Sensor

Abstract

The purpose of the present invention is to provide a sensor for measuring a biological component concentration with a means for improving the reaction efficiency.　The present invention provides a sensor for measuring a biological component concentration, the sensor comprising: an electrode that includes at least a working electrode; and a reagent layer that is formed on the electrode. The reagent layer includes an oxidoreductase, an electron transfer body having an electron transfer material bonded to a terminal of branched-chain polyethylene glycol, and a single-walled carbon nanotube.

Description

Sensor

[0001] The present invention relates to a sensor used to measure the concentration of a biological constituent.

[0002] Enzyme-based electrochemical sensors are used in clinical chemistry testing to continuously measure bioconstituents in blood, such as glucose, lactate, cholesterol, bilirubin, and amino acids. The concentrations of these bioconstituents in blood are important for diagnosing and monitoring diseases.

[0003] One known method for measuring biological components using enzymes is to use hydrogen peroxide as an electron carrier. In this method, hydrogen peroxide produced by an enzymatic reaction of a biological component is oxidized on the surface of a working electrode by applying a potential, generating an oxidation current. While this method allows for a relatively simple reagent layer to be constructed, oxygen is required for the reaction, and oxygen blocking due to sensor encapsulation can significantly affect the stability of sensor quality over time. Furthermore, the need to apply a relatively high potential to the working electrode can lead to problems such as interference with substances other than the biological component being the analyte in the living body.

[0004] In contrast, a method exists in which an oxidoreductase and an artificial electron mediator are encapsulated in the sensor's reagent layer. In this method, the biological component to be analyte reacts with the oxidoreductase and is oxidized, and the artificial electron mediator accepts the electrons generated by the oxidation. The artificial electron mediator, which accepts the electrons and is reduced, is electrochemically oxidized at an electrode, and the concentration of the biological component in a sample such as blood can be measured from the resulting current value and charge magnitude. This method is advantageous for the sensor's long-term stability because oxygen is not required for the reaction. Furthermore, the reaction proceeds at a relatively low working electrode potential, which has the advantage of being less affected by interfering substances. On the other hand, because artificial electron mediators are generally small molecules, they may permeate through membranes such as a biological component diffusion control membrane, even when placed on the sensor's reagent layer. This makes it difficult to retain the artificial electron mediator in the reagent layer, resulting in insufficient reaction efficiency.

[0005] In response to this, JP-A 2003-514924 discloses a polymer transition metal complex in which a metal complex serving as an artificial electron transfer substance is bound to a polymer. The use of such a polymer transition metal complex makes it easier to retain the artificial electron transfer substance in the reagent layer.

[0006] However, according to the investigations of the present inventors, it has been found that even when the technique described in JP-A-2003-514924 is used, sufficient reaction efficiency may not be obtained in some cases.

[0007] SUMMARY OF THE INVENTION It is therefore an object of the present invention to provide a means for improving the reaction efficiency of a sensor for measuring the concentration of a biological constituent.

[0008] The present inventors have conducted extensive research to solve the above-mentioned problems, and as a result, have found that the above-mentioned problems can be solved by incorporating, into the reagent layer of a sensor for measuring the concentration of a biological component, the sensor comprising an electrode including at least a working electrode and a reagent layer formed on the electrode, an oxidoreductase, an electron carrier having an electron carrier bound to the end of branched polyethylene glycol, and single-walled carbon nanotubes, thereby completing the present invention.

[0009] That is, the above object can be achieved by the following: 1. A sensor for measuring the concentration of a biological component, comprising an electrode including at least a working electrode and a reagent layer formed on the electrode, wherein the reagent layer comprises: an oxidoreductase; an electron carrier in which an electron carrier is bound to an end of a branched-chain polyethylene glycol; and single-walled carbon nanotubes. 2. The sensor according to 1., wherein the branched-chain polyethylene glycol has a weight-average molecular weight of 10,000 to 20,000. 3. The sensor according to 1. or 2., wherein the branched-chain polyethylene glycol is a 4-branched polyethylene glycol. 4. The sensor according to any one of 1. to 3., wherein the electron carrier is a compound represented by the following chemical formula (1):

[0010]

[0011] In chemical formula (1), R １ , R ２ , R３ , and R ４ At least one of the above is a phenazine derivative represented by the following chemical formula (2):

[0012]

[0013] In chemical formula (2), R ５ are each independently an alkyl group having 1 to 4 carbon atoms; １ are each independently -(CH ２ ) ａ —C(O)—(CH ２ ) ｂ -, -(CH ２ ) ａ -C(O)-NH-(CH ２ ) ｂ -, -(CH ２ ) ａ -NH-C(O)-(CH ２ ) ｂ -, -C(O)-(CH ２ ) ａ -C(O)-, -(CH ２ ) ａ -NH-(CH ２ ) ｂ - and -(CH ２ ) ａ -O-(CH ２ ) ｂ wherein each a is independently an integer from 1 to 6, each b is independently an integer from 1 to 6, * is a bonding position, and X － are each independently an anion, and R １ , R ２ , R ３ , and R ４ Among these, those other than the phenazine derivatives each independently represent H, —CH ２ -NH ２ , -(CH ２ ) ２ -NH ２ , and -(CH ２ ) ３ -NH ２and n is independently an integer of 1 to 1,000. 5. The sensor according to 4., wherein the reagent layer further comprises a compound in which one or more selected from the group consisting of nicotinamide adenine dinucleotide (NAD), pyrroloquinoline quinone (PQQ), and derivatives thereof are bound to an end of a branched-chain polyethylene glycol. 6. The sensor according to any one of 1. to 5., wherein 1 to 10 molecules of the electron carrier are bound to one molecule of the branched-chain polyethylene glycol. 7. The sensor according to 1. to 6., wherein the oxidoreductase is glucose dehydrogenase (GDH), glucose oxidase (GOD), glucose-6-phosphate dehydrogenase, cholesterol dehydrogenase, cholesterol oxidase, glycerophosphate dehydrogenase, glycerophosphate oxidase, lactate dehydrogenase (LDH), lactate oxidase, alcohol dehydrogenase, alcohol oxidase, uricase, 3-hydroxybutyrate dehydrogenase, or urate dehydrogenase. 8. The sensor according to any one of 1. to 7., further comprising an enzyme protective film and / or a biological component diffusion control film on the reagent layer. 9. The sensor according to any one of 1. to 8., for measuring the concentration of glucose, 3-hydroxybutyric acid, lactic acid, cholesterol, neutral fat, or uric acid in blood or a body fluid. 10. An electron carrier which is a compound represented by the following chemical formula (1):

[0014]

[0015] In chemical formula (1), R １ , R ２ , R ３ , and R ４ At least one of the above is a phenazine derivative represented by the following chemical formula (2):

[0016]

[0017] In chemical formula (2), R ５ are each independently an alkyl group having 1 to 4 carbon atoms; １ are each independently -(CH ２ ) ａ —C(O)—(CH２ ) ｂ -, -(CH ２ ) ａ -C(O)-NH-(CH ２ ) ｂ -, -(CH ２ ) ａ -NH-C(O)-(CH ２ ) ｂ -, -C(O)-(CH ２ ) ａ -C(O)-, -(CH ２ ) ａ -NH-(CH ２ ) ｂ - and -(CH ２ ) ａ -O-(CH ２ ) ｂ wherein each a is independently an integer from 1 to 6, each b is independently an integer from 1 to 6, * is a bonding position, and X － are each independently an anion, and R １ , R ２ , R ３ , and R ４ Among these, those other than the phenazine derivatives each independently represent H, —CH ２ -NH ２ , -(CH ２ ) ２ -NH ２ , and -(CH ２ ) ３ -NH ２ and n is independently an integer from 1 to 1,000.

[0018] FIG. 1 is an exploded perspective view showing one embodiment of the sensor of the present invention. FIG. 2 is a cross-sectional view of the sensor taken along line a-a in FIG. 1. FIG. 3 is a graph showing the results of electrochemical measurements using sensors prepared in Example 1 and Comparative Example 1. FIG. 3(a) shows the results of measurements of glucose concentration-dependent oxidation current, and FIG. 3(b) shows the results of cyclic voltammetry measurements. FIG. 4 is a graph showing the results of continuous glucose measurements using the sensor prepared in Example 2. FIG. 5 is a graph showing the results of measurements of glucose concentrations on the first, third, and seventh days after implantation of the sensor prepared in Example 2 into an animal. FIG. 6 is a graph showing the results of continuous lactate measurements using the sensor prepared in Example 3. FIG. 7 is a graph showing the results of measurements of 3-hydroxybutyrate concentration-dependent oxidation current using the sensors prepared in Example 4, Comparative Example 4, and Comparative Example 5.

[0019] Hereinafter, embodiments of the present disclosure will be described, but the present disclosure is not limited to the following embodiments.

