A hydrogel electrode for long-term bioelectric signal monitoring and a preparation method and application thereof

Abstract

This invention belongs to the field of biomedical sensing and flexible electronics technology, and proposes a hydrogel electrode for long-term bioelectrical signal monitoring, its preparation method, and its application. The hydrogel electrode of this invention comprises a porous hydrogel and a network conductive layer embedded within the hydrogel, forming a three-dimensional mechanically interlocked structure. The average pore size of the hydrogel is 100 μm-500 μm, and its porosity is 60%-90%. The raw materials for preparing the hydrogel include the following components in parts by weight: 5-30 parts of a thermosensitive polymer, 0.5-15 parts of a conductive component, 5-40 parts of a water-retaining agent, 0.1-10 parts of an adhesive, and 0.01-2 parts of a crosslinking agent. Through multi-scale structural design and the synergistic effect of multiple components, the hydrogel electrode of this invention constructs a structure integrating a through-pore framework and a three-dimensional interlocked conductive network, achieving excellent comprehensive performance in terms of breathability, moisture retention, interfacial mechanics, and bioelectrical signal acquisition.

Description

Technical Field

[0001] This invention relates to the field of biomedical sensing and flexible electronics technology, specifically to a hydrogel electrode for long-term bioelectrical signal monitoring, its preparation method, and its application. Background Technology

[0002] Monitoring technologies for bioelectrical signals from the body surface (such as ECG, EMG, EEG, and EOG) are crucial foundational technologies for clinical diagnosis, health management, sports rehabilitation, and brain-computer interfaces. Their monitoring accuracy and long-term stability directly determine the reliability and effectiveness of related technological applications. However, these bioelectrical signals are extremely weak, typically ranging from microvolts to millivolts, and are highly sensitive to the interfacial contact state between the electrode and the skin. Even minute fluctuations in interfacial impedance and changes in contact stability can significantly affect signal acquisition quality and even lead to monitoring failure.

[0003] In long-term, continuous daily wearable monitoring scenarios, such as wearing for several hours or even across day and night, the bioelectrical signal monitoring electrodes on the skin surface must meet comprehensive performance requirements across multiple dimensions: First, the electrode-skin interface must have low and long-term stable interfacial impedance to ensure effective conduction of physiological electrical signals; second, the electrode must have good adhesion stability, able to fit tightly against the skin surface, avoiding poor contact due to loosening; third, under physiological conditions such as sweating and sebum secretion, the electrode must have good wearing comfort and breathability to reduce skin irritation and discomfort; fourth, it must effectively suppress interfacial micro-motion noise caused by daily activities to avoid motion artifacts interfering with signal acquisition. Furthermore, in monitoring areas covered by hair, such as the scalp, hair significantly reduces the effective contact area between the electrode and the skin, significantly increasing interfacial impedance. Simultaneously, the minute displacements and shear forces generated during wear can easily cause fluctuations in interfacial impedance, leading to severe motion artifacts and seriously affecting the stability and accuracy of the monitoring signal. Therefore, how to balance hair compatibility, breathability and moisture dissipation, and anti-motion artifact performance has become a core technical challenge that needs to be tackled collaboratively for wearable bioelectrical signal electrode systems.

[0004] Currently, hydrogel electrodes have become one of the mainstream types of electrodes for monitoring bioelectrical signals on the body surface due to their excellent biocompatibility, interfacial adhesion, and ionic conductivity. Their ability to retain moisture during long-term wear directly determines the electrode's lifespan and monitoring stability. Affected by factors such as surface temperature, sweat evaporation, and external environmental temperature and humidity, hydrogel materials are prone to water loss and shrinkage, leading to decreased ionic conductivity, a significant increase in the impedance at the electrode-skin interface, and discomfort such as skin tightness and itching. This severely restricts their application in long-term continuous monitoring scenarios. While some studies have attempted to improve water retention by adding polyols and polymeric moisturizers to the hydrogel system, in practical applications, maintaining both good breathability and moisture dissipation to meet long-term wear requirements while simultaneously ensuring long-lasting moisturizing performance and stable conductivity remains a challenging technical balance. Therefore, constructing an electrode structure that achieves synergistic optimization of breathability, water retention, and conductivity is a key breakthrough for improving the stability of long-term continuous monitoring of bioelectrical signals on the body surface.

[0005] To improve the skin adhesion and breathability of electrodes, existing technologies have developed a variety of technical routes for wearable bioelectrical signal electrodes, mainly including: (1) Flexible hydrogel electrodes: Electrodes are prepared by bonding and solidifying a composite conductive layer with a hydrogel interface layer. By optimizing the hydrogel formulation and the structure of the composite conductive layer, the skin adhesion, dehydration resistance and wearing comfort of the electrodes are taken into account to meet the attachment requirements in dynamic monitoring scenarios; (2) Electrode patches with open-pore structures: By designing a flexible electrode substrate with an open-pore structure (such as honeycomb openings), the open-pore substrate is bonded to the interface conductive hydrogel. The open-pore structure promotes sweat excretion and improves the breathability and moisture dissipation performance of the electrodes. At the same time, the conductive hydrogel reduces the contact impedance between the electrodes and the skin and improves the signal transmission effect; (3) Preparation of porous gel materials: By introducing air bubbles into the gel prepolymer solution and performing a rapid solidification and shaping process, a gel material with a porous structure is prepared. The focus is on solving the problems of pore collapse and contamination of the porous structure. The aim is to obtain a porous gel with controllable pore size and stable structure, providing material support for improving electrode performance.

[0006] In general, existing technologies are mainly improved and optimized in the following aspects: First, by designing the bonding between the composite conductive layer and the hydrogel interface layer, the flexibility and conductivity stability of the electrode are enhanced; second, by designing an open-pore or honeycomb substrate structure, the moisture dissipation and breathability of the electrode are improved, and the interface contact state is optimized by using conductive hydrogel; third, by introducing air bubbles and using a rapid curing and shaping process, the controllable preparation of porous gel materials is achieved, thereby improving the stability of the porous structure.

