Graphene hydrogel micro-current stable wearable bioelectrode and preparation method thereof

Abstract

The present application relates to the technical field of electrocardio electrode, and relates to a graphene hydrogel micro-current stable wearable bioelectrode and a preparation method thereof. The electrode is sequentially stacked from bottom to top with a flexible substrate layer, a conductive layer, an interface modification layer, a water locking protection layer and a high water retention layer. The conductive layer is covered on the flexible substrate layer. The interface modification layer is formed on the surface of the conductive layer by sequentially performing oxygen plasma treatment and APTES silanization treatment. The water locking protection layer is a hydrogel pre-solidified on the interface modification layer by covalent bonding. The high water retention layer is a hydrogel co-solidified with the water locking protection layer. The cross-linking density of the high water retention layer is lower than that of the water locking protection layer. The electrode is further provided with an encapsulation layer, which is a porous encapsulation film covering the upper surfaces of the flexible substrate layer and the conductive layer, and a window exposing the composite hydrogel is arranged in the middle of the encapsulation layer. The bioelectric signal, especially the micro-current signal, can be continuously and stably collected with high fidelity in dynamic and long-term monitoring scenarios.

Description

Technical Field

[0001] This invention relates to the field of electrocardiogram electrode technology, and specifically to a graphene hydrogel microcurrent-stabilized wearable bioelectrode and its preparation method. Background Technology

[0002] Wearable bioelectrical signal monitoring technology is an important development direction in the healthcare field. The acquisition of signals such as electrocardiogram (ECG), electromyography (EMG), and electroencephalography (EEG) relies on bioelectrodes that come into direct contact with the skin. Currently, mainstream electrodes include traditional Ag / AgCl gel wet electrodes and emerging dry electrodes. Wet electrodes reduce skin interfacial impedance through conductive gel, resulting in better signal quality and mature technology. Dry electrodes, on the other hand, directly contact the skin through metal or carbon-based materials, avoiding gel, offering better durability, and eliminating issues of gel allergies or dryness, making them more suitable for long-term wear.

[0003] Despite the aforementioned advancements in existing technologies, the following inherent limitations remain in achieving dynamic, long-term, high-fidelity bioelectrical signal monitoring: The core function of the conductive gel in wet electrodes relies on its internal moisture. However, in ambient air, the moisture in the gel inevitably evaporates continuously. As moisture is lost, the gel's conductivity decreases, and its contact impedance with the skin increases significantly, leading to a sharp decline in signal quality over time. This rapid deterioration typically limits its effective operation to only a few hours, failing to meet the demands of continuous monitoring for several days. Furthermore, the gel may cause skin allergies. Dry electrodes, on the other hand, have high and unstable contact impedance with the skin, making them highly susceptible to motion artifacts. Their signal quality, particularly their ability to acquire weak signals, is far inferior to wet electrodes in dynamic environments, making them unsuitable for medical-grade diagnostics.

[0004] In summary, current technologies lack a wearable electrode that can simultaneously achieve the low impedance / high signal quality of wet electrodes and the long-term durability / non-allergenicity of dry electrodes. The fundamental problem lies in the inability to construct and maintain a stable skin-electrode interface with excellent ion-electron conversion efficiency under dynamic usage conditions. This severely restricts the development of wearable medical devices from short-term monitoring to long-term management.

[0005] The information disclosed in this background section is intended only to enhance the understanding of the overall background of the present invention and should not be construed as an admission or in any way implying that the information constitutes prior art known to those skilled in the art. Summary of the Invention

[0006] The first objective of this invention is to provide a graphene hydrogel microcurrent-stabilized wearable bioelectrode that maintains a stable microcurrent interface impedance, enabling continuous high-fidelity and stable acquisition of bioelectrical signals, especially microcurrent signals, in dynamic and long-term monitoring scenarios.

[0007] The above-mentioned technical objective of the present invention is achieved through the following technical solution:

[0008] A graphene hydrogel microcurrent-stabilized wearable bioelectrode is provided with a flexible substrate layer, a conductive layer, an interface modification layer, a water-locking protective layer and a high water retention layer stacked sequentially from bottom to top.

[0009] The conductive layer is a graphene-based conductive film covering a flexible substrate; for example, a screen-printed reduced graphene oxide film.

[0010] The interface modification layer is a silane molecular layer formed on the surface of the conductive layer through sequential oxygen plasma treatment and APTES silanization treatment; this completely avoids the interlayer delamination problem that may exist in physical adhesion, and significantly improves the mechanical stability and service life of the electrode under dynamic use such as bending and stretching. At the same time, it provides a low-impedance and stable current transmission channel.

[0011] The water-locking protective layer is a composite hydrogel pre-cured on the interface modification layer through covalent bonding;

[0012] The high water retention layer is a composite hydrogel formed by co-curing with the water-locking protective layer; the cross-linking density of the high water retention layer is lower than that of the water-locking protective layer; the water-locking protective layer and the high water retention layer form a gradient structure with water gradient regulation function; the mass percentage content of glycerol and sodium hyaluronate in the high water retention layer is higher than that in the water-locking protective layer, and the high water retention layer does not contain sodium alginate, and its cross-linking density is lower than that of the water-locking protective layer.

[0013] This invention employs a process of pre-curing a water-locking protective layer followed by co-curing a high-water-retention layer, creating an interpenetrating network or strong hydrogen bonds between the two layers, forming a unified whole with gradient functionality. The water-locking protective layer has a high cross-linking density and contains sodium alginate, forming a dense network that effectively prevents water from evaporating upwards. The high-water-retention layer has a low cross-linking density and is rich in hydrophilic components, allowing for targeted moisture replenishment to the skin through a concentration gradient, achieving active moisture management and long-term maintenance. The synergy of these two components significantly extends the effective working time of the electrode in dry environments and maintains long-term stability of the interfacial impedance.

[0014] The electrode also features an encapsulation layer, which is a porous polyurethane film covering the flexible substrate and the conductive layer, with a central window exposing the composite hydrogel. The porous encapsulation film covers the non-contact area, allowing skin breathability and improving comfort, while protecting the vulnerable conductive layer. The central window precisely exposes the composite hydrogel, ensuring direct and effective contact with the skin, enhancing product reliability and wearing comfort.

[0015] This invention significantly reduces signal baseline drift and motion artifacts caused by changes in interface state, thereby achieving high signal-to-noise ratio and high stability acquisition of weak bioelectrical signals under conditions of human dynamic activity and long-term wear.

