT:PSS aqueous solution (3.75 ml, containing 0.045 g of PEDOT:PSS) for 6 hours; then adding the polar solvent ethylene glycol (EG) to the PEDOT:PSS aqueous solution to 25 vol% of the final solution (i.e., the volume fraction of the polar solvent in the final solution is 25%); stirring and reacting at room temperature for 24 hours to obtain a fourth mixed solution; directly pouring the fourth mixed solution into a template and drying it at 60 °C for 24 hours; then annealing it multiple times at 130 °C (three annealings, each time raising the temperature to 130 °C and then allowing it to cool naturally at room temperature for 30 minutes before repeating the above annealing process) and rehydrating it to obtain pure PEDOT:PSS conductive polymer hydrogel (PCH@EG) with a yield of 85.48%.

[0136] Example 11

[0137] The method for preparing the pure conductive polymer hydrogel (PCH@DMSO) is basically the same as that in Example 10, except that the polar solvent is dimethyl sulfoxide (DMSO).

[0138] Example 12

[0139] The method for preparing pure conductive polymer hydrogel (PCH@Glycerol) is basically the same as that in Example 10, except that the polar solvent is glycerol.

[0140] Example 13

[0141] The method for preparing pure conductive polymer hydrogel (PCH@DMF) is basically the same as that in Example 10, except that the polar solvent is N,N-dimethylacetamide (DMF).

[0142] Example 14

[0143] The method for preparing pure conductive polymer hydrogel (PCH@THF) is basically the same as that in Example 10, except that the polar solvent is tetrahydrofuran (THF).

[0144] Related tests:

[0145] 1. Mechanism Characterization

[0146] (1) In Examples 1 to 4 for the preparation of interpenetrating network conductive polymer hydrogels (IPNCH), taking the interpenetrating network conductive polymer hydrogel (IPNCH@CAPAM) in Example 1 as an example, the crosslinking mechanism in the formation process of the conductive polymer hydrogel (IPNCH) is described:

[0147] Reference Figure 1 In (A), the hydrogel IPNCH@CAPAM is an interpenetrating polymer network (IPN) hydrogel. Dopamine (DA) was initially polymerized in an alkaline carrageenan (CA) solution containing NaOH to generate polydopamine (PDA). Then, freeze-dried PEDOT:PSS was mixed with acrylamide (AAm, monomer), NN-methylenebisacrylamide (BIS, crosslinking agent), and sodium persulfate (SPS, initiator) and spontaneously polymerized within 3 minutes to form a conductive polymer hydrogel IPNCH@CAPAM with an interpenetrating network. (See reference...) Figure 1 In (B), scanning electron microscopy (SEM, scale bar: 10 μm) images of freeze-dried IPNCH@CAPAM show a typical three-dimensional (3D) interconnected porous microstructure, which provides the hydrogel with the conditions to withstand large deformations.

[0148] Reference Figure 1The successful synthesis of IPNCH@CAPAM hydrogel was also confirmed by Fourier transform infrared spectroscopy (FTIR) in the middle (C), which can be achieved by catechol groups (at 3500 cm⁻¹). -1 The broad hydroxyl stretching vibration absorption peak at 1450-1600 cm⁻¹, and the absorption peak at 1450-1600 cm⁻¹. -1 The absorption peaks at the benzene ring stretching vibrations (at 1122 cm⁻¹) and poly(3,4-ethylenedioxythiophene): poly(styrenesulfonic acid) (at 1122 cm⁻¹) -1 -SO3 at the location - Characteristic peaks, and at 1458 cm⁻¹ -1 This is reflected in the additional peaks of the PEDOT characteristic peak (at which point).

[0149] Reference Figure 1 The characteristic absorption peak of the catechol group observed at a wavelength of 280 nm in the medium (D) ultraviolet-visible spectroscopy (UV-Vis) further confirms that PDA was successfully incorporated into IPNCH@CAPAM.

[0150] Furthermore, XRD patterns revealed a broad diffusion peak around 20–24° after PEDOT was added to the IPNCH@CAPAM system, indicating enhanced interaction between PEDOT and the hydrogel matrix, leading to a more disordered structure. The resulting amorphous properties facilitate the formation of a flexible conductive network, thereby improving the electrical and mechanical properties of the system. The conductive polymer is dispersed within the hydrogel network composed of CA and PAM chains, forming conductive channels.

[0151] Furthermore, the binding energy of the CS bonds in PEDOT shows a slight shift in IPNCH@CAPAM during XPS spectroscopy. This shift indicates the existence of physical interactions between PEDOT and components such as CA, PAM, or PDA, such as hydrogen bonds or π-π interactions.

[0152] The aforementioned interactions alter the local environment of PEDOT, further enhancing the system's conductivity and mechanical strength. PDA and PEDOT form an electron donor-acceptor system, where PEDOT promotes electron transfer to quinones, preventing further oxidation of catechols. The dynamic redox-active system within the IPNCH@CAPAM hydrogel network retains sufficient catechols, contributing to the adhesiveness of the IPNCH@CAPAM hydrogel. Simultaneously, the polymerization inhibition of catechol groups in PDA also suppresses PAM chain growth, reducing entanglement between long PAM chains. The physical interaction between the linear CA chains and PAM chains improves the tensile strength and toughness of the IPNCH@CAPAM hydrogel.

[0153] (2) For Examples 5 to 9 of the preparation of phase-separated conductive polymer hydrogels (PSCH), taking the conductive polymer hydrogel (PSCH@SBMA) in Example 8 as an example, the crosslinking mechanism in the formation process of the phase-separated conductive polymer hydrogel (PSCH) is described:

[0154] Reference Figure 2 In (A), by introducing ionic compounds, a PSCH@SBMA hydrogel composed of a conductive polymer network and a multifunctional polyzwitterionic poly(sulfobetaine) (pSB) network was successfully constructed. SBMA can self-polymerize in water to form SBMA@water via a free radical mechanism. This induced association polymerization effect enables the formation of nanoscale associations in concentrated aqueous solutions of SBMA monomers, thereby increasing the effective monomer concentration and facilitating successful polymerization. Ammonium persulfate (APS) was used as a thermal initiator, and N,N-methylenebisacrylamide (BIS) was used as a crosslinking agent. After heating at high temperature for approximately 45 minutes, a covalently crosslinked pSB network was formed.

[0155] Reference Figure 2 In section (B), the formation of the covalently crosslinked pSB network in PSCH@SBMA was studied using Fourier transform infrared spectroscopy. The SBMA monomer was observed at 1305 cm⁻¹. -1 A characteristic CN absorption peak is observed at this point, but this peak is almost invisible in the spectra of SBMA@water and PSCH@SBMA. This is because after polymerization... Group P-SO3 - This is due to the electrostatic shielding effect of the radical.

[0156] Reference Figure 2 In the Raman spectra of SBMA@water and PSCH@SBMA (C), the peak of pSB appears at 1731 cm⁻¹. -1 (OC = O oscillation), 1160cm -1 and 1036cm -1 At (S=O asymmetric oscillation). In the spectra of SBMA@water and PSCH@SBMA, 1637cm -1 The Raman peak at the point also leads to the disappearance of C=C bonds. Characterization of PSCH@SBMA with different SBMA contents reveals that the electrostatic shielding provided by the pSB network reduces the attraction between PEDOT and PSS, promotes the transition from a compact core-shell structure to a more extended PEDOT chain, and ultimately promotes the formation of a percolating conductive network.

