An enzyme-free uric acid electrochemical sensor, a uric acid electrochemical detection method and a wearable wireless electrochemical sensing system

Abstract

The application discloses an enzyme-free uric acid electrochemical sensor, a uric acid electrochemical detection method and a wearable wireless electrochemical sensing system. The working electrode of the sensor comprises a graphene film and a continuous single-layer graphene oxide active layer with a coverage of no less than 94.9% and a thickness of 1.4-1.6 nm on the graphene film. The electrode has inherent activity at a lower potential for detecting uric acid by differential pulse voltammetry without uric acid oxidase and noble metal, with the minimum interface tunneling distance and n-type narrow band gap level arrangement. The detection method covers 0-5000 micromoles per liter and has a detection limit of sub-micromolar level by lithium carbonate solubilization, two-stage splicing calibration, and specific recognition of uric acid in the presence of ascorbic acid and dopamine by second-order derivative deconvolution. The system integrates the electrode into a laser patterned three-electrode array, PET packaging, PDMS microfluidic and wireless flexible circuit board, and can be used for non-invasive real-time monitoring of uric acid in body fluids, and can be used for uric acid monitoring of livestock, poultry and aquatic animals in intelligent agriculture.

Description

Technical Field

[0001] This invention belongs to the field of electrochemical biosensing technology, specifically relating to an enzyme-free uric acid electrochemical sensor, an electrochemical detection method for uric acid, and a wearable radiochemical sensing system, which can be used in smart agriculture for monitoring uric acid metabolism in livestock, poultry, and aquaculture animals. Background Technology

[0002] Uric acid is the end product of purine metabolism. Abnormal concentrations of uric acid in body fluids are closely related to pathological conditions such as gout and visceral gout (which is particularly prevalent in poultry). Therefore, highly sensitive, wide-range, and interference-resistant electrochemical monitoring of uric acid is of great significance.

[0003] In the electrochemical sensing of uric acid, enzymatic sensing (such as that based on uricase) is easily inactivated by environmental factors such as temperature and pH, exhibiting poor stability and high cost. Existing non-enzymatic electrocatalytic strategies mostly rely on noble metals or their composite nanostructures as catalytic sites, which suffer from high cost and limited resources, or require high oxidation overpotentials due to insufficient inherent activity at the electrode interface. Therefore, how to achieve highly sensitive electrochemical oxidation detection of uric acid at a lower potential using the inherent activity of the electrode interface without using uricase or relying on noble metal catalytic sites is a problem that urgently needs to be solved by current technologies.

[0004] On the other hand, the concentration range of uric acid in physiological samples is very wide, extending from low concentrations to high clinical concentrations. Existing uric acid sensing platforms generally have inherent trade-offs: to achieve sub-micromolar detection limits, the upper limit of the linear range is often compressed, making it difficult to balance a low detection limit with a wide linear dynamic range; furthermore, the inherently limited solubility of uric acid in conventional phosphate buffer is insufficient to cover the high clinical concentration range, further restricting the upper limit of the detection range. Therefore, how to maintain sub-micromolar detection limits while ensuring the linear detection range covers the high clinical concentration range remains a challenge in this field.

[0005] Furthermore, in real samples, uric acid often coexists with electroactive substances such as ascorbic acid and dopamine. The oxidation potentials of these three substances are close, and their voltammetric responses overlap significantly. In particular, the strong oxidative background of ascorbic acid can easily mask the independent signals of dopamine and uric acid. A key challenge in achieving interference-resistant detection of real samples is how to specifically distinguish and identify uric acid from these severely overlapping voltammetric responses without using enzymes or introducing additional modifications.

[0006] Furthermore, transforming a desktop electrochemical detection platform into a wearable terminal capable of continuously monitoring non-invasive bodily fluids such as sweat and urine in real time by conforming to the human or animal body requires deep integration of the sensing interface with flexible microelectronics and microfluidic sampling technology. It also necessitates addressing issues such as rapid collection of bodily fluids and uniform concentration within the cavity, wireless transmission of electrochemical signals, and maintaining the intrinsic sensing performance of the active interface after system-level engineering transformations involving flexible packaging, microfluidic integration, and wireless transmission. Current technologies still have shortcomings in integrating high-performance sensing interfaces into a complete wearable system that combines flexibility, microfluidic sampling, and wireless transmission capabilities. Summary of the Invention

[0007] The purpose of this invention is to address the shortcomings of the aforementioned background technology by providing an enzyme-free uric acid electrochemical sensor, an electrochemical uric acid detection method, and a wearable radiochemical sensing system. This enables enzyme-free and precious metal-independent low overpotential uric acid detection, while maintaining a sub-micromolar detection limit and a wide linear range. It also achieves specific identification of uric acid under strong interference conditions and facilitates the engineering transformation of high-performance sensing interfaces into flexible wearable systems.

[0008] The technical solution adopted in this invention is: an enzyme-free uric acid electrochemical sensor, comprising a three-electrode system consisting of a working electrode, a counter electrode, and a reference electrode; the working electrode comprises a graphene thin film conductive layer and a continuous monolayer graphene oxide active layer located on the graphene thin film conductive layer; the continuous monolayer graphene oxide active layer is formed by embedding monolayer graphene oxide nanosheets in a single-layer flat arrangement, with their edges connected and substantially non-overlapping, with a thickness of 1.4-1.6 nm, a coverage of not less than 94.9% on the surface of the graphene thin film conductive layer, and direct contact with the graphene thin film conductive layer via π-π conjugation interaction; the enzyme-free uric acid electrochemical sensor is configured to electrochemically detect uric acid in body fluid samples using differential pulse voltammetry without using uric acid oxidase and noble metal catalytic sites.

[0009] In the above technical solution, the conductive layer of the graphene film is a highly oriented graphene film patterned by laser micromachining to define the electrochemical active region of the working electrode.

[0010] In the above technical solution, in a 0.1 mol / L phosphate buffer solution at pH 7.4, using a silver / silver chloride reference electrode filled with saturated potassium chloride solution as the potential reference, the differential pulse voltammetric anodic oxidation peak of the enzyme-free uric acid electrochemical sensor for uric acid is located at 0.36 V. Under the same test conditions, the anodic oxidation peak is negatively shifted by no less than 0.3 V relative to the anodic oxidation peak of the bare graphene film conductive layer without the deposited continuous monolayer graphene oxide active layer for uric acid. Within the range of uric acid concentration greater than 0 and not exceeding 1000 μmol / L, the peak current of the anodic oxidation peak increases monotonically with the increase of uric acid concentration.