[0016] Preferably, the graphene-based conductive film is a screen-printed reduced graphene oxide film. By mass percentage, the raw materials of the reduced graphene oxide film include: 10-15 wt% rGO powder, 3-5 wt% polyvinyl alcohol, 0.5-1 wt% sodium dodecylbenzenesulfonate, and the balance being deionized water.

[0017] Reduced graphene oxide (rGO) provides the main conductive pathways, and its two-dimensional sheet structure facilitates the formation of a continuous conductive network. Polyvinyl alcohol (PVA) acts as a binder and film-forming agent. The hydroxyl groups on its molecular chain form strong hydrogen bonds with the residual oxygen-containing functional groups on the rGO sheets, enhancing the film's mechanical toughness and adhesion to flexible substrates, preventing the rGO film from becoming brittle or detaching from the substrate. Sodium dodecylbenzenesulfonate, as an anionic surfactant, has its hydrophobic ends adsorbed on the rGO sheets and its hydrophilic ends extending into the water. Through steric hindrance, it effectively prevents the rGO powder from re-agglomerating and settling in the slurry, ensuring the uniformity of the slurry and its printability, thereby obtaining a conductive film with uniform thickness and stable performance.

[0018] Preferably, the interface modification layer introduces abundant oxygen-containing functional groups such as hydroxyl (-OH) and carboxyl (-COOH) groups into the graphene-based conductive film through oxygen plasma treatment, thereby activating a chemically inert hydrophobic surface into a hydrophilic surface rich in reaction sites. Subsequently, the activated graphene is immersed in an APTES solution. The ethoxy groups of APTES first hydrolyze to generate highly reactive silanol groups (Si-OH). These silanol groups undergo a condensation reaction with the aforementioned oxygen-containing functional groups to form Si-OC covalent bonds. The other... The amino group (-NH2) at one end faces outward, resulting in a silane molecule capping layer with an amino group at the end, awaiting reaction with the hydrogel. Then, a hydrogel precursor liquid containing methacrylamide gelatin (GelMA) is coated onto the interface-modified layer. Under the action of a photoinitiator (LAP) and ultraviolet light, the exposed nucleophilic amino groups on the interface-modified layer actively attack the electron-deficient methacrylamide double bonds (C=C) in the GelMA molecules at the interface, undergoing a Michael addition reaction to form stable carbon-nitrogen (CN) covalent bonds. Thus, through covalent bonding, the graphene conductive layer and the hydrogel layer are firmly integrated into a unified whole.

[0019] It is important to note that there is a competitive relationship between cross-linking between hydrogel molecules and interfacial bonding. The LAP photoinitiator absorbs light energy to generate free radicals, which rapidly initiates the free radical polymerization of the methacryloyl groups on the GelMA molecular chain—a highly efficient chain reaction. In contrast, the nucleophilic attack of the amino groups on the methacryloyl C=C double bond by the amino groups in the interfacial modification layer is a relatively slow, stepwise process. Both compete for the double bond reaction sites.

[0020] This invention involves coating a water-locking protective layer precursor liquid onto the interface, and then using a low light intensity (5-15 mW / cm²). 2 The system is subjected to short-term irradiation. At this time, the dominant reaction is bulk radical polymerization, forming a preliminary three-dimensional network. During the process, the viscosity of the system increases sharply, and the molecular mobility decreases. Although amino groups are present at the interface, their Michael addition reaction with double bonds is much slower at room temperature than that of radical polymerization, and is significantly suppressed during the pre-curing stage.

[0021] Subsequently, a high water-retention precursor is coated onto the surface of the water-locking protective layer, using a higher light intensity (10-20 mW / cm²). 2 The irradiation was prolonged (5-10 min). In the bulk region, the high concentration of free radicals dominated the polymerization reaction, promoting the fusion of the two hydrogel layers into a whole. In the interfacial region, due to steric hindrance, the free radical concentration was relatively low, but the local concentration of amino groups was extremely high. In this microenvironment of high amino groups and relatively low free radical concentration, the Michael addition reaction gained a competitive advantage. The amino groups nucleophilically attacked unreacted double bonds in the nearby GelMA network, forming CN covalent bonds. In addition, the microthermal environment provided by ultraviolet light irradiation further accelerated the reaction, ultimately achieving a stable interfacial connection.

[0022] Preferably, the pre-curing degree of the water-locking protective layer is controlled between 50% and 80%. At this stage, the water-locking protective layer has initially formed a three-dimensional network, but the surface still contains a large number of unreacted methacryloyl groups and a certain degree of fluidity. When the precursor liquid for the high water-retention layer is applied, the uncured portion interpenetrates with the monomers of the high water-retention layer, forming an interpenetrating network or strong covalent bonds during the second curing, thus firmly bonding the two layers and preventing delamination. Pre-curing is performed using ultraviolet light with a wavelength of 365-405 nm and a light intensity of 5-15 mW / cm². 2 The irradiation is completed in 1-3 minutes.

[0023] The pre-curing degree of the water-locking protective layer is achieved by controlling the energy density of the UV curing process. Energy density is the product of UV light intensity and irradiation time, and its calculation formula is: Energy density (mJ / cm²) 2 = Light intensity (mW / cm) 2 () × time (s). The inventors, through extensive experimental verification, found that when the energy density is controlled between 1100-1800 mJ / cm³... 2Within the specified range, the pre-curing degree of the water-locking protective layer can be kept within the ideal range of 50%-80%, and ultimately form an excellent gradient interpenetrating network structure with the high water retention layer.

[0024] Preferably, the water-locking protective layer is formed on the interface-modified layer by photocuring with a LAP photoinitiator using methacrylamide gelatin, glycerin, betaine, KCl, sodium hyaluronate, and sodium alginate as raw materials. By mass percentage, the raw materials of the water-locking protective layer include: 15-20 wt% methacrylamide gelatin, 8-10 wt% glycerin, 2-3 wt% betaine, 0.4-0.6 wt% KCl, 0.1-0.15 wt% sodium hyaluronate, 0.1-0.3 wt% LAP photoinitiator, 0.2-0.5 wt% sodium alginate, and the balance being deionized water.

