Glucose, Cortisol, Sodium Ion and pH Detection Sensors and Methods of Preparation

Abstract

The application discloses a glucose, cortisol, sodium ion and pH detection sensor and a preparation method, belongs to the technical field of sensing and detection, and comprises a substrate layer, a detection module arranged above the substrate layer, the detection module comprising a glucose detection module, a cortisol detection module, a sodium ion detection module and a pH detection module, and a bionic microfluidic unit and a bionic sweat guiding fabric layer.The glucose detection module, the cortisol detection module, the sodium ion detection module and the pH detection module are used to realize multi-parameter synchronous real-time detection of glucose, cortisol, sodium ion concentration and pH in sweat; meanwhile, the bionic sweat guiding fabric layer is used to construct a wettability gradient between a hydrophilic side and a hydrophobic side, realize one-way wetting transmission of sweat from a skin side to a detection side, effectively improve the collection efficiency of sweat, inhibit the back seepage of sweat, and improve the accuracy of glucose, cortisol, sodium ion concentration and pH detection in sweat.

Description

Technical Field

[0001] This invention belongs to the field of sensor technology, specifically relating to glucose, cortisol, sodium ion and pH detection sensors and their preparation methods. Background Technology

[0002] Sweat contains a variety of biochemical indicators that reflect the body's physiological state. Among them, glucose characterizes the body's energy metabolism, cortisol characterizes the body's stress level, sodium ions are closely related to the body's water-electrolyte balance and neuromuscular function regulation, and pH reflects the acid-base environment of sweat and local metabolic status. Therefore, the combined detection of glucose, cortisol, sodium ions, and pH in sweat is of great significance for assessing the body's physiological state.

[0003] In existing technologies, the detection of glucose, cortisol, sodium ions, and pH largely relies on instruments such as biochemical analyzers, enzyme-linked immunosorbent assay (ELISA) devices, phosphoric acid analyzers, pH meters, or electrochemical workstations. These detection methods typically require offline sampling and suffer from problems such as large equipment size, complex operating procedures, high sample pretreatment requirements, and difficulty in achieving real-time continuous detection. Furthermore, existing detection equipment is mostly designed for single indicators, making it difficult to achieve integrated detection of multiple indicators on a single flexible platform, thus failing to meet the needs of wearable applications.

[0004] During an athlete's training, the body's sweat secretion rate, stress level, electrolyte loss, and metabolic environment are all in a dynamic state of flux. Traditional offline monitoring methods struggle to reflect real-time physiological information during training and also fail to adequately address wearing comfort, flexible fit, and resistance to sweat interference. Summary of the Invention

[0005] In response to one or more of the above-mentioned defects or improvement needs of the prior art, the present invention provides a glucose, cortisol, sodium ion and pH detection sensor to solve the problem that existing detection equipment cannot meet the real-time monitoring of sweat secretion rate, stress level, electrolyte loss and metabolic environment during athletes' training.

[0006] To achieve the above objectives, the present invention provides a glucose, cortisol, sodium ion, and pH detection sensor, comprising: basal layer; A detection module is disposed above the substrate layer; the detection module includes a glucose detection module, a cortisol detection module, a sodium ion detection module, and a pH detection module sequentially arranged on the substrate layer. A biomimetic microfluidic unit is provided, which covers the glucose detection module, cortisol detection module, sodium ion detection module and pH detection module. The biomimetic microfluidic unit has multiple flow channels arranged vertically, and each flow channel is provided with a biomimetic capillary pump drive structure to guide the liquid to be tested to each detection module of the detection module. A biomimetic sweat-wicking fabric layer is disposed on the side of the biomimetic microfluidic unit away from the detection module. The biomimetic sweat-wicking fabric layer is a double-sided fabric with unidirectional moisture-wicking properties. The side of the biomimetic sweat-wicking fabric layer away from the biomimetic microfluidic unit is a hydrophobic back-seepage barrier layer, and the side of the biomimetic sweat-wicking fabric layer close to the biomimetic microfluidic unit is a hydrophilic moisture-wicking layer. There is a flow channel between the hydrophilic moisture-wicking layer and the hydrophobic back-seepage barrier layer. The flow channel contains hydrophilic groups, and the hydrophobic back-seepage barrier layer to the hydrophilic moisture-wicking layer forms a gradient structure with increasing wettability.

[0007] As a further improvement of the present invention, the hydrophobic backflow barrier layer comprises a coating layer made of fluoropolymer or siloxane material, and the surface contact angle of the hydrophobic backflow barrier layer is not less than 110°. The hydrophilic moisture-wicking layer includes a surface-modified fiber substrate or a functional layer with hydrophilic functional groups, and the surface contact angle of the hydrophilic moisture-wicking layer is not greater than 30°.

[0008] As a further improvement of the present invention, the pH detection module is coupled to the glucose detection module; The pH detection module is used to acquire the acidity or alkalinity information of the liquid to be tested, so as to calibrate or compensate the detection signal of the glucose detection module. The glucose detection module integrates an enzyme activity dynamic compensation circuit based on the feedback from the pH detection module, which is used to receive the acidity or alkalinity information detected by the pH detection module and adjust the glucose concentration detection parameters accordingly.

[0009] As a further improvement of the present invention, the glucose detection module includes: A gold electrode, wherein a Prussian blue electrocatalytic layer and an enzyme immobilization functional layer are sequentially disposed on the gold electrode; The Prussian blue electrocatalytic layer was formed by cyclic voltammetry deposition. The enzyme immobilization functional layer is a composite membrane of chitosan, carbon nanotubes, and glucose oxidase.

[0010] As a further improvement of the present invention, the glucose detection module includes: A gold nanostructured electrode, wherein the surface of the gold nanostructured electrode is provided with a Prussian blue electrocatalytic layer, a nickel hexacyanide iron layer and an enzyme immobilization layer; The enzyme immobilization layer is formed by using a porous membrane system of alumina nanoparticles and agarose or chitosan matrix material, which is then mixed with glucose oxidase solution and drop-coated onto the surface of gold nanostructure electrode.

[0011] As a further improvement of the present invention, the pH detection module includes: A gold nanostructure electrode, wherein a conductive polymer sensitive layer is disposed on the surface of the gold nanostructure electrode, and the conductive polymer sensitive layer is a polyaniline film formed by electrodeposition.

[0012] As a further improvement of the present invention, the pH detection module includes: A metal oxide modified electrode, wherein the electrode surface has a hydrated metal oxide film, and the hydrated metal oxide film is a hydrated iridium oxide film.

[0013] As a further improvement of the present invention, both the glucose detection module and the cortisol detection module are provided with a counter electrode, and the glucose detection module, the cortisol detection module, the sodium ion detection module and the pH detection module share the same common reference electrode; The common reference electrode is a multilayer composite structure, which includes a silver nano-conductive layer formed by inkjet printing, a silver chloride layer formed by electrochemical chlorination, and an ion-selective permeation membrane covering the silver chloride layer.

[0014] As a further improvement of the present invention, the silver chloride layer has a porous structure with a porosity of 40-60%. The ion-selective permeation membrane is a Nafion membrane with a thickness of 2~5μm.