[0011] This invention provides an electrochemical detection method for uric acid, employing the enzyme-free uric acid electrochemical sensor described above, to electrochemically quantify uric acid in the sample without using uric acid oxidase, comprising: (1) Establish a two-segment splicing calibration relationship covering 0-5000 μmol / L: The differential pulse voltammetric anodic oxidation peak current of the working electrode is collected with a series of standard solutions with known uric acid concentrations. Based on this, a first linear calibration curve for quantification of uric acid not exceeding 1000 μmol / L is established in a phosphate buffer without lithium carbonate in the range of 0-1000 μmol / L uric acid concentration. A second linear calibration curve for quantification of uric acid greater than 1000 μmol / L is established in a phosphate buffer with lithium carbonate added as a solubility enhancer in the range of greater than 1000-5000 μmol / L uric acid concentration. The first linear calibration curve and the second linear calibration curve are connected at a connection concentration of 1000 μmol / L, and the current response deviation between the two at the connection concentration is less than 3%. (2) Acquire a differential pulse voltammetric response curve for the sample to be tested in a phosphate buffer solution without lithium carbonate, and obtain the anodic oxidation peak current of the sample to be tested from the differential pulse voltammetric response curve; when the obtained anodic oxidation peak current exceeds the current response corresponding to 1000 μmol / L of the first linear calibration curve, re-acquire a differential pulse voltammetric response curve for the sample to be tested in a phosphate buffer solution with added lithium carbonate, and obtain the anodic oxidation peak current of the sample to be tested from it; (3) When the sample to be tested is collected in a phosphate buffer without lithium carbonate, the obtained anodic oxidation peak current is converted into the uric acid concentration in the sample to be tested according to the first linear calibration curve; when the sample to be tested is collected in a phosphate buffer with added lithium carbonate, the obtained anodic oxidation peak current is converted into the uric acid concentration in the sample to be tested according to the second linear calibration curve.

[0012] In the above technical solution, the sample to be tested is a sample containing ascorbic acid, dopamine, and uric acid. The differential pulse voltammetric response curve obtained from the sample is smoothed using Savitzky-Golay second-order numerical differentiation, so that the oxidation contributions of the coexisting ascorbic acid, dopamine, and uric acid are expressed as independent minimum points on the resulting second-order derivative curve. Based on this, ascorbic acid, dopamine, and uric acid can be distinguished without the use of enzymes, and the oxidation contribution corresponding to uric acid can be identified, thus specifically detecting uric acid in the presence of ascorbic acid and dopamine. Furthermore, even when the concentration of ascorbic acid is not less than 20 times the concentration of dopamine, the minimum points corresponding to uric acid and dopamine can still be distinguished and identified.

[0013] In the above technical solution, the detection limit of uric acid corresponding to the first linear calibration curve is no greater than 0.078 μmol / L, and the detection limit of uric acid corresponding to the second linear calibration curve is no greater than 0.087 μmol / L.

[0014] This invention provides a wearable radiochemical sensing system, configured to be detachably worn or attached to the body surface of a test subject for non-invasive real-time electrochemical sensing of the test subject's bodily fluids. The system comprises: a three-electrode array, including a working electrode, a counter electrode, and a reference electrode, patterned on a graphene film by laser micromachining. The working electrode surface has a continuous monolayer graphene oxide active layer, which is formed by embedding monolayer graphene oxide nanosheets in a single-layer, flat, edge-connected, and substantially non-overlapping manner, with a coverage of not less than 94.9% and a thickness of 1.4-1.6 nm on the working electrode surface. The reference electrode is formed by... The system comprises: a saturated silver / silver chloride paste coated on the graphene film; a top polyethylene terephthalate encapsulation layer and a bottom polyethylene terephthalate substrate encapsulating the three-electrode array therein and having electrochemical windows exposing the working area of ​​the three-electrode array; an elastomeric microfluidic chip covering the electrochemical windows for collecting bodily fluids from the test subject and bringing the collected bodily fluids into contact with the working electrodes; and a flexible printed circuit board electrically connected to the three-electrode array via a flexible connector and integrating an electrochemical measurement circuit, a low-power wireless communication module, and a power management unit. The flexible printed circuit board is configured to wirelessly transmit the electrochemical signals collected by the three-electrode array to an external smart terminal.

[0015] In the above technical solution, the elastomer microfluidic chip is a polydimethylsiloxane microfluidic chip, which includes a semi-circular main collection chamber, multiple inlet channels communicating with the semi-circular main collection chamber, and an outlet channel located at the bottom of the semi-circular main collection chamber; the semi-circular main collection chamber is configured such that after the body fluid of the test subject enters through the multiple inlet channels, the concentration of the analyte in the semi-circular main collection chamber tends to be uniform within 15 seconds.

[0016] In the above technical solution, the wearable radiochemical sensing system is configured to perform electrochemical sensing of uric acid in the sweat of the test subject using wireless timing amperometry, and the concentration range of the electrochemical sensing covers 30-150 μmol / L; or it is configured to perform electrochemical sensing of uric acid in the urine of the test subject using differential pulse voltammetry, and the concentration range of the electrochemical sensing covers 2-5 mmol / L.

[0017] The application of the wearable radiochemical sensing system described above in the electrochemical detection of uric acid involves using the wearable radiochemical sensing system to perform enzyme-free electrochemical quantitative analysis of uric acid in a sample without the use of uric acid oxidase. The sample is an artificial simulated body fluid or animal body fluid, wherein the artificial simulated body fluid includes at least one of artificial sweat and artificial urine, and the animal body fluid is the body fluid of livestock, poultry, or aquatic animals. This is used for monitoring uric acid metabolism in livestock, poultry, or aquatic animals in smart agriculture.

[0018] The beneficial effects of this invention are as follows: The working electrode of the enzyme-free uric acid electrochemical sensor of this invention uses a continuous monolayer of graphene oxide with a coverage of not less than 94.9% and a thickness of 1.4-1.6 nm as the active layer. By minimizing the interfacial electron tunneling distance and achieving favorable energy level matching, the inherent activity of the electrode drives the electrochemical oxidation of uric acid at a lower potential. Since the detection does not require uric acid oxidase, it avoids the problems of susceptibility to environmental inactivation and poor stability of enzymatic sensing. Furthermore, it significantly reduces costs because it does not rely on noble metal catalytic sites, thus achieving highly sensitive uric acid detection that combines stability and economy. This invention defines the active region of the working electrode using a laser-patterned flexible, highly oriented graphene film, providing a flexible and mass-producible electrode basis for sensors. This invention quantifies the aforementioned low overpotential advantage—the differential pulse voltammetric anodic oxidation peak of uric acid in pH 7.4 phosphate buffer is located at approximately 0.36 V, a negative shift of no less than 0.3 V compared to the bare graphene film, and the peak current monotonically increases with concentration within a range not exceeding 1000 μmol / L, providing a reliable and monotonic quantitative response basis in the low concentration range.

