Glucose biosensor based on two-electrode configuration

Abstract

The present application relates to a glucose biosensor based on a two-electrode structure, and relates to the technical field of biosensors, wherein the sensor adopts a ceramic substrate, a two-electrode system composed of a platinum working electrode and a silver / silver chloride reference electrode is integrated on the surface of the ceramic substrate, a non-functional area is covered by a blue epoxy resin isolation layer, and a glucose oxidase layer cured by cross-linked amino resin is fixed on the surface of the working electrode. The preparation process comprises the following steps: laser cutting a 4*4 inch wafer to a 134-unit profile, sequentially screen printing Ag / Pd conductive paths, Pt / AgCl electrodes and an isolation layer, and finally dispensing an enzyme layer in a ten-thousand-level clean room and curing at room temperature. The scheme eliminates bubble interference through a two-electrode structure, combines the anti-hemolysis characteristics of the isolation layer and the wafer-level mass production process, realizes the core advantages of reducing the detection error rate by 60%, reserving enzyme activity by greater than or equal to 95%, and reducing the unit cost by 50%, and provides a high-precision, high-throughput glucose dynamic monitoring solution for clinical use.

Description

Technical Field

[0001] This invention relates to the field of biosensor technology, specifically to a glucose biosensor based on a dual-electrode structure. Background Technology

[0002] Current glucose biosensors generally employ a three-electrode system (working electrode / reference electrode / counter electrode). This system's dielectric design is susceptible to interference from air bubbles in the fluid, leading to potential drift >15%. Especially in dynamic detection scenarios, electrode spacing deviations cause electrolyte shunting, significantly reducing signal stability. The three-electrode detection process requires an additional dielectric layer to maintain potential balance, increasing structural complexity and failure rate.

[0003] Insufficient mass production efficiency and biocompatibility: Current production processes mainly rely on single-piece screen printing, yielding less than 30 sensors per wafer. Furthermore, enzyme immobilization processes depend on high-temperature sintering, resulting in a glucose oxidase activity loss of >20%. The isolation layer material has weak resistance to biocontamination; after contact with hemolyzed samples, the swelling rate exceeds 10%, and resistance fluctuation reaches 5%, severely shortening its lifespan. Traditional sensor dimensional tolerance control is loose (±0.3mm), further reducing detection consistency.

[0004] Furthermore, the clinical validation system is not yet closed-loop: traditional protocols lack systematic performance validation, resulting in: slow response speed (>10 seconds) which is insufficient for emergency needs; weak anti-interference capability (uric acid (1400 μmol / L) signal deviation >10%, exceeding the clinical threshold); reliance on cold chain for transportation (failed to overcome the 2-8℃ storage limit, increasing the cost of application in primary healthcare). High-concentration sample rinsing time >180 seconds severely restricts testing throughput.

[0005] Therefore, a glucose biosensor based on a dual-electrode structure is needed to solve the above problems. Summary of the Invention

[0006] To address the problems of the prior art, this invention provides a glucose biosensor based on a dual-electrode structure.

[0007] To solve the above-mentioned technical problems, the present invention is achieved through the following technical solution: In the first aspect, a glucose biosensor based on a dual-electrode structure includes: a ceramic substrate with dimensions of 25.4±0.1mm×7.00±0.1mm×0.63±0.05mm;

[0008] A dual-electrode system screen-printed onto a ceramic substrate:

[0009] The dual-electrode system includes: an area of ​​1.02 mm². 2 ±5% platinum working electrode,

[0010] Silver / silver chloride reference electrode;

[0011] The platinum working electrode and the silver / silver chloride reference electrode are respectively connected to a silver-palladium alloy conductive path;

[0012] The platinum working electrode and the silver / silver chloride reference electrode are arranged adjacent to each other, with an edge spacing tolerance of ±0.2 mm;

[0013] A blue epoxy resin isolation layer covering the non-electrode area; the blue epoxy resin isolation layer material is type 118-08 epoxy resin.

[0014] The glucose oxidase layer immobilized on the surface of the working electrode is made of water-based acrylic-modified polyester. AY5537w / 35WA cross-linked amino resin is cured and formed;

[0015] The dual-electrode system is configured for current densities ≤10 nA / mm². 2 The response time for detecting hydrogen peroxide is ≤4 seconds.

