Sensor electrode structure and gas sensor using the same

Abstract

To provide a sensor electrode structure that has high sensitivity and can perform reliable gas detection at low cost, and a gas sensor using the same.SOLUTION: A sensor electrode structure is to detect target gas, and the sensor electrode structure comprises: a substrate layer formed of hydrophilic nonwoven fabric; an electrode layer composed of a plurality of electrodes carried on the substrate layer and formed of carbon nanotube; and a reaction layer formed of an enzyme and carried on one electrode of the plurality of electrodes, in a state where a lamination state of the three layers is maintained. The target gas is transferred based on the hydrophilicity of the substrate layer. Electrons generated by a reaction of the target gas and the enzyme are transferred from the reaction layer to the electrode layer. The sensor electrode structure detects the target gas based on the electrons.SELECTED DRAWING: Figure 1

Description

[Technical Field]

[0001] The present invention relates to a sensor electrode structure for detecting various gases, and more particularly to a sensor electrode structure that enables highly accurate gas detection, and a gas sensor using the same. [Background technology]

[0002] Gas sensors that detect gases have a wide range of applications. In particular, it is becoming clear that the components and concentrations of biogases (transcutaneous gases) are closely related to people's illnesses and health conditions.

[0003] For example, a gas sensor that can detect transcutaneous gases emitted from the human skin can measure the blood alcohol concentration by detecting alcohol as a transcutaneous gas, which is highly versatile and convenient, and can also help prevent drunk driving.

[0004] The detection results of such gas sensors can have a significant impact on the lifestyle and behavior of test subjects. Therefore, highly sensitive and reliable gas sensors are required. In particular, improvements and development of the sensor electrode structure, which plays a central role in gas sensors, has been actively pursued.

[0005] Currently, biogas measurement principles under development can be broadly categorized into semiconductor, fluorometric, quartz crystal microbalance, and surface plasmon types. Of these, the semiconductor and fluorometric types are the most practical. However, semiconductor types have difficulty with gas detection specificity. Fluorometric types are solution-based, which makes them difficult to wear and portable, making it difficult to realize mobile sensors.

[0006] In this regard, biogas measurement using enzyme reactions is highly sensitive because it utilizes electrochemical detection based on enzyme reactions, allowing it to detect only specific gases based on the substrate properties of the enzyme and converting the enzyme reaction directly into an electrical signal.

[0007] A known example of a conventional sensor electrode structure that enables biogas measurement using such an enzyme reaction is an enzyme-functional electrode that includes an electrode and an enzyme that oxidizes a specific substrate, in which electrons are transferred from the substrate to the electrode via the enzyme, where the enzyme is membrane-bound alcohol dehydrogenase type III that uses PQQ as a coenzyme, and is composed of subunits I and III but does not contain subunit II containing cytochrome C (see Patent Document 1).

[0008] Furthermore, as a conventional sensor electrode structure, for example, although it is in a technical field different from the sensor field, an electrode for an enzyme-catalyzed fuel cell is known, which is constructed by adding an enzyme to a sheet made by kneading and integrating carbon nanotubes and cellulose nanofibers (see Non-Patent Document 1). [Prior art documents] [Patent documents]

[0009] [Patent Document 1] Japanese Patent Application Laid-Open No. 2008-237099 [Non-patent literature]

[0010] [Non-Patent Document 1] The 30th Annual Meeting of the MRS Japan (Web) (Invited Lecture) 2020 / 12 / 9-11 Summary of the Invention [Problem to be solved by the invention]

[0011] However, in conventional sensor electrode structures, there are enzyme-functional electrodes in which electrons are transferred from a substrate to an electrode via an enzyme, as in Patent Document 1, but because the target gas must first be sufficiently dissolved in an electrolyte, it is not possible to capture and instantly detect sequentially generated gases, such as transdermal gases, in real time.

[0012] Furthermore, even if an electrode constructed by adding an enzyme to a sheet made by kneading and integrating carbon nanotubes and cellulose nanofibers, as in Non-Patent Document 1, were to be used as a gas sensor, it would be necessary to first dissolve the target gas in the electrolyte, and therefore it would not be possible to capture and instantly detect sequentially generated gases, such as transdermal gases, in real time.

[0013] In particular, biogas components and concentrations vary greatly from person to person, just like blood pressure and body temperature. The greatest effectiveness of such a device is achieved by its ease of use on subjects on a daily basis. On the other hand, when enzymes are used in the sensor electrode structure, they cannot withstand long-term use and therefore require daily replacement. For this reason, conventional sensor electrode structures have not achieved the cost reduction that is crucial for practical application.

