Noninvasive glucose detection system

Abstract

The invention relates to the technical field of biomedical sensing and health monitoring, and discloses a noninvasive glucose detection system. The system comprises an interstitial fluid extraction device, a photoelectric sensing module and a main control circuit which are integrated on a flexible substrate. The interstitial fluid extraction device drives subcutaneous interstitial fluid to migrate to the hydrogel interface layer under the action of reverse ion electroosmosis; the photoelectric sensing module receives interstitial fluid through a hydrogel interface layer, a working electrode of the photoelectric sensing module comprises a Cu monatomic imprinting recognition layer, and imprinting holes formed by Cu monatomic sites are specifically coordinated with glucose molecules. The master control circuit controls sampling and determines the glucose concentration based on the photocurrent signal. Noninvasive sampling and monatomic imprinting photoelectric detection are fused, sampling-detection integration is achieved, and the device has the advantages of being noninvasive, painless, high in selectivity, high in interference resistance, comfortable to wear and the like and is suitable for continuous monitoring of blood glucose.

Description

Technical Field

[0001] This application relates to the field of biomedical sensing and health monitoring technology, and in particular to non-invasive glucose detection technology. Background Technology

[0002] Glucose (Glc) is the body's primary energy substrate, and its concentration level is closely related to diabetes and its chronic complications, metabolic syndrome, cardiovascular disease, and many other major diseases. Long-term, accurate monitoring of glucose levels is fundamental to diabetes prevention and metabolic management. Traditional blood glucose monitoring often involves finger-prick blood sampling combined with test strips or benchtop biochemical analyzers. While simple and intuitive, this method requires repeated skin punctures, causing pain and psychological burden, and carries the risk of cross-infection. It fails to meet patients' needs for high-frequency or even continuous dynamic monitoring, resulting in limited adherence.

[0003] To improve monitoring frequency and obtain dynamic information, various continuous glucose monitoring (CGM) systems have emerged in recent years. Most existing CGM products involve inserting microelectrodes into subcutaneous tissue to detect glucose concentration in tissue fluid, thus indirectly reflecting blood glucose levels. While these systems reduce repeated finger-prick blood sampling to some extent, they are still inherently invasive or minimally invasive implants, posing risks of skin and subcutaneous tissue damage during electrode implantation. Long-term use can easily lead to local inflammation, foreign body reactions, and discomfort. Furthermore, commonly used enzyme-based detection methods, such as glucose oxidase or glucose dehydrogenase, are susceptible to temperature, pH, and storage conditions, with enzyme activity decreasing over time, requiring periodic calibration and limiting long-term stable use.

[0004] Unlike the traditional approach of "directly detecting blood," interstitial fluid (ISF), located between capillaries and tissue cells, is a local microenvironment that cells directly "sensor." Under steady-state conditions, glucose concentration in ISF correlates well with blood glucose levels, exhibiting only a physiological lag of a few minutes. Therefore, it is widely regarded internationally as an effective alternative fluid for blood glucose monitoring. Compared to directly detecting blood, detecting glucose in ISF more accurately reflects the actual glucose levels exposed to tissues and cells, providing a more realistic picture of local metabolic status. Furthermore, it avoids direct entry into the vascular system, reducing risks of infection and coagulation. It also facilitates integration with flexible substrates and skin patches, enabling long-term continuous monitoring without frequent skin punctures, thus improving patient acceptance and compliance. Therefore, non-invasive or minimally invasive glucose monitoring in ISF has significant clinical value and promising application prospects.

[0005] On the one hand, ISF is located below the stratum corneum of the skin, making it difficult to obtain. Although existing reverse iontophoresis technology can achieve non-invasive transdermal extraction using microcurrents, it is often limited by the high variability of skin impedance and the complex physiological environment, resulting in problems such as low sampling throughput and unstable extraction efficiency.

[0006] On the other hand, most existing publicly available solutions still use enzyme-catalyzed electrochemical detection modules. The electrode interface is easily affected by coexisting substances such as ascorbic acid, uric acid, and skin secretions. Furthermore, interface contamination and enzyme activity decay are prone to occur during long-term operation, leading to signal drift and frequent calibration. Some detection methods that use optical or spectroscopic means often rely on bulky optical devices, resulting in limited device integration and making it difficult to create a truly thin and flexible patch. Summary of the Invention

[0007] One objective of this application is to provide a non-invasive glucose detection system that overcomes the shortcomings of existing systems, such as low sampling efficiency, poor anti-interference ability, and insufficient stability, and achieves non-invasive, continuous, and accurate detection of glucose.

[0008] This application discloses a non-invasive glucose detection system, including an interstitial fluid extraction device and a photoelectric sensing module integrated on the same flexible substrate, as well as a main control circuit; The interstitial fluid extraction device includes an electroosmotic electrode and a hydrogel interface layer; the hydrogel interface layer is adapted to adhere to the skin surface; the electroosmotic electrode is used to apply a microcurrent to generate reverse ion electroosmosis, thereby extracting subcutaneous interstitial fluid into the hydrogel interface layer. The photoelectric sensing module is in fluid communication with the interstitial fluid extraction device through the hydrogel interface layer to receive the interstitial fluid. It includes: a light window for allowing excitation light transmission; and a working electrode located on the optical path of the light window, comprising a photosensitive semiconductor substrate and a single-atom imprinted recognition layer. The single-atom imprinted recognition layer contains Cu single-atom active sites, which form imprinted holes matching the spatial configuration and electronic structure of glucose molecules, for specifically capturing glucose molecules through multi-site coordination with the hydroxyl and aldehyde groups of glucose molecules. The main control circuit is electrically connected to the electroosmotic electrode and the working electrode, and is used to: control the execution of the reverse ion electroosmosis; collect the photocurrent signal generated by the photoelectric sensing module in response to changes in glucose concentration under illumination; and determine the glucose concentration based on the photocurrent signal.

[0009] In a preferred embodiment, the photosensitive semiconductor substrate is titanium dioxide.

[0010] In a preferred embodiment, the single-atom imprinted recognition layer is prepared by the following method: Glucose template molecules are mixed with copper ion precursors in solution to form a coordination precursor solution; The photosensitive semiconductor substrate is immersed in the coordination precursor solution for incubation. The reaction system containing the photosensitive semiconductor substrate is frozen to a solid state; In a frozen state, a reduction potential is applied to the photosensitive semiconductor substrate to perform electrochemical deposition, thereby forming dispersed Cu single-atom active sites on its surface; Thaw the reaction system and remove the glucose template molecules.

[0011] In a preferred embodiment, the copper ion precursor is copper sulfate, the molar ratio of glucose template molecules to copper ions is 1:6, the freezing temperature range to solid state is -80°C to -20°C, the electrochemical deposition is performed by a constant potential method, and the reduction potential is -9V to -11V.

