WO2016140543A1 - Enzyme-based potentiometric glucose detection sensor and method for manufacturing same - Google Patents

Abstract

The present invention relates to an enzyme-based potentiometric glucose detection sensor and a method for manufacturing same, the glucose detection sensor detecting glucose by measuring a potential difference according to an enzyme reaction without the application of an external voltage, and comprising: an electrode which has the surface thereof modified by a gold-zinc alloyed dendritic oxide; a hydrogen ion sensitive conductive polymer which is formed on the oxide; and a glucose degrading enzyme which is bound to the conductive polymer through a covalent bond.

Description

Enzyme-based potentiometric glucose detection sensor and manufacturing method thereof

BACKGROUND OF THE INVENTION 1. Field of the Invention The present invention relates to a sensor for detecting potentiometric glucose and a method for manufacturing the same, and more particularly, to an enzyme-based glucose sensor and a method for manufacturing the same, including a metal alloy oxide and an electrically conductive polymer.

A sensor is a device that selectively captures a physical quantity or a chemical quantity of a measurement target and converts it into a useful signal. Electrochemical biosensors are used as devices for detecting various analytes in the environment, medicine, and food-related fields by using bioreceptors such as enzymes, microorganisms, and immune bodies. The high response specificity and fast responsiveness make it possible to measure target substances selectively, quickly and accurately.

The blood glucose sensor is a type of electrochemical biosensor which measures the concentration of glucose in blood or urine. The conventional commercially available blood glucose sensor is a second generation biosensor using an electrochemical method (time vs. current method).

The conventional blood glucose sensor has the advantages of high sensitivity, high selectivity to monosaccharides and easy portability, but has ascorbic acid, dopamine, acetaminophen having an oxidation potential similar to the glucose detection potential. Interfering substances, such as acetaminophen, react to each other, resulting in interference (so-called interfering effects) and inferior sensitivity.

On the contrary, the enzymatic potentiometric glucose sensor detects a substance produced by the reaction between glucose and an enzyme and does not need to apply a specific potential, thereby avoiding the interference of the above-mentioned interfering substances. However, the blood glucose sensor using the potentiometric method has not been studied much compared to the blood glucose sensor using the current measurement method.

On the other hand, metal oxides having various nanostructures (nanowire, nanorod, nanotube, nanoparticle, nanofiber, etc.) have been widely applied to enzymatic or non-enzymatic glucose sensors. Early work on metal oxides for glucose detection focused on the use of common electrode materials such as Cu, Ni, Fe, Pt and Au. However, these materials have resulted in fundamental disadvantages such as selectivity, low efficiency and contamination of metal electrode surfaces due to chemisorbed intermediates.

As an alternative, it is possible to use a single and multiple metal alloy oxides of various tree types (dendrite structures) with new structures using various metals (Au, Cu, Pt, Pd-Pt, Pd-Ag, Cu-Co, etc.). Although research is ongoing, studies on multi-metal alloy dendrites and practical applications using them have been insufficient.

In addition, many researches on functional electrically conductive polymers suitable for biological materials (enzymes, proteins, DNA, etc.) have been conducted to solve the problem of contaminants on the surface of metal oxides and selectivity of monosaccharides. As a representative example, studies on chemical and biosensor devices using polypyrrole and polyaniline have already been reported, but these materials have a big problem that they are easily oxidized and decomposed in the air despite the excellent electrical conductivity, thereby significantly reducing their stability.

The present invention provides a sensor for detecting glucose having a simple and high sensitivity by measuring a potential difference according to an enzyme reaction without applying an external voltage.

The present invention also provides a method of manufacturing an enzyme-based potentiometric glucose detection sensor.

The present invention also provides a method for simply measuring glucose concentration by measuring the potential difference according to the enzyme reaction without applying an external voltage.

In one aspect of the invention, the electrode is surface modified with a multi-metal alloy dendrite oxide; A hydrogen ion sensitive conductive polymer formed on the oxide; And a glucose detecting sensor covalently bound to the conductive polymer, wherein the glucose is detected by measuring a potential difference according to an enzyme reaction without applying an external voltage. It provides a sensor for detection.

The sensor of the present invention is a sensor having a two-electrode structure and the electrode can be manufactured using a carbon-based material as the working electrode. For example, carbon ink, graphite, glassy carbon, graphene, or the like may be used as a material of the working electrode, but is not limited thereto. The reference electrode may be a silver electrode (Ag / AgCl), but is not limited thereto.