[0007] However, despite some progress in improving individual electrode performance indicators, existing technologies still have many shortcomings and limitations for challenging applications such as long-term stable monitoring and dynamic anti-motion artifacts in hair-covered scalp environments. Specifically: First, regarding breathable and moisture-wicking structural design, existing technologies primarily focus on optimizing the pore structure of the electrode substrate. However, the hydrogel layer, the key interface layer for direct contact between the electrode and skin, often lacks a three-dimensional network of interconnected pores throughout its thickness. In scenarios with dense hair, high sweating, and prolonged wear, sweat tends to accumulate and soak at the hydrogel-skin interface, failing to evaporate quickly. This leads to interface impedance drift, making it difficult to balance breathability and moisture wicking with stable conductivity, severely impacting the stability of long-term monitoring. Second, regarding the structural stability between electrode layers, existing electrodes often use simple physical bonding or conventional encapsulation methods to bond the conductive core (such as metal conductive layers or composite conductive layers) to the hydrogel interface layer, lacking a deep three-dimensional mechanical interlocking structure. During long-term wear, the dynamic shear loads generated by daily human activities can easily cause slight relative slippage or local deadhesion between the conductive substrate and the hydrogel interface layer. This leads to significant fluctuations in the contact impedance between the electrode and the skin, increasing motion artifacts and severely affecting the stability and accuracy of long-term monitoring of bioelectrical signals on the body surface. Furthermore, in terms of combining the preparation of porous gel materials with electrode applications, existing technologies mainly focus on the controllability of the pore-forming process of the porous gel material itself and the anti-collapse design of the pore structure, failing to systematically integrate the preparation of porous gel materials with the core performance required for bioelectrical signal electrodes on the body surface. Especially for monitoring needs in hair-covered areas, existing technologies lack specific designs for engineering applications such as how to effectively shape and tightly adhere electrodes between hair gaps, how to maintain long-term electrode adhesion stability in sebum and sweat environments, and how to achieve a stable integration of porous gels and conductive frameworks. This makes it difficult to fully leverage the advantages of porous gel materials. Finally, in terms of dynamic motion artifact resistance, existing technological improvements mainly focus on the electrode adhesion performance and interfacial impedance stability under static or quasi-static conditions. For low-frequency mechanical disturbances related to step frequency and its harmonics generated by dynamic activities such as walking, running, and nodding on the hair-covered scalp, there is a lack of inherent passive anti-artifact mechanisms from the perspective of electrode interface structural mechanics design. This makes it difficult to meet the actual needs of long-term continuous monitoring, especially for surface bioelectrical signals such as EEG signals that are extremely sensitive to micro-motions, in real dynamic monitoring scenarios.

[0008] In summary, while existing technologies for wearable bioelectrical signal electrodes have made some attempts to improve individual properties such as electrode adhesion, breathability and moisture dissipation, and porous structure construction, they have failed to provide electrode materials, structures, and preparation methods that can systematically meet the long-term stable monitoring requirements of hair-covered areas. They cannot achieve synergistic optimization of core properties such as breathability and moisture dissipation, long-term adhesion, stable conductivity, and resistance to motion artifacts, making it difficult to support the reliability requirements of long-term continuous monitoring of bioelectrical signals on the body surface. Summary of the Invention

[0009] The purpose of this invention is to overcome the shortcomings of the prior art and provide a hydrogel electrode for long-term bioelectric signal monitoring, its preparation method, and its application.

[0010] To achieve the above objectives, the technical solution adopted by the present invention is as follows: In a first aspect, the present invention provides a hydrogel electrode for long-term bioelectric signal monitoring, comprising a hydrogel having a porous structure and a mesh conductive layer embedded inside the hydrogel and forming a three-dimensional mechanically interlocked structure with the hydrogel. The hydrogel has an average pore size of 100 μm-500 μm and a porosity of 60%-90%. The raw materials for preparing the hydrogel include the following components in parts by weight: 5-30 parts of thermosensitive polymer, 0.5-15 parts of conductive component, 5-40 parts of water-retaining agent, 0.1-10 parts of adhesive and 0.01-2 parts of crosslinking agent.

[0011] The hydrogel electrode of this invention is based on an integrated molding strategy of "phase transformation stabilizing pores - gelation window embedding - simultaneous curing and shaping," constructing an integrated electrode structure with a three-dimensional interlocking porous hydrogel and a mesh conductive layer. This structure, on the one hand, provides a low-resistance transmission path for water vapor diffusion and sweat conduction through the interconnected channels, achieving synergistic functions of breathability and water retention; on the other hand, it significantly improves the positioning stability of the conductive framework through the interlocking interface, physically enhancing the interfacial bonding strength between the conductive phase and the matrix, and endowing the overall material with stress dissipation and micro-motion suppression capabilities, thereby… This invention suppresses relative slippage during dynamic wearing at the mechanical source. Simultaneously, it constructs a synergistic system combining interfacial wettability, ion conduction pathways, and long-term water retention capacity through a multi-component compound of thermosensitive polymers, conductive components, water-retaining agents, adhesives, and crosslinking agents. This allows the hydrogel to maintain strong interfacial bonding, persistent water content, and stable ion conductivity even under complex body surface environments such as wet conditions, oiliness, sweat, and dynamic deformation. It enables low-noise, low-drift electrophysiological signal acquisition under complex conditions such as hair coverage and dynamic deformation, laying the material and structural foundation for long-term continuous physiological electrophysiological signal acquisition.

[0012] As a preferred embodiment of the hydrogel electrode for long-term bioelectric signal monitoring described in this invention, the hydrogel has an average pore size of 365 μm-405 μm and a porosity of 72%-81%.

[0013] Preferably, the hydrogel has an average pore size of 405 μm and a porosity of 81%.

