A glucose sensor having a three-dimensional microcavity and a method of manufacturing the same

Abstract

The application relates to the technical field of biosensors, and discloses a glucose sensor with a three-dimensional microcavity and a manufacturing method thereof, wherein the sensor comprises a substrate provided with at least one microcavity penetrating through the substrate; a conductive layer covering at least one side surface of the substrate and an inner wall of the microcavity, wherein the conductive layer located on the inner wall of the microcavity forms a three-dimensional microcavity sensing area; and a sensing film arranged in the three-dimensional microcavity sensing area and used for reacting with glucose to generate an electrochemical signal; the three-dimensional microcavity sensing area simultaneously provides an electrical connection path of the conductive layers on both sides of the substrate and a loading space of the sensing film. The sensor provided by the application has a high specific surface area feature, can improve the accuracy and long-term stability of the sensor, and is suitable for continuous monitoring of glucose and other physiological markers. Moreover, the manufacturing steps are highly standardized, batch production is easy to realize, and the consistency and stability of product performance are good.

Description

Technical Field

[0001] This invention relates to the field of biosensor technology, specifically to a glucose sensor with a three-dimensional microcavity and its manufacturing method. Background Technology

[0002] Continuous glucose monitoring (CGM) systems provide long-term, comprehensive, and dynamic monitoring of glucose levels, demonstrating clear clinical value in glycemic management for patients with type 1 and type 2 diabetes. Commercially available CGM systems typically employ percutaneous or transdermal glucose sensors. These sensors operate continuously for several days to approximately one week after implantation, monitoring glucose concentration changes in real time and providing data feedback to the patient. This allows for timely intervention, effectively controlling blood glucose fluctuations and preventing significant shifts in glucose levels. Currently, most glucose sensors are mass-produced and are single-use devices. They should be removed / replaced immediately after reaching a preset usage period (e.g., 7 days, 14 days, or other specified durations) to ensure safety and data reliability.

[0003] Currently, mainstream continuous glucose monitoring (CGM) systems on the market often use enzymes as core sensors to provide high specificity for analyte measurements. To minimize implantation trauma and improve patient compliance, the physical size of these sensors is designed to be extremely small, with the width of their functional area typically only a few hundred micrometers. Under such limited space, maximizing the effective sensing area and increasing the amount of bioactive enzyme immobilized becomes crucial for improving the overall performance of the sensor.

[0004] However, most existing sensors employ planar stacked electrode structures. This "smooth plate" design reveals significant limitations at the microscale. Its effective specific surface area is very limited, resulting in insufficient sites for enzyme molecule attachment, thus restricting the overall loading capacity of the sensing membrane. Because biosensitive molecules such as glucose oxidase are difficult to load in large quantities and stably onto the electrode surface, the supply of "raw materials" for the sensing reaction is limited, potentially leading to a weak initial intensity of the monitoring signal and affecting the accurate capture of glucose signals in low blood glucose concentration ranges. More critically, even if enzyme molecules are immobilized on planar electrodes through surface modification, the long-term stability of these active substances in the body fluid environment faces severe challenges. Molecular interference in the body fluid microenvironment, mechanical stress from tissue movement, and the natural inactivation of the enzymes themselves can all easily cause active substances to detach or lose their catalytic function during use. This directly leads to a shortened effective working life of the sensor, making it difficult to meet the urgent clinical need for longer replacement cycles for implantable / invasive devices.

[0005] Furthermore, traditional electrochemical glucose sensors typically employ a two- or three-electrode configuration, including at least one working electrode and one reference electrode, or an additional counter electrode. When the sensor operates, an oxidoreductase immobilized on the working electrode specifically catalyzes the redox reaction of glucose. The resulting electron transfer directly alters the current signal on the working electrode surface, and the signal intensity is proportional to the analyte concentration, thus enabling quantitative detection. This relatively fixed electrode configuration means that the sensor can usually only detect and analyze glucose as an analyte on a single physical entity. However, in clinical practice for diabetes management, in addition to blood glucose levels, other physiological parameters such as lactate and uric acid are also crucial for assessing a patient's metabolic status and predicting complication risks. To obtain this information simultaneously, patients often need to wear two or more different sensors, which not only increases physical burden and economic costs but also reduces the convenience and compliance of comprehensive monitoring. Simultaneously, in the complex environment of bodily fluids, numerous electrochemically active interfering substances (such as ascorbic acid, uric acid, and acetaminophen) can generate non-specific background currents on the working electrode, severely interfering with the accurate measurement of glucose signals. Although interference can be suppressed to some extent by selecting a low operating potential or using a selective membrane, in practical applications, the subtraction of interference signals still relies on complex algorithms or additional calibration steps, making it difficult to achieve real-time and accurate interference cancellation.

[0006] Therefore, how to achieve both multi-parameter detection capability and high anti-interference performance on a single micro-sensor platform through structural innovation has become a technical challenge that urgently needs to be solved in this field. Summary of the Invention

[0007] The present invention aims to provide a glucose sensor with a three-dimensional microcavity and its manufacturing method, in order to resolve the inherent contradiction between miniaturization and high performance of traditional planar electrodes.

[0008] To achieve the above objectives, the present invention provides the following basic solution.

[0009] Option 1 A glucose sensor with a three-dimensional microcavity, comprising: A substrate having at least one microcavity penetrating the substrate; A conductive layer covers at least one side surface of the substrate and the inner wall of the microcavity, wherein the conductive layer located on the inner wall of the microcavity forms a three-dimensional microcavity sensing region. A sensing membrane is disposed within the three-dimensional microcavity sensing area for reacting with glucose to generate an electrochemical signal. The three-dimensional microcavity sensing region simultaneously provides the electrical connection path between the conductive layers on both sides of the substrate and the load space of the sensing film.

[0010] The working principle and advantages of this invention are as follows: This invention discloses a glucose sensor with a three-dimensional microcavity, characterized by a high specific surface area, which improves sensor accuracy and long-term stability, making it suitable for continuous monitoring of glucose and other physiological markers. The key features are: This solution achieves functional integration and spatial efficiency enhancement of sensing performance through structural innovation, fundamentally breaking through the inherent contradiction between miniaturization and high performance in traditional planar electrodes, and realizing a design thinking shift from "surface" to "volume". In existing technologies, whether it is a double-sided planar electrode or a surface mesh structure, it is limited to electrode layout or surface morphology modification on a two-dimensional plane. Its essence is to increase the projected area or surface roughness without changing the basic electrode configuration. However, this solution creatively proposes a three-dimensional through-hole microcavity structure, which enables the conductive layer on the inner wall of the microcavity to simultaneously undertake the dual functions of electrical connection and sensing film load; the substrate itself is transformed from a simple supporting insulating layer into a functional sensor, and the originally isolated electrode surfaces on both sides are connected through the microcavity penetrating the substrate, so that the thickness direction of the substrate also becomes part of the sensing area.

[0011] This design integrates previously separate conductive pathways and enzyme loading spaces within a single microstructure, achieving a high degree of functional integration. From a space utilization perspective, the three-dimensional inner walls of the microcavity extend the sensing area from a planar plane to a three-dimensional space, resulting in an order-of-magnitude increase in the effective sensing area per unit volume with the same or even smaller projected area. More importantly, the concave shape of the microcavity creates a unique physical anchoring effect on the sensing membrane, fundamentally different from the single fixation method relying on chemical cross-linking in planar structures. This structural anchoring significantly enhances the adhesion and long-term sensing activity of the sensing membrane under conditions of bodily fluid disturbance and mechanical motion, thereby significantly extending the effective working life of the sensor. The synergistic effect of the above-mentioned functional integration and space efficiency enhancement allows this solution to simultaneously improve sensitivity and stability while maintaining the advantages of miniaturization.

[0012] Option 2 A method for manufacturing a glucose sensor with a three-dimensional microcavity, used to prepare a glucose sensor with a three-dimensional microcavity as described in Scheme 1, includes the following steps: S1, forming at least one microcavity penetrating the substrate on a substrate; S2, a conductive layer is formed on at least one side surface of the substrate and the inner wall of the microcavity, so that the conductive layer on the inner wall of the microcavity constitutes a three-dimensional microcavity sensing region. S3, Under vacuum conditions, the sensing membrane solution containing glucose oxidase is immersed into the sensing area of ​​the three-dimensional microcavity, so that the sensing membrane is loaded on the inner wall of the microcavity. S4, crosslinking and curing the sensing membrane.