[0015] The present invention also includes a method for preparing a glucose, cortisol, sodium ion, and pH detection sensor, which includes the following steps: S1. Cut and clean the substrate material to obtain the substrate layer; S2. Working electrodes and counter electrodes for glucose detection module, cortisol detection module, sodium ion detection module and pH detection module are formed on the substrate by electron beam evaporation or sputtering, respectively, and a common reference electrode is formed by microelectronic printing. S3. Modify the working electrode of the glucose detection module; S4. Modify the working electrode of the cortisol detection module; S5. Modify the working electrode of the sodium ion detection module; S6. Modify the working electrode of the pH detection module; S7. A biomimetic microfluidic unit is formed above the glucose detection module, cortisol detection module, sodium ion detection module and pH detection module, and a biomimetic sweat-guiding fabric layer is formed on the surface of the biomimetic microfluidic unit.

[0016] The aforementioned improved technical features can be combined with each other as long as they do not conflict with each other.

[0017] In summary, the beneficial effects of the above-described technical solutions conceived by this invention compared with the prior art include: (1) The glucose, cortisol, sodium ion and pH detection sensor of the present invention realizes the synchronous real-time detection of multiple parameters such as glucose, cortisol, sodium ion concentration and pH in sweat through glucose detection module, cortisol detection module, sodium ion detection module and pH detection module set on the base layer; at the same time, the present invention sets a biomimetic sweat guiding fabric layer, and constructs a wettability gradient between the hydrophilic side and the hydrophobic side through the biomimetic sweat guiding fabric layer, realizes the unidirectional moisture guiding transmission of sweat from the skin side to the detection side, which can effectively improve the sweat collection efficiency, inhibit sweat back seepage, and improve the accuracy of glucose, cortisol, sodium ion concentration and pH detection in sweat. Attached Figure Description

[0018] Figure 1 This is a schematic diagram of the preparation process of glucose, cortisol, sodium ion and pH detection sensors in an embodiment of the present invention; Figure 2 This is a schematic diagram of the biomimetic microfluidic unit structure of the glucose, cortisol, sodium ion and pH detection sensors in the embodiments of the present invention; Figure 3 This is a schematic diagram of the electrode system structure of the glucose, cortisol, sodium ion and pH detection sensors in an embodiment of the present invention; Figure 4 This is a graph showing the open-circuit potential test results of the sodium ion detection module for different concentrations of test liquids in this embodiment of the invention. Figure 5 This is a current response diagram of the cortisol detection module in this embodiment of the invention, which uses the chronoamperometry method to test the current response of the test liquid at different concentrations. Figure 6 This is a current response diagram of the glucose detection module in this embodiment of the invention using the chronoamperometry method for different concentrations of test liquid; Figure 7 This is a graph showing the electrode potential response of the pH detection module in this embodiment of the invention to buffer solutions of different pH values ​​using the open-circuit potential method. Detailed Implementation

[0019] To make the objectives, technical solutions, and advantages of this invention clearer, the invention will be further described in detail below with reference to the accompanying drawings and embodiments. It should be understood that the specific embodiments described herein are merely illustrative and not intended to limit the invention. Furthermore, the technical features involved in the various embodiments of this invention described below can be combined with each other as long as they do not conflict with each other.

[0020] In the description of this invention, it should be understood that, unless otherwise stated, the terms "center," "longitudinal," "lateral," "length," "width," "thickness," "upper," "lower," "front," "rear," "left," "right," "vertical," "horizontal," "top," "bottom," "inner," "outer," "clockwise," "counterclockwise," "axial," "radial," and "circumferential" indicate the orientation or positional relationship based on the orientation or positional relationship shown in the accompanying drawings. They are used only for the convenience of describing this invention and simplifying the description, and are not intended to indicate or imply that the device or element referred to must have a specific orientation, or be constructed and operated in a specific orientation. Therefore, they should not be construed as limiting this invention.

[0021] Furthermore, unless otherwise stated, the terms "first" and "second" are used for descriptive purposes only and should not be construed as indicating or implying relative importance or implicitly specifying the number of indicated technical features. Thus, a feature defined as "first" or "second" may explicitly or implicitly include at least one of that feature. In the description of this invention, "a plurality of" means at least two, such as two, three, etc., unless otherwise explicitly specified.

[0022] In this invention, unless otherwise explicitly specified and limited, the terms "installation," "connection," "linking," and "fixing," etc., should be interpreted broadly. For example, they can refer to a fixed connection, a detachable connection, or an integral part; they can refer to a mechanical connection or an electrical connection; they can refer to a direct connection or an indirect connection through an intermediate medium; they can refer to the internal communication of two components or the interaction between two components, unless otherwise explicitly limited. Those skilled in the art can understand the specific meaning of the above terms in this invention according to the specific circumstances.

[0023] In this invention, unless otherwise explicitly specified and limited, "above" or "below" the second feature can mean that the first feature is in direct contact with the second feature, or that the first feature is in indirect contact with the second feature through an intermediate medium. Furthermore, "above," "over," and "on top" of the second feature can mean that the first feature is directly above or diagonally above the second feature, or simply that the first feature is at a higher horizontal level than the second feature. "Below," "below," and "under" the second feature can mean that the first feature is directly below or diagonally below the second feature, or simply that the first feature is at a lower horizontal level than the second feature.

[0024] Example: Please see Figures 1-7 In a preferred embodiment of the present invention, the glucose, cortisol, sodium ion, and pH detection sensor includes a substrate layer and a detection module disposed above the substrate layer. The detection module includes a glucose detection module, a cortisol detection module, a sodium ion detection module, and a pH detection module sequentially disposed on the substrate layer. The glucose detection module is used to detect the glucose concentration in the test liquid, the cortisol detection module is used to detect the cortisol concentration in the test liquid, the sodium ion detection module is used to detect the sodium ion concentration in the test liquid, and the pH detection module is used to detect the pH of the test liquid. A biomimetic microfluidic unit covers the glucose detection module, the cortisol detection module, the sodium ion detection module, and the pH detection module. Above the module, the biomimetic microfluidic unit has multiple vertically arranged flow channels, each containing a biomimetic capillary pump drive structure to guide the liquid to be tested to each detection module of the detection module; and a biomimetic sweat-guiding fabric layer, located on the side of the biomimetic microfluidic unit away from the detection module; the biomimetic sweat-guiding fabric layer has a double-sided fabric with unidirectional moisture-wicking properties; wherein, the side of the biomimetic sweat-guiding fabric layer away from the biomimetic microfluidic unit is a hydrophobic back-permeability layer, and the side of the biomimetic sweat-guiding fabric layer near the biomimetic microfluidic unit is a hydrophilic moisture-wicking layer, and there is a flow channel between the hydrophilic moisture-wicking layer and the hydrophobic back-permeability layer, the flow channel contains hydrophilic groups, and a gradient structure with increasing wettability is formed between the hydrophobic back-permeability layer and the hydrophilic moisture-wicking layer.