[0014] As a preferred embodiment of the hydrogel electrode for long-term bioelectric signal monitoring according to the present invention, the raw materials for preparing the hydrogel include the following components in parts by weight: 6-28 parts of thermosensitive polymer, 1-7 parts of conductive component, 12-35 parts of water-retaining agent, 0.5-6 parts of adhesive and 0.8-1.5 parts of crosslinking agent; and / or, the raw materials for preparing the hydrogel also include 10-60 parts of solvent; the solvent includes water.

[0015] Preferably, the raw materials for preparing the hydrogel include the following components in parts by weight: 24 parts of thermosensitive polymer, 2 parts of conductive component, 30 parts of water-retaining agent, 4 parts of adhesive, 1 part of crosslinking agent and 39 parts of solvent.

[0016] As a preferred embodiment of the hydrogel electrode for long-term bioelectric signal monitoring according to the present invention, the thermosensitive polymer includes at least one of gelatin, agar, agarose, κ-carrageenan, poly-N-isopropylacrylamide, and polyoxyethylene-polyoxypropylene block copolymer; and / or, the conductive component includes at least one of conductive inorganic salt, proton-conductive material, metal or carbon-based conductive filler; and / or, the water-retaining agent includes polyol; and / or, the adhesive includes dopamine hydrochloride and / or gallic acid; and / or, the crosslinking agent includes tannic acid and / or citric acid.

[0017] Preferably, the temperature-sensitive polymer includes gelatin.

[0018] Preferably, the conductive component comprises a conductive inorganic salt.

[0019] Preferably, the water-retaining agent includes glycerin and / or sorbitol.

[0020] More preferably, the water-retaining agent includes glycerin.

[0021] Preferably, the adhesive comprises dopamine hydrochloride.

[0022] Preferably, the crosslinking agent includes tannic acid.

[0023] Preferably, the conductive inorganic salt includes at least one of potassium chloride, sodium chloride, lithium chloride, and sodium sulfate; and / or, the proton-conducting material includes at least one of phytic acid, sulfonic acid polymer, and choline-based ionic liquid; and / or, the metal or carbon-based conductive filler includes at least one of silver nanowires, conductive carbon black, graphene, and reduced graphene oxide.

[0024] More preferably, the conductive inorganic salt includes sodium chloride.

[0025] More preferably, the proton-conducting material includes phytic acid.

[0026] As a preferred embodiment of the hydrogel electrode for long-term bioelectric signal monitoring described in this invention, the mesh conductive layer is a silver / silver chloride mesh conductive layer with a pore size greater than 0.02 mm; the opening structure of the mesh conductive layer includes at least one of the following: grid structure, honeycomb structure, serpentine structure, circular pore array structure, leaf vein structure, slit / narrow slit structure, herringbone structure, fishbone structure, expanded metal mesh structure, triangular array structure, square array structure, rhombic array structure, Voronoi cell network structure, random pore array structure, quasi-periodic pore array structure, pore size gradient structure, and pore density gradient structure.

[0027] Preferably, the mesh size of the conductive mesh layer is 0.05 mm-0.5 mm, and the thickness is 0.03 mm-1 mm.

[0028] Preferably, the mesh conductive layer has a mesh aperture of 0.2 mm and a thickness of 0.2 mm.

[0029] Secondly, the present invention provides a method for preparing the hydrogel electrode for long-term bioelectrical signal monitoring, comprising the following steps: S1. Mix the thermosensitive polymer, conductive component, water-retaining agent, adhesive, crosslinking agent and solvent to obtain hydrogel precursor solution; S2. The hydrogel precursor liquid is sheared and foamed to form a foaming system; S3. The foaming system is coated, injected or printed onto the target site, and then a mesh conductive layer is embedded therein. The foaming system is solidified through phase transformation to form a three-dimensional mechanical interlocking structure, thereby obtaining the hydrogel electrode for long-term bioelectric signal monitoring.

[0030] The preparation method of the hydrogel electrode of this invention is based on an integrated molding process, and achieves the construction of a triple structure of "pore locking - framework interlocking - functional integration" through multi-step synergy. First, this invention forms a homogeneous precursor system with thermosensitive phase change characteristics, ionic conductivity, interfacial adhesion and long-lasting water retention by compounding a thermosensitive polymer, conductive component, water-retaining agent, adhesive and crosslinking agent. This achieves pre-dispersion and functional coupling of multiple components at the molecular level, providing a homogeneous and rheologically controllable reaction precursor for subsequent foaming, embedding and curing processes. Subsequently, gas is introduced into the precursor liquid under high-speed shear conditions, and the shear force is used to disperse the bubbles into uniformly sized and densely distributed micron-sized pores, forming a metastable foaming system. This allows the bubbles to remain stably and interconnect, providing a foundation for subsequent processes. The construction of the continuous pore network lays the geometric framework. Subsequently, during the gelation window period before the foaming system is fully cured, the mesh conductive layer is embedded by immersion, pressing, or laying. The foaming system fully wets the pores and covers the mesh fibers with the help of low viscosity fluidity. Then, the temperature-induced phase transformation triggers the sol-gel transition of the thermosensitive polymer, so that the system changes from a fluid dynamic to a solid state. The bubble network is locked in situ into a continuous macroporous structure, realizing highly controllable shaping and batch reproducibility of the pore structure. At the same time, the conductive layer is anchored in the hydrogel matrix as a whole, forming a three-dimensional mechanical interlocking structure that runs through the whole and has a strong interface bond.

[0031] In a preferred embodiment of the preparation method of the present invention, the mixing temperature in step S1 is 45℃-95℃.

[0032] Preferably, in step S1, the mixing temperature is 62 ℃-85 ℃.

[0033] In a preferred embodiment of the preparation method of the present invention, in step S2, the rotation speed of the shear foaming is 300 rpm-5000 rpm, and the time is 0.5 min-20 min.