[0020] In addition, in this specification, the range "X to Y" includes X and Y and means "X or more and Y or less." "M" means mol / L. Unless otherwise specified, operations and measurements of physical properties are performed under conditions of room temperature (20 to 25°C) and a relative humidity of 40 to 50% RH.

[0021] A first aspect of the present disclosure is a sensor for measuring the concentration of a biological component, comprising an electrode including at least a working electrode and a reagent layer formed on the electrode, wherein the reagent layer comprises an oxidoreductase, an electron carrier having an electron carrier bound to the end of branched polyethylene glycol, and single-walled carbon nanotubes.

[0022] A second aspect of the present disclosure is an electron carrier, which is a compound represented by the following chemical formula (1):

[0023]

[0024] In chemical formula (1), R １ , R ２ , R ３ , and R ４At least one of the above is a phenazine derivative represented by the following chemical formula (2):

[0025]

[0026] In chemical formula (2), R ５ are each independently an alkyl group having 1 to 4 carbon atoms; １ are each independently -(CH ２ ) ａ —C(O)—(CH ２ ) ｂ -, -(CH ２ ) ａ -C(O)-NH-(CH ２ ) ｂ -, -(CH ２ ) ａ -NH-C(O)-(CH ２ ) ｂ -, -C(O)-(CH ２ ) ａ -C(O)-, -(CH ２ ) ａ -NH-(CH ２ ) ｂ - and -(CH ２ ) ａ -O-(CH ２ ) ｂ wherein each a is independently an integer from 1 to 6, each b is independently an integer from 1 to 6, * is a bonding position, and X － are each independently an anion, and R １ , R ２ , R ３ , and R ４ Among these, those other than the phenazine derivatives each independently represent H, —CH ２ -NH ２ , -(CH ２ ) ２ -NH ２ , and -(CH ２ ) ３ -NH ２ and n is independently an integer from 1 to 1000.

[0027] According to the present disclosure, it is possible to improve the reaction efficiency in a sensor for measuring the concentration of a biological constituent.

[0028] A first aspect of the present disclosure is a sensor for measuring the concentration of a biological component, the sensor comprising an electrode including at least a working electrode and a reagent layer formed on the electrode, the reagent layer comprising an oxidoreductase, an electron carrier having a branched-chain polyethylene glycol with an electron carrier bound to an end of the branched-chain polyethylene glycol, and single-walled carbon nanotubes. According to the present disclosure, the reaction efficiency of the sensor for measuring the concentration of a biological component can be improved.

[0029] The sensor of this embodiment uses an electron carrier in which an electron carrier is bound to the end of branched polyethylene glycol. By polymerizing the electron carrier in this way, the electron carrier can be easily retained in the reagent layer of the sensor. Branched polyethylene glycol provides more active sites for introducing the electron carrier than linear polyethylene glycol. Furthermore, because branched polyethylene glycol is hydrated, the electron carrier can be polymerized while maintaining its flexibility in aqueous solution. Therefore, the electron carrier can be polymerized without reducing its reaction efficiency. Additionally, adding single-walled carbon nanotubes, which have excellent electronic conductivity, to the reagent layer can improve the efficiency of electron transfer, receiving electrons from the electron carrier and sending them to the electrode. These effects are believed to improve the reaction efficiency of sensors for measuring biological component concentrations.

[0030] Hereinafter, embodiments of the sensor of the present invention will be described in detail with reference to the drawings. Note that the present invention is not limited to the following embodiments as long as it does not deviate from the concept defined in the claims below. Also, the dimensional proportions in the drawings are exaggerated for the convenience of explanation and may differ from the actual proportions.

[0031] Figure 1 is an exploded perspective view showing one embodiment of a sensor 20 of the present invention. Figure 2 is a cross-sectional view of the sensor taken along line aa in Figure 1. Sensor 20 is fabricated as a subcutaneously implantable sensor and is used to measure an analyte in interstitial fluid.

[0032] As shown in FIGS. 1 and 2, electrodes including a working electrode 2, a reference electrode 3, and a counter electrode 4 are formed on an insulating substrate 1 (also simply referred to as a "substrate" in this specification).

[0033] An insulating layer 5 is formed on the working electrode 2, reference electrode 3, and counter electrode 4 formed on the insulating substrate 1 so as to expose the respective electrodes.

[0034] A working electrode working portion 2-1, a reference electrode working portion 3-1, and a counter electrode working portion 4-1 are formed at the tip of the sensor 20. The working electrode working portion 2-1, the reference electrode working portion 3-1, and the counter electrode working portion 4-1 are exposed from the insulating layer 5. A reagent layer 8, an enzyme protective film 9, a biological component diffusion-controlling film 10, and a biocompatible film 11 are sequentially formed on the working electrode working portion 2-1, the reference electrode working portion 3-1, and the counter electrode working portion 4-1. Note that the sensor shown in FIGS. 1 and 2 is formed so that the sample can pass through the reagent layer 8, the enzyme protective film 9, the biological component diffusion-controlling film 10, and the biocompatible film 11 to reach the electrodes.

[0035] The main components of the sensor 20 will now be described in detail.

[0036] <Insulating Substrate> The insulating substrate 1 is not particularly limited, and a conventionally known substrate can be used. One example is a flexible substrate formed from an insulating material. The shape and size of the insulating substrate 1 are also not particularly limited. Examples of resins that can be used include polyethylene terephthalate (PET), polyester, polystyrene, polypropylene, polycarbonate, polyimide, acrylic resin, PEEK (polyether ether ketone), and PAEK (polyaryl ether ketone).

[0037] <Electrodes> The electrodes function as a means for electrically connecting the sensor. In the sensor of this embodiment, the electrodes include at least a working electrode 2. The electrodes are not particularly limited as long as they can electrochemically detect the reaction between the sample (measurement object) and the oxidoreductase, and examples of the electrodes that can be used include carbon electrodes, gold electrodes, silver electrodes, silver / silver chloride electrodes, platinum electrodes, palladium electrodes, and copper electrodes. From the viewpoints of corrosion resistance and cost, carbon electrodes, silver / silver chloride electrodes, and gold electrodes are preferred.

[0038] As the electrodes, a two-electrode system consisting of only a working electrode 2 and a counter electrode 4, or a three-electrode system further including a reference electrode 3 can be preferably used. From the viewpoint of more sensitive potential control, a three-electrode system is more preferable. Furthermore, a sensing electrode for sensing the liquid volume may be further included.

[0039] Furthermore, the portion that comes into contact with the sample supply unit (working portion) may be made of a different material from the other electrode portions. For example, if the reference electrode 3 is made of carbon, the reference electrode working portion 3-1 may be made of silver / silver chloride. Disposable electrodes may also be used.

[0040] <Insulating Layer> The insulating layer 5 functions as an insulating means for preventing short circuits between the electrodes. There are no particular limitations on the material that constitutes the insulating layer 5, but examples thereof include resist ink, resins such as PET and polyethylene, glass, and ceramics.

[0041] <Reagent Layer> The reagent layer 8 is formed on the electrode. The reagent layer 8 has an oxidoreductase, an electron carrier in which an electron carrier is bound to the end of branched polyethylene glycol, and carbon nanotubes.

[0042] (Oxidoreductase) The reagent layer contains an oxidoreductase.

[0043] The oxidoreductase is not particularly limited and can be appropriately selected depending on the type of biological component to be measured. For example, glucose dehydrogenase (GDH), glucose oxidase (GOD), glucose-6-phosphate dehydrogenase, cholesterol dehydrogenase, cholesterol oxidase, glycerophosphate dehydrogenase, glycerophosphate oxidase, lactate dehydrogenase (LDH), lactate oxidase, alcohol dehydrogenase, alcohol oxidase, uricase, 3-hydroxybutyrate dehydrogenase, or urate dehydrogenase can be preferably used.

[0044] The oxidoreductase may contain pyrroloquinoline quinone (PQQ), flavin adenine dinucleotide (FAD), nicotinamide adenine dinucleotide (NAD), nicotine adenine dinucleotide phosphate (NADP), or flavin mononucleotide (FMN) as a coenzyme. For example, glucose dehydrogenases (GDHs) such as glucose dehydrogenase using pyrroloquinoline quinone (PQQ) as a coenzyme (PQQ-GDH), glucose dehydrogenase using flavin adenine dinucleotide (FAD) as a coenzyme (FAD-GDH), glucose dehydrogenase using nicotinamide adenine dinucleotide (NAD) as a coenzyme (NAD-GDH), and glucose dehydrogenase using nicotinamide adenine dinucleotide phosphate (NADP) as a coenzyme (NADP-GDH), and 3-hydroxybutyrate dehydrogenase using nicotinamide adenine dinucleotide (NAD) as a coenzyme may be used. When the oxidoreductase is an oxidoreductase that requires a coenzyme, the reagent layer contains the oxidoreductase and its coenzyme.