[0012] In a preferred embodiment, the system employs a planar multilayer stacked structure, with the interstitial fluid extraction device and the photoelectric sensing module arranged side by side on the flexible substrate; the hydrogel interface layer covers the sensing area of ​​the interstitial fluid extraction device and the sensing area of ​​the photoelectric sensing module, forming a continuous liquid channel so that the interstitial fluid diffuses laterally from the interstitial fluid extraction device to the photoelectric sensing module.

[0013] In a preferred embodiment, the flexible substrate is made of one or more materials selected from polydimethylsiloxane and thermoplastic polyurethane; a silver paste conductive circuit layer formed by screen printing is disposed on the flexible substrate for connecting the electroosmotic electrode and the working electrode respectively.

[0014] In a preferred embodiment, the main control circuit is configured to control the microcurrent applied by the electroosmotic electrode within the range of 0.2 mA to 0.4 mA.

[0015] In a preferred embodiment, an environmental sensor is also included for detecting at least one parameter among ambient temperature, humidity, and skin contact resistance; the main control circuit is configured to establish a calibration model using the detected parameters to correct the photocurrent signal.

[0016] In a preferred embodiment, the main control circuit further includes a wireless communication module for transmitting the determined glucose concentration to an external terminal device via Bluetooth Low Energy or Near Field Communication.

[0017] In a preferred embodiment, the system is encapsulated in a flexible wearable patch, the patch having an overall thickness of no more than 1 mm and a breathable and sweat-resistant membrane structure on its surface.

[0018] This application integrates an interstitial fluid extraction device based on reverse iontophoresis technology with a photoelectric sensing module containing a Cu single-atom imprinted recognition layer on the same flexible substrate. By utilizing the multi-site coordination between Cu single-atom sites and the hydroxyl and aldehyde groups in glucose molecules to form a specific "coordination anchor," true non-invasive blood glucose monitoring can be achieved, effectively solving the pain and infection risks of traditional invasive blood collection. At the same time, the single-atom imprinted structure precisely controls the spatial configuration and electronic environment of the recognition sites at the atomic scale, significantly improving the selective recognition ability of glucose molecules in complex body fluid environments. It effectively suppresses non-specific interference and signal drift caused by coexisting electroactive substances such as ascorbic acid and uric acid, thereby constructing an integrated closed-loop system of "sampling-transportation-highly selective detection," ensuring low-noise, high-sensitivity real-time dynamic monitoring results.

[0019] Furthermore, by using titanium dioxide as the photosensitive semiconductor substrate, the excellent photoelectric conversion properties and chemical stability of titanium dioxide can be utilized to provide a stable carrier support for the single-atom imprint recognition layer, ensuring the signal sensitivity and long-term stability of photoelectrochemical detection.

[0020] Furthermore, by mixing glucose template molecules with copper ion precursors in solution to form a coordination precursor solution, immersing the photosensitive semiconductor substrate in the solution for incubation, freezing it to a solid state, applying a reduction potential in the frozen state for electrochemical deposition, and removing the glucose template molecules after thawing, dispersed Cu single-atom active sites can be constructed in situ on the surface of the photosensitive semiconductor substrate. This forms imprint recognition holes whose spatial configuration and electronic environment are highly complementary to those of glucose molecules, achieving precise distribution of recognition sites at the atomic scale.

[0021] Furthermore, by using copper sulfate as a copper ion precursor and setting the molar ratio of glucose template molecules to copper ions to 1:6, controlling the freezing temperature within the range of -80℃ to -20℃, and setting the reduction potential of electrochemical deposition to -9V to -11V, the aggregation of metal atoms during the reduction process can be effectively suppressed, ensuring that Cu single atoms are uniformly dispersed on the carrier surface, thereby improving the uniformity of recognition site distribution and electron transport efficiency.

[0022] Furthermore, by adopting a planar multilayer stacked structure, the interstitial fluid extraction device and the photoelectric sensing module are arranged side by side on a flexible substrate, and the hydrogel interface layer covers the sensing area of ​​the two modules to form a continuous liquid channel. This allows the interstitial fluid extracted by electroosmosis to diffuse laterally from the interstitial fluid extraction device to the photoelectric sensing module. Spatial partitioning and liquid path connection between the sampling area and the detection area are realized in the same patch, simplifying the system structure and shortening the detection path.

[0023] Furthermore, by using polydimethylsiloxane or thermoplastic polyurethane to make a flexible substrate, and setting a silver paste conductive circuit layer formed by screen printing on the flexible substrate, a reliable electrical connection between the electroosmotic electrode and the working electrode can be achieved while ensuring the mechanical flexibility and skin compatibility of the patch, so that the patch can be tightly attached to the skin surface and adapt to human movement.

[0024] Furthermore, by controlling the microcurrent applied by the electroosmotic electrode within the range of 0.2 mA to 0.4 mA, the safety and comfort of the detection process can be ensured while effectively driving charged small molecules to migrate across the stratum corneum, avoiding irritation or damage to the skin.

[0025] Furthermore, by setting up environmental sensors to detect at least one parameter among ambient temperature, humidity, and skin contact impedance, and using the detected parameters to establish a calibration model to correct the photocurrent signal, measurement errors caused by environmental changes and individual differences can be eliminated, thereby improving the accuracy and repeatability of glucose concentration detection.

[0026] Furthermore, by setting up a wireless communication module in the main control circuit and using low-power Bluetooth or near-field communication to transmit the determined glucose concentration to an external terminal device, real-time wireless transmission of detection data can be achieved, making it convenient for users to conduct health management and risk assessment via mobile phones or computing platforms.

[0027] Furthermore, by encapsulating the system in a flexible wearable patch with a total thickness of no more than 1 mm and setting a breathable and sweat-proof membrane structure on the surface of the patch, the impact of the patch on daily activities can be reduced, while reducing the interference of sweat on the detection signal and preventing poor electrical contact caused by moisture evaporation, thereby improving wearing comfort and detection stability.

[0028] The various technical features disclosed in the above-described invention, the various technical features disclosed in the following embodiments and examples, and the various technical features disclosed in the accompanying drawings can be freely combined to form various new technical solutions (all of which should be considered as having been recorded in this specification), unless such a combination of technical features is technically infeasible. For example, in one example, feature A+B+C is disclosed, and in another example, feature A+B+D+E is disclosed. Features C and D are equivalent technical means that serve the same function, and technically only one needs to be used; it is impossible to use both simultaneously. Feature E can be technically combined with feature C. Therefore, the solution A+B+C+D should not be considered as having been recorded because it is technically infeasible, while the solution A+B+C+E should be considered as having been recorded. Attached Figure Description

[0029] Figure 1This is a schematic diagram of an integrated reverse ion electroosmosis-photoelectric detection patch structure according to an embodiment of this application.

[0030] Figure 2 This is a schematic diagram of the fabrication process of a single-atom imprint recognition layer according to an embodiment of this application.

[0031] Figure 3 This is a schematic diagram of the molecular-level structure of a Cu single atom coordinating with a glucose molecule at multiple sites according to an embodiment of this application.

