KR100464141B1 - Biosensor - Google Patents

Abstract

The present invention relates to a biosensor, the biosensor comprising an electrode, an enzyme layer formed on the electrode and an enzyme protective layer formed on the enzyme layer. The enzyme protective layer comprises a first layer comprising polytrifluorochloroethylene formed on the enzyme layer, a second layer comprising polytetrafluoroethylene formed on the first layer, and a polytrifluoro formed on the second layer. And a third layer comprising rochloroethylene, and a fourth layer comprising fluorine-based ion exchange resin formed on the third layer.

Description

Bio sensor {BIOSENSOR}

[Industrial use]

The present invention relates to a biosensor, and more particularly, to a biosensor having a wide linear reactivity and excellent sensitivity.

[Prior art]

A biosensor refers to a sensor in which a biomaterial is organically coupled with an electronic measuring device to convert chemical information in an arbitrary sample into an easily processed electrical signal. The biosensors are widely developed and widely used in the medical field because of the advantage that the quantitative information of the desired chemical species can be obtained very selectively in real time without the need for complicated chemical and biological treatment. In particular, there are many studies on biosensors for measuring glucose in diseases in which blood glucose should be frequently monitored, such as diabetes.

The biosensor for glucose measurement measures the concentration of glucose, an analyte, by using a sensitive membrane (enzyme layer) immobilized with enzymes that react with glucose, which is a biological substance, to a limited site, generally an electrode.

The method of forming an enzyme layer on the electrode includes a solvent coating method for immersing and drying the electrode in a polymer solution and a method using electrical electropolymerization. Among them, the electrical polymerization method can be applied regardless of the geometry and area of the electrode to be used, and it is mainly used in comparison with the solvent coating method because it has the advantage of controlling the thickness of the enzyme layer or forming a uniform enzyme layer. have. Such an enzyme layer was formed by using a method of immobilizing an enzyme on an electrode while performing electrical polymerization by using a glucose oxidase and a monomer to form a polymer as a monomer. Recently, poly-L-lysine and glutaaldehyde are further used to increase the amount of enzyme added to the biosensor. In addition, a conductive polymer is used as the polymer, and a non-conductive polymer having excellent perm-selectivity and reproducibility is mainly used recently.

The electrode structure and operating principle of the biosensor for glucose measurement are briefly shown in FIG. 1. As shown in FIG. 1, the biosensor is composed of an inner layer, an enzyme layer and an outer layer on a platinum electrode, and is composed of glucose oxidase (GOD) in the enzyme layer by oxygen. The redox reaction takes place and the oxygen is converted to H ₂ O ₂ , which is then converted into H ₂ O ₂ , giving electrons to the platinum electrode as it is converted into glucose gluconic acid present in the sample to be measured according to the redox reaction. Again converted to oxygen and H ⁺ ions, this oxygen is subsequently used to induce the redox reaction of glucose. As a result, the concentration of glucose is measured by measuring a change in current as electrons are generated by the H ₂ O ₂ . At this time, there may be interference of substances that may be oxidized in addition to glucose, such as ascorbic acid and acetaminophen.

As described above, the biosensor has a large influence of oxygen on the sensitivity of the sensor as the reaction occurs by the oxygen present in the enzyme layer. Accordingly, a study to increase the oxygen supply has been studied by Gough [Gough, D. A., Lucisano, J. Y., Tse, P. H. S. Two-Dimentional Enzyme Electrode Sensor for Glusoce, Anal. Chem. 57, 2351, 1985]. Wang and Lu also used polytrifluorochloroethylene with very high oxygen solubility as the oxygen internal source [Wang, J., Lu, F., Oxygen-Rich Oxidase Enzyme Electrodes for Operation in Oxygen-Free Solutions. J. Am. Chen. Soc. 120, 1048, 1998]. However, this type of sensor has not yet been successfully implanted into the human body to measure glucose levels in diabetics.

Zhang et al. Also adjusted the sensitivity and linearity to produce a glucose sensor having a polyurethane outer layer with little effect of oxygen. This polyurethane layer acts as a barrier to prevent glucose diffusion only without affecting oxygen diffusion, thus increasing the oxygen supply relatively.

However, until recently, a glucose sensor using a non-conductive polymer exhibiting a linear linear reactivity and excellent oxygen sensitivity has not been developed. Moreover, the biosensors developed so far have a narrow linear reaction range, and the enzyme layer has a problem in that it is easy to lose activity by an organic solvent.

