KR101856841B1 - Enzyme immobilized glucose biosensor and the manufacturing method thereof - Google Patents

Abstract

The present invention relates to a bioassay for glucose measurement, comprising a substrate comprising a plurality of electrodes and a porous matrix layer disposed on the substrate, wherein the porous matrix layer comprises nanofibers coated with glucose oxidase The present invention relates to a sensor and a method of manufacturing the same, and it is possible to quickly and accurately measure the glucose concentration in the blood without affecting the hematocrit level using the biosensor of the present invention.

Description

TECHNICAL FIELD The present invention relates to an enzyme-immobilized glucose biosensor and a manufacturing method thereof,

The present invention relates to a biosensor for measuring glucose immobilized glucose oxidase and a method for producing the same.

Point-of-care (POC) blood glucose biosensors have been widely used as a primary tool to manage disease in diabetic neonates and caregivers. Blood glucose meters generally require sensitive and accurate results when used in stable adults. It is relatively simple to use and relatively inexpensive, but it has been reported inaccurate in a variety of clinical situations, such as sampling errors, hematocrit level effects in patients with poor peripheral circulation, and changes in oxygen intensity when using the same amount of blood used in the POC glucose meter [Weiner et al., 1991; Louie et al., 2000; Tang et al., 2000; Tang et al., 2001]. In particular, many studies conclude that hematocrit fluctuations can significantly affect the accuracy of POC glucose measurements. This is because the blood cell in the blood contacts the electrode surface and the electro-electrode reaction is recognized. In order to improve the glucose biosensor capability, an enzyme-based nanofiber layer was used in the biosensor to prevent adhesion of the hemocyte and increase the sensitivity of the assay. Nanofiber membranes as enzyme carriers have many advantages over other nanostructured materials, such as high pore size, high surface area ratio, interconnectivity, and ease of fabrication through easy electrospinning techniques, so electrospun nanofiber membranes Has been widely used in the manufacture of biosensors that require sensitive and rapid analysis.

Because of its high hydrophilicity, chemical stability, low toxicity and good biocompatibility, poly (vinyl alcohol) (PVA), which is a semicrystalline polymer, can be used as a blood dialysis membrane, artificial skin, And has played an important role in the hydrogel category, which is applied biomedically.

Many strategies have been explored to reduce the effect of hematocrit on glucose electrodes. Forrow et al. Attempted to exclude erythrocytes using a porous carbon electrode impregnated with an activating reagent, but the analysis time was long, and there was a disadvantage that glucose in the blood cells was oxidized especially at high glucose concentrations.

Glucose oxidase (GOD, EC 1.1.3.4) converts glucose to gluconolactone and spontaneously hydrolyzes to gluconic acid with the simultaneous reduction of dioxygen with hydrogen peroxide. It is accompanied by the reduction of flavinadenine dinucleotide (FAD). FAD is responsible for enzyme reductivity and is tightly bound to the polypeptide protein of the enzyme but not covalently.

It is an object of the present invention to provide a method for measuring glucose, comprising a substrate comprising a plurality of electrodes and a porous matrix layer disposed on the substrate, wherein the porous matrix layer comprises nanofibers coated with glucose oxidase. The present invention provides a biosensor capable of quickly and accurately measuring the glucose concentration without affecting the hematocrit level in the blood.

However, the problems to be solved by the present invention are not limited to the above-mentioned problems, and other matters not mentioned can be clearly understood by those skilled in the art from the following description.

A first aspect of the present invention provides a biosensor for measuring glucose comprising a substrate comprising a plurality of electrodes and a porous matrix layer, wherein the porous matrix layer comprises nanofibers coated with glucose oxidase .

In one embodiment, the nanofibers comprise PVA / PAA, and the molar ratio of PVA to PAA may be from 2: 3 to 3: 2.

In one embodiment, the diameter of the nanofiber coated with glucose oxidase may be 400 to 600 nm.

