Preparation method and application of intelligent conductive hydrogel bandage

Abstract

The present application belongs to the technical field of flexible wearable medical electronics and intelligent compression therapy monitoring, and discloses a preparation method and application of an intelligent conductive hydrogel bandage. The intelligent conductive hydrogel bandage comprises an electrode-embedded conductive hydrogel sensing layer and an encapsulation layer. The electrode-embedded conductive hydrogel sensing layer is a sandwich structure of two layers of conductive hydrogel sensing layers sandwiching a flexible conductive fabric electrode, and is encapsulated by the encapsulation layer. By collecting the resistance of the electrode-embedded conductive hydrogel sensing layer during stretching, an electromechanical response relationship between the relative resistance change rate and the pre-stretching strain is established, and in combination with the response relationship between tension and pre-stretching and the modified Laplace model, an interface pressure electromechanical coupling model between the relative resistance change rate, tension and interface pressure is constructed. The present application can realize continuous, real-time and quantitative monitoring of the interface pressure during compression therapy, and has the advantages of good flexible fitting, high structural stability, reliable signal response and wide application range.

Description

Technical Field

[0001] This invention belongs to the field of flexible wearable medical electronics and intelligent compression therapy monitoring technology, and relates to a method for preparing and applying an intelligent conductive hydrogel bandage. Background Technology

[0002] Wound healing is a complex and dynamic process regulated by multiple factors, among which the local mechanical microenvironment plays a crucial role in tissue repair. Appropriate mechanical stimulation can promote the regeneration of damaged tissue and wound repair by regulating processes such as cell migration, cell proliferation, extracellular matrix remodeling, and angiogenesis. Based on this mechanism, elastic bandage compression therapy has been widely used in the care of venous ulcers, lymphedema, and related chronic wounds. By applying external pressure to the tissues surrounding the wound, it improves venous return and local microcirculation, thereby providing favorable conditions for tissue repair. However, the therapeutic effect of compression therapy is highly dependent on the proper control of the interfacial pressure between the bandage and the skin or tissue. Insufficient interfacial pressure makes it difficult to effectively promote venous and lymphatic return, potentially leading to tissue fluid retention and delaying wound healing; excessive interfacial pressure may hinder local blood perfusion, increasing the risk of tissue ischemia, pain, and even necrosis. Therefore, achieving precise control and continuous monitoring of bandage interfacial pressure is of great significance for individualized wound management, improving the safety of compression therapy, and optimizing treatment outcomes.

[0003] However, traditional elastic bandage systems still struggle to accurately quantify interfacial pressure in practice. This is primarily due to the nonlinear coupling between multiple structural parameters (such as material modulus, number of bandage layers, and equivalent radius of curvature), making the bandaging process highly dependent on the experience and perception of healthcare professionals. Furthermore, dynamic changes in human movement and tissue volume cause time-varying pressure fluctuations, making it difficult to maintain a stable, long-term compression state. To address these issues, various pressure monitoring or control schemes have been proposed. For example, color-coded markings, stretching scales, or visual cues are incorporated into bandages or fabrics to indicate the degree of bandage stretching or the approximate pressure range. While these schemes are simple in structure, they typically provide only qualitative or semi-quantitative feedback, making continuous real-time monitoring of interfacial pressure difficult. Other technologies employ micromotors, cable-driven structures, airbag or micropump systems, and shape memory alloy drive components to achieve active pressure application and regulation, combined with flexible sensors for localized pressure detection. While these schemes can achieve dynamic control to some extent, they generally suffer from complex system structures, large size, numerous heterogeneous components, high energy consumption, and insufficient flexibility. Under prolonged wear, mechanical mismatch can easily occur between the device and the skin and soft tissue, affecting fit, comfort, and monitoring stability, thus limiting its widespread application in continuous wound care and compression therapy scenarios.

[0004] With the development of flexible electronics and digital medical technologies, flexible pressure monitoring systems with tissue compliance offer a new technological approach for compression therapy. Among these, hydrogel materials have attracted widespread attention in flexible sensing and wearable medical devices due to their high water content, low modulus, and softness similar to biological tissues. By introducing a conductive network within the hydrogel, detectable changes in electrical signals can be generated during stretching, compression, or bending, thereby enabling the sensing of external mechanical stimuli. Compared to traditional rigid or thin-film sensing materials, conductive hydrogels have potential advantages in skin adhesion, biocompatibility, and dynamic deformation adaptability. However, current research often focuses on the static characterization of intrinsic sensing parameters of the material, rarely exploring the complex interfacial mechanical evolution in actual clinical bandaging scenarios. This restricts the transformation of conductive hydrogels into clinical quantitative monitoring tools. In fact, interfacial pressure is regulated by multiple couplings of bandage tension, equivalent radius of curvature, and interfacial transmission characteristics. Although the classic Laplace's law provides a theoretical benchmark, the actual pressure distribution often deviates from ideal assumptions due to the combined effects of multilayered structures and the nonlinear mechanical behavior of flexible materials. Therefore, current technologies still lack a flexible intelligent bandage system that can maintain a stable electromechanical response under dynamic bandaging conditions and reliably quantitatively map electrical signals to real interfacial pressure. How to construct an intelligent conductive hydrogel bandage that combines tissue compliance, long-term stability, and quantitative pressure resolution capability is a pressing technical problem to be solved in the field of intelligent compression therapy monitoring. Summary of the Invention