[0157] X-ray diffraction (XRD) results showed that the characteristic peaks of the monomer in the PSCH@SBMA spectrum disappeared, and a characteristic peak consistent with that of the SBMA@WATER polymer appeared at 2θ≈19°. Meanwhile, the large peak may have overshadowed the characteristic peaks of PEDOT. For PSCH@SBMA containing SBMA, a new peak appeared at 2θ=20-30°, which may be due to changes in the crystal structure, more pronounced phase separation, and improved charge transport properties.

[0158] Reference Figure 2 In (D), the current image measured by conductivity atomic force microscopy also shows phase separation between PEDOT-rich domains (bright) and PSS-rich domains (dark). The PEDOT:PSS transitions to a longer chain, and PSCH@SBMA shows a clearer conductive path, which is conducive to the formation of a more efficient electron transfer path.

[0159] Furthermore, the covalently cross-linked pSB network and the conductive polymer network are in a semi-interpenetrating state, further endowing PSCH@SBMA with tissue-like mechanical properties. In addition, the strong electrostatic interaction between pSB and charged groups on the tissue surface endows PSCH@SBMA with bioadhesive properties.

[0160] Therefore, phase-separated conductive polymer hydrogels overcome the inherent brittleness of conductive polymers and possess high flexibility and extensibility.

[0161] (3) For Examples 5 to 9, which prepared pure conductive polymer hydrogels (PCH), the crosslinking mechanism during the formation of the phase-separated conductive polymer hydrogel (PCH) is described using the conductive polymer hydrogel (PCH@EG) in Example 10 as an example:

[0162] Reference Figure 3 In section (A), a pure conductive polymer hydrogel was prepared by introducing a polar solvent. Ethylene glycol (EG) was added to the conductive polymer PEDOT:PSS and dried for 24 hours. Gelation was essentially completed within about 4 hours, followed by controlled drying, multiple annealing, and rehydration to obtain the PCH@EG hydrogel.

[0163] Reference Figure 3 In (B), the characteristic Fourier transform infrared absorption peak of EG (3100-3500 cm⁻¹) can be clearly observed in the PEDOT:PSS aqueous solution containing EG (corresponding curve PEDOT:PSS+25%EG). -1(hydroxyl absorption peak). However, these absorption peaks are significantly weakened in the Fourier transform infrared spectrum of dried annealed PEDOT:PSS (corresponding curve PCH@EG, wet), and further disappear in freeze-dried PCH@EG (corresponding curve PCH@EG, free-drying). The nuclear magnetic resonance (NMR) spectrum of PCH@EG further confirms the removal of EG. In aqueous colloidal dispersions, PEDOT:PSS tends to form micellar microstructures composed of a hydrophobic PEDOT core and a hydrophilic PSS shell. In the original solution, due to the lack of interconnected conductive polymer nanofibers, it dissociates into fragmented microgels in a humid environment. The addition of polar solvents such as ethylene glycol effectively transforms the conductive polymer microgel particles from a closed and / or folded state into a linear long-chain structure. This process facilitates the formation of larger PEDOT-rich nanofibers and interchain entanglements between PSS chains during drying.

[0164] Reference Figure 3 Further XRD analysis was performed on PCH@EG in (C) of Example 10. The results showed that varying the amount of ethylene glycol (EG) in Example 10 resulted in PEDOT diffraction peaks at 2θ = 7°, 18°, and 26° for PCH@EG with different EG contents. The peak at 2θ = 18° was likely obscured by a broad background peak at 26°, while the sub-peak varied with EG content, indicating a structural rearrangement in the hydrogel due to changes in doping levels. For 25 vol.% PCH@EG, a new peak appeared at 2θ = 30–35°, attributed to changes in crystal structure, more pronounced phase separation, and improved charge transport properties. Pull-out spectroscopy further confirmed these findings, revealing enhanced π-π stacking interactions between PEDOT chains. This structural refinement correlated with atomic force microscopy images, showing brighter, more continuous PEDOT-rich domains in the 25 vol.% PCH@EG sample, indicating well-defined nanoscale phase separation and improved electrical pathways.

[0165] Reference Figure 3 In the (D) atomic force microscopy morphology image, an increased surface roughness (Ra) is also observed, which is consistent with the presence of obvious nanofiber features.

[0166] Furthermore, thermogravimetric-mass spectrometry (TG-MS) and high-performance liquid chromatography (HPLC) analyses confirmed that EG was effectively removed after annealing and soaking, verifying the role of polar solvents as transient processing aids, which can promote chain rearrangement without leaving residual solvents that may affect performance.

[0167] Redesigning the phase distribution of conductive polymers is a key strategy for transforming them into flexible, stretchable hydrogels; the tight connection between conductive and hydrophobic domains provides excellent conductivity and water stability. The structural analysis above confirms that specific preparation methods result in unique microstructures, which in turn lead to variations in the mechanical and electrical properties of hydrogels constructed using different methods.

[0168] 2. Mechanical and electrical properties:

[0169] Schematic diagrams of the interpenetrating network conductive polymer hydrogels (IPNCH) prepared in Examples 1-4 can be found in [reference]. Figure 1 (A) ; Schematic diagrams of the phase-separated conductive polymer hydrogels (PSCH) prepared in Examples 5-9 can be found in (A). Figure 2 (A) ; Schematic diagrams of the pure conductive polymer hydrogels (PCH) prepared in Examples 10-14 can be found in [reference]. Figure 3 (A) in the middle.

[0170] The conductive polymer hydrogels prepared in Examples 1-14 were subjected to oscillation tests in a frequency range of 0.1 to 10 Hz to test the modulus (Pa) and conductivity (S / m) of the hydrogels. The test results for the elastic modulus (G') and viscous modulus (G") of the interpenetrating network conductive polymer hydrogels (IPNCH) prepared in Examples 1-4 can be found in [reference missing]. Figure 4 (A) In the diagram, the conductivity test results can be found in [reference]. Figure 4 (B); The test results of the elastic modulus (G') and viscous modulus (G") of the phase-separated conductive polymer hydrogels (PSCH) prepared in Examples 5-9 can be found in (B). Figure 4 (C) In the diagram, the conductivity test results can be found in [reference]. Figure 4 (D); The test results of the elastic modulus (G') and viscous modulus (G") of the pure conductive polymer hydrogels (PCH) prepared in Examples 10-14 can be found in (D). Figure 4 (E) The conductivity test results can be found in [reference]. Figure 4 (F) in the middle.