[0025] Preferably, the high water retention layer is formed on the water-locking protective layer by photocuring with a LAP photoinitiator using methacrylamide gelatin, glycerin, betaine, KCl, and sodium hyaluronate as raw materials. By mass percentage, the raw materials of the high water retention layer include: 15-20 wt% methacrylamide gelatin, 10-12 wt% glycerin, 2-3 wt% betaine, 0.4-0.6 wt% KCl, 0.15-0.2 wt% sodium hyaluronate, 0.1-0.3 wt% LAP photoinitiator, and the balance being deionized water.

[0026] In this invention, sodium alginate is only incorporated into the water-locking protective layer. The addition of sodium alginate, especially after pre-curing, imparts higher initial mechanical strength and modulus to this layer. This allows it to form a more stable and less easily damaged interface when covalently bonded with the interface modification layer. Secondly, sodium alginate, as a natural high-molecular-weight polysaccharide, is extremely sensitive to calcium ions in human sweat. When the electrode is worn on the skin, trace amounts of calcium ions in the sweat penetrate into the water-locking protective layer and undergo in-situ secondary cross-linking with sodium alginate. This greatly increases the durability of the layer. Simultaneously, the slightly lower glycerol and sodium hyaluronate content ensures the density and dimensional stability of the network structure.

[0027] The high-water-retention layer maintains a low cross-linking density and higher softness, allowing it to better adhere to the skin. A higher proportion of glycerin and sodium hyaluronate significantly enhances the layer's water absorption and moisture content, maintaining long-term moisture retention at the electrode-skin interface and ensuring stable impedance through continuous moisture release. The synergistic effect of the two layers establishes and maintains a stable moisture gradient from the skin to the electrode, thus guaranteeing long-term stability of the interfacial impedance. The water-locking protective layer, with its dense network, effectively blocks the path of moisture loss towards the electrode substrate. All moisture evaporation occurs only on the surface of the high-water-retention layer in contact with air / skin, thereby maintaining the moisture of the working interface with extremely high efficiency.

[0028] In this formula, betaine, an amphoteric compound, stabilizes the structure of proteins and biomolecules, reducing damage from external stressors. When the electrodes are worn for extended periods, betaine effectively protects skin cells, reducing irritation and itching caused by moisture imbalance, thus fundamentally lowering the risk of allergic reactions. The use of a low concentration of KCl (0.4-0.6 wt%) greatly avoids skin irritation caused by high osmotic pressure, fundamentally eliminating the risk of contact discomfort and allergic dermatitis due to excessive electrolyte concentration. All core components are derived from natural and highly biocompatible raw materials, resulting in a gentle overall formula that significantly reduces the probability of sensitization.

[0029] Preferably, the porous polyurethane film has a porosity of 30-50% and a pore size of 5-10 μm. This allows perspiration from skin respiration to escape promptly while preventing the intrusion of external liquid water, thus improving the electrode's durability.

[0030] Preferably, the edge of the encapsulation layer extends outward to form a wing-like structure. A medical pressure-sensitive adhesive layer is disposed on the lower surface of the wing-like structure, and release paper is adhered to the lower surfaces of the medical pressure-sensitive adhesive layer and the hydrogel layer. This ensures the electrodes remain stable and do not fall off during dynamic activities. The hydrogel layer provides conductivity and wet adhesion to the skin, while the pressure-sensitive adhesive of the wing-like structure provides strong mechanical dry adhesion. These two elements complement each other, ensuring secure wear under various movement conditions. In normal use, where frequent stress is not required, hydrogel adhesion alone can be used.

[0031] Preferably, the medical pressure-sensitive adhesive is an acrylic pressure-sensitive adhesive with a thickness of 5-10 μm, a 180° peel strength ≥0.8 N / cm, and a skin sensitization rate ≤0.5%.

[0032] Preferably, the thickness of the flexible substrate layer is 25-50 μm; the thickness of the conductive layer is 5-15 μm; the thickness of the interface modification layer is <0.1 μm; the thickness of the hydrogel layer is 80-150 μm; and the thickness of the encapsulation layer is 25-50 μm. Optimized matching of the thicknesses of each layer achieves an optimal balance between electrode flexibility, mechanical strength, electrical properties, and water retention capacity.

[0033] The second objective of this invention is to provide a method for preparing a graphene hydrogel microcurrent-stabilized wearable bioelectrode, which combines the high conductivity of graphene, the biocompatibility of hydrogel, and the flexible wearability, making it particularly suitable for long-term stable monitoring of bioelectrical signals (such as electrocardiogram (ECG), electromyography (EMG), and electroencephalography (EEG)).

[0034] The above-mentioned technical objective of the present invention is achieved through the following technical solution:

[0035] A method for preparing a graphene hydrogel microcurrent-stabilized wearable bioelectrode includes the following steps:

[0036] S1 provides a flexible substrate layer;

[0037] Preferably, in step S1, a PET film with a thickness of 25-50 μm is used as a flexible substrate layer. After wiping the surface with 75% anhydrous ethanol, it is dried at 60°C for 10-15 minutes. The ethanol wiping can effectively remove organic contaminants, grease, and dust from the surface of the PET film, providing a clean and highly active surface for subsequent printing of the conductive layer and ensuring the adhesion between the conductive layer and the substrate. Drying at 60°C thoroughly removes residual ethanol and moisture, avoiding the formation of bubbles or a decrease in adhesion during the subsequent high-temperature curing step.

[0038] S2 prints graphene-based conductive paste onto the upper surface of a flexible substrate layer, which is then thermo-cured to form a conductive layer.

[0039] Preferably, in step S2, rGO powder, polyvinyl alcohol, sodium dodecylbenzenesulfonate, and deionized water are mixed according to the specified ratio and ultrasonically dispersed at 300W for 20-30 minutes to ensure that the rGO powder is fully dispersed and deagglomerated, resulting in a uniform and stable screen printing paste. The paste is then coated onto the upper surface of a flexible substrate using a screen printing process, controlling the wet film thickness to 20-30μm. After curing at 120-150℃ for 30-45 minutes, a screen-printed reduced graphene oxide film with a thickness of 5-15μm and a sheet resistance ≤30Ω / sq is formed. This ensures the mechanical flexibility, conductivity, and adhesion to the substrate of the conductive layer. The sheet resistance ≤30Ω / sq enables efficient transmission of microcurrent signals without generating significant heat or voltage drop, thus guaranteeing signal fidelity.

