Microneedle Electrochemical Biosensor and its Preparation Method

Abstract

This invention proposes a microneedle electrochemical biosensor and its fabrication method. The sensor comprises: a substrate and a glucose sensing electrode, a sodium penicillin sensing electrode, a sodium chloride sensing electrode, a reference electrode, and a counter electrode disposed on the substrate. Specifically: the glucose sensing electrode comprises a first graphite layer, a Prussian blue layer, and an enzyme immobilization layer stacked from bottom to top; the sodium penicillin sensing electrode comprises a second graphite layer and a β-lactamase immobilization layer stacked from bottom to top; and the sodium chloride sensing electrode comprises a third graphite layer, a signal conversion layer, and an ion recognition membrane layer stacked from bottom to top. This sensor selectively measures the concentrations of glucose, sodium penicillin, and sodium chloride, achieving high sensitivity and accuracy in detection.

Description

Technical Field

[0001] This invention relates to the field of biosensor technology, and in particular to microneedle electrochemical biosensors and their preparation methods. Background Technology

[0002] In clinical treatment, glucose serves as an energy source, and monitoring its concentration is crucial for patient health, especially in the treatment of diabetes. Similarly, precise concentration monitoring of penicillin sodium in intravenous infusions is essential to ensure treatment efficacy and patient safety, as it is a commonly used antibiotic.

[0003] Existing microneedle electrochemical biosensors are typically single-purpose sensors, such as those for glucose. This limits their application in a wider range of fields and cannot meet the demand for a single microneedle electrochemical biosensor to detect multiple compounds or biomarkers. Furthermore, if the detection of multiple compounds or ions is required, multiple single-purpose sensors are needed, which increases the cost of the device. In addition, each sensor requires individual calibration and maintenance, further increasing the cost of use. Moreover, single-purpose sensors may mean that multiple devices need to be carried to meet different detection needs, which is detrimental to portability, especially in scenarios requiring mobile monitoring.

[0004] Therefore, it is necessary to design a multi-purpose microneedle electrochemical biosensor to meet the requirement that a single microneedle electrochemical biosensor can detect a variety of compounds or biomarkers, such as the concentration of glucose solution in an infusion bag, the concentration of sodium penicillin solution, etc. Summary of the Invention

[0005] To address the aforementioned shortcomings, this invention proposes a microneedle electrochemical biosensor and its preparation method, aiming to enable a single microneedle electrochemical biosensor to detect multiple compounds or biomarkers.

[0006] This invention provides a microneedle electrochemical biosensor, comprising: a substrate and a glucose sensing electrode, a sodium penicillin sensing electrode, a sodium chloride sensing electrode, a reference electrode, and a counter electrode disposed on the substrate; wherein:

[0007] The glucose sensing working electrode comprises a first graphite layer, a Prussian blue layer, and an enzyme immobilization layer stacked from bottom to top. The enzyme immobilization layer is composed of bovine serum albumin, glucose oxidase, and glutaraldehyde, and the glucose oxidase and glutaraldehyde cross-link to form a three-dimensional network structure.

[0008] The sodium penicillin sensing working electrode comprises a second graphite layer and a β-lactamase immobilization layer stacked from bottom to top. The β-lactamase immobilization layer is composed of β-lactamase and the electron medium cobalt phthalocyanine.

[0009] The sodium chloride sensing working electrode comprises a third graphite layer, a signal conversion layer, and an ion recognition film layer stacked from bottom to top. The signal conversion layer is composed of 3,4-ethylenedioxythiophene and sodium salt polystyrene sulfonic acid.

[0010] In one embodiment of the present invention, the mass ratio of bovine serum albumin, glucose oxidase, and glutaraldehyde is (8-12):(29-36):(0.16-0.20).

[0011] In one embodiment of the present invention, the mass ratio of the β-lactamase to the electron medium cobalt phthalocyanine is 10:12 to 20.

[0012] In one embodiment of the present invention, the mass ratio of 3,4-ethylenedioxythiophene to sodium polystyrene sulfonic acid is (0.05-0.09):(0.9-1.2).

[0013] In one embodiment of the present invention, the ion recognition membrane is formed by a selective membrane precursor solution coated on the outside of the signal conversion layer. The selective membrane precursor solution includes sodium tetrakis(3,5-bis(trifluoromethyl)phenyl)borate, high molecular weight polyvinyl chloride, di(2-ethylhexyl) sebacate, and tetraethyl 4-tert-butylcalixanetetraacetate, with the mass ratio of the four substances being (0.4-0.6):(30-35):(60-70):(0.8-1.2).

[0014] In one embodiment of the present invention, the Prussian blue layer is prepared by dissolving K3Fe(CN)6, FeCl3 and KCl in an acidic HCl solution, and the thickness of the Prussian blue layer is 5 μm to 30 μm and the pore size is 0.1 to 100 nm; the thickness of the enzyme immobilization layer is 0.1 μm to 40 μm and the pore size is 0.1 to 100 nm.

[0015] In one embodiment of the present invention, the first graphite layer, the second graphite layer, and the third graphite layer are all made of graphite powder or a mixture of the graphite powder and N-cyclohexyl-N-(2-morpholinoethyl)carbodiimide.

[0016] In one embodiment of the present invention, the thickness of the first graphite layer is 0.1 μm to 40 μm and the pore size is 0.1 to 100 nm; the thickness of the second graphite layer is 0.1 μm to 40 μm and the pore size is 0.1 to 100 nm; the thickness of the third graphite layer is 0.1 μm to 40 μm and the pore size is 0.1 to 100 nm.

[0017] In one embodiment of the present invention, the thickness of the β-lactamase immobilization layer is 0.1–40 μm, and the pore size is 0.1–100 nm.