[0032] Figure 4 This is a timing flowchart of electroosmotic sampling photoelectric detection according to an embodiment of this application.

[0033] Figure 5 This is a physical diagram of the construction of a reverse ion electroosmosis and photoelectric detection circuit according to an embodiment of this application. Detailed Implementation

[0034] To make the objectives, technical solutions, and advantages of this application clearer, the embodiments of this application will be described in further detail below with reference to the accompanying drawings.

[0035] Example 1 This embodiment provides a non-invasive glucose detection system, the overall structure of which is as follows: Figure 1 As shown, the system includes an interstitial fluid extraction device and a photoelectric sensing module integrated on the same flexible substrate, as well as a main control circuit electrically connected to both. The entire system is attached to the surface of human skin in the form of a patch, enabling real-time detection and quantitative analysis of glucose concentration in interstitial fluid without puncturing the skin.

[0036] The interstitial fluid extraction device consists of an electroosmotic electrode and a hydrogel interface layer. The hydrogel interface layer is made of polyacrylamide hydrogel material, with a thickness of approximately 0.2 mm to 0.3 mm, exhibiting good biocompatibility and high water content. During use, the hydrogel interface layer is directly attached to the skin surface, maintaining the moisture of the skin microenvironment through its three-dimensional network structure and providing ion transport channels. The electroosmotic electrode uses a screen-printed Ag / AgCl electrode pair, including a working electrode and a counter electrode, arranged concentrically or in parallel strips on a flexible substrate, with an electrode spacing of approximately 2 mm to 5 mm. When a microcurrent is applied to the electroosmotic electrode by the main control circuit, the electric field drives charged ions and small molecules in the subcutaneous interstitial fluid to migrate across the stratum corneum to the skin surface; this process is called reverse iontophoresis. Because glucose molecules can migrate alongside ions under physiological pH conditions, they can be extracted along with the interstitial fluid into the hydrogel interface layer. The sampling process typically lasts 10 to 20 minutes, enriching a sufficient amount of glucose-containing interstitial fluid sample in the hydrogel.

[0037] The photoelectric sensing module is fluidly connected to the interstitial fluid extraction device via a hydrogel interface layer to receive interstitial fluid extracted from the subcutaneous tissue. The core structure of this module includes a light window and a working electrode. The light window, located in the encapsulation layer on the upper surface of the patch, is made of a transparent polyethylene terephthalate film with a thickness of approximately 50 micrometers and a transmittance greater than 85% in the wavelength range of 300 nm to 400 nm, allowing external excitation light to be transmitted to the surface of the working electrode. The opening area of ​​the light window is approximately 4 square millimeters to 9 square millimeters, and its position is strictly aligned with the photosensitive area of ​​the working electrode below to ensure uniform illumination.

[0038] The working electrode is located on the optical path of the optical window and consists of two parts: a photosensitive semiconductor substrate and a single-atom imprinted recognition layer. In this embodiment, the photosensitive semiconductor substrate is made of titanium dioxide (TiO2). Titanium dioxide is a wide bandgap semiconductor with a bandgap of approximately 3.2 eV, capable of generating electron-hole pairs under ultraviolet light (wavelength less than 387 nm). Specifically, the TiO2 nanotube photonic crystal photosensitive material is prepared in situ on titanium foil using a two-step anodic oxidation method. The titanium foil is ultrasonically cleaned sequentially with acetone, ethanol, and deionized water, and then dried for later use. The anodic oxidation uses a dual-electrode system (titanium foil as the anode and Pt as the cathode), with an electrolyte of ethylene glycol containing NH4F (e.g., 0.32 wt% NH4F, 2.7 vol% water / EG), and is performed under cooling conditions. The first step involved oxidation at 60 V for 0.5 h to form a nanotube layer, followed by ultrasonic removal in water to obtain an ordered pit template. The second step involved further oxidation at 30 V for 0.5 h to regrow and form an ordered TiO2 nanotube structure. After cleaning and drying, the resulting sample was annealed at 500 °C for 1 h to obtain a TiO2 NTPCs photoelectrode.

[0039] A single-atom imprinted recognition layer covers the surface of a photosensitive semiconductor substrate, containing Cu single-atom active sites. These Cu single atoms are anchored monodispersely on the TiO2 support surface, forming a Cu-O coordination structure. The selection of copper ions is based on their good coordination ability with hydroxyl groups. In other embodiments, other metal ion precursors can be selected according to the functional group characteristics of the target molecule, such as manganese chloride, ferric chloride, cobalt sulfate, nickel sulfate, zinc sulfate, ruthenium chloride, potassium chloropalladium, cadmium chloride, potassium chloroiridate, potassium chloroplatinate, chloroauric acid, or bismuth nitrate metal salts as metal precursors for manganese, iron, cobalt, nickel, zinc, ruthenium, palladium, cadmium, iridium, platinum, gold, or bismuth, respectively. For complex target molecules containing multiple different functional groups, combinations of multiple metal ions can be used to achieve multi-metal, multi-site synergistic recognition. More importantly, the Cu single-atom active sites are imprinted using glucose molecules as templates during the preparation process, thereby forming imprinted holes that match the spatial configuration and electronic structure of glucose molecules. Glucose is a six-carbon aldose, and its molecular structure contains one aldehyde group (-CHO) and five hydroxyl groups (-OH). In the imprinted cavity, the spatial arrangement of Cu single-atom active sites precisely corresponds to the positions of the coordinateable functional groups in the glucose molecule. This allows for specific capture when the glucose molecule re-enters the detection interface, as Cu single atoms can coordinate with the hydroxyl and aldehyde groups of the glucose molecule at multiple sites.

[0040] The coordination mechanism between Cu single atoms and glucose molecules is as follows: Figure 3 As shown, the aldehyde oxygen atom at the C-1 position and the hydroxyl oxygen atoms at the C-2 and C-3 positions in the glucose molecule can simultaneously form coordination bonds with multiple adjacent Cu single atoms. This multi-site cooperative coordination mode produces a "coordination anchoring effect," giving the imprinted hole a very high binding constant for the glucose molecule, while preventing effective coordination with structurally similar interfering molecules (such as fructose, galactose, and maltose) due to differences in the spatial positions of their functional groups, thus achieving high selective recognition. When the glucose molecule is specifically captured and fills the imprinted hole, the charge transfer between the Cu single atom and the glucose significantly alters the electronic structure of the TiO2 / Cu interface, thereby affecting the recombination efficiency of photogenerated electron-hole pairs. Specifically, the coordination binding of the glucose molecule promotes the transmission of photogenerated electrons to the external circuit, increasing the photocurrent signal, and the increase is positively correlated with the glucose concentration.