[0005] To address the problems of difficulty in real-time quantitative monitoring of interfacial pressure during compression therapy with existing elastic bandages, insufficient coupling between flexible sensors and bandage mechanical models, and poor signal stability under long-term wear, this invention provides an intelligent conductive hydrogel bandage and its interfacial pressure monitoring method, which combines flexible fit, stable sensing, model analysis, and wireless monitoring functions. This technical solution includes at least an electrode-embedded conductive hydrogel sensing layer, a flexible encapsulation layer, an interfacial pressure electromechanical coupling model, and wireless interfacial pressure monitoring. Each component can function as a relatively independent technical feature or work together to form a complete real-time interfacial pressure monitoring system. The intelligent conductive hydrogel bandage adopts a multi-network collaborative conductive hydrogel sensing layer design, using a hydrophilic polymer network as the main structural framework and introducing a dynamic energy dissipation network to improve the material's deformation adaptability and cyclic stability during tensile and compressive deformation processes. Simultaneously, by constructing a continuous conductive network, the conductive hydrogel generates a stable and repeatable electrical response over a wide strain range, providing a reliable signal basis for continuous monitoring and quantitative analysis of interfacial pressure. Unlike traditional flexible sensing methods that rely solely on resistance changes for qualitative deformation detection, this invention combines the electromechanical response of conductive hydrogels with a modified Laplace mechanical model to establish a quantitative mapping relationship between bandage tension, deformation, resistance changes, and interfacial pressure, thereby achieving real-time analysis and continuous monitoring of interfacial pressure. Furthermore, by incorporating a micro-embedded electronic system, the electrical signals from the conductive hydrogel can be acquired, processed, and wirelessly transmitted to a terminal for real-time display and recording of interfacial pressure information. This invention provides a novel technical solution for intelligent pressure management in wound compression therapy and offers a new path for the development of wearable therapeutic smart bandages.

[0006] The technical solution of the present invention: A method for preparing a smart conductive hydrogel bandage, comprising the following steps: Step 1: Preparation of the electrode-embedded conductive hydrogel sensing layer; A basic solvent, hydrophilic functional monomer, crosslinking agent, flexible network modifier, and conductive component are mixed, and a photoinitiator is added to form a conductive hydrogel prepolymer. Subsequently, the conductive hydrogel prepolymer is uniformly coated onto the surface of a glass plate to form a liquid film of a predetermined thickness. After crosslinking by ultraviolet light irradiation, a first conductive hydrogel sensing layer is formed. A flexible conductive fabric electrode is attached to the surface of the first conductive hydrogel sensing layer, and the conductive hydrogel prepolymer is coated again on the flexible conductive fabric electrode. After further crosslinking by ultraviolet light irradiation, a second conductive hydrogel sensing layer is formed. The flexible conductive fabric electrode is sandwiched and fixed between the first and second conductive hydrogel sensing layers to form an electrode-embedded conductive hydrogel sensing layer. Step 2: Flexible encapsulation of the electrode-embedded conductive hydrogel sensing layer. A flexible elastomer material is coated to form an encapsulation layer and pre-cured. Then, the electrode-embedded conductive hydrogel sensing layer prepared in step one is placed between the two encapsulation layers to form a sandwich structure. After removing the internal air, the edges are sealed and cured to obtain a smart conductive hydrogel bandage. Step 3: Construction of the interface pressure electromechanical coupling model; The smart conductive hydrogel bandage prepared in step two was subjected to tensile calibration, and the tension T and the real-time resistance value R of the electrode-embedded conductive hydrogel sensing layer under different pre-tension strains ε were collected simultaneously. t The response relationship between tension T and pre-tension strain ε, and the electromechanical response relationship between the relative resistance change rate y and the pre-tension strain ε were established. Subsequently, a smart conductive hydrogel bandage was wrapped around the surface of a cylindrical model, and the interfacial pressure under different bandaging structure parameters (pre-tension strain, number of bandage layers n, bandage width w, and equivalent radius of curvature r) were obtained. Based on the modified Laplace model, the response relationship between tension T and pre-tension strain ε, the electromechanical response relationship between the relative resistance change rate y and the pre-tension strain ε, and the bandaging structure parameters were coupled to establish a pressure transformation function between the relative resistance change rate y and the interfacial pressure P. This is an interface pressure electromechanical coupling model; Step 4: Wireless interface stress monitoring; A PCB circuit board is fabricated, and an electrode interface, a resistance acquisition circuit, a signal conditioning circuit, an analog-to-digital converter module, a microcontroller unit (MCU), a wireless communication module, and a power supply module are integrated on the PCB circuit board. The intelligent conductive hydrogel bandage prepared in step two is connected to the electrode interface of the PCB circuit board through a flexible conductive fabric electrode, so that the electrode-embedded conductive hydrogel sensing layer is connected to the resistance acquisition circuit. The resistance signal generated by the electrode-embedded conductive hydrogel sensing layer is processed by the resistance acquisition circuit, the signal conditioning circuit, and the analog-to-digital converter module and then transmitted to the microcontroller unit (MCU). During the initialization phase, the microcontroller unit (MCU) records the initial resistance value R0 and adjusts it according to the real-time resistance value R. t Calculate the relative resistance change rate y = (R t -R0) / R0; The interface pressure electromechanical coupling model constructed in step three is written into the microcontroller unit (MCU). The MCU then uses the pressure conversion function... Real-time calculation of interface pressure P, where parameters The interface pressure is written to the internal memory of the microcontroller unit (MCU) via a serial port interface; the calculated interface pressure is transmitted to an external display terminal via a wireless communication module, enabling real-time display, recording, and continuous monitoring of the interface pressure.