[0014] The present invention has been made to solve the above-mentioned problems, and aims to provide a sensor electrode structure that enables highly sensitive and reliable gas detection at low cost, and a gas sensor using the same. [Means for solving the problem]

[0015] The sensor electrode structure disclosed in the present application is a sensor electrode structure for detecting a target gas, and comprises a substrate layer made of a hydrophilic nonwoven fabric, an electrode layer consisting of a plurality of electrodes made of carbon nanotubes supported on the substrate layer, and a reaction layer made of an enzyme and supported on one of the plurality of electrodes, all of which are maintained in a three-layer stacked state.The target gas is transmitted based on the hydrophilicity of the substrate layer, and electrons generated by the reaction between the target gas and the enzyme are transmitted from the reaction layer to the electrode layer, and the target gas is detected based on the electrons.

[0016] In this way, the device comprises a substrate layer made of hydrophilic nonwoven fabric, an electrode layer consisting of a plurality of electrodes made of carbon nanotubes supported on the substrate layer, and a reaction layer made of an enzyme supported on one of the plurality of electrodes, all of which are maintained in a three-layer stacked state. The target gas is transmitted based on the hydrophilicity of the substrate layer, electrons generated by the reaction between the target gas and the enzyme are transmitted from the reaction layer to the electrode layer, and the target gas is detected based on the electrons. With this three-layer structure, the target gas is transmitted based on the hydrophilicity of the substrate layer, electrons generated by the reaction between the target gas and the enzyme are transmitted from the reaction layer to the electrode layer, and the target gas is detected based on the electrons. Gas generated at the bottom (for example, transdermal gas generated from the skin of the human body) is transmitted upward due to the hydrophilicity of the substrate layer, passes through the electrode layer, and is absorbed into the reaction layer. Water constantly evaporates from the electrode layer and the reaction layer, making gas diffusion smoother, allowing the direction of gas diffusion to be controlled in one direction, and achieving quick responsiveness. Furthermore, because the substrate layer and the electrode layer consisting of multiple electrodes are integrated, the manufacturing process is extremely simple, realizing low costs.Furthermore, by simply selecting the appropriate enzyme corresponding to the target gas, the same structure and materials for the substrate layer and electrode layer can be used for any target gas, allowing for the construction of a more versatile sensor at low cost.

[0017] In the sensor electrode structure disclosed herein, the hydrophilic nonwoven fabric may be cellulose nanofiber, if necessary. Because the hydrophilic nonwoven fabric is cellulose nanofiber, the cellulose nanofibers in the substrate layer and the carbon nanotubes in the electrode layer both have nano-sized voids, which allow these voids to be spatially continuous and smoothly connected. The three-layer structure allows gas generated below (e.g., transdermal gas generated from human skin) to be absorbed by the reaction layer, and moisture optimally evaporates from the porous electrode layer and reaction layer. This smooth gas diffusion allows the gas diffusion to be controlled in one direction, resulting in rapid responsiveness. Furthermore, because the sensor electrode structure is made of biodegradable cellulose nanofiber and nanocarbon materials, it can be used as a disposable device, achieving both high convenience and reduced environmental impact.

[0018] In the sensor electrode structure disclosed herein, the enzyme is, as necessary, alcohol dehydrogenase (ADH) using pyrroloquinoline quinone (PQQ) as a coenzyme or alcohol dehydrogenase (GDH) using flavin adenine dinucleotide (FAD) as a coenzyme. Since the enzyme is alcohol dehydrogenase (ADH) using pyrroloquinoline quinone (PQQ) as a coenzyme or alcohol dehydrogenase (GDH) using flavin adenine dinucleotide (FAD) as a coenzyme, the alcohol dehydrogenase exhibits substrate specificity, which allows easy detection of alcohol vaporized from a target gas, thereby realizing quick response in alcohol detection, particularly when the target gas is a transdermal gas.

[0019] In the sensor electrode structure disclosed in the present application, the electrode spacing between the electrode layers is 0.01 to 10 mm as needed. Since the electrode spacing between the electrode layers is 0.01 to 10 mm, inter-electrode leakage caused by an inter-electrode distance being too close is suppressed, and an increase in inter-electrode resistance caused by an inter-electrode distance being too far is suppressed, thereby enabling sensor detection with an optimal inter-electrode distance, thereby achieving quick responsiveness and improved sensor accuracy.

[0020] In the sensor electrode structure disclosed in the present application, if necessary, the concentration of the carbon nanotubes per unit area of ​​the substrate layer is 0.01 to 20 μg / cm 2 In this way, the concentration of the carbon nanotubes per unit area of ​​the substrate layer is 0.01 to 20 μg / cm 2 Therefore, the occurrence of electrical leakage between electrodes caused by an excessively high concentration of carbon nanotubes is suppressed, and the decrease in sensor sensitivity caused by an excessively low concentration of carbon nanotubes is suppressed, thereby enabling sensor detection under optimal electrode conditions, thereby realizing quick response and improved sensor accuracy. Furthermore, the flow rate of the target gas transmitted upward from the substrate layer by gas diffusion and the flow rate of the target gas transmitted to the electrode layer are optimally adjusted in the electrode layer by the optimal concentration of carbon nanotubes, thereby further optimizing sensor accuracy.

