Method for preparing a self-powered lung impedance sensor monitoring device based on pyroelectric hydrogel

Abstract

The present application relates to the field of wearable medical electronics, and discloses a preparation method of a self-powered lung impedance sensor monitoring device based on thermoelectric hydrogel, which comprises a flexible substrate, a measurement electrode, a signal acquisition and transmission module, a thermoelectric hydrogel DC generator, an energy management module and a flexible packaging layer; the thermoelectric hydrogel DC generator generates electricity by utilizing the temperature difference between the human body and the environment, thereby providing self-power supply for the device; the core ion thermoelectric hydrogel comprises an interpenetrating network of polyacrylamide and sodium carboxymethyl cellulose, and single-layer graphene oxide and ion salt are introduced, so that the ion thermoelectric hydrogel has both thermoelectric performance and mechanical toughness; the components are integrated into a flexible structure through the flexible packaging layer; and the preparation method of the device is also provided.The self-power supply function of the present application prolongs the endurance time, the integrated flexible structure improves the wearing comfort, suppresses motion artifacts, guarantees the signal quality under dynamic monitoring, and is suitable for long-term and continuous non-invasive respiratory monitoring.

Description

Technical Field

[0001] This invention relates to the field of wearable medical electronics technology, specifically to a method for preparing a self-powered lung impedance sensor monitoring device based on thermoelectric hydrogel. Background Technology

[0002] Lung function monitoring is an important means of assessing respiratory health, and non-invasive monitoring technology based on bioelectrical impedance analysis shows great potential in the field of wearable health devices due to its convenience and safety. Currently, wearable devices for lung impedance monitoring are available on the market and in research. These devices power signal acquisition, processing, and wireless transmission modules via built-in batteries. However, due to limited battery capacity and the high power consumption requirements of continuous monitoring, these devices generally face the bottleneck of insufficient battery life. Users need to frequently remove the device for charging or battery replacement, which not only affects the convenience of use but also interrupts data continuity, making it difficult to meet the needs of long-term, uninterrupted clinical and health management monitoring.

[0003] Secondly, in order to integrate complex electronic components, many existing wearable monitoring devices adopt a rigid-flexible hybrid structure that combines rigid printed circuit boards with flexible electrode patches. This results in an overall rigid device that is difficult to conform to the contours of the human body. Prolonged wear can easily cause local pressure and foreign body sensations, leading to poor comfort. Furthermore, when users perform daily activities such as turning or bending over, the rigid parts will shift relative to the skin, causing the contact interface between the electrodes and the skin to become unstable, thus introducing motion artifacts.

[0004] To address the aforementioned energy issues, researchers have utilized human body heat for energy harvesting. However, traditional thermoelectric materials are mostly inorganic semiconductors, whose inherent rigidity, brittleness, and weight make them unsuitable for the basic requirements of wearable devices for softness, lightness, and stretchability. Although some flexible thermoelectric materials have emerged in recent years, they struggle to balance thermoelectric conversion efficiency with mechanical robustness. Under conditions where the temperature difference between the human body and the environment is only a few degrees Celsius, the power generation efficiency of many flexible materials is too low to meet the practical needs of the circuit. Therefore, this invention proposes a method for fabricating a self-powered lung impedance sensor monitoring device based on thermoelectric hydrogels to overcome the shortcomings of existing technologies. Summary of the Invention

[0005] To address the shortcomings of existing technologies, this invention provides a method for fabricating a self-powered lung impedance sensor monitoring device based on thermoelectric hydrogel, which solves the problems of insufficient battery life, poor wearing comfort, and susceptibility to signal interference during dynamic monitoring in existing wearable respiratory monitoring devices.

[0006] To achieve the above objectives, the present invention provides a self-powered lung impedance sensor monitoring device based on thermoelectric hydrogel, comprising: A flexible substrate has at least one pair of measuring electrodes on its surface, which are electrically connected to a signal acquisition and transmission module. A thermoelectric hydrogel DC generator is used to convert temperature difference into electrical energy. The thermoelectric hydrogel DC generator is electrically connected to an energy management module, which manages the electrical energy. A flexible encapsulation layer is provided on the surface of the energy management module, and a battery is disposed inside the flexible encapsulation layer. The flexible encapsulation layer encapsulates the flexible substrate, the measuring electrodes, the signal acquisition and transmission module, the battery, the thermoelectric hydrogel DC generator, and the energy management module as a whole, forming an integrated flexible structure.

[0007] By adopting the above technical solution, the device of this invention, through the synergistic design of a built-in thermoelectric hydrogel DC generator and an integrated flexible structure, simultaneously solves the core technical challenges of energy endurance and dynamic signal fidelity. Its innovative principle is reflected on two levels: Firstly, the device achieves self-powered operation. The thermoelectric hydrogel DC generator utilizes the minute temperature difference between human skin and the external environment, resulting in ion-selective thermal diffusion within it. The specific process is as follows: Step 1: After the device is worn, a stable temperature gradient is formed at both ends of the thermoelectric hydrogel; Step 2: Temperature difference drives thermal diffusion between lithium cations and chloride anions in the hydrogel network. Due to the inherent difference in the hydrated ionic radius and mobility between the two, the migration rate of cations is higher than that of anions. Step 3: This differentiated ion migration establishes a continuous and stable ion concentration gradient inside the hydrogel, thereby generating a macroscopic internal potential difference, i.e., thermoelectric potential; Step 4: The thermoelectric potential outputs DC power through electrodes integrated with the hydrogel, which is managed by the energy management module and used to charge the device's energy storage unit or directly power it.

[0008] Therefore, the device can continuously convert low-quality human body heat energy into electrical energy, which replenishes the energy consumption during operation and extends its continuous working time.

[0009] Secondly, it ensures signal quality and wearing comfort under dynamic monitoring. The device adopts an integrated flexible structure, allowing all components, including the substrate, circuitry, and encapsulation layer, to conformally fit the surface of human skin as a mechanical whole. When the human body moves, this structure can deform in sync with the skin, thereby maintaining the physical stability of the interface between the measuring electrodes and the skin, suppressing motion artifacts caused by relative displacement of the interface, and ensuring the reliability of acquiring high-quality bioimpedance signals in dynamic scenarios. At the same time, its soft and conforming characteristics also improve the user's long-term wearing comfort. Preferably, the thermoelectric hydrogel DC generator includes an ionic thermoelectric hydrogel, which is polymerized from the following components by weight: 100 parts by weight of acrylamide monomer; 5-40 parts by weight of sodium carboxymethyl cellulose; 0.01-1.0 parts by weight of monolayer graphene oxide; 0.01-0.2 parts by weight of crosslinking agent; 0.1-1.0 parts by weight of initiator; and an ionic salt, wherein the final molar concentration of the ionic salt in the hydrogel system is 1-4M.

[0010] By employing the above-mentioned technical solution, the hydrogel prepared with this specific component formulation exhibits both excellent thermoelectric conversion efficiency and mechanical robustness. The mechanism is as follows: First, by introducing linear sodium carboxymethyl cellulose polymer chains into a chemically cross-linked polyacrylamide network, an interpenetrating polymer network structure is formed. In this structure, the rigid polyacrylamide network provides strength, while the flexible sodium carboxymethyl cellulose chains play a role in energy dissipation and stress transfer. The synergistic effect of these two components endows the hydrogel with excellent toughness, characterized by high tensile strength and high elongation at break, ensuring the structural integrity of the thermoelectric unit under repeated deformation. Second, the introduced monolayer graphene oxide serves as a nano-functional filler. Its surface-rich oxygen-containing functional groups can electrostatically interact with lithium cations in the hydrogel network, constructing ordered ion transport channels. This enhances the selective migration efficiency of ions under temperature gradients, amplifying the Solette effect and thus improving the Seebeck coefficient and output power density of the material.

[0011] Preferably, the crosslinking agent is N,N-methylenebisacrylamide crosslinking agent, and the ionic salt is lithium chloride; the ionic thermoelectric hydrogel is polymerized from raw materials comprising the following components: 100 parts by weight of acrylamide monomer; 10 parts by weight of sodium carboxymethyl cellulose; 0.1 parts by weight of monolayer graphene oxide; 0.05 parts by weight of N,N-methylenebisacrylamide crosslinking agent; 0.5 parts by weight of ammonium persulfate initiator; and lithium chloride, wherein the final molar concentration of lithium chloride in the hydrogel system is 2M.

[0012] By adopting the above technical solution, the preferred ratio is a balance point obtained through a large number of experiments. It can achieve ideal mechanical properties and gelation effect while ensuring excellent thermoelectric performance, and is the best implementation of the present invention.

[0013] Preferably, the flexible substrate is a thermoplastic polyurethane film; the measuring electrode is formed by printing silver nanowires onto the thermoplastic polyurethane film.

[0014] By adopting the above technical solutions, thermoplastic polyurethane film has excellent biocompatibility, elasticity and wear resistance, making it an ideal flexible substrate. The nano-silver wire electrodes formed by printing technology are firmly bonded to the substrate, have stable conductivity, and can be stretched and deformed together with the substrate without easily breaking, ensuring the reliability of electrical connections.

[0015] Preferably, the flexible encapsulation layer is a silicone rubber encapsulation layer; the flexible encapsulation layer forms a cavity in the surface region of the measuring electrode.