[0013] The working principle and advantages of this invention are as follows: This invention discloses a method for manufacturing a glucose sensor with a three-dimensional microcavity. The manufacturing steps are highly standardized, the process parameters are well controllable, and it is easy to achieve mass production, ensuring the consistency and stability of performance across batches. The key points are: This solution achieves deep synergy between structural design and process methods, overcoming the limitations of conventional coating techniques in three-dimensional microcavity structures. Specifically, conventional methods for loading sensing films, such as drop coating, spin coating, or atmospheric pressure impregnation, struggle to penetrate deeply into microcavities with high aspect ratios due to the combined effects of air resistance and capillary effects. Often, only a liquid film seal is formed at the opening, preventing effective coverage of the conductive layer on the inner wall of the microcavity and hindering the full utilization of the area advantage of the three-dimensional structure. This solution addresses this challenge by creatively introducing a vacuum-assisted impregnation environment. By using negative pressure to overcome air resistance within the microcavity, the sensing film solution can penetrate unimpeded into every corner, achieving complete filling and uniform distribution of the sensing film within the three-dimensional microcavity. Through process reconfiguration, the microcavity becomes not only an electrical connection channel but also a truly efficient loading space for the sensing film.

[0014] The resulting technical benefits are: maximized contact area between the sensing membrane and the conductive layer on the inner wall of the microcavity, significantly increased enzyme loading, and the concave shape of the microcavity physically anchors the sensing membrane, enhancing its adhesion. Experimental data (see Comparative Example 3) show that the sensor prepared using the vacuum impregnation process has a near 100% filling rate of the sensing membrane within the microcavity, while a significant hollow region exists under atmospheric pressure. Correspondingly, in long-term stability tests, the performance degradation rate of the sensor prepared using the vacuum impregnation process is significantly reduced. Attached Figure Description

[0015] Figure 1 This is a cross-sectional view of a glucose sensor comprising multiple sensing regions, according to Embodiment 1 of the present invention, which discloses a glucose sensor with a three-dimensional microcavity and its manufacturing method. Figure 2a This is a top view of a glucose sensor according to Embodiment 1 of the present invention, which is a glucose sensor with a three-dimensional microcavity and a method for manufacturing the same. Figure 2b This is an overall bottom view of a glucose sensor according to Embodiment 1 of the present invention, which is a glucose sensor with a three-dimensional microcavity and a method for manufacturing the same. Figure 3a This is a top view of another glucose sensor according to Embodiment 1 of the present invention, which describes a glucose sensor with a three-dimensional microcavity and its manufacturing method. Figure 3b This is an overall bottom view of another glucose sensor according to Embodiment 1 of the present invention, which describes a glucose sensor with a three-dimensional microcavity and its manufacturing method. Figure 4 This is a top view of the sensor functional area in Embodiment 1 of the present invention, which describes a glucose sensor with a three-dimensional microcavity and its manufacturing method. Figure 5 This is a bottom plan view of the sensor functional area in Embodiment 1 of the present invention, which describes a glucose sensor with a three-dimensional microcavity and its manufacturing method. Figure 6 For along Figure 5 The side cross-sectional view of the sensor obtained by cutting line A is shown in the figure. Figure 7 For along Figure 5 The side cross-sectional view of the sensor obtained by cutting line B is shown in the figure. Figure 8a , Figure 8b , Figure 8c The images show the top view of the first working electrode sensing area when the aperture shape of the microcavity structure is square, circular, and rectangular, respectively. Figure 9 This is a cross-sectional view of the functional region of another glucose sensor according to Embodiment 1 of the present invention, which describes a glucose sensor with a three-dimensional microcavity and its manufacturing method. Figure 10 This is a bottom plan view of another functional area of ​​a glucose sensor according to Embodiment 1 of the present invention, which describes a glucose sensor with a three-dimensional microcavity and its manufacturing method. Figure 11 This is a flowchart of a second embodiment of the glucose sensor with a three-dimensional microcavity and its manufacturing method according to the present invention. Figure 12 This is a current-time response characteristic diagram of a glucose sensor with a three-dimensional microcavity and its manufacturing method according to Embodiment 3 of the present invention.

[0016] The markings in the accompanying drawings include: substrate 100, first conductive layer 210, second conductive layer 220, first working electrode 211, reference electrode 212, counter electrode 221, second working electrode 222, independent conductive region 223, protective layer 300, first surface of the first working electrode sensing region 401, second surface of the first working electrode sensing region 402, microcavity 403, sensing film 500, reference electrode sensing region 501, counter electrode sensing region 601, second working electrode sensing region 701, outer film 800, first working electrode connection contact point 910, reference electrode connection contact point 920, second working electrode connection contact point 930, conductive via 931, counter electrode connection contact point 940, conductive via 941, first conductive layer side functional region S1, and second conductive layer side functional region S2. Detailed Implementation

[0017] The following detailed explanation illustrates the specific implementation methods: Example 1 The basic implementation examples are as follows: Figure 1 As shown: A glucose sensor with a three-dimensional microcavity 403 includes: a substrate 100, a conductive layer covering the surface of the substrate 100 and the inner wall of the microcavity 403, a protective layer 300, a sensing membrane 500, and an outer membrane 800. The sensor is generally elongated and strip-shaped, with its end being a functional area for implantation in subcutaneous interstitial fluid for glucose concentration detection; its proximal end is a contact point area for electrical connection with external signal processing circuitry.

[0018] The substrate 100 and the protective layer 300 are made of polyimide, polytetrafluoroethylene, polyethylene, polycarbonate, polyethylene terephthalate or polydimethylsiloxane; the conductive material includes one or more of glassy carbon, graphite, silver, silver chloride, platinum, palladium, titanium, gold, copper, iridium, polypyrrole or poly3,4-ethylenedioxythiophene.

[0019] The functional area of ​​the sensor is located within a length range of 0.5~4mm from the end of the sensor. The width of the functional area is 0.2~2mm, and the thickness is 50~300μm. The overall length of the sensor is 4~50mm. In this embodiment, the functional area is 3mm long and 1mm wide. The overall length of the sensor is 15mm.

[0020] The glucose sensor with a three-dimensional microcavity 403 has a continuous glucose detection function. In some embodiments, the sensor can be placed in interstitial fluid to detect glucose levels, and the detected glucose can be used to infer the glucose level in the patient's bloodstream. Alternatively, the sensor can be inserted into a vein, artery, or other body part containing fluid. The basic function of the glucose sensor disclosed in this invention is to continuously monitor glucose levels over a period of time, which can be several minutes, hours, days, weeks, or longer, generating a glucose level-related signal. Exemplarily, the generated signal can be an electrical signal, the magnitude of which is proportional to the glucose concentration value.

[0021] It is understood that the miniature glucose sensor disclosed in this invention should be combined with several necessary components and / or devices to form a system to complete glucose monitoring or function. For example, the glucose sensor may be combined with a connector, control unit, receiver, and processor to form a complete sensing system, wherein only the functional area of ​​the sensor is in contact with bodily fluids, while the rest is located outside the body. In this system, the sensor functional area is the core part, generating a signal; the remaining parts provide the sensor with operating conditions (such as potential) and signal acquisition and processing.

[0022] In some embodiments, the necessary elements and / or devices further include communication elements, which can communicate directly with the terminal device via a communication path / link, or indirectly with a remote terminal device via a network (e.g., a mobile phone network, the Internet, or a cloud server). The communication path / link can also be wired or wireless, one-way or two-way, and encrypted or unencrypted. Any suitable electronic communication protocol can be used for each communication path or link, such as Near Field Communication (NFC), Radio Frequency Identification (RFID), or Low Energy Protocol (LEP), WiFi, etc. Exemplarily, the terminal device can be a computer, mobile phone, or watch, etc.

[0023] Specifically, it can be divided into the following core structures: 1. Three-dimensional microcavity 403 sensing area.

[0024] 1. Microcavity 403 structure.

[0025] like Figure 6 and Figures 8a-8cAs shown, at least one microcavity 403 penetrating the substrate 100 is provided on the substrate 100. There are multiple microcavities 403, arranged in a regular or random array; the aperture of each microcavity 403 is 50-800 μm, and the spacing between adjacent microcavities 403 is 50-800 μm; the aspect ratio of each microcavity 403 is 0.5-4. In this embodiment, the microcavities 403 are arranged in a 5×3 rectangular array. The design of multiple arrays significantly increases the sensing area, and the regularity of the array distribution helps ensure the consistency of each sensor batch. Furthermore, the specific geometric parameters of the microcavities 403 are as follows: Aperture (2r): Set to 300 μm. This aperture selection is based on a comprehensive consideration of mass transfer efficiency and structural strength. According to hydrodynamic calculations, when the aperture is less than 50 μm, the diffusion time constant of glucose molecules in the microcavity 403 will exceed 500 ms, resulting in a significant increase in sensor response time; while when the aperture is greater than 800 μm, excessive porosity will cause a sharp drop in the mechanical strength of the substrate 100.