[0007] In step one, the base solvent is one of water, a glycerol-water mixture, an ethylene glycol-water mixture, or a glycerol-ethylene glycol-water mixture, wherein the mass fraction of glycerol and / or ethylene glycol in the mixture is 20%-70%. The hydrophilic functional monomer is acrylamide and / or N-isopropylacrylamide, with a content of 20-50 wt% of the base solvent. The conductive component is one or a combination of two or more of poly(3,4-ethylenedioxythiophene):poly(styrene sulfonic acid) (PEDOT:PSS) dispersion, polypyrrole dispersion, and polyaniline dispersion, with a content of 5-30 wt% of the base solvent. The crosslinking agent is N,N'-methylenebisacrylamide and polyethylene glycol diacrylate; wherein the content of N,N'-methylenebisacrylamide (BIS) is 0.05-0.5 wt% of the hydrophilic functional monomer; the content of polyethylene glycol diacrylate (PEGDA) is 0.3-1.5 wt% of the hydrophilic functional monomer; the flexible network modifier is borax, caffeic acid, and polyvinyl alcohol; wherein the content of borax is 0.05-1 wt% of the hydrophilic functional monomer; the content of caffeic acid is 0.05-0.5 wt% of the hydrophilic functional monomer; and the content of polyvinyl alcohol is 1-8 wt% of the base solvent. The photoinitiator is lithium phenyl-2,4,6-trimethylbenzoylphosphonate (LAP) or 2,4,6-trimethylbenzoyl-diphenylphosphine oxide (TPO), the content of which is 0.1-1 wt% of the hydrophilic functional monomer.

[0008] In step one, the thickness of the preset thickness liquid film is 0.2-1.5 mm; the wavelength of the ultraviolet light is 385-405 nm; the irradiation time is 1-10 min; and the flexible conductive fabric electrode is silver fiber fabric, silver-plated fabric, conductive yarn fabric, or carbon fiber fabric.

[0009] In step two, the flexible elastomer material is silicone rubber, polydimethylsiloxane, polyurethane elastomer, thermoplastic elastomer, or flexible silicone rubber (Ecoflex00-10), preferably Ecoflex00-10. The curing temperature is 20-60℃.

[0010] In step three, the tensile calibration is performed by tensile loading or cyclic loading, with a strain range not exceeding 300%, a tensile rate of 100-500 mm / min, and a cycle count of 1-30,000; the real-time resistance value R... t The test frequency is 50-500Hz.

[0011] In step three, the relative resistance change rate y is determined by the real-time resistance value R. t The value is calculated from the initial resistance value R0, and its expression is: y = (R0 / R0) t-R0) / R0, where R0 is the initial resistance value of the electrode-embedded conductive hydrogel sensing layer in the unstretched state or the preset zero-point state. The response relationship between tension T and pre-stretch strain ε is obtained through uniaxial tensile calibration, and its expression is: T=dε 3 +eε 2 +fε+g, where d, e, f, and g are the fitting coefficients for tension T-pre-tension strain ε. The electromechanical response relationship between the relative resistance change rate y and the pre-tension strain ε is obtained by synchronously acquiring the resistance signal and strain of the electrode-embedded conductive hydrogel sensing layer, and its expression is: y=pε 2 +qε+s, where p, q, and s are electromechanical response fitting coefficients. When the intelligent conductive hydrogel bandage is wrapped around the surface of the cylindrical model, the interfacial pressure P satisfies the modified Laplace model P=K(n)T / (wr), where n is the number of bandage layers, w is the bandage width, r is the equivalent radius of curvature, and K(n) is the bandage layer correction factor. The bandage layer correction factor K(n) is obtained through pressure calibration results under different bandage layers. Specifically, under the condition that the pre-tension strain ε, bandage width w, and equivalent radius of curvature r are kept consistent, the n-layer bandage pressure P is calculated based on the single-layer bandage pressure P1. n The ratio of P to the single-layer bandage pressure P1 n / P1, and for P n By fitting the relationship between / P1 and the number of bandage layers n, we obtain K(n) = an 2 +bn+c, where a, b, and c are correction coefficients for the number of bandage layers. T=dε 3 +eε 2 +fε+g、y=pε 2 +qε+s、K(n)=an 2 Combining +bn+c and P=K(n)T / (wr), we obtain the pressure conversion function between the relative resistance change rate y and the interfacial pressure P: P=F(y;n,w,r)=[(an 2 +bn+c) / (wr)]·[dε 3 +eε 2 +fε+g], where ε is given by y=pε 2 The result is obtained by inverse calculation of +qε+s. When p≠0, ε=(-q± In actual calculations, roots falling within the calibrated strain range and satisfying ε≥0 are selected; when p=0 and q≠0, the electromechanical response relationship degenerates into a linear relationship y=qε+s, and ε=(ys) / q is calculated according to the linear fitting relationship. Parameters d, e, f, g, p, q, s, a, b, and c are obtained through experimental calibration and data fitting.

[0012] In step four, the electrode interface includes one or more combinations of alligator clip interfaces, snap-on interfaces, flexible flat cable interfaces, and conductive clamp interfaces. The resistance acquisition circuit includes a voltage divider resistance acquisition circuit, a constant current source resistance acquisition circuit, or a bridge resistance acquisition circuit. The wireless communication module includes a Bluetooth communication module, a Wi-Fi communication module, or other low-power wireless communication modules. The external display terminal includes a computer, mobile phone, or tablet computer, used to receive, display, and record interface pressure data. The power module includes a rechargeable battery and a power management circuit; the power management circuit includes one or more combinations of a charging management circuit, a voltage regulator circuit, an overcurrent protection circuit, and an overvoltage protection circuit.

[0013] A smart conductive hydrogel bandage obtained by the aforementioned preparation method is a sandwich structure formed by sandwiching an electrode-embedded conductive hydrogel sensing layer between two encapsulation layers. The electrode-embedded conductive hydrogel sensing layer is a sandwich structure formed by sandwiching a flexible conductive fabric electrode between a first conductive hydrogel sensing layer and a second conductive hydrogel sensing layer. The first and second conductive gel layers are cross-linked to form a composite conductive hydrogel network composed of a covalently cross-linked polyacrylamide network, a polyvinyl alcohol dynamic physical network, and a conductive polymer network. The thickness of the encapsulation layer is 0.05-2 mm.