[0016] By adopting the above technical solution, medical-grade silicone rubber has excellent biocompatibility, high tensile strength and chemical inertness. As an encapsulation layer, it can effectively protect internal components and ensure wearing safety. The cavity reserved in the electrode area is a necessary structure for signal acquisition, which ensures that the measuring electrode can directly and stably contact the skin.

[0017] Secondly, the present invention provides a method for fabricating a self-powered lung impedance sensor monitoring device based on thermoelectric hydrogel, comprising the following steps: S1: Preparation of ion-thermoelectric hydrogel; S2: Fabrication of a flexible substrate containing measurement electrodes, a signal acquisition and transmission module, and an energy management module; S3: Integrate the ion thermoelectric hydrogel prepared in step S1 with the conductive electrode to form a thermoelectric hydrogel DC generator, and assemble and electrically interconnect the thermoelectric hydrogel DC generator, the flexible substrate containing the measuring electrode, the signal acquisition and transmission module, and the energy management module; S4: Perform flexible encapsulation on the assembled structure from step S3 to form an integrated flexible structure.

[0018] By adopting the above technical solution, the preparation method has systematically designed the process of core material preparation, modular integration and overall flexible packaging. Through clear step division, it can reliably and repeatedly manufacture integrated flexible devices with the aforementioned excellent performance, ensuring product consistency and reliability, and has strong potential for industrial implementation.

[0019] Preferably, in step S1, the step of preparing the ion thermoelectric hydrogel includes: dissolving acrylamide monomer, sodium carboxymethyl cellulose, monolayer graphene oxide, ion salt, crosslinking agent and initiator in water to form a pregel solution, and then subjecting the pregel solution to a polymerization reaction at a temperature of 50℃-80℃ for 2-8 hours.

[0020] By adopting the above technical solution, the process parameters are the key to forming high-performance hydrogels. The temperature and time range ensure that the polymerization reaction can proceed fully, forming a complete and stable interpenetrating network structure, thereby guaranteeing the final mechanical and thermoelectric properties of the material.

[0021] Preferably, in step S2, the step of preparing the flexible substrate containing the measuring electrode includes: using direct ink writing technology to print conductive ink containing silver nanowires onto the surface of the flexible substrate to form the measuring electrode.

[0022] By adopting the above technical solution and using direct ink writing additive manufacturing technology, complex electrode circuits can be directly constructed on flexible substrates with high precision and without masks. It has the advantages of high design freedom, short process flow and high material utilization, and is very suitable for the manufacturing of personalized and integrated flexible electronic devices.

[0023] Preferably, in step S3, the step of performing electrical interconnection includes: using an anisotropic conductive film hot pressing process to connect the flexible lead wire to the thermoelectric hydrogel DC generator, the measuring electrode, the signal acquisition and transmission module, and the energy management module.

[0024] By adopting the above technical solution, the anisotropic conductive film hot pressing process is a mature and reliable flexible circuit connection technology. It can achieve high-density electrical interconnection at low temperature and low pressure, while maintaining the flexibility of the connection parts. This is crucial for maintaining the integrated flexibility of the entire device.

[0025] Preferably, in step S4, the step of flexibly encapsulating the assembled whole in step S3 includes: placing the assembled whole in step S3 in a mold, injecting liquid silicone rubber using a vacuum-assisted casting process, and heating and curing it; through the structural design of the mold, the cured silicone rubber forms a cavity in the surface area of ​​the measuring electrode.

[0026] By adopting the above technical solutions, the vacuum-assisted casting process can effectively eliminate air bubbles during the encapsulation process, ensuring that the encapsulation layer is dense and uniform, thereby improving the reliability of the encapsulation and the durability of the device. The electrode cavity is formed by integral molding using a customized mold structure, which is simpler and has a higher yield compared to subsequent processing methods.

[0027] This invention provides a method for fabricating a self-powered lung impedance sensor monitoring device based on thermoelectric hydrogel. It has the following beneficial effects: 1. This invention achieves the self-powered function of the device by integrating a thermoelectric hydrogel DC generator and a corresponding energy management module. The thermoelectric hydrogel DC generator can continuously convert heat energy into electrical energy by utilizing the small temperature difference between human skin and the external environment, and charge the internal battery of the device through the energy management module, thereby extending the single-charge battery life of the device and getting rid of the limitation of frequent charging required by traditional wearable devices. This provides a reliable energy guarantee for achieving long-term, uninterrupted continuous monitoring of lung function.

[0028] 2. This invention adopts an integrated flexible structure design, using a flexible substrate to support functional units and a flexible encapsulation layer to encapsulate all components as a whole, giving the device mechanical flexibility and wearing comfort. This structure allows the device to conformally fit the surface of human skin and deform in tandem with the skin during body movement, thereby maintaining the stability of the interface between the measuring electrodes and the skin, suppressing motion artifacts from interfering with bioimpedance signals, and ensuring that high-quality, high signal-to-noise ratio physiological signals can still be obtained in dynamic scenarios.

[0029] 3. The ion thermoelectric hydrogel used in this invention has excellent thermoelectric conversion efficiency and mechanical properties. By constructing an interpenetrating network structure of polyacrylamide and sodium carboxymethyl cellulose, the hydrogel achieves high tensile strength and high elongation at break, ensuring the structural integrity of the thermoelectric unit under repeated deformation. At the same time, by introducing monolayer graphene oxide as a functional filler into the network, the selective thermal diffusion effect of ions is synergistically enhanced, enabling the hydrogel to exhibit a high Seebeck coefficient and power density. This high-performance core material is the basis for achieving efficient and reliable self-powered function. Attached Figure Description

[0030] Figure 1 This is a system functional block diagram of the self-powered lung impedance sensor monitoring device of the present invention; Figure 2 This is a schematic cross-sectional view of the self-powered lung impedance sensor monitoring device of the present invention; Figure 3 This is a flowchart illustrating the working principle of the self-powered lung impedance sensor monitoring device of the present invention. Figure 4 This is a graph showing the tensile strength versus relative resistance of the self-powered lung impedance sensor monitoring device of the present invention. Figure 5 This is a first-view wearing image of the self-powered lung impedance sensor monitoring device of the present invention; Figure 6 This is a second-view wearing diagram of the self-powered lung impedance sensor monitoring device of the present invention; Figure 7 This is a flowchart illustrating the fabrication process of the self-powered lung impedance sensor monitoring device of the present invention. Figure 8 This is a schematic diagram of the self-driven power supply function test of the present invention; Figure 9 This is a schematic diagram of the continuous lung impedance signal of the present invention; Figure 10 This is a schematic diagram of the original electrical test of the present invention; Figure 11 This is a schematic diagram of the mechanical curves in Plot7-1 of the present invention; Figure 12 This is a schematic diagram of the mechanical curves in Plot 18-1 of the present invention; Figure 13 This is a schematic diagram of the mechanical curves in Plot9-1 of the present invention; Figure 14 This is a schematic diagram showing the comparison curves of weight retention rates between the packaged sample group and the bare sample group of the present invention. Figure 15 This is a schematic diagram showing the waveform comparison of the output signal changing over time under the normal breathing mode of the present invention.

[0031] Legend 1. Measuring electrode; 2. Signal acquisition and transmission module; 3. Battery; 4. Encapsulation layer; 5. Thermoelectric hydrogel DC generator; 6. Silver nanowire. Detailed Implementation

[0032] The technical solutions in the embodiments of the present invention will be clearly and completely described below with reference to the preparation examples, comparative examples, and test examples. Obviously, the described embodiments are only some embodiments of the present invention, and not all embodiments. Based on the embodiments of the present invention, all other embodiments obtained by those skilled in the art without creative effort are within the scope of protection of the present invention.

[0033] Please see Figure 1 and Figure 2 The self-powered lung impedance sensor monitoring device of the present invention has a core working principle that can be divided into two paths: energy flow and signal flow.

[0034] In the energy flow path: the device generates DC power by using the temperature difference between human skin and the external environment through its internally integrated thermoelectric hydrogel DC generator 5. The DC power is delivered to the energy management module (including the energy harvesting chip). After being processed by the module such as voltage boosting and stabilization, it charges the built-in battery 3, or directly powers the entire system in a specific mode. The battery 3 is an energy storage unit that provides energy security for the long-term stable operation of the device.

[0035] In the signal flow path: at least one pair of measuring electrodes 1 (physically silver nanowires 6 on a flexible substrate) are in contact with human skin. The signal acquisition and transmission module 2 controls them to apply a small, high-frequency excitation current to the human chest cavity. Simultaneously, the measuring electrodes 1 collect voltage signals of changes in chest cavity impedance caused by respiratory activity. After amplification and filtering within the signal acquisition and transmission module 2, the voltage signals are calculated by the microcontroller (MCU) to obtain real-time lung impedance data. Finally, the data is wirelessly transmitted to a terminal device (such as a mobile phone or computer) via Bluetooth (BLE) for user viewing or professional analysis.

[0036] like Figure 2 As shown, these functional modules are physically integrated into a single flexible structure. Specifically, the measuring electrode 1 is printed on a flexible substrate, while the signal acquisition and transmission module 2, energy management module, thermoelectric hydrogel DC generator 5, and battery 3 are all integrated and fixed on the flexible substrate. Finally, all components are encapsulated by a flexible encapsulation layer 4. At the same time, a cavity is reserved on the surface area of ​​the measuring electrode 1 through mold design to ensure that the electrode can make good contact with the skin.