[0026] Microcavity 403 spacing (w): set to 400 μm. The distance between the centers of adjacent microcavities 403 determines the density of the microcavity 403 array. This spacing is selected based on the results of finite element analysis (FEA) of materials mechanics: when the spacing is less than 50 μm, the solution of the sensing film 500 is difficult to fully penetrate into the interior of the microcavities 403, and the thin wall of the substrate 100 between adjacent microcavities 403 is prone to fracture under bending stress; while when the spacing is greater than 800 μm, the number of microcavities 403 per unit area is too small, the three-dimensional gain effect is not significant, and excessively large holes may weaken the mechanical strength of the substrate 100.

[0027] The depth (h) of microcavity 403 is equal to the thickness of substrate 100, which is 100 μm. The calculated aspect ratio (h / 2r) is 0.33. This aspect ratio is near the lower limit of the preferred range of 0.5 to 4, which is beneficial for sufficient impregnation of the sensing membrane 500 solution and rapid mass transfer of glucose molecules.

[0028] In particular, the structural parameters (pore size, spacing, porosity) of the microcavity 403 in this embodiment are not arbitrarily selected, but are optimized based on the following theoretical model to achieve the best balance between sensitivity and mechanical strength.

[0029] Equation 1: Sensitivity model.

[0030] .

[0031] This formula quantitatively describes the relationship between sensor sensitivity (S) and microcavity structure parameters: S: Sensor sensitivity (unit: nA / mM).

[0032] A: The projected area of ​​the sensing region of the first working electrode (in plane) ).

[0033] N, r, h: number of microcavities, radius (μm), and depth (μm), respectively.

[0034] Porosity .

[0035] First item: The contribution of the planar region decreases as porosity (φ) increases.

[0036] Second item: The contribution of all microcavity inner walls (three-dimensional microcavities) is proportional to the number of microcavities (N), radius (r), and depth (h).

[0037] k: Sensitivity constant (calibrated experimentally, approximately 0.95 in this embodiment). The connectivity coefficient (approximately 0.2 in this embodiment) and These are the effective coefficients of the plane and the effective coefficients of the hole wall (obtained through finite element simulation, approximately 0.8 and 0.9 respectively in this embodiment).

[0038] For a given planar area A, there exists an optimal porosity ( This maximizes the combined contribution of the planar region and the three-dimensional microcavity region. When the pore size is 300 μm and the spacing is 400 μm, the calculated porosity is... The value is approximately 0.44, at which point the sensitivity S reaches its peak.

[0039] Equation 2: Stiffness model.

[0040] .

[0041] This formula describes the effective bending strength of the sensor ( Relationship with microcavity structure: in, Effective bending strength (MPa).

[0042] f(r / w): Aperture-to-spacing ratio function, obtained by fitting from finite element analysis. .

[0043] Strength varies with porosity ( The value decreases linearly as the aperture-to-spacing ratio (r / w) increases, while it decreases nonlinearly as the aperture-to-spacing ratio (r / w) increases.

[0044] Based on the above model, the present invention limits the pore size of the microcavity 403 to 50-800 μm and the spacing to 50-800 μm. Preferably, the pore size is 300 μm and the spacing is 400 μm.

[0045] Furthermore, it is understood that the shape of the holes in the microcavity 403 structure is flexible and varied, such as rectangular, circular, square, elliptical or other polygonal shapes, preferably circular.

[0046] 2. A conductive layer covering at least one surface of the substrate 100 and the inner wall of the microcavity 403. The conductive layer located on the inner wall of the microcavity 403 forms a three-dimensional sensing region of the microcavity 403.

[0047] Specifically, the conductive layer is formed by a printing process (e.g., the formation method can be any one or more of electroplating, chemical deposition, mechanical rolling, screen printing, sputtering, or photolithography). In this embodiment, a conductive layer is deposited on both sides of the substrate 100 and the inner wall of the microcavity 403 by magnetron sputtering to obtain a first conductive layer 210 located on the first side of the substrate 100 and a second conductive layer 220 located on the second side of the substrate 100. A separation groove is formed between the first conductive layer 210 and the second conductive layer 220, and the width of the separation groove is in the range of 50~800μm.

[0048] The double-sided structure design with conductive layers on both sides of the substrate 100 allows for more independent electrodes to be accommodated within the same sensor planar projection area, thereby achieving miniaturization of the sensor and reducing implantation trauma. At the same time, the double-sided structure effectively utilizes the space in the thickness direction of the substrate 100, providing a structural basis for multi-electrode configuration, enabling the sensor to achieve multi-parameter detection or differential measurement functions while maintaining a small size.

[0049] like Figure 2a and Figure 2b The image shows the overall top and bottom views of the sensor. In simple terms, the sensor can be functionally divided into a connection contact area, electrodes, and functional areas. The connection contacts include a first working electrode connection contact 910, a reference electrode connection contact 920, a second working electrode connection contact 930, and a counter electrode connection contact 940. Each connection contact is an exposed conductive material area located in the first conductive layer 210, used for current transmission with other components and / or devices, and electrically connected to the second conductive layer 220 through conductive vias 931 and 941.

[0050] Figure 3a and Figure 3bThe image shows a top and bottom view of another glucose sensor. Its locally improved connection contact design allows for more flexible two-dimensional sensor design, a smaller area, and different sensor shapes. It should be understood that other connection contact and electrode layout designs are also included within the scope of this invention, including, but not limited to, different designs of stacked, parallel, and twisted connection contacts and electrodes.

[0051] Figure 4 The diagram shown is a top view of the sensor's functional area. The first working electrode 211 has an enlarged sensing area 401 at its end, and the reference electrode 212 has an enlarged sensing area 501 at its end. Figure 5 The diagram shows a bottom view of the sensor's functional area. The counter electrode 221 has an enlarged counter electrode sensing area 601 at its end, and the second working electrode 222 has an enlarged second working electrode sensing area 701 at its end.

[0052] The first conductive layer 210 is divided along its length into a first working electrode 211 and a reference electrode 212, which are independent of each other.

[0053] Independent electrodes are formed by dividing the conductive layer along its length, so that each electrode is arranged longitudinally, which meets the slender shape requirements of implantable sensors and facilitates percutaneous puncture and implantation.

[0054] In this embodiment, the terminal region of the first working electrode 211 corresponds to the position of the microcavity 403 array. A layer of platinum black (Pt-black) inert conductive material is deposited on the surface of this region. The high catalytic activity of platinum black is used for the efficient oxidation of hydrogen peroxide produced by the glucose oxidation reaction. An Ag / AgCl reference material layer is formed on the surface of the terminal sensing region of the reference electrode 212. This reference electrode 212 can provide a constant reference potential (approximately +0.222 V relative to a standard hydrogen electrode) in a body fluid environment with a relatively stable chloride ion concentration, ensuring the stability of the excitation voltage of the working electrode. The stability of the working voltage directly determines the efficiency of the glucose oxidation reaction and the repeatability of the current signal, thereby ensuring the accuracy and consistency of the detection results during long-term monitoring.

[0055] The second conductive layer 220 is divided along its length into a second working electrode 222 and a counter electrode 221, which are independent of each other. The microcavity 403 connects the region at the end of the first working electrode 211 with the independent conductive region 223 on the second conductive layer 220 corresponding to that region, such as... Figure 5As shown. It is understood that in some embodiments, the functions of each electrode can be flexibly allocated. For example, sensing region 601 can be configured as second working electrode sensing region 701, and sensing region 701 can be configured as counter electrode sensing region 601.

[0056] Specifically, the independent conductive region 223 is located at the end of the second conductive layer 220. This region is electrically connected to the first working electrode 211 through the conductive layer of the inner wall of the microcavity 403, while being electrically isolated from other electrodes on the second conductive layer 220 (such as the second working electrode 222 and the counter electrode 221).

[0057] The end sensing area of ​​the second working electrode 222 can be configured differently according to functional requirements: Function a, differential detection mode: forms a differential electrode pair with the first working electrode 211 to cancel the background current signal generated by coexisting interfering substances.

[0058] Specifically, when the second working electrode 222 is configured to perform differential signal detection in conjunction with the first working electrode 211, its sensing area is provided with an enzyme-free sensing membrane 500 (i.e., only covered with inert conductive material platinum black, without enzymes). In this mode, the two working electrodes are applied with the same working potential, and electrochemically active interfering substances in the body fluid (such as one or more of ascorbic acid, uric acid, fructose, lactic acid, acetylamino acid, or bilirubin) will generate similar interference currents on the surfaces of the two electrodes. By subtracting the interference signal of the second working electrode 222 from the detection signal of the first working electrode 211, the background current can be effectively canceled, significantly improving the signal-to-noise ratio.

[0059] Function b, multi-parameter detection mode: independently detects a second target substance other than glucose, which is selected from lactic acid, uric acid, alcohol or ascorbic acid.

[0060] When the second working electrode 222 is configured to independently detect a second target substance (such as lactic acid) other than glucose, its sensing area is provided with a second substance sensing membrane 500. This sensing membrane 500 contains lactate oxidase (LOx), BSA, potassium ferricyanide, and glutaraldehyde for the specific detection of lactic acid. LOx catalyzes the oxidation of lactic acid to produce hydrogen peroxide, the concentration of which is proportional to the lactic acid concentration.