[0025] The glucose, cortisol, sodium ion, and pH detection sensors of this invention achieve simultaneous multi-parameter detection of glucose, cortisol, sodium ion, and pH in the test liquid through glucose detection modules, cortisol detection modules, sodium ion detection modules, and pH detection modules, respectively. Furthermore, by cooperating with a biomimetic microfluidic unit, utilizing the embedded capillary structure within the biomimetic microfluidic unit, sweat spontaneously enters the detection module areas without external driving force, achieving efficient detection of physiological indicators such as human stress hormones, electrolytes, and metabolites. Moreover, this invention utilizes the unidirectional moisture-wicking and breathable properties of the biomimetic sweat-guiding fabric layer to achieve unidirectional transport of sweat from the skin side to the detection side, effectively suppressing backflow of the detection liquid and external interference, thereby improving the sensor's detection efficiency.

[0026] Furthermore, as a preferred embodiment of the present invention, the glucose detection module of the present invention includes a gold electrode, and a Prussian blue electrocatalytic layer and an enzyme immobilization functional layer are sequentially disposed on the surface of the gold electrode to achieve electrochemical detection of glucose. Specifically, the Prussian blue electrocatalytic layer and the enzyme immobilization functional layer in the present invention have a synergistic effect. Prussian blue is used to catalyze the hydrogen peroxide reduction reaction to reduce the detection potential and reduce interference; the composite membrane structure of the enzyme immobilization functional layer is used to provide an enzyme immobilization environment and a conductive network structure, promote electron transport and improve substrate diffusion efficiency, thereby improving the sensor's sensitivity, response rate and detection stability.

[0027] Preferably, in this invention, Prussian blue is Fe4[Fe(CN)6]3, and the Prussian blue electrocatalytic layer is formed by electrochemical deposition. Specifically, this invention obtains a functional interface with electrocatalytic activity by depositing a Prussian blue layer on the surface of a gold electrode in an electrolyte containing FeCl3, K3[Fe(CN)6], KCl, and HCl using cyclic voltammetry.

[0028] Preferably, the enzyme immobilization functional layer is a chitosan / carbon nanotube / glucose oxidase composite membrane. Specifically, in this invention, chitosan is dissolved in acetic acid solution to form a matrix solution, which is then ultrasonically mixed with a carbon nanotube dispersion, and then mixed with a glucose oxidase solution in a predetermined ratio. Finally, the mixed solution is drop-coated onto the surface of a Prussian blue layer and dried to form the enzyme immobilization layer.

[0029] Further, as another optional embodiment of the present invention, the glucose detection module of the present invention includes a gold electrode with a nanostructure, and a multilayer functional structure is constructed on the surface of the gold electrode. The gold electrode with the nanostructure is a dendritic nanostructure, which is formed by periodic voltage deposition in a solution containing HAuCl4 and HCl, thereby obtaining a high specific surface area electrode structure. Next, a Prussian blue electrocatalytic layer is deposited on the surface of the gold nanostructure electrode, and a nickel hexacyanine iron (NiHCF) functional layer is further deposited to form a composite electrocatalytic interface. The NiHCF layer is formed by deposition in an electrolyte containing Ni(NO3)2 and K3[Fe(CN)6] using cyclic voltammetry. Further, the enzyme immobilization layer is a porous enzymatic membrane structure. It is formed by mixing alumina nanoparticles with agarose or chitosan matrix material to form a porous membrane system, which is then mixed with a glucose oxidase solution and drop-coated onto the electrode surface, and dried to form a porous enzyme immobilization membrane. In this embodiment, the glucose detection module utilizes a gold nanostructure to provide a high specific surface area and active sites; Prussian blue and NiHCF synergistically enhance electrocatalytic performance; and a porous enzyme immobilization membrane increases enzyme loading and provides mass transfer channels, thereby achieving rapid and highly sensitive detection of glucose and improving the stability and anti-interference capability of the sensor.

[0030] Furthermore, as an optional embodiment of the present invention, the cortisol detection module of the present invention includes a gold nanostructure electrode, preferably a gold nanopillar array electrode, and a cysteine ​​modification layer is disposed on the surface of the gold nanostructure electrode, and the cysteine ​​modification layer is formed by molecular self-assembly. The cysteine ​​modification layer is preferably an organic molecule modification layer containing thiol functional groups, which can form stable bonds with the gold surface through thiol groups, thereby constructing an ordered interface structure on the electrode surface. This interface structure helps to regulate the charge transport behavior of the electrode surface, reduce interfacial impedance, and improve electrochemical reaction kinetics performance; at the same time, the high specific surface area of ​​the gold nanostructure electrode can provide abundant active sites, which, in synergy with the organic modification layer, can improve the detection signal intensity and response sensitivity of the target molecule.

[0031] As another optional embodiment of the present invention, the cortisol detection module may further include a gold nanoparticle electrode, which is formed by electrohydrodynamic printing or other methods, and a recognition functional membrane is constructed on its surface. Optionally, the recognition functional membrane is preferably a conductive polymer composite membrane formed by an electrochemical method. The conductive polymer composite membrane contains structural units with molecular recognition functions, and specific recognition sites are formed within the membrane after removing template molecules. This recognition functional membrane has a porous structure, which is conducive to the diffusion and binding of target molecules, thereby achieving selective recognition and detection of cortisol.

[0032] Furthermore, as an optional embodiment of the present invention, the sodium ion detection module of the present invention includes a working electrode, which is preferably a carbon nanotube / silver wire composite electrode. A conductive interface layer is disposed on the surface of the carbon nanotube / silver wire composite electrode, and the conductive interface layer is formed by electrochemical deposition.

[0033] Preferably, the conductive interface layer is a conductive polymer layer, which can be constructed with a gradient structure through a pulsed electrochemical deposition process. The conductive interface layer formed in this way has low interfacial impedance and can achieve efficient ion-electron signal conversion.

[0034] The sodium ion detection module of this invention features a vertically oriented three-dimensional conductive network structure with nanopores in its conductive interface layer. This three-dimensional conductive network significantly increases the specific surface area and active site density of the working electrode, enhancing the enrichment of target ions on the electrode surface. Simultaneously, the nanopore structure serves as an ion transport channel, shortening the diffusion path and thus improving the detection response speed. The particle-electron hybrid conductivity of the conductive polymer helps improve interfacial charge transport efficiency and reduce interfacial impedance, thereby improving the sensor's signal-to-noise ratio and detection sensitivity. Furthermore, the continuous three-dimensional conductive network optimizes the charge transport path, reduces internal resistance, and contributes to improving the stability and anti-interference capability of the detection process.

[0035] Optionally, the working electrode in this invention can also be a doped conductive electrode, preferably a boron-doped diamond electrode, and an ion-selective sensitive film is disposed on the surface of the doped conductive electrode. The ion-selective sensitive film is preferably a plasticized polymer film, comprising an ion carrier, an ion-pairing reagent, a plasticizer, and a conductivity-enhancing component. This sensitive film is used to achieve selective recognition of sodium ions and provide a stable ion-response interface.