[0037] Please see Figure 3 The device's workflow is shown in more detail. When the device is worn on the human body, the temperature difference between the body surface and the environment is captured by the thermoelectric hydrogel DC generator 5 and converted into electrical energy. After being processed by the energy management circuit, the electrical energy is stored in the battery 3. When the device is working, the battery 3 powers the entire circuit. The signal acquisition module drives the measuring electrode 1 to work and processes and digitizes the acquired bioimpedance signal. The processed data is wirelessly transmitted to the terminal device and finally displayed on the client in the form of lung impedance waveform or respiratory parameters, realizing non-invasive and continuous monitoring of the user's respiratory function.

[0038] Preparation Examples 1-5: Preparation Example 1: This preparation example describes a method for preparing an ion-thermoelectric hydrogel. First, 1.0 g of sodium carboxymethyl cellulose was completely dissolved in 99.0 g of deionized water, and the solution was magnetically stirred to form a homogeneous 1.0 wt% sodium carboxymethyl cellulose solution. Then, 10.0 g of acrylamide monomer was added to this solution, along with a 0.1 wt% aqueous dispersion of monolayer graphene oxide relative to the total mass of the acrylamide monomer. The mixture was placed on a magnetic stirrer and ultrasonically dispersed for 30 minutes to ensure uniform dispersion of the monolayer graphene oxide. Subsequently, sufficient lithium chloride was added to the solution to achieve the desired final molar concentration in the system. The solution was prepared at a concentration of 2M. Then, 0.05 wt% N,N-methylenebisacrylamide crosslinking agent was added relative to the total mass of acrylamide monomers. Stirring continued until the solution became clear and transparent. Finally, 0.5 wt% ammonium persulfate initiator was added relative to the total mass of acrylamide monomers. After rapid and even mixing, the mixture was immediately vacuumed to remove air bubbles. The degassed pregel solution was poured into a mold, sealed, and placed in a 60°C constant temperature oven for polymerization for 6 hours. After polymerization, the solution was removed from the mold to obtain an ion-thermoelectric hydrogel, denoted as HG-1.

[0039] Preparation Example 2: This preparation example aims to illustrate the effect of different monolayer graphene oxide contents on the preparation of hydrogels. The preparation steps are basically the same as those in Preparation Example 1, except that the amount of monolayer graphene oxide added is changed to 0.02 wt% relative to the total mass of acrylamide monomers, and the final hydrogel is denoted as HG-2.

[0040] Preparation Example 3: This preparation example aims to illustrate the effect of different lithium chloride contents on the preparation of hydrogels. The preparation steps are basically the same as those in Preparation Example 1, except that the amount of lithium chloride added is adjusted to achieve a final molar concentration of 4M in the system. The final hydrogel is denoted as HG-3.

[0041] Preparation Example 4: This preparation example aims to illustrate the effect of different sodium carboxymethyl cellulose contents on the preparation of hydrogels. The preparation steps are basically the same as those in Preparation Example 1, except that in the initial step, 3.0 g of sodium carboxymethyl cellulose was dissolved in 97.0 g of deionized water to prepare a 3.0 wt% sodium carboxymethyl cellulose solution. In subsequent steps, the amount of other components added was the same as in Preparation Example 1. The final hydrogel was denoted as HG-4.

[0042] Preparation Example 5: This preparation example aims to illustrate the influence of different polymerization process parameters on the preparation of hydrogels. The preparation steps are basically the same as those in Preparation Example 1, except that the final polymerization and curing step is carried out in a constant temperature oven at 70°C for 3 hours. The final hydrogel is designated as HG-5.

[0043] Examples 1-4: Example 1: This embodiment provides a method for preparing a self-powered lung impedance sensor monitoring device, which utilizes the ion thermoelectric hydrogel prepared in Preparation Example 1, and specifically includes the following steps: Functional module preparation: Medical-grade TPU film was selected as the substrate, and silver nanowires 6 were precisely printed on its surface using direct ink writing (DIW) technology to form a sensor circuit containing 3 pairs of measuring electrodes 1, with the electrode spacing set to 10 mm.

[0044] A flexible thin-film thermistor is fixed to a predetermined position on a TPU substrate using medical-grade cyanoacrylate adhesive to form a temperature-sensing unit.

[0045] Surface mount technology (SMT) is used to mount electronic components such as energy management chips, micro lithium batteries 3, instrumentation amplifiers, microcontrollers (MCUs), and Bluetooth (BLE5) modules onto a flexible PCB board to create a common circuit module and a signal acquisition and transmission module 2.

[0046] Sensor assembly and interconnection: The ion thermoelectric hydrogel (HG-1) prepared in Preparation Example 1 was integrated with conductive carbon cloth by pressing and bonding with conductive adhesive to form a thermoelectric hydrogel DC generator 5.

[0047] Using anisotropic conductive film (ACF) hot pressing process, flexible FPC leads are reliably connected to the leads of the thermoelectric generator, measuring electrode 1, and temperature sensing unit, respectively.

[0048] All the prepared common circuit modules and signal acquisition and transmission module 2 are electrically interconnected according to the circuit design diagram using flexible board connectors and manual micro-soldering, and then fixed to a medical-grade silicone sheet as a structural substrate using biocompatible double-sided tape.

[0049] Overall packaging: The substrate with all components fixed is placed into a custom mold, and two-component liquid silicone rubber is injected through a vacuum-assisted casting process. After heating and curing, an overall encapsulation layer 4 with a thickness of 0.1mm is formed. The measuring electrode area has a cavity left by the mold design so that its surface is not covered by silicone. The pin connection is filled and protected with epoxy resin.

[0050] Example 2: This embodiment provides a method for fabricating a self-powered lung impedance sensor monitoring device using different core materials, specifically including the following steps: The preparation method of this embodiment is basically the same as that of Example 1. The main difference is that the thermoelectric hydrogel DC generator 5 used in the overall assembly and interconnection of the sensor uses the ion thermoelectric hydrogel (HG-3, with a high lithium chloride concentration) prepared in Example 3 as its core material. All other preparation steps, material specifications and process parameters are exactly the same as those in Example 1.

[0051] Example 3: This embodiment provides a method for fabricating a self-powered lung impedance sensor monitoring device with different electrode layouts, specifically including the following steps: The preparation method of this embodiment is basically the same as that of Example 1. The main difference is that in the preparation of the functional module, when preparing the measuring electrode 1 on the TPU film, the electrode spacing of the three pairs of measuring electrodes 1 is set to 5 mm. All other preparation steps, material specifications and process parameters are exactly the same as those in Example 1.

[0052] Example 4: This embodiment provides a method for fabricating a self-powered lung impedance sensor monitoring device with different packaging parameters, specifically including the following steps: The preparation method in this embodiment is basically the same as that in Example 1. The main difference lies in the overall encapsulation process, where the amount of injected liquid silicone rubber is controlled to ensure that the final cured encapsulation layer 4 has a thickness of 0.2 mm. All other preparation steps, material specifications, and process parameters are exactly the same as in Example 1.

[0053] Comparative Examples 1-5: Comparative Example 1: Compared with Example 1, the difference is that the ion thermoelectric hydrogel used to prepare the sensor follows all the steps of Preparation Example 1, but monolayer graphene oxide (GO) is not added when preparing the pregel solution. All other preparation steps, material specifications and process parameters are the same as in Example 1.

[0054] Comparative Example 2: Compared with Example 1, the difference is that the ion thermoelectric hydrogel used to prepare the sensor was prepared without adding sodium carboxymethyl cellulose (sodium carboxymethyl cellulose) when preparing the pregel solution. Instead, acrylamide, lithium chloride, N,N-methylenebisacrylamide, ammonium persulfate and monolayer graphene oxide were directly dissolved in deionized water for polymerization. All other preparation steps, material specifications and process parameters were the same as in Example 1.

[0055] Comparative Example 3: Compared with Example 1, the difference is that the ion thermoelectric hydrogel used to prepare the sensor follows all the steps of Preparation Example 1, but lithium chloride is not added when preparing the pregel solution. All other preparation steps, material specifications and process parameters are the same as in Example 1.

[0056] Comparative Example 4: Compared with Example 1, the difference is that the sensor device does not have a self-powered function. Specifically, the DC generator made of thermoelectric hydrogel and the corresponding energy harvesting chip are removed from the structure of the device. The device relies entirely on its built-in micro lithium battery 3 for power. The battery 3 needs to be charged through a reserved external interface. All other preparation steps, material specifications and process parameters are the same as in Example 1.

[0057] Comparative Example 5: Compared with Example 1, the difference is that the monitoring device adopts a traditional rigid structure. Specifically, the microcontroller, Bluetooth module and other electronic components of the device are all integrated on a rigid printed circuit board. The lung impedance signal is not acquired through an integrated flexible electrode, but through a wire connected to a traditional, disposable silver chloride patch electrode. All other preparation steps, material specifications and process parameters are the same as in Example 1.

[0058] Test Examples 1-7: Test Example 1: Self-Driven Power Supply Function Test Experimental Description: This test aims to verify the self-driving function of the sensor device prepared in Example 1, which generates electrical energy by utilizing temperature difference and charges its internal battery 3 under simulated human wearing conditions.

[0059] The experimental steps are as follows: Take a complete sensor device prepared in Example 1, pre-discharge the built-in lithium battery 3 of the sensor device to 50% of its nominal capacity, and use a high-precision multimeter to measure and record its initial open-circuit voltage.