In one embodiment, the porous matrix layer may be such that the pore diameter before the introduction of the sample and after the introduction of the sample is maintained at 2.5 to 5.0 mu m.

In one embodiment, the thickness of the porous matrix layer may be 10 to 30 [mu] m.

In one embodiment, the biosensor is capable of constantly measuring the glucose concentration measured in an environment of a hematocrit level of 35% to 60% within an error range of 5% or less.

In one embodiment, the glucose concentration, which is constantly measured within an error range of 5% in the biosensor, may be in the range of 37.1 mg / dL to 544.7 mg / dL.

In one embodiment, the biosensor may measure the glucose concentration within 1 to 3 seconds after the sample is introduced into the porous matrix.

According to a second aspect of the present invention, there is provided a method of manufacturing a semiconductor device, comprising: preparing a substrate including a plurality of electrodes; Preparing a hydrophilic polymer solution; Electrospinning the hydrophilic polymer solution on the substrate to produce a porous matrix of nanofibers; Heat treating (crosslinking) the porous matrix; And coating the porous matrix with glucose oxidase. The present invention also provides a method for manufacturing a biosensor for glucose measurement.

In one embodiment, the hydrophilic polymer is PVA and PAA, and the molar ratio of PVA and PAA is from 2: 3 to 3: 2.

In one embodiment, the step of electrospinning the hydrophilic polymer solution on the substrate may comprise: providing a solution having a concentration of the hydrophilic polymer solution of 8 to 20 wt%, a voltage of the electrospinning of 10 to 80 kV, 5 to 30 ml / hr, and the spinning distance may be 10 to 30 cm.

In one embodiment, the step of coating the porous matrix of nanofibers with glucose oxidase may be performed using a screen-impregnation method, a foam impregnation method, a spray method, or a dispensing method.

In one embodiment, the porous matrix layer may be such that the pore diameter before the introduction of the sample and after the introduction of the sample is maintained at 2.5 to 5.0 mu m.

The biosensor of the present invention can quickly and accurately measure the glucose concentration in the blood without affecting the hematocrit level by including the porous matrix layer containing the enzyme-immobilized nanofibers.

1 is a schematic diagram of a biosensor according to an embodiment of the present invention.
2 is an SEM photograph of nanofibers prepared by electrospinning with PVA having different degrees of hydrolysis.
3 is a graph showing the diameter of the PVA according to the degree of hydrolysis.
4 is a SEM photograph of the composite nanofiber according to the content of PVA / PAA.
Figure 5 illustrates cross-linking of nanofibers, according to one embodiment of the present invention.
6 illustrates oxidation and reduction reactions of a biosensor according to an embodiment of the present invention.
FIG. 7 shows a glucose oxidase coating method performed in a conventional biosensor.
8 shows a biosensor according to an embodiment of the present invention.
9 is a photograph showing the surface of a conventional biosensor and a biosensor according to an embodiment of the present invention, respectively.
FIG. 10 is a graph of current versus glucose concentration in blood, as measured by a biosensor according to an embodiment of the present invention.
11 is a graph showing a current versus glucose concentration in blood, as measured by a biosensor according to an embodiment of the present invention.
FIG. 12 is a graph showing the glucose concentration measured according to the hematocrit level, which is measured with a conventional biosensor.
13 is a graph showing a glucose concentration measured according to a hematocrit level, which is measured with a biosensor according to an embodiment of the present invention.
14 is a flowchart of a method of manufacturing a biosensor according to an embodiment of the present invention.
15 is a schematic view showing a method of manufacturing an electrode of a biosensor according to an embodiment of the present invention.
16 is a schematic diagram showing the structure of a biosensor according to an embodiment of the present invention.
17 is a photograph of a step of cutting and attaching a matrix of a biosensor according to an embodiment of the present invention.

In the following, embodiments will be described in detail with reference to the accompanying drawings. Like reference symbols in the drawings denote like elements.