[0019] This invention uses lithium carbonate as a solubility enhancer to break through the inherent solubility limit of uric acid in phosphate buffer, and uses a two-stage splicing calibration with a current response deviation of less than 3% at a 1000 μmol / L connection concentration to cover an ultra-wide range of 0-5000 μmol / L. Thus, while maintaining the sub-micromolar detection limit, the linear range covers the clinical high concentration range, fundamentally breaking the inherent trade-off between "detection limit and linear range" in uric acid sensing. This invention uses Savitzky-Golay smoothed second-order numerical differentiation to deconvolve the severely overlapping voltammetric responses of ascorbic acid, dopamine, and uric acid into independent minimum points. Even in a strong imbalance background where the ascorbic acid concentration is 20 times that of dopamine, it can still distinguish and identify uric acid, thus meeting the anti-interference requirements of real samples without the use of enzymes or the introduction of additional modifications. Furthermore, this invention demonstrates the sensitivity advantage of this method with low detection limits of no more than 0.078 μmol / L (first segment) and 0.087 μmol / L (second segment), achieving highly sensitive quantification at the sub-0.1 μmol / L level.

[0020] The wearable radiochemical sensing system of the present invention combines the above-mentioned continuous single-layer sensing interface with a laser patterned three-electrode array, polyethylene terephthalate sandwich encapsulation, an elastomer microfluidic sampling module, and a flexible printed circuit board integrating electrochemical measurement circuitry, a low-power wireless communication module, and a power management unit. This transforms a desktop detection platform into a wearable terminal that can conform to the body surface, continuously monitor non-invasive bodily fluids in real time, and wirelessly transmit data to a smart terminal. Furthermore, its intrinsic sensing performance is effectively maintained even after undergoing system-level engineering transformations such as flexible encapsulation, microfluidic integration, and wireless transmission. The polydimethylsiloxane microfluidic chip of this invention uses a semi-circular main collection cavity combined with multiple inlet channels and a bottom outlet to ensure that the concentration of analytes in the cavity tends to be uniform within 15 seconds, meeting the requirements of second-level rapid detection in wearable scenarios for sufficient mass transfer and concentration uniformity of samples. This invention enables the system to be configured to monitor uric acid in sweat at 30-150 μmol / L using wireless timing amperometry or uric acid in urine at 2-5 mmol / L using differential pulse voltammetry, thereby adapting to the real-time monitoring needs of different body fluids and concentration ranges.

[0021] This invention applies the enzyme-free uric acid electrochemical sensor or wearable wireless radiochemical sensing system to the enzyme-free electrochemical quantification of uric acid in artificial simulated body fluids such as artificial sweat and artificial urine, verifying the effectiveness of its transformation from a desktop platform to a wearable system. Considering the objective existence of uric acid as the end product of animal purine metabolism and the intrinsic electrochemical response of this sensing interface to uric acid, this application can further be used for uric acid metabolism monitoring in livestock, poultry, and aquaculture animals in smart agriculture, providing an enzyme-free, portable, and wireless monitoring method for real-time detection and group health management at the breeding site.

[0022] Overall, this invention provides a comprehensive uric acid sensing solution that is enzyme-free, free of precious metals, has low overpotential, covers an ultra-wide range of 0-5000 μmol / L, has sub-micromolar detection limits, and can specifically identify uric acid and perform wearable wireless real-time monitoring under strong interference. This solution comprises an enzyme-free uric acid electrochemical sensor, an uric acid electrochemical detection method, a wearable wireless radiochemical sensing system, and their applications. Its beneficial effects are concentrated in the performance advantages of uric acid detection in terms of sensitivity, linear range, anti-interference selectivity, and wearable integration. In other words, it provides downstream sensor devices, detection methods, and systems for uric acid monitoring. The beneficial effects of this invention lie in the performance of uric acid detection and wearable monitoring, not in the manufacturing mechanism of the sensing interface itself, nor in the methodology of transforming the interface electronic structure into a sensor target screening method. Attached Figure Description

[0023] Figure 1 The images show the multi-scale morphology characterization of the GO-ML working electrode and the bare graphene film electrode using scanning electron microscopy (SEM) and atomic force microscopy-infrared (AFM-IR). Figure 2 AFM-IR characterization of bare graphene thin film electrodes (control); Figure 3 This is a cross-section of a high-resolution transmission electron microscope (HRTEM) image of GO-ML. Figure 4 The elemental distribution of the GO-ML energy dispersive spectroscopy (EDS) spectrum; Figure 5 The NEXAFS (Near-edge absorption fine structure) spectrum of carbon K-side in GO-ML and graphene films; Figure 6 The above figures are the X-ray photoelectron spectroscopy (XPS) C1s spectra of the two components; these figures are used to characterize the single-layer graphene oxide working electrode interface used in the sensor of this invention.

[0024] Figure 7 A comparison of the apparent standard heterogeneous electron transfer rate constants between GO-ML and control electrodes (GF, GO-PC, GO-FL); Figure 8 This is a complete band alignment diagram for four interface configurations with reference to the vacuum energy level. Figure 9 This is a pH-normalized effective energy level phase diagram (used to quantitatively screen uric acid as the optimal sensing target); the above diagram is used to illustrate the kinetics and energy level basis of the enzyme-free, low overpotential detection of uric acid by the sensor of this invention.

[0025] Figure 10 An evaluation diagram of the electrochemical double-layer capacitance (Cdl) of electrodes with different covering configurations; Figure 11The differential pulse voltammetry (DPV) response curve (0-5000 μM) of the GO-ML electrode to uric acid (UA) and the two-segment linear calibration curve are shown. Figure 12 The DPV response and calibration curves of three control electrodes (GF, GO-PC, and GO-FL) to 0-1000 μM UA are shown. Figure 13 The graph shows the selectivity of the GO-ML electrode (in a system where UA coexists with six potential interfering substances) and its 15-day stability evaluation. Figure 14 For urea, tyrosine, K + Na + Lactic acid, glucose, Ca²⁺ + Individual DPV response curves of each potential interfering substance in lithium carbonate; Figure 15 The original DPV curves and the second derivative decoupling curves after Savitzky–Golay smoothing are shown for the ternary mixture containing ascorbic acid (AA), dopamine (DA) and UA. Figure 16 This is a scatter plot comparing the performance of the sensor of this invention with that of the existing UA sensing platform. Figure 17 This is a diagram of the architecture of a wearable radio-chemical sensing system (system functional block diagram and actual flexible printed circuit board FPCB). Figure 18 Figures illustrating the structural design and mass production of flexible sensors (wafer-level array, PET sandwich encapsulation, bending state); Figure 19 The structure of the PDMS microfluidic sampling module and the finite element (FEA) simulation diagram of uric acid mass transfer are shown. Figure 20 Real-time UA sensing verification diagrams in artificial sweat (wireless timing amperometry) and artificial urine (DPV). Detailed Implementation

[0026] The present invention will be further described in detail below with reference to the accompanying drawings and specific embodiments to facilitate a clear understanding of the present invention, but these descriptions do not constitute a limitation on the present invention.