[0060] The sensor device's preset skin contact surface is tightly fitted to a flat temperature-controlled platform, whose temperature is constantly set at 35.0°C to simulate the surface temperature of human skin.

[0061] The entire test system was placed in a test chamber with an ambient temperature of 25.0℃, thereby establishing a stable temperature gradient of 10.0℃ across the thermoelectric hydrogel unit of the sensor device.

[0062] Turn on the power to the sensor device so that its signal acquisition and wireless transmission modules can enter normal working condition and continuously consume power.

[0063] The test started with a timer. Every 0.5 hours thereafter, the voltage across the battery 3 inside the sensor device was measured and recorded using a high-precision multimeter. The total test duration was 8 hours.

[0064] Experimental data Table 1. Record of battery voltage change in the device of Example 1 under temperature difference drive.

[0065] Conclusion: Table 1 and Figure 8 The test data shows that, under continuous operation, the terminal voltage of the internal battery 3 of the sensor device increased from the initial 3.70V to 3.88V within 8 hours, showing a net increase trend. This result confirms that the total electrical energy generated by the sensor device exceeds the total electrical energy consumed by its working circuit.

[0066] The mechanism of this phenomenon lies in the sensor device's core ionic hydrogel material, which uses polyacrylamide and sodium carboxymethyl cellulose as the network framework, lithium chloride as the ion conductor, and graphene as the functional enhancement filler. This material can convert temperature difference into electrical energy. Under a temperature gradient of 10.0℃, lithium cations and chloride anions in the hydrogel network undergo differential thermal diffusion due to differences in their hydrated ionic radius and mobility. This selective migration of ions establishes a stable ion concentration gradient inside the hydrogel, thereby generating a continuous internal potential difference (thermoelectric potential). The thermoelectric potential outputs DC power through electrodes integrated with the hydrogel, which is managed and boosted by an energy harvesting chip, ultimately charging the built-in lithium battery 3.

[0067] Therefore, the results of this test case verify the feasibility of the technical solution: the prepared sensor device can harvest energy by utilizing the low-amplitude temperature difference between the human body and the environment, realize self-powered power supply, and provide energy guarantee for the long-term continuous monitoring function of the sensor device.

[0068] Test Example 2: Lung Impedance Signal Acquisition Function Test Experimental Description: This test aims to verify the ability of the sensor device prepared in Example 1 to collect and transmit lung bioimpedance signals, and to confirm the correspondence between the signals and different breathing patterns.

[0069] The experimental steps are as follows: Take one of the sensor devices prepared in Example 1 and wear it on a predetermined position on the chest of a healthy adult male volunteer, ensuring good contact between the electrodes and the skin.

[0070] The device establishes a wireless data connection with the terminal device (personal computer) via Bluetooth Low Energy (BLE) and starts the data logging software.

[0071] Instruct volunteers to remain seated and follow the instructions to cycle through the following breathing patterns: 30 seconds of calm breathing, 30 seconds of deep breathing, and 10 seconds of breath-holding at the end of exhalation.

[0072] Throughout the test, the terminal device continuously recorded the lung impedance values ​​transmitted by the sensor device at a sampling rate of 20Hz.

[0073] The raw data was organized into a time series for subsequent analysis.

[0074] Experimental data: Table 2. Selected data from pulmonary impedance monitoring under different breathing modes

[0075] Note: "..." indicates the omission of data.

[0076] Conclusion: Table 2 and Figure 9 The test data shows that the transthoracic impedance value collected by the sensor device changes with the breathing pattern. During the quiet breathing phase, the impedance exhibits small periodic fluctuations; during the deep breathing phase, the amplitude of the impedance fluctuations increases significantly; and during the end-expiratory breath-hold phase, the impedance value remains at a low and relatively stable level. This result indicates that the sensor device can successfully monitor the impedance changes caused by changes in lung air content due to respiratory activity.

[0077] The mechanism for achieving this function stems from the structural design of this solution. First, respiratory movements cause periodic changes in thoracic cavity volume and lung air content. Since the resistivity of air is much higher than that of body fluids and tissues, an increase in lung air content (inhalation) will increase transthoracic impedance, and vice versa (exhalation). Second, the sensor device adopts an integrated flexible structure, combined with medical-grade silicone encapsulation, which allows it to fit closely to the surface of human skin. This stable physical contact reduces the change in electrode-skin interface impedance caused by micro-movements or respiratory fluctuations, thus suppressing motion artifacts and ensuring the baseline stability of signal acquisition. Furthermore, the sensor device highly integrates the measurement electrode array 1 with the signal processing circuit, shortening the signal transmission path, reducing external electromagnetic interference, and improving the signal-to-noise ratio.

[0078] Therefore, the results of this test case verify that the structural design of the sensor device can convert the physiological changes caused by respiration into reliable bioimpedance signals, and confirm the feasibility and effectiveness of this technical solution for non-invasive respiratory function monitoring.

[0079] Test Example 3: Mechanical Flexibility and Electrical Stability Test Experimental description: This test aims to evaluate the electrical performance stability of the internal conductive circuit of the sensor device prepared in Example 1 after repeated mechanical bending deformation.

[0080] The experimental steps are as follows: Take a complete sensor device prepared in Example 1, and use the four-probe method to measure the initial resistance value between any pair of measuring electrodes 1 on the device, denoted as . .

[0081] The sensor device is fixed on the fixture of the bending fatigue tester, and the bending mode is set to reciprocating bending around the axial centerline.

[0082] The radius of curvature for the bending test was set to 1 cm, and the cycle frequency was set to 0.5 Hz.

[0083] Start the tester and perform 1000 consecutive bending and flattening cycles on the sensor device.

[0084] After the cyclic test is completed, remove the sensor device from the fixture. Under the same conditions as in step 1, measure the resistance between the same pair of electrodes again using the four-probe method, and record it as . .

[0085] According to the formula Calculate the rate of change of resistance.

[0086] Experimental data: Table 3. Resistance changes of the device in Example 1 before and after the cyclic bending test

[0087] in conclusion: Figure 4 The test data in Table 3 show that after 1000 bending cycles with a radius of curvature of 1 cm, the resistance of the internal electrode circuit of the sensor device increased by only 2.26%, which is an extremely low level, proving that the sensor device can maintain the stability of its electrical performance when subjected to significant mechanical deformation.

[0088] The stability mechanism stems from the integrated, fully flexible structural design employed in this solution. First, the device uses a highly elastic TPU film as the substrate, which can withstand a wide range of deformation. Second, the conductive circuit is directly printed onto the flexible substrate, while the overall encapsulation uses high-strength medical-grade silicone. This creates a synergistic whole in terms of mechanical performance. When the sensor device bends, stress is dispersed throughout the structure via the flexible encapsulation layer 4 and the substrate, preventing stress concentration at weak points such as rigid-flexible interfaces or wire solder joints. This stress dispersion mechanism reduces the likelihood of microcracks forming in the conductive silver wires due to stretching or compression.

[0089] Therefore, the results of this test case verify that the structural design of this scheme can ensure the reliability of electrical connections and the stability of signal transmission of the device in dynamic application scenarios such as simulating human joint movements.

[0090] Test Example 4: Thermoelectric Performance Comparison Test Experimental Description: This test aims to quantitatively compare the thermoelectric properties of hydrogels with different components prepared in the examples and comparative examples. The core indicators are Seebeck coefficient and maximum output power density.

[0091] The experimental steps are as follows: Hydrogel samples prepared in Preparation Example 1, Comparative Example 1, Comparative Example 2 and Comparative Example 3 were taken respectively, and each sample was cut into standard specimens with dimensions of 20mm×10mm×2mm (length×width×height).

[0092] The test sample was placed on a self-built thermoelectric performance testing platform, which consisted of two Peltier temperature control elements. The sample was sandwiched between two copper plate electrodes, with the copper plates in contact with the upper and lower temperature control elements respectively. A type K thermocouple was used to measure the hot junction temperature by being in close contact with the upper and lower surfaces of the sample. ) and cold end temperature ( ).

[0093] By adjusting the current of the Peltier element, a temperature gradient varying from 2K to 10K is established across the sample. During this process, a high internal resistance voltmeter is used to synchronously record the open-circuit voltage generated by the sample. ), through the - The slope of the linear fit of the data is the Seebeck coefficient of the material. .

[0094] Temperature difference The sample was fixed at 10K and connected in series with a variable resistive load. The voltage across the load was measured while the resistance of the load was adjusted. With the current flowing through Calculate the output power And find its maximum value ( The maximum power density is determined by It is calculated by dividing by the cross-sectional area of ​​the sample.

[0095] Experimental data: Table 4. Comparison of thermoelectric properties of hydrogels in each embodiment and the comparative example.

[0096] Conclusion: Table 4 shows the effects of different components on the thermoelectric properties of the hydrogel. The sample in Example 1 exhibited the highest Seebeck coefficient (24.3 mV / K) and power density (4.17 μW / cm²). 2 ).

[0097] Compared with Example 1, the performance of the sample in Comparative Example 1 (without graphene) was significantly reduced. This confirms that the introduction of graphene is a key factor in improving thermoelectric conversion efficiency. The mechanism is that the oxygen-containing functional groups on the surface of monolayer graphene oxide can interact with lithium cations in the hydrogel network to construct ordered ion transport channels, enhance the selective migration of ions driven by temperature gradient (i.e., the Solette effect), and thus amplify the thermoelectric potential.