[0171] The results above show that the elastic modulus (G') of all conductive polymer hydrogels exceeds the viscous modulus (G"), consistent with their gel-like properties. Specifically, for interpenetrating network conductive polymer hydrogels (IPNCH), in the PEDOT:PSS hydrogel system forming the IPN, except for the PVA-based hydrogel, the elastic modulus of the other three systems is below 3.0 kPa, close to the modulus of brain soft tissue (0.2-3.1 kPa). Figure 4(A)). Among them, the elastic modulus of IPNCH@CAPAM hydrogel (corresponding to curve CA-PAM G') is only 0.1 kPa, exhibiting ultra-soft properties. The reason why this hydrogel has brain-level softness is that the catechol groups in polydopamine have a polymerization inhibitory effect, which can inhibit the growth of PAM chains and reduce the entanglement between long PAM chains. The uniformly dispersed conductive polymer interacts non-covalently with the CA and PAM polymer chains in the hydrogel, dissipating energy during deformation and maintaining the integrity of the main network of the hydrogel, thus forming a soft and tough property. This shows that it can serve as a bridge for mechanical matching and improve the long-term biomechanical interaction between bioelectronic devices and biological tissues; for phase-separated conductive polymer hydrogels (PSCH), the elastic modulus of phase-separated PEDOT:PSS hydrogels constructed from ionic compounds is between 12 and 77 kPa, covering the modulus range (1-100 kPa) of most related tissues. Figure 4 In (C), ionic compounds cause the conjugated polymer portion to aggregate while simultaneously softening the doped PSS domains, effectively reducing the elastic modulus of the conductive polymer hydrogel and giving it high ductility. For pure conductive polymer hydrogels (PCH), pure conductive polymer hydrogel systems constructed with polar solvents typically have a high elastic modulus, between 15-130 kPa, due to the lack of a supporting matrix. Figure 4 While the elastic modulus (E) is present, it is still several orders of magnitude lower than that of conventional rigid materials in bioelectronic devices, and even lower than the elastic modulus (1-10 MPa) of soft elastomers such as polydimethylsiloxane (PDMS). This is because, during dry annealing, the polar solvent, acting as an additive, promotes the recrystallization of PEDOT-rich nanofibers and the chain rearrangement of PEDOT:PSS. The uniform distribution of the PEDOT-rich rigid conjugated crystalline region and the PSS-rich soft region can generate good mechanical compliance in biological tissues. These results all confirm the advantages of the conductive polymer hydrogel of the present invention in matching the mechanical properties of biological tissues.

[0172] Regarding conductivity, the electrical properties of hydrogels that match those of biological tissues in a physiological environment are also crucial for ensuring stable bidirectional communication between devices and tissues. The conductivity of the conductive polymer hydrogels in Examples 1-14 above was tested. Figure 4In (B), (D), and (F) of the above, except for the hydrogels (PSCH@IL1 and PSCH@IL3) which have poor mechanical properties and are extremely brittle and difficult to test for conductivity, the conductivity of hydrogels IPNCH and PSCH ranges from 0.19 to 0.87 S / m. Some hydrogels PCH exhibit significantly higher conductivity, reaching 5.21 S / m (PCH@DMSO), which is attributed to the removal of non-conductive insulating components. Furthermore, the introduction of polar solvents promotes the coiling of conjugated polymer chains, increases their crystallinity, and establishes strong bonds between conjugated polymers.

[0173] The three types of conductive polymer hydrogels mentioned above complement each other: IPNCH hydrogel forms a soft base, PSCH hydrogel perfects the conductive pathway, and PCH hydrogel maximizes conductivity, thus collectively covering a broad mechatronic range and can be tailored for various bioelectronic applications. Different formulations within each hydrogel offer different advantages and trade-offs. Based on the test results of the elastic modulus (Pa), conductivity (S / m), and impedance (Ω) of the 14 conductive polymer hydrogel formulations in the above examples (e.g., ...), Figure 5 Among the formulations in A), three stand out due to their superior performance: IPNCH@CAPAM, PSCH@SBMA, and PCH@EG. The conductivity of these three conductive polymer hydrogels fully meets the requirements of electrophysiological signal monitoring and other bioelectronic applications.

[0174] Reference Figure 5 For components B, C, and D in Example 1, which exhibited superior performance, the optimal formulations were subsequently selected and optimized to achieve synergistic effects: IPNCH@CAPAM (with variations in the PEDOT:PSS content in Example 1), PSCH@SBMA (with variations in the SBMA content in Example 8), and PCH@EG (with variations in the EG content in Example 10, with volume fractions of 10%, 15%, 25%, 30%, and 35%, respectively). Overall, the relationship between hydrogel performance and component ratios was consistent. Specifically, higher PEDOT:PSS content, or better rearrangement of conjugated polymer nanofibers due to phase separation, resulted in stronger hydrogel conductivity and lower impedance at physiologically relevant frequencies. However, the introduction of more PEDOT-rich rigid conjugated regions inevitably led to an increase in hydrogel modulus.

[0175] According to the above Figure 5The experimental results show that the raw material formulations with high conductivity, low impedance, and low elastic modulus are as follows: IPNCH@CAPAM (containing 0.06 wt.% PEDOT:PSS), PSCH@SBMA (containing 25 wt.% SBMA), and PCH@EG (containing 25 wt.% EG). These three hydrogels collectively cover a wide range of mechatronics: IPNCH@CAPAM occupies the ultrasoft region (0.28 kPa, 1.99 S / m) and is ideal for neural interfaces; PSCH@SBMA bridges the gap between medium modulus (0.44 kPa) and peak conductivity (5.25 S / m), making it suitable for cardiac / muscle implants; PCH@EG achieves a modulus (15 kPa) that matches the skin while retaining high conductivity (3.79 S / m), making it the best choice for epidermal electronics.

[0176] 3. Mechanical properties of IPNCH@CAPAM (Example 1), PSCH@SBMA (Example 8), and PCH@EG (Example 10)

[0177] Rheological analysis was used to study the relationship between the elastic modulus (G') and viscous modulus (G") of hydrogels IPNCH@CAPAM, PSCH@SBMA, and PCH@EG and the frequency (0.1-10 Hz). Figure 6 As shown in Figure A, IPNCH@CAPAM hydrogel is one of the softest hydrogels. This ultra-soft material has an elastic modulus of approximately 280 Pa, well within the range of ultra-soft neural tissue (100-1500 Pa). The catechol groups of PDA effectively reduce the modulus by inhibiting PAM chain entanglement. This ultra-soft property makes IPNCH@CAPAM an ideal material for implantable bioelectronics, especially soft neural interfaces. Moreover, expansion further reduces its modulus, highlighting the importance of understanding how water absorption affects the overall properties of hydrogels in practical applications. PSCH@SBMA, prepared using zwitterionic modification, also exhibits good mechanical properties, but its modulus is slightly higher than that of IPNCH@CAPAM. PSCH@SBMA has an elastic modulus below 500 Pa, making it suitable for soft tissue applications, but slightly harder than neural tissue. The elastic modulus of PSCH@SBMA, combined with its flexibility, makes it suitable for implantable and wearable bioelectronic devices. The unsupported, purely conductive polymer hydrogel PCH@EG exhibits the highest elastic modulus of the three hydrogels, approximately 15 kPa, which is very close to the modulus of human skin (approximately 20 kPa). This makes PCH@EG ideally suited for epidermal bioelectronics, which requires extremely high mechanical compatibility with skin.

[0178] Cyclic shear strain (Strain, %) tests were performed on the hydrogel IPNCH@CAPAM. The G' and -G" of IPNCH@CAPAM were subjected to alternating step strain transitions from small strain (0.5%, 200 s) to large strain (300%, 100 s) at a fixed frequency (1 Hz). The test results are as follows: Figure 6 As shown in Figure B, this structure demonstrates that, in addition to its flexibility, IPNCH@CAPAM exhibits excellent self-recovery capabilities. In cyclic shear strain testing, the hydrogel rapidly recovers its original modulus after undergoing large strain deformation. When subjected to 300% strain followed by a 0.5% strain, the IPNCH@CAPAM hydrogel recovered to its initial modulus within 30 seconds. This rapid gel-sol transition highlights its potential in dynamic applications requiring reversible deformation. Similar to IPNCH@CAPAM, PSCH@SBMA exhibits excellent self-recovery properties in cyclic shear strain testing, responding quickly to strain changes.