SUMMARY OF THE INVENTION The present invention has been made to solve the above problems, and an object of the present invention is to provide a biosensor that exhibits a wide linear reactivity and exhibits excellent sensitivity by smoothly supplying oxygen.

Another object of the present invention is to provide an implantable biosensor.

1 is a simplified diagram showing the structure and operating principle of the biosensor for glucose measurement.

2 is a cross-sectional view schematically showing the structure of a biosensor for measuring glucose as an example of the present invention.

Figure 3 is a SEM photograph showing a cross section of the biosensor of the present invention.

4 is in vitro glucose concentration measurement data of the glucose measurement biosensors of Comparative Examples 1 and 1 of the present invention and the glucose measurement biosensor without a conventional enzyme protective layer.

5a to 5f are SEM photographs showing surface morphology change according to the manufacturing process sequence of the biosensor of the present invention.

Figure 6a is a graph showing the in vitro current-time response in serum and phosphate buffer of the biosensor of the present invention.

6B is a graph showing in vitro current-time response in serum containing glucose 6 mM.

7A and 7B are SEM photographs showing surface shapes of the biosensors of Examples 1 and 2 of the present invention, respectively.

8 is a graph showing a calibration curve under various oxygen tensions to determine the oxygen effects of a biosensor having an enzyme protective layer of the present invention and a biosensor without a conventional enzyme protective layer.

9 is a graph showing the life stability of the biosensor of the present invention.

10 is a graph showing a sensor signal measurement result obtained in the in vivo test of the biosensor of the present invention and a graph showing glucose concentration measured by a glucometer.

In order to achieve the above object, the present invention is a biosensor comprising an electrode, an enzyme layer formed on the electrode and an enzyme protective layer formed on the enzyme layer, the enzyme protective layer is polytrifluorochloroethylene formed on the enzyme layer A first layer comprising a; A second layer comprising polytetrafluoroethylene formed on the first layer; A third layer comprising polytrifluorochloroethylene formed on the second layer; It provides a biosensor comprising a fourth layer comprising a fluorine-based ion exchange resin formed on the third layer.

Hereinafter, the present invention will be described in more detail.

TECHNICAL FIELD The present invention relates to a biosensor, and relates to a biosensor that exhibits a broad linear reactivity using a nonconductive polymer.

The biosensor of the present invention comprises an electrode, an enzyme layer formed on the electrode, and an enzyme protective layer formed on the enzyme layer. The enzyme protective layer comprises a first layer comprising polytrifluorochloroethylene (Kel-F Oil) formed on the enzyme layer, a second layer comprising polytetrafluoroethylene formed on the first layer, and the second layer. And a third layer comprising polytrifluorochloroethylene formed on the layer, and a fourth layer comprising perfluorinated ion exchange resin formed on the third layer.

As the electrode, a Teflon-coated platinum (Pt) / iridium (Ir) alloy electrode, which is generally used as an electrode in a glucose measuring sensor, may be used, but is not limited thereto. It can be applied to all kinds of electrodes such as carbon electrode.

The enzyme layer includes enzymes, glutaraldehyde, poly-L-lysine and nonconductive polymers. The enzyme may include, but is not limited to, glucose oxidase, urease, lactate oxidase, lactate dehydrogenase, chlorinese ( enzymes such as creatinase) may be used. As the non-conductive polymer, aminobenzene-based polymers such as 1,3-diaminobenzene, 1,2-diaminobenzene, 1,4-diaminobenzene, or naphthalene-based polymers such as 2,3-diaminonaphthalene may be used. have.

The polytrifluorochloroethylene used in the enzyme protective layer can increase the oxygen supply to the enzyme layer since it does not affect the oxygen diffusion while preventing glucose from diffusing into the enzyme layer. Therefore, the redox reaction of the enzyme may occur actively to increase the sensitivity of the biosensor.

The polytetrafluoroethylene serves to improve the uniformity of the enzyme protective layer, and the polytrifluorochloroethylene of the third layer is well coated, and the fluorine-based ion exchange resin prevents glucose diffusion It protects enzymes while improving biocompatibility and improving biosensor life.

The enzyme protective layer is preferable because the fourth layer structure exhibits a wider linear reactivity than a single layer, and improves the lifespan and biocompatibility of the biosensor.