In one embodiment of the present invention, in the case of a sensor used for disposable, carbon ink may be used as the material of the working electrode and silver ink may be used as the material of the reference electrode.

In the present invention, the working electrode may be modified on the electrode surface of the multi-metal alloy dendrite oxide. In the multi-metal alloy dendrite oxide, the term "multi-metal" means two or more metals, preferably an alloy of two or three metals, more preferably a copper-cobalt alloy or a gold-zinc alloy. Can be.

In one embodiment of the present invention, the working electrode may be an electrode whose surface is modified with gold-zinc alloy dendrites.

The gold-zinc (Au-Zn) alloy dendrites of the present invention have a dendrite structure and can be applied as a preferred electrode in electrochemical devices because of the unique hierarchical structure having a large number of active sites and an extremely high surface area.

The gold-zinc alloy dendrites may have an atomic metal content ratio of Au and Zn of 70-90: 30-10.

Looking at the principle of the enzyme-type potentiometric blood glucose detection, the reaction mechanism for detecting H ⁺ produced by the reaction between glucose and glucose oxidase is as follows.

Glucose + FAD-GOx → Gluconolactone + FADH ₂ -GOx ------------ (1)

FADH ₂ -GOx + O ₂ → FAD-GOx + H ₂ O ₂ ----------------------------- (2)

^{_{H 2 O 2 → 2H + +}} O 2 + 2e - ------------------------------------- ---- (3)

In the present invention, the gold-zinc alloy dendrite oxide film serves as an electrode for detecting hydrogen ions to detect glucose. When hydrogen ions are generated by oxidizing glucose in the blood, which is a glucose degrading enzyme, the gold-zinc alloy dendrites act as probes of hydrogen ions. This mechanism uses the reaction between oxygen and hydrogen ions in the dendrites. Is as follows.

The sensor of the present invention can measure the concentration of glucose by measuring the difference in potential between the reference electrode and the dendrite oxide electrode according to the reaction. As described above, the sensor of the present invention uses an enzymatic potentiometric method, so it is not necessary to apply a specific potential from the outside, so that the measurement can be performed without an external constant voltage meter and an ammeter, so that the measurement is simple and the manufacturing of the sensor is simple. More importantly, since there is no need to apply a specific potential from the outside, interference of the interfering substances reacting to these potentials can be avoided, thereby preventing the interference effect.

The hydrogen ion sensitive conductive polymer may be a conductive polymer including a -COOH or -NH ₂ functional group, and a conductive polymer including a carboxyl group is preferable. More preferably, it may be a conductive polymer including terthiophene having excellent physical, chemical, mechanical and electrical properties. Most preferably, the hydrogen ion-sensitive conductive polymer is terthiophene benzoic acid (TTBA), terthiophene carboxylic acid (TTCA), and disthienyl pyrrole benzoic acid ( dithienyl pirol benzoic acid, DTPBA).

In the present invention, the gold-zinc alloy dendrite oxide film is used as a main material for detecting hydrogen ions, and the hydrogen-sensitive conductive polymer is introduced therein to improve the sensitivity of the dendrite oxide film to hydrogen ions and at the same time, the metal oxide surface. To prevent contamination and to stably stabilize the glucose degrading enzymes through peptide bonds.

Specifically, the hydrogen-ion-sensitive conductive polymer having a carboxylic acid dissociates or associates with -COO ^- and H ⁺ in an aqueous solution to act as a receptor for H ⁺ , thereby forming a hydrogen-ion-sensitive polymer on the gold-zinc alloy dendrites. Incorporation can significantly improve the sensitivity to potential change with respect to hydrogen ion concentration than with only a dendrite oxide film. In addition, when the glycolytic enzyme is in direct contact with the metal, the enzyme is inferior in stability. In the present invention, the conductive polymer is disposed between the metal and the enzyme to protect the surface of the electrode as well as to protect the enzyme, thereby enhancing the enzyme stability. Can be.

In one embodiment of the present invention, through the covalent bond between the carboxyl group of the conductive polymer and the amine group of the enzyme can be coupled to the polymer, thereby improving the stability of the enzyme in the sensor to maintain a high selectivity .

In the present invention, the glucose degrading enzyme is glucose oxidase, glucose dehydrogenase, glucose hexokinase, glutamic oxalacetic transminase and glutamic pyruvic transaminase. It may be selected from the group consisting of.