[0027] In all embodiments, GO refers to graphene oxide, the conductive layer (GF) of the graphene film refers to a flexible, highly oriented graphene film, the continuous monolayer graphene oxide active layer / GO-ML refers to a macroscopically continuous monolayer graphene oxide formed on a graphene film by limited proton source confinement electrophoresis, DPV refers to differential pulse voltammetry, PBS refers to phosphate buffered saline, PET refers to polyethylene terephthalate, and PDMS refers to polydimethylsiloxane. This invention presents four embodiments: an enzyme-free uric acid electrochemical sensor (Example 1), a uric acid electrochemical detection method (Example 2), a wearable radiochemical sensing system (Example 3), and its application (Example 4), each corresponding to four independent technical solutions, progressively unfolding from device structure and detection method to system integration.

[0028] All the above embodiments share the same working electrode interface, namely, a continuous monolayer graphene oxide active layer obtained by limited proton source confined anodic electrophoretic deposition on a flexible, highly oriented graphene film using the nonionic surfactant 2-naphthol polyoxyethylene ether (BNO-12). Its preparation involves: patterning the graphene film using laser micromachining to define the electrochemically active area (e.g., 1 × 1 cm²); impregnating it with a BNO-12 solution (5‰ mass / volume ratio) in ethanol for 8 h, followed by vacuum drying at 40 °C for 4 h to form an adsorption layer; using the pretreated graphene film as the anode, platinum foil as the cathode, and saturated potassium chloride silver / silver chloride as the reference electrode, at 0.5 mg·mL⁻¹… -1 Graphene oxide was deposited in a dispersion at a constant potential of 0.8 V (vs. Ag / AgCl) for 10 min. During the deposition process, irreversible anodic oxidation of BNO-12 released protons at a 1:1 proton-electron ratio as a finite and depletable proton source, and the anodic current decayed to near zero, thereby self-limiting the formation of a continuous monolayer of graphene oxide with a thickness of approximately 1.5 nm and a coverage of 94.9%.

[0029] The feasibility of the working electrode interface was confirmed through multi-scale characterization: SEM statistical surface coverage was 94.9% (see [link to relevant documentation]). Figure 1 (in coordinate system b); AFM-IR measurements showed nanosheet thicknesses of approximately 1.48 nm and 1.59 nm, with an average of approximately 1.5 nm (see coordinate system b). Figure 1 The coordinate system is c–d), while the bare graphene film control shows no significant morphological features or infrared absorption in the corresponding spectral region (see coordinate system c–d). Figure 2 ); HRTEM cross-sections show an approximately 1.5 nm amorphous interlayer with oxygen signals strictly confined within it (see...). Figure 3 , Figure 4 NEXAFS and XPS confirm that while introducing oxygen-containing functional groups such as C=O / O–C=O, the underlying graphene framework remains intact (see [link]). Figure 5 , Figure 6This interface compresses the heterogeneous electron transfer tunneling distance to the single-layer limit, which is the physical basis for the enzyme-free, low overpotential, and wide-range uric acid detection in various embodiments.

[0030] Example 1: Enzyme-free uric acid electrochemical sensor This embodiment of the enzyme-free uric acid electrochemical sensor comprises a three-electrode system consisting of a working electrode, a counter electrode, and a reference electrode. The working electrode includes a graphene film conductive layer and a continuous monolayer graphene oxide active layer thereon. The continuous monolayer graphene oxide active layer is formed by embedding monolayer graphene oxide nanosheets in a single-layer, flat, edge-connected, and substantially non-overlapping manner, with a thickness of 1.4-1.6 nm. Its coverage on the surface of the graphene film conductive layer is not less than 94.9%, and it directly contacts the graphene film conductive layer via π-π conjugation interactions. The sensor is configured to electrochemically detect uric acid in body fluid samples using differential pulse voltammetry without the use of uric acid oxidase and noble metal catalytic sites. Its working electrode interface is as described above in the common basic description and... Figure 1 , Figures 3-6 Representer.

[0031] Enzymatic uric acid sensing is susceptible to environmental inactivation and has poor stability, while most enzyme-free strategies rely on noble metal catalytic sites and are costly. According to the Marcus–Hush theory, the heterogeneous electron transfer rate decreases exponentially with tunneling distance (ln k0 ∝ -βd). Monolayer graphene oxide compresses the interfacial electron tunneling distance to its physical limit, increasing the apparent standard heterogeneous electron transfer rate constant kapp0 by approximately 44.9 times compared to bare graphene films (see [reference]). Figure 7 Simultaneously, its n-type narrow bandgap energy level arrangement (work function 4.61 eV, conduction band bottom) 3.94 eV, transmission bandgap 2.07 eV, see [reference] Figure 8 This provides a low-barrier electron-accepting channel for the highest occupied molecular orbital (HOMO) of uric acid, thereby driving uric acid oxidation through the inherent activity of the electrode. It achieves highly sensitive electrochemical oxidation detection of uric acid at lower potentials without the need for uricase or noble metal catalytic sites, combining stability and low cost.

[0032] The conductive layer of the graphene film is a highly oriented graphene film patterned by laser micromachining to define the electrochemical active region of the working electrode.

[0033] In a 0.1 mol / L phosphate buffer solution at pH 7.4, using a silver / silver chloride reference electrode filled with saturated potassium chloride solution as the potential reference, the differential pulse voltammetry anodic oxidation peak of the sensor for uric acid is located at 0.36 V. Under the same test conditions, the anodic oxidation peak is negatively shifted by no less than 0.3 V relative to the anodic oxidation peak of the bare graphene film conductive layer without a continuous monolayer graphene oxide active layer. Within the range of uric acid concentration greater than 0 and not exceeding 1000 μmol / L, the peak current of the anodic oxidation peak increases monotonically with increasing uric acid concentration.

[0034] The basis for this preferred performance is as follows: Figure 11 and Figure 12 The anodic oxidation peak of uric acid in GO-ML is located at approximately 0.36 V (vs. Ag / AgCl), while that of bare graphene film is approximately 0.67 V, GO-PC approximately 0.62 V (its positive bias is related to incomplete coverage leading to partial exposure of bare graphene film areas), and GO-FL approximately 0.35 V (although the peak position is close to GO-ML, the peak current increase in the high concentration range is slightly lower, reflecting the additional interlayer resistance introduced by multilayer stacking). GO-ML shows a negative shift of approximately 0.31 V (≥0.3 V) relative to bare graphene film, indicating a significant reduction in the overpotential driving force required for uric acid oxidation at this monolayer interface. In lithium carbonate-free PBS, the peak current monotonically increases with concentration in the range of 0-1000 μmol / L. (See calibration curves). Figure 12 The detection limits of the control electrodes obtained by df) showed an increasing sequence of GF < GO-PC < GO-FL < GO-ML, with GO-ML showing a significantly better detection limit of 0.078 μM than all control electrodes. This negative peak shift is consistent with the prediction of the n-type narrow bandgap energy level arrangement described in the common basis, and also with the pH-normalized effective energy level phase diagram (see [reference]). Figure 9 This confirms the conclusion that the uric acid kinetic overpotential η is the smallest (0.14 eV) and was quantitatively screened as the optimal sensing target.