[0018] In one embodiment of the present invention, the thickness of the signal conversion layer is 0.1 μm to 40 μm and the pore size is 0.1 to 100 nm, and the thickness of the ion recognition film layer is 0.1 μm to 40 μm and the pore size is 0.1 to 100 nm.

[0019] In one embodiment of the present invention, the microneedle electrochemical biosensor also integrates a processor chip.

[0020] In one embodiment of the present invention, the microneedle electrochemical biosensor also integrates a wireless sensor module, which uses Bluetooth, WiFi or Zigbee wireless communication technology.

[0021] In another aspect, the present invention provides a method for preparing a microneedle electrochemical biosensor, used to prepare the aforementioned microneedle electrochemical biosensor, characterized in that:

[0022] Prussian blue electrodeposition solution was coated onto the surface of the first graphite layer to obtain a Prussian blue layer; a glucose oxidase solution was coated onto the surface of the Prussian blue layer and bound together by covalent bonds to obtain an enzyme immobilization layer, thereby obtaining the glucose sensing working electrode.

[0023] The mixture obtained by adding cobalt phthalocyanine, an electronic medium, to β-lactamase and then adding it to an acetate buffer solution and mixing it with mineral oil is coated onto the surface of the second graphite layer to form a β-lactamase immobilization layer, thereby obtaining the sodium penicillin sensing working electrode.

[0024] A mixture of 3,4-ethylenedioxythiophene and sodium polystyrene sulfonic acid is electrodeposited on the surface of a third graphite layer to form a signal conversion layer. A selective membrane precursor solution is then coated on the surface of the signal conversion layer to form an ion recognition membrane layer, thereby obtaining the sodium chloride sensing working electrode.

[0025] In one embodiment of the present invention, the selective membrane precursor solution comprises sodium tetrakis(3,5-bis(trifluoromethyl)phenyl)borate, high molecular weight polyvinyl chloride, di(2-ethylhexyl) sebacate, and tetraethyl 4-tert-butylcalixanetetraacetate.

[0026] In one embodiment of the present invention, the signal conversion layer is generated by electrodeposition of a mixture of 3,4-ethylenedioxythiophene and sodium polystyrene sulfonic acid via constant current electrochemical polymerization using a reference electrode and a counter electrode.

[0027] In one embodiment of the present invention, K3Fe(CN)6, FeCl3 and KCl are dissolved in an acidic HCl solution and mixed evenly to obtain the Prussian blue electrodeposition solution.

[0028] Bovine serum albumin and glucose oxidase were each dissolved in a phosphate buffer solution to obtain the glucose oxidase solution.

[0029] Glutaraldehyde was used as a cross-linking agent to cross-link with glucose oxidase to form covalent bonds, thereby immobilizing glucose oxidase and bovine serum albumin on the Prussian blue layer to form a three-dimensional network structure.

[0030] In one embodiment of the present invention, the method further includes: applying graphite powder or a mixture of graphite powder and N-cyclohexyl-N-(2-morpholinoethyl)carbodiimide to the ends of the conductive traces of the glucose sensing working electrode, the sodium penicillin sensing working electrode, and the sodium chloride sensing working electrode, and then screen printing the mixture to obtain the first graphite layer, the second graphite layer, and the third graphite layer.

[0031] In one embodiment of the present invention, the method further includes: screen printing Ag / AgCl ink at the end of the conductive trace of the reference electrode to obtain the reference electrode;

[0032] The counter electrode is obtained by screen printing platinum wire ink at the end of the conductive trace of the counter electrode.

[0033] The advantages of this invention are:

[0034] The microneedle electrochemical biosensor provided by this invention integrates three working electrodes: a glucose sensing electrode, a penicillin sodium sensing electrode, and a sodium chloride sensing electrode. By matching different working electrodes with the counter electrode and reference electrode, it can selectively measure the concentrations of glucose, penicillin sodium, and sodium chloride. Through modification of the working electrodes, it can achieve high sensitivity and a wide linear detection range. This microneedle electrochemical biosensor has high integration and is characterized by intelligence, flexibility, compactness, and stable performance. Furthermore, this sensor has a wide range of applications, such as rapid and accurate monitoring of the composition and concentration of chemical substances during infusion, improving the safety and effectiveness of infusion therapy.

[0035] To provide a better understanding of the above and other aspects of the present invention, specific embodiments are described below in conjunction with the accompanying drawings. Attached Figure Description

[0036] Figure 1 A schematic diagram of the structure of a microneedle electrochemical biosensor provided according to an embodiment of the present invention is shown.

[0037] Figure 2 A schematic diagram of the working electrode for the glucose sensor is shown.

[0038] Figure 3 A schematic diagram of the working electrode of the sodium penicillin sensor is shown.

[0039] Figure 4 A schematic diagram of the working electrode of the sodium chloride sensor is shown.

[0040] Figure 5 A schematic diagram illustrating the overall process of fabricating a microneedle electrochemical biosensor according to an embodiment of the present invention is shown.

[0041] Figures 6(a)-6(b) The results of monitoring glucose concentration in infusion bags using a microneedle electrochemical biosensor are shown in Figure 6(a), which shows the time-ampere test results, and Figure 6(b), which shows the linear relationship between ampere current and glucose concentration.

[0042] Figures 7(a)-7(b) The results of monitoring the concentration of sodium penicillin in infusion bags using a microneedle electrochemical biosensor are shown in Figure 7(a), which shows the results of the differential pulse voltammetry test, and Figure 7(b) shows the linear relationship between peak current and sodium penicillin concentration.

[0043] Figures 8(a)-8(b) The results of monitoring sodium chloride concentration in infusion bags using a microneedle electrochemical biosensor are shown in Figure 8(a). Figure 8(a) shows the open-circuit potential test results of the Na+ sensor in an electrolyte solution with a concentration of (10-320)×10-3mM Na+. Figure 8(b) shows the linear relationship between open-circuit potential and sodium ion concentration.