[0021] The sensor electrode structure disclosed in the present application may be configured so that the concentration of the enzyme is 0.01 to 100 pg / cm. 2 Thus, the concentration of the enzyme is 0.01 to 100 pg / cm 2 As a result, it is possible to suppress a delay in reaction caused by an excessively high concentration of the enzyme, and to suppress a decrease in sensor sensitivity caused by an excessively low concentration of the enzyme. Furthermore, the surface of one of the electrodes in the electrode layer is optimally covered with a single layer of enzyme molecules, and the enzyme is distributed over the entire surface of one of the electrodes in the electrode layer, allowing the enzyme to optimally react with the target gas transported upward from the electrode layer by gas diffusion, thereby optimizing sensor accuracy and reducing manufacturing and maintenance costs.

[0022] The gas sensor disclosed in the present application is configured with the above-described sensor electrode structure. As a result of being configured with the above-described sensor electrode structure, the enzyme reacts optimally to the target gas smoothly transmitted upward from the electrode layer by gas diffusion, optimizing the sensor accuracy. Optimizing the enzyme reduces manufacturing and maintenance costs and achieves quick response.

[0023] The gas sensor disclosed in the present application may, as necessary, comprise an upper surface layer made of a thin synthetic resin film and disposed on the upper surface side of the electrode, and a lower surface layer made of a thin synthetic resin film and disposed on the lower surface side of the substrate, the lower surface layer having an opening. Thus, since the gas sensor comprises an upper surface layer made of a thin synthetic resin film and disposed on the upper surface side of the electrode, and a lower surface layer made of a thin synthetic resin film and disposed on the lower surface side of the substrate, the opening forms a minute space at the position where the gas sensor and the skin come into contact when, for example, human skin is placed against the lower surface layer, and the gas diffusion is guided in one upward direction without dispersing the diffusion of percutaneous gas generated from the skin in multiple directions and destabilizing the detected concentration, thereby achieving more responsiveness and highly accurate gas detection. [Brief explanation of the drawings]

[0024] [Figure 1] 1 shows a configuration diagram of a sensor electrode according to a first embodiment of the present invention. [Figure 2] 1A and 1B are explanatory diagrams of an enzyme reaction in a sensor electrode according to a first embodiment of the present invention. [Figure 3] 1A and 1B are schematic diagrams illustrating the mechanism of a sensor electrode according to a first embodiment of the present invention. [Figure 4] FIG. 4 shows a configuration diagram of a gas sensor according to a second embodiment of the present invention. [Figure 5] 1A and 1B are a perspective view and a cross-sectional view of a gas sensor according to a second embodiment of the present invention. [Figure 6] 5A to 5C are explanatory views showing a method for manufacturing a gas sensor according to a second embodiment of the present invention. [Figure 7]1 shows SEM images of the surface (a) and cross section (b) of a CNF sheet constituting a sensor electrode according to Example 1 of the present invention. [Figure 8] 6 shows the results of CV measurement at each potential sweep rate of the sensor electrode according to Example 1 of the present invention. [Figure 9] 1 shows a plot of oxidation peak current versus the square root of each sweep rate for a sensor electrode according to Example 1 of the present invention. [Figure 10] 10 shows the CV results of ethanol detection using a sensor electrode according to Example 2 of the present invention. [Figure 11] 10 shows the CV results of ethanol detection for various concentrations of ethanol using a sensor electrode according to Example 2 of the present invention. [Figure 12] 10 shows current values ​​corresponding to ethanol detection for various concentrations of ethanol in the sensor electrode according to Example 2 of the present invention. [Figure 13] 10 shows the sensor measurement results of a subject who had drunk alcohol using a gas sensor according to Example 3 of the present invention. [Figure 14] 10 shows the sensor measurement results of a subject who had drunk alcohol using a gas sensor according to Example 3 of the present invention. [Figure 15] 10 shows the measurement results of vibration of the gas sensor according to Example 4 of the present invention. [Figure 16] 10 shows the measurement results of vibration of the gas sensor according to Example 4 of the present invention. [Figure 17] 10 shows measurement results (enlarged view) of vibrations of the gas sensor according to Example 4 of the present invention. DETAILED DESCRIPTION OF THE INVENTION

[0025] (First embodiment) As shown in FIG. 1, the sensor electrode structure according to the first embodiment of the present application is a sensor electrode structure for detecting a target gas, and comprises a substrate layer 1 made of a hydrophilic nonwoven fabric, an electrode layer 2 composed of a plurality of electrodes made of carbon nanotubes supported on the substrate layer 1, and a reaction layer 3 made of an enzyme and supported on one of the plurality of electrodes, all of which are maintained in a stacked three-layer state. The target gas is transmitted based on the hydrophilicity of the substrate layer 1, and electrons generated by the reaction between the target gas and the enzyme are transmitted from the reaction layer 3 to the electrode layer 2, and the target gas is detected based on these electrons.