[0036] This invention selects a doped conductive electrode as the working electrode, which gives it a wider electrochemical working window and can effectively avoid interference from side reactions such as water decomposition. At the same time, its low background current helps to improve the signal-to-noise ratio of the detection signal, and the structural stability of the sensitive membrane can suppress phase separation of materials during long-term use, thereby improving the service life and detection reliability of the sensor.

[0037] Further, as an optional embodiment of the present invention, the pH detection module of the present invention includes a working electrode, which is preferably a gold nanostructure electrode, and a pH-sensitive layer is disposed on the surface of the gold nanostructure electrode. Preferably, the pH-sensitive layer is a conductive polymer film, which is formed by electrochemical deposition, and the conductive polymer film can form a continuous and dense functional layer structure on the surface of the working electrode through an electrodeposition process. The conductive polymer film is preferably a polyaniline film. The conductive polymer film has reversible protonation / deprotonation characteristics, and its surface potential changes under different pH environments, thereby realizing the potential response detection of hydrogen ion activity. By constructing this conductive polymer film on the surface of the gold nanostructure electrode, a stable interfacial electrochemical response system can be formed.

[0038] The high specific surface area of ​​gold nanostructured electrodes provides abundant active sites, which is beneficial for enhancing interfacial reactivity. The continuous structure of the conductive polymer film helps to improve charge transport efficiency and reduce interfacial impedance, thereby improving the sensor's response sensitivity and stability. In addition, the conductive polymer film has a rapid response to pH changes, enabling real-time monitoring of the acidity and alkalinity of the analyte liquid.

[0039] Furthermore, as another optional embodiment of the present invention, the pH detection module includes a working electrode, which is a metal or carbon-based electrode, and has a metal oxide sensitive layer disposed on its surface. The metal oxide sensitive layer is preferably a hydrated metal oxide thin film, which is formed on the surface of the working electrode by electrochemical deposition or physical deposition to construct a stable pH-responsive interface. The metal oxide thin film can achieve potential response detection of pH changes through the reversible reaction between surface hydroxyl groups and hydrogen ions in the solution. This hydrated metal oxide thin film has good chemical and electrochemical stability and can maintain stable response characteristics in complex environments.

[0040] The continuous structure of the metal oxide sensitive layer in this invention helps improve the stability of the electrode interface, while its surface active sites can promote the proton exchange process, thereby improving the response speed. Furthermore, the metal oxide sensitive layer has low background current and good anti-interference ability, which can effectively improve the accuracy and repeatability of the detection signal.

[0041] Preferably, the pH detection module is also used to comprehensively calibrate the enzyme activity of the glucose detection module. Fluctuations in the pH value of the test liquid significantly affect the activity of glucose oxidase and electron transfer efficiency, thus interfering with the glucose current response signal. Therefore, to achieve high-precision glucose detection, a calibration mechanism based on real-time pH measurement must be introduced.

[0042] Specifically, the pH-glucose calibration procedure is as follows: Constructing a pH-glucose calibration surface: (1) Preparation of a series of calibration solutions: Using phosphate buffer as the matrix, a series of calibration solutions covering the pH fluctuation range that may occur in actual application scenarios are prepared. Based on the physiological conditions of cortisol and glucose coexisting in human sweat, the pH value is usually set between 4.0 and 8.0. Based on this, seven calibration points are set with pH values ​​of 4.0, 5.0, 6.0, 6.5, 7.0, 7.5, and 8.0. At the same time, at each pH level, a glucose calibration solution with a concentration gradient covering the physiologically relevant range is prepared.

[0043] (2) Measurement of standard curve and sensitivity matrix: For each fixed pH value, glucose solutions of different known concentrations were added to the electrodes of the glucose detection module of the sensor in sequence, and the corresponding steady-state current response was measured. The glucose concentration was linearly regressed against the response current to obtain the glucose response curve under that pH condition.

[0044] (3) Constructing calibration equations: The sensitivity and baseline measured under different pH conditions are fitted with second-order polynomials to obtain the sensitivity calibration function and the baseline calibration function. A pH-glucose calibration surface is established based on the two calibration functions to describe the mapping relationship between glucose detection response and concentration under any pH condition.

[0045] The calibration equations are constructed as follows: Based on the established pH-glucose calibration surface, this invention uses the following calibration algorithm to perform pH compensation on the glucose measurement values ​​in actual samples.

[0046] Basic calibration formula: Among them, I glu The initial response current of the glucose electrode (unit: nA or μA); C gluThe glucose concentration to be calibrated; I0 (pH) and S glu (pH) are all functions of pH, expressed as second-order polynomials, as follows:

[0047]

[0048] The polynomial coefficients a0, a1, a2 and b0, b1, b2 were obtained through regression analysis of the factory calibration experiment and were pre-stored in the sensor's processing chip.

[0049] In actual testing, the sensor is only responsible for collecting the raw current signal and pH value data of the glucose electrode, and then, based on the preset calibration equation and coefficients, the pH-compensated glucose concentration is calculated.

[0050] Furthermore, such as Figure 3 As shown, in an optional embodiment of the present invention, the sensor includes an electrode system, which comprises a working electrode, a counter electrode, and a common reference electrode. The sodium ion detection module and the cortisol detection module each have a corresponding counter electrode, while the glucose detection module and the pH detection module only have working electrodes and no separate counter electrodes. Furthermore, the glucose detection module, cortisol detection module, sodium ion detection module, and pH detection module share the same common reference electrode.

[0051] More preferably, the common reference electrode has a multilayer composite structure, comprising a conductive layer, a chloride layer, and an ion-selective permeation layer sequentially disposed thereon. The conductive layer is preferably a silver-based conductive layer, the chloride layer is a silver / silver chloride composite layer formed on the surface of the conductive layer by an electrochemical method, and the ion-selective permeation layer covers the surface of the chloride layer to provide a stable electrochemical reference potential.

[0052] More preferably, the silver-based conductive layer is formed by inkjet printing, with a line width and spacing of 50 μm; and the chlorination layer has a porous structure with a porosity of 40-60%; the ion-selective permeation layer is preferably a Nafion membrane with a thickness of 2-5 μm.

[0053] Furthermore, such as Figure 2As shown in the optional embodiment of the present invention, the biomimetic microfluidic unit includes a flow channel structure constructed using 3D printed PDMS material, or a multi-layer flow channel structure formed by laser cutting double-sided adhesive tape. The inner wall surface of the multi-layer flow channel structure is modified with silica nanostructures. These silica nanostructures, by regulating the surface energy gradient, work in conjunction with the capillary drive structure to achieve stable spreading and directional transport of low-surface-tension sweat. An embedded capillary drive structure is also provided within the flow channel. This capillary drive structure includes a hydrophilic region and a hydrophobic flow-limiting structure, which together form a directionally selective capillary transport unit. This unit enables self-driven transport of the test liquid within the flow channel without external driving force, with a transport flow rate of 0.2~0.5 μL / min. Through this structural design, directional transport and stable supply of the test liquid can be achieved, ensuring uniform input to each detection module and improving the stability and repeatability of multi-parameter detection.