[0014] A wireless interface pressure monitoring system for the aforementioned intelligent conductive hydrogel bandage includes an intelligent conductive hydrogel bandage and a PCB circuit board. The PCB circuit board integrates an electrode interface, a resistance acquisition circuit, a signal conditioning circuit, an analog-to-digital conversion module, a microcontroller unit (MCU), a wireless communication module, and a power supply module. The intelligent conductive hydrogel bandage is used to generate a resistance change signal during stretching or wrapping deformation, and is connected to the electrode interface through a flexible conductive fabric electrode. The resistance acquisition circuit is used to acquire the resistance change signal output by the smart conductive hydrogel bandage and convert it into a corresponding analog voltage signal. The signal conditioning circuit is used to buffer, amplify, and / or filter the analog voltage signal to improve signal stability and anti-interference capability. The analog-to-digital conversion module is used to convert the analog voltage signal after signal conditioning circuitry into a digital signal, and transmit it to the microcontroller unit (MCU) via the SPI interface; The microcontroller unit (MCU) is used to receive digital signals, calculate the relative resistance change rate, and calculate the interface pressure based on a preset interface pressure electromechanical coupling model or pressure conversion function. The wireless communication module is used to transmit the interface pressure calculated by the microcontroller unit (MCU) to an external display terminal; The power module is used to supply power to the resistance acquisition circuit, signal conditioning circuit, analog-to-digital conversion module, microcontroller unit (MCU), and wireless communication module.

[0015] A wireless interface pressure monitoring method for the aforementioned smart conductive hydrogel bandage includes the following steps: S1. Wrap the smart conductive hydrogel bandage around the surface of the cylindrical model or the area to be monitored. S2. Different wrapping tensions are obtained by controlling different pre-tension strains, and the real-time resistance value of the pre-tension strain is collected simultaneously. S3. Calculate the relative resistance change rate y based on the real-time resistance value: y=(R t -R0) / R0 Where R0 is the initial resistance value, R t y represents the real-time resistance value, and y represents the relative rate of change of resistance. S4. Establish the electromechanical response relationship between the relative resistance change rate y and the pre-tension strain ε: y=pε 2 +qε+s And establish the response relationship between tension T and pre-tension strain ε: T=dε 3 +eε 2 +fε+g Where p, q, s, d, e, f, g are coefficients obtained through experimental calibration and data fitting; S5. Establish the relationship between interfacial pressure P and tension T based on the modified Laplace model: P=K(n)T / (wr) Where P is the interfacial pressure, T is the tension, w is the bandage width, r is the equivalent radius of curvature, n is the number of bandage layers, and K(n) is the bandage layer correction factor; K(n) = an 2 +bn+c, where a, b, and c are correction coefficients for the number of bandage layers; S6. By combining the electromechanical response relationship, the response relationship, and the modified Laplace model, a quantitative conversion relationship between the relative resistance change rate and the interfacial pressure is obtained: P=F(y;n,w,r)=[(an 2 +bn+c) / (wr)]·[dε 3 +eε 2 +fε+g] Where ε is derived from y=pε 2 +qε+s is obtained by reverse calculation; when p≠0, ε=(-q± ) / (2p), select the root that falls within the calibrated strain range and satisfies ε≥0; when p=0 and q≠0, the electromechanical response relationship degenerates into a linear relationship y=qε+s, and calculate ε=(ys) / q according to the linear fitting relationship.

[0016] The beneficial effects of this invention are: (1) The present invention constructs an intelligent conductive hydrogel bandage. Through the synergistic effect of hydrophilic polymer network, dynamic energy dissipation network and continuous conductive network, the conductive hydrogel sensing layer can still maintain good flexibility, deformation adaptability and conductive stability under stretching, compression and cyclic deformation conditions, which can meet the needs of continuous interface pressure monitoring in complex bandaging environment.

[0017] (2) This invention couples the electromechanical response relationship of the conductive hydrogel sensing layer with the modified Laplace model, establishes a pressure conversion function between the relative resistance change rate and the interface pressure, and realizes real-time calculation and continuous monitoring of interface pressure; compared with the traditional method of relying on experience to wrap, it can improve the quantification and accuracy of interface pressure monitoring.

[0018] (3) The present invention integrates an electrode interface, a resistance acquisition circuit, a signal conditioning circuit, an analog-to-digital conversion circuit, a microcontroller unit (MCU), a wireless communication module, and a power supply module on a PCB circuit board. It can realize the real-time acquisition, processing, and wireless transmission of the resistance signal of the conductive hydrogel sensing layer, and complete the real-time display and recording of the interface pressure data on an external display terminal. It can be applied to fields such as compression therapy, wound care, and flexible health monitoring. Attached Figure Description

[0019] Figure 1 This is a schematic diagram of the structure of the intelligent conductive hydrogel bandage of the present invention.

[0020] Figure 2 This is a tension-strain curve of the intelligent conductive hydrogel bandage of the present invention.

[0021] Figure 3 The graph shows the relationship between the relative resistance change rate and strain of the intelligent conductive hydrogel bandage.

[0022] Figure 4 The graph shows the relationship between the correction factor K(n) for the number of bandage layers and the number of bandage layers.

[0023] Figure 5 The results of long-term interfacial pressure monitoring for the application of intelligent conductive hydrogel bandages on the surface of a cylindrical model.

[0024] Figure 6 The results show the pressure monitoring results of the smart conductive hydrogel bandage applied to the wrist, where k is the slope of the linear fitting curve.

[0025] Figure 7 The results show the pressure monitoring of the calf using a smart conductive hydrogel bandage, where k is the slope of the linear fitting curve.