[0016] This application addresses three core pain points of traditional glucose biosensors in clinical testing through a triple breakthrough: innovative dual-electrode structure, composite functional material design, and efficient mass production process. These pain points include insufficient signal stability (fluid bubble interference), weak resistance to biocontamination (interference from hemolyzed samples), and excessively high unit production cost.

[0017] Abandoning the traditional three-electrode structure, a dual-electrode design is adopted, consisting of a platinum working electrode and a silver / silver chloride reference electrode, achieving nA / mm 2 At low current density (≤10nA / mm) 2 This maintains potential stability and significantly reduces the risk of bubble interference in fluid systems (60% reduction in error rate compared to a three-electrode structure). The underlying technology is as follows:

[0018] Simplified detection circuit: The electrode and dielectric system are omitted, and the electrode spacing is precisely controlled to ±0.2mm to avoid signal drift caused by electrolyte shunting;

[0019] Rapid response mechanism: Enzyme-amperometry for detecting hydrogen peroxide byproducts, with a response time of ≤4 seconds, meets the needs of high-throughput scenarios such as emergency testing.

[0020] Blue epoxy resin isolation layer (type 118-08) and The combination of enzyme carriers forms a dual barrier against hemolytic contamination and protects enzyme activity. Key properties of the isolation layer include: resistance to biocorrosion: no swelling / exudation after 24 hours of contact with hemolyzed blood at 37°C; ease of operation: blue markings reduce the error rate; and electrical stability: contact resistance change with the system solution ≤1%.

[0021] Innovation in enzyme immobilization technology: Waterborne acrylic modified polyester crosslinked amino resin, cured in a clean room environment at 25±2℃ (humidity 45±5%), with an enzyme activity retention rate of ≥95%, supporting ≥100 consecutive tests.

[0022] High-efficiency mass production of 134 sensors per wafer is achieved through 4×4-inch wafer laser cutting and three-layer screen printing processes:

[0023] Precision manufacturing control:

[0024] Laser cutting depth is 0.3mm ± 0.05mm to ensure the integrity of the substrate structure;

[0025] The alignment accuracy between screen printing layers is ±0.05mm, and the sequential printing of the conductive path from the electrode layer to the isolation layer ensures the compatibility of the functional layers.

[0026] Automated enzyme layer distribution: Distribution error in Class 10,000 cleanroom is ≤2%, enzyme layer thickness uniformity deviation is <5%, eliminating batch differences from the source.

[0027] In one specific embodiment of the first aspect, the isolation layer satisfies the following condition: no swelling or exudate occurs after 24 hours of contact with a hemolyzed blood sample at 37°C.

[0028] The technical solution achieves clinical-grade reliability through a rigorous validation system.

[0029] Anti-interference performance: Uric acid (1400μmol / L) signal deviation ≤5.5%, significantly better than the industry 10% threshold;

[0030] Convenience of transportation: The blister packaging can withstand transportation at 40℃ for 48 hours, reducing application costs;

[0031] Accuracy Validation: Correlation between test results of 80 patient samples and the YSI-2300 reference instrument R 2 =0.992.

[0032] In one specific embodiment of the first aspect, the curing conditions for the glucose oxidase layer are: a clean room environment at 25±2℃, a relative humidity of 45±5%, and a time of ≤2 hours.

[0033] Secondly, the fabrication method of a glucose biosensor based on a dual-electrode structure includes:

[0034] S1: Laser pre-cutting of 134 sensor unit outlines on a 4×4 inch ceramic wafer;

[0035] S2: Sequential screen printing:

[0036] First layer: Silver-palladium alloy conductive path material, model number: EL44-040.

[0037] The second layer consists of platinum working electrode paste of model E1192 and silver / silver chloride reference electrode paste of model 5876.

[0038] Third layer: Blue epoxy resin isolation layer covers non-functional areas;

[0039] S3: Distribute the enzyme-polymer mixture onto the surface of the working electrode in a Class 10,000 cleanroom, and solidify it at room temperature to form an enzyme layer.

[0040] In one specific embodiment of the second aspect, the alignment accuracy between screen printing layers is ±0.05mm, and the enzyme layer coverage area tolerance is ±2%.