[0054] Furthermore, as an optional embodiment of the present invention, the biomimetic sweat-wicking fabric layer of the present invention is a double-sided structure with unidirectional moisture-wicking function. It includes a hydrophilic moisture-wicking layer and a hydrophobic back-seepage-blocking layer arranged sequentially. A wettability gradient is formed between the hydrophilic moisture-wicking layer and the hydrophobic back-seepage-blocking layer, thereby realizing the unidirectional transmission of the test liquid from the hydrophobic back-seepage-blocking layer facing the skin to the hydrophilic moisture-wicking layer away from the skin, so as to improve the liquid collection efficiency and suppress back-seepage. Among them, the hydrophilic moisture-wicking layer is a functional layer formed by surface modification of fiber substrate or introduction of hydrophilic functional groups, and its surface contact angle is not greater than 30°, which is used to quickly absorb and diffuse the test liquid. The hydrophobic back-seepage-blocking layer is a functional layer formed by modification with low surface energy materials, preferably a coating layer composed of fluoropolymers or siloxane materials, and its surface contact angle is not less than 110°, which is used to prevent liquid back-permeation. Through the structural design of the biomimetic sweat-guiding fabric layer, it is possible to avoid external interference while ensuring efficient liquid delivery. Furthermore, the biomimetic sweat-guiding fabric layer has a flexible and breathable structure, which allows the sensor to have good comfort and adaptability when attached to the surface, thereby meeting the needs of long-term continuous monitoring.

[0055] Furthermore, regarding the glucose, cortisol, sodium ion, and pH detection sensor of the present invention, the present invention also includes a method for preparing the glucose, cortisol, sodium ion, and pH detection sensor, which includes the following steps: S1. Cut and clean the substrate material to obtain the substrate layer; S2. Working electrodes for glucose detection module, cortisol detection module, sodium ion detection module, pH detection module, counter electrode, and common reference electrode are respectively set on the substrate layer. S3. Modify the working electrode of the glucose detection module; S4. Modify the working electrode of the cortisol detection module; S5. Modify the working electrode of the sodium ion detection module; S6. Modify the working electrode of the pH detection module; S7. A biomimetic microfluidic unit is formed above the glucose, cortisol, sodium ion and pH detection modules, and a biomimetic sweat-guiding fabric layer is formed on the surface of the biomimetic microfluidic unit.

[0056] As an optional embodiment of the present invention, the preparation of the substrate layer in step S1 includes cutting and cleaning the substrate material. Specifically, a flexible polymer film, preferably a PET film, is cut to a preset size as the substrate material; subsequently, the substrate material is cleaned sequentially with an organic solvent and deionized water, preferably acetone and ethanol. Oils, particulate matter, and polar contaminants on the substrate surface are removed using an ultrasonic-assisted method. After cleaning, the substrate surface is dried using a gas purging method. PET film has good flexibility and chemical stability, making it suitable for wearable sensor substrates attached to the skin surface; multi-stage solvent cleaning can effectively remove different types of surface contaminants, thereby improving the uniformity and stability of subsequent functional layer adhesion; the gas purging process avoids liquid residue and secondary contamination, improving the cleanliness of the substrate surface.

[0057] As a further preferred embodiment, the acetone cleaning temperature is 40°C, the ethanol cleaning temperature is 30°C, and the deionized water cleaning temperature is room temperature. Nitrogen gas is preferably used for gas purging at a pressure of 0.3 MPa.

[0058] Furthermore, as an optional embodiment of the present invention, the method for preparing the working electrode in step S2 is applicable to the glucose detection module, cortisol detection module, and pH detection module, all three employing a unified electrode preparation process. Specifically, it includes the following steps: A gold working electrode is prepared on a substrate layer using photolithography. The gold working electrode has a circular structure with a diameter of 3 mm, a thickness of 100 nm, and a linewidth of 0.5 μm. Photolithography enables high-precision pattern definition, thereby obtaining a uniform and repeatable electrode interface.

[0059] As another optional embodiment of the present invention, the working electrodes of the glucose detection module, cortisol detection module, and pH detection module in step S2 are prepared as follows: gold nanoparticle conductive ink is deposited on the surface of the substrate layer by inkjet printing, and a circular working electrode with a diameter of 3.5 mm is formed by laser sintering. The gold nanoparticle ink has a particle size of 20±2 nm. A laser source with a wavelength of 532 nm and a power of 0.5 W is selected during the laser sintering process to achieve dense bonding of nanoparticles, thereby forming a gold electrode with good conductivity. The surface roughness of the working electrode formed by this method is less than 50 nm.

[0060] Further, as an optional embodiment of the present invention, the preparation of the working electrode of the sodium ion detection module in step S2 includes: depositing a conductive paste on the surface of the substrate layer using microelectronic printing to form a working electrode structure of a preset size. The working electrode is preferably circular, with a size on the order of millimeters and a thickness on the order of micrometers, to obtain an electrode interface with good conductivity and structural stability. Microelectronic printing offers high positioning accuracy and enables micrometer-level pattern control, thereby ensuring the consistency of the electrode pattern. After formation, the working electrode is preferably subjected to a staged curing process to reduce thermal stress generated during curing, prevent deformation of the flexible substrate or interface cracking, and thus improve the bonding stability between the electrode and the substrate.

[0061] Optionally, the working electrode of the sodium ion detection module in this invention can also be fabricated using printed electronics. This involves forming an electrode pattern on the substrate surface using conductive composite ink. The conductive composite ink contains conductive nanomaterials and metal components, and after molding, it is densified through energy processing, thereby forming a continuous conductive network structure. The conductive composite ink has low resistivity, which is beneficial for improving the conductivity and detection sensitivity of the working electrode. Preferably, the curing process employs multi-stage temperature gradual curing. By designing the temperature changes as a gradient structure, internal stress is gradually released, and the density and stability of the electrode structure are improved. Using multi-stage temperature curing results in a low defect density on the cured electrode surface, which is beneficial for the uniform construction of subsequent functional layers, thereby improving the response stability and repeatability of the sodium ion detection module.

[0062] Furthermore, as an optional embodiment of the present invention, the preparation of the counter electrode in step S2 of the present invention includes: using a microelectronic printer to spray and print platinum paste on the substrate layer to form a circular counter electrode with a diameter of 3 mm and a thickness of 20 ± 2 μm.

[0063] Optionally, the preparation of the counter electrode in step S2 of the present invention can also be carried out as follows: a conductive composite material is deposited on the surface of a substrate layer to form the counter electrode. The conductive composite material preferably contains a metal catalytic component and a conductive carbon material. A porous electrode interface is formed through film formation and curing. The porous structure helps to increase the specific surface area of ​​the electrode, thereby enhancing the electrochemical reaction activity and improving catalytic performance. After formation, the counter electrode can be cured by heat treatment or radiation heating to promote the bonding between materials and form a stable conductive network structure. By controlling the curing conditions, an electrode interface with a good pore structure can be obtained, thereby improving the reaction efficiency and long-term stability of the electrode.

[0064] Furthermore, the preparation of the common reference electrode in step S2 of this invention includes: forming a silver-based conductive layer on the surface of the substrate layer, and constructing a chloride layer on the surface of the silver-based conductive layer by an electrochemical method to form a silver / silver chloride composite reference electrode structure. The electrochemical treatment is preferably performed using a constant potential method to obtain a potential-stable reference interface.