[0045] For example, when the analyte is glucose, the oxidoreductase is preferably glucose dehydrogenase or glucose oxidase. When the analyte is cholesterol, the oxidoreductase is preferably cholesterol dehydrogenase or cholesterol oxidase. When the analyte is lactic acid, the oxidoreductase is preferably lactate dehydrogenase (LDH). When the analyte is 3-hydroxybutyric acid, the oxidoreductase is preferably 3-hydroxybutyrate dehydrogenase, which uses nicotinamide adenine dinucleotide (NAD) as a coenzyme.

[0046] Here, the oxidoreductase may be used alone or in combination of two or more kinds.

[0047] The amount of oxidoreductase contained is not particularly limited and can be appropriately selected depending on the type of sample to be measured, the amount of sample added, the type and amount of electron carrier added, and the like.

[0048] (Electron Carrier) The reagent layer contains an electron carrier, which has a structure in which an electron carrier substance is bound to the end of branched polyethylene glycol.

[0049] (Branched Polyethylene Glycol) The branched polyethylene glycol is not particularly limited, and conventionally known branched polyethylene glycols can be used. The branched structure can increase the active sites for binding the electron transfer substance. In addition, a hydration layer can be easily obtained, and the activity of the electron transfer substance bound to the branched polyethylene glycol is less likely to decrease. In addition, reducing the viscosity makes it easier to apply to electrodes.

[0050] The structure of the branched-chain polyethylene glycol is not particularly limited. Preferably, the branched-chain polyethylene glycol is one that allows an electron transfer substance to be bound to each end of the branched chain. The number of branches is also not particularly limited, and is, for example, 3 to 10, preferably 4 or 8. That is, in a preferred embodiment, the branched-chain polyethylene glycol is a 4-branched polyethylene glycol or an 8-branched polyethylene glycol, more preferably a 4-branched polyethylene glycol. This allows at least one molecule of electron transfer substance to be bound per molecule, and up to 4 or 8 molecules of electron transfer substance can be bound. As a result, the efficiency of the enzyme reaction can be further improved. Therefore, a sensor with higher sensitivity can be obtained.

[0051] In a preferred embodiment, the branched polyethylene glycol has the structure represented by the following formula:

[0052]

[0053] In the formula, n's are each independently an integer of 1 to 1,000, and * is a bonding position. Preferably, n's are each independently an integer of 20 to 500, more preferably an integer of 40 to 300.

[0054] In another preferred embodiment, the branched polyethylene glycol has a structure represented by the following formula:

[0055]

[0056] In the formula, n's are each independently an integer of 1 to 1,000, and * is a bonding position. Preferably, n's are each independently an integer of 20 to 500, more preferably an integer of 40 to 300.

[0057] The weight-average molecular weight of the branched-chain polyethylene glycol is not particularly limited, but is preferably 10,000 to 20,000. Within this range, the viscosity can be easily adjusted to an appropriate range. From the same viewpoint, the weight-average molecular weight of the electron carrier is preferably 15,000 to 30,000.

[0058] The weight-average molecular weight of the branched polyethylene glycol or the electron carrier can be measured by gel permeation chromatography (GPC), and in this specification, the value is measured by GPC. Measurement conditions for the GPC method include, for example, the following: (1) Pretreatment: Filtration with a 0.45 μm PTFE cartridge filter (2) Apparatus: HLC-8420GPC (manufactured by Tosoh Corporation) (3) Separation column: TSKgel Super AWM-H (6.0 mm I.D. × 15 cm) (4) Measurement temperature: 40°C (5) Carrier: hexafluoroisopropanol (+10 mM CF ３ (COONa) (6) Flow rate: 0.3 mL / min (7) Injection volume: 20 μL (8) Detector: differential refractometer (RI detector), polarity = (+) (9) Concentration: 1 mg / mL (10) Molecular weight standard: standard polymethyl methacrylate.

[0059] (Electron Transfer Substance) The electron transfer substance is not particularly limited, and known electron transfer substances can be used. Specific examples include phenazine derivatives, diaphorase, nicotinamide adenine dinucleotide phosphate (NADPH), nicotinamide adenine dinucleotide (NADH), or derivatives thereof. The electron transfer substances may be used alone or in combination of two or more.

[0060] The electron carrier is not particularly limited, but from the viewpoint of suppressing aggregation of the electron carrier, it is preferable that the electron carrier does not contain metal complexes such as potassium ferricyanide, sodium ferricyanide, ferrocene and its derivatives, osmium complexes, and ruthenium complexes.

[0061] In the sensor of this embodiment, the electron mediator is preferably a compound having a phenazine structure represented by the following chemical formula: Compounds having a phenazine structure have a redox potential more negative than 0 V (vs. Ag / AgCl saturated KCl), and are therefore less susceptible to the influence of contaminants such as ascorbic acid (vitamin C) and uric acid in biological samples, making it easier to achieve highly accurate detection of biological components.

[0062]

[0063] In the above formula, R ５ is an alkyl group having 1 to 4 carbon atoms, and X － is an anion and * is a binding site.

[0064] Examples of the alkyl group having 1 to 4 carbon atoms include a methyl group, an ethyl group, an n-propyl group, an isopropyl group, an n-butyl group, an isobutyl group, a sec-butyl group, and a tert-butyl group, and preferably an ethyl group.

[0065] Examples of anions include halogen ions, ions of compounds containing halogen, hydroxide ions, carboxylate ions (e.g., acetate ions), nitrate ions, nitrite ions, hydrogen carbonate ions, dihydrogen phosphate ions, hydrogen sulfate ions, alkyl sulfonate ions, hydrogen sulfide ions, hydrogen oxalate ions, cyanate ions, and thiocyanate ions. Examples of ions of compounds containing halogen include hexafluoroantimonate (SbF ６ － ), tetrafluoroborate (BF ４ － ), hexafluorophosphate (PF ６ － ), hexafluoroarsenate (AsF ６ － ), hexachloroantimonate (SbCl ６ － ), trifluoromethanesulfonate ion (CF ３ SO ３ － ), trifluoromethanesulfonate ion (CF ３ SO ３ － ), fluorosulfonate ion (FSO ３ － ), P.F. ３ (C ２ F ５ ) ３ － These include, but are not limited to:

[0066] A specific example of such an electron transfer substance is phenazine ethosulfite (PES) represented by the following formula: wherein * indicates a bonding position.

[0067]

[0068] In a preferred embodiment of the present invention, the electron carrier is a branched polyethylene glycol having an electron carrier covalently bonded to its terminal, either directly or via a linker. The covalent bond may be, for example, a -C(O)-, -C(O)-NH-, -NH-, or -O- bond, and is preferably a -C(O)-NH- (amide bond). According to one embodiment, the phenazine ethosulfite (PES) is a -(CH) ２ The branched polyethylene glycol is linked by a group (linker) selected from the group consisting of -, -C(O)-, -C(O)-NH-, -NH-, -O- and combinations thereof.

[0069] In addition to those containing phenazine ethosulfite as an electron carrier, for example, the following phenazine skeletons into which a substituent may be introduced can be used as electron carriers.

[0070]

[0071] In the above formula, R ６ ~R １２ are each independently a hydrogen atom, an alkyl group having 1 to 6 carbon atoms, a halogenated alkyl group having 1 to 6 carbon atoms, an aryl group having 6 to 30 carbon atoms, a halogen atom, an amino group, an alkylamino group having 1 to 6 carbon atoms, a hydroxyl group (—OH), an alkoxy group having 1 to 6 carbon atoms, an aldehyde group (—CHO), a carboxyl group (—COOH), a nitro group (—NO ２ ), cyano group (-CN), sulfo group (-SO ３ a group having an ester bond such as a mercapto group (-SH), an alkoxycarbonyl group having 2 to 7 carbon atoms or an acyloxy group having 2 to 7 carbon atoms, or an amide group (-CONH ２ ) or a group having an amide bond such as a carbamoylalkyl group having 2 to 7 carbon atoms, and R ５ is a hydrogen atom or an alkyl group having 1 to 4 carbon atoms, and L １ is -(CH ２ ) ａ —C(O)—(CH ２ ) ｂ -, -(CH２ ) ａ -C(O)-NH-(CH ２ ) ｂ -, -(CH ２ ) ａ -NH-C(O)-(CH ２ ) ｂ -, -C(O)-(CH ２ ) ａ -C(O)-, -(CH ２ ) ａ -NH-(CH ２ ) ｂ - and -(CH ２ ) ａ -O-(CH ２ ) ｂ wherein each a is independently an integer from 1 to 6, each b is independently an integer from 1 to 6, * is a bonding position, and X － is an anion.