[0035] Example 2: Electrochemical Detection Method for Uric Acid This embodiment provides an electrochemical detection method for uric acid, employing the enzyme-free uric acid electrochemical sensor described in Embodiment 1. It performs electrochemical quantification of uric acid in the sample without using uric acid oxidase, including: (1) Establish a two-segment splicing calibration relationship covering 0-5000 μmol / L: The differential pulse voltammetric anodic oxidation peak current of the working electrode was collected with a series of standard solutions with known uric acid concentrations. Based on this, a first linear calibration curve for quantification of uric acid not exceeding 1000 μmol / L was established in a phosphate buffer without lithium carbonate in the range of 0-1000 μmol / L. A second linear calibration curve for quantification of uric acid greater than 1000 μmol / L was established in a phosphate buffer with lithium carbonate added as a solubility enhancer in the range of greater than 1000-5000 μmol / L. The first and second calibration curves were connected at a junction concentration of 1000 μmol / L, and the current response deviation between the two at this junction concentration was less than 3%.

[0036] (2) Acquire the differential pulse voltammetric response curve of the sample to be tested in a phosphate buffer without lithium carbonate and obtain its anodic oxidation peak current; when the obtained peak current exceeds the current response corresponding to 1000 μmol / L of the first calibration curve, re-acquire the sample to be tested in a phosphate buffer with added lithium carbonate and obtain its anodic oxidation peak current.

[0037] (3) When the sample to be tested is collected in phosphate buffer without lithium carbonate, the peak current obtained is converted into uric acid concentration according to the first calibration curve; when the sample to be tested is collected in phosphate buffer with added lithium carbonate, the uric acid concentration is converted according to the second calibration curve.

[0038] Differential pulse voltammetry (DPV) has a high efficiency in distinguishing between Faraday current and charging current, and exhibits a significant signal-to-noise ratio advantage in the quantitative detection of low-concentration electroactive species, making it the high-sensitivity standard testing method adopted in this invention. The preferred DPV test parameters are: initial potential 0 V, termination potential 0.8 V, potential increment 0.004 V, pulse amplitude 0.05 V, pulse width 0.05 s, sampling width 0.0167 s, and pulse period 0.5 s. The test is conducted in a standard three-electrode system at room temperature, using PBS as the sensing electrolyte to simulate the physiological testing environment. Both calibration curves are linear fits, and their slopes represent the detection sensitivity of the corresponding segments. The main reagents used are listed in Table 1.

[0039] Table 1 Main experimental reagents The basis for the two-stage calibration is as follows: Figure 11As uric acid concentration increased from 0 to 5000 μM, the anodic oxidation peak current at approximately 0.36 V showed a continuous and monotonically increasing trend. Since the inherent solubility of uric acid in standard phosphate buffer is insufficient to cover the high clinical concentration range (>1000 μM), lithium carbonate was introduced as a solubility enhancer in the >1000 μM range. At the 1000 μM cross-concentration point, the current response deviation between the original PBS system and the lithium carbonate-enhanced system was less than 3%, confirming that the introduction of lithium carbonate did not reduce the detection confidence. Furthermore, the electrochemical active area (ECSA) of each electrode was quantitatively evaluated using the electrochemical double-layer capacitance method (see [link to relevant documentation]). Figure 10 Each electrode is positioned within a non-Radidatic potential window at a rate of 20-100 mV·s. - ¹ Cyclic voltammetry scans were performed at a scan rate. The slope obtained by linearly fitting the current density difference ΔJ to the scan rate is the Cdl of the electrode. The Cdl of GO-ML is 16.02 μF·cm. - ² (R² > 0.999), significantly higher than GO-PC and GO-FL, indicating that the conformal monolayer has the highest ECSA and maximizes the proportion of electrochemically accessible active sites, which is the structural basis for its high sensitivity and quantification.

[0040] The test sample is a sample containing ascorbic acid, dopamine, and uric acid. The differential pulse voltammetric response curve is smoothed using Savitzky-Golay second-order numerical differentiation, so that the oxidation contributions of the coexisting ascorbic acid, dopamine, and uric acid are expressed as independent minimum points on the resulting second-derivative curves. Based on this, the three substances can be distinguished without the use of enzymes, and the oxidation contribution corresponding to uric acid can be identified. This allows for the specific detection of uric acid in the presence of ascorbic acid and dopamine. Furthermore, even when the ascorbic acid concentration is not less than 20 times the dopamine concentration, the minimum points corresponding to uric acid and dopamine can still be distinguished and identified.

[0041] According to Figure 15 In a ternary system containing 2 mM ascorbic acid, 0.1 mM dopamine, and 0.2 mM uric acid, the original DPV curve shows a single broadened envelope peak at approximately 0.28 V, with the independent contributions of dopamine and uric acid masked by the strong background of ascorbic acid. After Savitzky-Golay smoothing and second-order numerical differentiation, the three components exhibit distinct and independent minimum points at approximately 0.21 V, 0.30 V, and 0.47 V, respectively, and are accurately distinguished. Even under extreme imbalances where the ascorbic acid concentration is 20 times that of dopamine, the dopamine and uric acid signals can still be independently identified. Each independent oxidation process exhibits superimposed Gaussian peaks in the current-potential relationship. The amplification effect of the second derivative on the peak curvature ensures that each process corresponds to a distinct minimum, thereby deconvolving severely overlapping peaks. Specific recognition under strong interference backgrounds is achieved without enzymes or additional modifications, meeting the anti-interference requirements of real physiological samples.

[0042] The detection limit for uric acid corresponding to the first linear calibration curve is no greater than 0.078 μmol / L, and the detection limit for uric acid corresponding to the second linear calibration curve is no greater than 0.087 μmol / L.

[0043] Further evidence of selectivity and stability: At the 50 μM level, the peak current of the DPV for uric acid was approximately 18 μA, while that for glucose, lactate, and sodium... + K + Ca² + The responses to urea were all below 3.5 μA and showed no significant difference from the baseline (see [reference]). Figure 13 a, Figure 14 Although tyrosine has an independent oxidation peak at approximately 0.65 V, it is far from the uric acid detection window and does not constitute substantial overlap; the peak current retention rate was higher than 93% and the relative standard deviation was only 2.60% in repeated tests over 15 consecutive days (see [reference]). Figure 13 (b) The stability stems from the strong π-π conjugation between the conformal monolayer GO and the graphene film, which inhibits nanosheet desorption and reconstruction. A comparison with the UA sensing platform in recent literature is shown in Table 3-3. Figure 16 This invention achieves a detection limit of 0.078 μM and a linearity limit of 5000 μM, both within the optimal range, with a concentration span of nearly five orders of magnitude, forming a significant Pareto frontier advantage.