[0034] In the high-speed shear foaming process of this invention, the local negative pressure and liquid surface disturbance generated by high-speed shearing can naturally entrain ambient air into the precursor liquid system. Under otherwise identical conditions, moderately enhancing shear-induced air entrainment can help improve porosity; however, excessive shearing or over-foaming can lead to bubble rupture or collapse, resulting in increased pore size or wider pore size distribution. Preferably, in step S2, the rotational speed of the shear foaming is 400 rpm-4900 rpm, and the time is 1 min-10 min.

[0035] More preferably, in step S2, the rotation speed of the shear foaming is 2000 rpm and the time is 5 min.

[0036] In a preferred embodiment of the preparation method of the present invention, in step S3, the phase transformation is carried out by temperature induction, so that the foaming system completes the sol-gel transformation; the temperature-induced cooling rate is 3 ℃ / min-20 ℃ / min.

[0037] Preferably, the temperature-induced cooling rate is 7.5 ℃ / min-20 ℃ / min.

[0038] Thirdly, the present invention provides the application of the hydrogel electrode for long-term bioelectrical signal monitoring in electroencephalogram (EEG) electrodes, electrocardiogram (ECG) electrodes, electromyogram (EMG) electrodes, electrooculogram (EOG) electrodes, or skin-touch biosensors that are worn continuously by the human body in daily life.

[0039] The hydrogel electrode of this invention can perform continuous electrophysiological signal monitoring for up to 8 days, achieving a systematic improvement in performance in terms of breathability, moisture retention, interfacial mechanics, and electrophysiological signal acquisition.

[0040] Compared with existing technologies, the beneficial effects of this invention are as follows: The hydrogel electrode of this invention solves the common problems of easy collapse and uneven pores in traditional foamed gels through an integrated keyhole molding process. The resulting interconnected macroporous skeleton can remain stable during curing and long-term wear, and can still maintain high water content and ionic conductivity under open air and body temperature conditions, significantly suppressing impedance drift caused by interfacial water loss, and achieving a synergistic unity of breathability and water retention. Secondly, the hydrogel electrode of this invention can be infiltrated and molded in complex morphologies such as hair-covered scalp, and can achieve close adhesion with the scalp microstructure to form a continuously attached and strongly bonded electrode-skin interface. Furthermore, its three-dimensional interconnected macroporous hydrogel and mesh conductive layer can further synergistically construct a continuous sweat-wicking and moisture-dissipating channel, effectively alleviating interfacial immersion and performance degradation caused by sweat retention, and ensuring low amplitude fluctuations in interfacial impedance and stable signal quality during long-term monitoring. Furthermore, the hydrogel electrode of this invention provides mechanical buffering and dissipation pathways through viscoelastic porous hydrogel, and combines the intrinsic constraints of interface micro-movements on the three-dimensional interlocking conductive framework to attenuate low-frequency disturbance artifacts in dynamic scenarios such as walking, running, and head rotation from the mechanical source, significantly improving the signal-to-noise ratio and data stability under motion conditions, and greatly extending the effective monitoring time, providing an engineering solution for the acquisition of physiological electrical signals in dynamic, long-term, and real-world scenarios. Detailed Implementation

[0041] To better illustrate the objectives, technical solutions, and advantages of this invention, the invention will be further described below with reference to specific embodiments. Those skilled in the art should understand that the specific embodiments described herein are merely illustrative of the invention and are not intended to limit the invention.

[0042] The following description, in conjunction with specific embodiments, illustrates the practical effects of the present invention.

[0043] Unless otherwise specified, the experimental methods used in the examples are conventional methods; the materials, reagents, equipment, etc. used are all commercially available unless otherwise specified.

[0044] The preparation method of the mesh Ag / AgCl conductive layer used in the following examples and comparative examples includes the following steps: A 0.2 mm silver sheet is cut into a mesh structure using a laser cutter, and then chlorinated until the surface is a uniform light gray, forming an AgCl layer on its surface, resulting in a mesh Ag / AgCl conductive layer with a pore size of 0.2 mm and a thickness of 0.2 mm.

[0045] Example 1: This embodiment provides a breathable, water-retaining, hair-compatible, and highly adhesive hydrogel electrode for long-term bioelectric signal monitoring. Its preparation method includes the following steps: (1) Preparation of precursor solution In a constant temperature water bath at 75 ℃, the raw materials were mixed in the following weight proportions and stirred continuously until completely dissolved to obtain a homogeneous and clear hydrogel precursor solution: Thermosensitive polymer: 24 parts gelatin; conductive inorganic salt: 2 parts sodium chloride (NaCl); water-retaining agent: 30 parts glycerin; adhesive: 4 parts dopamine hydrochloride; crosslinking agent: 1 part tannic acid; solvent: 39 parts deionized water.

[0046] (2) High-speed shear foaming The above hydrogel precursor solution was transferred to a high-speed emulsifier and sheared and foamed at 2000 rpm for 5 min.

[0047] (3) Gelatinized window embedding The foaming system of step (2) is coated, injected or printed on the target area. Then, within the gelation window period of 1-3 minutes after the foaming is completed, a mesh Ag / AgCl conductive layer with a pore size of 0.2 mm and a thickness of 0.2 mm is quickly pressed into the middle of the foaming layer and gently rolled to make the precursor system penetrate and wet the mesh and cover the mesh wire.

[0048] (4) Phase transformation and solidification Within 2 minutes, the system of step (3) is rapidly cooled from 40 °C to 25 °C. By utilizing the thermosensitive phase transition properties of gelatin, the foam structure is locked in situ, forming a macroporous hydrogel with interconnected channels. At the same time, the gel network and the mesh Ag / AgCl conductive layer are tightly combined and interlocked in three-dimensional space to form an integrated three-dimensional mechanical interlocking structure, thus obtaining the target electrode.