[0044] In the attached figures, the following labels are used:

[0045] 10: Microneedle electrochemical biosensors;

[0046] 11: Glucose sensing working electrode;

[0047] 111: First graphite layer;

[0048] 112: Prussian Blue Layer;

[0049] 113: Enzyme immobilization layer;

[0050] 12: Sodium penicillin sensor working electrode;

[0051] 121: Second graphite layer;

[0052] 122: β-lactamase immobilization layer;

[0053] 13: Sodium chloride sensor working electrode;

[0054] 131: Third graphite layer:

[0055] 132: Signal conversion layer;

[0056] 133: Ion recognition membrane layer;

[0057] 14: Reference electrode;

[0058] 15: Counter electrode;

[0059] 16: Base. Detailed Implementation

[0060] like Figure 1 As shown, Figure 1 A schematic diagram of a microneedle electrochemical biosensor structure provided in an embodiment of the present invention is shown. The microneedle electrochemical biosensor 10 specifically includes a substrate 16 and a glucose sensing working electrode 11, a penicillin sodium sensing working electrode 12, a sodium chloride sensing working electrode 13, a reference electrode 14, and a counter electrode 15 disposed on the substrate.

[0061] In one embodiment, the microneedle electrochemical biosensor 10 uses stainless steel as a substrate. A glucose sensing electrode 11, a sodium penicillin sensing electrode 12, a sodium chloride sensing electrode 13, a reference electrode 14, and a counter electrode 15 work together on the microneedle array structure to achieve highly sensitive and selective detection of biomolecules. The glucose sensing electrode 11, sodium penicillin sensing electrode 12, and sodium chloride sensing electrode 13 serve as sites for electrochemical reactions, reacting with the analyte compound and generating a current signal. The reference electrode provides a stable potential reference to ensure measurement accuracy. The counter electrode balances the current, providing equal but opposite currents to allow the electrochemical reaction to continue.

[0062] Specifically, refer to Figure 2 As shown, in one embodiment, the glucose sensing working electrode 11 comprises a first graphite layer 111, a Prussian blue layer 112, and an enzyme immobilization layer 113 stacked from bottom to top. In one embodiment, the thickness of the first graphite layer is 0.1 μm to 40 μm, and the pore size is 0.1 to 100 nm; the thickness of the Prussian blue layer is 5 μm to 30 μm, and the pore size is 0.1 to 100 nm; the thickness of the enzyme immobilization layer is 0.1 μm to 40 μm, and the pore size is 0.1 to 100 nm.

[0063] In this embodiment, the first graphite layer is made of graphite powder (GP) or a mixture of the graphite powder and N-cyclohexyl-N-(2-morpholinoethyl)carbodiimide (CDI).

[0064] The Prussian blue layer is prepared by electrodeposition of K3Fe(CN)6, FeCl3, and KCl dissolved in an acidic HCl solution. Specifically, K3Fe(CN)6, FeCl3, and KCl are dissolved in an acidic HCl solution and mixed thoroughly to obtain a Prussian blue electrodeposition solution. The Prussian blue electrodeposition solution is then coated onto the surface of a first graphite layer, and multiple cyclic voltammetry cycles are performed within a given potential range to deposit the Prussian blue layer.

[0065] The enzyme immobilization layer is formed by coating the surface of the Prussian blue layer with an oxidase solution and binding them together through covalent bonds. Specifically, the enzyme immobilization layer comprises bovine serum albumin (BSA), glucose oxidase, and glutaraldehyde. BSA and GOx are each dissolved in phosphate-buffered saline (PBS) to obtain a glucose oxidase solution. This glucose oxidase solution is then coated onto the Prussian blue layer and dried. Glutaraldehyde is used as a cross-linking agent to react with the glucose oxidase (GOx) to form covalent bonds, thereby firmly immobilizing the glucose oxidase and BSA on the Prussian blue layer, forming a three-dimensional network structure. In one embodiment, the preferred mass ratio of BSA, glucose oxidase, and glutaraldehyde is (8–12):(29–36):(0.16–0.20), more preferably 10:30:0.18.

[0066] refer to Figure 3 As shown, in one embodiment, the penicillin sodium sensing working electrode 12 includes a second graphite layer 121 and a β-lactamase immobilization layer 122 stacked from bottom to top. In one embodiment, the thickness of the second graphite layer 121 is 0.1 μm to 40 μm, and the pore size is 0.1 to 100 nm. The thickness of the β-lactamase immobilization layer 122 is 0.1 to 40 μm, and the pore size is 0.1 to 100 nm.

[0067] In this embodiment, the second graphite layer, like the first graphite layer, is made of graphite powder (GP) or a mixture of the graphite powder and N-cyclohexyl-N-(2-morpholinoethyl)carbodiimide (CDI).

[0068] The β-lactamase immobilization layer comprises β-lactamase and the electron mediator cobalt phthalocyanine. Specifically, the β-lactamase is mixed with cobalt phthalocyanine in an acetate buffer solution and then coated onto the surface of a second graphite layer to form the β-lactamase immobilization layer. In one embodiment, the mass ratio of β-lactamase to cobalt phthalocyanine is 10:12 to 20, more preferably 10:15.

[0069] refer to Figure 4 As shown, in one embodiment, the sodium chloride sensing working electrode 13 includes a third graphite layer 131, a signal conversion layer 132, and an ion recognition film layer 133 stacked from bottom to top. In one embodiment, the thickness of the third graphite layer 131 is 0.1 μm to 40 μm, and the pore size is 0.1 to 100 nm; the thickness of the signal conversion layer 132 is 0.1 μm to 40 μm, and the pore size is 0.1 to 100 nm; the thickness of the ion recognition film layer 133 is 0.1 μm to 40 μm, and the pore size is 0.1 to 100 nm.