[0044] Table 2 Comparison of detection performance of this work with existing technologies using UA electrochemical sensors (partial) Existing UA sensor platforms generally suffer from an inherent trade-off—sub-micromolar detection limits often come at the cost of compressing the upper limit of the linear range. This paper overcomes the solubility limitations of uric acid in PBS by using lithium carbonate as a solubility enhancer, and achieves coverage of 0-5000 μM with two segments spliced ​​together with a deviation of less than 3% at 1000 μM. Simultaneously, it achieves a sub-0.1 μM detection limit thanks to the high active site density and low mass transfer resistance of the conformal monolayer. By simultaneously achieving a sub-micromolar detection limit and an ultra-wide linear range covering the high concentration range in clinical applications without relying on noble metal catalysis or enzymatic amplification, this paper fundamentally breaks the aforementioned trade-off.

[0045] Example 3: Wearable Radio Chemical Sensing System The wearable radiochemical sensing system of this embodiment is configured to be detachably worn or attached to the body surface of the subject for non-invasive real-time electrochemical sensing of the subject's bodily fluids. It includes a three-electrode array consisting of a working electrode, a counter electrode, and a reference electrode, patterned on a graphene film via laser micromachining. The working electrode surface has a continuous monolayer graphene oxide active layer (composed of monolayer graphene oxide nanosheets laid flat in a single layer, with edges connected and substantially non-overlapping, forming a continuous layer with a coverage of not less than 94.9% and a thickness of 1.4-1.6). The reference electrode is composed of saturated silver / silver chloride paste coated on a graphene film; a top PET encapsulation layer and a bottom PET substrate encapsulate the three-electrode array and have electrochemical windows that expose the working area of ​​the three-electrode array; an elastomer microfluidic chip covering the electrochemical windows for collecting body fluids of the test object and bringing them into contact with the working electrode; and a flexible printed circuit board electrically connected to the three-electrode array via a flexible connector and integrating an electrochemical measurement circuit, a low-power wireless communication module and a power management unit, wherein the flexible printed circuit board is configured to wirelessly transmit the electrochemical signals collected by the three-electrode array to an external smart terminal.

[0046] Specifically (see) Figure 17 , Figure 18 The integrated flexible three-electrode device is fabricated on a graphene film using laser patterning. The contact pad width is 2.5 mm, the spacing between adjacent contact pads is 3.2 mm, the working electrode diameter is 4 mm, and the widths of the counter electrode and reference electrode are both 1 mm. The reference electrode is constructed by coating a saturated silver / silver chloride paste on a specific patterned area. Each sensor is encapsulated using a sandwich layer of top PET 1 (defining the electrode exposure window), a graphene film electrode layer modified with GO-ML, and a bottom PET 2 transparent substrate. The three-electrode array is connected to a miniaturized flexible printed circuit board via a flexible connector. This board integrates an analog front-end (AFE), an MS03 electrochemical chip (including I2C communication and EEPROM storage), a Bluetooth Low Energy System-on-Chip (BLE SoC), an RF module, and a power management unit (PM + Battery, including the battery). Data is wirelessly transmitted to a smartphone application via Bluetooth Low Energy, forming a complete data link from electrochemical signal acquisition to wireless transmission. The aforementioned signal acquisition and wireless transmission can be achieved using a miniature Bluetooth electrochemical workstation from Shenzhen Refresh Sensing Technology Co., Ltd.

[0047] Transforming desktop testing platforms into continuous monitoring terminals that conform to the contours of the human body is key to bridging the gap between basic interfaces and practical physiological monitoring, requiring that the intrinsic properties of the active interface be maintained after system-level integration. The inherent mechanical flexibility of graphene films provides a natural foundation for wearable technology; PET sandwich encapsulation protects the active interface from contamination while preserving the device's flexibility and conforming to the skin's curves (see...). Figure 18 (in the medium coordinate system c); laser etching technology can fabricate dozens of sensor arrays in parallel on a single graphene film in one go (see...) Figure 18 The GO-ML interface, using a mid-coordinate system (a), possesses significant scalability potential. After undergoing system-level engineering transformations such as flexible packaging, microfluidic integration, and wireless transmission, the wide dynamic range detection capabilities exhibited by its desktop platform have been effectively maintained.

[0048] The elastomer microfluidic chip is a polydimethylsiloxane microfluidic chip, which includes a semi-circular main collection chamber, multiple inlet channels communicating with the main collection chamber, and an outlet channel located at the bottom of the main collection chamber; the semi-circular main collection chamber is configured to make the concentration of the analyte in the chamber tend to be uniform within 15 seconds after the body fluid of the test object enters through the multiple inlet channels.

[0049] According to Figure 19 The PDMS module, formed by laser etching, can be directly attached to the skin to collect sweat and guide waste liquid out through the bottom outlet channel. Time-dependent finite element analysis was performed on the diffusion process of uric acid in the chamber (at four times: 0.85, 1.80, 3.65, and 14.90 s). At 0.85 s, the concentration gradient was concentrated near the channel opening. As time progressed (1.80 s and 3.65 s), the concentration front gradually expanded towards the center and bottom of the chamber. By 14.90 s, the entire chamber achieved a basically uniform concentration distribution (close to the inlet concentration of 0.08 μM). This indicates that the microfluidic module can complete the full mass transfer and concentration homogenization of the analyte within 15 s, meeting the requirements of second-level rapid detection in wearable scenarios.

[0050] The system is configured to perform electrochemical sensing of uric acid in the sweat of the test subject using a wireless time-amperometric method, with a concentration range covering 30-150 μmol / L, or to perform electrochemical sensing of uric acid in the urine of the test subject using a differential pulse voltammetry method, with a concentration range covering 2-5 mmol / L.

[0051] According to Figure 20 In the artificial sweat scenario, a wireless timed amperometry method was used to sequentially inject 30, 60, 90, 120, and 150 μM uric acid from a blank baseline. The current-time curves clearly showed the current transitions corresponding to each concentration step, and the steady-state plateaus between each step were well resolved (see [link]). Figure 20(a) The hyperuricemia risk threshold is marked with a dashed line as a physiological reference level. It maintains a distinguishable response even at concentrations significantly exceeding the typical physiological upper limit of sweat, providing ample detection margin for screening. In artificial urine scenarios, DPV is calibrated within a concentration range of 2-5 mM (see [reference]). Figure 20 (b, dashed line indicates the typical upper limit of clinical concentration) The current response is well linear with concentration (R²=0.965), confirming that the flexible sensor maintains reliable quantitative capability in the millimolecular high concentration range.

[0052] Example 4: Application in the electrochemical detection of uric acid This embodiment provides the application of the wearable radiochemical sensing system in the electrochemical detection of uric acid. The wearable radiochemical sensing system performs enzyme-free electrochemical quantitative analysis of uric acid in a sample without using uric acid oxidase. The sample is an artificial simulated body fluid or animal body fluid. The artificial simulated body fluid includes at least one of artificial sweat and artificial urine, and the animal body fluid is the body fluid of livestock, poultry, or aquaculture animals. This is used for monitoring uric acid metabolism in livestock, poultry, or aquaculture animals in smart agriculture.