[0179] A compression test was conducted on the hydrogel, and the test results are as follows: Figure 6 As shown in C and D, Figure 6 The illustration in C illustrates a schematic diagram of the experimental setup for the compression test. Figure 6 In the figure, C represents the compressive stress (Pa)-strain (%) curves for IPNCH@CAPAM, PSCH@SBMA, and PCH@EG. Figure 6D in the figure represents the compressive stress (Pa) curves of PSCH@SBMA and PCH@EG under 30% strain for 1000 consecutive compression load-unload cycles. In compression tests, the hydrogel IPNCH@CAPAM withstood up to 80% strain, maintaining a compressive stress of 100.12 kPa at 80% strain. Cyclic testing further confirmed the fatigue resistance of IPNCH@CAPAM during compression. In cyclic testing, the hydrogel IPNCH@CAPAM exhibited minimal stress change and maintained its original strength after 100 load-unload cycles at 20% strain. This robust performance indicates that IPNCH@CAPAM is suitable for stable, long-term bioelectronic applications, minimizing signal baseline drift during extended use. The hydrogel PSCH@SBMA exhibited a compressive stress of 1161.75 kPa at 82% strain, demonstrating higher compressive strength than IPNCH@CAPAM. This moderate compressive strength and the ability to maintain structural integrity after repeated compression indicate its suitability for applications subject to moderate mechanical stress. Despite its high modulus, PCH@EG still exhibits strong compressive strength, capable of withstanding up to 81% strain with a compressive stress of 1689.11 kPa, far exceeding that of IPNCH@CAPAM and PSCH@SBMA. Figure 6 (C in the text). Furthermore, PCH@EG showed a 16.9% reduction in stress after 1000 compression cycles (…). Figure 6 (D in the diagram). This indicates that EG enhances the flexibility and conductivity of PEDOT. The early stress decrease may be due to the weak initial interaction between PEDOT and EG, leading to molecular rearrangement. However, as EG improves flexibility, the system exhibits better stability in later cycles. Energy dissipation in the PCH@EG system also gradually decreases by 12.7%, eventually stabilizing. This suggests that the internal molecular chains reach equilibrium after the initial rearrangement, thereby reducing energy dissipation.

[0180] The tensile stress (Pa)-strain (%) curves of IPNCH@CAPAM and PSCH@SBMA were tested, as shown below. Figure 6 As shown in Figure E, the inset represents the tensile testing apparatus; the stress curve of IPNCH@CAPAM under 30% strain conditions after 1000 consecutive tensile loading-unloading cycles is shown in the figure. Figure 6As shown in F in the figure. In tensile tests, IPNCH@CAPAM exhibited remarkable elasticity, with an elongation at break of 530% and a tensile strength of 8.02 kPa. In 1000 tensile cycles of the IPNCH@CAPAM system, the stress remained relatively stable throughout the entire cycle. To comprehensively evaluate its fatigue resistance and obtain key parameters in the presence of microcracks, fatigue crack propagation tests were inevitably conducted. A 3000-cycle fatigue test under 30% strain conditions further confirmed its long-term stability. This combination of flexibility and strength makes hydrogel IPNCH@CAPAM ideally suited for flexible electronic devices requiring high flexibility limits, such as implantable bioelectronic devices. Tensile tests of hydrogel PSCH@SBMA showed an elongation at break of up to 789% and a tensile strength of 5.85 kPa. The combination of elongation and sufficient strength indicates that PSCH@SBMA can be used in wearable bioelectronic devices requiring high flexibility and skin adhesion.

[0181] A schematic diagram illustrating the interaction between IPNCH@CAPAM and PSCH@SBMA hydrogels and biological tissues can be found here. Figure 6 The catechol redox active system of G, IPNCH@CAPAM, can undergo covalent bonding and non-covalent interactions with various surfaces. Peel strength measurements of IPNCH@CAPAM and PSCH@SBMA on porcine skin are shown below. Figure 6 The results of the H, IPNCH@CAPAM peel strength test on pig skin after 15 repetitions can be found in [reference]. Figure 6 The I in the middle. Figure 6 The H in the figure demonstrates the excellent adhesion properties of IPNCH@CAPAM and PSCH@SBMA to pigskin, ensuring stable adhesion and conformal contact with various polar and non-polar substrates. Even after 15 consecutive peels on pigskin, the adhesion strength of IPNCH@CAPAM remains above 50% of the initial adhesion strength. Figure 6 The presence of I in this image highlights its strong, repeatable, and durable adhesion capabilities. This characteristic is crucial for bioelectronic devices, effectively reducing motion artifacts and ensuring stable signal acquisition during extended use.

[0182] In terms of adhesion, PSCH@SBMA also exhibits moderate bioadhesion. The zwitterions promote strong interactions with biological surfaces through hydrogen bonding, electrostatic attraction, and dipole-dipole interactions. Figure 6 (G and H in the text). These adhesive properties make it suitable for applications requiring close and stable contact with biological tissue for reliable signal detection.

[0183] In summary, the above analysis shows that all three types of hydrogels exhibit unique and stable mechanical properties, making them suitable for various bioelectronic applications. The swelling of hydrogels also significantly affects their mechanical properties. As observed in the tests, the shear modulus of the hydrogels decreased substantially after swelling, indicating a reduction of several orders of magnitude due to the softening of the hydrogel network structure. This reduction in mechanical strength is crucial because it affects the hydrogel's ability to retain stress and strain under operating conditions. Since typical skin deformation is less than 30%, the elasticity of the hydrogels is considered sufficient to meet the application requirements of wearable or implantable bioelectronic devices.

[0184] 4. Electrical properties of IPNCH@CAPAM (Example 1), PSCH@SBMA (Example 8), and PCH@EG (Example 10)

[0185] Figure 7-1 Figure A shows the conductivity measurements of hydrogels IPNCH@CAPAM, PSCH@SBMA, and PCH@EG, as well as a commercial gel electrode (CGE). Figure 7-1 Figure B shows a schematic diagram of a device in which hydrogel acts as a partial conductor in a closed circuit to light up an LED, along with optical images of the corresponding hydrogel IPNCH@CAPAM, PSCH@SBMA, and PCH@EG patches. Figure 7-1 Figure C shows the electrical impedance spectra of hydrogels IPNCH@CAPAM, PSCH@SBMA, and PCH@EG, as well as a commercial gel electrode (CGE), after complete swelling in PBS in the physiologically relevant frequency range of 1–100 Hz. The inset shows the spectrum for a larger frequency range (0.1–10 Hz). 4 Impedance spectrum within Hz. Figure 7-1 The figure shows the interfacial resistance of the skin surface with hydrogels IPNCH@CAPAM, PSCH@SBMA, and PCH@EG, as well as the commercial gel electrode (CGE). Figure 7-2 E in the figure shows a schematic diagram of the three-electrode setup used in cyclic voltammetry (CV) measurements, including the reference electrode RE (Ag / AgCl), the counter electrode CE (platinum Pt), the working electrode WE, and the electrolyte (0.1M Na2SO4). Figure 7-2 F in the figure shows the CV curve of IPNCH@CAPAM after 10,000 cycles (x-axis: electrode potential (V), y-axis: current (μA). -2 )). Figure 7-2 G in the figure shows the CV curve of PSCH@SBMA after 10,000 cycles. Figure 7-2 H in the figure shows the CV curve of PCH@EG after 10,000 cycles. Figure 7-2 The 'I' in the figure represents the quantification of charge storage capacity (CSC) over 10,000 cycles, illustrating the capacity retention of different hydrogels during long-term cycling.