[0049] Example 2: This embodiment provides a breathable, water-retaining, hair-compatible hydrogel electrode for long-term stable monitoring of physiological electrical signals. Its preparation method includes the following steps: (1) Preparation of precursor solution In a constant temperature water bath at 75 ℃, the raw materials were mixed in the following weight proportions and stirred continuously until completely dissolved to obtain a homogeneous and clear hydrogel precursor solution: Thermosensitive polymer: 6 parts gelatin; conductive inorganic salt: 1 part sodium chloride (NaCl); water-retaining agent: 12 parts glycerin; adhesive: 0.5 parts dopamine hydrochloride; crosslinking agent: 0.8 parts tannic acid; solvent: 39 parts deionized water.

[0050] (2) High-speed shear foaming The above hydrogel precursor solution was transferred to a high-speed emulsifier and sheared and foamed at 2000 rpm for 5 min.

[0051] (3) Gelatinized window embedding The foaming system of step (2) is coated, injected or printed on the target area. Then, within the gelation window period of 1-3 minutes after the foaming is completed, a mesh Ag / AgCl conductive layer with a pore size of 0.2 mm and a thickness of 0.2 mm is quickly applied and gently rolled to allow the precursor system to penetrate and wet the mesh and cover the mesh wire.

[0052] (4) Phase transformation and solidification Within 5 minutes, the system of step (3) is rapidly cooled from 40 °C to 25 °C. The thermosensitive phase transition properties of gelatin are used to lock the foam structure in situ, forming a macroporous hydrogel with interconnected channels. At the same time, the gel network and the mesh Ag / AgCl conductive layer are tightly combined and interlocked in three-dimensional space to form an integrated three-dimensional mechanical interlocking structure, thus obtaining the target electrode.

[0053] Example 3: The only difference between the preparation method of the hydrogel electrode in this embodiment and that in Example 1 is: In step (1), the raw materials and proportions of the precursor solution are different, as follows: Thermosensitive polymer: 28 parts gelatin; conductive inorganic salt: 5 parts sodium chloride (NaCl); proton conductive agent: 2 parts phytic acid; water-retaining agent: 35 parts glycerol; adhesive: 6 parts dopamine hydrochloride; crosslinking agent: 1.5 parts tannic acid; solvent: 22.5 parts deionized water; The remaining steps are the same as in Example 1.

[0054] Example 4: The only difference between the preparation method of the hydrogel electrode in this embodiment and that in Example 1 is: In step (1), the raw materials and proportions of the precursor solution are different, as follows: Thermosensitive polymer: 15 parts agarose; conductive inorganic salt: 2 parts sodium chloride (NaCl); proton conductive agent: 1 part phytic acid; water-retaining agent: 25 parts glycerol; adhesive: 1 part gallic acid; crosslinking agent: 1 part tannic acid; solvent: 55 parts deionized water; The remaining steps are the same as in Example 1.

[0055] Example 5: The only difference between the preparation method of the hydrogel electrode in this embodiment and that in Example 1 is: In step (2), the shearing speed is 4900 rpm; The remaining steps are the same as in Example 1.

[0056] Example 6: The only difference between the preparation method of the hydrogel electrode in this embodiment and that in Example 1 is: In step (2), the shearing speed is 400 rpm; The remaining steps are the same as in Example 1.

[0057] Comparative Example 1: The only difference between the preparation method of this comparative hydrogel electrode and that of Example 1 is: In step (2), high-speed shear foaming is not performed; the hydrogel precursor liquid is simply stirred to remove large air bubbles and ensure uniform mixing. The remaining steps are the same as in Example 1.

[0058] Comparative Example 2: The only difference between the preparation method of this comparative hydrogel electrode and that of Example 1 is: In step (1), the mass fraction of gelatin in the precursor liquid is 4 parts; In step (4), the system from step (3) is slowly cooled from 40 °C to 25 °C, which takes more than 1 hour; The remaining steps are the same as in Example 1.

[0059] Comparative Example 3: The only difference between the preparation method of this comparative hydrogel electrode and that of Example 1 is: In step (1), gelatin is replaced with an equal amount of chitosan in the raw materials of the precursor solution, and deionized water is replaced with an aqueous acetic acid solution (the aqueous acetic acid solution includes 0.8 parts of glacial acetic acid and 38.2 parts of deionized water) to promote the dissolution of chitosan. The remaining steps are the same as in Example 1.

[0060] Comparative Example 4: The only difference between the preparation method of this comparative hydrogel electrode and that of Example 1 is: In step (1), the raw materials and proportions of the precursor solution are different, as follows: The precursor solution consists of the following components in parts by weight: Thermosensitive polymer: 24 parts gelatin; conductive inorganic salt: 2 parts sodium chloride (NaCl); water-retaining agent: 30 parts glycerin; solvent: 44 parts deionized water; The remaining steps are the same as in Example 1.

[0061] Comparative Example 5: The preparation method of this comparative hydrogel electrode includes the following steps: (1) Preparation of precursor solution In a constant temperature water bath at 75 ℃, the raw materials were mixed in the following weight proportions and stirred continuously until completely dissolved to obtain a homogeneous and clear hydrogel precursor solution: Thermosensitive polymer: 24 parts gelatin; conductive inorganic salt: 2 parts sodium chloride (NaCl); water-retaining agent: 30 parts glycerin; adhesive: 4 parts dopamine hydrochloride; crosslinking agent: 1 part tannic acid; solvent: 39 parts deionized water.

[0062] (2) High-speed shear foaming The above hydrogel precursor solution was transferred to a high-speed emulsifier and sheared and foamed at 2000 rpm for 5 min.

[0063] (3) Layered bonding The foaming system from step (2) is first coated into a layer at the target location and then cured by gelation to form a hydrogel layer. Subsequently, a mesh Ag / AgCl conductive layer with a pore size of 0.2 mm and a thickness of 0.2 mm is covered on the surface of the hydrogel layer at room temperature, and then lightly pressed and bonded with a pressure of 0.02 MPa-0.10 MPa for 1 min-3 min to obtain a layered bonded hydrogel electrode.