In the biosensor for glucose measurement of the present invention having such a structure, the electrode is first subjected to anodization at a voltage of 1.6 to 1.8 V in sulfuric acid, and then subjected to surface treatment using cyclic voltammetry, followed by enzymes and writing. Immerse in a buffer comprising rutaaldehyde, poly-L-lysine, and a non-conductive polymer. In this case, when using the Teflon-coated electrode as the electrode, it is preferable to remove the Teflon by the appropriate length from the end of the electrode before use.

The nonconductive polymer is then subjected to electropolymerization at a scanning rate of 0.002 V / s in the range of 0.2 V to 1.0 V. The concentration of enzyme in the buffer is 4.4 uM to 8.8 uM, the concentration of glutaaldehyde is 0.234 mM to 0.468 mM, the concentration of poly-L- lysine is 0.221 mM to 0.442 mM, the concentration of the non-conductive polymer is 1 mM to 5 mM is preferred. If the concentrations of the enzymes, glutaraldehyde, poly-L-lysine and non-conductive polymer are outside of the above-mentioned ranges, the reaction may occur due to the excess of glutaaldehyde or poly-L-rasine before forming a sensitive film on the electrode. Too much progress may result in precipitation in the solution, which lowers the uniformity of the enzyme membrane itself, which is undesirable.

The electrode on which the enzyme layer is formed is immersed in a polytrifluorochloroethylene solution to form a polytrifluorochloroethylene first layer that is an enzyme protective layer on the enzyme layer. In the polytrifluorochloroethylene solution, an alcohol such as ethanol or 2-propanol may be used as a solvent, and the concentration is preferably 5 to 40%. If the concentration of the polytrifluorochloroethylene solution is less than 5%, the membrane is insufficiently coated, which is undesirable. If the concentration of the polytrifluorochloroethylene solution is more than 40%, the membrane is unevenly formed or the enzyme is inactivated, which is not preferable.

The obtained electrode was again coated with a polytetrafluoroethylene dispersion, and then coated with a polytrifluorochloroethylene solution, and the polytrifluorochloroethylene first layer, the polytetrafluoroethylene second layer, and polytrifluor on the enzyme layer A third layer of chloroethylene is formed in turn. The polytetrafluoroethylene dispersion liquid can be used that is dispersed in water, the concentration is preferably 10 to 60%. The polytrifluorochloroethylene solution used to form the third layer uses a solution of the same concentration as when forming the first layer. The thickness of the second polytetrafluoroethylene layer is preferably 2 to 4 µm. When the thickness of this second layer is thicker than 4 mu m, the reproducibility is greatly degraded and the film itself is difficult to be uniformly coated. Also, when the thickness is too thin, the thickness is less than 2 mu m. It is not preferable because it is not appropriate.

In this way, the third layer is formed, and then coated with a fluorine ion exchange resin solution to form a fourth layer of fluorine ion exchange resin to manufacture a biosensor. The concentration of the fluorine-based ion exchange resin is preferably 1 to 5%, and a mixed solvent of water and 2-propanol (for example, 1: 1 volume ratio) may be used as the solvent.

After each coating process, the process of drying at 25-35 degreeC for 1-3 hours is performed.

Although the biosensor of the present invention has been described using glucose measurement as an example, it is not necessarily limited thereto. That is, the present invention can be applied to the fabrication of biosensors in which all kinds of water soluble enzymes and biological elements are immobilized on a solid electrode.

Hereinafter, preferred examples and comparative examples of the present invention are described. However, the following examples are only one preferred embodiment of the present invention and the present invention is not limited to the following examples.

(Comparative Example 1 and Example 1)

1. Material used

Glucose oxidase (EC 1.1.3.4. High purity, 250,000 units / g, without O ₂ type II-S, Argellius niger ), D-(+)-glucose and acetaminophen (AP) was used with Sigma. L-ascorbic acid, 1,3-diaminobenzene (99%), glutaraldehyde (25% by weight solution in water), polytetrafluoroethylene (60%) dispersed in water and fluorine ion exchange resin (Aldrich) ) Was used as it is without commercially available purification. Poly-L-lysine (hydrobromide salt with a molecular weight of 23400) was used after dialysis in 0.1M NaCl. The poly-L-lysine solution was used after storage at 4 ° C. As the polytrifluorochloroethanol oil, Powdercheck (Fremont, Canada) was used.