[0098] The performance of the sample in Comparative Example 2 also showed a significant decline, indicating that the interpenetrating network structure formed by sodium carboxymethyl cellulose and polyacrylamide is crucial for optimizing thermoelectric performance. The introduction of sodium carboxymethyl cellulose not only improves the mechanical properties of the hydrogel, but the carboxyl functional groups on its chain can also selectively bind cations through electrostatic interaction, further enhancing the mobility difference between cations and anions, thereby producing a gain effect on the Seebeck coefficient.

[0099] The sample in Comparative Example 3 showed almost no thermoelectric effect. This result directly proves the indispensability of lithium chloride as an ion source. The thermoelectric effect of the hydrogel is due to the thermal diffusion of mobile ions. Without lithium ions and chloride ions as charge carriers, the thermoelectric conversion mechanism cannot be established.

[0100] In summary, the data from this test case validates the design principle of this scheme: by constructing a polyacrylamide-carboxymethyl cellulose sodium interpenetrating network as a framework, introducing lithium chloride as an ion source, and utilizing graphene as a functional filler to synergistically enhance ion selective transport, the thermoelectric performance of the hydrogel is ultimately improved.

[0101] Test Example 5: Comparison Test of Mechanical Properties Experimental Description: This test aims to quantitatively evaluate the macroscopic mechanical properties of hydrogels with different components prepared in the examples and comparative examples, mainly examining their tensile strength and elongation at break.

[0102] The experimental steps are as follows: Take the hydrogel samples prepared by Preparation Example 1, Comparative Example 1 and Comparative Example 2 respectively, let each sample stand for 24 hours under standard conditions, and then use a custom mold to cut them into dumbbell-shaped standard test specimens with a gauge length of 10 mm in the test area.

[0103] The dumbbell-shaped specimen is clamped at both ends onto the fixtures of the universal testing machine.

[0104] The tensile rate of the testing machine was set to 100 mm / min. The equipment was started to perform uniaxial tensile testing on the specimen until it completely fractured. The testing machine simultaneously recorded the load and displacement data during the tensile process.

[0105] Based on the recorded load-displacement curves, calculate and extract the tensile strength (maximum stress at fracture) and elongation at fracture (the ratio of the elongation of the gauge length at fracture to the initial gauge length) of each specimen.

[0106] Experimental data: Table 5. Comparison of mechanical property parameters of hydrogels in each example and the comparative example

[0107] in conclusion: Figure 4 The test data in Table 5 show that the sample of Example 1 exhibited the highest tensile strength (1.13 MPa) and elongation at break (857%). In contrast, the mechanical properties of the sample of Comparative Example 2 (without sodium carboxymethyl cellulose) deteriorated, with its tensile strength and elongation at break being much lower than those of Example 1. The mechanical properties of the sample of Comparative Example 1 (without graphene) were slightly lower than those of Example 1, but still remained at a high level.

[0108] The results reveal the decisive role of the interpenetrating polymer network structure in the mechanical properties of the material in this scheme. The mechanism lies in the fact that the polyacrylamide and sodium carboxymethyl cellulose in Example 1 together constitute an interpenetrating polymer network. In this structure, polyacrylamide forms a relatively rigid first layer network, while long-chain sodium carboxymethyl cellulose interpenetrates to form a second flexible network. When the material is subjected to external tensile force, the rigid network first bears the stress, and the breaking of its internal chemical bonds can serve as an energy dissipation mechanism. The flexible network, on the other hand, is responsible for maintaining the continuity of the overall structure and transferring stress to a wider area of ​​the material, thereby delaying the initiation and propagation of macroscopic cracks. This synergistic effect endows the hydrogel with excellent toughness, that is, a combination of high strength and high tensile strength.

[0109] Comparative Example 2, lacking sodium carboxymethyl cellulose, only formed a single polyacrylamide network. This network lacked toughness and was prone to brittle fracture under external force, resulting in the worst mechanical properties. The introduction of graphene (comparative Example 1 and Comparative Example 1) as a nanofiller had a certain reinforcing effect on the polymer matrix, but its contribution to mechanical properties was far less than that of the construction of the interpenetrating network.

[0110] In summary, this test case demonstrates that by constructing a polyacrylamide-sodium carboxymethyl cellulose interpenetrating network, this method can prepare hydrogel materials with both high strength and high flexibility. This mechanical property ensures that the thermoelectric unit in the sensor device can maintain its structural integrity when subjected to repeated stretching, bending and torsion caused by human activities, which is the physical basis for achieving long-term stable operation of wearable devices.

[0111] Test Example 6: Comparison of Wearing Comfort and Dynamic Signal Quality Experimental Description: This test aims to compare the wearing comfort of the fully flexible sensor device prepared in Example 1 with the conventional rigid device in Comparative Example 5 during simulated daily activities and the anti-interference ability of the collected bioimpedance signals.

[0112] The experimental steps are as follows: Select a healthy adult volunteer, first wear the fully flexible sensor device of Example 1 on the volunteer's chest at a preset position, and start signal recording.

[0113] Volunteers were instructed to perform a set of standard dynamic movements, including: marching in place for 1 minute and turning the upper body left and right 10 times. During this process, lung impedance signals were continuously recorded.

[0114] After the dynamic test, the volunteers sat quietly and gave a subjective rating of the comfort of the fully flexible sensor device based on their wearing experience (1-10 points, where 1 point represents extreme discomfort and 10 points represents no foreign body sensation).

[0115] After removing the device from Example 1 and allowing the volunteer's skin to rest for 15 minutes, the rigid device of Comparative Example 5 (connected to silver / silver chloride patch electrodes via wires) was worn in the same position, and the above steps were repeated.

[0116] The impedance signal data recorded from the two tests were analyzed to extract the maximum amplitude of the signal baseline deviation from the resting state during the dynamic action, which was defined as the maximum amplitude of motion artifact.

[0117] Experimental data: Table 6. Performance comparison of Example 1 and Comparative Example 5 in dynamic testing

[0118] in conclusion The test results in Table 6 show that, compared with the rigid device of Comparative Example 5, the fully flexible device of Example 1 achieved a significantly higher subjective comfort score (9.0 points vs 3.5 points), and its motion artifact amplitude in dynamic testing was also significantly lower than that of the former (1.1 ohms vs 17.6 ohms).

[0119] This result directly verifies the superiority of the integrated, fully flexible design of this scheme. The mechanism is reflected in two aspects: Firstly, in terms of wearing comfort, the device in Example 1 uses a thermoplastic polyurethane base and medical-grade silicone for overall encapsulation. The material itself is soft and highly ductile, which can conformally fit the surface of human skin and deform in tandem with body movement, thereby reducing the feeling of foreign objects and local pressure. In contrast, the rigid circuit board in Comparative Example 5 cannot conform to the curvature of the human body, which easily creates pressure concentration points, and the external electrode wires will cause a pulling sensation when moving, resulting in low comfort.

[0120] Secondly, regarding dynamic signal quality, motion artifacts mainly originate from the abrupt changes in interface impedance caused by the relative displacement between the electrode and the skin interface. The integrated design of Example 1 encapsulates the electrode, wires, and circuitry into a flexible whole, which can move with the skin, maintaining high stability of the electrode-skin interface and thus suppressing the generation of motion artifacts. In the structure of Comparative Example 5, the rigid circuit board, wires, and patch electrodes are separate. Body movement can easily cause the electrodes to slide relative to the skin or the wires to sway, resulting in significant electrical noise and interfering with the real physiological signal.

[0121] In summary, this test case demonstrates that the fully flexible integrated structure adopted in this solution not only improves the long-term wearing comfort of wearable devices, but also solves the signal distortion problem of traditional rigid-flexible hybrid structures in dynamic monitoring applications, ensuring the reliability of obtaining high-quality physiological signals in real-world activity scenarios.

[0122] Test Example 7: Comparison of Continuous Working Battery Life Experimental Description: This test aims to quantify and compare the continuous working battery life of the self-powered sensor device of Example 1 and the sensor device of Comparative Example 4 without the self-powered function under simulated wearing conditions.

[0123] The experimental steps are as follows: Take one sensor device from Example 1 and one sensor device from Comparative Example 4. The structure of the sensor device from Comparative Example 4 is the same as that from Example 1, but it does not contain a thermoelectric hydrogel unit. Its energy source is only the built-in battery 3.

[0124] Charge the built-in lithium batteries 3 of both sensor devices to full power and confirm that their initial voltages are the same.

[0125] The two sensor devices were placed simultaneously in the simulated wearing environment described in Test Example 1-1, that is, the skin contact surface of the sensor devices was attached to a temperature control platform at 35.0°C, and the whole device was exposed to a constant temperature environment at 25.0°C.

[0126] Simultaneously turn on the power to both sensor devices, enabling their signal acquisition and wireless transmission modules to enter normal working condition and continuously consume power.

[0127] Continuously monitor and record the total time elapsed from the start of the test until the sensor device automatically shuts down due to the battery voltage falling below the operating threshold.

[0128] Experimental data: Table 7. Comparison of continuous operating battery life between Example 1 and Comparative Example 4

[0129] Conclusion: The data in Table 7 show that the continuous working time of the device in Example 1 (58.7 hours) is more than 2.3 times that of the device in Comparative Example 4 (25.3 hours). This significant difference stems from the core innovative mechanism of this solution: self-driven power supply function.