[0070] In this embodiment, the third graphite layer is also made of graphite powder (GP) or a mixture of the graphite powder and N-cyclohexyl-N-(2-morpholinoethyl)carbodiimide (CDI).

[0071] The signal conversion layer comprises 3,4-ethylenedioxythiophene and sodium polystyrene sulfonic acid. Specifically, using a reference electrode and a counter electrode, a mixture of 3,4-ethylenedioxythiophene (EDOT) and sodium polystyrene sulfonic acid (NaPSS) is electrodeposited onto the surface of the third graphite layer via constant current electrochemical polymerization to generate the signal conversion layer. Then, a mixture containing sodium tetrakis(3,5-bis(trifluoromethyl)phenyl)borate (Na-TFPB), high molecular weight polyvinyl chloride (PVC), di(2-ethylhexyl) sebacate (DOS), and tetraethyl 4-tert-butylcalixanetetraacetate (sodium ion carrier X) is dissolved in tetrahydrofuran to obtain a selective membrane precursor solution. This selective membrane precursor solution is then coated onto the surface of the signal conversion layer to form an ion recognition membrane layer.

[0072] In one embodiment, the mass ratio of 3,4-ethylenedioxythiophene to sodium polystyrene sulfonic acid is preferably (0.05-0.09):(0.9-1.2), more preferably 0.07:1.

[0073] Sodium tetra(3,5-bis(trifluoromethyl)phenyl)borate, high molecular weight polyvinyl chloride, di(2-ethylhexyl) sebacate, and tetraethyl tetraacetate of 4-tert-butylcalixane-tetraacetate, wherein the mass ratio of sodium tetra(3,5-bis(trifluoromethyl)phenyl)borate, high molecular weight polyvinyl chloride, di(2-ethylhexyl) sebacate, and tetraethyl tetraacetate of 4-tert-butylcalixane-tetraacetate is preferably (0.4-0.6):(30-35):(60-70):(0.8-1.2), more preferably 0.55:33:65.45:1.

[0074] In addition, in one embodiment, the reference electrode is prepared based on Ag / AgCl filling, and the counter electrode is prepared based on platinum wire filling.

[0075] In addition, in one embodiment, the microneedle electrochemical biosensor also integrates a wireless sensor module, which uses Bluetooth, WiFi or Zigbee wireless communication technology.

[0076] In specific applications of microneedle electrochemical biosensors, such as detecting the concentration of glucose solution in infusion bags, the microneedle electrochemical biosensor is connected to an external electrochemical analyzer. In some embodiments, the electrochemical analyzer should have electrode interfaces compatible with the microneedle electrochemical biosensor. These interfaces typically include a working electrode interface, a reference electrode interface, and a counter electrode interface. Each electrode of the microneedle electrochemical biosensor is connected to the corresponding interface of the electrochemical analyzer via cable or wireless means to ensure that the electrochemical analyzer can accurately receive and process signals from the microneedle electrochemical biosensor.

[0077] Furthermore, for different electrochemical analyzers, various methods such as potential analysis, current analysis, conductivity analysis, and electrochemical impedance spectroscopy can be employed to achieve accurate concentration monitoring. For example, potential analysis infers the concentration or properties of the analyte by measuring the potential difference between the working electrode and the reference electrode. This method relies on the Nernst equation, which states that the potential is linearly related to the logarithm of the analyte concentration. Current analysis measures the current flowing through the working electrode at a constant potential; this current is proportional to the concentration of the analyte. Conductivity analysis infers the concentration or properties of the analyte by measuring the conductivity of the solution; conductivity is related to the concentration and mobility of ions in the solution. Electrochemical impedance spectroscopy applies a small-amplitude AC signal to the electrochemical cell and measures the system's impedance response. By analyzing the impedance spectrum, useful information about electrode surface processes, electrolyte properties, and the interaction between the analyte and the electrode can be obtained.

[0078] In addition, in one embodiment, the microneedle electrochemical biosensor also integrates a processor chip to process and analyze the raw current signal generated by the electrochemical biosensor, extract useful biomarker information, realize real-time monitoring data, and transmit it to a smartphone or cloud server via wireless communication.

[0079] Furthermore, in one embodiment, the electrochemical biosensor is fabricated using CNC (Computer Numerical Control) micromachining technology. CNC micromachining is a high-precision machining technique suitable for fabricating micro- or nanoscale components and structures. In sensor fabrication, CNC micromachining technology can achieve high-precision cutting, engraving, and shaping of materials through precise numerical control programming. CNC micromachining technology has significant advantages in manufacturing microneedle electrochemical biosensors. Its application in the biosensor field not only improves manufacturing accuracy and consistency but also enhances the ability to manufacture complex sensor structures, production efficiency, while reducing costs and improving performance. Specifically, firstly, a structural model of the microneedle array is designed using computer-aided design software, including the layout and shape of the working electrode, counter electrode, and reference electrode. Simultaneously, parameters such as the size, spacing, and array density of the microneedles are determined. Suitable conductive materials are selected as electrode materials, such as metals (gold, silver, platinum, etc.) or conductive polymers. The materials are pretreated, such as by cleaning, degreasing, and drying, to ensure processing quality and electrode performance. Next, the designed microneedle array structural model is imported into the CNC machining equipment for programming and settings. CNC micromachining technology is then used to precisely cut or carve the material to form the microneedle array structure. During this process, precise control of machining parameters, such as cutting speed, feed rate, and depth of cut, is required to ensure the dimensional accuracy and surface quality of the microneedles. Finally, the machined microneedle array is assembled with the sensor substrate to form a complete microneedle electrochemical biosensor.