2. Device

Electrochemical experiments were performed using a model cDAQ-1604 multichannel constant potential supply system (potentiostat, ELBIO Co., Seoul), and Ag / AgCl (3M KCl) and platinum lines were used as reference and counter electrodes, respectively. The surface morphology of the prepared electrode was observed using a scanning electron microscope (SEM) (model JSM 840-A).

3. Manufacturing of the sensor

Teflon-coated platinum-iridium (Pt / Ir) alloy wire was used as the electrode of the needle-type sensor. The Teflon coating was removed at the end of the wire and the removed area was 1.72 mm 2 (3 mm long). The Teflon-coated portion was oxidized with 0.5 M sulfuric acid at +1.6 V for 30 seconds, followed by 25 cycles of potential sweep at 100 mV / s at 0 to +1.2 V. The obtained electrode was washed with nanopure deionized water, and the washed electrode was washed with 0.8 mg / ml glucose oxidase, 2.5 × 10 ⁻⁴ M glutaaldehyde, 2.5 × 10 ⁻⁴ M poly-L-lysine and 5.0 × 10 ^{−. It} was immersed in a phosphate buffer solution containing 3M 1,3-diaminobenzene. The phosphate buffer solution contained 0.1M Na ₂ HPO ₄ , 0.15M NaCl and 0.1g / l NaN ₃ , pH was 7.4. Electropolymerization was carried out in five consecutive cycles at a rate of 2 mV / sec between +0.2 and 1.0V. At this time, the enzyme layer was immobilized on the electrode while 1,3-diaminobenzene was polymerized.

The electrode on which the enzyme layer was formed was then washed with deionized water, which was then immersed in an ethanol solution containing 20% polytrifluorochloroethylene oil and immersed in a 30% polytetrafluoroethylene solution. The electrode was again immersed in an ethanol solution containing 20% polytrifluorochloroethylene oil to prepare a biosensor for glucose measurement of Comparative Example 1.

In addition, the electrode on which the enzyme layer was formed was then washed with deionized water, which was then immersed in an ethanol solution containing 20% polytrifluorochloroethylene oil and immersed in a 30% polytetrafluoroethylene solution. The electrode was again immersed in an ethanol solution containing 20% polytrifluorochloroethylene oil, and finally, the electrode was coated with a 1.5% fluorine-based ion exchange resin solution, and the glucose measurement bio of Example 1 having the structure shown in FIG. The sensor was prepared.

After each coating process, the electrode was dried under vacuum at 30 ° C. for 1 hour.

4. Measurement of in vitro and in vivo physical properties of the sensor

1) in vitro measurement

① Electrode surface shape

3 shows an SEM photograph of the biosensor of Example 1. FIG. The total thickness of the enzyme protective layer of the biosensor was about 3 μm. A relatively bright, 400 nm thin film layer formed on Pt is considered the first polytrifluorochloroethylene layer and the other bright layer formed on the top layer is considered the second polytrifluorochloroethylene and fluorine-based ion exchange resin layer. The dark layer between the two layers is considered polytetrafluoroethylene and the thickness of this layer was about 2 μm.

② Glucose Diffusion Control Effect of Enzyme Protective Layer

In order to determine the glucose diffusion control effect according to the enzyme protective layer structure, the current reactivity of Comparative Example 1 and Example 1 and the biosensor with only the conventional polytetrafluoroethylene enzyme protective layer was measured and the results are shown in FIG. 4. (b), (c) and (a) are shown, respectively.

The experiment immersed the biosensor in a temperature-controlled cell containing 15 ml of phosphate buffer (pH 7.4 37 ° C.). As the phosphate buffer, 0.1 M phosphate buffer containing 0.15 M NaCl and 0.1 g / L NaN ₃ was used.

The biosensor was immersed in the phosphate buffer until the base current stabilized while applying the same potential for current measurement. Subsequently, 1, 3, 6, 9, 12, 15, 18, and 21 mM glucose was added to the cell per 200 seconds to measure the current reactivity of the biosensor. In addition, ascorbic acid (AA in FIG. 4) and acetaminophenone (AP in FIG. 4) were added to examine the effect on the interference of other oxidizable materials.