[0041] The beneficial effects of this invention are as follows:

[0042] 1. By employing a dual-electrode system with a platinum working electrode and a silver / silver chloride reference electrode, at a current density ≤10 nA / mm², 2 This method achieves stable potential control under certain conditions, solving the signal drift problem caused by bubble interference in traditional three-electrode structures. Its core advantages are: improved response speed (≤4 seconds for enzyme-amperometric detection of hydrogen peroxide byproducts, meeting the needs of high-throughput scenarios such as emergency medicine); simplified structure leading to increased stability (precisely controlled electrode spacing within ±0.2 mm, avoiding signal distortion caused by electrolyte shunting); and a wide linear range covering clinical needs (linear response within the 1-50 mmol / L glucose concentration range (R0). 2 >0.99), effectively covering high-value pathological samples. This design has been validated in multiple batches, with signal stability error <3% at a high concentration of 40 mmol / L, providing a reliable data foundation for intensive care.

[0043] 2. Blue epoxy resin isolation layer and The innovative combination of enzyme immobilization carriers provides a dual guarantee of biocompatibility and mass production efficiency: breakthroughs in anti-biocontamination performance, with the isolation layer showing no swelling or exudation after 24 hours of immersion in hemolyzed blood at 37°C, a resistance change of ≤1%, and a 300% extension of service life; long-term enzyme activity retention technology, cured in a cleanroom at 25±2°C (humidity 45±5%), ensures an enzyme activity retention rate of ≥95%, supporting ≥100 consecutive tests; 4×4-inch wafer laser cutting of 134 units, combined with a three-layer screen printing sequence, achieves interlayer alignment accuracy of ±0.05mm, reducing unit production costs by 60%. This process automatically controls enzyme layer thickness deviation to <5%, eliminating batch differences from the source and ensuring consistency in large-scale production;

[0044] 3. The feasibility of its clinical application was verified through multi-dimensional validation: Excellent anti-interference performance, with a uric acid (1400 μmol / L) signal deviation ≤5.5%, significantly better than the 10% clinical threshold; Breakthrough tolerance to extreme environments, with signal attenuation ≤5% after 48 hours of transport at 40℃ under blister packaging; Empirical evidence of diagnostic accuracy, with correlation R between the test results of 80 patient samples and the YSI-2300 reference instrument. 2 =0.992, with an error distribution conforming to CLSI EP09-A3 guidelines. Combined with rapid wetting in 30 minutes and a sample throughput of >100 samples per hour, this sensor provides an easy-to-use solution for primary healthcare and home-based dynamic monitoring, helping to improve the efficiency of diabetes management. Attached Figure Description

[0045] Figure 1 This is a schematic diagram of the structure of the present invention.

[0046] Figure 2 This is a dimensional schematic diagram of the present invention.

[0047] Figure 3 This is a schematic diagram of the enzyme amperometric biosensor principle of the present invention.

[0048] Figure 4 This is a schematic diagram of the signal drift of the four sensors in this invention.

[0049] Figure 5 This is a schematic diagram of the basic current signals of the four sensors of the present invention during the initial wetting stage.

[0050] Figure 6 This is a schematic diagram of the sensor response for a glucose sample within the normal range (5.0 mmol / L) of this invention.

[0051] Figure 7 This is a schematic diagram of the rinsing time for sensor recovery after measuring 5 mmol / L glucose according to the present invention.

[0052] Figure 8 This is a schematic diagram of the sensor recovery flushing time after measuring 40 mmol / L glucose according to the present invention.

[0053] Figure 9 This is a schematic diagram of the linearity of the four sensors of the present invention, wherein (a) shows the relationship between the current change rate of sensor 1 and glucose concentration, (b) shows the relationship between the current change rate of sensor 2 and glucose concentration, (c) shows the relationship between the current change rate of sensor 3 and glucose concentration, and (d) shows the relationship between the current change rate of sensor 4 and glucose concentration.

[0054] Figure 10This is a schematic diagram of the VS YSI-2300 measurement values ​​of the control samples for batches #CNB001, #CNB002 and #CNB003 of the present invention.

[0055] Figure 11 This is a schematic diagram of the patient sample measurement values ​​VS YSI-2300 for batch #CNB001 of the present invention.

[0056] Figure 12 This is a schematic diagram of the patient sample measurement VS YSI-2300 for batch #CNB002 of the present invention.

[0057] Figure 13 This is a schematic diagram of the patient sample measurement values ​​VS YSI-2300 for batch #CNB003 of the present invention.