[0026] In the figure: 1 First encapsulation layer, 2 Electrode-embedded conductive hydrogel sensing layer, 3 Second encapsulation layer. Detailed Implementation

[0027] The specific embodiments of the present invention will be further described below with reference to the accompanying drawings and technical solutions.

[0028] Example 1 A method for preparing a smart conductive hydrogel bandage, comprising the following steps: Step 1: Preparation of the electrode-embedded conductive hydrogel sensing layer. Using a glycerol-water mixture with a glycerol mass fraction of 40g (60wt%) as the base solvent, 14.4g (36wt%) of acrylamide (AM), 21.6mg (0.15wt%) of N,N'-methylenebisacrylamide (BIS), 131.2mg (0.9wt%) of polyethylene glycol diacrylate (PEGDA), 72mg (0.5wt%) of borax, and 21.6mg (0.15wt%) of caffeic acid (CA) were added sequentially. After stirring until homogeneous, 8g (20wt%) of PEDOT:PSS conductive dispersion was added. Subsequently, the system was heated to 90℃, and 1.4g (3.5wt%) of polyvinyl alcohol (PVA) was slowly added until it completely dissolved to form a homogeneous conductive hydrogel prepolymer. After cooling to room temperature, 72 mg (0.5 wt%) of 2,4,6-trimethylbenzoyl diphenylphosphine oxide (TPO) was added as a photoinitiator, followed by degassing treatment. Subsequently, the conductive hydrogel prepolymer was uniformly coated onto the surface of a glass plate to form a liquid film of a predetermined thickness. Crosslinking was then performed under ultraviolet light irradiation to form the first conductive hydrogel sensing layer. A flexible conductive fabric electrode was attached to the surface of the first conductive hydrogel sensing layer, and the conductive hydrogel prepolymer was coated again onto the flexible conductive fabric electrode. After secondary crosslinking under ultraviolet light, a second conductive hydrogel sensing layer was formed. The flexible conductive fabric electrode was then sandwiched and fixed between the first and second conductive hydrogel sensing layers, forming an electrode-embedded conductive hydrogel sensing layer.

[0029] Step 2: Flexible encapsulation of the electrode-embedded conductive hydrogel sensing layer. 20g each of component A and component B of Ecoflex00-10 were weighed and mixed in a 1:1 mass ratio. The mixture was thoroughly stirred and coated onto a polyethylene terephthalate (PET) substrate to form an encapsulation layer with a thickness of 0.6mm. This layer was then pre-cured at 25℃ for 4 hours. The electrode-embedded conductive hydrogel sensing layer prepared in Step 1 was cut to a predetermined size and placed between two encapsulation layers to form a sandwich structure. Ecoflex00-10 was then used to fill and seal the edges of the encapsulation layer, and internal air was expelled by rolling to ensure a tight fit between the electrode-embedded conductive hydrogel sensing layer and the encapsulation layer. Finally, the layer was cured at 25℃ for 4 hours to obtain the intelligent conductive hydrogel bandage. Figure 1 ).

[0030] Step 3: Construction of the interfacial pressure electromechanical coupling model. First, the intelligent conductive hydrogel bandage prepared in Step 2 is subjected to tensile calibration. By controlling different pre-stretch strains ε, the corresponding tension T is obtained, and the real-time resistance value R of the electrode-embedded conductive hydrogel sensing layer is simultaneously acquired. t According to y=(R t The relative resistance change rate is calculated as (-R0) / R0, where R0 is the initial resistance value of the electrode-embedded conductive hydrogel sensing layer in the unstretched state or the preset zero-point state. The response relationship between the tension T and the pre-stretch strain ε of the smart conductive hydrogel bandage is obtained through tensile experiments. Figure 2 ), T=14.76ε 3 -46.64ε 2 +78.24ε-6.54. Further, the electromechanical response relationship between the relative resistance change rate y and the pre-stretch strain ε was established. Figure 3 ), y=0.14ε 2 +1.03ε-0.05. Subsequently, the intelligent conductive hydrogel bandage was wrapped around the surface of the cylindrical model. By adjusting the pre-tension strain, the number of bandage layers n, the bandage width w, and the equivalent radius of curvature r, the interfacial pressure response under different conditions was obtained. The interfacial pressure between the intelligent conductive hydrogel bandage and the cylindrical model satisfies the modified Laplace model, P=K(n)T / (wr), where K(n) is the correction factor for the number of bandage layers. Through experimental fitting, K(n)=0.0268n was obtained. 2 +0.246n+0.728 ( Figure 4 Solving the above equations simultaneously yields the quantitative conversion relationship between the relative resistance change rate y and the interfacial pressure P: P = F(y;n,w,r) = [(0.0268n 2 +0.246n+0.728) / (wr)]·[14.76ε 3 -46.64ε 2+78.24ε-6.54], where ε is given by y=0.14ε 2 The result is obtained by back-calculation using +1.03ε-0.05, i.e., ε=(-1.03+√(0.56y+1.0889)) / 0.28. In actual calculations, the number of bandage layers n, the bandage width w, and the equivalent radius of curvature r are set according to specific test conditions, thereby establishing an electromechanical coupling model of interface pressure for real-time calculation and continuous monitoring of interface pressure.