[0041] The main control circuit is electrically connected to the electroosmotic electrode and the working electrode to achieve three core functions. First, the main control circuit controls the execution of reverse iontophoresis, specifically by outputting a set microcurrent to the electroosmotic electrode through a constant current source module to drive the interstitial fluid extraction process. Second, the main control circuit acquires the photocurrent signal generated by the photoelectric sensor module in response to changes in glucose concentration under illumination through a transimpedance amplifier. The transimpedance amplifier converts the weak photocurrent (nanoampere to microampere level) output from the working electrode into a voltage signal, which is then low-pass filtered and digitally sampled by an analog-to-digital converter. Finally, the main control circuit determines the glucose concentration based on the photocurrent signal. The microprocessor built into the main control circuit runs a pre-calibrated concentration conversion algorithm, substituting the digitized photocurrent data into the calibration curve to calculate the glucose concentration value in the interstitial fluid. Since there is a stable correlation between the interstitial fluid glucose concentration and blood glucose concentration (with a time delay of approximately 5 to 15 minutes), the current blood glucose level can be further calculated.

[0042] In this embodiment, the relationship between the photocurrent signal and the glucose concentration can be described by the following mathematical model. Let the photocurrent be... The glucose concentration is Within a certain concentration range, the two satisfy the following: in, This represents the baseline photocurrent in the absence of glucose. This is the sensitivity coefficient. These are parameters related to the binding constant of Cu single-atom active sites. In practical applications, the system is calibrated at multiple points using a glucose standard solution of known concentration. After fitting the above parameter values, the glucose concentration can be deduced from the measured photocurrent.

[0043] Experimental results demonstrate that the non-invasive glucose detection system of this embodiment exhibits good linear response within a glucose concentration range of 0.1 mmol / L to 20 mmol / L, with a detection limit of approximately 0.05 mmol / L. Compared to existing glucose detection technologies, this system eliminates the need for blood sampling, is easy to operate, and allows for direct quantitative detection of glucose on the skin surface. It is suitable for home health monitoring and continuous blood glucose management for diabetic patients and high-risk groups.

[0044] Example 2 This embodiment describes in detail the preparation method of the single-atom imprint recognition layer, and its process flow is as follows: Figure 2 As shown, this preparation method involves three core steps: template molecule coordination, cryo-electrochemical deposition, and template removal, to construct Cu single-atom active sites with glucose molecule complementary structures in situ on the surface of a photosensitive semiconductor substrate.

[0045] First, the glucose template molecule is mixed with the copper ion precursor in solution to form a coordination precursor solution. In this embodiment, copper sulfate (CuSO4) is used as the copper ion precursor, and the solvent is a 0.5 mol / L dilute sulfuric acid aqueous solution. The reason for choosing dilute sulfuric acid as the solvent is that protons (H+)... + In frozen solutions, protons play a crucial role in electron transfer. Experimental data show that the current density increases significantly with increasing proton concentration. Acidic solutions with a pH of 1 can be used in the preparation method of this application, with a preferred concentration of 0.5 mol / L sulfuric acid solution. The molar ratio of glucose template molecules to copper ions is set at 1:6, meaning that six Cu ions are arranged around each glucose molecule. 2+ Ions. This ratio is based on the fact that a glucose molecule contains six coordinateable oxygen atoms (one aldehyde oxygen and five hydroxyl oxygens), and six Cu atoms. 2+ The ions can coordinate with these oxygen atoms, thereby pre-constructing in solution a structure with glucose molecules as the core and Cu atoms as the core. 2+ The ion serves as the template for the coordination shell – a metal coordination complex. Specifically, 18 mg (1 mmol) of glucose and 15 mg (6 mmol) of copper sulfate pentahydrate are weighed and dissolved sequentially in 100 mL of 0.5 mol / L dilute sulfuric acid solution.

[0046] Subsequently, the pretreated photosensitive semiconductor substrate is immersed in a coordination precursor solution for incubation. Before immersion in the substrate, the prepared precursor solution needs to be stirred for approximately 12 hours to promote sufficient coordination between glucose molecules and copper ions. The stirring process should be conducted in the dark to prevent possible photochemical reactions from affecting the stability of the coordination complex. The photosensitive semiconductor incubation process is carried out with stirring at room temperature (20°C to 25°C) for 2 to 6 hours, preferably 4 hours. During this period, the glucose-Cu in the solution... 2+ The coordination complex is adsorbed onto the TiO2 substrate surface through electrostatic interactions and hydrogen bonding, forming a uniformly distributed precursor adsorption layer.

[0047] Next, the entire reaction system containing the photosensitive semiconductor substrate is frozen to a solid state. The freezing temperature range is -80°C to -20°C, preferably -78°C (the temperature of a dry ice-ethanol bath). Before freezing, the electrodes of the electrochemical system must be pre-placed in the reaction system. Specifically, a polyethylene plastic cup containing the substrate and precursor solution is placed in a cooling medium, and two platinum plates are inserted as the working electrode and counter electrode, respectively. The TiO2 substrate is placed between the two platinum plates with the photosensitive side facing the working electrode. The constructed reaction system is then completely immersed in liquid nitrogen for about 3 minutes to rapidly freeze it. After removal, a small amount of deionized water is added to the liquid surface, and the system is placed back into liquid nitrogen for 5 minutes to ensure that the reaction system is completely frozen into a solid ice block. Rapid freezing with liquid nitrogen allows the solution to freeze quickly within 5 to 15 minutes, resulting in a more uniform ice crystal structure compared to slow cooling. The core function of freezing is "space locking": when the solution freezes, solvent molecules form an ice crystal framework, while glucose-Cu 2+ The coordination complex is confined within the tiny spaces between ice crystals, Cu 2+ Diffusion and aggregation between ions are physically blocked. This low-temperature solid environment provides spatial constraints for the monodispersion of Cu atoms during the subsequent electrochemical reduction process.

[0048] Subsequently, electrochemical deposition was performed on the photosensitive semiconductor substrate under frozen conditions by applying a reduction potential to form dispersed Cu single-atom active sites on its surface. The electrochemical reduction process was carried out in a cryogenic bath at -78°C, controlled by an electrochemical workstation (such as the Chenhua electrochemical workstation system). A three-electrode system was used for electrochemical deposition, with a TiO2 substrate coated with a precursor adsorption layer as the working electrode, a platinum sheet as the counter electrode, and an Ag / AgCl electrode as the reference electrode. Electrochemical deposition was performed at a constant potential, with the reduction potential set from -9V to -11V, preferably -10V (relative to the Ag / AgCl reference electrode), and the sensitivity set to 1e. -1 At this strong reduction potential, the spatially confined Cu 2+ Ions are reduced to Cu single atoms in situ. Due to the hindered ion migration in the frozen state, the reduced Cu atoms can only remain near their original coordination sites and cannot aggregate into nanoparticles or clusters, thus achieving single-atom-level dispersion. The electrochemical deposition time is determined by controlling the cumulative charge. The charge is monitored in real time, and the reduction process is automatically stopped when the cumulative charge reaches a preset value (approximately 60 coulombs per square centimeter of electrode area in this embodiment) to control the Cu single-atom loading within a suitable range. There is a linear relationship between charge and single-atom loading, and the single-atom loading can be measured and verified using inductively coupled plasma atomic emission spectrometry (ICP-AES).