[0058] Figures 1 to 13 In the middle: 1. Ceramic substrate; 2. Blue epoxy resin isolation layer; 3. Silver / silver chloride reference electrode; 4. Platinum working electrode; 5. Silver-palladium alloy conductive path. Detailed Implementation

[0059] The technical solutions of the present invention will be clearly and completely described below with reference to the embodiments of the present invention. Obviously, the described embodiments are only some embodiments of the present invention, and not all embodiments. Based on the embodiments of the present invention, all other embodiments obtained by those of ordinary skill in the art without creative effort are within the scope of protection of the present invention.

[0060] like Figures 1 to 13 The glucose biosensor based on a dual-electrode structure shown includes: a ceramic substrate 1 with dimensions of 25.4±0.1mm×7.00±0.1mm×0.63±0.05mm;

[0061] A dual-electrode system screen-printed on the surface of ceramic substrate 1:

[0062] The dual-electrode system includes: an area of ​​1.02 mm². 2 A platinum working electrode 4 with ±5% platinum chloride and a silver / silver chloride reference electrode 3 are connected to the platinum working electrode 4 and the silver / silver chloride reference electrode 3 respectively by a silver-palladium alloy conductive path 5. The platinum working electrode 4 and the silver / silver chloride reference electrode 3 are arranged adjacent to each other with an edge spacing tolerance of ±0.2mm. A blue epoxy resin isolation layer 2 covers the non-electrode area. The material of the blue epoxy resin isolation layer 2 is 118-08 type epoxy resin.

[0063] The glucose oxidase layer immobilized on the surface of the working electrode is made of water-based acrylic-modified polyester. AY5537w / 35WA cross-linked amino resin is cured and formed;

[0064] The dual-electrode system is configured for current densities ≤10 nA / mm². 2 The response time for detecting hydrogen peroxide is ≤4 seconds.

[0065] The principle of this multi-purpose biosensor.

[0066] Please refer to Figure 3 As shown, the measurement cycle includes: 1. Contact between the sensor and the liquid sample; 2. Transfer of glucose to the enzyme layer; 3. Enzyme-catalyzed reaction of glucose under aerobic conditions; detection of hydrogen peroxide as a byproduct of the reaction on the detector; 4. Recovery of the sensor by rinsing and removing the reaction components. Figure 3 The entire measurement cycle is displayed.

[0067] The main feature of multi-purpose biosensors is the stable immobilization of receptor components, enabling multi-cycle measurements.

[0068] The electrochemical detector consists of a dual-electrode structure, including a platinum working electrode 4 and a silver / silver chloride reference electrode 3.

[0069] A simple two-electrode structure is a more suitable detection tool because the current density of biosensors is very small (a few nA / mm). 2 (within a certain range), it will not affect the electrode potential of the Ag / AgCl reference electrode.

[0070] Please refer to Figure 2 The geometry and dimensions are shown.

[0071] The sensor format specifications are as follows: The sensor should be screen-printed on a 4x4 inch wafer, which is the most common standard substrate format for screen printing. The required sensor geometry results in 13x4 individual sensors on a laser-cut wafer.

[0072] Sensor geometry:

[0073] Length: 25.4mm / Tolerance: ±0.1mm after separation;

[0074] Width: 7.00mm ± 0.1mm / Tolerance: ± 0.1mm after separation;

[0075] Thickness: 0.63mm ± 0.05mm;

[0076] The area of ​​the working electrode is 1.02 mm². 2 ;

[0077] Screen printing tolerances:

[0078] Measurement: ±0.05 mm;

[0079] Minimum distance between two sides: ±0.2mm;

[0080] Structural thickness: specified by the screen printing ink manufacturer;

[0081] Material Selection: The isolator layer is defined as blue. This will affect the compatibility of the isolator print adhesive with all other printed components.

[0082] Ceramic substrate material: Ceramic substrate 1 is a standard material for thick-film screen printing, formatted as 4x4 inches (101.6 mm) with a thickness of 0.630 mm. This size is optimal because most screen printing equipment accommodates this format. The substrate employs a laser-etched external sensor geometry, comprising 13x4 sensors, each measuring 25.40x7.00 mm.

[0083] Conductor path: The conductor path and contact pads are composed of a silver-palladium alloy (Ag / Pd). This is a common standard technology with no special requirements for further development.

[0084] Metallic Electrode: The described sensor is based on a two-electrode measurement system. Unlike the 3-electrode measurement system where there is no current load on the reference electrode, in the 2-electrode system, current flows through the Ag / AgCl electrode. However, at extremely small nA / mm... 2 Low current density loads within the range do not change the reference electrode potential. Therefore, the advantages of the simple and robust 2-elctrode structure in the flow system become apparent.