[0080] This embodiment discloses a microneedle electrochemical biosensor that integrates three working electrodes: a glucose sensing electrode, a penicillin sodium sensing electrode, and a sodium chloride sensing electrode. By matching different working electrodes with the counter electrode and reference electrode, it can selectively measure the concentrations of glucose, penicillin sodium, and sodium chloride. This microneedle electrochemical biosensor has a high degree of integration, integrating electrodes, a stable power supply and processor, a wireless sensing module, and a signal receiving and conversion module into a single intelligent, flexible, compact, and stable microneedle electrochemical biosensor.

[0081] Another embodiment of the present invention also provides a method for preparing the above-mentioned microneedle electrochemical biosensor, which can be found in the following description. Figure 5 As shown, Figure 5 A schematic diagram of the overall process for fabricating a microneedle electrochemical biosensor is shown.

[0082] A method for preparing a microneedle electrochemical biosensor specifically includes the following steps:

[0083] Step S1: Coat the surface of the first graphite layer with Prussian blue electrodeposition solution to obtain a Prussian blue layer; coat the surface of the Prussian blue layer with glucose oxidase solution and bind it with covalent bonds to obtain an enzyme immobilization layer, so as to obtain the glucose sensing working electrode.

[0084] Specifically, in step S1, graphite powder or a mixture of graphite powder and N-cyclohexyl-N-(2-morpholinoethyl)carbodiimide is first coated onto the end of the conductive trace of the glucose sensing working electrode and screen-printed to obtain the first graphite layer. In one embodiment, the solvent is water, acetic acid, etc. Then, K3Fe(CN)6, FeCl3, and KCl are dissolved in an acidic HCl solution and mixed evenly to obtain a Prussian blue electrodeposition solution. The Prussian blue electrodeposition solution is coated onto the surface of the first graphite layer, and multiple cyclic voltammetry cycles are performed within a given potential range (e.g., -0.5V to 0.6V) to deposit the Prussian blue layer. For example, in one embodiment, 2.5 mM K3Fe(CN)6, 2.5 mM FeCl3, and 0.1 M KCl were dissolved in 0.1 M HCl acidic solution and mixed thoroughly to obtain a PB electrodeposition solution. The modified electrode was immersed in the Prussian blue electrodeposition solution and subjected to 6 cyclic voltammetry (CV) cycles within a given potential range of -0.5 V to 0.6 V to deposit the Prussian blue layer as a sensing medium. However, the contents of K3Fe(CN)6, FeCl3, KCl, and HCl are not limited to this. Finally, bovine serum albumin (BSA) and glucose oxidase (GOx) were each dissolved in phosphate-buffered saline (PBS) to obtain a glucose oxidase solution. The glucose oxidase solution was coated onto the Prussian blue layer and dried to obtain an enzyme immobilization layer. Glutaraldehyde is used as a cross-linking agent to react with glucose oxidase (GOx) to form covalent bonds, thereby firmly immobilizing glucose oxidase and bovine serum albumin (BSA) on the Prussian blue layer to form a three-dimensional network structure, ultimately obtaining the glucose sensing working electrode. In one embodiment, the preferred mass ratio of BSA, glucose oxidase, and glutaraldehyde is (8-12):(29-36):(0.16-0.20), more preferably 10:30:0.18. For example, in one embodiment, 10 mg of BSA and 30 mg of glucose oxidase (GOx) are each dissolved in 1 mL of phosphate-buffered saline (PBS) to obtain a glucose oxidase solution. The prepared glucose oxidase solution is then coated onto the first graphite layer and dried in an oven at 40°C. 18 mg of glutaraldehyde (GA) is then used to react with glucose oxidase (GOx) to form a three-dimensional network structure, resulting in the enzyme immobilization layer.

[0085] This embodiment uses an enzyme immobilization layer formed by Prussian blue, bovine serum albumin (BSA), and glucose oxidase to modify the glucose sensing working electrode, which significantly improves the electron transfer efficiency and electron conduction capability of the electrode, increases the redox current of the sensor, and improves the sensitivity of the sensor in detecting glucose.

[0086] Step S2: After adding cobalt phthalocyanine, an electronic medium, to the β-lactamase, add it to an acetate buffer solution and mix it with mineral oil to form a carbon paste. Coat the carbon paste onto the surface of the second graphite layer to form a β-lactamase immobilization layer, thereby obtaining the sodium penicillin sensing working electrode.

[0087] Specifically, in step S2, graphite powder or a mixture of graphite powder and N-cyclohexyl-N-(2-morpholinoethyl)carbodiimide is first screen-printed onto the ends of the conductive traces of the sodium penicillin sensing working electrode to obtain a second graphite layer. Then, a mixture of β-lactamase with added cobalt phthalocyanine (electron mediator) is added to an acetate buffer solution and mixed with mineral oil, and coated onto the surface of the second graphite layer to form a β-lactamase immobilization layer, ultimately obtaining the sodium penicillin sensing working electrode. In one embodiment, the mass ratio of β-lactamase to cobalt phthalocyanine is 10:12 to 20, more preferably 10:15. For example, 7.5 mg of cobalt phthalocyanine is added to 10 U of β-lactamase and mixed with 40 μL of mineral oil in an acetate buffer solution.

[0088] This embodiment uses β-lactamase, which is fixed by covalent bonding, to modify the working electrode of the penicillin sodium sensor. This significantly improves the electron transfer efficiency and electron conduction capability of the electrode, increases the redox current of the sensor, and enhances the sensitivity of the sensor in detecting penicillin sodium.

[0089] Step S3: Electrodeposit a mixture of 3,4-ethylenedioxythiophene and sodium salt polystyrene sulfonic acid onto the third graphite layer to generate a signal conversion layer. Coat the surface of the signal conversion layer with a selective membrane precursor solution to form an ion recognition membrane layer, thereby obtaining the sodium chloride sensing working electrode.