[0072] In a preferred embodiment of the present invention, the electron carrier is one in which 1 to 10 molecules of an electron carrier are bound to one molecule of branched-chain polyethylene glycol. This configuration can further improve detection sensitivity. When two or more molecules of an electron carrier are bound, they may be the same or different. More preferably, the electron carrier is one in which one molecule of an electron carrier is bound to each end of a branched-chain polyethylene glycol. In a preferred embodiment of the present invention, the electron carrier is one in which 1 to 4 molecules of an electron carrier are bound to one molecule of four-branched polyethylene glycol at the ends thereof.

[0073] In a preferred embodiment of the present invention, the electron carrier is a compound represented by the following chemical formula (1): This makes it possible to obtain the effects of the present invention more significantly.

[0074]

[0075] In chemical formula (1), R １ , R ２ , R ３ , and R ４ At least one of the above is a phenazine derivative represented by the following chemical formula (2):

[0076]

[0077] In chemical formula (2), R ５ are each independently an alkyl group having 1 to 4 carbon atoms; １ are each independently -(CH ２ ) ａ —C(O)—(CH ２ ) ｂ -, -(CH ２ ) ａ -C(O)-NH-(CH ２ ) ｂ -, -(CH ２ ) ａ -NH-C(O)-(CH ２ ) ｂ -, -C(O)-(CH ２ ) ａ -C(O)-, -(CH ２ ) ａ -NH-(CH ２ ) ｂ - and -(CH ２ ) ａ -O-(CH ２ ) ｂ wherein each a is independently an integer from 1 to 6, each b is independently an integer from 1 to 6, * is a bonding position, and X － are each independently an anion, and R １ , R ２ , R ３ , and R ４ Among these, those other than the phenazine derivatives each independently represent H, —CH ２ -NH ２ , -(CH ２ ) ２ -NH ２ , and -(CH ２ ) ３ -NH ２ and n is independently an integer from 1 to 1,000.

[0078] Another aspect of the present invention is an electron carrier represented by the above chemical formula (1).

[0079] In chemical formula (1), R １ , R ２ , R ３ , and R ４ At least one of R １ , R ２ , R ３ , and R ４ It is preferable that two or more of them are phenazine derivatives represented by chemical formula (2), more preferable that three or more of them are phenazine derivatives represented by chemical formula (2), and it is particularly preferable that all of them are phenazine derivatives represented by chemical formula (2).

[0080] In a preferred embodiment, the phenazine derivative is represented by the chemical formula (2), ５ In a preferred embodiment, L in chemical formula (2) is an ethyl group. １ is -(CH ２ ) ａ -NH-C(O)-(CH ２ ) ｂ - or - (CH ２ ) ａ -NH-C(O)-(CH ２ ) ｂ - is.

[0081] R in chemical formula (2) ５ Specific examples of alkyl groups having 1 to 4 carbon atoms represented by the formula: － The specific form of the anion represented by the formula (I) is the same as that described above.

[0082] The method for obtaining the electron carrier is not particularly limited. For example, a compound having an amino group or a carboxyl group and a compound in which a carboxyl group or an amino group is introduced at the end of a branched-chain polyethylene glycol are prepared as derivatives of the electron carrier. The electron carrier can be obtained by reacting these to form an amide bond between the amino group or the carboxyl group bound to the electron carrier and the carboxyl group or the amino group bound to the end of the branched-chain polyethylene glycol. Note that the carboxyl group may be activated with N-hydroxysuccinimide to improve reactivity, to form a group represented by the following formula:

[0083]

[0084] The reaction conditions for reacting a compound having an amino group or a carboxyl group as a derivative of an electron transfer substance with a branched-chain polyethylene glycol having a carboxyl group or an amino group at its terminal are not particularly limited. Preferably, the raw materials are mixed so that the amino group and the carboxyl group are in equal amounts, and the reaction is carried out in a phosphate buffer solution. The pH of the reaction solution is not particularly limited, but is preferably 7.5 to 9, and more preferably 8 to 8.5. The reaction temperature is, for example, 20 to 40°C, and the reaction time is, for example, 1 to 10 hours. The resulting reaction product is preferably purified, for example, by ultrafiltration.

[0085] Although there are no particular limitations on the electron carrier according to this embodiment, it is preferable that the viscosity of an aqueous solution containing the electron carrier at a concentration of 1 mM is 10 to 100 cps as measured at 25° C. using a rotational viscometer. When the viscosity is in this range, the handling property is excellent.

[0086] The content of the electron carrier in the reagent layer is not particularly limited and can be set appropriately.

[0087] (Polymerized Coenzyme) The sensor of this embodiment may further include in the reagent layer a compound (polymerized coenzyme) in which one or more selected from the group consisting of nicotinamide adenine dinucleotide (NAD), pyrroloquinoline quinone (PQQ), and derivatives thereof are bonded to the end of a branched polyethylene glycol. By polymerizing the coenzyme, it becomes easier to stably retain the coenzyme in the reagent layer when the coenzyme is required for the reaction of the enzyme used. This improves the reaction efficiency of the enzyme reaction, and the effects of the present invention can be achieved even more significantly.

[0088] As a specific form of the branched-chain polyethylene glycol (also referred to as second branched-chain polyethylene glycol) to which NAD, PQQ or a derivative thereof is bound at the end, the same branched-chain polyethylene glycol (also referred to as first branched-chain polyethylene glycol) as in the above-mentioned electron carrier can be preferably used. Note that the first branched-chain polyethylene glycol and the second branched-chain polyethylene glycol may be the same or different from each other.

[0089] In a preferred embodiment, the polymerized coenzyme is one in which 1 to 10 molecules of NAD, PQQ, or a derivative thereof are bound to one molecule of branched-chain polyethylene glycol. This can further improve detection sensitivity. When two or more molecules of NAD, PQQ, or a derivative thereof are bound, they may be the same or different. More preferably, the polymerized coenzyme is one in which one molecule of NAD, PQQ, or a derivative thereof is bound to each end of a branched-chain polyethylene glycol. In a preferred embodiment of the present invention, the polymerized coenzyme is one in which 1 to 4 molecules of NAD, PQQ, or a derivative thereof are bound to one molecule of four-branched polyethylene glycol at the end of the four-branched polyethylene glycol.

[0090] In a preferred embodiment, the reagent layer contains the electron carrier represented by the above chemical formula (1) and a polymerized coenzyme, thereby making it possible to obtain the effects of the present invention more significantly.

[0091] The method for obtaining the polymerized coenzyme is not particularly limited. The polymerized coenzyme can be obtained, for example, by reacting a compound having an amino group or a carboxyl group, which is NAD, PQQ, or a derivative thereof, with a compound in which a carboxyl group or an amino group has been introduced at the end of a branched-chain polyethylene glycol, to form an amide bond between the amino group or the carboxyl group of NAD, PQQ, or a derivative thereof and the carboxyl group or the amino group bonded to the end of the branched-chain polyethylene glycol. In this case, the carboxyl group may be activated with N-hydroxysuccinimide to improve reactivity.

[0092] For example, the reaction conditions for reacting NAD, PQQ, or derivatives thereof having an amino or carboxyl group with branched polyethylene glycol having a terminal carboxyl or amino group are not particularly limited. Preferably, the raw materials are mixed so that the amino and carboxyl groups are in equal amounts, and the reaction is carried out in a phosphate buffer solution. The reaction temperature is, for example, 20 to 40°C, and the reaction time is, for example, 1 to 10 hours. The resulting reaction product is preferably purified, for example, by ultrafiltration.

[0093] The content of the polymerized coenzyme in the reagent layer is not particularly limited and can be set appropriately.

[0094] (Carbon Nanotubes) The reagent layer contains carbon nanotubes. The carbon nanotubes act as a conductive material for transferring electrons received from the electron carrier to the electrode. The carbon nanotubes used are single-walled carbon nanotubes, which have a structure in which planar graphite is wound into a single layer into a cylindrical shape. The single-walled carbon nanotubes may be manufactured by, for example, but are not limited to, laser ablation, arc discharge, thermal CVD, plasma CVD, or combustion methods.

[0095] The fiber diameter (average outer diameter on the minor axis side) of the single-walled carbon nanotubes is not particularly limited, but is, for example, 1 to 500 nm. The fiber diameter of the single-walled carbon nanotubes can be obtained by randomly selecting 100 single-walled carbon nanotubes from an observation image of the single-walled carbon nanotubes obtained using a transmission electron microscope, measuring the outer diameter of each minor axis, and calculating the number average value. The fiber length (average length on the major axis side) of the single-walled carbon nanotubes is not particularly limited, but is, for example, 0.1 to 150 μm. The fiber length of the single-walled carbon nanotubes can be obtained by randomly selecting 100 single-walled carbon nanotubes from an observation image of the single-walled carbon nanotubes obtained using a transmission electron microscope, measuring the length of each major axis, and calculating the number average value.