[0085] The metal electrodes used in the dual-electrode electrochemical detector structure are made of platinum (working electrode) and silver / silver chloride (reference electrode), respectively. The corresponding materials were selected based on long-term experience.

[0086] Isolation layer: The isolator covers the conductive path and other areas of the thick film structure, making it easier for users to operate.

[0087] The most important requirements for the isolator are: (a) good printability; (b) a clearly defined geometry and smooth surface after printing; (c) blue color; (d) high solvent resistance; (e) chemical resistance in contact with the applied liquids, such as system solutions, calibrators, hemolyzed blood samples, and other biologically derived samples; (f) no swelling during use; (g) no leakage of any contents during use; and (h) stable resistance in contact with the applied liquid. Finally, a 2k epoxy resin was selected as the printing ink for the isolator. Performance (d)-(h) was tested iteratively, and the performance of the biosensor in use was also investigated.

[0088] Enzyme: Glucose oxidase (GOx) from Aspergillus niger (EC1.1.3.4.) was used as a receptor for glucose biosensors.

[0089] Polymer: To immobilize GOx enzymes, waterborne acrylic-modified polyesters crosslinked with amino resins were used in waterborne emulsions.

[0090] Description of the production process: The production process consists of two main parts: 1. Manufacturing of the basic sensor structure in the screen printing process; 2. Fixing the receptor components during the formulation process.

[0091] Screen printing process: The first production step is a third-party manufacturer service, based on customized production according to specifications. The basic sensor structure undergoes pre-delivery inspection and is then delivered in a 4x4-inch format.

[0092] Enzyme Immobilization Process: The second production step involves the immobilization of the biosensor receptor assembly. A measured mass of lyophilized GOx is mixed with a polymer-based emulsion. The liquid prepared in this way is then dispensed using an automated dispenser under cleanroom conditions. The drying and hardening of the enzyme-polymer layer are also controlled under cleanroom conditions. Preliminary tests are then performed on the sensor wettability and sensor signal (sensitivity). After confirmation, the sensor is packaged in blister packs and then stored in a refrigerator.

[0093] Bill of Materials (BOM)

[0094] Table 1 shows the bill of materials.

[0095]

[0096] Table 1: Bill of Materials (BOM)

[0097] Biosensor performance testing: Biosensor performance testing encompasses all processes following sensor manufacturing. Typically, these processes cover storage and key functional testing.

[0098] The investigation included three different sensor batches (#CNB001, #CNB002, #CNB003), each containing a total of 51 sensors. Seven samples were randomly selected from each batch for different tests. The sensor was designed for use with hemolyzed whole blood samples pre-diluted 1:51 with system fluid. Concentrations given herein always refer to this dilution rate. Unless otherwise stated.

[0099] Wetting Time: Wetting time is the necessary time span between initial insertion of the sensor into the analyzer and operational readiness. This "delay" is due to the conditioning process that begins from the sensor's first contact with the system solution. The sensor layer absorbs water and electrolyte components and expands. During this period, the sensor is not yet ready for measurement because the signal is not stable enough to properly meet sensor specifications. A 30-minute time span is reasonable for wetting.

[0100] Please refer to Figure 4 The graph shows the signal drift of four sensors (batch number #CNB003). It compares the normalized signals (current change rate dI / dt) of the four sensors at time points of 180 seconds, 480 seconds, 780 seconds, 1080 seconds, 1380 seconds, 1680 seconds, 1980 seconds, and 2280 seconds. It can be used to observe the consistency and stability of the sensor signals.

[0101] The wetting process of the glucose biosensor (four sensors, numbers 1-4, from the same batch, batch #CNB003) is exemplarily shown in the figure. Figure 4 In the initial wetting phase of 3 minutes, samples of 12.0 mmol / L glucose were measured every 5 minutes, and signal drift was assessed. A key criterion for proper wetting was the signal drift at the end of the wetting period. The signal was set to 100% after 3 minutes, with a signal drift of 35 minutes. After wetting, the drift should not exceed 3%, for more than 20 samples (one complete sample rack), or drift within 10 minutes individually. Clearly, after 1680 seconds (28 minutes), the signal drift was already within the <2% range for the next 10 minutes, as shown in all four sensors.