As shown in FIG. 4, the linear response range was diffused to 21 mM, while the sensitivity was somewhat lowered, and the effect of Example 1 was superior.

③ Surface shape change in electrode manufacturing process

5a to 5f show the shape change of the electrode surface during the sensor manufacturing process. FIG. 5A is a SEM photograph of a basic Pt electrode without a coating layer showing scratches of several tens of nanometers on its surface. An SEM image of the enzyme layer formed on the basic Pt electrode is shown in FIG. 5B. As shown in Figure 5b, it can be seen that the enzyme deposition due to the electropolymerization of 1,3-diaminobenzene fixes the glucose oxidase into the particulate film. In addition, the enzyme layer is very thin, and it can be seen that the scratches of several tens of nanometers formed on the basic Pt surface do not disappear and are partially visible. This enzyme layer was about 54 nm.

A SEM photograph of the polytrifluorochloroethylene oil layer formed on the enzyme layer is shown in FIG. 5C. Since the polytrifluorochloroethylene oil is also coated in the form of particles, it can be seen that the poly (1,3-diaminobenzene) layer does not cover the entire surface.

When the electrode having the structure of FIG. 5C is coated with polytetrafluoroethylene, since the polytetrafluoroethylene is very fine particles, the polytrifluorochloroethylene layer can be uniformly coated to form a very uniform coating layer. This can be done (FIG. 5D). The polytetrafluoroethylene layer also serves as an excellent underlayer for polytrifluorochloroethylene deposition.

The SEM photograph shown in FIG. 5E shows a polytrifluorochloroethylene layer coated on the polytetrafluoroethylene layer, the first polytrifluorochloroethylene layer formed on the poly (1,3-diaminobenzene) As can be seen, it forms a more dense layer. It can thus be expected that a sensor coated with a polytrifluorochloroethylene layer will exhibit a reduced oxygen effect.

Figure 5f is a SEM photograph showing a fluorine-based ion exchange resin is coated on the polytrifluorochloroethylene layer, it can be seen that more compactly deposited with a smaller particle size than the polytrifluorochloroethylene layer.

As such, in the biosensor of Example 1, as the outer layer, which is the enzyme protective layer, has a dense structure, it is expected that the mass transfer rate will be reduced and thus the oxygen effect will be reduced. In addition, the compact structure can impart safety against aggregation of serum.

④ Measurement of sensitivity and reaction time

Table 1 shows the biosensor having the polytrifluorochloroethylene / polytetrafluoroethylene / polytrifluorochloroethylene enzyme protective layer of Comparative Example 1 and the polytrifluorochloroethylene / polytetrafluoroethylene of Example 1 The sensitivity, current, and reaction time for the biosensor with a polytrifluorochloroethylene / fluorine-based ion exchange resin enzyme protective layer were measured and the results are shown. Each base current stabilized within 30 minutes and was found to be reproducible within 1.0 nA.

Enzyme protective layer Basic current ^a [nA / ㎡] Sensitivity ^b [nA / mM / mm2] Response time (T _90% ) Interference ^e / nA / ㎡ Ascorbic Acid (0.11 mM) Acetaminophen (0.17 mM) Comparative Example 1 0.8 ± 0.4 (n = 7) 2.8 ± 0.8 (n = 7) <5 seconds ^c <60 seconds ^d 0.5 0.0 Example 1 0.5 ± 0.2 (n = 7) 1.6 ± 0.6 (n = 7) <5 seconds ^c <60 seconds ^d 0.2 0.0

In Table 1 above,

a: The base current stabilized within 30 minutes.

b: 1-6 mM concentration range exceeded

c: stirring

d: stationary

e: Interference added at maximum physiological concentration in the blood

As shown in Table 1, the sensitivity of the biosensors of Comparative Example 1 and Example 1 is 2.8 ± 0.8 and 1.6 ± 0.6 nA / mm 2 mm 2, respectively, it can be seen that the sensitivity of Example 1 is better. The reaction time (t _90% ) of the sensor was less than 60 seconds in fixed solution and less than 5 seconds in constantly stirred solution.

⑤ serum test

In FIG. 6A, the current time responsiveness in the phosphate buffer (a) and undiluted serum (b) for the biosensor of Example 1 was measured and shown. The base current for the sensor stabilized within 30 seconds. The stabilization time in serum was longer compared to the results in phosphate buffer. Although the reactivity was somewhat reduced, the current increased in proportion to the glucose concentration with the addition of the concentrated glucose solution and was stable during the serum test.