[0064] Comparative Example 6: The only difference between the preparation method of this comparative hydrogel electrode and that of Example 1 is: In step (1), the raw materials and proportions of the precursor solution are different, as follows: The precursor solution consists of the following components in parts by weight: Thermosensitive polymer: 32 parts gelatin; conductive inorganic salt: 2 parts sodium chloride (NaCl); water-retaining agent: 20 parts glycerin; adhesive: 11 parts dopamine hydrochloride; crosslinking agent: 3 parts tannic acid; solvent: 32 parts deionized water; The remaining steps are the same as in Example 1.

[0065] Test example: The hydrogel electrodes prepared in the above embodiments and comparative examples were subjected to the following performance tests and characterizations.

[0066] (1) Measurement of average pore size and morphology of hydrogel Test method: Cut a hydrogel sample of appropriate size from the effective area (such as the middle) of the hydrogel electrode, observe its cross-sectional morphology and take pictures using an optical microscope, and use ImageJ software to randomly count the diameter of at least 200 pores and calculate the average pore diameter of the hydrogel.

[0067] (2) Hydrogel porosity test The determination was performed using the immersion-Archimedes gravity method.

[0068] The specific testing method is as follows: the hydrogel sample to be tested is dried to constant weight, and the mass W is recorded. d Then, immerse it in anhydrous ethanol under vacuum until saturated, weigh it, and record the saturated mass W. s The saturated sample was then immersed in ethanol, and the immersion mass W was recorded. i ; Porosity (P) is calculated using the following formula: P=(W s -W d ) / (W s -W i )×100%.

[0069] (3) Water vapor transmission rate test The water vapor transmission rate of the hydrogel samples was determined by referring to ASTM E96 standard and incorporating intrinsic water exchange correction for the hydrogel samples.

[0070] The specific testing method is as follows: Sample preparation: Prepare a hydrogel sample with a thickness of 1 mm and cut it into a square that matches the mouth of the test bottle; at the same time, prepare a reference hydrogel sample with the same initial water content, thickness and exposed area as the test sample for environmental exchange calibration. Test conditions: The test was conducted in a constant temperature and humidity environment (temperature 25±0.5 ℃, relative humidity 30±2% RH) for a period of 2 days. A permeation cup or weighing bottle was used for the test. The inner diameter d of the bottle opening was measured, and the effective exposure area A (unit: m²) was calculated based on the inner diameter of the bottle opening. 2 ); Test Procedure: Add 10 mL of deionized water to the permeation cup, tightly cover the hydrogel sample with the bottle opening, and seal the edges using clamps, sealing rings, or sealing films to ensure that water vapor is transferred only through the exposed area of ​​the bottle opening. Record the initial mass m of the assembled test system. 0,mea The test system was placed in the above-mentioned constant temperature and humidity environment. After 2 days, it was removed and weighed quickly, and the endpoint mass m was recorded. t,mea ; The mass of moisture loss from the test sample is calculated using the following formula: W mea =m 0,mea - m t,mea ; The reference hydrogel sample was exposed in parallel to the test group under the same environmental conditions (25 ℃, 30% RH) for 2 days. Record the initial mass m of the reference sample 0,ref and final mass m t,ref ; The change in the mass of the reference sample is calculated using the following formula: W ref =m 0,ref - m t,ref ; Among them, when the reference sample absorbs moisture and gains weight during the test, W ref It can be a negative value; To eliminate the influence of intrinsic water exchange between the hydrogel and the environment, the mass loss of the test sample minus the change in mass of the reference sample is used as the corrected water loss, and the corrected water vapor transmission rate (WVTR) is calculated using the following formula: ; Among them, W mea The mass of water loss (g) of the test sample; W ref A represents the change in mass of the reference sample (g); A represents the effective exposure area (m²). 2 ); t represents the test time (day).

[0071] (4) Contact impedance test Three-electrode AC impedance measurements were performed using a Princeton electrochemical workstation (VersaSTAT 3F) to evaluate the contact impedance at the electrode-skin interface.

[0072] Test method: The test area was the scalp covered by hair of the subject. The hydrogel electrode to be tested (a circle with a thickness of 1 mm and a diameter of 1 cm) was used as the working electrode (WE). Two other identical hydrogel electrodes of the present invention were attached to both ears as the reference electrode (RE) and the counter electrode (CE), respectively, to form a three-electrode test system. The test was conducted at room temperature, and the subject remained seated. The test frequency was set to 10 Hz, and the contact impedance of the electrode-skin interface at this frequency (unit: kΩ) was measured. The initial impedance value after electrode attachment and the impedance value at each time point during long-term wear were recorded.

[0073] (5) Interface peel strength test Test method: An electronic tensile testing machine (model MTS E43.504) with a 90° peel fixture was used to perform 90° peel tests on the adhesion performance of the hydrogel electrode of the present invention at the skin-hydrogel interface and the hydrogel-electrode interface. The sample size was 1.5 cm long, 1.0 cm wide, and 1 mm thick. The peel rate was 50 mm / min. The force-displacement curves were recorded, and the interfacial peel strength (G) was calculated using the following formula: G=F / w; Where F is the steady-state peeling force and w is the effective peeling width of the peeled sample.

[0074] (6) Signal-to-noise ratio test Electroencephalogram (EEG) testing methods: Hydrogel electrodes were attached to the occipital scalp (hair-covered area) of the subjects, and EEG signals were monitored at room temperature in both resting and exercise states (treadmill speed 2 km / h). During the test, accelerometers and EEG acquisition devices were used to record signals synchronously for subsequent cadence recognition and motion artifact analysis. The test parameters are set as follows: sampling rate: 250 Hz; single monitoring duration: no less than 2 min; subject status control: keep the eyes looking forward and minimize talking when at rest; maintain a stable gait and do not hold onto handrails when moving.

[0075] In addition, the long-term wear stability test (8 days) was conducted by repeating the above test on the subject while continuously wearing the matching hydrogel electrode.