[0026] Target gases include transdermal gases. By selecting the right enzyme, it is possible to detect alcohol gas, an example of a transdermal gas. This makes it possible to check for and prevent drunk driving, for example. In addition, by selecting the right enzyme, it is possible to construct an extremely simple gas sensor that can detect various types of gases while using the same electrode chip structure and materials.

[0027] The hydrophilic nonwoven fabric constituting the substrate layer 1 is not particularly limited as long as it is a hydrophilic nonwoven fabric, but it is preferable to use cellulosic fibers. Examples of such cellulosic fibers include naturally occurring cellulosic fibers such as pulp, hemp, and cotton, and more preferably, cellulose nanofibers (CNF).

[0028] Cellulose nanofiber (CNF) is a biomass material made by breaking down plant fibers to nano-size, and is characterized by low environmental impact, light weight, and flexibility. CNF is not electrically conductive, but it is hygroscopic, flexible, and has a large specific surface area.

[0029] Other synthetic fibers may also be used, such as polyethylene, polypropylene, polyethylene terephthalate, polytrimethylene terephthalate, vinylon, etc. Hydrophilically treated synthetic fibers may also be used.

[0030] The carbon nanotubes that make up the electrodes of the electrode layer 2 are tube-like cylindrical substances that are composed only of carbon elements and have a diameter on the order of nanometers. There are no particular limitations on the type of carbon nanotubes, and examples include well-known carbon nanotubes such as single-walled carbon nanotubes (SWCNTs) and multi-walled carbon nanotubes (MWCNTs). However, multi-walled carbon nanotubes (MWCNTs) are more suitable in terms of their high conductivity and large surface area.

[0031] The concentration of the carbon nanotubes is not particularly limited, but is preferably 0.01 to 20 μg / cm 3 per unit area of ​​the substrate layer 1. 2 For example, 2 μg / cm 2 This suppresses electrical leakage between electrodes caused by an excessively high carbon nanotube concentration, and also suppresses a decrease in sensor sensitivity caused by an excessively low carbon nanotube concentration, thereby enabling sensor detection under optimal electrode conditions and achieving quick response and improved sensor accuracy. Furthermore, the flow rate of the target gas transmitted upward from the substrate layer 1 by gas diffusion and the flow rate of the target gas transmitted to the electrode layer 2 are optimally adjusted in the electrode layer 2 by the optimal carbon nanotube concentration, thereby further optimizing sensor accuracy.

[0032] At a minimum, the multiple electrodes of this electrode layer 2 can be composed of two electrodes, a first electrode 21 containing an enzyme and a second electrode 22 not containing an enzyme, which serve as an anode and a cathode, as shown in FIG. 1(a). Alternatively, as shown in FIG. 1(b), it can be composed of three electrodes, a first electrode 21 containing an enzyme, and a second electrode 22 and a third electrode 23 not containing an enzyme. The material added to the second electrode 22 and the third electrode 23 not containing an enzyme is not particularly limited, but it is preferable to add polybenzimidazole (PBI) to them. In this way, multiple electrodes can be produced integrally on this substrate layer 1, which significantly simplifies the manufacturing process and reduces manufacturing costs.

[0033] Furthermore, when used as various sensors, it is possible to combine the functions of two of the working electrode (WE), counter electrode (CE), and reference electrode (RE) into a single electrode, thereby forming a sensor from a total of two electrodes.

[0034] When used as various sensors, the electrode chip is made disposable by using biodegradable CNF as the substrate layer 1 and nanocarbon material as the electrode layer 2, thereby increasing convenience.

[0035] CNFs are biodegradable, flexible, environmentally friendly, disposable, and small in size, making them suitable for future wearable sensors. Skin gases emitted by the human body are a mixture of various organic substances, such as ethanol, acetone, methane, and acetaldehyde, and are present at very low concentrations (ppb level). The gas sensor according to this embodiment can detect ethanol at an appropriate range of 75 to 112 ppb, which is required for individuals with drinking problems (see the examples below). Real-time ethanol skin gas measurements are necessary to verify sensor performance in daily life. The gas sensor according to this embodiment, with its simple configuration, can enable real-time remote medical monitoring, individual health status assessment, and feedback from remote facilities.