[0049] After electrochemical reduction, the reaction system was thawed and glucose template molecules were removed. The reaction vessel was removed from the cryogenic environment and allowed to thaw naturally at room temperature for approximately 30 minutes. After thawing, the substrate was repeatedly rinsed with plenty of deionized water to remove residual glucose template molecules and unreduced Cu. 2+ Ions and other possible impurities.

[0050] The single-atom imprinted recognition layer prepared by the above method contains a large number of imprinted holes on its surface that are complementary to the structure of glucose molecules. Since the spatial arrangement of Cu single-atom active sites remains unchanged after template removal, the geometry and electronic environment of the imprinted holes are highly matched with glucose molecules. When a glucose-containing interstitial fluid sample comes into contact with this recognition layer, glucose molecules can precisely "embed" into the imprinted holes and form multi-site coordination with Cu single atoms, achieving specific recognition and capture. X-ray photoelectron spectroscopy analysis confirms that Cu is present on the support surface in the form of Cu... + The valence state exists, and there are no obvious Cu nanoparticle peaks; the elemental mapping results of scanning transmission electron microscopy-energy dispersive spectroscopy show that Cu atoms are uniformly dispersed on the TiO2 surface, further verifying the successful realization of single-atom dispersion.

[0051] The preparation method in this embodiment combines the advantages of high activity and high atomic utilization of single-atom materials with the structural memory and selectivity advantages of molecular imprinting technology, overcoming the problems of template residue, uneven site distribution and blocked electron transport that exist in traditional molecular imprinted polymers, and providing an effective way to construct a highly sensitive and selective glucose photoelectrochemical sensing interface.

[0052] To further illustrate the universality of this preparation method, other target molecules besides glucose can also be used for single-atom imprinting. Taking 1,2-ethanedithiol (a dithiol separated by two methylene groups) as an example, since gold and thiol groups can form strong bonds, chloroauric acid (HAuCl4) was chosen as the metal ion precursor. 1,2-ethanedithiol and chloroauric acid were prepared in a 1:2 molar ratio and dissolved in a 0.5 mol / L dilute sulfuric acid solution. The remaining steps were the same as those for glucose imprinting. Photoelectric analysis results showed that the single-atom imprinted material only responded significantly to 1,2-ethanedithiol, while it did not respond to 1,3-propanedithiol (separated by three methylene groups) and 1,6-hexanedithiol (separated by six methylene groups), which also have thiol groups at both ends but with different spacing. This demonstrates that the single-atom imprinting method can precisely control the interatomic distance to achieve the recognition of specific molecular structures. Similarly, when 1,2-ethylenediamine was used as the target molecule, zinc sulfate was used as the metal precursor due to the good coordination ability of zinc with the amine group. The precursor solution was prepared by mixing 1,2-ethylenediamine and zinc sulfate in a 1:2 ratio. The resulting single-atom imprinted material also responded only to 1,2-ethylenediamine and not to 1,3-propanediamine or 1,6-hexanediamine. These experimental results fully demonstrate the good versatility of the single-atom imprinting preparation method of this application, allowing for the selection of appropriate metal ion precursors based on the structural characteristics of different target molecules, thus enabling the customization of highly selective molecular recognition materials.

[0053] Example 3 This embodiment describes in detail the planar multilayer stacked structure design of the non-invasive glucose detection system, focusing on the spatial layout of the interstitial fluid extraction device and the photoelectric sensing module, the construction of the hydrogel liquid channel, and the implementation of the flexible substrate and conductive circuit.

[0054] The system in this embodiment adopts a planar multi-layer stacked structure, such as Figure 1 As shown, the interstitial fluid extraction device and the photoelectric sensing module are arranged side by side on a flexible substrate, adjacent to each other in the horizontal direction, occupying the left and right regions of the patch, respectively. This side-by-side layout allows the patch to maintain a relatively thin overall thickness (facilitating skin adhesion), while the two functional modules are functionally independent yet continuously connected in the fluid path.

[0055] The flexible substrate, serving as the support carrier for the entire patch, is made of polydimethylsiloxane (PDMS) in this embodiment. PDMS is an organosilicon elastomer with excellent flexibility, biocompatibility, and breathability. Its Shore hardness is approximately 40A to 50A, allowing it to deform with the bending and stretching of human skin without causing discomfort or slippage during wear. The flexible substrate has a thickness of approximately 0.3 mm to 0.5 mm and planar dimensions of approximately 25 mm × 15 mm. In other embodiments, the flexible substrate can also be made of thermoplastic polyurethane (TPU), which also possesses good elasticity and skin affinity, and is easier to process and mold. Depending on the specific application requirements, the flexible substrate can also employ a composite structure of PDMS and TPU, for example, using TPU as the base layer and PDMS as the surface coating, combining the performance advantages of both.

[0056] A conductive silver paste layer is formed on a flexible substrate to connect the electroosmotic electrode and the working electrode. The conductive silver paste layer is formed using a screen printing process. Specifically, a pre-designed screen printing stencil is first placed on the surface of the flexible substrate. The pattern includes the leads for the electroosmotic electrode, the leads for the photoelectric working electrode, and pads for connection to external control circuitry. Then, conductive silver paste (a dispersion of nano-silver particles with a silver content of approximately 70% to 80%) is applied through the cutout areas of the screen printing stencil, depositing the silver paste onto the substrate surface to form conductive lines. Finally, the layer is baked and cured at 80°C to 120°C for 20 to 30 minutes, allowing the organic solvents in the silver paste to evaporate and the silver particles to sinter into a dense conductive layer. The cured silver paste circuit layer is approximately 10 to 20 micrometers thick, with a linewidth of approximately 0.2 to 0.5 millimeters and a sheet resistance of less than 0.1 ohms per square, meeting the low impedance requirements for microcurrent transmission and weak photocurrent acquisition.

[0057] In the layout design, the electroosmotic electrode leads and photoelectric working electrode leads of the silver paste conductive circuit layer are placed on the two side edges of the surface mount, maintaining an insulation gap of more than 2 mm in between, to avoid electric field interference or signal crosstalk between the two functional modules during operation. The electroosmotic electrode leads extend from the left edge of the surface mount to the Ag / AgCl electrode pair in the left functional area, while the photoelectric working electrode leads extend from the right edge of the surface mount to the TiO2 photosensitive electrode in the right functional area. The two sets of leads converge at the pad area at the bottom of the surface mount and are connected to the external main control circuit board via flexible ribbon cables or conductive adhesive.