[0076] Signal-to-noise ratio (SNR) calculation method: The Welch method was used to estimate the power spectral density (PSD) of EEG signals acquired in resting and moving states (2 km / h), with the analysis frequency band being 1.5–30 Hz; the stride frequency f was determined based on the synchronous acceleration signal (or gait information). step The spectral peaks from the step frequency up to the sixth harmonic are defined as motion noise peaks. The noise power P is obtained by integrating the frequency bands surrounding each noise peak over a preset bandwidth. noise The total power within the 1.5-30 Hz frequency band is integrated to obtain the total power P. total ; Effective signal power (P) signal Calculate using the following formula: ; The signal-to-noise ratio is calculated using the following formula: ; Signal-to-noise ratio tests for electrooculography (EOG), electrocardiography (ECG), and electromyography (EMG) signals were all performed using conventional lead methods at the respective test sites. The specific test methods are as follows: Electrooculography (EOG) signal testing method: Electromyography (EMG) signals were acquired using an electromyography-evoked potential (MEB-2312C, NIHON KOHDEN, Japan) instrument. The hydrogel electrodes were deployed in a dual-channel configuration: vertical EMG electrodes were attached to the upper eyelid and lower eyelid of the ipsilateral eye, respectively; reference / ground electrodes were attached to the forehead or mastoid process. Subjects were tested in a seated position at room temperature, keeping their heads as stable as possible and minimizing talking.

[0077] The sampling rate was set to 250 Hz, with a bandpass filter of 0.05–30 Hz, and the continuous recording time was no less than 2 minutes. The test tasks included resting gaze and vertical saccades as prompted, used to form baseline and event segments.

[0078] Signal-to-noise ratio (SNR) is calculated using the event-baseline method: the RMS of the resting gaze segment is used as the noise amplitude. The RMS value of the scanned event segment is used as the signal amplitude. And calculate using the following formula: ; The SNR of the electrooculogram (EOG) signal was obtained, and the average value was used to characterize the EOG signal quality.

[0079] Electrocardiogram (ECG) testing method (resting): The hydrogel electrode was attached to the test subject's chest according to the standard ECG lead method. ECG signals were collected in a resting state (for no less than 2 minutes), and then bandpass filtered (e.g., 0.5-40 Hz) before being used for signal-to-noise ratio (SNR) analysis.

[0080] Signal-to-noise ratio (SNR) calculation method: Detect the R-peak and extract multiple heartbeat waveform segments (e.g., each heartbeat takes a fixed time window); align and average the multiple heartbeat segments to obtain the average heartbeat template, denoted as . (Consider it a valid signal); subtract each heartbeat segment from the template to obtain the residual. (Consider it as noise); calculate the signal power using the following formulas respectively. and noise power : ; And calculate the electrocardiogram signal-to-noise ratio using the following formula: ; Electromyography (EMG) testing method (resting): Hydrogel electrodes were attached to the surface of the target muscle (such as the forearm flexor muscles) of the subject, and electromyography (EMG) signals were acquired in a resting state, with a single recording time of no less than 2 minutes. The power obtained after the resting state signal underwent the same preprocessing and power calculation was recorded as the noise power. The subjects were then instructed to hold a 10 kg hand grip dynamometer and maintain this position (or repeat the grip strength task in a standardized manner), and electromyographic (EMG) signals were collected. Each recording session lasted at least 2 minutes. The power obtained after the signal under this state underwent the same preprocessing and power calculation was recorded as the effective signal power. ; And calculate the myoelectric signal-to-noise ratio using the following formula: .

[0081] Table 1 Performance test results of the hydrogel electrode in the embodiment. Table 2 Performance test results of comparative hydrogel electrodes As shown in Tables 1 and 2, the hydrogel electrodes of this invention, through multi-scale structural design and the synergy of multiple components, successfully constructed an integrated structure of a through-pore foam skeleton and a three-dimensional interlocking conductive network, achieving a systematic improvement in performance in terms of breathability, moisture retention, interfacial mechanics, and electrophysiological signal acquisition. Specifically, from the perspective of microstructure and breathability, the hydrogel electrodes of this invention all form a foam network structure with uniform pore size and through-pores, which provides an efficient and low-resistance channel for water vapor diffusion, thereby achieving excellent breathability while maintaining good water retention capacity. For example, the average pore size of the foam in Example 1 is 405 μm, the porosity is 81%, and its water vapor transmission rate is as high as 2809 g / m³. 2 ·day -1 Comparative Example 1, on the other hand, has a dense structure (pore size ≤ 5 μm) and a water vapor permeability of only 650 g / m³. 2 ·day -1 The average pore size of the foams in Comparative Examples 2-4 was less than 60 μm, the porosity was less than 25%, and the WVTR was less than 800 g / m³. 2 ·day -1 Its breathability is severely limited.

[0082] Secondly, from the perspective of interfacial mechanical properties, the embodiments of the present invention utilize the brief gelation window after foaming to precisely embed the mesh conductive layer into the foaming system. The rapid phase transition of the temperature-sensitive matrix enables the gel network to in-situ encapsulate and three-dimensionally lock the mesh fibers, forming an integrated mechanically interlocked structure. This structure endows the hydrogel-electrode interface with extremely high peel resistance. Simultaneously, the adhesive functional components significantly enhance the skin-gel interface bonding strength through multiple hydrogen bonds and coordination interactions. For example, the hydrogel-electrode peel force in Example 1 is as high as 36.1 N / cm, and the skin-hydrogel peel force is as high as 15.5 N / cm. In contrast, Comparative Example 5, although having the same formulation as Example 1, did not embed the mesh during the gelation window, only forming a layered adhesive structure. The hydrogel-electrode peel force plummeted to 0.9 N / cm, and the electrode layer detached after only 2 days of wear. The skin-hydrogel peel force in Comparative Example 4 was only 0.9 N / cm, and the peel force in Comparative Example 3 also decreased significantly.