Various modifications may be made to the embodiments described below. It is to be understood that the embodiments described below are not intended to limit the embodiments, but include all modifications, equivalents, and alternatives to them.

The terms used in the examples are used only to illustrate specific embodiments and are not intended to limit the embodiments. The singular expressions include plural expressions unless the context clearly dictates otherwise. In this specification, the terms "comprises" or "having" and the like refer to the presence of stated features, integers, steps, operations, elements, components, or combinations thereof, But do not preclude the presence or addition of one or more other features, integers, steps, operations, elements, components, or combinations thereof.

Unless defined otherwise, all terms used herein, including technical or scientific terms, have the same meaning as commonly understood by one of ordinary skill in the art to which this embodiment belongs. Terms such as those defined in commonly used dictionaries are to be interpreted as having a meaning consistent with the contextual meaning of the related art and are to be interpreted as either ideal or overly formal in the sense of the present application Do not.

In the following description of the present invention with reference to the accompanying drawings, the same components are denoted by the same reference numerals regardless of the reference numerals, and redundant explanations thereof will be omitted. In the following description of the embodiments, a detailed description of related arts will be omitted if it is determined that the gist of the embodiments may be unnecessarily blurred.

Blood is classified into blood cells, which are cellular components, and plasma, which is a liquid component. Blood cells consist of red blood cells carrying oxygen, white blood cells acting as phagocytosis, and platelets involved in blood coagulation. Plasma is composed of 90-92% 7 ~ 8% and 1% of salt. In addition, lipids, saccharides, inorganic salts and nonproteinous nitrogen compounds contain urea, amino acid, and uric acid. A biosensor for measuring glucose measures the concentration of glucose contained in blood plasma. However, due to the interference of substances other than glucose in the blood, the conventional biosensor has a problem that it takes a long time to measure the concentration of glucose or the measured value may not be accurate.

1 is a schematic diagram of a biosensor according to an embodiment of the present invention.

Referring to FIG. 1, the biosensor of the present invention includes a substrate 100 including a plurality of electrodes and a porous matrix layer 200, wherein the porous matrix layer 200 is formed of nanofibers coated with glucose oxidase .

The electrodes are for measuring an electrochemical reaction. The electrodes include a reference electrode for applying power and a working electrode. If necessary, a pair of recognition electrodes for checking the amount of a sample required for analysis may be additionally included have. That is, when a sample is confirmed on a pair of the recognition electrodes, power is applied to the reference electrode and the working electrode to measure the electrochemical reaction. The number and structure of the electrodes can be changed as needed.

The porous matrix layer is attached to the upper surface of the electrodes so as to cover the electrodes so that the blood does not come into contact with the electrodes but can contact the electrodes after passing through the porous matrix layer. The hematocrit, which interferes with the glucose measurement, can be filtered as blood passes through the porous matrix layer.

According to an embodiment of the present invention, the porous matrix layer includes nanofibers coated with glucose oxidase. The pore size of the porous matrix made of nanofibers prepared through electrospinning can be controlled according to the electrospinning condition, and thus glucose can be selectively transmitted through the blood. The nanofibers constituting the porous matrix are coated with glucose oxidase. The glucose oxidase is brought into contact with glucose in the blood to cause a redox reaction, and the biosensor electrode of the present invention is capable of transferring electrons and current The glucose concentration can be measured. In order to increase the contact area with blood, glucose oxidase is preferably coated on the surface of the nanofiber constituting the matrix. In addition to glucose oxidase, the coating layer on the surface of the nanofibers may include an electron transporting agent or a surfactant, and the electron transporting material may be ruthenium or ferrocyanide.