As is known, glucose oxidase requires cofactor flavin adenine dinucleotide (FAD) in order to act as a catalyst for glucose oxidation reaction. It can be used in the form of an enzyme (FAD-GOx). In the catalytic reaction by glucose oxidase, FAD acts as an electron acceptor. In addition, glucose dehydrogenase, glucose hexokinase, glutamic oxalacetic transaminase, glutamic pyruvic transaminase, and the like, which are involved in glucose metabolism, can be used as glucose degrading enzymes.

The glucose detecting sensor of the present invention may further include a Nafion membrane coated on the layer on which the glucose degrading enzyme is formed.

The Nafion membrane serves to protect the sensor surface. For example, the Nafion membrane protects the surface from factors affecting pH, thereby increasing the stability of the sensor. Sensors protected by Nafion membranes have the advantage of being stored for a long time.

In another aspect, the present invention provides a method for manufacturing a sensor for detecting glucose by measuring the potential difference according to the enzymatic reaction, comprising the steps of: modifying the gold-zinc alloy dendrites on the electrode surface through voltage application; Forming an oxide in an alkaline solution of the gold-zinc alloy dendrites; Modifying the hydrogen-ion (H ⁺ ) sensitized conductive polymer via potential injection on the gold-zinc alloy dendrites; It provides a method of manufacturing a sensor for detecting the potentiometric method comprising the step of binding a glucose degrading enzyme to the conductive polymer through a covalent bond.

In one embodiment of the present invention, the dendrites may be formed by applying a voltage for 150 to 250 seconds at -0.2V to -0.7V.

The alkaline solution may be used without limitation as long as it includes a -OH group, but preferably NaOH, KOH, NH ₄ OH, LiOH, Mg (OH) ₂ , Ca (OH) ₂ , Ba (OH) ₂ , Al (OH) ₃ And the like can be used.

The hydrogen-ion-sensitive conductive polymer and glucose degrading enzyme used in the sensor manufacturing method of the present invention are as described above, and the method of preparing the polymer film of TTCA and DTPBA on the metal dendrites is the same as that of TTBA.

The sensor manufacturing method of the present invention may further include a step of activating the carboxyl group of the conductive polymer by treating the catalyst with the polymer-modified electrode after the step of modifying the polymer and before the step of binding the glucose degrading enzyme.

The catalyst can be used without limitation so long as it is a catalyst for making a peptide bond, and is used as a crosslinking agent. Preferably, at least one from the group consisting of EDC (1-Ethyl-3- [3- (dimethylamino) propyl] carbodiimide hydrochloride), NHS (N-hydroxysuccinimide), and glutaldehyde (glutaldehyde) may be used. no.

In one embodiment of the present invention, the step of coupling the glucose degrading enzyme to the conductive polymer through a covalent bond, by reacting the sugar oxidase on the electrode having the carboxyl group activated by the carboxyl group and the amine group of the sugar oxidase of the conductive polymer It can be made in such a way to form a covalent bond between them.

The method of manufacturing the glucose detecting sensor may further include coating a Nafion membrane on the layer on which the glucose degrading enzyme is formed.

By using the potentiometric glucose detection sensor thus prepared, sugar in blood can be detected easily.

In yet another aspect, the present invention provides a method for measuring glucose concentration by measuring a potential difference according to an enzyme reaction without applying an external voltage, wherein the enzyme is a sugar oxidase and is a gold-modified electrode on a surface of an electrode as a probe of hydrogen ions. It provides a method for measuring glucose concentration, characterized in that using zinc alloy dendrites oxide and hydrogen ion sensitive conductive polymer.

By using the glucose concentration measuring method, it is not necessary to apply an external voltage, thereby eliminating the influence of an interfering substance reacting to the same voltage, and measuring the potential difference using a natural oxidation method through an enzymatic reaction rather than an artificial oxidation method. This makes the measurement simpler and more precise.

The sensor for detecting glucose of the present invention uses the potentiometric method to overcome the disadvantage that the external constant voltage meter and the current meter used by the conventional sensor are degraded by the interference material, so that the application of the external voltage is unnecessary, so that the blocking effect of the interference material is not required. Excellent, using a gold-zinc alloy dendrite oxide film as a hydrogen ion probe, and forming a polymer film between the oxide film and the sugar oxidase to improve the stability of the enzyme reaction to ensure both excellent sensitivity and excellent life and durability It works.

Figure 1 shows a schematic diagram of the production of potentiometric blood glucose sensor.

Figure 2 shows (A) Au-Zn alloy dendrite electrodeposition curve through voltage application, (B) electrochemical polymerization curve of pTTBA using cyclic voltammetry.