[0040] S3 performs oxygen plasma treatment on the conductive layer, followed by APTES silanization treatment to form an interface modification layer on its surface.

[0041] Preferably, in step S3, the conductive layer obtained in step S1 is subjected to oxygen plasma treatment: the power is set to 200-300W and the treatment time is 5-10min, introducing oxygen-containing functional groups such as -OH and -COOH on the surface of the conductive layer; a 1.5-2wt% APTES ethanol solution with pH 4.5-5.0 is prepared, and the plasma-treated conductive layer is immersed in the solution and left to stand at room temperature for 1-1.5h; after removal, it is vacuum dried at 70-80℃ for 20-30min, then soaked in 75% anhydrous ethanol for 10-15min to remove unreacted APTES, rinsed with deionized water 2-3 times, and dried at 60-70℃ for 15-20min to form a silane molecular layer with a thickness of <0.1μm on the surface of the conductive layer.

[0042] S4 is used to prepare a water-locking protective layer precursor liquid containing sodium alginate, which is then coated onto the interface modification layer and pre-cured under first-intensity ultraviolet light.

[0043] Preferably, in step S4, methacrylamide gelatin, glycerin, betaine, KCl, sodium hyaluronate, sodium alginate and deionized water are mixed in mass ratio, stirred at 50-60℃ for 15-20 min until completely dissolved, cooled to room temperature and then LAP photoinitiator is added to avoid premature decomposition and failure of LAP at high temperature, and stirred for 5-10 min to obtain the precursor liquid of the water-locking protective layer.

[0044] The precursor liquid for the water-locking protective layer is poured onto the upper surface of the interface-modified layer obtained in step S3, and light is applied at a wavelength of 365-405nm and an intensity of 5-15mW / cm. 2 Expose the GelMA to ultraviolet light for 1-3 minutes to allow it to pre-crosslink and cure.

[0045] S5 applies a high water retention layer precursor liquid to the pre-cured water-locking protective layer and then fully cures it under second intensity ultraviolet light to form a hydrogel layer with a gradient structure.

[0046] Preferably, in step S5, methacrylamide gelatin, glycerin, betaine, KCl, sodium hyaluronate and deionized water are mixed in mass ratio, stirred at 50-60°C for 15-20 min until completely dissolved, cooled to room temperature and then LAP photoinitiator is added, and stirred for 5-10 min to obtain a high water-retaining layer precursor liquid.

[0047] The precursor liquid for the high water retention layer is poured onto the upper surface of the water-locking protective layer obtained in step S4, and light is applied at a wavelength of 365-405nm and an intensity of 10-20mW / cm². 2 Irradiation with ultraviolet light for 5-10 minutes causes GelMA to crosslink and solidify, forming a hydrogel layer with a gradient structure, while simultaneously promoting the formation of covalent bonds between the amino groups of the water-locking protective layer and the interface modification layer.

[0048] S6 covers the flexible substrate and the upper surface of the conductive layer with a porous encapsulation film with a window, and exposes the hydrogel layer from the window to obtain the electrode;

[0049] Preferably, in step S6, a porous polyurethane film with a porosity of 30-50%, a pore size of 5-10 μm, and a thickness of 25-50 μm is taken, cut according to the electrode size, and a window matching the hydrogel layer is reserved in the middle to obtain the encapsulation layer; the encapsulation layer is covered on the upper surface of the flexible substrate layer and the conductive layer, so that the window is aligned with the hydrogel layer and its lower surface is exposed, and the edge of the encapsulation layer is attached with medical acrylate adhesive, and the core electrode structure is obtained after curing.

[0050] S7 sets a medical pressure-sensitive adhesive layer on the lower surface of the wing-shaped structure formed by extending the edge of the encapsulation layer of the electrode pre-finished product, and attaches release paper to the lower surface of the medical pressure-sensitive adhesive layer to obtain a wearable bioelectrode.

[0051] Preferably, in step S7, an acrylic medical pressure-sensitive adhesive with a thickness of 5-10 μm is coated on the lower surface of the wing-shaped structure extending outward from the encapsulation layer, and release paper is attached to complete the preparation of the wearable bioelectrode.

[0052] Compared with the prior art, the beneficial effects of the present invention are as follows:

[0053] This invention introduces abundant oxygen-containing functional groups at the edges and defects of graphene lattice through bombardment and oxidation by high-energy reactive oxygen species, thereby transforming the chemically inert sp... 2 The carbon surface is transformed into a highly reactive surface, and then a silane molecular layer is coated onto the graphene surface via covalent grafting of APTES silane molecules. One end of the silane is fixed to the graphene through Si-OC bonds, while the other end exposes an active amino-linked hydrogel. The covalently bonded interface forms a strong chemical bond, ensuring the continuity and stability of the electron transmission channel. This effectively avoids interlayer delamination or relative displacement caused by human movement, sweat erosion, etc., thus significantly suppressing the generation of motion artifacts and ensuring that the signal baseline remains stable over a wearing period of several days. The hydrogel layer forms a tight, moist, low-impedance contact with the skin, ensuring high ion-electron conversion efficiency; laying the physical foundation for achieving ultra-low noise acquisition of microcurrent-level signals.

[0054] This invention controls the pre-curing degree of the water-locking protective layer, enabling it to form a unique gradient interpenetrating network structure with the high water-retention layer during co-curing. This structure increases the bonding force between the two hydrogel layers from the level of ordinary physical adsorption to the level of chemical covalent bonding, effectively improving the interfacial shear strength. This ensures that the internal interface does not break down when the electrode is subjected to repeated bending, stretching, and other dynamic mechanical loads, thus significantly improving the anti-interference capability of dynamic signal acquisition and the product's lifespan. Simultaneously, the gradient hydrogel structure effectively blocks the directional replenishment of internal moisture and external moisture evaporation, significantly extending the electrode's working life in dry environments and preventing a sharp increase in impedance due to hydrogel dehydration. The lower flexible substrate layer and the upper composite hydrogel layer provide excellent mechanical matching, conforming to the human body's curved surface, reducing motion artifacts, and achieving high signal-to-noise ratio and high stability biosignal monitoring in microcurrent environments. Attached Figure Description

[0055] To more clearly illustrate the technical solutions in the embodiments of the present invention or the prior art, the drawings used in the description of the embodiments or the prior art will be briefly introduced below. Obviously, the drawings described below are only some embodiments recorded in the present invention. For those skilled in the art, other drawings can be obtained based on these drawings without creative effort.