[0061] An inert conductive material (platinum black, gold, silver, and carbon) is disposed on the surface of the end sensing area of ​​the counter electrode 221, forming a complete current loop with the first working electrode 211 and the second working electrode 222 to handle the current flux generated by the electrochemical reaction. Its chemical stability ensures that the counter electrode 221 will not corrode or dissolve during long-term use, and will not release harmful substances into bodily fluids. Simultaneously, the high conductivity and good electrochemical activity of the inert material enable the counter electrode 221 to efficiently complete charge transfer, preventing polarization of the counter electrode 221 from affecting the normal operation of the working electrodes, thus ensuring the stable conduct of the electrochemical reaction and the reliability of signal acquisition.

[0062] It should be noted that, Figure 4 and Figure 5 The schematic diagrams of the sensing regions and their corresponding conductive layers shown are not intended to represent the only embodiments disclosed in this invention. Other possible designs for the shape, size, and position of the sensing regions and their corresponding conductive layers are also within the scope of this invention. However, it should be noted that the area of ​​the sensing region located on the surface of the conductive layer should be smaller than that of the conductive layer; that is, the sensing region should be completely disposed on the conductive layer.

[0063] Figure 6 The image shows the sensor along... Figure 5 The image shows a cross-sectional view obtained by taking a section from line A. A microcavity 403 structure is provided between the first surface 401 and the second surface 402 of the aforementioned first working electrode sensing area to connect them.

[0064] Figure 7 The image shows the sensor along... Figure 5 The image shows a cross-sectional view obtained by taking a section from line B. It can be seen that the first conductive layer 210 is divided into a first working electrode 211 and a reference electrode 212, wherein a reference electrode sensing region 501 is formed on a large area of ​​the surface of the reference electrode 212. Similarly, a counter electrode sensing region 701 is formed on a large area of ​​the surface of the counter electrode 221 located in the second conductive layer 220. Although a cross-sectional view of the second working electrode sensing region 701 is not shown, it should have a similar structure and characteristics. Figure 8a , Figure 8b , Figure 8c A top view of the sensing area of ​​the first working electrode 211 is shown. As can be seen from the figure, the aforementioned microcavity 403 structure is not unique in number and has an array feature.

[0065] Furthermore, the sensing region of the three-dimensional microcavity 403 simultaneously provides an electrical connection path between the conductive layers on both sides of the substrate 100 and a load space for the sensing film 500. Specifically, this region has the following characteristics: Electrical connectivity: The conductive layer is continuously covered by the inner wall of the microcavity 403, realizing electrical connectivity between the conductive layers on the first side (top surface) and the second side (bottom surface) of the substrate 100. This allows different electrode regions located on both sides of the substrate 100 (such as the end of the first working electrode 211 and the independent conductive region 223 on the second conductive layer 220) to form an electrical connection through the microcavity 403.

[0066] The sensing membrane 500's loading function: The hollow region inside the microcavity 403 and the conductive layer on its inner wall together constitute the loading space of the sensing membrane 500. This space not only provides a surface area far exceeding that of a planar structure, but its concave shape also physically anchors the sensing membrane 500. Compared to a planar structure, this significantly increases the effective attachment area of ​​the sensing membrane 500, allowing more glucose oxidase to be stably fixed to the electrode surface, thereby improving the sensor's sensitivity and signal output strength.

[0067] Interface characteristics: By controlling the sputtering process parameters, the conductive layer deposited on the inner wall of the microcavity 403 has a nanoscale rough surface with an average surface roughness (Ra) of 10~100 nm. At the same time, the conductive layer is dense and non-porous, with a porosity of less than 5%, which can effectively prevent electrolyte penetration and corrosion.

[0068] In some implementations, such as Figure 9 and Figure 10 As shown, the sensor can be configured with three electrodes, including a working electrode 211, a reference electrode 212, and a counter electrode 221. The first conductive layer 210 is divided into the first working electrode 211 and the reference electrode 212, and the second conductive layer 220 is divided into the counter electrode 221 and the second surface 402 of the sensing area of ​​the first working electrode. It should be understood that when the sensor is configured with three electrodes, the corresponding sensing area should also be three, and the corresponding contact points should also be only three.

[0069] 3. A sensing film 500 is disposed within the sensing area of ​​the three-dimensional microcavity 403. Specifically, it is a conductive layer loaded on the surface of the inner wall of the microcavity 403 and fills the internal space of the microcavity 403 (e.g., Figure 6 (As shown). The sensing membrane 500 is used to specifically react with glucose to generate a signal that can be detected electrochemically.

[0070] The sensing membrane 500 comprises glucose oxidase, a protective agent, an electron mediator, and a crosslinking agent; the protective agent is selected from bovine serum albumin, trehalose, or polyethylene glycol; the electron mediator is selected from ferricyanide, methylene blue, ferrocene, or transition metal complexes.

[0071] The protective agent is used to reduce the possibility of excessive cross-linking and structural damage of glucose oxidase during the cross-linking process, thereby improving the long-term stability of glucose oxidase. The electron mediator is used to promote the transfer of electrons related to the glucose oxidation reaction to the surface of the first working electrode 211. Specifically, it includes directly acting as an electron mediator in the glucose oxidation reaction, replacing oxygen as an oxidant; in addition, it includes catalyzing the hydrogen peroxide produced by glucose oxidase catalyzing glucose, indirectly completing the transfer of electrons related to the glucose oxidation reaction. It is understood that both of the aforementioned forms use an electron mediator to help transfer electrons related to glucose oxidation to the electrode, and the two forms should be able to coexist.

[0072] The crosslinking agent is selected from glyoxal, glutaraldehyde, adipaldehyde, 1,3-dioxane, etc. Butylene dialdehyde, terephthalaldehyde, 2,3-dihydroxyterephthalaldehyde or 2,4-dihydroxyterephthalaldehyde Hexadienal. All of the above dialdehyde compounds contain two aldehyde groups, which can simultaneously undergo Schiff base reactions with two amino groups to form covalent bonds, thereby connecting amino-containing enzyme molecules, protective agent molecules, and electrode surfaces into a three-dimensional cross-linked network. The density of the cross-linked network formed by aldehyde cross-linking agents of different carbon chain lengths varies: short-chain cross-linking agents (such as glyoxal) form a tight network, which is beneficial to the mechanical stability of the sensing membrane 500; long-chain cross-linking agents (such as glyoxal) form a looser network, which is beneficial to maintaining the conformational flexibility and catalytic activity of enzyme molecules. By selecting or mixing different cross-linking agents, the balance between the stability and activity of the sensing membrane 500 can be optimized.

[0073] In this embodiment, the sensing membrane 500 is composed of glucose oxidase (GOx), a protective agent (bovine serum albumin, BSA), an electron mediator (potassium ferricyanide), and a crosslinking agent (glutaraldehyde). The specific ratio is as follows: GOx concentration is 10 mg / mL, BSA concentration is 20 mg / mL (GOx:BSA mass ratio 1:2), potassium ferricyanide concentration is 0.5 mg / mL, and the final glutaraldehyde concentration is 0.5% (v / v).

[0074] The aforementioned components, through the cross-linking action of glutaraldehyde, undergo a Schiff base reaction with enzyme molecules, protective agents, and amino groups on the electrode surface to form a stable three-dimensional cross-linked network, thereby firmly anchoring to the conductive layer on the inner wall of the microcavity 403. This fixation method ensures the mechanical stability of the sensing membrane 500 during long-term use.

[0075] The sensing membrane 500 can react with glucose, thereby causing the sensor to generate a signal related to glucose concentration. For example, the signal generated by the first working electrode 211 of the sensor that is positively correlated with the glucose concentration level can be a current signal or a voltage signal. In some embodiments, a stable voltage excitation must be provided to the first working electrode 211 to generate the signal; for example, a common voltage excitation range is -1 to 1V, specifically, the voltage can be 0.6V, 0.1V, or -0.3V, etc.

[0076] 4. A protective layer 300 is disposed on the surface of the non-sensing area of ​​the first conductive layer 210 and the second conductive layer 220 (covering all surfaces of the first conductive layer 210 and the second conductive layer 220 except for the connecting contact point and the end sensing area), and in the partition groove between each electrode (filling the partition groove between each electrode), for insulating and isolating each electrode, preventing short circuits and signal crosstalk, and protecting the non-sensing area from body fluid corrosion and mechanical damage.

[0077] In this embodiment, the protective layer 300 is formed by coating, exposing, developing and curing polyimide photoresist, and has a thickness of about 30 μm.