[0058] A hydrogel interface layer covers the silver paste circuit layer and the functional electrode, simultaneously covering the sensing areas of both the interstitial fluid extraction device and the photoelectric sensing module, forming a continuous liquid channel. Specifically, the hydrogel interface layer is elongated or dumbbell-shaped, with its left end covering the area of ​​the electroosmotic electrode (electroosmotic sensing area) and its right end covering the area of ​​the photoelectric working electrode (photoelectric sensing area), connected by a hydrogel "bridge" approximately 3 to 5 millimeters wide. This continuous coverage design allows the interstitial fluid to diffuse laterally along the three-dimensional network structure of the hydrogel into the photoelectric sensing area on the right after electroosmosis extracts it into the hydrogel in the left sensing area, where it is detected by the single-atom imprint recognition layer on the surface of the working electrode. The lateral diffusion distance of the liquid in the hydrogel is approximately 5 to 10 millimeters, and the diffusion time is approximately 5 to 10 minutes. The total thickness of the hydrogel interface layer is approximately 0.2 to 0.4 millimeters, with a water content greater than 90%, ensuring both the continuity of liquid transport and maintaining a moist environment at the skin contact surface.

[0059] The top layer of the patch is an encapsulation and protective layer, made of a flexible material (such as a PDMS film) that is the same as or similar to the flexible substrate, with a thickness of approximately 0.1 mm to 0.2 mm. This encapsulation layer covers the entire patch surface, providing mechanical protection and electrical insulation to prevent external contaminants from entering the sensing area. A light window is created at the location of the photoelectric sensing module on the encapsulation layer. This light window is a rectangular or circular opening, with its area matching the active area of ​​the underlying TiO2 photoelectrode. The light window can be filled with transparent encapsulation material or left open to allow excitation light to pass through. The encapsulation layer above the electroosmotic module remains completely sealed, exposing only the hydrogel interface layer on the bottom surface in contact with the skin, ensuring that the electroosmotic current can form a stable circuit through the hydrogel-skin interface.

[0060] Through the aforementioned planar multilayer stacked structure design, this embodiment achieves spatial separation and fluid flow between the electroosmotic sampling functional area and the photoelectric detection functional area on the same flexible patch, enabling the entire "sampling-transfer-detection" process to be completed in a closed loop within a single device. This compact and integrated structural design avoids loss and contamination during sample transfer, shortens the detection response time, and keeps the overall patch thickness within 0.8 mm, providing excellent wearability and user comfort.

[0061] Example 4 This embodiment describes in detail the control strategy of the main control circuit for the electroosmotic sampling and photoelectric detection processes, as well as the implementation method of signal correction using environmental sensors.

[0062] The main control circuit is configured to control the microcurrent applied by the electroosmotic electrode within the range of 0.2 mA to 0.4 mA, preferably 0.3 mA. This microcurrent range is set by comprehensively considering both sampling efficiency and user safety. From the perspective of sampling efficiency, the extraction rate of reverse iontophoresis is positively correlated with the applied current intensity; the higher the current, the higher the migration rate of charged small molecules across the stratum corneum, and the shorter the sampling time. From a safety perspective, excessively high current can cause skin irritation or burns, and prolonged application may also damage the skin barrier function. Human trial data shows that applying a constant current of 0.3 mA to the forearm skin continuously for 15 to 20 minutes resulted in no significant discomfort in the subjects, with no adverse reactions such as redness, swelling, or blisters, while simultaneously extracting sufficient interstitial fluid for subsequent testing. The constant current source module in the main control circuit is built based on an operational amplifier and a reference voltage source, with an output current accuracy better than ±5% and overcurrent protection. It automatically cuts off the output when an abnormal increase in the electrode-skin circuit impedance is detected to prevent accidental injury.

[0063] The system in this embodiment also includes an environmental sensor for detecting at least one parameter among ambient temperature, humidity, and skin contact resistance. The environmental sensor is integrated with the main control circuit, and the parameter data it collects is transmitted in real time to the microprocessor of the main control circuit for processing.

[0064] Ambient temperature detection is achieved using a thermistor temperature sensor. This embodiment uses a negative temperature coefficient (NTC) thermistor with a nominal resistance of 10 kΩ (25°C) and a B value of 3950K. The NTC thermistor is connected to the analog-to-digital converter input of the microprocessor via a voltage divider circuit. The current temperature is obtained by looking up a table or calculating based on the divided voltage value. The temperature sensor is positioned close to the skin contact surface of the patch, providing a more accurate reflection of the temperature conditions near the detection interface. Ambient temperature has a significant impact on the photoelectrochemical reaction rate and photocurrent signal: as temperature increases, the carrier mobility of the semiconductor increases and electrolyte ion diffusion accelerates, leading to an increase in the photocurrent signal; conversely, as temperature decreases, the photocurrent decreases. Without temperature correction, fluctuations in ambient temperature can result in measurement errors of several percentage points.

[0065] Ambient humidity detection is achieved using a capacitive humidity sensor. The sensing medium of the capacitive humidity sensor is a polyimide film. When the relative humidity changes, the film absorbs or releases water molecules, altering its dielectric constant and thus changing the sensing capacitance. The capacitance signal output by the sensor is processed by a capacitance-to-voltage conversion circuit before being input to the microprocessor. Ambient humidity primarily affects the water content and ionic conductivity of the hydrogel interface layer: in high humidity environments, the hydrogel is less prone to water loss, ion transport is smooth, and the sampling and detection process is stable; in low humidity environments, the hydrogel surface is easily evaporated and dried, leading to a decrease in ionic conductivity, which may result in reduced electroosmotic efficiency and photocurrent baseline drift.

[0066] Skin contact impedance detection is achieved using an AC impedance measurement circuit. The main control circuit applies a weak AC excitation signal (frequency approximately 1 kHz, amplitude approximately 10 mV) to the electroosmotic electrodes while simultaneously detecting the response current, thereby calculating the impedance value of the skin-hydrogel interface. Skin contact impedance reflects the adhesion quality between the patch and the skin and the skin condition: excessively high impedance indicates poor adhesion or excessively dry skin, while excessively low impedance may indicate the presence of sweat or skin damage. Under normal adhesion conditions, the skin contact impedance is typically in the range of 5 kΩ to 50 kΩ.

[0067] The main control circuit is configured to establish a calibration model using the detected environmental parameters to correct the photocurrent signal. The basic form of the calibration model is as follows: in, The corrected photocurrent value. The original measured photocurrent value. , , These are the temperature correction function, humidity correction function, and impedance correction function, respectively. , , These represent the measured temperature, humidity, and impedance values, respectively. Each correction function was determined through extensive calibration experiments: the photocurrent response of a glucose standard solution of known concentration was measured under different temperature, humidity, and impedance conditions, and a multidimensional correction lookup table or fitted polynomial function was established. For example, the temperature correction function can be expressed as: in, This is a reference temperature (usually 25℃ or 37℃). , These are the coefficients obtained through experimental fitting.