3 shows (A) Au-Zn DOx, (B) Au-Zn DOx / pTTBA, (C) Au-Zn DOx / pTTBA / FAD-GOx, and Au-Zn DOx / pTTBA / FAD-GOx / Nafion SEM surface photo.

4 shows the calibration curve for glucose concentration change using the potentiometric method.

5 shows a graph of the effects of temperature and relative humidity.

6 shows a glucose detection sensitivity graph for interfering substances.

Figure 7 shows the calibration curve according to the addition of glucose to the specific simulated blood (medium).

Figure 8 shows the calibration curve using three concentrations of simulated blood (low, medium, high) for the detection of glucose in real blood.

Hereinafter, the configuration and effects of the present invention will be described in more detail with reference to test examples and examples. However, these test examples and examples are provided only for the purpose of illustration in order to facilitate understanding of the present invention, and the scope and scope of the present invention is not limited by the following examples.

[ Example  1] using the multi-metal oxide and the hydrogen-sensitive conductive polymer according to the present invention Potentiometric method  Manufacturing of Blood Glucose Sensors

1 shows a process of preparing a potentiometric glucose sensor. On screen printed carbon electrodes To form a gold-zinc alloy dendrite, 30mM Gold (III) Chloride trihydrate and 30mM Zinc chloride were dissolved together in 0.1M sodium sulfate, and then a precursor solution was prepared by adjusting the pH to 4.0 with 0.1M sodium hydroxide solution. After applying voltage at -0.5V for 200 seconds (FIG. 2A) to form gold-zinc alloy dendrites on the screen printed carbon electrode, the electrode surface was washed with ethanol and tertiary distilled water. Gold-zinc alloy dendrite oxide film is formed by scanning 7 times from + 0.0V to + 1.5V using linear current-voltage method to form gold-zinc alloy dendrite oxide film, and then distilled water and ethanol Washed. 1 mM TTBA, a H ⁺ sensitized polymer monomer, was added to a mixed solvent of di (propylene glycol) methyl ether and tri (propylene glycol) methyl ether (1: 1). After melting, 5 μL was dropped onto the gold-zinc alloy dendrites, dried at 40 ° C., and then electrochemically polymerized in a 0.1 M phosphate buffer (pH 7.4) by a cyclic current-voltage method. The scanning potential range was 0.0V to + 1.6V, and the scanning speed was sweep three times at 100 mv / s to form a polymer film (FIG. 2B).

The electrode of the potentiometric blood glucose sensor was used as a screen print electrode composed of a working electrode and a reference electrode, the working electrode using carbon ink, and the reference electrode was prepared using silver ink.

The pTTBA-formed electrode was added to a solution of 10.0 mM EDC (1-Ethyl-3- [3- (dimethylamino) propyl] carbodiimide hydrochloride) and 10.0 mM NHS (N-hydroxysuccinimide) in 0.1 M phosphate buffer (pH 7.4). The reaction was carried out at 30 ° C. for 6 hours to activate the carboxyl acid group of pTTBA. A 10 μL solution of FAD-GOx (6 mg / mL) was dropped on the surface of pTTBA and reacted for 12 hours at 4 ° C. to form a covalent bond between the carboxyl acid group of pTTBA and the amine group of FAD-GOx. Enzyme cluster formation was also used to increase the amount of enzyme adsorption on the sensor surface. The solution to be used for the enzyme cluster formation method is glutaaldehyde as a crosslinking agent between ammonium sulfate (ammonium sulfate (0.55 mg / mL)) and enzymes (FAD-GOx) as a precipitant of enzyme in FAD-GOx (6mg / mL) solution. , 0.5%) was prepared. 10 μL of the enzyme cluster solution was dropped onto the surface of the screen-printed carbon electrode modified to FAD-GOx and reacted at 4 ° C. at 12 hr. Enzyme adsorption was increased through covalent bonds between amine groups between FAD-GOx. Finally, 1 μL of Nafion (1.0%) was coated on the enzyme layer to prepare a blood glucose sensor.

[ Test Example  1] of each denaturation step of the sensor SEM  Surface image

SEM analysis was performed to confirm the surface of each denaturing step of the novel potentiometric blood glucose sensor according to the present invention.