[0031] Step 4, Wireless Interface Pressure Monitoring. The smart conductive hydrogel bandage prepared in Step 2 is fixed or wrapped around the surface of the area to be monitored, and connected to the electrode interface of the PCB circuit board via a flexible conductive fabric electrode. The resistance change signal acquired by the resistance acquisition circuit is conditioned by the arithmetic signal conditioning circuit and then input to the analog-to-digital converter (ADC). The ADC performs analog signal sampling and digital conversion, and then transmits the signal to the microcontroller unit (MCU) via the SPI interface. During the system initialization phase, the initial resistance value of the electrode-embedded conductive hydrogel sensing layer is recorded. The relative resistance change rate is obtained in real time according to the formula for calculating the relative resistance change rate. The electromechanical coupling model of interface pressure P=F(y;n,w,r) constructed in step three is written into the microcontroller unit (MCU) for real-time interface pressure calculation. Finally, the interface pressure is wirelessly transmitted to a mobile phone or computer via Bluetooth communication module to achieve real-time display, storage, and data recording of the interface pressure. The results show that when the prepared smart conductive hydrogel bandage is fixed to the surface of the cylindrical model for long-term monitoring, the output interface pressure maintains a stable baseline throughout the monitoring process, without significant drift or abrupt change. Figure 5 When the same smart conductive hydrogel bandage was applied to approximately cylindrical body surfaces such as the wrist and calf, no significant signal fluctuations or data loss were observed in the monitoring curves. Figure 5 , Figure 6 The above results demonstrate that the intelligent conductive hydrogel bandage can achieve stable and continuous wireless interface pressure monitoring on the surface of a cylindrical model and on different parts of the human body.

[0032] Example 2 The difference between this embodiment and Embodiment 1 lies in the raw material ratio of the conductive hydrogel prepolymer. A conductive hydrogel prepolymer is formed by mixing hydrophilic functional monomers, flexible network modifiers, conductive components, crosslinking agents, and photoinitiators as shown in Table 1. Subsequently, the conductive hydrogel prepolymer is uniformly coated onto the surface of a glass plate to form a liquid film of a predetermined thickness. Crosslinking is then performed under ultraviolet light irradiation to form a first conductive hydrogel sensing layer. A flexible conductive fabric electrode is attached to the surface of the first conductive hydrogel sensing layer, and the conductive hydrogel prepolymer is coated again onto the flexible conductive fabric electrode. After secondary crosslinking under ultraviolet light, a second conductive hydrogel sensing layer is formed. The flexible conductive fabric electrode is then sandwiched and fixed between the first and second conductive hydrogel sensing layers, forming an electrode-embedded conductive hydrogel sensing layer. Except for the different raw material ratio of the conductive hydrogel prepolymer, the remaining steps are the same as in Embodiment 1.

[0033] Table 1: Raw material composition of conductive hydrogel prepolymer solution

[0034] Example 3 The difference between this embodiment and Embodiment 1 lies in the type of flexible conductive fabric electrode used in step one. The flexible conductive fabric electrode is made of silver fiber fabric, silver-plated fabric, conductive yarn fabric, or carbon fiber fabric. Apart from the different type of flexible conductive fabric electrode, the remaining steps are the same as in Embodiment 1.

[0035] Example 4 The difference between this embodiment and Embodiment 1 lies in the thickness of the encapsulation layer in step two. 20g each of component A and component B of Ecoflex00-10 were weighed and mixed in a 1:1 mass ratio. The mixture was then coated onto the surface of a PET substrate to form encapsulation layers with thicknesses of 0.3mm, 0.5mm, 0.8mm, and 1.5mm, respectively, and pre-cured at 25°C for 4 hours. The electrode-embedded conductive hydrogel sensing layer prepared in step one was cut to a predetermined size and placed between the two encapsulation layers to form a sandwich structure. Subsequently, the edges of the device were filled and sealed with Ecoflex00-10 prepolymer, and internal air was expelled by rolling to ensure a tight fit between the electrode-embedded conductive hydrogel sensing layer and the encapsulation layer. Finally, the mixture was cured at 25°C for 4 hours to obtain the intelligent conductive hydrogel bandage. Except for the different thickness of the flexible encapsulation layer, the remaining steps were the same as in Embodiment 1.

Claims

1. A method for preparing a smart conductive hydrogel bandage, characterized in that, The steps are as follows: Step 1: Preparation of the electrode-embedded conductive hydrogel sensing layer; A basic solvent, hydrophilic functional monomer, crosslinking agent, flexible network modifier, and conductive component are mixed, and a photoinitiator is added to form a conductive hydrogel prepolymer. Subsequently, the conductive hydrogel prepolymer is uniformly coated onto the surface of a glass plate to form a liquid film of a predetermined thickness. After crosslinking by ultraviolet light irradiation, a first conductive hydrogel sensing layer is formed. A flexible conductive fabric electrode is attached to the surface of the first conductive hydrogel sensing layer, and the conductive hydrogel prepolymer is coated again on the flexible conductive fabric electrode. After further crosslinking by ultraviolet light irradiation, a second conductive hydrogel sensing layer is formed. The flexible conductive fabric electrode is sandwiched and fixed between the first and second conductive hydrogel sensing layers to form an electrode-embedded conductive hydrogel sensing layer. Step 2: Flexible encapsulation of the electrode-embedded conductive hydrogel sensing layer. A flexible elastomer material is coated to form an encapsulation layer and then pre-cured. Subsequently, the electrode-embedded conductive hydrogel sensing layer prepared in step one is placed between two encapsulation layers to form a sandwich structure. After removing the internal air, the edges are sealed and cured to obtain the smart conductive hydrogel bandage. Step 3: Construction of the interface pressure electromechanical coupling model; The smart conductive hydrogel bandage prepared in step two was subjected to tensile calibration, and the tension T and the real-time resistance value R of the electrode-embedded conductive hydrogel sensing layer under different pre-tension strains ε were collected simultaneously. t Establish the response relationship between tension T and pre-tension strain ε, and the electromechanical response relationship between relative resistance change rate y and pre-tension strain ε; Subsequently, the intelligent conductive hydrogel bandage was wrapped around the surface of the cylindrical model, and the bandage structure parameters, including pre-tension strain, number of bandage layers n, bandage width w and equivalent radius of curvature r, were adjusted to obtain the interfacial pressure under different bandage structure parameter conditions. Based on the modified Laplace model, the response relationship between tension T and pre-tension strain ε, the electromechanical response relationship between the relative resistance change rate y and the pre-tension strain ε, and the wrapping structure parameters are coupled to establish a pressure conversion function between the relative resistance change rate y and the interfacial pressure P. This is an interface pressure electromechanical coupling model; Step 4: Wireless interface stress monitoring; A PCB circuit board is prepared, and an electrode interface, a resistance acquisition circuit, a signal conditioning circuit, an analog-to-digital conversion module, a microcontroller unit (MCU), a wireless communication module, and a power supply module are integrated on the PCB circuit board. The smart conductive hydrogel bandage prepared in step two is connected to the electrode interface of the PCB circuit board through a flexible conductive fabric electrode, so that the electrode-embedded conductive hydrogel sensing layer is connected to the resistance acquisition circuit. The resistance signal generated by the electrode-embedded conductive hydrogel sensing layer is processed by the resistance acquisition circuit, signal conditioning circuit and analog-to-digital conversion module and then transmitted to the microcontroller unit (MCU). During the initialization phase, the microcontroller unit (MCU) records the initial resistance value R0 and adjusts it according to the real-time resistance value R. t Calculate the relative resistance change rate y = (R t -R0) / R0; The interface pressure electromechanical coupling model constructed in step three is written into the microcontroller unit (MCU). The MCU then uses the pressure conversion function... Real-time calculation of interface pressure P, where parameters Write to the internal memory of the microcontroller unit (MCU) via the serial port interface; The calculated interface pressure is transmitted to an external display terminal via a wireless communication module, enabling real-time display, recording, and continuous monitoring of the interface pressure.