[0078] like Figure 6 and Figure 7 As shown, a protective layer 300 should also be provided in the area where no conductive layer is formed at the edge of the sensor to provide mechanical and solution protection for the conductive layer of the sensor and to maintain electrical insulation.

[0079] 5. An outer membrane 800, covering the outer side of the sensing membrane 500. The outer membrane 800 comprises at least one polymer selected from polyamide, polysulfone, polydimethylsiloxane, nylon, polyvinyl chloride, polyvinylimidazolium, polyacrylate, polyurethane, acrylate derivatives, polyetherurethane, silicone, polyvinylidene fluoride, or chitosan.

[0080] In some embodiments, the outer membrane 800 further includes enzymes or catalysts. For example, it may contain catalase or a catalyst to catalyze the escape of hydrogen peroxide and replenish the oxygen supply; or it may contain an enzyme that catalyzes the degradation of interfering substances (such as ascorbic acid oxidase) to reduce interference signals.

[0081] In this embodiment, the outer membrane 800 is formed of polyurethane (PU) by dip-coating and covers the outer side of all sensing areas at the sensor tip, with a thickness of 5~30 μm. This ensures that each sensing area receives the same level of material transport restriction, allowing glucose and other target substances to reach the electrode surface at a controllable and uniform rate. The complete coverage provides physical protection for the sensing membrane 500, preventing damage to the sensing membrane 500 due to tissue contact during implantation and slowing down the deposition of contaminants such as proteins in body fluids.

[0082] The outer membrane 800 has glucose and interfering substance flux limiting functions and biocompatibility characteristics. Specifically: Glucose flux limitation: By adjusting the film-forming conditions (solvent ratio, drying rate), a thin film with a controllable microcavity 403 structure is formed. Based on the dissolution-diffusion mechanism, this film limits the rate of glucose transport to the electrode surface, making glucose supply the rate-limiting step of the reaction, thereby extending the linear detection range of the sensor to the physiological concentration range of 0-30 mM.

[0083] Biocompatibility: The low surface energy of polyurethane materials reduces non-specific adsorption of proteins and aggregation of inflammatory cells, thereby reducing foreign body reactions after implantation and extending the effective working life of the sensor.

[0084] This embodiment provides a glucose sensor with a three-dimensional microcavity 403, which is no longer a single planar stacked structure, but has an optimized three-dimensional depth structure. By forming the microcavity 403 structure in the sensing area, the sensor has a larger relative sensing surface, improving sensor sensitivity. Specifically, the three-dimensional inner wall surface of the microcavity 403 structure increases the sensing area several times compared to the planar structure, allowing more glucose oxidase to be immobilized in the expanded sensing area, generating a stronger electrochemical response signal and significantly improving sensing sensitivity. In addition, the microcavity 403 structure can also load sensing material, effectively increasing the amount of sensing material immobilized and improving the long-term sensitivity of the sensor; the concave shape of the microcavity 403 physically anchors the sensing membrane 500, enhancing the adhesion of the sensing membrane 500 and extending the effective working life of the sensor; or, with the same amount of sensing material immobilized, significantly reducing the required sensor electrode size, thereby reducing user trauma. The dual-working-electrode differential design can suppress interference currents generated by interfering substances in body fluids, effectively improving the accuracy of sensing signals; or it can achieve simultaneous detection of multiple parameters, completing the monitoring of multiple physiological indicators such as glucose and lactate in a single implantation.

[0085] In summary, the glucose sensor provided in this embodiment allows for excellent long-term stability, high sensitivity, and accurate glucose detection with minimal trauma. Specifically, the miniaturized design of the sensor reduces patient implantation pain and foreign body discomfort, improving compliance; high sensitivity ensures reliable detection in the hypoglycemic range, helping to promptly detect life-threatening hypoglycemic events; long-term stability extends the effective lifespan of a single sensor, reducing replacement frequency and operating costs; and differential detection or multi-parameter detection capabilities expand the sensor's application range and clinical value.

[0086] To further demonstrate the superiority of the present invention, several comparative examples are introduced below to verify the technical effects of the present invention: Comparative Example 1 Comparison with a single-sided planar blind hole electrode sensor (verifying the effect of the "three-dimensional microcavity 403 sensing area" of this invention).

[0087] Comparison object: The control sensor has only a conductive layer and sensing film 500 on one side, and the microcavity 403 is a blind hole (closed at the bottom), and does not form a through "three-dimensional microcavity 403 sensing area".

[0088] Specifically, a polytetrafluoroethylene substrate with a width of about 1 mm and a thickness of about 0.1 mm was used to prepare the double-sided planar microcavity 403 interconnect sensor of the present invention and the single-sided planar blind hole sensor (i.e., control sensor) used as comparative example 1.

[0089] Both sensor microcavities 403 are 1×4, with a pore diameter and pore gap of 400 μm. A platinum electrode is used as the counter electrode 221, and Ag / AgCl is used as the reference electrode 212. The sensing membrane 500 and the outer membrane 800 are fixed in the same way.

[0090] Test results: The voltammetric test results in a 5 mM glucose solution show that the peak current of the sensor of this invention is approximately 2030~2550 nA, while the peak current of the control sensor is approximately 1090~1260 nA. This result indicates that the sensor structure proposed in this invention effectively increases the sensing area, and the response current for the same glucose concentration is 1.6~2.3 times that of a single-sided planar blind-hole sensor.

[0091] Data source: This data comes from the chronoamperometry test results of the electrochemical workstation. Each condition was measured 5 times and the average value was taken.

[0092] Conclusion: The data directly demonstrate that the three-dimensional microcavity 403 sensing region, due to the conductive layer on its inner wall simultaneously achieving electrical connection and sensing membrane 500 load, expands the effective sensing area from a single surface to the first surface + second surface + inner wall of microcavity 403, thereby significantly improving the enzyme loading and signal response per unit projected area.

[0093] Comparative Example 2 Compared with a fully filled sensor with a microcavity 403 structure (to verify the effectiveness of the microcavity 403 as the load space of the sensing membrane 500 in this invention).

[0094] Comparison object: The control sensor microcavity 403 is completely filled with epoxy resin, serving only as an electrical connection channel and not as a load space for the sensing membrane 500.

[0095] Specifically, a polytetrafluoroethylene substrate with a width of about 1 mm and a thickness of about 0.1 mm was used to prepare the dual-sided planar microcavity 403 interconnect sensor of the present invention and the microcavity 403 fully filled sensor used as Comparative Example 2 (i.e., control sensor).

[0096] The main difference between the two is that the microcavity 403 structure of the fully filled microcavity 403 sensor only serves as an electrical conductor and is no longer used as a sensing load. Instead, it is completely filled with epoxy resin. Both sensors use a 1×4 array of microcavities 403, with a hole diameter and hole spacing of 400 μm. A platinum electrode is used as the counter electrode 221, and Ag / AgCl is used as the reference electrode 212. The sensing membrane 500 and the outer membrane 800 are fixed in the same way.

[0097] This result shows that using the microcavity 403 structure of the sensor as part of the sensing load helps to increase the sensing area, and the response current can increase by up to 1.67 times; it also helps to improve the long-term stability of the sensing material, which in turn is beneficial to the long-term stability of the sensor performance.

[0098] Test results: In a 5 mM glucose solution, the peak current of the dual-sided planar microcavity 403 interconnect sensor of this invention was approximately 2030~2550 nA, and the peak current of the fully filled microcavity 403 structure sensor was approximately 1520~1800 nA (this invention improves upon the initial current by 1.13~1.67 times). After 7 days, the response current of the dual-sided planar microcavity 403 interconnect sensor of this invention remained at 89~92%, while the response current of the fully filled microcavity 403 structure sensor was 78~83% of the initial current (this invention improves upon the initial current by 9~14 percentage points).

[0099] Data source: Stability data was obtained by immersing the sensor in PBS at 35°C for 7 consecutive days, with tests conducted every 24 hours.

[0100] Conclusion: Using the microcavity 403 structure of the sensor as part of the sensing load helps to increase the sensing area, and the response current can be increased by up to 1.67 times. Furthermore, the microcavity 403, as a load space, not only provides a larger initial enzyme loading capacity, but its concave shape also creates a "physical anchoring" effect that significantly enhances the adhesion of the sensing membrane 500, preventing enzyme detachment during long-term use and contributing to the long-term stability of the sensing material, thereby benefiting the long-term stability of the sensor performance.

[0101] Example 2 like Figure 11 As shown, a method for manufacturing a glucose sensor with a three-dimensional microcavity 403, used to prepare a glucose sensor with a three-dimensional microcavity 403 as described in Example 1, includes the following steps: S1, at least one microcavity 403 penetrating the substrate 100 is formed on a substrate 100.

[0102] The present invention does not limit the method of forming the microcavity 403; any method known to those skilled in the art can be used, such as mechanical drilling or laser cutting.