[0036] The electrode spacing of this electrode layer 2 is not particularly limited, but is preferably 0.01 to 10 mm, and can be, for example, 2 to 3 mm. This suppresses inter-electrode leakage caused by an inter-electrode distance being too close, and also suppresses an increase in inter-electrode resistance caused by an inter-electrode distance being too far, thereby enabling sensor detection with an optimal inter-electrode distance, thereby achieving quick responsiveness and improved sensor accuracy.

[0037] The enzyme constituting the electrode of the reaction layer 3 is not particularly limited, but it is preferable to use, for example, alcohol dehydrogenase (ADH) using pyrroloquinoline quinone (PQQ) as a coenzyme or alcohol dehydrogenase (GDH) using flavin adenine dinucleotide (FAD) as a coenzyme. As a result, as shown in Figure 2, alcohol dehydrogenase exhibits substrate specificity, so that alcohol vaporized from the target gas can be easily detected based on direct electron transfer by the enzymatic reaction, and quick response is achieved in alcohol detection, particularly when the target gas is a transdermal gas.

[0038] Because it uses enzymes, it cannot withstand long-term use and must be replaced daily. For this reason, cost reduction is extremely important for practical application. In this regard, it has been confirmed that practical application is possible by fabricating a device that is separated into 1. a disposable electrode chip portion (this embodiment) as a consumable item, and 2. a reusable electrode chip connection portion and a signal detection portion (see the examples below).

[0039] The concentration of this enzyme is not particularly limited, but is preferably 0.01 to 100 pg / cm 2 For example, it is preferable that the concentration is 0.1 pg / cm 2 This suppresses a delay in reaction caused by an excessively high concentration of the enzyme, and also suppresses a decrease in sensor sensitivity caused by an excessively low concentration of the enzyme, and further, the surface of one of the electrodes of the electrode layer 2 is optimally covered with a single layer of enzyme molecules, and the enzyme spreads over the entire surface of one of the electrodes of the electrode layer 2, allowing the enzyme to optimally react with the target gas transported upward from the electrode layer 2 by gas diffusion, thereby optimizing sensor accuracy and reducing manufacturing and maintenance costs.

[0040] Furthermore, the state in which the three layers are maintained in a stacked state means that on one electrode carrying an enzyme, the three layers of the substrate layer 1, electrode layer 2, and reaction layer 3 are not formed as a single, mixed layer, but are maintained in a state in which they are individually separated and stacked.

[0041] The target gas is transmitted due to the hydrophilicity of the substrate layer 1, and electrons generated by the reaction between the target gas and the enzyme are transmitted from the reaction layer 3 to the electrode layer 2, and the target gas is detected based on these electrons.

[0042] It is more preferable that the pores of cellulose nanofibers (CNF) are larger than those of carbon nanotubes, which allows the transdermal gas to pass upward more easily and allows the transdermal gas to react stably at the sensor electrode in the carbon nanotubes, thereby achieving high sensor accuracy.

[0043] The mechanism behind this excellent effect is thought to be as follows: First, carbon nanotubes generally have a size of 1 to 10 nm, and cellulose nanofibers (CNFs) have a size of 10 to 15 nm when they form bundles with each fiber being approximately 2 nm.

[0044] Regarding the voids of each material, cellulose nanofibers (CNFs) generally have voids of 50–100 nm, depending on their density, as shown in the SEM image in Figure 7 (described later). Because carbon nanotubes have a high degree of cohesion, their voids are smaller than those of cellulose nanofibers (CNFs). As shown in Figure 3(a), it is presumed that carbon nanotubes form a large number of clumps (2a) of several hundred microns in size. The voids (L2) within these clumps are estimated to be several nm to 20 nm, while the voids (L1) between the clumps are estimated to be even larger, at several hundred microns. As shown in Figure 3(b), it is presumed that the target gas (B) (e.g., ethanol) diffuses throughout the carbon nanotube layer through the large voids (L1) between the clumps and then enters the small voids (L2) within the clumps, thereby promoting the reaction with the enzyme in the reaction layer (3).

[0045] In the sensor electrode structure according to this embodiment, the disposable electrode chip can be made of, for example, cellulose nanofiber, nanocarbon, and enzymes, making it biodegradable and environmentally friendly. Furthermore, this electrode chip can be manufactured roll-to-roll, facilitating cost reduction through mass production and enabling it to be priced at a level that subjects can use on a daily basis (assuming 100 yen per chip, similar to a mobile blood glucose sensor). Thus, the sensor electrode structure according to this embodiment makes it possible to develop a biogas detection device that offers the best balance between cost and detection sensitivity.