[0096] The carbon purity of the single-walled carbon nanotubes is not particularly limited, but is preferably 80% by mass or more, more preferably 90% by mass or more, even more preferably 95% by mass or more, and particularly preferably 97% by mass or more, based on 100% by mass of the single-walled carbon nanotubes. Here, the single-walled carbon nanotubes may contain a catalyst component, and the carbon purity is expressed as the value obtained by subtracting the ash content from the single-walled carbon nanotubes. The carbon purity value can be calculated according to formula (1) by measuring the ash content after calcining in an air environment at 900°C for 5 hours: Carbon purity of single-walled carbon nanotubes (mass%) = 1 - (mass of ash content (g) after calcination / mass of single-walled carbon nanotubes (g) before calcination) × 100... formula (1).

[0097] In the sensor of this embodiment, the single-walled carbon nanotubes may exist as secondary particles, which may be, for example, in the form of a complex entanglement of single-walled carbon nanotubes, which are primary particles, or an aggregate of linear single-walled carbon nanotubes.

[0098] The single-walled carbon nanotubes may be surface-treated or may have a functional group such as a carboxyl group. Single-walled carbon nanotubes containing organic compounds, metal atoms, fullerenes, etc. may also be used. When surface-treated or functionally-containing single-walled carbon nanotubes are used, the content of the single-walled carbon nanotubes is calculated based on the mass including the surface treatment agent and the functional group. When the single-walled carbon nanotubes contain inclusions, the content of the single-walled carbon nanotubes is calculated based on the mass including the mass of the inclusions.

[0099] The content of single-walled carbon nanotubes in the reagent layer is not particularly limited. The content of single-walled carbon nanotubes in the reagent layer is preferably 50 to 1000 mass % based on the total solid content of the reagent layer. Within this range, the effects of the present invention can be more significantly obtained.

[0100] In the sensor of this embodiment, the reagent layer preferably further contains a dispersant. By using a dispersant, the single-walled carbon nanotubes can be uniformly dispersed. As a result, the effects of the present invention can be obtained even more significantly. The dispersant is not particularly limited, but examples thereof include carboxymethyl cellulose (CMC), carboxyethyl cellulose, gelatin, polyvinylpyrrolidone (PVP), polylysine, etc. The content of the dispersant in the reagent layer is not particularly limited, but is, for example, 10 to 1,000 mass % relative to the total solid content of the reagent layer.

[0101] In the sensor of this embodiment, the reagent layer may further contain a conductive material other than single-walled carbon nanotubes. Examples of conductive materials other than single-walled carbon nanotubes include metal materials such as aluminum, stainless steel (SUS), silver, gold, copper, titanium, and the like, alloys containing these metals, and metal oxides; carbon fibers (specifically, vapor-grown carbon fibers (VGCF), polyacrylonitrile-based carbon fibers, pitch-based carbon fibers, rayon-based carbon fibers, activated carbon fibers, etc.); multi-walled carbon nanotubes; carbon black (specifically, acetylene black, Ketjen Black (registered trademark), furnace black, channel black, thermal lamp black, etc.); graphite; activated carbon; and other carbon materials.

[0102] However, in the sensor of this embodiment, the single-walled carbon nanotubes preferably account for 90% by mass or more of the conductive material contained in the reagent layer, more preferably 95% by mass or more, even more preferably 97% by mass or more, even more preferably 99% by mass or more, particularly preferably 99.5% by mass or more, and most preferably 100% by mass.

[0103] (Other Components) The reagent layer may contain other components (optional components) in addition to the oxidoreductase, electron carrier, polymerized coenzyme, carbon nanotubes, and dispersant. Examples of the other components include surfactants, inorganic salts, and immunosuppressants.

[0104] Examples of surfactants that can be used include nonionic surfactants, amphoteric surfactants, and anionic surfactants. Nonionic surfactants are preferred from the viewpoint of not inhibiting the activity of oxidoreductase. The surfactants may be used alone or in combination of two or more.

[0105] The nonionic surfactant is not particularly limited, but examples thereof include polyoxyethylene alkyl ethers and nonylphenyl polyethylene glycols in which the average number of moles of oxyethylene groups added is 1 to 150 and the number of carbon atoms in the alkyl group is 1 to 18. Such nonionic surfactants may be synthesized or commercially available products may be used. Examples of commercially available products include polyoxyethylene (9) octylphenyl ether (octylphenoxypoly(ethyleneoxy)ethanol or octylphenyl-polyethylene glycol) (Sigma-Aldrich, Nonidet). ＴＭ Polyoxyethylene p-t-octylphenyl ethers (Triton surfactants) such as Triton® X-100 (polyoxyethylene (10) octylphenyl ether), Triton® X-114 (polyoxyethylene (8) octylphenyl ether); polyoxyethylene sorbitan fatty acid esters such as Tween® 85; dodecyl-β-D-maltose; Octyl-β-D-glucoside; Nonidet® P-40 (octylphenoxypoly(ethyleneoxy)ethanol) and Nonidet® P-40 substitutes; Tergitol® NP-10 Surfactant (Nonylphenol Ethoxylate); IGEPAL® CA-630 (octylphenoxypoly(ethyleneoxy)ethanol); Emulgen® 108 (polyoxyethylene lauryl ether), Emulgen® 109P; Brij® 96 polyethylene glycol monooleyl ether (n=approximately 2), and the like can be used.

[0106] Amphoteric surfactants and anionic surfactants may also be synthesized or commercially available. Commercially available products include CHAPS (3-(3-cholamidepropyl)dimethylammonio-1-propanesulphonate), alkylpolyaminoethylglycine chloride, and sodium dodecyl sulfate. Saponin may also be used.

[0107] The total content of the surfactants is, for example, 0.1 to 5 parts by mass relative to 100 parts by mass of the total solid content of the reagent layer.

[0108] The reagent layer is not particularly limited and may be in any form of solid, gel, sol, or liquid, but is preferably solid. The reagent layer may further contain water, a buffer (pH buffer), etc. Here, the buffer is not particularly limited, and buffers generally used when measuring the concentration of a biological component can be used in the same manner. Specifically, phosphate buffer, citrate buffer, citrate-phosphate buffer, trishydroxymethylaminomethane-HCl buffer (trishydrochloric acid buffer), MES buffer (2-morpholinoethanesulfonic acid buffer), TES buffer (N-tris(hydroxymethyl)methyl-2-aminoethanesulfonic acid buffer), acetate buffer, MOPS buffer (3-morpholinopropanesulfonic acid buffer), MOPS-NaOH buffer, HEPES Examples of buffers that can be used include: GOOD buffers such as 4-(2-hydroxyethyl)-1-piperazineethanesulfonic acid buffer, HEPES-NaOH buffer, and amino acid-based buffers such as glycine-hydrochloric acid buffer, glycine-NaOH buffer, glycylglycine-NaOH buffer, and glycylglycine-KOH buffer; boric acid-based buffers such as Tris-borate buffer, boric acid-NaOH buffer, and boric acid buffer; and imidazole buffer. Among these, phosphate buffer, citrate buffer, citrate-phosphate buffer, Tris-hydrochloric acid buffer, MES buffer, acetate buffer, MOPS buffer, and HEPES-NaOH buffer are preferred.

[0109] The method for producing the reagent layer is not particularly limited, and the reagent layer can be formed, for example, by preparing a coating solution containing an oxidoreductase, an electron carrier, and single-walled carbon nanotubes, and optionally containing a dispersant, a coenzyme, a polymerized coenzyme, and other optional components, and coating the coating solution on the electrode.

[0110] The method for preparing the coating solution is not particularly limited. The coating solution may use a buffer solution containing the above-mentioned buffering agent as a medium. In a preferred embodiment, the buffer solution may be phosphate buffered saline (PBS). The concentration of the electron carrier in the coating solution is not particularly limited, but may be adjusted to a concentration of 1 to 5 mM, for example.

[0111] The method for applying the coating liquid is not particularly limited, and methods that can be used include dropping, dip coating, and applying using a coating tool such as a spray coater, a bar coater, a die coater, a reverse coater, a comma coater, a gravure coater, a doctor knife, etc. After applying the coating liquid, it is preferable to dry the coating film.

[0112] The thickness of the reagent layer is not particularly limited, but is, for example, 0.01 to 25 μm, and preferably 0.025 to 10 μm. The method for controlling the thickness is also not particularly limited, but it can be controlled, for example, by adjusting the amount of coating liquid applied to form the reagent layer.