After testing the sensor in serum, the surface morphology was examined using SEM (Comparative Example 1: FIG. 7A, Example 1: 7B). As a result, the decrease in reactivity is believed to be due to the attachment of the biocomponent to the surface of the sensor as shown in FIGS. 7A and 7B.

After the serum test, the stability for continuous monitoring in vitro was performed in serum containing 6 mM glucose and the results are shown in FIG. 6B. The current response was stable since there was little decrease in the current value for more than 15 hours (FIG. 6B).

⑥ oxygen effect

The experiment immersed the biosensor without Example 1 and the conventional enzyme protective layer in a temperature-controlled cell containing 15 ml of phosphate buffer (pH 7.4 37 ° C.). As the phosphate buffer, 0.1 M phosphate buffer containing 0.15 M NaCl and 0.1 g / L NaN ₃ was used.

The oxygen effect was measured by bubbling various partial pressures of oxygen-nitrogen mixed gas into the phosphate buffer, and the measured calibration curve results are shown in FIG. 8. During the experiment, a constant oxygen partial pressure was maintained through the atmosphere and solution of the cell. Oxygen partial pressure was measured by a DO meter (Pentagon 777, BioTel Co.,).

In Figure 8, it can be seen that the sensor (○ (air), □ (40mmHg), △ (15mmHg) without the enzyme protective layer has a slight oxygen dependence as the difference in the slope of the air, 40 and 15mmHg is large, On the other hand, the sensor (● (air), ■ (40mmHg), ▲ (15mmHg)) having the outer layer is clearly shown to have relatively oxygen independence as the inclination difference between air, 40 and 15mmHg is small.

⑦ Life test

In order to examine the life characteristics of the biosensor according to the enzyme protective layer, the biosensor of Example 1 having the enzyme protective layer and the conventional biosensor without the enzyme protective layer were used in 15 ml of phosphate buffer containing 6 mM glucose (pH 7.4 37 ° C.). It was immersed in a temperature-controlled cell containing). 9 shows the measured amount of current.

2) in vivo experiment

① Sensor sensitivity

Ten female Sprague Dawley rats of 400-500 g were anesthetized by intramuscular injection of 50 mg of ketamine. After shaving the anesthetized rat neck, an indwelling catheter was inserted into the cervical vessels and carotid arteries. A 1.5 cm incision was made in the back, and the glucose sensor of Example 1 was inserted into the subcutaneous tissue through this area. After insertion of the sensor, a 50% glucose solution (300 mg / kg body weight) was injected into the catheter of the cervical vein at 20-30 second intervals and blood samples were collected from the carotid artery. All blood samples were analyzed twice using Glucose Analyzer II (Backman).

The current signal was measured with increasing glucose concentration, which was defined as sensor sensitivity. This result is shown in FIG. The current value measured by the sensor in vivo (the glucose concentration measured by applying a constant potential while the sensor is implanted in the body) is shown by a line in FIG. 10, and the glucose concentration result measured by the glucose meter (in rats). The blood glucose value, which is the blood glucose value measured by taking blood when the above-described constant potential is applied, is shown in FIG. In FIG. 10, the sensor current was stabilized in response to infusion within 60 minutes and 1 ml of 50% glucose, and a signal decrease gradually appeared.

The present invention provides a process for the preparation of new outer layers based on polytrifluorochloroethylene thin films to extend the linear response range of glucose sensors produced by electropolymerization. The sensitivity of the biosensor of the present invention was 1.6 ± 0.6 nA / mA mm 2, which is consistent with the optimal sensitivity for glucose sensors in the known in vivo . The biocompatibility is improved according to the fluorine-based ion exchange resin used in the present invention, and is not affected by the change in oxygen value.

Claims (2)

electrode; An enzyme layer formed on the electrode; An enzyme protective layer formed on the enzyme layer As a biosensor comprising: The enzyme protective layer comprises a first layer comprising polytrifluorochloroethylene formed on the enzyme layer; A second layer comprising polytetrafluoroethylene formed on the first layer; A third layer comprising polytrifluorochloroethylene formed on the second layer; And A fourth layer comprising a fluorine-based ion exchange resin formed on the third layer, And The enzyme is a glucose sensor for measuring glucose oxidase. delete