[0046] The sensor electrode structure according to the first embodiment of the present invention can be used in a wide range of applications, for example, as a gas sensor capable of detecting various gases.

[0047] (Second embodiment) The gas sensor according to the second embodiment of the present application is configured with the sensor electrode structure according to the first embodiment. By configuring this gas sensor with this sensor electrode structure, the enzyme reacts optimally to the target gas smoothly transmitted upward from the electrode layer 2 by gas diffusion, thereby optimizing the sensor accuracy, and this optimization of the enzyme reduces manufacturing and maintenance costs and achieves quick responsiveness.

[0048] The gas sensor according to this embodiment can be configured by adding other members to the sensor electrode structure. For example, as shown in Fig. 4, it can be configured to include an upper surface layer 41 made of a thin synthetic resin film and disposed on the upper surface of the electrode, and a lower surface layer 42 made of a thin synthetic resin film and disposed on the lower surface of the substrate, with an opening.

[0049] The multiple electrodes of the electrode layer 2 can consist of a minimum of two electrodes, an anode and a cathode, as shown in Figure 4(a). Furthermore, the multiple electrodes of the electrode layer 2 may also consist of three electrodes. For example, as shown in Figure 4(b), in the case of three electrodes, an integrated three-layer electrode is constructed in which a working electrode (WE) as a first electrode 21, which has a three-layer structure consisting of a substrate layer 1 (e.g., CNF), an electrode layer 2 (carbon nanotubes), and an enzyme, a counter electrode (CE) as a second electrode 22 that does not contain an enzyme, and a reference electrode (RE) as a third electrode 23 are arranged on the same CNF substrate. This can be used as a variety of sensors. The materials added to the enzyme-free counter electrode (CE) and reference electrode (RE) are not particularly limited, but polybenzimidazole (PBI) is preferably added. Because multiple electrodes can be integrated onto the substrate layer 1 in this way, the manufacturing process is extremely simple, reducing manufacturing costs.

[0050] The thin-film synthetic resin is not particularly limited, but for example, a transparent plastic can be used. The size of the opening is not particularly limited, but in a transcutaneous gas sensor application, it can be a rectangular shape with a length and width of about 1 mm so that it can contact a human arm. This leaves a minute space at the site that comes into contact with the skin of the subject 100.

[0051] 5(a), the gas sensor according to this embodiment can connect and fix each electrode to an electrode support part 300 made of a carbon material. The electrode support part 300 enables easy current detection and can be configured in a shape that is particularly easy to wear on a human arm.

[0052] As described above, the gas sensor of this embodiment comprises an upper surface layer 41 made of a thin-film synthetic resin and disposed on the upper surface of the electrode, and a lower surface layer 42 made of a thin-film synthetic resin and disposed on the lower surface of the substrate, and having an opening. Therefore, when human skin is placed against the lower surface layer 42, as shown in Figure 5(b), this opening A forms a minute space at the position where the gas sensor and the skin of the subject 100 come into contact, and the diffusion of the transdermal gas B generated from the skin is guided in one upward direction without dispersing it in multiple directions and making the detected concentration unstable, thereby achieving even more responsiveness and highly accurate gas detection.

[0053] An example of a method for manufacturing the gas sensor according to this embodiment will be described below with reference to FIG.

[0054] First, as shown in Figure 6(a), cellulose nanofibers (CNF) are dispersed in water by ultrasonic agitation or other means. This CNF dispersion is spread on a Teflon® plate 200 and heated until dry to obtain a CNF sheet. Next, as shown in Figure 6(b), a multi-walled carbon nanotube (MWCNT) solution that will form the electrode layer 2 is dropped and dispersed on the surface of the CNF sheet that will become the substrate layer 1. Next, an enzyme that will become the reaction layer 3 is dropwise coated on the surface of the first electrode 21, modifying the first electrode 21 with the enzyme. Next, as shown in Figure 6(c), the CNF sheet electrode composed of the enzyme-modified MWCNT is sandwiched between the upper and lower layers 41 and 42. Each electrode is connected to an electrode support 300 made of a carbon material, allowing for easy current detection.

[0055] It has been confirmed that the gas sensor according to this embodiment obtained in this way is capable of measuring transcutaneous gases (skin gases) at low levels and can certainly achieve the cost required for daily use by subjects (assuming 100 yen per sensor, like mobile blood glucose level sensors) (see Examples below). In other words, it is expected that commercialization will develop into a global market of 1 trillion yen per year in the future, which is equal to or greater than that of blood glucose level sensors, and this will have an extremely large impact.