[0113] (Enzyme protective film) The sensor of this embodiment preferably has an enzyme protective film provided on the reagent layer. The enzyme protective film suppresses the reaction between interfering substances in the sample and the oxidoreductase. It also has the effect of protecting the enzyme from the organic solvent used during production. Preferably, the enzyme protective film contains a water-soluble polymer. By providing an enzyme protective film containing a water-soluble polymer, adsorption of adsorptive substances such as proteins contained in the sample solution to the electrode is suppressed, and the measurement sensitivity and reliability of the sensor can be further improved. Examples of water-soluble polymers include cellulose, guar gum, carboxymethyl cellulose, hydroxyethyl cellulose, hydroxypropyl cellulose, methyl cellulose, ethyl cellulose, carboxymethyl ethyl cellulose, polyvinylpyrrolidone, polyvinyl alcohol, polyamino acids such as polylysine, polystyrene sulfonic acid, gelatin, acrylic acid and its salts, methacrylic acid and its salts, starch, maleic anhydride and its salts, agarose gel, tannic acid, pectin, casein, carrageenan, furcellaran, pullulan, collagen, chitin, chitosan, sodium chondroitin sulfate, lignin sulfonic acid, polyvinyl alcohol, polyvinylpyrrolidone, polyacrylamide, polyethylene glycol, and derivatives thereof.

[0114] The enzyme protective film may further contain, in addition to the water-soluble polymer, the optional components described in the section on the reagent layer above.

[0115] The method for producing the enzyme protective film is not particularly limited, and for example, the film can be produced by dissolving a water-soluble polymer in an appropriate solvent, applying the solution to the surface of the reagent layer, and drying the solution.

[0116] The thickness of the enzyme protective film is not particularly limited, but is, for example, 0.01 to 25 μm, preferably 0.025 to 10 μm.

[0117] (Biological component diffusion control membrane) The sensor of this embodiment preferably has a biological component diffusion control membrane on the reagent layer. The sensor of this embodiment may have both an enzyme protection membrane and a biological component diffusion control membrane on the reagent layer. The sensor of this embodiment may have the enzyme protection membrane and the biological component diffusion control membrane on the reagent layer in this order.

[0118] The biological component diffusion control membrane serves to limit the diffusion of the biological component to be analyzed and deliver an appropriate amount of the biological component to the reagent layer. This allows for linearity in the signal from the sensor relative to changes in biological component concentration. Examples of the biological component diffusion control membrane include polymeric materials such as polycarbonate, poly(2-hydroxyethyl methacrylate) (polyHEMA), and urethane resin. Urethane resin is preferred due to its excellent biocompatibility, and aliphatic / polyether-based aqueous urethane resins, which can be dissolved in organic solvents, are even more preferred. The biological component diffusion control membrane can be fabricated, for example, by dissolving a polymeric material in an appropriate solvent, applying it to the surface of the reagent layer or enzyme protection layer, and then drying, or by immersing the coated sensor in a poor solvent and allowing phase separation to occur.

[0119] The thickness of the membrane for controlling the diffusion of biological components is not particularly limited, but is, for example, 0.01 to 25 μm, preferably 0.025 to 10 μm.

[0120] (Biocompatible Membrane) The sensor of this embodiment may have a biocompatible membrane on the reagent layer. When an enzyme protection film and / or a biological component diffusion control film are provided on the reagent layer, the biocompatible membrane may be disposed on these layers. The biocompatible membrane may improve the biocompatibility of the sensor.

[0121] The biocompatible membrane preferably contains a biocompatible polymer. The biocompatible polymer is not particularly limited, but natural or synthetic polymers can be used. Examples of natural polymers include proteins such as collagen, fibronectin, fibrinogen, laminin, and fibrin; and polysaccharides such as chitosan, calcium alginate, heparan sulfate, chondroitin sulfate, hyaluronic acid, heparin, starch, gellan gum, agarose, guar gum, xanthan gum, carrageenan, pectin, carboxymethylcellulose, polyglutamic acid, and polylysine. Examples of synthetic polymers include polyethylene glycol, polypropylene glycol, polyethylene terephthalate, polyvinyl alcohol, thermoplastic elastomers, polypropylene, polyethylene, polystyrene, polymethyl methacrylate, polycarbonate, polydimethylsiloxane, cycloolefin polymers, polylactic acid, polyglycolic acid, polycaprolactone, and polydioxanone. These biocompatible polymers may be used alone or in combination.

[0122] The biocompatible membrane can be prepared, for example, by dissolving a biocompatible polymer in an appropriate solvent, applying the solution to the surface of the reagent layer (or to the outermost surface of the laminate, if an enzyme protection layer or a biological component diffusion-controlling membrane is further provided), and then drying the solution.

[0123] The thickness of the biocompatible membrane is not particularly limited, but is, for example, 0.01 to 25 μm, preferably 0.025 to 10 μm.

[0124] <Application of the Sensor> The sensor of this embodiment can be used to measure the concentration of a specific biological component contained in a biological sample (a biological component measurement target) with high accuracy. The biological component measurement target is not particularly limited as long as it contains the target biological component. Specific examples include blood, as well as body fluids such as urine, saliva, and interstitial fluid.

[0125] The biological components are not particularly limited and include, for example, glucose, 3-hydroxybutyric acid, lactic acid, cholesterol, triglycerides, nicotinamide adenine dinucleotide phosphate (NADPH), nicotinamide adenine dinucleotide (NADH), uric acid, etc., with glucose, 3-hydroxybutyric acid, lactic acid, cholesterol, triglycerides, and uric acid being preferred. That is, according to a preferred embodiment of the present invention, the sensor of the present disclosure is used to measure the concentration of glucose, 3-hydroxybutyric acid, lactic acid, cholesterol, triglycerides, or uric acid in blood or a body fluid.

[0126] There are no particular limitations on the manner in which the sample is supplied to the sensor 20. For example, when the sensor 20 shown in Figures 1 and 2 is inserted into the measurement site, the sample containing the analyte (interstitial fluid when the sensor 20 is inserted subcutaneously) permeates the enzyme protective film or the like, and the analyte reaches the reagent layer 8.

[0127] When a sample is supplied to the reagent layer 8, the biological components in the sample are oxidized by the action of the oxidizing / reducing enzyme, releasing electrons simultaneously with their oxidation. The electrons released from the biological components are captured by the electron carrier, which then converts from an oxidized to a reduced form. When a predetermined potential is applied via the electrode, the reduced electron carrier is electrochemically oxidized and converted to an oxidized form. The amount of reduced electron carrier before the application of the potential (hereinafter also referred to as the "oxidation current") measured during this process can be calculated, and the amount of biological material reacting with the oxidizing / reducing enzyme can be quantified. The value of the potential applied when passing the oxidation current is not particularly limited and can be adjusted appropriately based on conventionally known knowledge. For example, a potential of approximately -200 to +700 mV, preferably -100 to +500 mV, and more preferably 0 to +500 mV is applied between the counter electrode 4 and the working electrode 2. The potential application means for applying the potential are also not particularly limited, and conventionally known potential application means can be used as appropriate.

[0128] The effects of the present invention will be explained using the following examples and comparative examples. However, the technical scope of the present invention is not limited to the following examples. In the following examples, unless otherwise specified, operations were performed at room temperature (25°C). Furthermore, unless otherwise specified, "%" and "parts" mean "% by mass" and "parts by mass", respectively.

[0129] Example 1 Synthesis of Electron Carrier Tetra-PEG (SUNBRIGHT (registered trademark) PTE-100PA, MW: 10,000, manufactured by NOF Corporation) having an n-propylamino group at each end of four branches was prepared as a branched PEG. Amine-reactive PES (1-[3(Succinimidyloxycarbonyl)propoxy]-5-ethylphenazinium triflate, manufactured by Dojindo Laboratories, was prepared as a derivative in which a reactive group was introduced into PES, an electron carrier. These were mixed in a phosphate buffer solution at pH 8.3 so that the amino groups of the branched PEG and the carboxyl groups of the Amine-reactive PES were in equal equivalent amounts, and the mixture was allowed to react at room temperature for 24 hours with stirring.

[0130]

[0131] The reaction solution was subjected to ultrafiltration using a centrifugal ultrafiltration filter. By the above procedure, an electron carrier (PEG-PES) was obtained in which at least one molecule of electron carrier was covalently bound to one molecule of Tetra-PEG. A PEG-PES solution (5 mM in PES equivalent) was obtained by the above procedure.

[0132] (Preparation of Enzyme Solution 1) Enzyme solution 1 was obtained by dissolving 2 mg of glucose dehydrogenase (FADGDH-AB, manufactured by Kikkoman Biochemifa Corporation) as an oxidoreductase in 50 μL of phosphate buffered saline (PBS).

[0133] (Preparation of CNT Dispersion 1) CNT dispersion 1 was obtained by stirring 0.5% w / v of single-walled carbon nanotubes (SWCNT, manufactured by Tokyo Chemical Industry Co., Ltd.) together with 0.1% w / v of carboxymethyl cellulose (CMC) as a dispersant.

[0134] (Sensor Fabrication) The electrodes used were disposable electrodes DEP Chip EP-PP manufactured by Biodevice Technology, Inc. As shown in Figure 1, DEP Chip EP-PP has a working electrode 2, reference electrode 3, and counter electrode 4, each made of carbon, formed on an insulating substrate 1, with an insulating layer 5 sandwiched between them, a working electrode working portion 2-1 made of carbon, a reference electrode working portion 3-1 made of silver / silver chloride, and a counter electrode working portion 4-1 made of carbon.