In one embodiment of the present invention, the nanofibers are made of PVA / PAA, and the molar ratio of the PVA and the PAA is 2: 3 to 3: 2. The PVA / PAA is made of hydrophilic polymers. Unlike the conventional polymers, crosslinking is possible even if only heat is used without the use of a crosslinking agent, so that hydrophilic hydrogel properties can be imparted. Particularly, the PVA / PAA produced at an appropriate molar ratio has superiority of maintaining the cut-off function and nanoscale characteristics even in the state of retaining the multi-pore property at the maximum moisture absorption amount, and thus the biosensor of the present invention is made of PVA / PAA ≪ / RTI >

On the other hand, since PVA is derived from the hydrolysis of poly (vinyl acetate: PVAc), the properties of PVA are influenced by degrees of hydrolysis (DH) PVA having a degree of decomposition has lower mechanical properties and water resistance than PVA having a degree of hydrolysis of between 98 and 99.9%, but is easily dissolved in water and excellent in electric radiation (refer to Example 2). Thus, the biosensor matrix PVA having a degree of hydrolysis of 85 to 90 is preferably used, and PVA having a degree of hydrolysis of 88 is most preferably used in order to reduce the diameter of the nanofibers constituting the layer and to form the pores smaller and more closely.

However, since PVA / PAA is bonded through formation of an ester between -OH group of PVA and -COOH group of PAA, there is a problem that it can be hydrolyzed in contact with moisture. In order to solve this problem, the PVA / PAA of the present invention was electrospun at a PVA: PAA molar ratio of 2: 3 to 3: 2 and crosslinked through a simple heat treatment process after electrospinning to increase the water resistance Example 2 and FIG. 4). When the molar ratio of PAA to PVA is less than 0.25 or more than 0.75, the morphology of the fiber is destroyed and it is difficult to maintain the porosity and the function as the nanofiber can not be performed properly. (Examples 2 and 3 4)

In one embodiment of the present invention, the diameter of the nanofiber coated with glucose oxidase may be 400 to 600 nm. The average diameter of the PVA / PAA fibers before coating with glucose oxidase was approximately 300 to 400 nm, but glucose oxidase was relatively closely and uniformly coated on the surface of the nanofibers in the form of crystals to increase the total diameter by approximately 100 to 200 nm . When the diameter of the nanofibers coated with glucose oxidase is less than 400 nm, the surface area is high but the productivity is low. When the diameter of the nanofibers is more than 600 nm, the specific surface area may be lowered.

In one embodiment of the present invention, the pore diameter of the porous matrix layer may be 2.5 to 5.0 m. In order to eliminate the hematocrit which interferes with the measurement of the glucose concentration in the blood, the pore of the porous matrix is preferably 5.0 m or less. Further, when the size of the pores is less than 2.5 占 퐉, permeation of blood does not occur smoothly, and measurement time can be delayed. More preferably, the pore diameter of the porous matrix layer may be between 2.5 and 3.0 m, and in order to more completely eliminate the hematocrit which interferes with the measurement of glucose concentration in the blood, the pore of the porous matrix may be 3.0 m or less May be more preferable.

In addition, the porous matrix layer may be such that the pore diameters thereof before and after the introduction of the sample are maintained at 2.5 to 5.0 mu m. That is, even if a sample such as blood is put into the porous matrix layer, the nanofiber constituting the porous matrix layer of the present invention has high water resistance and is not hydrolyzed, so that the morphology of the fiber can be maintained almost completely (Examples 2 and 3 4). Therefore, the biosensor of the present invention can constantly and accurately measure the glucose concentration due to the porous matrix which maintains a uniform pore size even after the sample is introduced.

In one embodiment of the present invention, the thickness of the porous matrix layer may be 10 to 30 탆. If the porous matrix layer is larger than 30 탆, the thickness of the porous matrix layer is preferably 30 탆 or less, because the porous matrix layer may be trapped or filtered with the hematocrit, and if the porous matrix layer is less than 10 탆, The probability of the emergence of hematocrit increases and the effect of eliminating the hematocrit may be reduced.