[0056] The promotion of independent and decentralized health management, which enables individuals to manage their own health and remotely manage information with medical institutions, is a technology that society needs right now. The gas sensor according to this embodiment can make a significant contribution to realizing a resilient and sustainable post-COVID society through medical applications as well as agricultural applications and remote measurement and communication of a wide variety of gases. As the global population becomes increasingly super-aged, the social burden of medical care and nursing care will accelerate. The gas sensor according to this embodiment will realize independent and decentralized health management, which allows medical institutions to remotely manage information on the health status of individuals, and there is a high demand for it on a global scale.

[0057] The present invention will be described in more detail below by showing examples, but the present invention is not limited to the following examples.

[0058] Example 1 - Fabrication of gas sensors The following materials were used: <Material> MWCNT: Multi-walled carbon nanotubes (manufactured by Aldrich) CNF (manufactured by Chuetsu Pulp Co., Ltd.) FQQ-ADH (Provided by Shirai Laboratory, Faculty of Agriculture, Kyoto University)

[0059] (1) Fabrication of MWCNT-modified CNF sheet electrodes CNF sheets were prepared by dispersing 1 mg of 2.35 wt% CNF in 20 mL of water obtained from an ultrapure water system (Milli-Q) using an ultrasonic homogenizer for 20 minutes. Next, 10 mL of this CNF dispersion was spread on a Teflon® plate (4 cm x 10 cm) and heated on a hot plate at 140 °C until dry. The electrical conductivity of the CNF sheet was measured by dropping a dispersed MWCNT solution onto its surface. Next, the working electrode was modified by dropwise coating the PQQ-ADH enzyme onto the electrode surface.

[0060] (2) Physical properties of MWCNT-modified CNF sheet electrodes SEM images of the surface and cross-section of a CNF sheet before modification with MWCNTs are shown in Figure 7(a). The surface area morphology was relatively flat, suitable for electrode modification. The cross-sectional area morphology shown in Figure 7(a) showed a layered structure with nanometer-scale thickness. This layered structure could potentially increase the horizontal conductivity by further modifying the CNF sheet with MWCNTs.

[0061] (3) Experimental conditions The dispersion was homogenized using an ultrasonic probe model BRANSON 5520 (Kanagawa, Japan). All potentials were confirmed against Ag | AgCl | saturated KCl at 25 °C (+199 mV vs. a normal hydrogen electrode). Electrochemical measurements were performed using an electrochemical analyzer (Model 660A, ALS Co., Ltd., Tokyo, Japan). pH was recorded using a pH meter (AUT-501, DKK-TOA Corp., Tokyo, Japan). Cyclic voltammetry (CV) measurements were performed using a three-electrode system included in the CNF electrode. All electrochemical measurements were performed under atmospheric conditions.

[0062] (4) Electrochemical performance of MWCNT-modified CNF sheet electrode To evaluate the electrochemical performance of the MWCNT-modified CNF sheet electrode, the CNF sheet electrode was treated with a pH 7 KCl solution (1 mol / dm 3 ) in Fe(CN)6 3- (1mmol / dm 3 ) and the results of CV measurements at each potential sweep rate are shown in Figure 8.

[0063] From the results obtained, the oxidation and reduction peak potentials of ferrocyanide were 100 mVs -1 The potentials were observed around +0.46 V and +0.08 V, respectively, at a potential sweep rate of 100 mV. The separation between the anodic and cathodic peak potentials was approximately 380 mV, indicating a slow redox reaction. A separation of the two peak potentials equal to approximately 60 mV per electron indicates a fast electron transfer rate. Despite the slow reaction, the CNF conductive sheet was able to transfer electrons and function satisfactorily as an electrode.

[0064] A plot of the oxidation peak current against the square root of each sweep rate is shown in Figure 9. The results confirm that the observed redox reaction was a diffusion-limited process, as the plot of cathodic peak current (ipc) versus the square root of the potential sweep rate is linear.

[0065] Thus, the CNF sheet fabricated in this study was confirmed to be extremely thin and lightweight, with an average thickness of approximately 60 μm. Given CNF's affinity with human skin, the fabricated CNF sheet is a promising framework for electrochemical sensors. Observation of redox reactions at the MWCNT-modified CNF sheet electrode using ferrocyanide demonstrated that the modified CNF sheet possesses sufficient electrochemical performance as an electrode. Furthermore, the PQQ-ADH-modified CNF sheet electrode was able to detect ethanol evaporated from a 0.1 M ethanol solution under ambient conditions. Oxidation catalytic current was detected at -0.1 V vs Ag | AgCl | saturated KCl, based on direct electron transfer between the enzyme and the electrode surface. Therefore, the CNF sheet electrode is a promising candidate for future wearable sensors.