[0065] Optionally, the common reference electrode in step S2 of this invention can also be prepared by printed electronics, that is, by depositing silver / silver chloride composite conductive ink on the surface of the substrate layer, and forming a stable chloride layer structure on the electrode surface through a post-processing process. Interface control components can be introduced into the composite ink to improve the stability and adhesion performance of the electrode structure, thereby obtaining a common reference electrode with good potential stability.

[0066] The counter electrode and common reference electrode prepared by the above method can form a stable electrode system structure, providing a reliable electrochemical testing basis for multiple detection modules, thereby improving the overall detection accuracy and repeatability of the sensor.

[0067] Furthermore, as an optional embodiment of the present invention, the modification of the working electrode of the glucose detection module in step S3 of this application includes the construction of an electrocatalytic layer and the formation of an enzyme immobilization functional layer. Step S3 specifically includes: S301. Construct a Prussian blue electrocatalytic layer on the surface of the working electrode. Specifically, in an electrolyte containing FeCl3, K3[Fe(CN)6], KCl, and HCl, a Prussian blue (PB) layer is deposited on the surface of a gold electrode using cyclic voltammetry. The scanning potential range is 0–0.5 V relative to an Ag / AgCl reference electrode, and the scanning rate is 20 mV / s, thereby forming a Fe4[Fe(CN)6]3 electrocatalytic layer on the electrode surface.

[0068] S302. Preparation of enzyme-immobilized composite membrane. Specifically, chitosan is dissolved in an acidic solution to form a chitosan matrix solution, which is then mixed with a carbon nanotube dispersion and ultrasonically treated to form a uniform conductive composite solution. The composite solution is then mixed with a glucose oxidase solution at a volume ratio to obtain an enzyme-functionalized composite solution.

[0069] S303. The enzyme-functionalized composite solution is drop-coated onto the surface of the Prussian blue-modified working electrode and allowed to dry naturally to form an enzyme-immobilized functional layer, thus completing the modification of the working electrode of the glucose detection module.

[0070] The glucose detection module prepared by this method has a Prussian blue electrocatalytic layer that can catalyze the reduction reaction of hydrogen peroxide at low potential, thereby reducing interference signals; chitosan provides a good biocompatible environment to stabilize and immobilize enzyme molecules, and carbon nanotubes form a conductive network to promote electron transport. The synergistic effect of the three can significantly improve the sensitivity, response speed and selectivity of glucose detection.

[0071] As another optional embodiment of the present invention, the modification of the working electrode of the glucose detection module in step S3 of this application further includes constructing a multilayer composite functional structure. Step S3 specifically includes: Dendritic gold nanostructures were formed on the electrode surface by electrodeposition through periodic voltage signals applied in an electrolyte containing HAuCl4 and HCl. A Prussian blue electrocatalytic layer was deposited on the surface of the nanostructured gold electrode, followed by the deposition of a nickel hexacyanine iron (NiHCF) functional layer to form a composite electrocatalytic interface. The Prussian blue layer was formed by periodic voltage deposition, and the NiHCF layer was formed by cyclic voltammetry in an electrolyte containing Ni(NO3)2 and K3[Fe(CN)6].

[0072] Constructing porous enzyme immobilized membranes. Alumina nanoparticles were dispersed in a polymer matrix solution to form a porous membrane emulsion, which was then mixed with a glucose oxidase solution and drop-coated onto the electrode surface. After drying, a porous enzymatic membrane structure was formed.

[0073] Furthermore, as an optional embodiment of the present invention, the modification of the working electrode of the cortisol detection module in step S4 includes electrode pretreatment and construction of the recognition functional layer. Step S4 specifically includes: S401. Pretreatment of the working electrode. Specifically, the working electrode is cleaned in an acidic solution, and impurities and oxide layers on the electrode surface are removed by electrochemical scanning. Subsequently, it is rinsed with deionized water and dried with gas to obtain a clean and active electrode interface, which is conducive to the uniform construction of subsequent functional layers.

[0074] S402. Construct a composite film structure with molecular recognition function on the surface of the working electrode. Specifically, an electropolymerization system containing conductive monomers, metal coordinating components and target molecular templates is introduced into the electrode interface. A conductive polymer composite film is formed on the electrode surface by electrochemical methods, and the target molecule is embedded during the film formation process, thereby constructing an initial recognition structure.

[0075] S403. After the composite membrane is formed, the template molecules are removed by electrochemical methods combined with solvent treatment, resulting in a recognition cavity structure inside the composite membrane that matches the target molecule. This recognition cavity has a specific size and spatial configuration, enabling selective recognition and binding of the target molecule.

[0076] Through the above modification process, a recognition functional film with a porous structure can be formed on the surface of the working electrode. This structure is conducive to the diffusion and capture of target molecules. At the same time, the conductive polymer network can improve the charge transfer efficiency, thereby improving the sensitivity and response speed of the detection signal and enhancing the selectivity and stability of the sensor.

[0077] As another optional embodiment of the present invention, the recognition functional membrane can also be directly constructed on the electrode surface by printing or electropolymerization. A conductivity-enhancing component can be further introduced into the electropolymerization system to improve the conductivity and structural stability of the composite membrane. After film formation, template molecules are removed through electrochemical treatment and solvent synergy, thereby forming a functional interface with a high density of recognition sites. The resulting recognition structure has good porosity characteristics and interfacial stability, which is beneficial for achieving highly selective detection of cortisol.

[0078] Furthermore, as an optional embodiment of the present invention, the modification of the working electrode of the sodium ion detection module in step S5 includes the construction of an ion-selective sensitive layer and the formation of an interface conductive layer. Step S5 specifically includes: S501. Preparation of a sodium ion-selective membrane. Specifically, a membrane component comprising an ion support, an ion-pairing reagent, a polymer matrix, and a plasticizer is dissolved in an organic solvent, and a homogeneous membrane solution is formed by ultrasonic treatment. The solution is then stored in the dark to maintain its stability. The ion support is used to achieve selective recognition of sodium ions. The polymer matrix and plasticizer together form a flexible membrane structure, thereby ensuring that the sensitive membrane has good ion conductivity and mechanical stability.

[0079] S502. Construct an interfacial conductive layer on the surface of the working electrode. Specifically, a conductive polymer layer is deposited on the surface of the working electrode using an electrochemical method to form an ion-electron conversion interface. The conductive polymer layer has good ion-electron mixed conductivity, which can effectively reduce interfacial impedance and suppress potential drift, thereby improving the stability of the detection signal.

[0080] S503. After the conductive interface layer is formed, the sodium ion-selective membrane solution is coated onto the surface of the working electrode and then naturally cured or mildly treated to form a continuous and dense sensitive membrane structure. The thickness of the sensitive membrane is in the micrometer range, which can achieve a stable electrochemical response while ensuring ion selectivity.