[0083] Furthermore, from the perspective of electrophysiological signal acquisition performance, the hydrogel electrode of this invention, with its large pores that physically accommodate hair, its high breathability that effectively inhibits sweat accumulation, and its strong interface adhesion that intrinsically constrains movement displacement, can maintain stable, low-noise electrode-skin contact for a long time even in hair-covered areas. This results in extremely low motion artifact interference and excellent long-term signal fidelity. For example, the hydrogel electrode of Example 1 has a contact impedance of less than 10 kΩ, and its EEG noise ratio remains at 18.3 dB during movement. Even after 8 days of continuous wear, the EEG noise ratio still reaches 10.3 dB, demonstrating excellent anti-motion artifact capability. In contrast, Comparative Examples 1 and 2 dropped to 4.2 dB and 4.1 dB after 3 days, Comparative Examples 3 and 4 fell below 3 dB after 1 day, and Comparative Example 5 completely failed due to electrode layer detachment.

[0084] Therefore, the hydrogel electrode of the present invention achieves breakthrough synergy in terms of unifying the contradiction between breathability and moisturization, deep integration of the conductive layer and the gel phase, and dynamic adaptation of the skin interface through multi-scale structural design and system integration of multiple components. It fundamentally solves the long-standing industry problems of poor breathability, easy shedding, hair incompatibility and significant motion artifacts of traditional hydrogel electrodes, and provides a feasible engineering path for monitoring physiological electrical signals in dynamic, long-term and real-world scenarios.

[0085] Finally, it should be noted that the above embodiments are only used to illustrate the technical solutions of the present invention and are not intended to limit the scope of protection of the present invention. Although the present invention has been described in detail with reference to preferred embodiments, those skilled in the art should understand that modifications or equivalent substitutions can be made to the technical solutions of the present invention without departing from the essence and scope of the technical solutions of the present invention.

Claims

1. A hydrogel electrode for long-term bioelectrical signal monitoring, characterized in that, It includes a hydrogel with a porous structure and a mesh conductive layer embedded inside the hydrogel and forming a three-dimensional mechanically interlocked structure with the hydrogel; The hydrogel has an average pore size of 100 μm-500 μm and a porosity of 60%-90%. The raw materials for preparing the hydrogel include the following components in parts by weight: 5-30 parts of thermosensitive polymer, 0.5-15 parts of conductive component, 5-40 parts of water-retaining agent, 0.1-10 parts of adhesive and 0.01-2 parts of crosslinking agent.

2. The hydrogel electrode for long-term bioelectrical signal monitoring as described in claim 1, characterized in that, The raw materials for preparing the hydrogel include the following components in parts by weight: 6-28 parts of thermosensitive polymer, 1-7 parts of conductive component, 12-35 parts of water-retaining agent, 0.5-6 parts of adhesive and 0.8-1.5 parts of crosslinking agent; and / or, the raw materials for preparing the hydrogel also include 10-60 parts of solvent; the solvent includes water.

3. The hydrogel electrode for long-term bioelectrical signal monitoring as described in claim 1, characterized in that, The thermosensitive polymer includes at least one of gelatin, agar, agarose, κ-carrageenan, poly-N-isopropylacrylamide, and polyoxyethylene-polyoxypropylene block copolymer; and / or, the conductive component includes at least one of conductive inorganic salt, proton-conductive material, metal or carbon-based conductive filler; and / or, the water-retaining agent includes polyol; and / or, the adhesive includes dopamine hydrochloride and / or gallic acid; and / or, the crosslinking agent includes tannic acid and / or citric acid.

4. The hydrogel electrode for long-term bioelectrical signal monitoring as described in claim 3, characterized in that, The conductive inorganic salt includes at least one of potassium chloride, sodium chloride, lithium chloride, and sodium sulfate; and / or, the proton-conducting material includes at least one of phytic acid, sulfonic acid polymer, and choline-based ionic liquid; and / or, the metal or carbon-based conductive filler includes at least one of silver nanowires, conductive carbon black, graphene, and reduced graphene oxide.

5. The hydrogel electrode for long-term bioelectrical signal monitoring as described in claim 1, characterized in that, The mesh conductive layer is a silver / silver chloride mesh conductive layer with a mesh pore size greater than 0.02 mm. The opening structure of the mesh conductive layer includes at least one of the following: grid structure, honeycomb structure, serpentine structure, circular hole array structure, leaf vein structure, slit / narrow slit structure, herringbone structure, fishbone structure, expanded metal mesh structure, triangular array structure, square array structure, rhombic array structure, Voronoi cell network structure, random pore array structure, quasi-periodic pore array structure, pore size gradient structure, and pore density gradient structure.

6. The method for preparing the hydrogel electrode for long-term bioelectrical signal monitoring according to any one of claims 1-5, characterized in that, Includes the following steps: S1. Mix the thermosensitive polymer, conductive component, water-retaining agent, adhesive and crosslinking agent to obtain hydrogel precursor solution; S2. The hydrogel precursor liquid is sheared and foamed to form a foaming system; S3. The foaming system is coated, injected or printed onto the target site, and then a mesh conductive layer is embedded therein. The foaming system is solidified through phase transformation to form a three-dimensional mechanical interlocking structure, thereby obtaining the hydrogel electrode for long-term bioelectric signal monitoring.

7. The preparation method according to claim 6, characterized in that, In step S1, the mixing temperature is 45 ℃-95 ℃.

8. The preparation method according to claim 6, characterized in that, In step S2, the rotation speed of the shear foaming is 300 rpm to 5000 rpm, and the time is 0.5 min to 20 min.

9. The preparation method according to claim 6, characterized in that, In step S3, the phase transformation is carried out by temperature induction, so that the foaming system completes the sol-gel transition; the temperature-induced cooling rate is 3 ℃ / min-20 ℃ / min.

10. The application of the hydrogel electrode for long-term bioelectrical signal monitoring as described in any one of claims 1-5 in electroencephalogram (EEG) electrodes, electrocardiogram (ECG) electrodes, electromyogram (EMG) electrodes, electrooculogram (EOG) electrodes, or skin-touch biosensors that are worn continuously in daily life.

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