[0130] With a stable temperature difference of 10K between the two ends of the sensor device, the thermoelectric hydrogel unit inside the device of Example 1 continuously converts thermal energy into electrical energy. This electrical energy is used to recharge the built-in battery 3 of the device via the energy management circuit. Therefore, the battery 3 of the device of Example 1 receives energy input while discharging, and its net energy consumption rate is much lower than that of the device of Comparative Example 4, which only has a unidirectional discharge process.

[0131] The results of this test case verify, from an application perspective, the practical effectiveness of self-powered power supply technology in extending the effective working time of wearable devices after a single charge, providing energy security for long-term, uninterrupted monitoring of physiological parameters.

[0132] Test Example 8: Validation of Gradient Optimization of Conductivity of Ion-Based Thermoelectric Hydrogel Based on Core Group Allocation Ratio Experimental Description: This test was used to verify the specific effects of the bivariate gradient changes in the concentrations of monolayer graphene oxide and lithium chloride on the conductivity of ion-thermoelectric hydrogel samples.

[0133] Experimental steps: According to the preset concentration gradient, samples of the molded ion-thermoelectric hydrogel were taken respectively. The concentration gradient of monolayer graphene oxide was set to 0wt%, 0.05wt%, 0.1wt%, and 0.5wt%, and the concentration gradient of lithium chloride was set to 1M, 1.5M, 2M, and 3M.

[0134] The ion thermoelectric hydrogel samples with different ratios were cut into regular block-shaped samples with a length of 5cm, a width of 1.3cm, and a height of 0.4cm.

[0135] The block sample is clamped between the test electrodes of the electrochemical workstation, and the current flow direction of the system is arranged along the 0.4 cm height of the block sample.

[0136] The scanning voltage range of the electrochemical workstation is set to 1.00V to 2.00V, and the electrochemical workstation system automatically records the current response data at the corresponding voltage.

[0137] According to Ohm's law, the slope of the linear fitting of the current-voltage characteristic curve is extracted to obtain the conductivity value. Using the conductivity value, the working area and thickness of the block sample, the final conductivity of the ion thermoelectric hydrogel sample with the corresponding ratio is calculated.

[0138] Table 8. Conductivity test results of ion thermoelectric hydrogel samples with different component ratios

[0139] Conclusion: According to the data in Table 8, the concentrations of monolayer graphene oxide and lithium chloride significantly affect the conductivity of the ion thermoelectric hydrogel samples. Comparison of the test results of samples T8-1, T8-2, and T8-3 shows that, under the condition of a constant lithium chloride concentration of 1M, the conductivity of the ion thermoelectric hydrogel samples increases as the concentration of monolayer graphene oxide increases from 0 wt% to 0.1 wt%. Comparison of the test results of samples T8-1 and T8-4 ​​confirms that increasing the lithium chloride concentration directly improves the conductivity of the ion thermoelectric hydrogel samples. Lithium chloride dissociates in the interpenetrating network structure composed of polyacrylamide and sodium carboxymethyl cellulose, providing free lithium cations and chloride anions as charge transport carriers. The increased concentration of mobile ions in the system leads to an increase in the macroscopic conductivity value.

[0140] To further quantify the ion transport characteristics under the optimal ratio, the original electrical test spectrum was extracted from the T8-5 sample, which exhibited the highest conductivity. Please refer to [reference needed]. Figure 10 The figure shows the IV Characteristic Curve with Linear Fitting for the T8-5 sample. The horizontal axis represents the applied voltage, the vertical axis represents the detected current, the blue dots represent the actual experimental data, and the orange solid line passing through the blue dots represents the calculated linear fitting curve.

[0141] According to the statistical information in the text box in the figure, within the test voltage range of 1.00 V to 2.00 V, the current range flowing within the ion-thermoelectric hydrogel sample is between 0.012938 A and 0.070579 A. The experimental data points highly overlap with the linear fitting curve, and the calculated coefficient of determination (R²) is... 2 The score was 0.739240, proving that the ion thermoelectric hydrogel sample conforms to a stable Ohm's law transport mechanism. Based on the slope of the linear fitting curve, the conductivity of sample T8-5 reached 0.0508 S, and the final converted conductivity reached 0.3126 S / m, exhibiting the highest peak value among all gradient test groups.

[0142] The internal mechanism for generating peak conductivity lies in the strong electrostatic interaction between the oxygen-containing functional groups on the surface of monolayer graphene oxide and the free lithium cations in the system. This electrostatic interaction constructs continuous and ordered ion transport channels within the three-dimensional polymer network of the ion thermoelectric hydrogel. These ordered ion transport channels reduce the spatial steric hindrance of lithium cations during thermal diffusion and directional migration driven by the electric field, thereby increasing the ion migration rate. The 2M lithium chloride carrier concentration combined with the high-efficiency transport channels constructed by 0.1wt% monolayer graphene oxide together achieves the optimization of macroscopic conductivity performance.

[0143] Comparing the test results of samples T8-5 and T8-6 reveals that when the concentration of monolayer graphene oxide is excessively high, reaching 0.5 wt%, the conductivity of the ion-thermoelectric hydrogel sample decreases in the opposite direction. Excessive monolayer graphene oxide nanosheets are prone to van der Waals aggregation in the pre-gel aqueous solution system. This aggregation of nanofillers disrupts the uniformity of the cross-linking of the ion-thermoelectric hydrogel polymer network, blocking continuous ion transport pathways. Comparing the test results of samples T8-5 and T8-7 indicates that when the lithium chloride concentration is further increased to 3 M, the conductivity of the ion-thermoelectric hydrogel sample does not maintain a continuous upward trend. Excessively high ion concentrations... The concentration of ions triggered the salting-out effect of the polymer chains. The salting-out effect caused the pore size of the polymer three-dimensional network to shrink and deform sharply, which limited the physical mobility range of ions in the free water environment of the hydrogel and increased the resistance to ion transport. The composition ratio parameters of the T8-5 sample achieved the physicochemical balance of the polymer network pore structure, the dispersion state of the nanofunctional filler and the internal carrier concentration. This provided the conditions for the non-destructive acquisition of bioimpedance signals by the signal acquisition and transmission module 2 inside the self-powered lung impedance sensor monitoring device, and for the efficient conversion of energy through a small temperature difference by the thermoelectric hydrogel DC generator 5. It also provided the conditions for the synergistic transport pathway of electrons and ions.

[0144] Test Example 9: Verification of the dynamic mechanical response of hydrogel structures under extreme deformation Experimental Description: This test was used to verify the dynamic stress response characteristics of the ion thermoelectric hydrogel prepared in Example 1 under continuous large-scale tensile deformation conditions, and to evaluate the stress dispersion and energy dissipation mechanism of the interpenetrating polymer network composed of polyacrylamide and sodium carboxymethyl cellulose during the tensile process.

[0145] Experimental steps: The ion thermoelectric hydrogel sample prepared in Example 1 was cut into dumbbell-shaped tensile specimens that met the testing standards using a punching die, and the initial length measurement baseline of the central test area of ​​the dumbbell-shaped tensile specimen was marked.

[0146] Fix both ends of the dumbbell-shaped tensile specimen in the upper and lower clamping fixtures of the high-precision electronic universal testing machine, and adjust the initial distance between the upper and lower clamping fixtures so that the dumbbell-shaped tensile specimen is in a relaxed state of natural straightening and no tension.

[0147] Set the tensile movement rate of the high-precision electronic universal testing machine to a constant 100 mm / min, and start the control program to perform unidirectional continuous tensile operation on the dumbbell-shaped tensile specimen.

[0148] The high-precision electronic universal testing machine has built-in sensors that synchronously collect and record the deformation data and corresponding tensile load data of the dumbbell-shaped tensile specimen throughout the tensile process. The unit of deformation data is set to mm, and the unit of tensile load data is set to N.

[0149] The tensile procedure is continuously executed until the dumbbell-shaped tensile specimen physically fractures. The deformation data and tensile load data of the entire test cycle are extracted and used to construct a numerical table of mechanical response.

[0150] Table 9. Record of Tensile Deformation and Corresponding Tensile Load Data for Thermoelectric Hydrogel Specimens

[0151] Conclusion: According to the data in Table 9, the ion-thermoelectric hydrogel sample prepared in Example 1 exhibited a smooth and non-linear dynamic increase in tensile load during extreme physical deformation, extending from 0.00 mm to 178.42 mm. The original tensile test spectrum automatically generated by the high-precision electronic universal testing machine was extracted to verify the numerical records in Table 9. After comparing the data points, and referring to… Figure 12 The mechanical curve with the Plot18-1 legend label represents the deformation value caused by the displacement of the fixture of the high-precision electronic universal testing machine on the horizontal axis, expressed in mm; the vertical axis represents the tensile load detected by the high-precision electronic universal testing machine in real time, expressed in N. The English Plot18-1 in the legend box belongs to the test program batch number generated by the control software of the high-precision electronic universal testing machine. The mechanical curve in the figure presents the nonlinear stress pattern of the ion thermoelectric hydrogel sample from zero strain to final fracture.