[0103] In this embodiment, an ultraviolet nanosecond laser processing system (wavelength 355 nm) is used to form a microcavity 403 array penetrating the substrate 100 in the sensing functional region (located at the end of the substrate 100). The array configuration is optimized based on the theoretical models of Equations 1 and 2. In this embodiment, the microcavities 403 are distributed in a 5×3 rectangular array. The aperture of the microcavities 403 is limited to 50-800 μm, and the spacing is limited to 50-800 μm. Preferably, the aperture is 300 μm and the spacing is 400 μm.

[0104] S2, a conductive layer is formed on at least one side surface of the substrate 100 and the inner wall of the microcavity 403, so that the conductive layer on the inner wall of the microcavity 403 constitutes a three-dimensional microcavity 403 sensing area.

[0105] In this invention, the method of forming the conductive layer is not limited, and any deposition method well known to those skilled in the art is acceptable. For example, the formation method can be any one or more of electroplating, chemical deposition, mechanical rolling, screen printing, sputtering, or photolithography. Furthermore, this invention does not limit the method of forming the conductive layer. For example, it can be formed entirely in one step, or it can be formed in multiple steps in different areas.

[0106] In addition, the conductive layer on the inner wall of the microcavity 403 should be dense and have nanoscale roughness (10~100 nm). The dense and low porosity of the conductive layer prevents electrolyte penetration that could lead to corrosion or stress concentration, improves mechanical stability, and reduces the risk of breakage; the nanoscale roughness (Ra 10–100 nm) on the surface of the conductive layer is beneficial for improving electrochemical performance and increasing enzyme immobilization sites.

[0107] This embodiment includes the following sub-steps: S201, conductive layer formed.

[0108] In this embodiment, a magnetron sputtering process is used to deposit conductive layers on both sides of the substrate 100 with the microcavity 403 already formed and on the inner wall of the microcavity 403, to obtain a first conductive layer 210 located on the first side of the substrate 100 and a second conductive layer 220 located on the second side of the substrate 100, thereby realizing the construction of the three-dimensional microcavity 403 sensing region.

[0109] Specifically, under an argon atmosphere, a titanium (Ti) layer of approximately 20 nm thickness is first sputtered as an adhesion layer, followed by a gold (Au) layer of approximately 200 nm thickness as a base conductive layer. During the sputtering process, the tilt and rotation of the substrate 100 support are adjusted to ensure that the inner wall of the microcavity 403 is uniformly covered.

[0110] By controlling the sputtering parameters (e.g., power 100 W, gas pressure 0.5 Pa), a dense microstructure with nanoscale roughness is formed on the surface of the deposited Au layer. At this point, the conductive layer on the inner wall of the microcavity 403 has formed a three-dimensional sensing region, achieving electrical connection (electrically connecting the conductive layers on both sides of the substrate 100, laying the foundation for subsequent electrode definition), and providing a hollow, high-surface-area three-dimensional space to accommodate the sensing film 500 as a sensing load space.

[0111] S202, conductive layer segmentation.

[0112] The first conductive layer 210 and the second conductive layer 220 are cut and separated to form a partition groove, which separates each conductive layer to form an independent electrode.

[0113] Furthermore, the first conductive layer 210 (top surface) is cut along the length direction to form a first working electrode 211 (end region) and a reference electrode 212 (proximal region) that are independent of each other.

[0114] The second conductive layer 220 (bottom surface) is cut along its length to form an independent second working electrode 222 and a counter electrode 221. Specifically, the region at the end of the second conductive layer 220 corresponding to the first working electrode 211 is cut into an independent conductive island. This independent conductive island is electrically connected to the first working electrode 211 only through the conductive layer on the inner wall of the microcavity 403, and is not connected to other electrodes on the second conductive layer 220 (such as the second working electrode 222).

[0115] The width of the separator groove is 50-800 μm, preferably 300 μm, which can ensure reliable insulation between adjacent electrodes without excessively increasing the overall size of the sensor.

[0116] Laser segmentation technology is used for the conductive layer segmentation operation. Specific parameters such as laser wavelength and pulse energy are not limited in this invention; any parameters well-known to those skilled in the art are acceptable. In this embodiment, ultraviolet laser cutting is selected. Furthermore, it is understood that the purpose of segmenting the conductive layer is to form mutually independent electrodes / conductive regions. When the conductive layer formed in some embodiments already satisfies the condition of mutual independence between the sensor and the electrode 221 / conductive region, the segmentation operation can be omitted.

[0117] S203, functionalization of the electrode sensing area.

[0118] To impart specific electrochemical functions to each electrode, a functional metal layer is formed in the corresponding region. This involves the following operations: (1) Modification of working electrode: An inert conductive material is formed in the end sensing regions of the first working electrode 211 and the second working electrode 222. (2) Modification of reference electrode 212: Ag / AgCl is formed in the end sensing region of the reference electrode 212. This reference electrode 212 can provide a stable reference potential in body fluids.

[0119] In this invention, the method of forming the metal layer is not limited; any deposition method well known to those skilled in the art is acceptable. For example, the formation method can be any one or more of electroplating, chemical deposition, mechanical rolling, screen printing, sputtering, or photolithography. Sputtering, electroplating, or chemical deposition is preferred. This invention does not limit the specific steps and parameters of the above methods; process parameters well known to those skilled in the art are acceptable. Furthermore, in some embodiments, the formed metal layer can be multiple layers of metal, for example, first depositing a nickel layer, and then depositing a gold layer on the nickel layer.

[0120] In this embodiment, a layer of platinum black (Pt-black) is deposited as an inert conductive material via electrochemical deposition. The platinum black layer provides a highly active catalytic surface for efficiently catalyzing the oxidation reaction of hydrogen peroxide, thereby amplifying the detection signal. An Ag / AgCl layer is formed via electrochemical chlorination. First, a silver (Ag) thin film is sputtered, and then in a 0.1 M HCl solution, a constant current (0.1 mA / L) is applied. The surface is anodicly polarized for 60 seconds to convert it to Ag / AgCl.

[0121] S204, protective layer 300 is formed.

[0122] A protective layer 300 is formed on all conductive layer surfaces except for the connection point (PAD) and end sensing area, and within all partition grooves.

[0123] In this invention, the formation method of the protective layer 300 is not limited, and any deposition method well known to those skilled in the art is acceptable. Preferably, it is one or more of substrate lamination, coating / printing, or chemical deposition. In this embodiment, the protective layer 300 is formed by coating polyimide.

[0124] S205, sensor unitization.

[0125] The layers are cut along the designed outline to form individual sensors. The cutting method can be mechanical cutting or laser cutting. This invention does not limit the specific steps and parameters of the above methods; process parameters well known to those skilled in the art are acceptable.

[0126] In this embodiment, a laser cutting system is used to completely cut the sensor from a single sheet, forming a single, independent sensor. The overall length of the sensor is 15 mm, the length of the functional area at the end is 3 mm, and the width is 1 mm.

[0127] S3, under vacuum conditions, a solution containing glucose oxidase, the sensing membrane 500, is immersed in the sensing region of the three-dimensional microcavity 403, so that the sensing membrane 500 is loaded onto the inner wall of the microcavity 403. Specifically, this includes the following sub-steps: S301, Prepare sensor membrane 500.

[0128] A glucose sensing film 500 is formed on the surface of the sensing area of ​​the first working electrode 211, and the operation is repeated 3 to 5 times under vacuum conditions, with an interval of 1 to 3 minutes between each operation.

[0129] Specifically, a phosphate buffered saline (PBS) solution is prepared for use. The specific formulation and preparation steps of the phosphate buffered saline solution are not required, as long as they are familiar to those skilled in the art.

[0130] The pH range of the phosphate buffer solution should be 6.0–8.0, preferably 7.2–7.6. Solution A is obtained by dissolving glucose oxidase, the protective agent, and the electron mediator using the phosphate buffer solution, while solution B (the working solution of the crosslinking agent) is obtained by dissolving and diluting the crosslinking agent using the same phosphate buffer solution. This invention does not limit the specific steps for solution preparation; methods well-known to those skilled in the art are acceptable.

[0131] In solution A, the mass percentage of glucose oxidase is 1% to 20%, preferably 5% to 10%; the mass ratio of glucose oxidase to protective agent is 1:1 to 1:20, preferably 1:3 to 1:10; and the mass ratio of glucose oxidase to electron medium is 20:1 to 100:1, preferably 50:1 to 100:1. In solution B, the mass percentage of crosslinking agent is 1% to 10%, preferably 1% to 5%.

[0132] The process of forming the sensing film 500 can be any one of dip coating, drop coating, and spray coating, preferably dip coating.

[0133] Under magnetic stirring, solution B is added dropwise to solution A while stirring, and the mixture is thoroughly mixed to obtain a glucose sensing membrane 500 solution. The volume ratio of solution A to solution B is controlled at 10:1 to 1:10, preferably 3:1 to 1:3.