[0090] Specifically, in step S3, graphite powder or a mixture of graphite powder and N-cyclohexyl-N-(2-morpholinoethyl)carbodiimide is first coated onto the end of the conductive trace of the sodium chloride sensing working electrode and screen-printed to obtain a third graphite layer. Using a reference electrode and a counter electrode, a mixture of 3,4-ethylenedioxythiophene (EDOT) and sodium polystyrene sulfonic acid (NaPSS) is electrodeposited onto the third graphite layer via constant current electrochemical polymerization to generate a signal conversion layer. Then, a mixture containing sodium tetrakis(3,5-di(trifluoromethyl)phenyl)borate (Na-TFPB), high molecular weight polyvinyl chloride (PVC), di(2-ethylhexyl) sebacate (DOS), and tetraethyl 4-tert-butylcalixarene-tetraacetate (sodium ion carrier X) is dissolved in tetrahydrofuran to obtain a selective membrane precursor solution. This selective membrane precursor solution is then coated onto the surface of the signal conversion layer to form an ion recognition membrane layer. In one embodiment, the mass ratio of 3,4-ethylenedioxythiophene to sodium polystyrene sulfonic acid is preferably (0.05–0.09):(0.9–1.2), more preferably 0.07:1. The mass ratio of sodium tetra(3,5-di(trifluoromethyl)phenyl)borate, high molecular weight polyvinyl chloride, di(2-ethylhexyl) sebacate, and tetraethyl 4-tert-butylcalixanetetraacetate is preferably (0.4–0.6):(30–35):(60–70):(0.8–1.2), more preferably 0.55:33:65.45:1. For example, in one embodiment, a mixture of 0.01 M 3,4-ethylenedioxythiophene (EDOT) and 0.1 M sodium polystyrene sulfonic acid (NaPSS) was deposited. 100 mg of a mixture containing Na-TFPB, high molecular weight PVC, DOS, and sodium ion carrier X in a mass ratio of 0.55:33:65.45:1 was dissolved in 660 μL of tetrahydrofuran.

[0091] In this embodiment, a signal conversion layer is generated by electrodeposition of a mixture of 3,4-ethylenedioxythiophene and sodium polystyrene sulfonic acid. An ion recognition film layer is generated by forming a sodium chloride sensing electrode using sodium tetrakis(3,5-bis(trifluoromethyl)phenyl)borate, high molecular weight polyvinyl chloride, di(2-ethylhexyl) sebacate, and tetraethyl 4-tert-butylcalixarene-tetraacetate. The change in the electronic structure of the signal conversion layer leads to a change in the electrode potential, resulting in a current signal proportional to the sodium ion concentration, which improves the sensitivity of the sensor in detecting sodium chloride.

[0092] In addition, in one embodiment, for the preparation of the reference electrode and the counter electrode, Ag / AgCl ink is screen-printed at the end of the conductive trace of the reference electrode to obtain the reference electrode; platinum wire ink is screen-printed at the end of the conductive trace of the counter electrode to obtain the counter electrode.

[0093] The microneedle electrochemical biosensor provided in this embodiment integrates three working electrodes: a glucose sensing working electrode, a penicillin sodium sensing working electrode, and a sodium chloride sensing working electrode. By matching different working electrodes with the counter electrode and reference electrode, it can flexibly and selectively measure the concentrations of glucose, penicillin sodium, and sodium chloride.

[0094] Taking the detection of drug concentration in an infusion bag using this microneedle electrochemical biosensor as an example:

[0095] Test Example 1

[0096] This test example demonstrates the results of using a microneedle electrochemical biosensor to monitor glucose concentration in an infusion bag.

[0097] When detecting glucose concentration in an infusion bag, a microneedle electrochemical biosensor is installed on the bag. The glucose sensing electrode, counter electrode, and reference electrode are immersed in the solution, while the penicillin sodium sensing electrode and sodium chloride sensing electrode remain in contact with the liquid. During infusion, glucose is oxidized by glucose oxidase through free diffusion to produce hydrogen peroxide (H₂O₂). This hydrogen peroxide rapidly decomposes on the surface of Prussian blue nanoparticles, reducing to water. Electrons are generated during this decomposition, and these electrons are transferred to the electrodes, causing a change in current. The electrochemical analyzer converts the recorded current signal into glucose concentration, enabling real-time continuous monitoring of the glucose level in the infusion bag.

[0098] In this embodiment, the sensor's response to 0.25 mM, 5 mM, 10 mM, 15 mM, 20 mM, and 36 mM glucose (Glu) was characterized using a chronoamperometry method in 0.01 M phosphate-buffered saline (PBS). The operating voltage was -0.2 V, the sampling interval was 0.1 s, and the settling time was 0 s. All electrochemical tests were performed at room temperature using a microneedle electrochemical biosensor.

[0099] The linear range and sensitivity of this sensor allow for continuous analysis of glucose concentration using a current-time curve (it curve) at a potential of -0.2V. Figure 6(a) shows the time-ampere test results, and Figure 6(b) shows the linear relationship between ampere current and glucose concentration. Figure 6(a) shows that the steady-state current response signal of the it curve increases with increasing glucose concentration. Figure 6(b) shows that the steady-state current response is linearly related to glucose concentration in the range of 0.25-36 mM (R² = 0.9962).

[0100] Test Example 2

[0101] Test Example 2 shows the results of using a microneedle electrochemical biosensor to monitor the concentration of sodium penicillin in an infusion bag.