[0081] Through the above modification process, a composite structure consisting of a conductive interface layer and an ion-selective membrane is formed on the surface of the working electrode. The conductive interface layer is used to realize charge transport and signal conversion, while the ion-selective membrane is used to realize the specific recognition of sodium ions. The two work together to significantly improve detection sensitivity, response speed and signal stability.

[0082] Optionally, the conductive interface layer in the sodium ion detection module can also be constructed using pulsed electrochemical deposition to form a conductive network with a gradient structure, thereby further reducing interfacial impedance and improving charge transport efficiency. The ion-selective membrane can also be constructed using inkjet printing or layer-by-layer deposition, and conductive enhancement components can be introduced to improve the membrane's conductivity and structural stability. The optimized sensitive membrane possesses excellent pore structure and interfacial bonding performance, thereby improving detection repeatability and anti-interference capabilities.

[0083] Furthermore, as an optional embodiment of the present invention, the modification of the working electrode of the pH detection module in step S6 includes the introduction of conductive nanomaterials and the construction of a pH-sensitive layer. Step S6 specifically includes: S601. Introduce conductive nanostructures on the surface of the working electrode. Specifically, disperse carbon nanotubes in an organic solvent, form a uniformly dispersed suspension by ultrasonic treatment, and modify the working electrode surface with the suspension to construct a conductive enhanced interface to improve the electrode's electron transport capability.

[0084] S602. Construct a pH-sensitive functional layer on the surface of the working electrode. Specifically, the modified working electrode is placed in an electrolyte containing an acidic medium and a conductive monomer, and a conductive polymer film is formed on the electrode surface by cyclic voltammetry; the scanning is performed multiple times within a preset potential range to form a uniform and dense functional layer structure on the electrode surface.

[0085] After electrodeposition, a continuous conductive polymer film is formed on the surface of the working electrode. The conductive polymer film has reversible response characteristics to hydrogen ions and can cause changes in electrode potential through protonation and deprotonation processes, thereby realizing the potential response detection of pH.

[0086] Through the above modification methods, carbon nanotubes provide a highly conductive network and enhance interfacial stability, while the conductive polymer film, as a pH-sensitive layer, enables rapid response; the synergistic effect of the two can significantly improve the sensitivity, response speed, and repeatability of pH detection.

[0087] As another optional embodiment of the present invention, the modification of the working electrode of the pH detection module in step S6 of this application further includes constructing a metal oxide sensitive layer. Specifically, a metal oxide thin film is formed on the surface of the working electrode by electrochemical deposition or physical deposition. The metal oxide thin film is preferably made of iridium oxide to construct a stable pH response interface. The surface of the metal oxide thin film has abundant hydroxyl active sites, which can undergo reversible reactions with hydrogen ions in the solution, thereby generating a stable potential response under different pH conditions. At the same time, this type of material has good chemical and electrochemical stability, and can maintain stable detection performance in complex environments.

[0088] As an optional embodiment of the present invention, the preparation of the biomimetic sweat-wicking fabric layer in step S7 of this application includes the construction of a hydrophilic moisture-wicking layer and a hydrophobic back-seepage-blocking layer. Step S7 specifically includes: S701. Select flexible fiber fabric as the base material, pre-treat it to remove surface impurities and improve the adhesion performance of subsequent functional layers.

[0089] S702. Construct a hydrophilic moisture-wicking layer on one side of the fabric. Specifically, hydrophilic functional groups are introduced onto the fiber surface through plasma treatment or chemical modification to enhance the wettability of the fabric surface, thereby forming a hydrophilic moisture-wicking interface with rapid liquid absorption and diffusion capabilities.

[0090] S703. Construct a hydrophobic backflow barrier layer on the other side of the fabric. Specifically, a hydrophobic functional layer is formed on the fabric surface by coating a low surface energy material. The hydrophobic layer can be formed by solution coating, spraying or dipping, and then dried or cured to obtain a stable structure.

[0091] Through the aforementioned double-sided structure, a wettability gradient is created along the fabric thickness, enabling unidirectional liquid transport from the hydrophilic side to the hydrophobic side. The hydrophilic layer rapidly absorbs and diffuses sweat, while the hydrophobic layer inhibits backflow, significantly improving sweat collection efficiency and transport stability. Furthermore, the fabric substrate possesses excellent flexibility and breathability, ensuring good comfort when the resulting biomimetic sweat-guiding fabric layer adheres to the skin surface, making it suitable for long-term continuous monitoring scenarios.

[0092] Furthermore, the present invention uses PDMS material to construct a biomimetic microfluidic unit with multiple branches (such as Y-shaped or tree-shaped channels), the outlet position of which is precisely aligned with the laser-cut microfluidic channel array on the biomimetic sweat-guiding fabric layer, covering the working electrodes of each detection module.

[0093] As an optional embodiment of the present invention, this application also includes characterization and analysis of the morphology and microstructure of each functional layer structure in the detection sensor.

[0094] Specifically, scanning electron microscopy is used to observe the morphology of the surface of each detection module to characterize the microstructure of the molecular recognition functional layer and the ion-selective sensitive membrane, in order to confirm whether the functional membrane layer forms a uniform and continuous porous structure, and to detect the compactness and integrity of the interface layer, thereby determining whether there are cracks, delamination or structural defects.

[0095] Furthermore, atomic force microscopy was used to perform three-dimensional morphology analysis on the surface of the sensitive membrane to obtain the surface roughness and micro / nano structure distribution of the membrane, thereby evaluating the structural uniformity of the membrane and its impact on ion or molecular transport behavior.

[0096] In addition, the structural state of the polymer-based sensitive membrane is analyzed by X-ray diffraction to determine the crystallization characteristics and internal structural stability of the membrane, thereby assessing whether structural changes occur during its preparation and use, and providing a basis for optimizing membrane performance.

[0097] The above characterization methods can be used to comprehensively evaluate the structural integrity and consistency of the functional layers in each detection module, thereby ensuring the stability and reliability of the sensor in the multi-parameter detection process.

[0098] Furthermore, as an optional embodiment of the present invention, the present invention also includes testing the electrochemical performance of each detection module of the glucose, cortisol, sodium ion and pH multifunctional detection sensor.

[0099] The sodium ion detection module uses the open-circuit potential method to test sodium ion solutions of different concentrations. The concentration gradient is preferably set to cover the range of 0.1 mM to 100 mM, with multiple logarithmic gradient concentration points. The test results are as follows: Figure 4 As shown, the potential response exhibits significant differences under different concentration conditions, and there is a good linear relationship between the potential signal and the sodium ion concentration, indicating that the module has stable detection capability over a wide concentration range.

[0100] The cortisol detection module uses chronoamperometry to test cortisol solutions of different concentrations. The concentration gradient is preferably set in the nanomolar range and includes multiple low-concentration test points. The test results are as follows: Figure 5 As shown, the current response corresponding to different concentrations has obvious differentiation, especially in the low concentration range where it still maintains high sensitivity, indicating that the module can effectively detect low concentrations of cortisol.

[0101] The glucose detection module uses the current response method to detect glucose solutions of different concentrations, with the concentration gradient preferably covering the micromolar to millimole range. Test results are as follows: Figure 6 As shown, the current signal exhibits a stable growth trend with increasing glucose concentration and a good linear fitting relationship, indicating that the module has high detection sensitivity and response stability.