[0056] Figure 1A schematic diagram of the structure of a graphene hydrogel microcurrent-stabilized wearable bioelectrode.

[0057] Figure 2 A cross-sectional view of a graphene hydrogel microcurrent-stabilized wearable bioelectrode.

[0058] Figure 3 A flowchart illustrating the steps involved in the fabrication of a graphene hydrogel microcurrent-stabilized wearable bioelectrode.

[0059] Reference numerals: 1. Flexible substrate layer; 2. Conductive layer; 3. Interface modification layer; 4. Water-locking protective layer; 5. High water retention layer; 6. Encapsulation layer. Detailed Implementation

[0060] To further illustrate the technical means and effects adopted by the present invention to achieve its intended purpose, the specific implementation methods, features, and effects of a graphene hydrogel microcurrent-stabilized wearable bioelectrode and its preparation method according to the present invention are described in detail below. It should be understood that the specific embodiments described herein are only for explaining the present invention and are not intended to limit the present invention.

[0061] Example 1:

[0062] The specific steps for fabricating an electrode for dynamic electrocardiogram (ECG) monitoring are as follows:

[0063] S1 Preparation of Flexible Substrate Layer 1:

[0064] Take a commercially available PET film with a thickness of 50μm and a smooth surface. Use a non-woven cloth dipped in 75% (v / v) anhydrous ethanol solution to carefully wipe its surface to remove grease and contaminants. Then place it in an oven at 60℃ for 15 minutes. After taking it out, cool it to room temperature for later use.

[0065] S2 is used to prepare conductive layer 2:

[0066] Slurry preparation: Weigh 12g of rGO powder, 4g of polyvinyl alcohol (PVA-1788), and 0.8g of sodium dodecylbenzenesulfonate, and add them to 83.2g of deionized water. Disperse the mixture ultrasonically at 300W for 25min to obtain a uniform and stable conductive slurry.

[0067] Printing and Curing: Using a 200-mesh polyester screen, the above paste is printed onto the surface of the treated PET film, controlling the wet film thickness to be approximately 25 μm. The printed film is then placed in a 135°C forced-air oven for 40 minutes to cure, and then allowed to cool naturally.

[0068] S3 constructs interface modification layer 3:

[0069] Oxygen plasma treatment: The PET film with conductive layer 2 was placed in an oxygen plasma cleaner and treated at 250W power for 8 minutes to fully activate the graphene surface and introduce oxygen-containing functional groups.

[0070] APTES silanization: A 1.8 wt% APTES ethanol solution was prepared, and the pH was adjusted to 4.8 with glacial acetic acid. The plasma-treated sample was immersed in this solution and allowed to stand at room temperature in the dark for 1.2 h. After removal, it was vacuum dried at 75 °C and -0.08 MPa for 25 min. Subsequently, the sample was immersed in 75% anhydrous ethanol for 12 min to remove physically adsorbed unreacted APTES molecules. Finally, it was rinsed three times with deionized water and dried at 65 °C for 18 min. A dense silane molecular layer with a thickness of less than 0.1 μm was finally formed on the surface of conductive layer 2.

[0071] S4 is used to prepare the water-locking protective layer 4:

[0072] Preparation of precursor solution: Weigh 18g of methacrylamide gelatin, 9g of glycerin, 2.5g of betaine, 0.5g of KCl, 0.12g of sodium hyaluronate, and 0.3g of sodium alginate, and add them to 69.38g of deionized water. Stir magnetically in a 55℃ water bath for 18 minutes until all components are completely dissolved. After cooling to room temperature, add 0.2g of LAP photoinitiator and stir for 8 minutes under light-protected conditions to ensure complete dissolution, obtaining a homogeneous and transparent precursor solution for the water-locking protective layer 4.

[0073] Pre-curing: The aforementioned precursor liquid was poured onto the interface modification layer 3, and the thickness was controlled to be approximately 60 μm using a scraper. It was then immediately placed in a UV curing chamber at a wavelength of 385 nm and a light intensity of 10 mW / cm². 2 Irradiate for 2 minutes (energy density = 1200 mJ / cm²) 2 At this point, the water-locking protective layer 4 has reached a semi-cured state, but there are still a large number of active groups inside.

[0074] S5 preparation of high water retention layer 5:

[0075] Preparation of precursor fluid: Weigh 18g of methacrylamide gelatin, 11g of glycerin, 2.5g of betaine, 0.5g of KCl, and 0.18g of sodium hyaluronate, and add them to 67.62g of deionized water. Dissolve the mixture by stirring in a 55℃ water bath. After cooling, add 0.2g of LAP photoinitiator and stir in the dark to obtain a high water-retaining layer 5 precursor fluid.

[0076] Co-curing: The precursor liquid of the high water retention layer 5 is poured onto the pre-cured water-locking protective layer 4, controlling the total hydrogel layer thickness to be 120 μm. A light intensity of 15 mW / cm² is used. 2The gel was irradiated with 385nm ultraviolet light for 8 minutes to complete the final curing. During this process, the two hydrogel layers interpenetrate and covalently cross-link at the interface, forming a strong gradient interpenetrating network structure.

[0077] S6 Packaging and Integration:

[0078] A commercially available porous polyurethane film with a thickness of 40 μm, a porosity of 40%, and an average pore size of 8 μm was used. A 10 mm diameter encapsulation layer 6 was punched out and placed over the electrode, precisely exposing the hydrogel layer through the window. The edges of the encapsulation layer 6 were then bonded and cured using medical-grade acrylic adhesive. Finally, an 8 μm thick medical-grade acrylic pressure-sensitive adhesive with a 180° peel strength of 1.0 N / cm was applied to the lower surface of the wing-like structure extending from the encapsulation layer 6, and release paper was attached. The electrode fabrication was thus completed. Figure 1 and Figure 2 The structure shown.

[0079] Example 2:

[0080] The specific steps for fabricating an electrode for dynamic electrocardiogram (ECG) monitoring are as follows:

[0081] S1 Preparation of Flexible Substrate Layer 1:

[0082] Take a commercially available PET film with a thickness of 40μm and a smooth surface. Use a non-woven cloth dipped in 75% (v / v) anhydrous ethanol solution to carefully wipe its surface to remove grease and contaminants. Then place it in an oven at 60℃ for 15 minutes. After taking it out, cool it to room temperature for later use.