In one embodiment of the present invention, the biosensor can measure the glucose concentration measured in an environment of a hematocrit level of 35% to 60% within a tolerance range of 5% or less. After the hematocrit concentration was varied in the samples having the same glucose concentration, glucose was measured by the biosensor of the present invention, and the measured glucose value was constant even when the hematocrit concentration changed from 35% to 60%. (See Example 9 and Fig. 13)

In one embodiment of the present invention, the concentration of glucose measured constantly within an error range of 5% in the biosensor may be 37.1 mg / dL to 544.7 mg / dL.

In one embodiment of the present invention, the biosensor may measure the glucose concentration within 1 to 3 seconds after the sample is introduced into the porous matrix. Conventional glucose biosensors require a measurement time of 5 seconds or more on average because they are subjected to a process of correcting the effect of hematocrit on the measured glucose concentration. However, since the biosensor of the present invention excludes the hematocrit from the porous matrix without the calibration process and measures the glucose concentration, it takes only 1 to 3 seconds to measure the concentration.

14 is a flowchart of a method of manufacturing a biosensor according to an embodiment of the present invention.

Referring to FIG. 14, a method of manufacturing a biosensor for glucose measurement according to the present invention includes: preparing a substrate including a plurality of electrodes (110); Preparing (120) a hydrophilic polymer solution; (130) electrospinning the hydrophilic polymer solution on the substrate to produce a porous matrix of nanofibers; Heat treating the porous matrix (140); And coating (150) the porous matrix with glucose oxidase.

In one embodiment of the present invention, the hydrophilic polymer is PVA and PAA, and the molar ratio of PVA and PAA is 2: 3 to 3: 2.

The step of preparing the enzyme mixture containing glucose oxidase may be further included before the step of coating the porous matrix of nanofibers with glucose oxidase. The enzyme mixture may contain, in addition to glucose oxidase, an electron donor or a surfactant. The electron transporting material may be ruthenium or ferrocyanide.

The step 140 of heat treating the porous matrix may be performed at a temperature of 60 to 200 DEG C for 10 to 120 minutes. When the heat treatment is carried out at a temperature lower than 60 ° C, cross-linking may not be properly formed. When the heat treatment is performed at a temperature of 200 ° C or higher, the polymer in the nanofiber may melt or decompose. Further, when the heat treatment process is performed for less than 10 minutes, sufficient reaction does not occur and partial crosslinking may not occur. If the heat treatment process is performed for 120 minutes or more, the reaction has already been sufficiently performed, There is no.

According to an embodiment of the present invention, the step of electrospinning the hydrophilic polymer solution on the substrate may include the steps of: providing a solution having a concentration of the hydrophilic polymer solution of 8 to 20 wt%, a voltage of the electrospinning of 1 to 80 kV, The spinning speed may be 5 to 30 ml / hr, and the spinning distance may be 10 to 30 cm.

In one embodiment of the present invention, the step of coating the porous matrix made of nanofibers with glucose oxidase may be carried out using a screen filling method, a foam impregnation method, a spray method, or a dispensing method, Any method known in the art as a matrix coating method can be used without limitation.

15 is a schematic view showing a method of manufacturing an electrode of a biosensor according to an embodiment of the present invention. For example, after preparing a PET substrate, the copper film may be adhered to the substrate by a laminating heat treatment, and an ultraviolet (UV) light sensitive dry film may be laminated thereon. UV exposure and wet etching can then be used with a reverse photo mask to form the copper electrode pattern. A nickel layer may be formed on the copper electrode pattern by electroless plating.

16 is a schematic diagram showing the structure of a biosensor according to an embodiment of the present invention. FIG. 16 illustrates an embodiment of the present invention, which may include an intermediate plate, an upper plate, or both on the biosensor electrode of the present invention. An inlet port through which the sample is injected may be formed in the upper substrate. The intermediate substrate may have a channel for guiding the sample to be intensively introduced into the porous matrix layer including the enzyme. In one embodiment of the present invention, the biosensor may include, but is not limited to, an intermediate substrate, an upper substrate, or both.