[0134] Under vacuum conditions, the sensing membrane 500 solution is uniformly coated onto the sensing area of ​​the first working electrode 211 of the sensor using an immersion-lift process. As described in the previous embodiment, the first working electrode 211 is located at the far end of the sensor. Therefore, during the immersion operation, the distance between the sensor and the sensing membrane 500 solution should be controlled to ensure that the sensing membrane 500 is coated only on the sensing area of ​​the first working electrode 211.

[0135] It should be noted that, as described in the previous embodiment, the sensing area of ​​the first working electrode 211 includes a microcavity 403 structure. Therefore, the immersion operation must be carried out in a vacuum environment to ensure that the sensing film 500 is fully immersed in the microcavity 403 structure.

[0136] The absolute pressure of the vacuum environment is 1~50 kPa, preferably 5~20 kPa. The sensor's terminal functional area is immersed in the solution and maintained at this vacuum level for 1 minute, allowing the solution to fully penetrate the inner wall of the microcavity 403 under negative pressure. Subsequently, the sensor is slowly pulled out of the liquid and left to stand in the air for 2 minutes to allow the solvent to initially evaporate.

[0137] Repeat the above "immersion-lifting-resting" operation 3 times to achieve the layer-by-layer deposition and complete filling of the sensing membrane 500 in the microcavity 403.

[0138] In some embodiments, the formation of the sensing membrane 500 may also be a single, non-repeatable operation, depending on the characteristics of the sensing membrane 500 solution and the required thickness of the sensing membrane 500.

[0139] S4, crosslinking and curing the sensing membrane 500.

[0140] The sensor that has been impregnated is transferred to a constant temperature and humidity chamber, with the temperature set at 30~38℃, preferably 35℃, and the humidity at 50~70%, preferably 60%.

[0141] Under these conditions, the crosslinking is allowed to proceed statically for 18–36 hours, preferably 24 hours. Under these temperature and humidity conditions, glutaraldehyde reacts with the amino groups of GOx and BSA via a Schiff base reaction, forming a stable three-dimensional crosslinked network. This crosslinking reaction firmly anchors the components of the sensing membrane 500 onto the nano-roughened conductive layer on the inner wall of the microcavity 403.

[0142] If the second working electrode 222 is configured as a differential electrode, then there is no need to prepare a sensing film 500 in this region.

[0143] If the second working electrode 222 is configured to detect lactate, the above steps need to be repeated, but the GOx in solution A is replaced with lactate oxidase (LOx, 5 mg / mL), the BSA concentration is 10 mg / mL, and the mass ratio of GOx to BSA is adjusted to 1:2. The second sensing membrane 500 solution is precisely drop-coated onto the sensing area of ​​the second working electrode 222 under vacuum using a vacuum drop-coating method.

[0144] S5, preparation of outer membrane 800: The thoroughly stirred outer membrane 800 solution is uniformly distributed in the functional area at the end of the sensor, and then stabilized in a constant temperature and humidity chamber at 30~38℃ and 50~70% humidity for 72h.

[0145] The formation of the sensing membrane 500 can be achieved through any of the following methods: dip coating, drop coating, and spray coating, with dip coating being preferred. Specifically, the aforementioned outer membrane 800 material is fully dissolved in a solvent to form an outer membrane 800 solution, with the mass fraction of the outer membrane 800 material ranging from 0.1% to 20%, preferably from 1% to 10%. Then, the sensor with the prepared sensing membrane 500 is immersed in and pulled out of the outer membrane 800 solution once or multiple times using a dip coating process. The immersion and pulling operation should ensure that the membrane solution fully covers the functional area at the sensor tip, so that the outer membrane 800 covers all sensing areas. Finally, it is stabilized for 72 hours in a constant temperature and humidity chamber at 30–38°C and 50–70% humidity. In some embodiments, the outer membrane 800 solution also includes enzymes or catalysts. It should be noted that the present invention does not limit the aforementioned solvents; any solvent familiar to those skilled in the art is acceptable, such as acetone, chloroform, dimethylformamide, or dimethyl sulfoxide.

[0146] In this embodiment, a polyurethane outer film 800 is formed on the outermost layer of all sensing areas at the end of the sensor by an impregnation-coating method.

[0147] Medical-grade polyurethane (PU) was dissolved in a mixed solvent of tetrahydrofuran (THF) and N,N-dimethylformamide (DMF) (volume ratio 1:1) to prepare a 5% (w / v) homogeneous solution as the outer membrane 800 solution. The sensor with the prepared sensing membrane 500 was then fixed again on an immersion lifter and vertically immersed in the PU solution at a speed of 0.5 mm / s. After holding for 10 seconds, it was slowly lifted out at a speed of 0.2 mm / s. The sensor was then transferred to a constant temperature and humidity chamber and stabilized for 72 hours at 35°C and 60% humidity. During this process, the solvent slowly evaporated, and the PU molecular chains rearranged, forming a uniformly thick film with a controllable microcavity 403 structure.

[0148] The PU outer membrane 800 is approximately 5-30 μm thick. As a glucose flux limiting layer, it controls the diffusion rate of glucose to the sensing membrane 500, extending the sensor's linear detection range to the physiological concentration range (0-30 mM). Simultaneously, its biocompatible surface can reduce post-implantation inflammatory responses.

[0149] This embodiment provides a method for manufacturing a glucose sensor with a three-dimensional microcavity 403. The manufacturing steps are highly standardized, the process parameters are well controllable, and it is easy to achieve mass production, ensuring the consistency and stability of the performance of each batch of products. By creatively introducing a vacuum environment to assist impregnation, and using negative pressure to overcome the air resistance within the microcavity 403, the sensing membrane 500 solution can penetrate into every corner of the microcavity 403 without obstruction. This achieves complete filling and uniform distribution of the sensing membrane 500 within the three-dimensional microcavity 403, fully utilizing the area gain of the three-dimensional structure.

[0150] To further demonstrate the superiority of the present invention, two comparative examples are introduced below to verify the technical effects of the present invention: Comparative Example 2 The experimental setup was the same as that of Comparative Example 2 described in Example 1.

[0151] The experimental results (initial response: the response current of the sensor of this invention is 2030~2550 nA, while that of the control sensor is 1520~1800 nA. The response current of the sensor of this invention is 1.13~1.67 times that of the control; 7-day stability: after continuous immersion in PBS at 35℃ for 7 days, the response current of the sensor of this invention still retains 89~92%, while that of the control sensor only retains 78~83%) demonstrate that: The synergistic effect of microcavity 403 as the loading space of sensing membrane 500 and vacuum impregnation method—the physical anchoring effect (concave shape) of microcavity 403 and the full filling effect brought by vacuum impregnation work together to significantly enhance the adhesion and mechanical stability of sensing membrane 500, and prevent enzymes from falling off and becoming inactive during long-term use.

[0152] Comparative Example 3 Compared with sensors fabricated without vacuum conditions.

[0153] Comparison object: The control sensor was impregnated with the sensing membrane at 500°C under normal pressure.

[0154] Specifically, a polytetrafluoroethylene (PTFE) substrate approximately 1 mm wide and 0.1 mm thick was used to fabricate the sensor. Specifically, the microcavities 403 were arranged in a 1×4 array, with a pore diameter and inter-pore spacing of 400 μm. A platinum electrode was used as the counter electrode 221, and Ag / AgCl as the reference electrode 212. The sensor membrane 500 was fixed under both vacuum and ambient pressure to compare the performance differences of the sensors obtained under different conditions.

[0155] Test Results: SEM observation showed that after drying, the microcavity 403 structure of the sensor obtained under normal pressure was noticeably hollow and not completely filled with the sensing membrane 500 material. In contrast, the microcavity 403 structure of the sensor obtained under vacuum was filled with a uniform sensing membrane 500 material. Voltammetry tests were performed on both sensors in 5 mM glucose solution at 7-day intervals. The results showed that the response current of the sensor obtained under normal pressure was 80-86% of the initial current after 7 days, while the response current of the sensor obtained under vacuum remained at 89-92%.

[0156] Data source: SEM images were captured using a field emission scanning electron microscope; fill rate was calculated statistically using image analysis software.

[0157] Conclusion: A vacuum environment is crucial for ensuring that the 500 solution in the sensing membrane overcomes surface tension and fully penetrates the high aspect ratio microcavity 403. Complete filling results in greater enzyme loading and more robust fixation, thus conferring superior long-term stability to the sensor.

[0158] Example 3 A method of using a glucose sensor with a three-dimensional microcavity 403, as described in Example 1, includes: exposing the sensor's functional area to body fluid and acquiring the sensor current intensity under low potential conditions.