[0083] S2 is used to prepare conductive layer 2:

[0084] Slurry preparation: Weigh 15g of rGO powder, 5g of polyvinyl alcohol (PVA-1788), and 1g of sodium dodecylbenzenesulfonate, and add them to 79g of deionized water. Disperse the mixture ultrasonically at 300W for 25 minutes to obtain a uniform and stable conductive slurry.

[0085] Printing and Curing: Using a 200-mesh polyester screen, the above paste is printed onto the surface of the treated PET film, controlling the wet film thickness to be approximately 25 μm. The printed film is then placed in a 135°C forced-air oven for 40 minutes to cure, and then allowed to cool naturally.

[0086] S3 constructs interface modification layer 3:

[0087] Oxygen plasma treatment: The PET film with conductive layer 2 was placed in an oxygen plasma cleaner and treated at 250W power for 8 minutes to fully activate the graphene surface and introduce oxygen-containing functional groups.

[0088] APTES silanization: A 1.8 wt% APTES ethanol solution was prepared, and the pH was adjusted to 4.8 with glacial acetic acid. The plasma-treated sample was immersed in this solution and allowed to stand at room temperature in the dark for 1.2 h. After removal, it was vacuum dried at 75 °C and -0.08 MPa for 25 min. Subsequently, the sample was immersed in 75% anhydrous ethanol for 12 min to remove physically adsorbed unreacted APTES molecules. Finally, it was rinsed three times with deionized water and dried at 65 °C for 18 min. A dense silane molecular layer with a thickness of less than 0.1 μm was finally formed on the surface of conductive layer 2.

[0089] S4 is used to prepare the water-locking protective layer 4:

[0090] Preparation of precursor fluid: Weigh 15g of methacrylamide gelatin, 8g of glycerin, 2g of betaine, 0.5g of KCl, 0.1g of sodium hyaluronate, and 0.3g of sodium alginate, and add them to 73.9g of deionized water. Stir magnetically in a 55℃ water bath for 18 minutes until all components are completely dissolved. After cooling to room temperature, add 0.2g of LAP photoinitiator and stir for 8 minutes under light-protected conditions to ensure complete dissolution, yielding a homogeneous and transparent precursor fluid for the water-locking protective layer 4.

[0091] Pre-curing: The aforementioned precursor liquid was poured onto the interface modification layer 3, and the thickness was controlled to be approximately 70 μm using a scraper. It was then immediately placed in a UV curing chamber at a wavelength of 385 nm and a light intensity of 9 mW / cm². 2 Irradiation for 2 minutes (energy density = 1080 mJ / cm²) 2 At this point, the water-locking protective layer 4 has reached a semi-cured state, but there are still a large number of active groups inside.

[0092] S5 preparation of high water retention layer 5:

[0093] Preparation of precursor fluid: Weigh 15g of methacrylamide gelatin, 10g of glycerin, 2g of betaine, 0.5g of KCl, and 0.1g of sodium hyaluronate, and add them to 72.2g of deionized water. Dissolve the mixture by stirring in a 55℃ water bath. After cooling, add 0.2g of LAP photoinitiator and stir in the dark to obtain a high water-retaining layer 5 precursor fluid.

[0094] Co-curing: The precursor liquid of the high water retention layer 5 is poured onto the pre-cured water-locking protective layer 4, controlling the total hydrogel layer thickness to be 120 μm. A light intensity of 15 mW / cm² is used. 2 The gel was irradiated with 385nm ultraviolet light for 8 minutes to complete the final curing. During this process, the two hydrogel layers interpenetrate and covalently cross-link at the interface, forming a strong gradient interpenetrating network structure.

[0095] S6 Packaging and Integration:

[0096] A commercially available porous polyurethane film with a thickness of 40 μm, a porosity of 40%, and an average pore size of 8 μm was used. A 10 mm diameter encapsulation layer 6 was punched out from this film and placed over the electrode, precisely exposing the hydrogel layer through the window. The edges of the encapsulation layer 6 were then bonded and cured using medical-grade acrylic adhesive. Finally, an 8 μm thick medical-grade acrylic pressure-sensitive adhesive with a 180° peel strength of 1.0 N / cm was applied to the lower surface of the wing-like structure extending from the encapsulation layer 6, and release paper was attached. The electrode fabrication was thus completed.

[0097] Example 3:

[0098] The specific steps for fabricating an electrode for dynamic electrocardiogram (ECG) monitoring are as follows:

[0099] S1 Preparation of Flexible Substrate Layer 1:

[0100] Take a commercially available PET film with a thickness of 30μm and a smooth surface. Use a non-woven cloth dipped in 75% (v / v) anhydrous ethanol solution to carefully wipe its surface to remove grease and contaminants. Then place it in an oven at 60℃ for 15 minutes. After taking it out, cool it to room temperature for later use.

[0101] S2 is used to prepare conductive layer 2:

[0102] Slurry preparation: Weigh 12g of rGO powder, 4g of polyvinyl alcohol (PVA-1788), and 0.8g of sodium dodecylbenzenesulfonate, and add them to 83.2g of deionized water. Disperse the mixture ultrasonically at 300W for 25min to obtain a uniform and stable conductive slurry.

[0103] Printing and Curing: Using a 200-mesh polyester screen, the above paste was printed onto the surface of the treated PET film, controlling the wet film thickness to be approximately 25 μm. The printed film was then placed in a 135°C forced-air oven for 40 minutes and allowed to cure naturally before cooling. The thickness of conductive layer 2 was measured to be approximately 10 μm, with a sheet resistivity of 18 Ω / sq.

[0104] S3 constructs interface modification layer 3:

[0105] Oxygen plasma treatment: The PET film with conductive layer 2 was placed in an oxygen plasma cleaner and treated at 250W power for 8 minutes to fully activate the graphene surface and introduce oxygen-containing functional groups.

[0106] APTES silanization: A 1.8 wt% APTES ethanol solution was prepared, and the pH was adjusted to 4.8 with glacial acetic acid. The plasma-treated sample was immersed in this solution and allowed to stand at room temperature in the dark for 1.2 h. After removal, it was vacuum dried at 75 °C and -0.08 MPa for 25 min. Subsequently, the sample was immersed in 75% anhydrous ethanol for 12 min to remove physically adsorbed unreacted APTES molecules. Finally, it was rinsed three times with deionized water and dried at 65 °C for 18 min. A dense silane molecular layer with a thickness of less than 0.1 μm was finally formed on the surface of conductive layer 2.