[0068] The working sequence of electroosmotic sampling and photoelectric detection is as follows: Figure 4As shown. After the system is powered on, the main control circuit first reads the environmental sensor data to determine whether the environmental conditions are within the allowable range (temperature 10℃ to 45℃, humidity 20% to 90%, impedance 5 kΩ to 100 kΩ). If the conditions are met, the sampling phase begins, and the main control circuit controls the constant current source to output a 0.3 mA current for 15 minutes. After the sampling phase, the system enters the detection phase. The main control circuit turns on the light source (which can be a built-in miniature ultraviolet LED or rely on external light) and simultaneously starts photocurrent acquisition, continuously acquiring data at a sampling rate of 10 times per second for 30 to 60 seconds, and taking the average value as the effective photocurrent signal. Subsequently, the main control circuit calls the calibration model to correct the original photocurrent and then substitutes it into the concentration conversion formula to calculate the glucose concentration. The entire sampling-detection cycle takes about 20 minutes, and the system can automatically repeat the process at set intervals to achieve continuous dynamic monitoring.

[0069] Through the above-mentioned multi-parameter environmental perception and correction mechanism, this embodiment effectively eliminates the measurement error introduced by environmental factors, enabling the system to maintain stable and reliable detection performance under different temperature and humidity conditions and different skin conditions.

[0070] Example 5 This embodiment describes in detail the implementation of the wireless communication module in the main control circuit and the wearable packaging design of the system.

[0071] The main control circuit also includes a wireless communication module for transmitting the determined glucose concentration to an external terminal device via Bluetooth Low Energy (BLE) or Near Field Communication (NFC). In one implementation, the wireless communication module uses a Bluetooth Low Energy (BLE) chip. The BLE chip is integrated on the main control circuit board and connects to the microprocessor via a serial peripheral interface or a universal asynchronous transceiver (UAT). The microprocessor packages the calculated glucose concentration value into data frames conforming to the BLE protocol and transmits them through the BLE chip to the paired external terminal device (such as a smartphone, tablet, or dedicated health management device). BLE operates in the 2.4 GHz band, with an effective communication distance of approximately 10 to 30 meters, and consumes only a few milliwatts per data transmission, making it ideal for the low-power requirements of wearable devices. The external terminal device runs a companion mobile application that receives and displays glucose concentration data, records historical trends, sets high and low blood sugar warning thresholds, and generates health reports.

[0072] In another implementation, the wireless communication module uses a Near Field Communication (NFC) chip. The NFC chip is also integrated on the main control circuit board. When a user brings an NFC-enabled smartphone close to the patch (within 4 cm), the NFC chip is activated by the radio frequency field emitted by the phone, transmitting the stored glucose concentration data to the phone via near-field coupling. The advantage of NFC is that the patch does not require a battery or only a very small one; data transmission relies entirely on the radio frequency energy provided by the phone, resulting in a thinner and lighter patch structure. However, the disadvantage of NFC is that it cannot achieve real-time continuous transmission; the user must actively bring the phone close to the patch to read the data. Therefore, for applications requiring continuous monitoring and real-time alerts, BLE is recommended; for applications requiring only periodic point measurements, NFC is simpler.

[0073] The system in this embodiment is encapsulated in a flexible wearable patch. The patch has a rounded rectangular or elliptical shape, with planar dimensions of approximately 25 mm × 15 mm to 30 mm × 20 mm, and an overall thickness not exceeding 1 mm, preferably 0.6 mm to 0.8 mm. Both the upper surface encapsulation layer and the lower surface skin contact layer of the patch are made of medical-grade PDMS or TPU material, which is soft, conforms well, and is less likely to cause allergies. The main control circuit board, battery (if any), environmental sensors, and other electronic components are encapsulated in the middle layer inside the patch and connected to the silver paste circuit layer via flexible ribbon cables. The battery can be an ultra-thin lithium polymer battery, approximately 0.3 mm to 0.5 mm thick, with a capacity of approximately 10 mAh to 30 mAh, capable of supporting continuous system operation for 12 to 48 hours. If NFC is used, the battery can be omitted, and the system can be powered by an external radio frequency field.

[0074] The patch surface features a breathable and sweat-resistant membrane structure to reduce signal interference from sweat and prevent poor electrical contact caused by moisture evaporation. This membrane is made of porous polytetrafluoroethylene (PTFE) film or polyurethane breathable membrane with pore sizes ranging from approximately 0.1 to 1 micrometer. It effectively blocks liquid water and sweat from penetrating while allowing water vapor to pass through, maintaining a humidity balance within the patch's internal microenvironment. The breathable and sweat-resistant membrane covers the outer surface of the patch's encapsulation layer, and its edges are heat-sealed to the encapsulation layer to prevent edge leakage. During strenuous exercise or in high-temperature and high-humidity environments, the breathable and sweat-resistant membrane effectively isolates sweat from the skin surface, preventing salt and metabolic products in sweat from interfering with the electroosmotic sampling and photoelectric detection processes.

[0075] The patch is secured to the skin using medical pressure-sensitive adhesive. This adhesive is applied to the non-functional area surrounding the hydrogel on the underside of the patch, forming a ring-shaped adhesive band. This adhesive is hypoallergenic, has moderate adhesion, and good breathability, allowing users to apply the patch like a band-aid to smooth skin areas such as the inner forearm, outer upper arm, or abdomen. Normal activity and light exercise will not cause the patch to fall off. A single patch can function stably for more than 10 testing cycles, with a cumulative wear time of 24 to 72 hours, after which it can be replaced with a new patch for continued use.

[0076] In summary, this embodiment, through a wireless communication module and a flexible wearable packaging design, enables the non-invasive glucose detection system to easily connect with the user's smart terminal device, achieving real-time transmission, long-term recording, and intelligent early warning of detection data. This provides a convenient, comfortable, and sustainable blood glucose monitoring solution for diabetic patients and high-risk groups.

[0077] Technical Effect Description Compared with existing glucose detection technologies and wearable body fluid sensing technologies, the embodiments of this application have significant advantages in terms of detection methods, identification mechanisms, signal stability, and system integration.

[0078] This application's embodiments achieve truly non-invasive detection. Traditional glucose testing requires venous blood collection and relies on laboratory analysis, a highly invasive process with a long testing cycle, making continuous monitoring impossible. Existing wearable devices are mostly limited to acquiring physical signals (such as heart rate and blood oxygenation), lacking the ability to quantitatively analyze chemical indicators. This application's embodiments, by constructing a non-invasive body fluid sampling layer on the skin surface and using interstitial fluid instead of blood as the detection medium, can complete the acquisition of biochemical indicators without puncture or blood collection, providing a new technical approach for molecular-level health monitoring.