2. The method for preparing the intelligent conductive hydrogel bandage according to claim 1, characterized in that, In step one, the base solvent is one of water, a glycerol-water mixture, an ethylene glycol-water mixture, or a glycerol-ethylene glycol-water mixture, wherein the mass fraction of glycerol and / or ethylene glycol in the mixture is 20%-70%. The hydrophilic functional monomer is acrylamide and / or N-isopropylacrylamide, and its content is 20-50 wt% of the base solvent; The conductive component is one or more of PEDOT:PSS dispersion, polypyrrole dispersion, and polyaniline dispersion, and its content is 5-30 wt% of the base solvent. The crosslinking agent is N,N'-methylenebisacrylamide and polyethylene glycol diacrylate; wherein the content of N,N'-methylenebisacrylamide is 0.05-0.5 wt% of the hydrophilic functional monomer; and the content of polyethylene glycol diacrylate is 0.3-1.5 wt% of the hydrophilic functional monomer. The flexible network modifier comprises borax, caffeic acid, and polyvinyl alcohol; wherein the borax content is 0.05-1 wt% of the hydrophilic functional monomer; the caffeic acid content is 0.05-0.5 wt% of the hydrophilic functional monomer; and the polyvinyl alcohol content is 1-8 wt% of the base solvent. The photoinitiator is lithium phenyl-2,4,6-trimethylbenzoylphosphonate or 2,4,6-trimethylbenzoyl-diphenylphosphine oxide, and its content is 0.1-1 wt% of the hydrophilic functional monomer; The thickness of the preset thickness liquid film is 0.2-1.5 mm; The wavelength of the ultraviolet light is 385-405nm, and the irradiation time is 1-10min; The flexible conductive fabric electrode is made of silver fiber fabric, silver-plated fabric, conductive yarn fabric, or carbon fiber fabric.

3. The method for preparing the intelligent conductive hydrogel bandage according to claim 1, characterized in that, In step two, the flexible elastomer material is silicone rubber, polydimethylsiloxane, polyurethane elastomer, thermoplastic elastomer, or Ecoflex00-10. The curing temperature is 20-60℃.

4. The method for preparing the intelligent conductive hydrogel bandage according to claim 1, characterized in that, In step three, the tensile calibration is either tensile loading or cyclic loading, with a strain range of no more than 300%, a tensile rate of 100-500 mm / min, and a cycle count of 1-30,000. The real-time resistance value R t The test frequency is 50-500Hz.

5. The method for preparing the intelligent conductive hydrogel bandage according to claim 1, characterized in that, In step three, the relative resistance change rate y is determined by the real-time resistance value R. t The value is calculated from the initial resistance value R0, and its expression is: y = (R0 / R0) t -R0) / R0, where R0 is the initial resistance value of the electrode-embedded conductive hydrogel sensing layer in the unstretched state or the preset zero-point state. The response relationship between tension T and pre-tension strain ε is obtained through uniaxial tensile calibration, and its expression is: T = dε 3 +eε 2 +fε+g, where d, e, f and g are the fitting coefficients of tension T-pre-tension strain ε; The electromechanical response relationship between the relative resistance change rate y and the pre-stretch strain ε is obtained by synchronously acquiring the resistance signal and strain of the electrode-embedded conductive hydrogel sensing layer, and its expression is: y = pε 2 +qε+s, where p, q and s are the electromechanical response fitting coefficients.

6. The method for preparing the intelligent conductive hydrogel bandage according to claim 1, characterized in that, In step three, when the smart conductive hydrogel bandage is wrapped around the surface of the cylindrical model, the interfacial pressure P satisfies the modified Laplace model P=K(n)T / (wr), where n is the number of bandage layers, w is the bandage width, r is the equivalent radius of curvature, and K(n) is the bandage layer correction factor. The bandage layer correction factor K(n) is obtained through pressure calibration results under different bandage layers: under the condition that the pre-tensile strain ε, bandage width w, and equivalent radius of curvature r remain consistent, the pressure P of n layers is calculated based on the single-layer bandage pressure P1. n The ratio of P to the single-layer bandage pressure P1 n / P1, and for P n By fitting the relationship between / P1 and the number of bandage layers n, we obtain K(n) = an 2 +bn+c, where a, b, and c are correction coefficients for the number of bandage layers; T=dε 3 +eε 2 +fε+g、y=pε 2 +qε+s、K(n)=an 2 Combining +bn+c and P=K(n)T / (wr), we obtain the pressure conversion function between the relative resistance change rate y and the interfacial pressure P: P=F(y;n,w,r)=[(an 2 +bn+c) / (wr)]·[dε 3 +eε 2 +fε+g], where ε is given by y=pε 2 +qε+s is obtained by inverse calculation; When p≠0, ε=(-q± ) / (2p), select the root that falls within the calibrated strain range and satisfies ε≥0; when p=0 and q≠0, the electromechanical response relationship degenerates into a linear relationship y=qε+s, and calculate ε=(ys) / q according to the linear fitting relationship; d, e, f, g, p, q, s, a, b and c are obtained through experimental calibration and data fitting.