[0107] S4 is used to prepare the water-locking protective layer 4:

[0108] Preparation of precursor solution: Weigh 18g of methacrylamide gelatin, 9g of glycerin, 2.5g of betaine, 0.5g of KCl, 0.12g of sodium hyaluronate, and 0.3g of sodium alginate, and add them to 69.38g of deionized water. Stir magnetically in a 55℃ water bath for 18 minutes until all components are completely dissolved. After cooling to room temperature, add 0.2g of LAP photoinitiator and stir for 8 minutes under light-protected conditions to ensure complete dissolution, obtaining a homogeneous and transparent precursor solution for the water-locking protective layer 4.

[0109] Pre-curing: The aforementioned precursor liquid was poured onto the interface modification layer 3, and the thickness was controlled to be approximately 60 μm using a scraper. It was then immediately placed in a UV curing chamber at a wavelength of 385 nm and a light intensity of 13 mW / cm². 2 Irradiation for 2 minutes (energy density = 1560 mJ / cm²) 2 At this point, the water-locking protective layer 4 has reached a semi-cured state. The surface has been set and is no longer sticky, but there are still a large number of active groups inside.

[0110] S5 preparation of high water retention layer 5:

[0111] Preparation of precursor fluid: Weigh 18g of methacrylamide gelatin, 11g of glycerin, 2.5g of betaine, 0.5g of KCl, and 0.18g of sodium hyaluronate, and add them to 67.62g of deionized water. Dissolve the mixture by stirring in a 55℃ water bath. After cooling, add 0.2g of LAP photoinitiator and stir in the dark to obtain a high water-retaining layer 5 precursor fluid.

[0112] Co-curing: The precursor liquid of the high water retention layer 5 is poured onto the pre-cured water-locking protective layer 4, controlling the total hydrogel layer thickness to be 120 μm. A light intensity of 15 mW / cm² is used. 2 The gel was irradiated with 385nm ultraviolet light for 8 minutes to complete the final curing. During this process, the two hydrogel layers interpenetrate and covalently cross-link at the interface, forming a strong gradient interpenetrating network structure.

[0113] S6 Packaging and Integration:

[0114] A commercially available porous polyurethane film with a thickness of 40 μm, a porosity of 40%, and an average pore size of 8 μm was used. A 10 mm diameter encapsulation layer 6 was punched out. This encapsulation layer was then placed over the electrode to precisely expose the hydrogel layer through the window. The edges of the encapsulation layer 6 were then bonded together using medical-grade acrylic adhesive and cured.

[0115] Comparative Example 1:

[0116] Except for the composite hydrogel in steps S4 and S5, the remaining steps in this comparative example are the same as in Example 1, and will not be repeated here. However, a single hydrogel layer is used to replace the water-locking protective layer 4 and the high-water-retention layer 5 in Example 1.

[0117] Preparation of the hydrogel precursor solution: Weigh 18g of methacrylamide gelatin, 10g of glycerin, 2.5g of betaine, 0.5g of KCl, 0.15g of sodium hyaluronate, and 0.3g of sodium alginate, and add them to 68.35g of deionized water. Stir magnetically in a 55℃ water bath for 18 minutes until all components are completely dissolved. After cooling to room temperature, add 0.2g of LAP photoinitiator and stir for 8 minutes under light-protected conditions to ensure complete dissolution, obtaining a homogeneous and transparent hydrogel precursor solution. A light intensity of 15mW / cm² was used. 2 It was irradiated with ultraviolet light for 8 minutes to completely cure it. The total thickness was the same as in Example 1.

[0118] Comparative Example 2:

[0119] Step S3 is omitted in this comparative example. The same process parameters as in Example 1 are used to directly coat the precursor liquid of the water-locking protective layer 4 onto the graphene conductive layer 2 without any surface treatment, and pre-curing and co-curing are performed. The rest is the same as in Example 1, and will not be repeated here.

[0120] Comparative Example 3:

[0121] Step S6 is omitted in this comparative example. After the hydrogel has cured, the porous polyurethane encapsulation film is no longer applied. The electrodes are directly exposed to the environment.

[0122] Performance testing:

[0123] The electrodes prepared in this embodiment and the comparative example were attached to the chests of volunteers and compared with commercial Ag / AgCl electrodes (3M RedDot). TM Simultaneous dynamic electrocardiogram (ECG) monitoring and comparison were performed. The test included:

[0124] Dynamic performance test: Volunteers walked / ran at a speed of 6km / h for 20 minutes on a treadmill to evaluate the electrodes' anti-interference ability and signal quality (dynamic signal-to-noise ratio) during motion.

[0125] Long-term stability test: Volunteers wore the electrode continuously for 72 hours while engaging in normal daily activities (including office work, walking, sleeping, etc.) to evaluate the long-term impedance stability and moisturizing performance of the electrode.

[0126] Data recording: Throughout the 72-hour test period, volunteers were periodically asked to maintain a seated position for 3 minutes every 12 hours, and the impedance value under this steady-state condition was recorded to calculate impedance fluctuations. In Example 3, the electrodes did not have medical pressure-sensitive adhesive, and dynamic performance was not tested.

[0127] The test results are shown in Table 1 below:

[0128] Table 1. Performance test results of the examples and comparative examples.

[0129]

[0130] The data shows that the performance of Examples 1-3 exhibits a regular change, with the pre-curing energy density increasing from 1200 mJ / cm³. 2 Increased to 1560 mJ / cm 2 The initial impedance increased from 3.8 kΩ to 4.3 kΩ, and the 24-hour impedance fluctuation increased from 9.5% to 13.5%. This trend indicates that increased energy density leads to a higher crosslinking density in the water-locking protective layer 4. While this results in better mechanical strength, it also slightly reduces ion mobility and the number of active groups available for interfacial interpenetration. Example 1 showed the best performance.

[0131] Comparative Example 1, lacking a moisture gradient control mechanism, experienced a significant increase in impedance fluctuation to 28.5% over 24 hours; its internal moisture evaporation dynamics were unidirectional and uniform. Regardless of formulation optimization, moisture continued to evaporate, leading to gradual drying and a steady increase in impedance.