[0135] A coating solution was obtained by mixing 2 μL of the enzyme solution 1, 2 μL of the PEG-PES solution (5 mM in PES equivalent), and 6 μL of the CNT dispersion 1. 2 μL of the resulting coating solution was dropped onto the working electrode of the electrode so as to cover the working portion, and then appropriately dried to obtain a reagent layer.

[0136] Thereafter, the surface of the reagent layer was dip-coated with a solution prepared by dissolving chitosan (manufactured by Tokyo Chemical Industry Co., Ltd.) in RO water with acetic acid to adjust the concentration to 5% w / v, thereby forming an enzyme protective film, thereby obtaining the sensor of this example.

[0137] Comparative Example 1 A sensor of this comparative example was produced in the same manner as in Example 1, except that CNT dispersion liquid 1 was not added when preparing the coating liquid for producing the reagent layer.

[0138] Comparative Example 2 Preparation of CNT Dispersion 2 CNT dispersion 2 was obtained by stirring 0.5% w / v of multi-walled carbon nanotubes (MWCNT, manufactured by Tokyo Chemical Industry Co., Ltd.) together with 0.1% w / v of carboxymethyl cellulose (CMC) as a dispersant.

[0139] The sensor of this comparative example was produced in the same manner as in Example 1, except that CNT dispersion liquid 2 was used instead of CNT dispersion liquid 1 in Example 1.

[0140] Comparative Example 3 Preparation of CNT Dispersion 3 A single-walled carbon nanotube dispersion (containing 0.4 mass% of single-walled carbon nanotubes and 0.1 mass% of CMC) (Lamfil (registered trademark) WPB-030 manufactured by Kusumoto Chemicals Co., Ltd.) was diluted three times with RO water to prepare CNT dispersion 3.

[0141] In Example 1, when preparing the coating liquid for producing the reagent layer, CNT Dispersion 1 was changed to CNT Dispersion 3, and the following electron transfer substance (SS-Os-1-57) was added instead of PEG-PES, but aggregation of the electron transfer substance occurred, and a sensor could not be produced. SS-Os-1-57 was prepared with reference to J. Am. Chem. Soc. 2003, 125, 16, 4951-4957.

[0142]

[0143] (Electrochemical Measurement) Electrochemical measurements were performed on the sensors prepared above using a potentiostat / galvanostat / impedance measurement device PalmSens 4 (manufactured by Palmsens). A potential of -100 mV was applied between the working electrode and the counter electrode relative to the reference electrode. The glucose concentration at the start of the measurement was set to 0 mg / dL. Every minute from the start of the measurement, the sensor was immersed in sample solutions with glucose concentrations of 0 mg / dL (at the start of the measurement), 50 mg / dL, 100 mg / dL, 200 mg / dL, 400 mg / dL, 200 mg / dL, 100 mg / dL, and 0 mg / dL in PBS, and the current flowing between the working electrode and the counter electrode was measured. The results for the sensors of Example 1 and Comparative Example 1 are shown in Figure 3(a).

[0144] 3(a), the sensor of Example 1 exhibits superior responsiveness at glucose concentrations of 0 to 400 mg / dL compared to the sensor of Comparative Example 1, which does not contain SWCNT. This result shows that the addition of SWCNT improves the efficiency of electron transfer and increases the signal level.

[0145] Furthermore, cyclic voltammetry was performed on the sensors of Example 1 and Comparative Example 1 by immersing them in a PBS solution containing 0 mg / dL of glucose. The results are shown in Figure 3(b). A larger redox current was observed in the sensor of Example 1 compared to the sensor of Comparative Example 1, which did not contain SWCNT, indicating an increased signal level.

[0146] In the sensor of Comparative Example 2, the signal intensity at a glucose concentration of 400 mg / dL was similar to that of the sensor of Comparative Example 1 (not shown), which revealed that the addition of MWCNT did not enhance the signal.

[0147] Example 2 A PEG-PES solution and an enzyme solution 1 were prepared in the same manner as in Example 1. In addition, a CNT dispersion 3 similar to that used in Comparative Example 3 was prepared.

[0148] A coating solution was obtained by mixing 40 μL of enzyme solution 1, 40 μL of PEG-PES solution (5 mM PES equivalent), and 120 μL of CNT dispersion solution 3. The sensor was fabricated by wiring electrodes using carbon paste on a PET substrate with a subcutaneous insertion portion 5 mm long, 0.35 mm wide, and 0.25-0.30 mm thick, and then partially covering it with resist. 12 nL of the above coating solution was applied to the working electrode of this sensor using an inkjet printer to obtain a reagent layer. Then, a 0.1% by mass CMC solution was dip-coated on the surface of the reagent layer and dried to produce an enzyme protective film. Next, a urethane resin dissolved in THF was dip-coated on the surface of the enzyme protective film and dried to produce a biological component diffusion-controlling film, and the sensor of this example was obtained.

[0149] (Electrochemical Measurement) Continuous in vitro solution measurements were performed on the sensor fabricated above using a PalmSens 4 potentiostat / galvanostat / impedance measurement system (manufactured by Palmsens). Glucose was continuously introduced into PBS dissolved in glucose, and an applied voltage of 0.1 V vs. Ag / AgCl was applied. Measurements were performed by immersing the sensor tip in glucose-PBS solutions prepared with glucose concentrations of 50, 200, and 400 mg / dL. The immersion solution was changed at predetermined intervals. Specifically, a sample solution with a glucose concentration of 200 mg / dL was used, and the signal intensity was monitored. Once a day, the solution was switched sequentially to 400 mg / dL and 50 mg / dL, each for one hour. Continuous measurements were performed for 28 days, confirming that a glucose concentration-dependent signal could be obtained with high efficiency over the 28 days. The results are shown in Figure 4.

[0150] (In vivo measurement) The same sensor was implanted into a domestic pig (male) and left in place for one week. The sensor was evaluated on the first, third, and seventh days after implantation. A sample solution of glucose injection (manufactured by Terumo Corporation) with a glucose concentration of 50% was infused intravenously into the pig's auricular vein at a flow rate of 1 mL / kg / h for approximately 30 minutes, followed by a 30-minute cessation of the sample infusion. This cycle was repeated twice to confirm the signal. While acquiring signals from the inserted sensor, blood samples were periodically taken and blood glucose levels were measured using a blood glucose meter (Accu-Check Guide, manufactured by Roche DC Japan Co., Ltd.) to confirm whether the sensor signal followed blood glucose level fluctuations. The sensor of this example was able to acquire signals with similar efficiency from the first day of implantation until one week later, and it was confirmed that the signal obtained from the sensor of this example tracked blood glucose level fluctuations even one week after subcutaneous insertion. In other words, it was confirmed that the sensor of this example did not deteriorate even after one week of indwelling. The results are shown in Figure 5. In FIG. 5, the black circles represent the values ​​measured by the blood glucose meter, and the solid lines represent the signals from the sensor.

[0151] Example 3 A PEG-PES solution and a CNT dispersion 3 were prepared in the same manner as in Example 2.

[0152] (Preparation of Enzyme Solution 2) Enzyme solution 2 was obtained by dissolving 2 mg of lactate dehydrogenase (LDH-E, manufactured by Kikkoman Biochemifa Corporation) as an oxidoreductase in 50 μL of phosphate buffered saline (PBS).

[0153] A coating solution was obtained by mixing 40 μL of enzyme solution 2, 40 μL of PEG-PES solution (5 mM PES equivalent), and 120 μL of CNT dispersion solution 3. The sensor was subcutaneously implantable and was fabricated by wiring electrodes using carbon paste on a PET substrate with a subcutaneous insertion portion measuring 5 mm in length, 0.35 mm in width, and 0.25-0.30 mm in thickness, and then partially covering the substrate with resist. 12 nL of the above coating solution was applied to the working electrode of the sensor using an inkjet printer to obtain a reagent layer. The surface of the reagent layer was then dip-coated with a 0.1% by mass CMC solution and dried to form an enzyme protective film. Next, a urethane resin dissolved in THF was dip-coated on the surface of the enzyme protective film and dried to form a biological component diffusion-controlling film, resulting in the sensor of this example.

[0154] (Electrochemical Measurement) Electrochemical measurements were performed in vitro on the sensor fabricated above using a PalmSens 4 potentiostat / galvanostat / impedance measurement system (manufactured by Palmsens). The applied voltage was 0.1 V vs. Ag / AgCl. Aqueous solutions with lactic acid concentrations of 10, 50, and 100 mg / dL were prepared as sample solutions, and the sensor was repeatedly immersed in each sample solution in turn. Specifically, the sensor was first immersed in a sample solution with a lactic acid concentration of 50 mg / dL to monitor the signal intensity. Once a day, the sample solution was switched to 100 mg / dL and 10 mg / dL for several hours each. Measurements were performed continuously for seven days. It was confirmed that lactic acid could be continuously measured over seven days, and a concentration-dependent signal was obtained. The results are shown in Figure 6. It was confirmed that the sensor of this example can continuously measure lactic acid with high efficiency.