[0053] Implementation basis see Figure 20 The flexible GO-ML sensor, integrated with a wireless circuit board and microfluidic patch, was tested in two simulated bodily fluids: artificial sweat (wireless timing amperometry, 30-150 μM) and artificial urine (DPV, 2-5 mM). The results demonstrate its effective transformation from a desktop platform to a wearable system. Uric acid is the end product of purine metabolism in livestock and poultry, and its metabolic abnormalities are associated with pathological conditions such as gout and visceral gout. Therefore, the following section further explains the application targets, sample types, and application scenarios of this sensor / system in the field of smart agriculture.

[0054] Within a reasonable scope, the application targets include livestock (such as pigs, cattle, and sheep), poultry (such as chickens, ducks, and geese), and aquaculture animals; among them, poultry mainly excretes uric acid, and its uric acid metabolism disorder is closely related to visceral gout (deposition of urate in the viscera and joints), making it particularly valuable for monitoring. The test sample can be body fluid samples such as blood or serum, urine, excrement extracts, and interstitial fluid from the above-mentioned animals, or it can be an extract / clarified liquid from aquaculture water, animal-derived food, or feed; for solid or high-matrix samples, conventional pretreatment such as dilution, filtration, or centrifugation can be performed before measurement. The sensor of this invention drives the electrochemical oxidation of uric acid through the inherent activity of the electrodes and has good universality to sample matrices, as demonstrated in the aforementioned artificial sweat and artificial urine (see...). Figure 20 This characterizes its ability to quantify uric acid in simulated physiological matrices, which can be extended to the detection of uric acid in the aforementioned agricultural body fluids and matrix samples.

[0055] In application, the differential pulse voltammetric response of the sample is collected according to the method described in Example 2 (or the wireless timing amperometry method is used for in situ / surface samples according to Example 3), and then the uric acid concentration is calculated according to the two-segment spliced ​​calibration curve. The ultra-wide linear range of 0-5000 μmol / L and the low detection limit of 0.078 μmol / L can cover the quantitative needs of uric acid in animal body fluids from low concentration to clinically high concentration. With the aforementioned flexible, patternable, laminated, and wirelessly transmittable device form, the sensor / system is particularly suitable for point-of-care testing (POCT) and wearable / portable continuous monitoring. Typical application scenarios include, but are not limited to: rapid on-site detection and continuous monitoring of uric acid levels in individual or group animals for early screening and assessment of the health status of groups of animals with hyperuricemia, gout, and visceral gout in poultry; precise feeding and nutritional regulation combined with purine intake; and quality monitoring of purine metabolites such as uric acid in aquaculture water, animal-derived food, and feed. It should be noted that the above situations are extensions of the application of the enzyme-free uric acid detection capability (including real-time detection verification in artificial simulated body fluids) verified by this invention in the field of smart agriculture. Its feasibility stems from the objective existence of uric acid as the end product of animal purine metabolism and the intrinsic electrochemical response of this sensing interface to uric acid.

[0056] As a specific application example for animal body fluids (based on the aforementioned validated enzyme-free uric acid detection capability, a feasible extension is made): Blood from livestock (e.g., pigs, cattle, sheep) or poultry (e.g., chickens, ducks, geese) is centrifuged to obtain serum, urine, or excrement. This serum is then extracted with phosphate buffer and filtered to obtain the extract, which serves as the test sample. If necessary, the sample is diluted with phosphate buffer at an appropriate ratio to ensure the uric acid concentration falls within the calibration range. Then, as described in Example 2: a differential pulse voltammetric response is first collected in phosphate buffer without lithium carbonate. If the obtained anodic oxidation peak current exceeds the current response corresponding to 1000 μmol / L in the first calibration curve, the sample is re-collected in phosphate buffer with added lithium carbonate. The uric acid concentration is then calculated based on the two-segment spliced ​​calibration curve and reduced by the dilution factor. For wearable or in-situ monitoring, continuous detection of body fluids collected from the body surface or in situ is performed using a wireless timing amperometric method, as described in Example 3. The obtained uric acid levels can be compared with the normal metabolic reference range for the corresponding species, and can be used for early screening of hyperuricemia, gout, and visceral gout in poultry, as well as for assessing the health status of the population. It should be noted that the detection of uric acid in animal body fluid samples described in this example is based on the enzyme-free uric acid detection capability of this invention (which has been verified in artificial simulated body fluids, see...). Figure 20 The feasible application of this technology in the field of smart agriculture is consistent with the aforementioned embodiments in terms of uric acid response mechanism and quantification method.

[0057] Based on the above embodiments, the differences between the present invention and the prior art are as follows: 1. An enzyme-free and precious metal-free uric acid electrochemical sensor was constructed using a monolayer graphene oxide interface. By minimizing the interfacial tunneling distance (44.9 times increase in kapp0) and the n-type narrow bandgap energy level arrangement, highly sensitive uric acid oxidation detection was achieved at a lower potential (approximately 0.36 V, approximately 0.31 V negative shift compared to the bare substrate) without using uric acid oxidase or precious metal catalytic sites, while also possessing stability and low cost.

[0058] 2. A two-stage splicing calibration method with lithium carbonate as a solubility enhancer (<3% deviation at 1000 μM) and Savitzky–Golay second derivative decoupling are proposed, which simultaneously achieves submicromolar detection limits (0.078 / 0.087 μM) and an ultra-wide linear range covering 0-5000 μM. It also maintains specificity under strong interference from ascorbic acid up to 20 times that of dopamine, breaking the inherent trade-off (Pareto front) between the detection limit and linear range of UA sensors.

[0059] 3. The single-layer interface was integrated into a flexible wearable system that includes PDMS microfluidic sampling (<15 s concentration homogenization) and low-power Bluetooth wireless transmission. This verified the effective retention of intrinsic sensing performance after system-level engineering transformation and enabled real-time wireless detection in artificial sweat (30-150 μM, chronoamperometry) and artificial urine (2-5 mM, DPV), completing the entire chain from atomic-level interface construction to wearable real-time monitoring.

[0060] The contents not described in detail in this specification are existing technologies known to those skilled in the art.

Claims

1. An enzyme-free uric acid electrochemical sensor, characterized in that, The system comprises a three-electrode system consisting of a working electrode, a counter electrode, and a reference electrode. The working electrode includes a graphene film conductive layer and a continuous monolayer graphene oxide active layer located on the graphene film conductive layer. The continuous monolayer graphene oxide active layer is formed by embedding monolayer graphene oxide nanosheets in a single-layer, flat, edge-connected, and substantially non-overlapping manner, with a thickness of 1.4-1.6 nm. The active layer has a coverage of not less than 94.9% on the surface of the graphene film conductive layer and is in direct contact with the graphene film conductive layer via π-π conjugation interactions. The enzyme-free uric acid electrochemical sensor is configured to perform electrochemical detection of uric acid in body fluid samples using differential pulse voltammetry without using uric acid oxidase and noble metal catalytic sites.