[0159] Specifically, a matching implantation device can be used to implant the functional area at the end of the sensor into the patient's subcutaneous interstitial fluid via percutaneous puncture. During implantation, ensure that the sensor is implanted horizontally along the superficial subcutaneous layer, with the functional area completely immersed in the interstitial fluid, and that the sensor tail and contact points remain outside the body. After implantation, the sensor is left to stand in the body for 30 minutes to 2 hours without applying any working potential, allowing the sensor to naturally equilibrate in the body fluid environment. The purpose of this stabilization period is twofold: firstly, to allow the sensing membrane 500 to reach ionic balance with the electrolytes in the interstitial fluid, eliminating initial drift; and secondly, to moderate the tissue stress response in the early stages of implantation, reducing the impact of inflammatory cell aggregation on the stability of the sensor interface.

[0160] After the stabilization period, the external signal processing system is activated to apply a predetermined operating potential relative to the reference electrode 212 to the first working electrode 211. Since ferricyanide electron mediators are incorporated into the sensing membrane 500, electrons generated by the glucose oxidation reaction can be transferred to the electrode surface through the mediator without relying on dissolved oxygen as a natural electron acceptor, thus the operating potential can be set in a lower range.

[0161] Under constant potential conditions, the system continuously collects the response current of the first working electrode 211 at a set sampling frequency (e.g., once every minute). The sampling time window is 100 milliseconds, and the average value of 5 measurements is taken for each sampling point to reduce noise.

[0162] When the sensor is configured in differential detection mode, the system synchronously applies the same operating potential to the second working electrode 222 as to the first working electrode 211. Since the sensing area 701 of the second working electrode is only covered by an inert conductive material and does not contain glucose oxidase, it is unresponsive to glucose. However, it still generates a background current similar to that of the first working electrode 211 in response to electrochemically active interfering substances in body fluids (such as ascorbic acid, uric acid, and acetaminophen). The system acquires the current signals from both working electrodes and calculates the difference in real time. This difference is used as the glucose-specific response signal. This differential processing directly subtracts common-mode interference at the hardware level, without relying on subsequent complex software algorithms, significantly improving the signal-to-noise ratio and measurement accuracy.

[0163] When the sensor is configured in multi-parameter detection mode and the second working electrode 222 is used for lactic acid detection, the system drives the two working electrodes separately using a time-sequential voltage scanning method. Specifically, in the first detection cycle (e.g., 0~500 ms), the system applies -0.1 V to the first working electrode 211 and collects the glucose response current; in the second detection cycle (e.g., 500~1000 ms), the system applies +0.4 V to the second working electrode 222 and collects the lactic acid response current. Lactate oxidase in the lactic acid sensing membrane 500 catalyzes the oxidation of lactic acid to produce hydrogen peroxide. At a potential of +0.4 V, hydrogen peroxide undergoes electrochemical oxidation on the surface of the platinum black electrode, and the resulting oxidation current is proportional to the lactic acid concentration. The two cycles alternate, with each parameter sampled once per minute. The two detection channels share the same reference electrode 212 and counter electrode 221, achieving synchronous real-time monitoring of glucose and lactic acid on a single sensor.

[0164] An external signal processing system wirelessly transmits the real-time calculated glucose concentration data to the patient's smart terminal device (such as a mobile phone, smartwatch, or dedicated receiver) via Bluetooth Low Energy protocol. The application on the terminal device stores the data, performs trend analysis, and displays it graphically, generating a dynamic blood glucose profile and issuing an alert when the blood glucose concentration exceeds a preset threshold (such as hypoglycemia 3.9 mmol / L or hyperglycemia 13.9 mmol / L).

[0165] like Figure 12As shown, the sensor's current intensity is directly proportional to the glucose solution concentration and exhibits a high current advantage, with a current response reaching the microampere level. Due to the sensor's microcavity 403 structure and the inclusion of electronic media, the sensor can achieve a large current response within a small implantation volume, which is beneficial for the accurate acquisition of glucose current signals.

[0166] The above descriptions are merely embodiments of the present invention. Commonly known structures and characteristics of the solutions are not described in detail here. Those skilled in the art are aware of all common technical knowledge in the field prior to the application date or priority date, are aware of all existing technologies in that field, and have the ability to apply conventional experimental methods prior to that date. Those skilled in the art can, under the guidance of this application, improve and implement this solution in combination with their own capabilities. Some typical known structures or methods should not be obstacles for those skilled in the art to implement this application. It should be noted that those skilled in the art can make several modifications and improvements without departing from the structure of the present invention. These should also be considered within the scope of protection of the present invention, and will not affect the effectiveness of the implementation of the present invention or the practicality of the patent.

Claims

1. A glucose sensor with a three-dimensional microcavity, characterized in that, include: A substrate having at least one microcavity penetrating the substrate; A conductive layer covers at least one side surface of the substrate and the inner wall of the microcavity, wherein the conductive layer located on the inner wall of the microcavity forms a three-dimensional microcavity sensing region. A sensing membrane is disposed within the three-dimensional microcavity sensing area for reacting with glucose to generate an electrochemical signal. The three-dimensional microcavity sensing region simultaneously provides the electrical connection path between the conductive layers on both sides of the substrate and the load space of the sensing film.

2. A glucose sensor with a three-dimensional microcavity according to claim 1, characterized in that, The conductive layer includes a first conductive layer located on a first side of the substrate and a second conductive layer located on a second side of the substrate; The first conductive layer is divided into a first working electrode and a reference electrode along its length, and the second conductive layer is divided into a second working electrode and a counter electrode along its length; the microcavity connects the region at the end of the first working electrode with the independent conductive region on the second conductive layer corresponding to that region.

3. A glucose sensor with a three-dimensional microcavity according to claim 2, characterized in that, The second working electrode is used to perform one of the following functions: Function a, together with the first working electrode, forms a differential electrode pair to cancel the background current signal generated by coexisting interfering substances; Function b, independently detects a second target substance other than glucose, the second target substance being selected from lactic acid, uric acid, alcohol or ascorbic acid.

4. A glucose sensor with a three-dimensional microcavity according to claim 1, characterized in that, The microcavities are multiple and arranged in a regular or random array; the aperture of the microcavities is 50~800μm, the spacing between adjacent microcavities is 50~800μm, and the depth-to-diameter ratio of the microcavities is 0.5~4.

5. A glucose sensor with a three-dimensional microcavity according to claim 1, characterized in that, The sensing membrane comprises glucose oxidase, a protective agent, an electron mediator, and a crosslinking agent; The protective agent is selected from bovine serum albumin, trehalose, or polyethylene glycol; The electron mediator is selected from ferricyanide, methylene blue, ferrocene or transition metal complexes; The crosslinking agent is selected from glyoxal, glutaraldehyde, adipaldehyde, 1,3-dioxane, etc. Butylene dialdehyde, terephthalaldehyde, 2,3-dihydroxyterephthalaldehyde or 2,4-dihydroxyterephthalaldehyde Hexadienal.

6. A glucose sensor with a three-dimensional microcavity according to claim 1, characterized in that, The conductive layer on the inner wall of the microcavity has a nanoscale rough surface with a surface roughness Ra of 10~100 nm and a porosity of less than 5%.

7. A glucose sensor with a three-dimensional microcavity according to claim 1, characterized in that, Also includes: An outer membrane is provided, covering the outside of the sensing membrane, the outer membrane having glucose and interfering substance flux limiting function and biocompatibility characteristics; the outer membrane comprises at least one polymer selected from polyamide, polysulfone, polydimethylsiloxane, nylon, polyvinyl chloride, polyvinylimidazolium, polyacrylate, polyurethane, acrylate derivatives, polyetherurethane, silicone, polyvinylidene fluoride or chitosan.

8. A glucose sensor with a three-dimensional microcavity according to claim 2, characterized in that, Also includes: A protective layer is disposed on the surface of the non-sensing area of ​​the first conductive layer and the second conductive layer, as well as in the partition groove between each electrode, for insulating and isolating each electrode and protecting the non-sensing area.

9. A glucose sensor with a three-dimensional microcavity according to claim 1, characterized in that, The functional area of ​​the sensor is located within a length range of 0.5~4mm at the end of the sensor, the width of the functional area is 0.2~2mm, and the thickness is 50~300μm; the overall length of the sensor is 2~50mm.

10. A method for manufacturing a glucose sensor with a three-dimensional microcavity, characterized in that, The method for preparing a glucose sensor with a three-dimensional microcavity as described in any one of claims 1-9 comprises the following steps: S1, forming at least one microcavity penetrating the substrate on a substrate; S2, a conductive layer is formed on at least one side surface of the substrate and the inner wall of the microcavity, so that the conductive layer on the inner wall of the microcavity constitutes a three-dimensional microcavity sensing region. S3, Under vacuum conditions, the sensing membrane solution containing glucose oxidase is immersed into the sensing area of ​​the three-dimensional microcavity, so that the sensing membrane is loaded on the inner wall of the microcavity. S4, crosslinking and curing the sensing membrane.

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