The method of manufacturing a biosensor of the present invention may further include cutting the porous matrix and attaching the porous matrix to the electrode.

17 is a photograph of a step of cutting and attaching a matrix of a biosensor according to an embodiment of the present invention. Referring to Fig. 17, a porous matrix membrane prepared using electrospinning and coated with glucose oxidase may be cut along a red line, and then adhered onto a substrate on which an electrode pattern is formed.

≪ Example 1: Material preparation >

All structural materials for sensor chips and reactants are commercially available. Polyvinyl alcohol (PVA) having a degree of polymerization (DP) of 1,700 and a degree of hydrolysis (DH) of 88%, 96%, and 99.9% was obtained from Kuraray Co. Ltd., Japan. Polyacrylic acid (PAA) having a molecular weight of 250 kDa was obtained from Wako Pure Chemical Industries, Ltd., Japan. Glucose oxidase (GOD, EC 1.1.3.4, 15,000-25,000 units g- ¹ Type II from Aspergillus niger ), 99% purity ruthenium (Ⅲ) chloride, Triton X-100 wetting agent and β-d - (+) - glucose were purchased from Sigma-Aldrich. Samples, buffers, and enzyme solutions were prepared and a 0.1 M phosphate buffer solution (PBS) was prepared as a mixed solution of KH ₂ PO ₄ and K ₂ HPO _4. The pH of the prepared solution was 7.03. And dissolved in a buffer solution.

≪ Example 2: PVA hydrolysis degree and PVA / PAA content ratio >

PVA having hydrolysis degrees of 88%, 92%, 96% and 99.9%, respectively, were prepared, and then their diameters were measured by electrospinning them under the same conditions.

2 is an SEM photograph of nanofibers prepared by electrospinning with PVA having different degrees of hydrolysis. Fig. 2 (a) shows a PVA nanofiber having a degree of hydrolysis of 88%, Fig. 2 (b) shows a PVA nanofiber with a degree of hydrolysis of 92%, Fig. 2 (c) shows a PVA nanofiber with a degree of hydrolysis of 96% SEM photograph of the fiber. It can be confirmed that the hydrolysis degree of 88% forms the thinnest fiber.

3 is a graph showing the diameter of the PVA according to the degree of hydrolysis. PVA nanofibers having an average diameter of 190 nm and an average degree of hydrolysis of 92% were found to have an average diameter of 200 nm and a degree of hydrolysis of 96%. The PVA nanofibers having an average diameter of 220 nm and a degree of hydrolysis of 99.9% The average diameter of the PVA nanofibers was 470 nm.

The degree of hydrolysis of PVA / PAA composite nanofibers according to the content ratio of PVA / PAA was confirmed. When the molar ratios (PAA) of PAA to PVA were 0.25, 0.5, and 0.75, composite nanofibers were prepared by electrospinning, and the water resistance and degree of hydrolysis of each nanofiber were observed.

4 is a SEM photograph of the composite nanofiber according to the content of PVA / PAA. (A) in the first row (upper line) of FIG. 4 is a nanofiber in the case of φPAA = 0.25, (b) is a nanofiber in the case of φPAA = 0.5, and (c) to be. A photograph of the second row (bottom line) in Fig. 4 is an SEM photograph taken after immersing each nanofiber in water for 1 hour. (a) is a nanofiber when φPAA = 0.25, (b) is a nanofiber with φPAA = 0.5, and (c) is a nanofiber with φPAA = 0.75. When φPAA = 0.5, the most fiber morphology was completely preserved and the water resistance was confirmed to be good.