[0152] To verify the innovative mechanical mechanism of the synergistic effect of polyacrylamide, sodium carboxymethyl cellulose, and monolayer graphene oxide, this test case simultaneously extracted the original tensile spectra of the control group lacking specific core components for horizontal comparison, referring to... Figure 11The mechanical curves labeled Plot7-1 correspond to the pure hydrogel test group containing only a single polyacrylamide network. The English Plot7-1 in the legend box also represents the batch number of the control group test procedure. As can be seen from the mechanical curves, the single-network pure hydrogel dropped vertically when the tensile deformation reached only about 10.5 mm and the tensile load increased to 0.74 N, indicating that the single-network pure hydrogel underwent extremely brittle macroscopic fracture. In the initial stage of tensile stress from 0.00 mm to 25.41 mm, the initial upward slope of the mechanical curve can be observed. The tensile load of the ion thermoelectric hydrogel sample prepared in Example 1 increased rapidly to 0.94 N, and the mechanical response showed obvious elastic bearing characteristics. In the elastic bearing stage, the rigid polymer cross-linked network composed of polyacrylamide mainly bears the tensile stress applied externally. The bond angle of the main chain segment of the polyacrylamide molecule changes and spatial extension, giving the ion thermoelectric hydrogel its basic tensile modulus, which is used to resist the physical pull during normal wear.

[0153] Reference Figure 13 The mechanical curves labeled Plot9-1 correspond to the binary hydrogel test group containing a dual network of polyacrylamide and sodium carboxymethyl cellulose but lacking a single layer of graphene oxide. The English "Plot9-1" in the legend box represents the batch number of the binary hydrogel test procedure. Observing the mechanical curves in the figure, it can be seen that the maximum tensile deformation of the binary hydrogel increases to approximately 94.5 mm, and it fractures after a tensile load reaches 2.45 N. The mechanical toughness is significantly improved compared to the single-network pure hydrogel, but it is still far lower than that of the ion-thermoelectric hydrogel sample prepared in Example 1. When the deformation continues to increase into the physical network dissipation stage and high strain slip stage (52.17 mm to 132.08 mm), the corresponding mechanical curve with the Plot18-1 legend shows a gradually slowing upward trajectory in the middle section, as seen in Example 1. The tensile load of the prepared ion thermoelectric hydrogel sample gradually increased from 1.38 N to 2.87 N, and the rate of increase in tensile load decreased significantly. A few fragile covalent bonds inside the rigid cross-linked network of polyacrylamide began to undergo sacrificial breakage, absorbing the destructive mechanical energy input from the outside. Simultaneously, the flexible long chains of sodium carboxymethyl cellulose interspersed inside the rigid cross-linked network also played a role. Since the flexible long chains of sodium carboxymethyl cellulose were not restricted by highly dense chemical cross-linking, they could undergo large-scale conformational adjustment and chain segment slip in the free water environment. The long chain slip behavior conducted and dispersed the locally concentrated mechanical stress to the global space of the three-dimensional network of the ion thermoelectric hydrogel. The sacrificial breakage of polyacrylamide chemical bonds and the physical slip of sodium carboxymethyl cellulose long chains synergistically constructed a dual energy dissipation mechanism.

[0154] When the deformation exceeded 165.75 mm, reaching the extreme macroscopic deformation strengthening state, the tensile load of the ion-thermoelectric hydrogel sample prepared in Example 1 increased to over 3.52 N. Before the vertical fracture and fall at the end of the mechanical curve labeled Plot 18-1 (where the deformation reached approximately 220 mm), the ion-thermoelectric hydrogel sample prepared in Example 1 maintained a continuous physical shape and did not experience sudden macroscopic brittle fracture. Comparison with the premature fracture characteristics of the mechanical curve labeled Plot 9-1 confirms the relationship between the single-layer graphene oxide two-dimensional nanofiller and the polymer matrix network. The presence of non-covalent physical entanglements provides interfacial frictional resistance during extreme deformation. The nanofiller significantly delays the occurrence and polymerization expansion of microstructural cracks. The multi-component composite network structure jointly constructed by polyacrylamide, sodium carboxymethyl cellulose, and monolayer graphene oxide ensures that the thermoelectric hydrogel DC generator 5 encapsulated inside the encapsulation layer 4 does not lose its physical integrity due to fatigue deformation when the self-powered lung impedance sensor monitoring device is closely fitted to the human chest and undergoes large-scale respiratory expansion movements with the torso. This verifies the mechanical feasibility of the self-powered lung impedance sensor monitoring device under long-term working conditions.

[0155] Test Example 10: Verification of Long-Term Water Retention and Packaging Reliability Experimental Description: This test is used to verify the water retention capacity of the flexible packaging structure of the self-powered lung impedance sensor monitoring device prepared in Example 1, as well as the retention rate of the thermoelectric and electrical properties of the self-powered lung impedance sensor monitoring device under long-term storage conditions.

[0156] Prepare the complete self-powered lung impedance sensor monitoring device prepared in Example 1 as the encapsulated sample, and prepare the ion thermoelectric hydrogel with the same internal specifications as in Example 1 but without silicone rubber encapsulation as the bare sample.

[0157] The initial weights of the packaged and bare samples were measured using a high-precision electronic balance. The packaged and bare samples were then connected to an electrochemical workstation and a thermoelectric performance testing platform to measure their initial conductivity and initial Seebeck coefficient.

[0158] The encapsulated samples and the bare samples were placed simultaneously on the test stage of the constant temperature and humidity test chamber. The internal environmental parameters of the constant temperature and humidity test chamber were set, with the operating temperature set to 25℃ and the relative humidity set to 50%RH.

[0159] Turn on the constant temperature and humidity test chamber. When the cumulative storage time reaches 7 days, 15 days, and 30 days, take out the packaged group samples and the bare group samples from the constant temperature and humidity test chamber, weigh them, and record the weight. Based on the ratio of each weighing data to the initial weight, calculate the weight retention rate of the packaged group samples and the bare group samples at different points.

[0160] After weighing the samples at the final point when the storage time reached 30 days, the packaged samples and the bare samples were connected to the electrochemical workstation and thermoelectric performance testing platform again to measure the conductivity and Seebeck coefficient of the packaged samples and the bare samples after 30 days of storage.

[0161] Table 10-1. Weight Retention Rate Test Data

[0162] Table 10-2: Test data of thermoelectric and electrical properties of encapsulated and exposed samples before and after 30 days of storage

[0163] Conclusions and Analysis: According to Table 10-1, Table 10-2 and Figure 14 According to the data, after 30 days in a constant temperature and humidity chamber, the weight retention rate of the packaged samples remained at 97.92%, while the weight retention rate of the exposed samples decreased to 24.15%. After 30 days of environmental exposure, the conductivity of the packaged samples slightly decreased from the initial 0.3126 S / m to 0.3015 S / m, and the Seebeck coefficient changed from 24.3 mV / K to 23.6 mV / K; the conductivity of the exposed samples rapidly dropped to 0.0051 S / m, and the Seebeck coefficient decreased to 1.2 mV / K.

[0164] The underlying mechanism causing the different data performance lies in the direct intervention of the water molecule content in the hydrogel system on the microstructure of the polymer network and the internal ion transport characteristics. The interpenetrating polymer network composed of polyacrylamide and sodium carboxymethyl cellulose requires sufficient water molecules as a dispersion medium and supporting volume. Without a physical barrier layer, the free water inside the exposed sample continuously evaporates and dissipates into the external test environment. The large loss of water causes the interpenetrating polymer network composed of polyacrylamide and sodium carboxymethyl cellulose to shrink in volume and collapse physically. The collapse of the network pores compresses the ion physical transport channels in the system. As the water molecules inside the system decrease sharply, the concentration of lithium chloride exceeds the solubility limit, and lithium chloride crystallizes and precipitates from the three-dimensional polymer network. The number of free lithium cations and chloride anions decreases sharply, the system loses its charge carrier, and the macroscopic conductivity of the exposed sample approaches the insulating state. The ion-selective thermal diffusion effect driven by temperature difference completely stops, and the Seebeck coefficient is basically zero.

[0165] The self-powered lung impedance sensor monitoring device prepared in Example 1 was injected with two-component liquid silicone rubber using a vacuum-assisted casting process, forming a flexible encapsulation layer covering the entire sensor assembly. Silicone rubber is a dense cross-linked elastomer with molecular chain gaps much smaller than the dynamic diameter of water molecules. The flexible encapsulation layer provides mechanical tensile and bending support while constructing a dense isolation barrier that cuts off the evaporation path of water molecules at the physical space level. The slight weight loss of the encapsulated sample is due to a small amount of vapor permeation at the interface of the flexible encapsulation layer, but it did not cause any destructive change to the macroscopic water content of the internal ion thermoelectric hydrogel during the 30-day test period.

[0166] The internal ion thermoelectric hydrogel maintains a stable hydrated liquid phase environment, lithium chloride remains in a free ionic state, and the ordered ion transport channels constructed by the monolayer graphene oxide are intact. The stable micro-liquid phase environment ensures that the self-powered lung impedance sensor monitoring device can continuously perform the tasks of collecting energy from small temperature differences and acquiring bioimpedance signals under normal atmospheric exposure conditions without liquid immersion. This confirms the physical guarantee role of the flexible silicone rubber encapsulation layer design in ensuring the long-term reliability of the self-powered lung impedance sensor monitoring device.

[0167] Test Example 11: Verification of Respiratory Function Consistency with Medical Gold Standard Equipment Experimental Description: This test is used to verify the accuracy of human respiratory rate data obtained by the self-powered lung impedance sensor monitoring device prepared in Example 1. The physiological parameters measured by the self-powered lung impedance sensor monitoring device are compared synchronously with the physiological parameters measured by the clinical medical piezoelectric respiratory chest band.