[0186] The above tests show that:

[0187] The hydrogel IPNCH@CAPAM exhibits excellent electrical properties, with a conductivity of 1.99 S / m, which is superior to that of commercial gel electrodes (CGE). Figure 7-1 (A) This high conductivity is attributed to the interpenetrating network of the conductive polymer, which ensures a unified path for electron transport. To visually demonstrate the conductivity, Figure 7-1 Figure B shows an LED circuit powered by a 3V battery and connected to IPNCH@CAPAM. In terms of impedance, all three hydrogels exhibited significantly lower impedance compared to commercially available hydrogels, indicating improved conductivity and reduced resistive loss at biologically relevant frequencies. Figure 7-1 (C in the text). This is crucial for minimizing resistive losses and ensuring stable electrical signal transmission at the bioelectronic interface. IPNCH@CAPAM also exhibits lower interfacial resistance on human skin than commercial hydrogels, confirming IPNCH@CAPAM's superior ability to form low-resistance interfaces with biological tissues. Figure 7-1 (D in the text). The electrochemical properties of the hydrogel are also crucial for its bioelectronic stimulation capabilities for further applications. Therefore, the electrochemical performance of IPNCH@CAPAM was evaluated using cyclic voltammetry (CV) for 10,000 cycles. The results showed that its charge storage capacity (CSC) decreased very little, maintaining nearly 95% of the original CSC. Figure 7-2 (I in the text). This highlights the excellent electrochemical stability of the hydrogel, making it an ideal candidate material for long-term bioelectronic applications, especially devices requiring continuous electrical stimulation.

[0188] The PSCH@SBMA hydrogel exhibits the highest conductivity among the three hydrogels, reaching 5.25 S / m. Figure 7-1 Similar to IPNCH@CAPAM, PSCH@SBMA also successfully lit up the LED in the closed-circuit test. Figure 7-1 The B in the figure demonstrates its effective electron transport capability. Regarding impedance, PSCH@SBMA exhibits a significant reduction in the physiological frequency range, comparable to IPNCH@CAPAM. Figure 7-1 (C in the original text). Reduced impedance is crucial for improving signal transmission efficiency in bioelectronic applications. PSCH@SBMA also outperformed CGE in interface resistance testing, indicating its ability to form a stable, low-resistance interface with biological tissues. Figure 7-1 (D in the middle).

[0189] The electrochemical performance of PSCH@SBMA also demonstrated excellent stability through CV testing. After 10,000 cycles, it exhibited a high CSC value (11.49 μC cm⁻¹). -2 The decrease was only 16.6%, indicating its electrochemical durability for long-term use. Figure 7-2 G and Figure 7-2 (I) After 50 days of water immersion, the specific surface area of ​​the hydrogel increased from 5.735 m² / g. 2 / g increased significantly to 15.044m 2 / g. This significant increase indicates that the structure of the hydrogel network gradually expands and reorganizes over time. This expansion behavior further enhances ion mobility by forming more continuous conductive pathways, consistent with the observed increase in CSC after prolonged immersion. This stability ensures that PSCH@SBMA is well-suited for bioelectronic devices requiring continuous electrical stimulation, such as neural interfaces and cardiac pacemakers.

[0190] The electrical conductivity of the PCH@EG hydrogel reached 3.79 S / m, which, although lower than that of PSCH@SBMA, is still a significant improvement compared to CGE. Figure 7-1 A) in this context can be used in systems such as skin electrodes or sensors requiring a rapid electrical signal response. Figure 7-1 In section B, the hydrogel effectively powers the LED circuit, further confirming its conductivity. Regarding impedance, PCH@EG shows improvement over CGE, but not as significantly as the other two hydrogels. Figure 7-1 (C in the text). Its interface resistance is also lower than CGE, ensuring effective electrical contact with biological tissues. Figure 7-1 (D in the text). Furthermore, the CSC value of PCH@EG is as high as 48.23 μC·cm⁻¹. -2 Compared to a bare platinum electrode of the same size (3.76 μC·cm⁻¹), -2 12.8 times higher Figure 7-2 H in Figure 7-2(I) Notably, the charge storage capacity (CSC) of PCH@EG increases with increasing CV cycles, showing a significant increase of over 95% after 10,000 cycles. This improvement in electrochemical performance can be attributed to the structure of the pure PEDOT:PSS hydrogel, which facilitates full exposure of the conductive polymer in the electrolyte, allowing the PEDOT chains to rearrange more freely under the influence of the electric field during CV cycling. This molecular rearrangement makes the PEDOT chain segments more ordered and improves the continuity of the conductive channels, thus significantly increasing the CSC, even after long-term cycling. This behavior indicates that PCH@EG hydrogel maintains excellent electrochemical stability and sustained charge storage capacity under prolonged electrochemical activation, making it a promising material for flexible bioelectronic devices, including electrophysiological signal recording and electrical nerve stimulation (TENS).

[0191] These three hydrogels were immersed in different bodily fluids, such as simulated bodily fluids, simulated sweat, and simulated cerebrospinal fluid, and their electrochemical stability was tested and verified under 10,000 to 100,000 cycles. These electrochemical performance results demonstrate that these three conductive polymer hydrogels possess excellent conductivity, low interfacial resistance, and significant stability. The powerful electrochemical properties of these three hydrogels enable the fabrication of more efficient bioelectronic devices, particularly in applications at tissue-electron interfaces, where long-term stability and low interfacial resistance are crucial.

[0192] From the perspective of the mechanical and electrical properties of hydrogels constructed using the three methods, these three methods form a hierarchical performance optimization framework: First, IPNCH serves as the foundation, verifying that decoupling the conductive network and the structural network can significantly improve mechanical compliance; then, PSCH introduces ionic additives to improve the conductive pathway, demonstrating that controlled microphase separation can simultaneously enhance stretchability and conductivity; finally, PCH simplifies the system by eliminating the non-conductive matrix, proving that solvent-mediated crystallization alone can achieve skin-compatible properties. Most importantly, each strategy (e.g., the polar effect in PCH provides a basis for the selection of ionic dopants in PSCH) mutually reinforces each other, thereby promoting the goal of achieving a harmonious unity between mechanics and electronics.

[0193] 5. Epidermal electrophysiological signal recordings of IPNCH@CAPAM (Example 1), PSCH@SBMA (Example 8), and PCH@EG (Example 10).