[0132] In Comparative Example 2, the lack of interface treatment resulted in direct failure due to delamination. In Comparative Example 2, the hydrogel and graphene were only bound by weak physical adsorption. This connection is extremely fragile and cannot withstand electrochemical stress, mechanical stress, and hydration stress, leading to interface delamination, interruption of electron transport channels, impedance fluctuations, and ultimately, complete failure.

[0133] Comparative Example 3: No encapsulation; rapid evaporation of moisture caused a 250% impedance fluctuation.

[0134] In this invention, the gradient hydrogel structure is key to solving the long-term stability problem, the covalently bonded interface is key to solving the interface reliability and motion artifact problems, and the encapsulation layer 6 is the fundamental guarantee for the long-term operation of the above two aspects. The combined effect of the above technical features enables continuous high-fidelity and stable acquisition of bioelectrical signals, especially microcurrent signals, in dynamic and long-term monitoring scenarios.

[0135] Those skilled in the art should understand that this invention is not limited to the above embodiments. The embodiments and descriptions in the specification are merely illustrative of the principles of the invention. Various changes and modifications can be made to this invention without departing from its spirit and scope, and all such changes and modifications fall within the scope of the invention as claimed. The scope of protection of this invention is defined by the appended claims and their equivalents.

Claims

1. A graphene hydrogel microcurrent-stabilized wearable bioelectrode, characterized in that, The layers are stacked sequentially from bottom to top: a flexible substrate layer, a conductive layer, an interface modification layer, a water-locking protective layer, and a high water-retention layer. The conductive layer is a graphene-based conductive film covering a flexible substrate layer; The interface modification layer is a silane molecular layer formed on the surface of the conductive layer by sequential oxygen plasma treatment and APTES silanization treatment. The water-locking protective layer is a composite hydrogel pre-cured on the interface modification layer by covalent bonding; The high water retention layer is a composite hydrogel formed by co-curing with the water-locking protective layer; the crosslinking density of the high water retention layer is lower than that of the water-locking protective layer. The electrode is further provided with an encapsulation layer, which is a porous encapsulation film covering the upper surface of the flexible substrate layer and the conductive layer, and has a window in the middle exposing the composite hydrogel; The raw materials of the water-locking protective layer, by weight percentage, include: 15-20 wt% methacrylamide gelatin, 8-10 wt% glycerin, 2-3 wt% betaine, 0.4-0.6 wt% KCl, 0.1-0.15 wt% sodium hyaluronate, 0.1-0.3 wt% LAP photoinitiator, 0.2-0.5 wt% sodium alginate, with the balance being deionized water; The raw materials of the high water retention layer, by weight percentage, include: 15-20 wt% methacrylamide gelatin, 10-12 wt% glycerin, 2-3 wt% betaine, 0.4-0.6 wt% KCl, 0.15-0.2 wt% sodium hyaluronate, 0.1-0.3 wt% LAP photoinitiator, and the balance being deionized water.

2. The graphene hydrogel microcurrent-stabilized wearable bioelectrode according to claim 1, characterized in that, The graphene-based conductive film is a screen-printed reduced graphene oxide film. By mass percentage, the raw materials of the reduced graphene oxide film include: 10-15 wt% rGO powder, 3-5 wt% polyvinyl alcohol, 0.5-1 wt% sodium dodecylbenzenesulfonate, and the balance being deionized water.

3. The graphene hydrogel microcurrent-stabilized wearable bioelectrode according to claim 1, characterized in that, The pre-curing degree of the water-locking protective layer is controlled between 50% and 80%, and the pre-curing degree of the water-locking protective layer is achieved by controlling the energy density of ultraviolet light curing.

4. The graphene hydrogel microcurrent-stabilized wearable bioelectrode according to claim 1, characterized in that, The thickness of the flexible substrate layer is 25-50 μm; the thickness of the conductive layer is 5-15 μm; the thickness of the interface modification layer is <0.1 μm; the thickness of the water-locking protective layer is 40-70 μm; the thickness of the high water retention layer is 40-70 μm; and the thickness of the encapsulation layer is 25-50 μm.

5. A method for preparing a graphene hydrogel microcurrent-stabilized wearable bioelectrode as described in any one of claims 1 to 4, characterized in that, Includes the following steps: S1 provides a flexible substrate layer; S2. Graphene-based conductive paste is printed onto the upper surface of the flexible substrate layer and then thermo-cured to form a conductive layer. S3 performs oxygen plasma treatment on the conductive layer, followed by APTES silanization treatment to form an interface modification layer on its surface. S4 is used to prepare a water-locking protective layer precursor liquid containing sodium alginate, which is then coated onto the interface modification layer and pre-cured under first-intensity ultraviolet light. S5 applies a high water retention layer precursor liquid to the pre-cured water-locking protective layer and then fully cures it under second intensity ultraviolet light to form a hydrogel layer with a gradient structure. S6 covers the flexible substrate and the upper surface of the conductive layer with a porous encapsulation film with a window, and exposes the hydrogel layer from the window to obtain the electrode.

6. The method for preparing a graphene hydrogel microcurrent-stabilized wearable bioelectrode according to claim 5, characterized in that, In step S2, the process parameters for thermosetting are: drying at 120-150℃ for 30-45 minutes.

7. The method for preparing a graphene hydrogel microcurrent-stabilized wearable bioelectrode according to claim 5, characterized in that, In step S3, the oxygen plasma treatment power is 200-300W, and the treatment time is 5-10min. The specific process of the APTES silanization treatment is as follows: the conductive layer treated with oxygen plasma is immersed in an APTES ethanol solution with a pH of 4.5-5.0 and left to stand at room temperature for 1-1.5h. After being taken out, it is vacuum dried at 70-80℃ for 20-30min, then soaked in 75% anhydrous ethanol for 10-15min, and finally rinsed with deionized water and dried.

8. The method for preparing a graphene hydrogel microcurrent-stabilized wearable bioelectrode according to claim 5, characterized in that, In step S4, the process parameters for the first intensity ultraviolet curing are: using ultraviolet light with a wavelength of 365-405nm, and a light intensity of 5-15mW / cm². 2 Irradiate for 1-3 minutes.

9. The method for preparing a graphene hydrogel microcurrent-stabilized wearable bioelectrode according to claim 5, characterized in that, In step S5, the process parameters for the second intensity ultraviolet curing are: using ultraviolet light with a wavelength of 365-405nm, and a light intensity of 10-20mW / cm². 2 Irradiate for 5-10 minutes.

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