2. The enzyme-free uric acid electrochemical sensor according to claim 1, characterized in that, The conductive layer of the graphene film is a highly oriented graphene film patterned by laser micromachining to define the electrochemical active region of the working electrode.

3. The enzyme-free uric acid electrochemical sensor according to claim 1, characterized in that, In a 0.1 mol / L phosphate buffer solution at pH 7.4, using a silver / silver chloride reference electrode filled with saturated potassium chloride solution as the potential reference, the differential pulse voltammetric anodic oxidation peak of the enzyme-free uric acid electrochemical sensor for uric acid is located at 0.36 V. Under the same test conditions, the anodic oxidation peak is negatively shifted by no less than 0.3 V relative to the anodic oxidation peak of the bare graphene film conductive layer without the deposited continuous monolayer graphene oxide active layer for uric acid. Within the range of uric acid concentration greater than 0 and not exceeding 1000 μmol / L, the peak current of the anodic oxidation peak increases monotonically with increasing uric acid concentration.

4. A method for electrochemical detection of uric acid, characterized in that, Using the enzyme-free uric acid electrochemical sensor as described in claim 1, electrochemical quantification of uric acid in the sample to be tested is performed without the use of uric acid oxidase, comprising: (1) Establish a two-segment splicing calibration relationship covering 0-5000 μmol / L: The differential pulse voltammetric anodic oxidation peak current of the working electrode is collected with a series of standard solutions with known uric acid concentrations. Based on this, a first linear calibration curve for quantification of uric acid not exceeding 1000 μmol / L is established in a phosphate buffer without lithium carbonate in the range of 0-1000 μmol / L uric acid concentration. A second linear calibration curve for quantification of uric acid greater than 1000 μmol / L is established in a phosphate buffer with lithium carbonate added as a solubility enhancer in the range of greater than 1000-5000 μmol / L uric acid concentration. The first linear calibration curve and the second linear calibration curve are connected at a connection concentration of 1000 μmol / L, and the current response deviation between the two at the connection concentration is less than 3%. (2) Acquire a differential pulse voltammetric response curve for the sample to be tested in a phosphate buffer solution without lithium carbonate, and obtain the anodic oxidation peak current of the sample to be tested from the differential pulse voltammetric response curve; when the obtained anodic oxidation peak current exceeds the current response corresponding to 1000 μmol / L of the first linear calibration curve, re-acquire a differential pulse voltammetric response curve for the sample to be tested in a phosphate buffer solution with added lithium carbonate, and obtain the anodic oxidation peak current of the sample to be tested from it; (3) When the sample to be tested is collected in a phosphate buffer without lithium carbonate, the obtained anodic oxidation peak current is converted into the uric acid concentration in the sample to be tested according to the first linear calibration curve; when the sample to be tested is collected in a phosphate buffer with added lithium carbonate, the obtained anodic oxidation peak current is converted into the uric acid concentration in the sample to be tested according to the second linear calibration curve.

5. The electrochemical detection method for uric acid according to claim 4, characterized in that, The test sample is a sample containing ascorbic acid, dopamine, and uric acid. The differential pulse voltammetric response curves obtained from the collected test samples are processed using Savitzky-Golay smoothing followed by second-order numerical differentiation. This allows the oxidation contributions of the coexisting ascorbic acid, dopamine, and uric acid to be represented as independent minimum points on the resulting second-derivative curves. Based on this, ascorbic acid, dopamine, and uric acid can be distinguished without the use of enzymes, and the oxidation contribution corresponding to uric acid can be identified. This enables the specific detection of uric acid in the presence of ascorbic acid and dopamine. Furthermore, even when the concentration of ascorbic acid is not less than 20 times the concentration of dopamine, the minimum points corresponding to uric acid and dopamine can still be distinguished and identified.

6. The electrochemical detection method for uric acid according to claim 4, characterized in that, The detection limit of uric acid corresponding to the first linear calibration curve is no greater than 0.078 μmol / L, and the detection limit of uric acid corresponding to the second linear calibration curve is no greater than 0.087 μmol / L.

7. A wearable radiochemical sensing system, wherein the wearable radiochemical sensing system is configured to be detachably worn or attached to the body surface of a subject to be tested, for non-invasive real-time electrochemical sensing of the body fluids of the subject to be tested, characterized in that, include: A three-electrode array, comprising a working electrode, a counter electrode, and a reference electrode, is patterned on a graphene film using laser micromachining. The working electrode has a continuous monolayer graphene oxide active layer, which is formed by embedding monolayer graphene oxide nanosheets in a single-layer, flat, edge-connected, and substantially non-overlapping manner, with a coverage of not less than 94.9% and a thickness of 1.4-1.6 nm on the working electrode surface. The reference electrode is composed of a saturated silver / silver chloride paste coated on the graphene film. A top polyethylene terephthalate encapsulation layer and a bottom polyethylene terephthalate substrate encapsulate the three-electrode array and provide electrochemical windows exposing the working area of ​​the three-electrode array. An elastomer microfluidic chip covers the electrochemical windows and is used to collect the body fluid of the test subject and bring the collected body fluid into contact with the working electrode. And a flexible printed circuit board that is electrically connected to the three-electrode array via a flexible connector and integrates an electrochemical measurement circuit, a low-power wireless communication module and a power management unit, the flexible printed circuit board being configured to wirelessly transmit the electrochemical signals collected by the three-electrode array to an external smart terminal.

8. The wearable radio-chemical sensing system according to claim 7, characterized in that, The elastomer microfluidic chip is a polydimethylsiloxane microfluidic chip, which includes a semi-circular main collection cavity, multiple inlet channels communicating with the semi-circular main collection cavity, and an outlet channel located at the bottom of the semi-circular main collection cavity. The semi-circular main collection chamber is configured such that after the body fluid of the test subject enters through the multiple inlet channels, the concentration of the analyte in the semi-circular main collection chamber tends to be uniform within 15 seconds.

9. The wearable radio-chemical sensing system according to claim 7, characterized in that, The wearable radiochemical sensing system is configured to electrochemically sense uric acid in the sweat of the test subject using a wireless time-amperometric method, with the concentration range of the electrochemical sensing covering 30-150 μmol / L, or configured to electrochemically sense uric acid in the urine of the test subject using a differential pulse voltammetry method, with the concentration range of the electrochemical sensing covering 2-5 mmol / L.

10. The application of the wearable radiochemical sensing system according to claim 7 in the electrochemical detection of uric acid, characterized in that, The wearable radiochemical sensing system described above performs enzyme-free electrochemical quantitative analysis of uric acid in a test sample without using uric acid oxidase. The test sample is an artificial simulated body fluid or animal body fluid, wherein the artificial simulated body fluid includes at least one of artificial sweat and artificial urine, and the animal body fluid is the body fluid of livestock, poultry, or aquatic animals. This system is used for monitoring uric acid metabolism in livestock, poultry, or aquatic animals in smart agriculture.

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