[0102] The pH detection module uses the potentiometric response method to test buffer solutions with different pH values. The pH range is preferably designed to cover the weak acid to weak base range, and multiple interval test points are set. Test results are as follows: Figure 7 As shown, the electrode potential exhibits a good linear response to pH changes, indicating that the module can achieve rapid response detection of acidity and alkalinity.

[0103] Based on the above test results, each detection module exhibits good response characteristics and linearity within its corresponding concentration range, and demonstrates high consistency and stability during multi-parameter detection. Furthermore, the pH detection module can assist in calibrating the glucose detection signal, thereby further improving the overall detection accuracy and reliability.

[0104] Those skilled in the art will readily understand that the above description is merely a preferred embodiment of the present invention and is not intended to limit the present invention. Any modifications, equivalent substitutions, and improvements made within the spirit and principles of the present invention should be included within the scope of protection of the present invention.

Claims

1. A glucose, Cortisol, sodium ion and pH detection sensor, characterized by, include: basal layer; A detection module is disposed above the substrate layer; the detection module includes a glucose detection module, a cortisol detection module, a sodium ion detection module, and a pH detection module sequentially arranged on the substrate layer. A biomimetic microfluidic unit is provided, which covers the glucose detection module, cortisol detection module, sodium ion detection module and pH detection module. The biomimetic microfluidic unit has multiple flow channels arranged vertically, and each flow channel is provided with a biomimetic capillary pump drive structure to guide the liquid to be tested to each detection module of the detection module. The pH detection module is coupled to the glucose detection module; The pH detection module is used to acquire the acidity or alkalinity information of the liquid to be tested, so as to calibrate or compensate the detection signal of the glucose detection module; the glucose detection module integrates an enzyme activity dynamic compensation circuit based on the feedback of the pH detection module, which is used to receive the acidity or alkalinity information detected by the pH detection module and adjust the glucose concentration detection parameters accordingly. The pH detection module calibrates or compensates for the detection signal of the glucose detection module, including: Prepare a series of calibration solutions: using phosphate buffer as the matrix, prepare calibration points with pH values ​​of 4.0, 5.0, 6.0, 6.5, 7.0, 7.5, and 8.0; at each pH level, prepare a glucose calibration solution with a concentration gradient covering the physiologically relevant range. Standard curve and sensitivity matrix measurement: For each fixed pH value, glucose solutions of different known concentrations are added sequentially to the electrodes of the glucose detection module of the sensor, and the corresponding steady-state current response is measured; the glucose concentration is linearly regressed against the response current to obtain the glucose response curve under that pH condition. Constructing calibration equations: The sensitivity and baseline measured under different pH conditions are fitted with second-order polynomials to obtain the sensitivity calibration function and the baseline calibration function; Based on the two calibration functions, a pH-glucose calibration surface is established to describe the mapping relationship between glucose detection response and concentration under any pH condition; The sensor is used to collect the raw current signal and pH value data of the glucose acquisition module, and calculate the pH-compensated glucose concentration according to the calibration equation. The glucose detection module includes: A gold nanostructured electrode, wherein the surface of the gold nanostructured electrode is provided with a Prussian blue electrocatalytic layer, a nickel hexacyanide iron layer and an enzyme immobilization layer; The enzyme immobilization layer is formed by using a porous membrane system of alumina nanoparticles and agarose or chitosan matrix material, and then mixing it with glucose oxidase solution and drop-coating it onto the surface of gold nanostructure electrode. A biomimetic sweat-wicking fabric layer is provided on the side of the biomimetic microfluidic unit away from the detection module. The biomimetic sweat-wicking fabric layer is a double-sided fabric with unidirectional moisture-wicking properties. The side of the biomimetic sweat-wicking fabric layer away from the biomimetic microfluidic unit is a hydrophobic back-seepage barrier layer, and the side of the biomimetic sweat-wicking fabric layer close to the biomimetic microfluidic unit is a hydrophilic moisture-wicking layer. There is a flow channel between the hydrophilic moisture-wicking layer and the hydrophobic back-seepage barrier layer, and a gradient structure with increasing wettability is formed between the hydrophobic back-seepage barrier layer and the hydrophilic moisture-wicking layer.

2. The glucose, cortisol, sodium ion, and pH detection sensor according to claim 1, characterized in that, The hydrophobic backflow barrier layer comprises a coating layer made of fluoropolymer or siloxane material, and the surface contact angle of the hydrophobic backflow barrier layer is not less than 110°. The hydrophilic moisture-wicking layer includes a surface-modified fiber substrate or a functional layer with hydrophilic functional groups, and the surface contact angle of the hydrophilic moisture-wicking layer is not greater than 30°.

3. The glucose, cortisol, sodium ion, and pH detection sensor according to claim 1, characterized in that, The pH detection module includes: A gold nanostructure electrode, wherein a conductive polymer sensitive layer is disposed on the surface of the gold nanostructure electrode, and the conductive polymer sensitive layer is a polyaniline film formed by electrodeposition.

4. The glucose, cortisol, sodium ion, and pH detection sensor according to claim 1, characterized in that, The pH detection module includes: A metal oxide modified electrode, wherein the electrode surface has a hydrated metal oxide film, and the hydrated metal oxide film is a hydrated iridium oxide film.

5. The glucose, cortisol, sodium ion, and pH detection sensor according to claim 1, characterized in that, Both the glucose detection module and the cortisol detection module are provided with a counter electrode, and the glucose detection module, the cortisol detection module, the sodium ion detection module, and the pH detection module share the same common reference electrode; The common reference electrode is a multilayer composite structure, which includes a silver nano-conductive layer formed by inkjet printing, a silver chloride layer formed by electrochemical chlorination, and an ion-selective permeation membrane covering the silver chloride layer.

6. The glucose, cortisol, sodium ion, and pH detection sensor according to claim 5, characterized in that, The silver chloride layer has a porous structure with a porosity of 40-60%. The ion-selective permeation membrane is a Nafion membrane with a thickness of 2~5μm.

7. A method for preparing a glucose, cortisol, sodium ion, and pH detection sensor, used to prepare the glucose, cortisol, sodium ion, and pH detection sensor as described in any one of claims 1 to 6, characterized in that, The steps include the following: S1. Cut and clean the substrate material to obtain the substrate layer; S2. Working electrodes and counter electrodes for glucose detection module, cortisol detection module, sodium ion detection module and pH detection module are formed on the substrate by electron beam evaporation or sputtering, respectively, and a common reference electrode is formed by microelectronic printing. S3. Modify the working electrode of the glucose detection module; S4. Modify the working electrode of the cortisol detection module; S5. Modify the working electrode of the sodium ion detection module; S6. Modify the working electrode of the pH detection module; S7. A biomimetic microfluidic unit is formed above the glucose detection module, cortisol detection module, sodium ion detection module and pH detection module, and a biomimetic sweat-guiding fabric layer is formed on the surface of the biomimetic microfluidic unit.

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