[0066] Example 2 - Ethanol gas detection To detect ethanol, the modified CNF sheet electrode modified with PQQ-ADH enzyme prepared above was used as the working electrode to construct the gas sensor shown in Figure 5 shown in the second embodiment. The target gas was generated by evaporating 100 μL of 0.1 M ethanol solution to simulate transdermal gas.

[0067] Figure 10 shows the CV results of ethanol detection at the PQQ-ADH enzyme / MWCNT-modified CNF sheet electrode in 0.1 M acetate buffer (pH 5.5) containing 2 mM CaCl2 under atmospheric conditions in the presence (solid line) and absence (dashed line) of 0.1 M ethanol (scan rate: 10 mV s -1 The estimated ethanol gas concentration generated from the 0.1M ethanol solution was 344 ppm.

[0068] The results showed that no catalytic oxidation current was detected at the MWCNT-modified CNF sheet electrode in the absence of alcohol. Next, a catalytic oxidation current was detected from -0.1 V (vs. Ag | AgCl | saturated KCl) in the presence of ethanol. This result confirmed that the enzyme physically adsorbed on the working electrode was active and could indeed transfer electrons when ethanol was oxidized.

[0069] Furthermore, the CV results of ethanol detection for various concentrations of ethanol are shown in Figure 11 (scan rate: 10 mV s -1 ) Figure 12 shows the current values ​​corresponding to the detection of ethanol for each concentration of ethanol. These results demonstrate that sensitive sensing according to the ethanol concentration is possible.

[0070] The backside of the CNF sheet directly faced the target component, allowing direct interaction between the hydrophilic components of the CNF and ethanol, resulting in the adsorption of the target component and reaction with the electrode surface. The stacked morphology of the CNF sheet shown in Figure 7(b) above also increased the amount of target component in the target gas adsorbed onto its surface. Ethanol was then oxidized by the enzyme via direct electron transfer between the enzyme and the MWCNT electrode surface. Overall, the CNF electrode of this example was confirmed to be capable of detecting ethanol.

[0071] Example 3 -Transcutaneous gas detection Using the gas sensor fabricated in Example 1, adult subjects were actually asked to drink beer (5%) and transdermal gas detection was performed. The results are shown in Figure 13. These results confirmed the surprising sensitivity of the sensor, with the sensor responding just a few minutes after the subject began drinking beer (5%).

[0072] Furthermore, we actually had adult subjects drink not only beer (5%), but also shochu (12%) and shochu (43%), and performed transdermal gas detection. The results are shown in Figure 14. These results confirmed the incredibly accurate sensor sensitivity, with the sensor responding in order according to the alcohol concentration of the subjects' drinks.

[0073] Example 4 -Other measurements The gas sensor fabricated in Example 1 was examined for the influence of vibration. As shown in FIG. 15, it was confirmed that vibrations in the arms and legs of the human body were sensed separately. The influence of vibrations due to blood pressure fluctuations was also examined. As shown in FIG. 16 and its enlarged view in FIG. 17, it was confirmed that sensing was performed in response to blood pressure fluctuations. [Explanation of symbols]

[0074] 1. Substrate Layer 2 electrode layer 2a Carbon nanotube clumps 21 First electrode 22 Second electrode 23 Third electrode 3. Reaction layer 41 Top layer 42 Bottom layer 100 subjects 200 Teflon plates 300 Electrode support part

Claims

1. A sensor electrode structure for detecting a target gas, a substrate layer made of a hydrophilic nonwoven fabric; an electrode layer including a plurality of electrodes made of carbon nanotubes supported on the substrate layer; a reaction layer made of an enzyme and supported on one of the plurality of electrodes; The three layers are maintained in a stacked state. The target gas is transferred based on the hydrophilicity of the substrate layer, and electrons generated by the reaction between the target gas and the enzyme are transferred from the reaction layer to the electrode layer, and the target gas is detected based on the electrons. Sensor electrode structure.

2. 2. The sensor electrode structure according to claim 1, The hydrophilic nonwoven fabric is characterized in that it is made of cellulose nanofiber. Sensor electrode structure.

3. 3. The sensor electrode structure according to claim 1, The electrode spacing of the electrode layer is 0.01 to 10 mm. Sensor electrode structure.

4. 3. The sensor electrode structure according to claim 1, The concentration of the carbon nanotubes per unit area of ​​the substrate layer is 0.01 to 20 μg / cm2. Sensor electrode structure.

5. The sensor electrode structure according to claim 1 or 2 is configured. Gas sensor.

6. 6. The gas sensor according to claim 5, an upper surface layer made of a thin film synthetic resin and disposed on an upper surface side of the electrode; a lower surface layer made of a thin-film synthetic resin, disposed on the lower surface side of the substrate layer, and having an opening. Gas sensor.

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