[0194] Figure 8-1 Figure A shows electromyography (EMG, x-axis: time, y-axis: amplitude) of the lateral dumbbell lifting process recorded by electrodes of hydrogels IPNCH@CAPAM, PSCH@SBMA and PCH@EG and commercial gel electrodes (CGE). Figure 8-1B in the figure shows the electromyographic signals recorded during lateral dumbbell raises. Figure 8-1 C in the figure shows the amplitude of electromyographic signals recorded by different hydrogel electrodes and commercial gel electrodes during lateral dumbbell lifts. Figure 8-1 D in the figure shows the signal-to-noise ratio (SNR) recorded by different hydrogels and commercial gel electrodes during lateral dumbbell lifts. Figure 8-1 The E-cell records the electromyography (EMG) generated by raising the palm at different angles. Figure 8-1 F in the figure shows the electromyography (EMG) generated when the angle of the palm is increased. Figure 8-1 G in the figure shows the amplitude of electromyographic signals recorded by different hydrogels and commercial gel electrodes when the palm is raised. Figure 8-1 H in the figure shows that the signal-to-noise ratio recorded by different hydrogels and commercial gel electrodes increases with increasing angle when the palm is raised. Figure 8-2 The I in the figure shows the epidermal electrocardiogram signals recorded by different hydrogels and commercial gel electrodes. Figure 8-2 J in the diagram is a schematic diagram of recording the electrocardiogram signal of the skin in a resting state. Figure 8-2 K in the figure represents the signal-to-noise ratio recorded by different hydrogels and commercial gel electrodes. Figure 8-2 L in the figure represents the baseline variability of epidermal electrocardiograms (ECGs) recorded by different hydrogels and commercial gel electrodes. Figure 8-2 M in the image shows magnified views of characteristic signals from electrocardiograms (ECGs) recorded using different hydrogels and commercially available gel electrodes. Figure 8-3 N in the figure represents the epidermal electroencephalogram (EEG) signals recorded by different hydrogels and commercial gel electrodes. Figure 8-3 P in the figure represents the signal-to-noise ratio recorded by different hydrogels and commercial gel electrodes.

[0195] The tests above demonstrate that hydrogels IPNCH@CAPAM, PSCH@SBMA, and PCH@EG possess excellent mechanical and electrical properties and can be used as bioelectrodes for detecting epidermal electrophysiological signals, including electromyography (EMG), electrocardiography (ECG), and electroencephalography (EEG).

[0196] Hydrogel electrode pads were attached to the flexor muscles (right arm) of a volunteer. The volunteer then laterally lifted a 2 kg dumbbell and lowered their arm for 2 seconds at 3-second intervals. Figure 8-1 (A and B in the original text). During biceps contraction, all hydrogels were able to record significant electromyographic signals, with improved signal amplitude and signal-to-noise ratio (SNR) compared to commercial gel electrodes. EMG ) are higher ( Figure 8-1 (C and D in the text). Furthermore, electromyography (EMG) can successfully display the action potential signal characteristics corresponding to small-amplitude gestures. Figure 8-1(E). Hydrogel electrodes were formed by attaching them to the volunteer's flexor muscles (outer forearm) and elbow. As the palm was gradually raised upwards, the contraction of the flexor muscle fibers gradually increased. (e.g., ...) Figure 8-1 As shown in F, similar to commercial gel electrodes, the hydrogel electrode of this invention captures increasingly larger amplitude electromyographic signals at higher palm elevation angles (10°, 30°, 50°). Muscle contraction causes significant changes in potential intensity, highlighting the hydrogel electrode's remarkable sensitivity in distinguishing the different forces generated by different palm elevation heights. Furthermore, at any palm elevation angle, the average amplitude and signal-to-noise ratio collected by the three hydrogels exceed those of commercial hydrogel CGE (Current Emissions Per Term). Figure 8-1 (G and H in the diagram). At the highest angle of the raised hand, the signal-to-noise ratio of IPNCH@CAPAM was 16.4 dB, more than twice that of the commercial hydrogel CGE (ENG, 8.1 dB). This result is attributed to the inherent softness of the hydrogel, which creates a conformal contact between the hydrogel electrode and the tissue, while its excellent electrical properties further ensure a stable and high-quality signal. Therefore, the stable and effective recording of epidermal electromyography signals preliminarily demonstrates the application potential of the conductive polymer hydrogel of this invention in bioelectronics.

[0197] Further utilizing the IPNCH@CAPAM, PSCH@SBMA, and PCH@EG hydrogel electrodes of this invention, along with commercially available hydrogels, to detect epidermal electrocardiogram (ECG) signals. Figure 8-2 (I) Electrocardiogram (ECG) is an effective tool for monitoring changes in cardiac electrical activity and can reveal certain cardiac abnormalities, such as arrhythmias, insufficient coronary blood flow, and electrolyte imbalances. For example... Figure 8-2 As shown in Figure J, three hydrogel electrode pads were attached to a volunteer's left / right arm and right thigh to record an electrocardiogram (ECG). The hydrogel electrode of this invention successfully captured a stable ECG signal similar to that of commercially available gel electrodes, displaying typical ECG waveforms, including the P wave (atrial depolarization), QRS complex (ventricular depolarization), and T wave (ventricular repolarization) (as shown in Figure J). Figure 8-2 The comparison results of electrocardiogram signals show that the hydrogel electrode of the present invention has a higher signal-to-noise ratio, while PCH@EG has significant advantages (M). Figure 8-2 In addition, the baseline drift of the collected ECG signals was corrected using a sliding window averaging filter in MATLAB, and the baseline drift (K) was calculated. Figure 8-2 (L in the text). Clearly, the baseline drift of the three hydrogel electrodes is significantly less than that of commercially available hydrogel CGE. The baseline drift of the hydrogel IPNCH@CAPAM is only about 15% of that of CGE. Stable ECG recordings and clear waveform identification demonstrate the potential significance of the hydrogel electrodes of this invention in health monitoring and disease diagnosis, including myocarditis and arrhythmias.

[0198] Monitoring of epidermal EEG signals further demonstrates the practicality of the hydrogel bioelectrode of this invention. Different hydrogels were used as bioelectrodes on the forehead and both earlobes of volunteers to capture EEG signals in a resting state. In the resting state of the subjects, the EEG signals obtained from the three hydrogels of this invention and commercially available gel electrodes showed similar patterns. This indicates that the hydrogel electrode of this invention exhibits the best signal stability, without any artifacts. Figure 8-3 (N in the context). Furthermore, the signal-to-noise ratio (SNR) of the EEG signal... EEG This also indicates that the electroencephalogram (EEG) signals generated by hydrogel electrodes are of higher quality. Figure 8-3 The P) in the figure was characterized at approximately 10 Hz by fast Fourier transform, reordering the resting state within 30 seconds, revealing specific resting activity in the frequency range of 8 to 13 Hz. Furthermore, the IPNCH@CAPAM and PCH@EG hydrogel electrodes exhibited a higher proportion of the alpha band, indicating that this was a genuine signal rather than a spurious signal. These results demonstrate that the hydrogel electrode pads of the present invention can accurately capture high-quality epidermal electrophysiological signals. Compared to commercial gel electrodes, the hydrogel bioelectrodes of the present invention exhibit superior performance due to their softness, excellent conductivity, and robust fit, making them ideal materials for epidermal bioelectronics.