≪ Example 3: Preparation of PVA / PAA cross-linked nanofiber >

PVA / PAA To prepare cross-linked nanofibers, PVA is first dissolved in water to prepare a solution for electrospinning at a concentration of approximately 10% by weight. In order to improve the solubility of the PVA polymer, the PVA solution is heated to a high temperature of 80 ° C to dissolve the strong internal bonds that may be present in the PVA polymer. The PVA / PAA solution for electrospinning was prepared by mixing PAA solution with PVA solution. The molar ratio of PVA / PAA solution was 0.50 and the concentration of PVA / PAA solution was adjusted to approximately 10 wt%. The PVA / PAA mixed solution is electrospinned on a rotating metal collector under the same conditions. PVA / PAA Mixed nanofibers are thermally treated at 160 ° C for 30 minutes to form an ester between the -OH group of the PVA and the -COOH group of the PAA to chemically cross-link.

Figure 5 illustrates cross-linking of nanofibers, according to one embodiment of the present invention.

≪ Example 4: Preparation of glucose sensor >

The enzyme solution was prepared by mixing 1.4 g of GOD and 600 mg of ruthenium (Ⅲ) chloride as an electric carrier in 20 mL of PBS. The enzyme is immobilized on the PVA / PAA using a physical absorption method. The sensor was dried at room temperature for 24 hours and stored for about one week for enzyme stabilization and maturation before measurement.

The electrochemical measurement was evaluated as a sensor chip composed of two electrode systems and was performed using an electrochemical analyzer. The introduced device included a cyclic voltammeter, amperometer (WonATech Co., Ltd), and WPCIPG software version 1.0. A Ni / Cu electrode was mounted on a polyethylene terephthalate (PET) substrate manufactured by Jaein Circuit Co., Ltd, Korea. All electrochemical devices were performed at 0.25 L in a 0.1 M phosphate buffer solution (PBS, pH 7.03) using Cu / Ni and Cu / Ni / Pt, AgCl and carbon electrodes as reference electrodes at room temperature.

≪ Example 5: Cu / Ni electrode >

The Cu / Ni electrode was prepared using a heat treatment to laminate a 20 탆 thick copper film on a 100 탆 thick PET substrate. An ultraviolet (UV) light sensitive dry film was then attached by the laminating method. A reverse photo mask was used to form the copper electrode pattern by UV exposure and wet etching. The copper electrode is metallic and not treated with water, but oxygen in the air will slowly react at room temperature to form a copper oxide layer on the copper metal. The copper electrode is subsequently electroless plated of nickel sequentially.

≪ Example 6: Measurement method >

SEM images were measured with a Hitachi S4800 scanning electron microscope at an accelerating voltage of 5 kV. The samples were coated with Pt. Fiber diameters were measured using Image J software. For each sample, the fiber diameter was measured at 50-fold and averaged. Electrochemical experiments were performed on a three-electrode system on an IM6ex electrochemical workstation (Zahner, Germany). Cyclic voltammetric measurements were performed in an unstirred electrochemical cell.

6 illustrates oxidation and reduction reactions of a biosensor according to an embodiment of the present invention.

Measurement of glucose in whole blood using an electrochemical biosensor begins when a drop of blood is introduced into the top and sides of the test strip. The electrons generated by the reaction generate an electric current. The current is generated by the electrons generated during the course of glucose oxidation. The current is measured to determine the glucose concentration in the whole-blood sample. Whole-blood glucose concentrations were determined on a YSI 2300 (Yellow Springs Instrument Inc., Yellow Springs, OH) standard blood glucose meter. The YSI standard blood glucose analyzer is used to measure the glucose concentration in duplicate with two glucose channels using glucose oxidase. The linearity is 0 to 1000 mg / dL. The YSI analyzer self-meters every 15 minutes or after 5 measurements. The hematocrit level of each blood sample was determined by centrifuging the sample at 10,000 rpm for 5 minutes with a micro-capillary centrifuge (Model MB, International Equipment Company, Needham Heights, Mass.).

Example 7: Morphology of PVA / PAA cross-linked nanofibers containing GOD.