[0102] When detecting the concentration of sodium penicillin in the infusion bag, a microneedle electrochemical biosensor is installed on the infusion bag. The working electrode, counter electrode, and reference electrode for the sodium penicillin sensor are immersed in the solution, while the working electrodes for the glucose and sodium chloride sensors remain in contact with the liquid. During the infusion process, β-lactamase catalyzes the hydrolysis of sodium penicillin to produce penicillic acid. Cobalt phthalocyanine, an electron mediator, participates in electron transfer, generating an oxidation current signal. The current signal recorded by the electrochemical analyzer is converted into the sodium penicillin concentration, enabling real-time continuous monitoring of the sodium penicillin concentration in the infusion bag.

[0103] In this embodiment, differential pulse voltammetry (DPV) was specifically used to qualitatively analyze the reduction peak potentials of 5 mM, 10 mM, 20 mM, 25 mM, 35 mM, and 50 mM penicillin sodium in 0.01 M phosphate buffered saline (PBS). The potential range was 0.3–0.6 V, the pulse width was 0.06 s, the pulse period was 0.5 s, and the amplitude was 50 mV. All electrochemical tests were performed at room temperature using a microneedle electrochemical biosensor.

[0104] DPV scans were recorded at different concentrations of penicillin sodium, with a potential range of -0.3 to 0.6 V, a pulse width of 0.06 s, a pulse period of 0.5 s, and an amplitude of 50 mV. Figure 7(a) shows the results of the differential pulse voltammetry test, and Figure 7(b) shows the linear relationship between peak current and penicillin sodium concentration. Figure 7(a) shows that there is a linear relationship between the peak current and the 10 different concentrations of penicillin sodium solution; the peak current of penicillin sodium increases linearly with the concentration increasing from 0 mM to 50 mM. Figure 7(b) shows that the calibration curve for detecting penicillin sodium concentration using differential pulse voltammetry (DPV) is linear between 0 mM and 50 mM.

[0105] Test Example 3

[0106] Test Example 3 shows the results of using a microneedle electrochemical biosensor to monitor the sodium chloride concentration in an infusion bag.

[0107] When detecting the sodium chloride concentration in the infusion bag, a microneedle electrochemical biosensor is installed in the infusion bag. The sodium chloride sensing working electrode, counter electrode, and reference electrode are immersed in the medication solution, while the glucose and penicillin sodium sensing working electrodes are not in contact with the liquid. The sodium ion concentration is detected using the open-circuit potential method. During the infusion process, when sodium ions come into contact with the ion recognition membrane, sodium tetrakis(3,5-bis(trifluoromethyl)phenylborate acts as an ion carrier and binds to the sodium ions, causing a change in the electronic structure of the signal conversion layer, resulting in a change in electrode potential and generating a current signal. The change in electrode potential caused by the change in the electronic structure of the signal conversion layer is recorded by an electrochemical analyzer, and measured using the open-circuit potential method (OCPT). A current signal proportional to the sodium ion concentration can be obtained, thus achieving real-time monitoring of the sodium ion concentration. In this embodiment, the open-circuit potential method (OCPT) is specifically used to detect 10 × 10⁻⁶ sodium ions. -3 mM, 50×10 -3 mM, 100×10 -3 mM, 150×10 -3 mM, 200×10 -3 mM, 320×10 -3 The open-circuit voltage (OCP) of mM Na+ was measured over a time of 100 s. All electrochemical measurements were performed at room temperature using a microneedle electrochemical biosensor. Figure 8(a) shows the concentration at (10-320)×10⁻⁶ mM Na+. -3 mM Na + The open-circuit potential test results of the Na+ sensor in the electrolyte solution are shown in Figure 8(b). Figure 8(b) shows the linear relationship between the open-circuit potential and the sodium ion concentration. As can be seen from Figures 8(a) and 8(b), the sensor has excellent sensing performance, with a concentration sensitivity of 53.08 mV dec. -1 .

[0108] In summary, the microneedle electrochemical biosensor provided by this invention integrates three working electrodes: a glucose sensing electrode, a penicillin sodium sensing electrode, and a sodium chloride sensing electrode. By matching different working electrodes with the counter electrode and reference electrode, it can selectively measure the concentrations of glucose, penicillin sodium, and sodium chloride. Through modification of the working electrodes, high sensitivity and a wide linear detection range can be achieved. This microneedle electrochemical biosensor has a high degree of integration, combining electrodes, a stable power supply and processor, a wireless sensing module, and a signal receiving and conversion module into a single, intelligent, flexible, compact, and stable microneedle electrochemical biosensor. Furthermore, this sensor has a wide range of applications, such as rapid and accurate monitoring of the composition and concentration of chemical substances during infusion, improving the safety and effectiveness of infusion therapy.

[0109] In summary, although this application has been disclosed above with reference to embodiments, it is not intended to limit this application. Those skilled in the art to which this application pertains can make various modifications and refinements without departing from the spirit and scope of this application. Therefore, the scope of protection of this application is determined by the claims.

Claims

1. A microneedle electrochemical biosensor, characterized by, include: A substrate and a glucose sensing electrode, a sodium penicillin sensing electrode, a sodium chloride sensing electrode, a reference electrode, and a counter electrode disposed on the substrate; wherein: The glucose sensing working electrode comprises a first graphite layer, a Prussian blue layer, and an enzyme immobilization layer stacked from bottom to top. The enzyme immobilization layer is composed of bovine serum albumin, glucose oxidase, and glutaraldehyde, and the glucose oxidase and glutaraldehyde cross-link to form a three-dimensional network structure. The penicillin sodium sensing working electrode comprises a second graphite layer and a β-lactamase immobilization layer stacked from bottom to top. The β-lactamase immobilization layer is composed of β-lactamase and the electron medium cobalt phthalocyanine. The sodium chloride sensing working electrode comprises a third graphite layer, a signal conversion layer, and an ion recognition film layer stacked from bottom to top. The signal conversion layer is composed of 3,4-ethylenedioxythiophene and sodium salt polystyrene sulfonic acid.