[0199] 6. Implantable electrophysiological signal recorders of IPNCH@CAPAM (Example 1) and PSCH@SBMA (Example 8)

[0200] Figure 9-1 Figure A shows a schematic diagram of a hydrogel used as a bioelectrode to connect the HD-X02 (a bioelectric telemetry implantation device for small animals manufactured by DSI) implant to biological tissue for implantable recording of electrocardiogram and electromyogram signals. Figure 9-1 B in the figure shows the cell viability of BHK21 cells cultured in hydrogel IPNCH@CAPAM and PSCH@SBMA extraction media for 24 hours and 72 hours, respectively. Data are presented as mean ± standard deviation (SD) (n=4 tests), with a threshold *p<0.05 (one-way ANOVA and Tukey test). The dashed line represents the 80% biocompatibility threshold. Figure 9-1 C in the figure represents representative H&E staining images of mouse tissues from the subcutaneous hydrogel implantation group (i.e., experimental group) and the non-hydrogel implantation group (i.e., control group, corresponding to Sham in the figure) within 2 weeks. There was no significant difference between the experimental group and the control group. The cells were intact and neatly arranged, and no obvious inflammatory cell aggregation was observed. Scale bar: 50 μm. Figure 9-1The DE in the image shows implanted ECG signals recorded over three consecutive weeks (Week 1, Week 2, Week 3) by different hydrogel bioelectrodes and a non-hydrogel control group (continuous signals within 30 seconds and amplified signals within 1 second, respectively). Figure 9-2 F in the figure represents the heart rate detected by different hydrogel bioelectrodes and the non-hydrogel control group. Figure 9-2 The GH in the image shows implanted electromyography signals recorded over three consecutive weeks by different hydrogel bioelectrodes and a non-hydrogel control group (continuous signals over 30 seconds and amplified signals over 1 second, respectively). Figure 9-2 I in the figure shows the electromyographic peak (Peak, mV) detected by different hydrogel bioelectrodes and the non-hydrogel control group.

[0201] In addition to wearable devices attached to the skin, the hydrogel of this invention also possesses excellent softness and good biocompatibility, making it suitable as a material for implantable bioelectronic devices. Biocompatibility is a key factor in implantable bioelectronic devices, ensuring accurate capture of physiological data and mitigating foreign body reactions during long-term use. PNCH@CAPAM and PSCH@SBMA, with their superior flexibility, were used in implantable bioelectronic devices, and both underwent in vitro cytotoxicity and biocompatibility testing prior to application. In vitro cytotoxicity assessment included measuring the cell viability of hamster Syrian kidney (BHK21) cells cultured in the extract and other human-derived cells, such as... Figure 9-1 As shown in Figure B, after 24 and 72 hours of culture, the cell viability remained above 80%, meeting the biocompatibility criteria. PSCH@SBMA showed better performance due to optimized elimination of residual monomers. The results indicate that the cytotoxicity of the two hydrogels of this invention to BHK21 cells is negligible. To further investigate the biocompatibility of the hydrogels, PNCH@CAPAM and PSCH@SBMA hydrogels were implanted subcutaneously in the dorsal region of mice for 2 weeks, and histological analysis was performed to detect the in vivo tissue immune response to the hydrogels. Figure 9-1The image C- shows a representative histological cross-section of mouse subcutaneous tissue stained with hematoxylin and eosin. No inflammatory response was observed in the tissue surrounding the hydrogel, and there was no significant difference in morphology between the test tissue and normal tissue. Furthermore, to ensure stable bonding between the hydrogel and biological tissue in vivo, we investigated the coefficient of friction of the hydrogel through repeated friction tests in a normal saline environment. The data showed that IPNCH@CAPAM and PSCH@SBMA maintained stable contact with tissue, which was also confirmed by the friction coefficient measurements. IPNCH@CAPAM exhibited strong mechanical properties and a stable coefficient of friction, making it particularly suitable for implantable electrode materials requiring high friction and wear resistance. It maintained structural integrity and performance stability even under changes in external mechanical stress and internal environmental conditions. In summary, the hydrogels IPNCH@CAPAM and PSCH@SBMA possess negligible cytotoxicity, good biocompatibility, and sufficient abrasion resistance, thus showing promise for use in implantable bioelectronics.

[0202] Evaluation of the efficacy of hydrogel bioelectrodes in vivo: Implantable electrocardiogram (ECG) recording showed significant advantages in detecting cardiac allogeneic transplant rejection, diagnosing sudden arrhythmias, and ECG imaging. In this test, the HD-X02 implant biopotential leads for acquiring electrical signals were embedded in a hydrogel precursor solution and co-molded with a hydrogel film in a mold to form a 1 cm × 1 cm × 1 mm thin-film bioelectrode. The ECG biopotential leads coated with the hydrogel electrodes were fixed to the periosteum of the rib in a lead-- configuration and sutured with non-absorbable polyester braided sutures to prevent dislodgement. The control group used leads directly as electrodes. The implants were activated weekly for 30 minutes each time to record ECG signals over three weeks. Figure 9-1 The F in the figure shows that the mouse heart rate detected by the hydrogel electrode and biopotential wire was approximately 750 bpm, and the data obtained by the hydrogel electrode was not significantly different from that of the control group. Figure 9-1 D in the figure represents the one-minute ECG signal in each test. Starting from the second week, the signal amplitude detected in the non-hydrogel electrode control group was significantly reduced. Conversely, in the hydrogel experimental group, instability / noise interference in the monitoring signal gradually appeared until the third week. Based on the amplified ECG signal within 1 second, PSCH@SBMA achieved continuous monitoring of ECG characteristic signals for three consecutive weeks (e.g., Figure 9-2 (As shown in E in the figure). IPNCH@CAPAM successfully monitored characteristic electrocardiogram signals in the first two weeks, while the non-hydrogel control group failed to obtain stable signals from the second week onwards. The results indicate that the hydrogel film bioelectrode of the present invention promotes stable contact between bioelectronic devices and tissues, unlike traditional bioelectrical leads which are difficult to establish a strong connection and may experience frictional slippage in the body.

[0203] Similarly, the non-hydrogel electrode control group began to show intermittent signals at week 3, while the IPNCH@CAPAM and PSCH@SBMA hydrogel membrane bioelectrodes maintained stable signal recordings throughout the three weeks (e.g., Figure 9-2 Comparing the G and H values ​​in the electromyography (EMG) peaks, there was no significant difference between the PSCH@SBMA group and the control group, indicating comparable signal intensity. However, the peak value of IPNCH@CAPAM was slightly lower. This difference may be due to the relatively low conductivity of IPNCH@CAPAM, which may lead to signal loss during transmission. Figure 9-2 (I in the middle).

[0204] The above results demonstrate that the hydrogel film of the present invention possesses synergistic mechanical and electrical properties, enabling the establishment of robust connections between bioelectronic devices and tissues, thus becoming an implantable bioelectrode. This ensures the long-term, stable acquisition of reliable and high-quality electrophysiological signals.

[0205] The tests described above demonstrate that the hydrogel IPNCH@CAPAM achieved an ultra-low modulus of 0.28 kPa, an elongation of 530%, and an electrical conductivity of 1.99 S / m; the hydrogel PSCH@SBMA balanced a modulus of 0.44 kPa, an elongation of 789%, and an electrical conductivity of 5.25 S / m; and the hydrogel PCH@EG, through a simple solvent annealing route, achieved skin-matched hardness (15 kPa) and an electrical conductivity of 3.79 S / m. The applications of these hydrogels include soft neural interfaces and high-fidelity epidermal sensors. In the tests, the hydrogel PCH@EG electrode recorded EMG / ECG / EEG with a signal-to-noise ratio as high as 20.0 dB, while the hydrogels IPNCH@CAPAM and PSCH@SBMA maintained stable implanted ECG / EMG readings for over three weeks. By combining multidimensional performance screening with systematic in vitro and in vivo validation, this research not only bridges the "performance islands" problem in the field of conductive hydrogels but also lays the blueprint for next-generation bioelectronic materials. Looking ahead, further incorporating these hydrogels into layered hybrid materials, integrating self-healing patterns, 3D-printed microelectrode arrays, and scaling up production will further expand the application of the hydrogels of this invention in personalized, wearable, and implantable devices.