3 is a SEM surface photograph of each modification step of the sensor. In Figure 3 (A) it was possible to observe the gold-zinc dendrite structure having a multi-tree of gold-zinc. In FIG. 3, (B) is a surface photograph of H ⁺ sensitive polymer electrodeposited by cyclic voltammetry on gold-zinc dendrites. H ⁺ sensitive polymer is formed between the branch trees of gold-zinc dendrites. It was confirmed that the size of the dry and the voids are reduced. (C) in Figure 3 is a surface photograph of the FAD-GDH immobilized on the H ⁺ sensitive polymer. Referring to (B) of FIG. 3, FAD-GDH is immobilized on the H ⁺ sensitized polymer so that no distinct dendrite multi-tree shape is observed and the surface is very rough. FIG. 3 (D) is a surface photograph coated with 0.5% Nafion, covering the FAD-GDH layer in the form of a net film, and has a surface that is softer than the surface photograph of FIG.

From the above results, the surface state of the denaturation step of the potentiometric glucose sensor of the present invention was confirmed.

[ Test Example  2] according to the present invention Potentiometric method  Blood glucose sensor Glucose  Performance Evaluation of Potential Changes with Concentration

Glucose solution of 30mg / dL, 100mg / dL, 200mg / dL, 300mg / dL, 400mg / dL, 500mg / dL dissolved in 0.1M phosphate buffer (pH 7.4) in the blood glucose sensor prepared according to Example 1 Was processed to measure the potential change according to the glucose concentration, and the results are shown in FIG. 4.

Each concentration interval was measured using 10 sensors. Excellent linearity was shown at the sugar concentration of 30-500 mg / dL. The correlation of this calibration curve is E (mv) = (1.80 ± 0.485) + (0.15 ± 0.002) [C] (mg / dL) and the correlation coefficient is 0.998. Relative standard deviation was 5.15% ( n = 10) in the sugar concentration of 200mg / dL. In addition, the relative standard deviation range with respect to the average potential difference value of each concentration range is within ± 5 ~ 13%.

[ Test Example  3] Glucose  Of temperature and relative humidity for measurement Impact  evaluation

The glucose sensor prepared in Example 1 was treated with a glucose solution having a concentration of 200 mg / dL dissolved in 0.1 M phosphate buffer solution (pH 7.4) to measure glucose potential at various temperatures (15, 16, 20, 25, 30, 35, 40 ° C.) and relative humidity (20, 30, 40, 50, 60, 70, 80%) The results are shown in FIG. 5.

Three blood glucose sensors were measured for each section of temperature and relative humidity. In FIG. 5A, the detection sensitivity was constant in the 20 to 40 ° C. range, but the detection sensitivity was reduced by 24% when compared to 25 ° C. at 15 ° C. FIG. It can be confirmed that the enzyme activity is lowered at a temperature of 20 ° C. or lower, and as a result, the glucose measurement sensitivity is lowered. In Figure 5 (A) it was confirmed to maintain a constant glucose detection sensitivity in the relative humidity 20 ~ 80% range.

[ Test Example  4] Glucose  Of disturbances to the measurement Impact  evaluation

6 is a graph comparing glucose detection sensitivity in the presence of interfering substances. Interfering substances used in this experiment were ascorbic acid (AA, 5mg / dL), dopamine (DA, 5mg / dL), acetaaminophen (AP, 15mg / dL), four monosaccharides ((mannose (10mg / dL) dL), lactose (10 mg / dL), xylose (10 mg / dL), fructose (30 mg / dL)).

Irradiation of the interfering substances was determined by including the respective interfering substances in a glucose solution of 200 mg / dL concentration dissolved in 0.1 M phosphate buffer (pH 7.4). In the case of monosaccharides, the four types of monosaccharides described above were included in a 200 mg / dL glucose solution, which was 1.2% lower than the mean potential difference value of glucose. Dopamine, acetaaminophen, and ascorbic acid decreased glucose detection sensitivity by 5.2%, 5.5%, and 4.8%, respectively, and these hinders (monosaccharides, DA, AP, and AA) were found in the error range for glucose measurement. It does not matter because it is distributed within.

[ Test Example  5] In mock blood Glucose  Concentration measurement

Dilute the hematocrit concentration to 10% with physiological saline (0.9% NaCl) and add glucose to the 6 different concentrations (29, 119, 159, 250, 348, 434mg). mock blood comprising glucose with (dL) was prepared.

FIG. 7 is a calibration curve for glucose concentration (29, 119, 159, 250, 348, 434 mg / dL) of 6 sections of simulated blood using the potentiometric method. Three concentrations were measured for each concentration section. Excellent linearity was observed at the glucose concentration of 29 ~ 434 mg / dL. The correlation of this calibration curve is E (mV) = (1.178 ± 0.615) + (0.137 ± 0.004) [C] (mg / dL) and the correlation coefficient is 0.99.