2. The microneedle electrochemical biosensor of claim 1, wherein, The mass ratio of bovine serum albumin, glucose oxidase, and glutaraldehyde is (8-12):(29-36):(0.16-0.20).

3. The microneedle electrochemical biosensor of claim 1, wherein, The mass ratio of the β-lactamase to the electron medium cobalt phthalocyanine is 10:12 to 20.

4. The microneedle electrochemical biosensor of claim 1, wherein, The mass ratio of 3,4-ethylenedioxythiophene to sodium polystyrene sulfonic acid is (0.05-0.09):(0.9-1.2).

5. The microneedle electrochemical biosensor of claim 1, wherein, The ion recognition membrane is formed by coating a selective membrane precursor solution on the outside of the signal conversion layer. The selective membrane precursor solution includes sodium tetrakis(3,5-bis(trifluoromethyl)phenyl)borate, high molecular weight polyvinyl chloride, di(2-ethylhexyl) sebacate, and tetraethyl 4-tert-butylcalixanetetraacetate, with the mass ratio of the four substances being (0.4-0.6):(30-35):(60-70):(0.8-1.2).

6. The microneedle electrochemical biosensor of claim 1, wherein, The Prussian blue layer is made by dissolving K3Fe(CN)6, FeCl3 and KCl in an acidic HCl solution. The thickness of the Prussian blue layer is 5 μm to 30 μm and the pore size is 0.1 to 100 nm. The thickness of the enzyme immobilization layer is 0.1 μm to 40 μm and the pore size is 0.1 to 100 nm.

7. The microneedle electrochemical biosensor of claim 1, wherein, The first graphite layer, the second graphite layer, and the third graphite layer are all made of graphite powder or a mixture of the graphite powder and N-cyclohexyl-N-(2-morpholinoethyl)carbodiimide.

8. The microneedle electrochemical biosensor according to claim 1 or 7, characterized in that, The thickness of the first graphite layer is 0.1 μm to 40 μm, and the pore size is 0.1 to 100 nm; The thickness of the second graphite layer is 0.1 μm to 40 μm, and the pore size is 0.1 to 100 nm; The thickness of the third graphite layer is 0.1 μm to 40 μm, and the pore size is 0.1 to 100 nm.

9. The microneedle electrochemical biosensor of claim 1, wherein, The thickness of the β-lactamase immobilization layer ranges from 0.1 to 40 μm, and the pore size ranges from 0.1 to 100 nm.

10. The microneedle electrochemical biosensor of claim 1, wherein, The thickness of the signal conversion layer is 0.1 μm to 40 μm, and the pore size is 0.1 to 100 nm. The thickness of the ion recognition film layer is 0.1 μm to 40 μm, and the pore size is 0.1 to 100 nm.

11. The microneedle electrochemical biosensor of claim 1, wherein, The microneedle electrochemical biosensor also integrates a processor chip.

12. The microneedle electrochemical biosensor of claim 1, wherein, The microneedle electrochemical biosensor also integrates a wireless sensor module, which uses Bluetooth, WiFi, or Zigbee wireless communication technology.

13. A method of fabricating a microneedle electrochemical biosensor, comprising: Include: Prussian blue electrodeposition solution was coated onto the surface of the first graphite layer to obtain a Prussian blue layer; a glucose oxidase solution was coated onto the surface of the Prussian blue layer and bound together by covalent bonds to obtain an enzyme immobilization layer, thereby obtaining the glucose sensing working electrode. The mixture obtained by adding cobalt phthalocyanine, an electronic medium, to β-lactamase and then adding it to an acetate buffer solution and mixing it with mineral oil is coated onto the surface of the second graphite layer to form a β-lactamase immobilization layer, thereby obtaining the sodium penicillin sensing working electrode. A mixture of 3,4-ethylenedioxythiophene and sodium polystyrene sulfonic acid is electrodeposited on the surface of a third graphite layer to form a signal conversion layer. A selective membrane precursor solution is then coated on the surface of the signal conversion layer to form an ion recognition membrane layer, thereby obtaining the sodium chloride sensing working electrode.

14. The method according to claim 13, characterized in that, The selective membrane precursor solution comprises sodium tetrakis(3,5-bis(trifluoromethyl)phenyl)borate, high molecular weight polyvinyl chloride, di(2-ethylhexyl) sebacate, and tetraethyl 4-tert-butylcalixanetetraacetate.

15. The method according to claim 13, characterized in that, K3Fe(CN)6, FeCl3 and KCl were dissolved in an acidic HCl solution and mixed thoroughly to obtain the Prussian blue electrodeposition solution. Bovine serum albumin and glucose oxidase were each dissolved in a phosphate buffer solution to obtain the glucose oxidase solution. Glutaraldehyde was used as a cross-linking agent to react with glucose oxidase to form covalent bonds, thereby immobilizing glucose oxidase and bovine serum albumin on the Prussian blue layer to form a three-dimensional network structure.

16. The method of claim 13, wherein, Also includes: Graphite powder or a mixture of graphite powder and N-cyclohexyl-N-(2-morpholinoethyl)carbodiimide is coated onto the ends of the conductive traces of the glucose sensing electrode, the sodium penicillin sensing electrode, and the sodium chloride sensing electrode, respectively, and then screen-printed to obtain the first graphite layer, the second graphite layer, and the third graphite layer.

17. The method of claim 13, wherein, Also includes: The reference electrode is obtained by screen printing Ag / AgCl ink at the end of the conductive trace of the reference electrode. The counter electrode is obtained by screen printing platinum wire ink at the end of the conductive trace of the counter electrode.

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