[0155] Example 4 A PEG-PES solution was prepared using the same procedure as in Example 2, except that the PEG-PES solution had a concentration of 1.25 mM in terms of PES equivalent. CNT dispersion 3 was also prepared using the same procedure as in Example 2.

[0156] (Preparation of Enzyme Solution 3) Enzyme solution 3 was obtained by dissolving 2 mg of 3-hydroxybutyrate dehydrogenase (HBD-301, manufactured by Toyobo Co., Ltd.) as an oxidoreductase in 50 μL of phosphate buffered saline (PBS).

[0157] (Preparation of PEG-NAD) As a branched chain PEG, Tetra-PEG (SUNBRIGHT (registered trademark) PTE-100HS, MW: 10,000, manufactured by NOF Corporation) having a succinimidyl group at each end of four branches was prepared. An NAD derivative (6-ADE-NAD, manufactured by Fujifilm Wako Co., Ltd.) was prepared. These were mixed in a phosphate buffer solution at pH 8.3 such that the succinimidyl groups of the branched chain PEG and the amino groups of the NAD derivative were in equal equivalent amounts, and the mixture was reacted at room temperature for 24 hours with stirring.

[0158]

[0159] The reaction solution was subjected to ultrafiltration using a centrifugal ultrafiltration filter, and a polymerized coenzyme (PEG-NAD) was obtained in which at least one NAD molecule was covalently bound to one Tetra-PEG molecule (1.25 mM in NAD equivalent).

[0160] (Sensor Fabrication) The electrode used was a disposable electrode DEP Chip EP-PP manufactured by Biodevice Technology, Inc. 2 μL of the above enzyme solution 3, 2 μL of PEG-PES solution (1.25 mM in PES equivalent), PEG-NAD solution (1.25 mM in NAD equivalent), and 6 μL of CNT dispersion solution 3 were mixed to obtain a coating solution. 2 μL of the obtained coating solution was dropped onto the working electrode of the electrode so as to cover the working portion, and the electrode was appropriately dried to obtain a reagent layer.

[0161] The surface of the reagent layer was then dip-coated with a 0.1% by mass CMC solution and dried to form an enzyme protective film. Next, the surface of the enzyme protective film was dip-coated with a polyurethane dispersion dissolved in THF and dried to form a biological component diffusion-controlling film, and the sensor of this example was obtained.

[0162] Comparative Example 4 The sensor of this comparative example was produced in the same manner as in Example 4, except that the PEG-PES solution was not added when preparing the coating solution for producing the reagent layer.

[0163] Comparative Example 5 The sensor of this comparative example was prepared in the same manner as in Example 4, except that the PEG-NAD solution was not added when preparing the coating solution for producing the reagent layer.

[0164] (Electrochemical Measurement) 3-hydroxybutyric acid was measured using the sensor fabricated above. 3-hydroxybutyric acid is known as a ketone body marker in vivo. Increased blood ketone body concentrations also increase 3-hydroxybutyric acid concentrations. The oxidoreductase used was 3-hydroxybutyrate dehydrogenase, which uses NAD as a coenzyme. Electrochemical measurements were performed using a PalmSens 4 potentiostat / galvanostat / impedance measurement system (PalmSens). A potential of −100 mV relative to the reference electrode was applied between the working and counter electrodes. The 3-hydroxybutyric acid concentration at the start of the measurement was set to 0. The sensor was then immersed in sample solutions containing PBS with 3-hydroxybutyric acid concentrations of 0 (at the start of the measurement), 10, 50, 10, and 0 mg / dL, and the current flowing between the working and counter electrodes was measured approximately every minute. The results for the sensors of Comparative Example 4, Comparative Example 5, and Example 4 are shown in Figure 7.

[0165] 7, it was found that the sensor of Example 4 was able to detect 3-hydroxybutyric acid with high efficiency. In contrast, no signal was detected in the sensors of Comparative Example 4, which did not use PEG-PES, or Comparative Example 5, which did not use an oxidoreductase (coenzyme).

[0166] This application is based on Japanese Patent Application No. 2024-055834, filed on March 29, 2024, the disclosure of which is incorporated herein by reference in its entirety.

[0167] 1 insulating substrate, 2 working electrode, 2-1 working electrode working portion, 3 reference electrode, 3-1 reference electrode working portion, 4 counter electrode, 4-1 counter electrode working portion, 5 insulating layer, 8 reagent layer, 9 enzyme protective film, 10 biological component diffusion control film, 11 biocompatible film, 20 sensor.

Claims

1. A sensor for measuring the concentration of a biological component, comprising an electrode including at least a working electrode and a reagent layer formed on the electrode, wherein the reagent layer comprises: an oxidoreductase; an electron carrier having an electron carrier bound to the end of branched polyethylene glycol; and single-walled carbon nanotubes.

2. The sensor according to claim 1, wherein the branched chain polyethylene glycol has a weight average molecular weight of 10,000 to 20,000.

3. The sensor of claim 1, wherein the branched polyethylene glycol is a four-branched polyethylene glycol.

4. The sensor according to claim 1, wherein the electron carrier is a compound represented by the following chemical formula (1): In chemical formula (1), R １ , R ２ , R ３ , and R ４ At least one of the above is a phenazine derivative represented by the following chemical formula (2): In chemical formula (2), R ５ are each independently an alkyl group having 1 to 4 carbon atoms; １ are each independently -(CH ２ ) ａ —C(O)—(CH ２ ) ｂ -, -(CH ２ ) ａ -C(O)-NH-(CH ２ ) ｂ -, -(CH ２ ) ａ -NH-C(O)-(CH ２ ) ｂ -, -C(O)-(CH ２ ) ａ -C(O)-, -(CH ２ ) ａ -NH-(CH ２ ) ｂ - and -(CH ２ ) ａ -O-(CH ２ ) ｂ wherein each a is independently an integer from 1 to 6, each b is independently an integer from 1 to 6, * is a bonding position, and X － are each independently an anion, and R １ , R ２ , R ３ , and R ４ Among these, those other than the phenazine derivatives each independently represent H, —CH ２ -NH ２ , -(CH ２ ) ２ -NH ２ , and -(CH ２ ) ３ -NH ２ and n is independently an integer from 1 to 1,000.

5. The sensor according to claim 4, wherein the reagent layer further comprises a compound having one or more selected from the group consisting of nicotinamide adenine dinucleotide (NAD), pyrroloquinoline quinone (PQQ), and derivatives thereof bound to the end of a branched polyethylene glycol.

6. The sensor according to claim 1, wherein 1 to 10 molecules of the electron mediator are bound to one molecule of the branched-chain polyethylene glycol.

7. The sensor of claim 1, wherein the oxidoreductase is glucose dehydrogenase (GDH), glucose oxidase (GOD), glucose-6-phosphate dehydrogenase, cholesterol dehydrogenase, cholesterol oxidase, glycerophosphate dehydrogenase, glycerophosphate oxidase, lactate dehydrogenase (LDH), lactate oxidase, alcohol dehydrogenase, alcohol oxidase, uricase, 3-hydroxybutyrate dehydrogenase, or urate dehydrogenase.

8. The sensor according to claim 1, further comprising an enzyme protection film and / or a biological component diffusion control film on the reagent layer.

9. The sensor of claim 1 for measuring the concentration of glucose, 3-hydroxybutyrate, lactic acid, cholesterol, triglycerides, or uric acid in blood or a body fluid.

10. An electron carrier which is a compound represented by the following chemical formula (1): In chemical formula (1), R １ , R ２ , R ３ , and R ４ At least one of the above is a phenazine derivative represented by the following chemical formula (2): In chemical formula (2), R ５ are each independently an alkyl group having 1 to 4 carbon atoms; １ are each independently -(CH ２ ) ａ —C(O)—(CH ２ ) ｂ -, -(CH ２ ) ａ -C(O)-NH-(CH ２ ) ｂ -, -(CH ２ ) ａ -NH-C(O)-(CH ２ ) ｂ -, -C(O)-(CH ２ ) ａ -C(O)-, -(CH ２ ) ａ -NH-(CH ２ ) ｂ - and -(CH ２ ) ａ -O-(CH ２ ) ｂ wherein each a is independently an integer from 1 to 6, each b is independently an integer from 1 to 6, * is a bonding position, and X － are each independently an anion, and R １ , R ２ , R ３ , and R ４ Among these, those other than the phenazine derivatives each independently represent H, —CH ２ -NH ２ , -(CH ２ ) ２ -NH ２ , and -(CH ２ ) ３ -NH ２ and n is independently an integer from 1 to 1,000.

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