[ Test Example  6] in real blood Glucose  Concentration measurement

Figure 8 is a calibration curve obtained using simulated blood (low, medium, high) having three different glucose concentrations for the actual blood glucose detection. Three concentrations were measured for each concentration section. The correlation between these calibration curves is E (mV) = (19.45 ± 0.626) + (0.137 ± 0.012) [C] (mg / dL). Using the above correlation, glucose was detected from real human blood and the concentration was 99 ± 8.6. In addition, the measurement results using a commercially available blood glucose sensor (OneTouch Ultra) is 105 ± 1.5. The two sensors were used to confirm that the actual blood glucose concentrations were in good agreement. Therefore, the sensor developed through the present invention can be used as a sensor for detecting glucose in real blood.

Claims (14)

1. An electrode whose surface is modified with gold-zinc alloy dendrites;

   A hydrogen ion sensitive conductive polymer formed on the oxide; And

   A sensor for detecting glucose, comprising a glucose degrading enzyme covalently bonded to the conductive polymer,

   The enzyme-based potentiometric glucose detection sensor, characterized in that for detecting glucose by measuring the potential difference according to the enzyme reaction without applying an external voltage.
2. The sensor of claim 1, wherein the electrode is a carbon-based electrode as a working electrode.
3. The method of claim 1, wherein the hydrogen-ion sensitive conductive polymer is selected from the group consisting of terthiophene benzoic acid (TTBA), terthiophene carboxylic acid (TTCA), and dithioenyl pyrrole benzoic acid (DTPBA). A sensor for potentiometric glucose detection, characterized in that it is selected.
4. According to claim 1, wherein the glucose degrading enzyme (glucose oxidase), glucose dehydrogenase (glucose dehydrogenase), glucose hexokinase (glucose hexokinase), glutamic oxalacetic transminase and glutamic pyruvic trans Potentiometric glucose detection sensor, characterized in that selected from the group consisting of minases.
5. The sensor of claim 1, further comprising a Nafion membrane coated on the glucose degrading enzyme.
6. As a method of manufacturing a sensor for detecting glucose by measuring the potential difference according to the enzyme reaction,

   Modifying the gold-zinc alloy dendrites on the electrode surface through voltage application;

   Forming an oxide in an alkaline solution of the gold-zinc alloy dendrites;

   Modifying the hydrogen-ion (H ⁺ ) sensitized conductive polymer via potential injection on the gold-zinc alloy dendrites;

   Comprising the step of coupling the glucose degrading enzyme to the conductive polymer through a covalent bond, the method of producing a sensor for detecting the potentiometric glucose.
7. The method of claim 6, wherein the alkaline solution is selected from the group consisting of NaOH, KOH, NH ₄ OH, LiOH, Mg (OH) ₂ , Ca (OH) ₂ , Ba (OH) ₂ , and Al (OH) ₃ . A method of manufacturing a sensor for detecting potential difference glucose.
8. 7. The method of claim 6, wherein the hydrogen ion sensitive conductive polymer is selected from the group consisting of terthiophene benzoic acid (TTBA), terthiophene carboxylic acid (TTCA), and disthioenyl pyrrole benzoic acid (DTPBA). A method of manufacturing a potentiometric glucose detection sensor, characterized in that it is selected.
9. According to claim 6, wherein the glucose degrading enzyme (glucose oxidase), glucose dehydrogenase (glucose dehydrogenase), glucose hexokinase (glucose hexokinase), glutamic oxalacetic transminase and glutamic pyruvic trans A method for producing a potentiometric glucose detection sensor, characterized in that it is selected from the group consisting of minases.
10. The method of claim 6, further comprising activating a carboxyl group of the conductive polymer by treating a catalyst with a polymer-modified electrode before the glucose degrading enzyme binding step. .
11. The method of claim 10, wherein the catalyst is a catalyst for making peptide bonds.
12. 7. The method of claim 6, further comprising coating a Nafion membrane on the glucose degrading enzyme.
13. A method for detecting sugar in blood using a potentiometric glucose detection sensor manufactured according to any one of claims 6 to 12.
14. As a method of measuring glucose concentration by measuring the potential difference according to the enzyme reaction without applying an external voltage,

    The enzyme is a sugar oxidase,

    A method of measuring glucose concentration, comprising using a gold-zinc alloy dendrite oxide and a hydrogen ion-sensitive conductive polymer modified on an electrode surface as probes of hydrogen ions.

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