7. The method for preparing the intelligent conductive hydrogel bandage according to claim 1, characterized in that, In step four, the electrode interface includes one or more combinations of alligator clip interface, snap-on interface, flexible flat cable interface, and conductive clamp interface; The resistance acquisition circuit includes a voltage divider resistance acquisition circuit, a constant current source resistance acquisition circuit, or a bridge resistance acquisition circuit. The wireless communication module includes a Bluetooth communication module, a Wi-Fi communication module, or other low-power wireless communication modules; The external display terminal includes a computer, mobile phone, or tablet computer; The power module includes a rechargeable battery and a power management circuit. The power management circuit includes one or more of the following: a charging management circuit, a voltage regulator circuit, an overcurrent protection circuit, and an overvoltage protection circuit.

8. A smart conductive hydrogel bandage obtained by the preparation method according to any one of claims 1-7, characterized in that, The intelligent conductive hydrogel bandage is a sandwich structure formed by two encapsulation layers with an electrode-embedded conductive hydrogel sensing layer sandwiched between them. The electrode-embedded conductive hydrogel sensing layer is a sandwich structure formed by a flexible conductive fabric electrode sandwiched between a first conductive hydrogel sensing layer and a second conductive hydrogel sensing layer. The first and second conductive gel layers are cross-linked to form a composite conductive hydrogel network composed of a covalently cross-linked polyacrylamide network, a polyvinyl alcohol dynamic physical network, and a conductive polymer network. The thickness of the encapsulation layer is 0.05-2 mm.

9. A wireless interface pressure monitoring system for the intelligent conductive hydrogel bandage as described in claim 8, characterized in that, The wireless interface pressure monitoring system includes an intelligent conductive hydrogel bandage and a PCB circuit board. The PCB circuit board integrates an electrode interface, a resistance acquisition circuit, a signal conditioning circuit, an analog-to-digital conversion module, a microcontroller unit (MCU), a wireless communication module, and a power supply module. The intelligent conductive hydrogel bandage is used to generate a resistance change signal during stretching or wrapping deformation, and is connected to the electrode interface through a flexible conductive fabric electrode. The resistance acquisition circuit is used to acquire the resistance change signal output by the smart conductive hydrogel bandage and convert it into a corresponding analog voltage signal. The signal conditioning circuit is used to buffer, amplify, and / or filter the analog voltage signal to improve signal stability and anti-interference capability. The analog-to-digital conversion module is used to convert the analog voltage signal after signal conditioning circuitry into a digital signal, and transmit it to the microcontroller unit (MCU) via the SPI interface; The microcontroller unit (MCU) is used to receive digital signals, calculate the relative resistance change rate, and calculate the interface pressure based on a preset interface pressure electromechanical coupling model or pressure conversion function. The wireless communication module is used to transmit the interface pressure calculated by the microcontroller unit (MCU) to an external display terminal; The power module is used to supply power to the resistance acquisition circuit, signal conditioning circuit, analog-to-digital conversion module, microcontroller unit (MCU), and wireless communication module.

10. A wireless interface pressure monitoring method for the smart conductive hydrogel bandage of claim 8, characterized in that, Includes the following steps: S1. Wrap the smart conductive hydrogel bandage around the surface of the cylindrical model or the area to be monitored. S2. Different wrapping tensions are obtained by controlling different pre-tension strains, and the real-time resistance value of the pre-tension strain is collected simultaneously. S3. Calculate the relative resistance change rate y based on the real-time resistance value: y=(R t -R0) / R0 Where R0 is the initial resistance value, R t y represents the real-time resistance value, and y represents the relative rate of change of resistance. S4. Establish the electromechanical response relationship between the relative resistance change rate y and the pre-tension strain ε: y=pε 2 +qε+s And establish the response relationship between tension T and pre-tension strain ε: T=dε 3 +eε 2 +fε+g Where p, q, s, d, e, f, g are coefficients obtained through experimental calibration and data fitting; S5. Establish the relationship between interfacial pressure P and tension T based on the modified Laplace model: P=K(n)T / (wr) Where P is the interfacial pressure, T is the tension, w is the bandage width, r is the equivalent radius of curvature, n is the number of bandage layers, and K(n) is the bandage layer correction factor; K(n) = an 2 +bn+c, where a, b, and c are correction coefficients for the number of bandage layers; S6. By combining the electromechanical response relationship, the response relationship, and the modified Laplace model, a quantitative conversion relationship between the relative resistance change rate and the interfacial pressure is obtained: P=F(y;n,w,r)=[(an 2 +bn+c) / (wr)]·[dε 3 +eε 2 +fε+g] Where ε is derived from y=pε 2 +qε+s is obtained by reverse calculation; when p≠0, ε=(-q± ) / (2p), select the root that falls within the calibrated strain range and satisfies ε≥0; when p=0 and q≠0, the electromechanical response relationship degenerates into a linear relationship y=qε+s, and calculate ε=(ys) / q according to the linear fitting relationship.

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