[0168] Prepare the complete self-powered lung impedance sensor monitoring device prepared in Example 1 and the clinically routine medical piezoelectric breathing chest band. Fix the bottom adhesion interface of the self-powered lung impedance sensor monitoring device to the skin surface of the left front side of the subject's chest. Horizontally wrap the medical piezoelectric breathing chest band around the subject's chest and lock the external straps.

[0169] Connect the lead wires at both ends of the self-powered lung impedance sensor monitoring device to the impedance signal acquisition front-end circuit board. Connect the output terminals of the medical piezoelectric breathing chest band to the multi-channel physiological recorder.

[0170] The test procedure includes three human breathing modes: slow breathing, normal breathing, and rapid breathing. The duration of the test for each breathing mode is set at 3 minutes.

[0171] Subjects were instructed to remain in a supine position. They then sequentially performed slow breathing mode, normal breathing mode, and rapid breathing mode. Simultaneously, the impedance signal acquisition front-end circuit board and the multi-channel physiological recorder were activated to input data.

[0172] After the test, the continuous impedance voltage waveform signal output by the self-powered lung impedance sensor monitoring device and the continuous piezoelectric force waveform signal output by the medical piezoelectric breathing chest band were captured within a 3-minute test cycle under the three breathing modes.

[0173] The number of troughs in the continuous impedance voltage waveform signal and the continuous piezoelectric force waveform signal within 3 minutes were counted separately as the total number of respiratory cycles. The total number of respiratory cycles was divided by time to convert it into a respiratory rate value per minute. The absolute difference between the respiratory rate value measured by the self-powered lung impedance sensor monitoring device and the respiratory rate value measured by the medical piezoelectric breathing chest band was calculated. The absolute difference was divided by the respiratory rate value measured by the medical piezoelectric breathing chest band to obtain the relative error rate.

[0174] Table 11. Comparison of respiratory rates

[0175] Conclusions and Analysis: According to Table 11 and Figure 15 According to the data, the respiratory rate measured by the self-powered lung impedance sensor monitoring device in slow breathing mode was 10.8 bpm, and the respiratory rate measured by the medical piezoelectric breathing chest band was 11.1 bpm, with a relative error of 2.70%. The respiratory rate measured by the self-powered lung impedance sensor monitoring device in normal breathing mode was 18.2 bpm, and the respiratory rate measured by the medical piezoelectric breathing chest band was 17.9 bpm, with a relative error of 1.68%. The respiratory rate measured by the self-powered lung impedance sensor monitoring device in rapid breathing mode was 31.4 bpm, and the respiratory rate measured by the medical piezoelectric breathing chest band was 31.9 bpm, with a relative error of 1.57%.

[0176] Human breathing involves the rhythmic expansion and contraction of the thoracic cavity. When air is inhaled into the alveoli, the volume of air inside the thoracic cavity increases significantly, leading to a decrease in the overall conductivity of the pleural tissues and an increase in local bioelectrical impedance. During exhalation, air is expelled from the alveoli, reducing the volume of air inside the thoracic cavity and decreasing the local bioelectrical impedance. The self-powered lung impedance sensor monitoring device relies on a highly adhesive flexible silicone rubber encapsulation layer to tightly adhere to the surface of the subject's chest skin, eliminating the contact gap between the skin and the self-powered lung impedance sensor monitoring device. Changes in the body's impedance are converted into periodic electrical signals extracted by the electrodes at both ends of the self-powered lung impedance sensor monitoring device.

[0177] Medical piezoelectric breathing chest bands rely on the physical tension generated by the rise and fall of the chest to stretch the internal piezoelectric material and generate signals. The self-powered lung impedance sensor monitoring device directly measures the changes in the electrical properties of the internal tissues. Compared with medical piezoelectric breathing chest bands that rely on the physical deformation of the external chest, the self-powered lung impedance sensor monitoring device has a direct electrical response to changes in lung volume.

[0178] The self-powered lung impedance sensor monitoring device uses an ion thermoelectric hydrogel built in to continuously establish an internal temperature gradient based on the heat from the human body surface, providing a stable driving voltage for the bioimpedance signal extraction circuit. It can complete continuous impedance monitoring tasks without the need for an external power supply. Data confirms that the physiological parameter values ​​extracted by the self-powered lung impedance sensor monitoring device can reach the same monitoring accuracy level as clinical medical piezoelectric respiratory chest bands, meeting the data consistency requirements of medical-grade respiratory monitoring equipment.

Claims

1. A self-powered lung impedance sensor monitoring device based on thermoelectric hydrogel, characterized in that, include: A flexible substrate is provided with at least one pair of measuring electrodes (1) on its surface. The measuring electrodes (1) are electrically connected to a signal acquisition and transmission module (2). A thermoelectric hydrogel DC generator (5) is provided to convert temperature difference into electrical energy. The thermoelectric hydrogel DC generator (5) is electrically connected to an energy management module. The energy management module is used to manage the electrical energy. A flexible encapsulation layer (4) is provided on the surface of the energy management module. A battery (3) is provided inside the flexible encapsulation layer (4). The flexible encapsulation layer (4) encapsulates the flexible substrate, the measuring electrodes (1), the signal acquisition and transmission module (2), the battery (3), the thermoelectric hydrogel DC generator (5), and the energy management module as a whole to form an integrated flexible structure.

2. The self-powered lung impedance sensor monitoring device based on thermoelectric hydrogel according to claim 1, characterized in that, The thermoelectric hydrogel DC generator (5) contains an ionized thermoelectric hydrogel, which is polymerized from the following raw materials in parts by weight: Based on 100 parts by weight of acrylamide monomer; 5-40 parts by weight of sodium carboxymethyl cellulose; 0.01-1.0 parts by weight of monolayer graphene oxide; 0.01-0.2 parts by weight of crosslinking agent; 0.1-1.0 parts by weight of initiator; and ionic salt, wherein the final molar concentration of the ionic salt in the hydrogel system is 1-4M.

3. The self-powered lung impedance sensor monitoring device based on thermoelectric hydrogel according to claim 2, characterized in that, The crosslinking agent is N,N-methylenebisacrylamide crosslinking agent, and the ionic salt is lithium chloride; The ion thermoelectric hydrogel is polymerized from raw materials comprising the following components: 100 parts by weight of acrylamide monomer; 10 parts by weight of sodium carboxymethyl cellulose; 0.1 parts by weight of monolayer graphene oxide; 0.05 parts by weight of N,N-methylenebisacrylamide crosslinking agent; 0.5 parts by weight of ammonium persulfate initiator; and lithium chloride, wherein the final molar concentration of lithium chloride in the hydrogel system is 2M.

4. The self-powered lung impedance sensor monitoring device based on thermoelectric hydrogel according to claim 1, characterized in that, The flexible substrate is a thermoplastic polyurethane film; the measuring electrode (1) is formed by printing silver nanowires (6) onto the thermoplastic polyurethane film.

5. The self-powered lung impedance sensor monitoring device based on thermoelectric hydrogel according to claim 1, characterized in that, The flexible encapsulation layer (4) is a silicone rubber encapsulation layer (4); the flexible encapsulation layer (4) has a cavity formed on the surface area of ​​the measuring electrode (1).

6. A method for fabricating a self-powered lung impedance sensor monitoring device based on thermoelectric hydrogel, characterized in that, Includes the following steps: S1: Preparation of ion-thermoelectric hydrogel; S2: Fabrication of a flexible substrate containing a measurement electrode (1), a signal acquisition and transmission module (2), and an energy management module; S3: Integrate the ion thermoelectric hydrogel prepared in step S1 with the conductive electrode to form a thermoelectric hydrogel DC generator (5), and assemble and electrically interconnect the thermoelectric hydrogel DC generator (5), the flexible substrate containing the measuring electrode (1), the signal acquisition and transmission module (2), and the energy management module. S4: Perform flexible encapsulation on the assembled structure from step S3 to form an integrated flexible structure.

7. The method for preparing the self-powered lung impedance sensor monitoring device based on thermoelectric hydrogel according to claim 6, characterized in that, In step S1, the step of preparing the ion thermoelectric hydrogel includes: dissolving acrylamide monomer, sodium carboxymethyl cellulose, monolayer graphene oxide, ion salt, crosslinking agent and initiator in water to form a pregel solution, and then subjecting the pregel solution to a polymerization reaction at a temperature of 50℃-80℃ for 2-8 hours.

8. The method for preparing the self-powered lung impedance sensor monitoring device based on thermoelectric hydrogel according to claim 6, characterized in that, In step S2, the step of preparing a flexible substrate containing a measuring electrode (1) includes: using direct ink writing technology to print conductive ink containing silver nanowires (6) onto the surface of the flexible substrate to form the measuring electrode (1).

9. The method for preparing the self-powered lung impedance sensor monitoring device based on thermoelectric hydrogel according to claim 6, characterized in that, In step S3, the step of electrical interconnection includes: using anisotropic conductive film hot pressing process to connect the flexible lead wire to the thermoelectric hydrogel DC generator (5), the measuring electrode (1), the signal acquisition and transmission module (2) and the energy management module.

10. The method for preparing the self-powered lung impedance sensor monitoring device based on thermoelectric hydrogel according to claim 6, characterized in that, In step S4, the step of flexibly encapsulating the assembled whole in step S3 includes: placing the assembled whole in step S3 into a mold, injecting liquid silicone rubber using a vacuum-assisted casting process, and then heating and curing it. Through the structural design of the mold, the cured silicone rubber forms a cavity in the surface area of ​​the measuring electrode (1).

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