[0206] The terms "first" and "second" used in this document are for descriptive purposes only and should not be construed as indicating or implying relative importance or implicitly specifying the number of technical features indicated. Therefore, a feature marked "first" or "second" may explicitly or implicitly include one or more of that feature. In the description of this application, "multiple" means two or more, unless otherwise explicitly specified.

[0207] In the description of this specification, the references to terms such as "one embodiment," "some embodiments," "example," "specific example," or "some examples," etc., refer to specific features, structures, materials, or characteristics described in connection with that embodiment or example, which are included in at least one embodiment or example of the present invention. In this specification, the illustrative expressions of the above terms do not necessarily refer to the same embodiment or example. Furthermore, the specific features, structures, materials, or characteristics described may be combined in any suitable manner in one or more embodiments or examples. Moreover, without contradiction, those skilled in the art can combine and integrate the different embodiments or examples described in this specification, as well as the features of different embodiments or examples.

[0208] Although embodiments of the present invention have been shown and described above, it is understood that the above embodiments are exemplary and should not be construed as limiting the present invention. Those skilled in the art can make changes, modifications, substitutions and variations to the above embodiments within the scope of the present invention.

Claims

1. A conductive polymer hydrogel, characterized in that, Includes poly(3,4-ethylenedioxythiophene): polystyrene sulfonate and water, Alternatively, it may comprise the poly(3,4-ethylenedioxythiophene): polystyrene sulfonate, additives, and water, wherein the additives include one or more of non-conductive polymers and ionic compounds. The added component and / or the poly(3,4-ethylenedioxythiophene):polystyrene sulfonate has a network structure.

2. The conductive polymer hydrogel according to claim 1, characterized in that, The added ingredients include the non-conductive polymer, which includes one or more of polyacrylamide, polyacrylamide-carrageenan, polyacrylic acid, and polyvinyl alcohol.

3. The conductive polymer hydrogel according to claim 2, characterized in that, Based on the total mass of the conductive polymer hydrogel, the conductive polymer hydrogel comprises, by mass percentage, 0.03 wt.% to 0.1 wt.% of the poly(3,4-ethylenedioxythiophene):polystyrene sulfonate and 12 wt.% to 20 wt.% of the non-conductive polymer.

4. The conductive polymer hydrogel according to claim 1, characterized in that, The added ingredients include the ionic compounds, which include one or more of the following: 1-ethyl-3-methylimidazolium ethyl sulfate; 4-(3-butyl-1-imidazolium)-1-butane sulfonate triester; 1-butyl-3-methylimidazolium tetrafluoroborate; [2-(methacryloyloxy)ethyl]dimethyl-(3-sulfopropyl)ammonium hydroxide; and 3-[dimethyl-[3-(2-methylprop-2-enoylamino)propyl]ammonium]propane-1-sulfonate.

5. The conductive polymer hydrogel according to claim 4, characterized in that, Based on the total mass of the conductive polymer hydrogel, the conductive polymer hydrogel comprises, by mass percentage, 0.8 wt.% to 1.60 wt.% of the poly(3,4-ethylenedioxythiophene):polystyrene sulfonate and 20 wt.% to 30 wt.% of the ionic compound.

6. The conductive polymer hydrogel according to claim 1, characterized in that, The conductive polymer hydrogel comprises poly(3,4-ethylenedioxythiophene):polystyrene sulfonate and water, and the conductive polymer hydrogel comprises, by mass percentage, 0.8 wt.% to 2 wt.% of the poly(3,4-ethylenedioxythiophene):polystyrene sulfonate.

7. The conductive polymer hydrogel according to any one of claims 1 to 6, characterized in that, The conductive polymer hydrogel has an elastic modulus of less than 130 kPa and / or an electrical conductivity of 0.15 to 6 S / m.

8. A method for preparing a conductive polymer hydrogel according to any one of claims 1 to 7, characterized in that, Includes one of the following methods: Method 1: Add dopamine hydrochloride or polyvinyl alcohol to the alkaline solution to obtain the first mixed solution; Add freeze-dried poly(3,4-ethylenedioxythiophene):polystyrene sulfonate to the first mixed solution to obtain a second mixed solution; Add the polymerizing monomer, crosslinking agent and initiator to the second mixed solution, and after the reaction, obtain the first reaction solution; The first reaction solution self-cures to obtain the conductive polymer hydrogel. Method 2: An ionic compound, the crosslinking agent, and an initiator are added to an aqueous solution of poly(3,4-ethylenedioxythiophene):polystyrene sulfonate, and the mixture is stirred to obtain a third mixed solution. The third mixed solution is polymerized at a predetermined temperature for a predetermined time to obtain the conductive polymer hydrogel. Method 3: A polar solvent was added to the aqueous solution of poly(3,4-ethylenedioxythiophene):polystyrene sulfonate and stirred to obtain a fourth mixed solution; The fourth mixed solution was subjected to drying, annealing and hydration treatments in sequence to obtain the conductive polymer hydrogel.

9. The method according to claim 8, characterized in that, In method one, at least one of the following conditions must be met: The concentration of the alkaline solution is 0.2–0.8 g / mL; The mass ratio of dopamine hydrochloride to the volume of the alkaline solution is 1 mg: (0.01-0.05) ml, or the mass ratio of polyvinyl alcohol to the volume of the alkaline solution is 1 mg: (0.0001-0.0005) ml; The mass ratio of dopamine hydrochloride to poly(3,4-ethylenedioxythiophene):polystyrene sulfonate is (1.5-5):1, or the mass ratio of polyvinyl alcohol to poly(3,4-ethylenedioxythiophene):polystyrene sulfonate is (200-300):

1. The mass ratio of the polymeric monomer to the poly(3,4-ethylenedioxythiophene):polystyrene sulfonate is (200-300):1, or (150-230):1; The crosslinking agent is added at a mass fraction of 0.03 wt.% to 0.09 wt.%. The initiator is added at a mass fraction of 0.1 wt.% to 0.2 wt.%.

10. The method according to claim 8 or 9, characterized in that, In Method 1, the alkaline solution also includes carrageenan, and the mass ratio of the carrageenan to the mass of the poly(3,4-ethylenedioxythiophene):polystyrene sulfonate is (12-20):

1.

11. The method according to claim 8, characterized in that, In method two, at least one of the following conditions must be met: The mass ratio of the ionic compound to the poly(3,4-ethylenedioxythiophene):polystyrene sulfonate is (15-25):1; The crosslinking agent is added at a mass fraction of 0.03 wt.% to 0.09 wt.%. The initiator is added at a mass fraction of 0.1 wt.% to 0.2 wt.%. The predetermined temperature is 75–110°C; The predetermined time is 30 minutes to 60 minutes. Alternatively, in method three, at least one of the following conditions must be met: In the fourth mixed solution, the volume fraction of the polar solvent is 20 vol% to 30 vol%. In the aqueous solution of poly(3,4-ethylenedioxythiophene):polystyrene sulfonate, the concentration of poly(3,4-ethylenedioxythiophene):polystyrene sulfonate is (8-16) mg / ml; The annealing process includes heating the fourth mixed solution, which has undergone the drying process, to the annealing temperature, and then cooling it naturally at room temperature, repeating the annealing process 2 to 5 times. The annealing temperature is 80–200°C; The natural cooling time is 10–120 minutes.

12. The application of the conductive polymer hydrogel according to any one of claims 1 to 7 or the conductive polymer hydrogel prepared by any one of claims 8 to 11 in biological tissues.

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