[0102] Sensor base current: The base current of glucose biosensors is typically in the low nA / mm² range. 2 Measurements are taken within the specified range. Only in the initial period after the sensor comes into contact with the system solution will the sensor base current show a higher value. This is due to charge reversal and oxidation processes in the dense electrode region. The first cycle is exemplarily as follows: Figure 5 As shown. After 3 minutes, the typical base current is <10 nA / mm. 2 .

[0103] Please refer to Figure 5 The figure shows the basic current signals (0 to 180s) of the four sensors during the initial wetting phase. It illustrates the current of sensor 1 (linear), sensor 2 (linear), sensor 3 (linear), and sensor 4 (linear) in batch #CNB003. The current of each sensor rises rapidly in the initial phase (approximately 0-50 seconds), then gradually decreases and tends to stabilize. The current change trends of different sensors vary to some extent.

[0104] Sensor signal value and accuracy: The signal strength of a biosensor is a crucial parameter for signal resolution and accuracy. Long-term experience clearly shows that a sensor signal strength between 0.5 and 3 nA / ((mmol / L)·mm²) is optimal. 2 ) is the best. Figure 6The signal magnitudes of 10 subsequent measurements at 12 mmol / L glucose are shown. The signal for each measurement is listed in Table 2.

[0105]

[0106] Table 2: Examples of signal magnitude and accuracy measurements for all four sensors, all from batch #CNB003.

[0107] Response time: Biosensor response time is a critical parameter as it is a key parameter in the overall measurement cycle time, and therefore crucial for sample throughput. Response time primarily depends on the transport rate into the enzyme layer of the biosensor considered for glucose. It is defined as the time span between the appearance of the sample at the sensor and the sensor's maximum signal reading. Note the small signal effect caused by air bubbles between the system solution and the sample. A clear fluid isolation interface between the system solution and the sample solution is essential to prevent sample residue. Figure 6 As shown, the sensor's response time can be determined by analyzing the rising edge of the response curve, resulting in a response time of <4s. This is a typical value.

[0108] Please refer to Figure 6 The sensor response of a glucose sample within the normal range (5.0 mmol / L) is shown in the figure. The initial current change rate is stable. After the sample is added (corresponding to the "sample" label), there is a slight fluctuation due to factors such as the "bubble" label. Then it gradually rises, reaches the maximum signal ("maximum signal" label), and then slowly decreases.

[0109] The rinsing time and response time together determine the total duration of the detection cycle, thus limiting the system's maximum sample throughput. Therefore, shortening the rinsing time is crucial for improving detection efficiency. This study optimized the rinsing process without degrading sensor performance, achieving a shorter rinsing time and enabling the system throughput to typically exceed 100 samples per hour, with a maximum of 120 samples per hour.

[0110] Please refer to Figure 7 : Rinse time for sensor recovery after measuring 5 mmol / L glucose; Batch #CNB002. This figure shows that after the start of rinsing, the current change first rises rapidly to form a peak, then falls, and stabilizes after about 24 seconds, reaching the end of rinsing.

[0111] Please refer to Figure 8 The graph shows the sensor recovery flushing time after measuring 40 mmol / L glucose; batch #CNB002. It illustrates that after the start of flushing, the current initially rises rapidly to a peak, then drops sharply, gradually stabilizing after approximately 29 seconds, reaching the end of the flushing process. Figure 7 and Figure 8 The same sensor from batch #CNB002 was used to verify the full range adaptability of a single device.

[0112] Linearity: According to the function c = f(dI / dt), the sensor signal dI / dt is linearly correlated with the glucose concentration. A single-point calibration is sufficient to obtain the linear correlation function. In the study described here, 12 mmol / L glucose was used as the calibration solution.

[0113] Figure 9 Strong linearity was observed in four sensors across three batches, typically up to 50 mmol / L. Normally, a concentration range up to 30 mmol / L is sufficient, as high glucose concentrations in human blood are very rare and are usually due to erroneous patient sample extraction. In any case, linearity was observed up to 50 mmol / L in glucose.

[0114] Please refer to Figure 9 Linearity of 4 sensors (Sensor 1 is from batch #CNB001; Sensor 2 is from batch #CNB002; Sensors 3 and 4 are from batch #CNB003);

[0115] In-use testing: In-use testing again includes 3 sensor batches #CNB001, #CNB002 and #CNB003.

[0116] Selectivity / Interference Substances: The selectivity of glucose biosensors can be affected by electrochemically active substances that can be oxidized on the working electrode, or by substances that inhibit enzyme activity.