[0079] This application's embodiments possess the dual advantages of structural precision and chemical selectivity in molecular recognition mechanisms. Existing glucose detection methods often rely on enzymatic reactions, immunorecognition, or non-specific electrochemical signals, which are significantly affected by environmental interference and lack selectivity. This application's embodiments employ a single-atom imprinting strategy, using the glucose molecule itself as a template, and constructing Cu single-atom sites in situ on a TiO2 photosensitive support through a cryo-electrochemical deposition process. This design can simultaneously utilize the coordination characteristics of glucose's hydroxyl and aldehyde groups to form "coordination anchoring" sites, which are highly matched to the target molecule in both spatial configuration and electronic environment, thereby achieving highly selective recognition of glucose in complex body fluids. Compared to traditional molecularly imprinted polymers, the recognition sites in this application's embodiments are uniformly distributed, have short electron transport paths, and high recognition efficiency, avoiding template residue and non-specific adsorption problems.

[0080] The embodiments of this application exhibit good anti-interference and environmental stability. In the biological fluid environment, in addition to glucose, there are many other interfering substances that may respond to the sensing interface, especially electroactive small molecules such as ascorbic acid, uric acid, and dopamine, as well as structurally similar sugars such as fructose, galactose, and maltose. The typical concentrations of the main interfering substances in human plasma are: ascorbic acid approximately 400 μmol / L, glutathione approximately 1.1 mmol / L, lactose (a disaccharide composed of glucose and galactose) approximately 15 μmol / L, and fructose (an isomer of glucose) approximately 10 μmol / L. These substances may undergo redox reactions on the electrode surface or participate in charge transfer in photoelectrochemical processes, leading to false positives or quantitative deviations in traditional detection systems. The embodiments of this application construct a Cu single-atom imprinted recognition structure through coordination template construction, which selectively constrains the target molecule in terms of local coordination environment and spatial configuration, effectively suppressing competitive adsorption and bypass reactions of non-target molecules. Photoelectrochemical testing results showed that none of the aforementioned interfering substances produced a significant response on the single-atom imprinted material at their physiological concentrations, with typical signal changes caused by the interfering substances being less than 10%, demonstrating good anti-interference performance. This single-atom imprinted material can be used to determine glucose in real samples such as human plasma or urine for monitoring blood glucose or urine glucose concentrations.

[0081] The detection interface and flexible packaging design of this application combine to form a wearable, reusable integrated system. The imprinted photoelectrode and the flexible hydrogel sampling layer are directly coupled in space and connected to the micro signal acquisition unit through a conductive channel. The entire system is encapsulated in a PDMS or TPU substrate, ensuring both mechanical flexibility and photoelectric response stability. This structure is suitable for adhesion to the skin surface and can operate stably in human environments such as 37°C and high humidity, achieving integrated "sampling-detection-transmission" operation. Compared with existing electrochemical patches or microneedle sensors, this application does not require skin penetration, offers stable signal, is safe to use, and has higher user compliance and clinical translation potential.

[0082] All references to this application are considered to be incorporated in their entirety into the disclosure of this application so that they can serve as a basis for modifications if necessary. Furthermore, it should be understood that after reading the contents of this application, those skilled in the art can make various alterations or modifications to this application, and these equivalent forms also fall within the scope of protection claimed in this application.

Claims

1. A non-invasive glucose detection system, characterized by, The system comprises an interstitial fluid extraction device and a photoelectric sensing module integrated on the same flexible substrate, and a master control circuit; The interstitial fluid extraction device comprises an electro-osmotic electrode and a hydrogel interface layer; the hydrogel interface layer is suitable for attaching to the skin surface; the electro-osmotic electrode is used to apply a micro-current to generate a reverse iontophoresis, thereby extracting subcutaneous interstitial fluid into the hydrogel interface layer; The photoelectric sensing module is in fluid communication with the interstitial fluid extraction device through the hydrogel interface layer to receive the interstitial fluid, and comprises: a light window for allowing excitation light transmission; a working electrode located on the light path of the light window, comprising a photosensitive semiconductor substrate and a single-atom imprint recognition layer; the single-atom imprint recognition layer contains Cu single-atom active sites, and the active sites form imprint cavities matched with the spatial configuration and electronic structure of glucose molecules, for specifically capturing glucose molecules by multi-site coordination with the hydroxyl and aldehyde groups of the glucose molecules; The master control circuit is electrically connected with the electro-osmotic electrode and the working electrode, and is used to: control the execution of the reverse iontophoresis; collect the photocurrent signal generated by the photoelectric sensing module in response to the change of glucose concentration under light; and determine the glucose concentration based on the photocurrent signal.

2. The noninvasive glucose detection system of claim 1, wherein, The photosensitive semiconductor substrate is a titanium dioxide material.

3. The noninvasive glucose detection system of claim 1, wherein, The single-atom imprint recognition layer is prepared by the following method: Mixing glucose template molecules and copper ion precursors in a solution to form a coordination precursor solution; Immersion of the photosensitive semiconductor substrate in the coordination precursor solution for incubation; Freezing the reaction system containing the photosensitive semiconductor substrate to a solid state; Applying a reduction potential to the photosensitive semiconductor substrate in the frozen state for electrochemical deposition to form dispersed Cu single-atom active sites on the surface thereof; Thawing the reaction system and removing the glucose template molecules.

4. The noninvasive glucose detection system of claim 3, wherein, The copper ion precursor is copper sulfate, and the molar ratio of the glucose template molecules to the copper ions is 1:6; the freezing temperature ranges from-80℃ to-20℃; the electrochemical deposition is performed by a constant potential method, and the reduction potential is-9V to-11V.

5. The noninvasive glucose detection system of claim 1, wherein, The system adopts a planar multi-layer laminated structure, and the interstitial fluid extraction device and the photoelectric sensing module are arranged side by side on the flexible substrate; the hydrogel interface layer covers the sensing area of the interstitial fluid extraction device and the sensing area of the photoelectric sensing module, forming a continuous liquid channel to enable the interstitial fluid to laterally diffuse from the interstitial fluid extraction device to the photoelectric sensing module.

6. The noninvasive glucose detection system of claim 1, wherein, The flexible substrate is made of one or more materials selected from polydimethylsiloxane and thermoplastic polyurethane; a silver paste conductive circuit layer formed by screen printing is provided on the flexible substrate for connecting the electro-osmotic electrode and the working electrode, respectively.

7. The noninvasive glucose detection system of claim 1, wherein, The master control circuit is configured to control the micro-current applied by the electro-osmotic electrode to be in the range of 0.2 to 0.4 milliamperes.

8. The noninvasive glucose detection system of any one of claims 1-7, wherein, The system further comprises an environmental sensor configured to detect at least one of an ambient temperature, a humidity, and a skin contact impedance; and the main control circuit is configured to establish a correction model using the detected parameter to correct the photocurrent signal.

9. The noninvasive glucose detection system of any one of claims 1-7, wherein, The main control circuit further comprises a wireless communication module configured to transmit the determined glucose concentration to an external terminal device via Bluetooth Low Energy or near field communication.

10. The noninvasive glucose detection system of any one of claims 1-7, wherein, The system is packaged in a flexible wearable patch, the total thickness of the patch is not more than 1 millimeter, and the surface layer is provided with a breathable and sweat-proof membrane structure.

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