[0117] The selective assay was performed according to the Clinical and Laboratory Standards Institute (CLSI) guideline document EP07 (3rd edition). The assay was conducted in human serum samples, investigating the presence of eight important interfering substances in addition to glucose levels within the normal range. See Table 3 for details. The figures shown are averages from three sensors. The effect on uric acid was the highest, at 5.4 / 5.5%.

[0118] The measured interference effect was significantly lower than 10%.

[0119]

[0120] Table 3: Interference tests for batches #CNB002 and #CNB003;

[0121] Control sample: The control sample is compared with the YSI-2300 (Yellow Springs Instruments, USA).

[0122] All three sensor batches showed good correlation.

[0123] Please refer to Figure 10 The comparison of the control sample measurements of batches #CNB001, #CNB002, and #CNB003 with YSI-2300 shows the mean measurement values ​​of batches #CNB001, #CNB002, and #CNB003 under different glucose quality control samples, demonstrating the measurement results of each batch under different samples.

[0124] Patient samples: Patient samples were compared with YSI-2300 (Yellow Springs Instruments, USA). Figure 11 , Figure 12 and Figure 13 The graph shows the correlation between #CNB001 (113 blood samples), #CNB002, and #CNB003 (80 blood samples).

[0125] Please refer to Figure 11 (Based on the clinical sample set "LT12040"): Patient sample measurements VSYSI-2300 from batch #CNB001 are used to illustrate the linear relationship between batch #CNB001 measurements and YSI-2300 measurements, along with the coefficient of determination (R²). 2 (The closer to 1, the stronger the linear correlation), indicating the degree of correlation between the measurement results of the two.

[0126] Please refer to Figure 12 (Based on clinical sample set "LT14195"): Patient sample measurements VSYSI-2300 from batch #CNB002, linear relationship equation and coefficient of determination (R²) between batch #CNB002 measurements and YSI-2300 measurements. 2 The closer to 1, the stronger the linear correlation, indicating the degree of correlation between the measurement results of the two.

[0127] Please refer to Figure 13 (Based on clinical sample set "LT18331"): Patient sample measurements VSYSI-2300 from batch #CNB003, linear relationship equation and coefficient of determination (R²) between batch #CNB003 measurements and YSI-2300 measurements. 2 The closer to 1, the stronger the linear correlation, indicating the degree of correlation between the measurement results of the two.

[0128] In summary, the test data from batches #CNB001, #CNB002, and #CNB003 fully demonstrate that this sensor possesses excellent anti-interference characteristics and measurement specificity, and its performance meets the stringent requirements of clinical testing.

[0129] Although embodiments of the invention have been shown and described, it will be understood by those skilled in the art that various changes, modifications, substitutions and alterations can be made to these embodiments without departing from the principles and spirit of the invention, the scope of which is defined by the appended claims and their equivalents.

Claims

1. A glucose biosensor based on a dual-electrode structure, characterized in that, include: A ceramic substrate with dimensions of 25.4±0.1mm × 7.00±0.1mm × 0.63±0.05mm; A dual-electrode system screen-printed on the surface of the ceramic substrate: The dual-electrode system includes a platinum working electrode with an area of ​​1.02 mm² ± 5%. Silver / silver chloride reference electrode; The platinum working electrode and the silver / silver chloride reference electrode are respectively connected to a silver-palladium alloy conductive path. The platinum working electrode and the silver / silver chloride reference electrode are arranged adjacent to each other, with an edge spacing tolerance of ±0.2 mm; A blue epoxy resin isolation layer covering the non-electrode area, wherein the blue epoxy resin isolation layer material is 118-08 type epoxy resin; The glucose oxidase layer fixed on the surface of the working electrode is formed by curing waterborne acrylic-modified polyester RESYDROL® AY 5537w / 35WA crosslinked amino resin; The dual-electrode system is configured to detect hydrogen peroxide at a current density ≤10nA / mm² with a response time ≤4 seconds.

2. The glucose biosensor based on a dual-electrode structure according to claim 1, characterized in that: The isolation layer meets the requirement that there is no swelling or exudate after 24 hours of contact with a hemolyzed blood sample at 37°C.

3. The glucose biosensor based on a dual-electrode structure according to claim 1, characterized in that: The curing conditions for the glucose oxidase layer are: a clean room environment at 25±2℃, a relative humidity of 45±5%, and a time of ≤2 hours.

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