In order to prevent dissolution, PVA / PAA composite nanofibers have been successfully prepared using heat treatment methods. The formation of a chemically crosslinked network through the formation of ester bonds between the hydroxyl groups of PVA and the carboxyl groups of PAA has been clearly demonstrated through the analysis of ¹³ C-NMR spectra in solid state and FT-IR.

Two kinds of samples (samples in which the GOD solution was directly dispensed on the Cu / Ni electrode and a sample in which the electrospun PVA / PAA nanofiber film was coated with the GOD solution) were prepared and the hematocrit effect of the PVA / PAA nanofiber membrane .

FIG. 7 illustrates a glucose oxidase coating method performed in a conventional biosensor, and FIG. 8 illustrates a biosensor according to an embodiment of the present invention.

9 is a photograph showing the surface of a conventional biosensor and a biosensor according to an embodiment of the present invention, respectively.

9 (a) and 9 (b) are SEM image photographs of a glucose oxidase complex subjected to plasma treatment on a PET film according to a conventional method, wherein GOD is aggregated into a large, non-uniform cluster on the PET film . 9 (c) and 9 (d), on the other hand, show SEM images after vacuum-drying the PVA / PAA crosslinked-bonded composite nanofibers by coating them in a glucose oxidase solution.

Referring to FIG. 9C, when the PVA / PAA nanofiber film is immersed in the GOD equivalent solution, the GOD is relatively uniformly deposited on the PVA / PAA nanofibers according to an embodiment of the present invention . In addition, the fiber diameter was increased from approximately 350 ± 40 nm to 510 ± 50 nm after treatment with GOD solution, and GOD was uniformly formed on PVA / PAA nanofibers. 9 (d)), it was confirmed that the GOD particles formed on the PVA / PAA nanofibers were relatively uniformly deposited on the surface of the nanofiber as nanocrystals in the structure. GOD particles were deposited on the PVA / PAA nanofibers, increasing the overall diameter by approximately 100-150 ± 20 nm, and successfully immobilized on the PVA / PAA nanofibers without agglomeration.

≪ Example 8: Electrochemical ability of PVA / PAA-GOD coated nanofiber >

Electrochemical reactions of Cu / Ni of PVA / PAA-GOD coated nanostructures were confirmed by using cyclic voltammetry.

FIG. 10 is a graph of current versus glucose concentration in blood, as measured by a biosensor according to an embodiment of the present invention. Fig. 10 shows that the cyclic bolt reaction of a PVA / PAA-GOD coated nanostructure layer on a Cu / Ni / PET film has a low electrical resistance (0.01 Ω or less), which allows a biosensor electrode having a uniform electrical resistance to be mass- It is possible to produce it. The low electrical resistance of the sensor is advantageous because the reduction peak occurs faster than the high resistance electrode.

The reproducibility of PVA / PAA-GOD coated nanofiber glucose biosensors was evaluated. The biosensor was tested using the amperometric method.

FIG. 11 is a graph showing current versus glucose concentration in blood, measured by a biosensor according to an embodiment of the present invention, showing the change in reaction current at each whole blood glucose concentration. Referring to FIG. 11, the PVA / PAA-GOD coated nanofibers on the Cu / Ni electrode had a linear correlation coefficient ( R ² ) of 0.9924 and a constant current of 3 or less. The values observed in Figure 11 were measured 10 times on disposable PVA / PAA-GOD coated nanofibers on a Cu / Ni electrode glucose biosensor, resulting in a disposable Cu / Ni electrode glucose biosensor with linearity and considerable reproducibility Respectively. For the PVA / PAA-GOD coated nanofiber Cu / Ni electrode, the overall nonlinear error was 2.15% (coefficient of variation (CV,%) = standard deviation / arithmetic average ㅧ 100). According to ISO 15197 [ISO 15197 First edition, 2003], the production standard allowed for blood glucose sensors for clinical applications is within 7% of the nonlinear error rate.

A whole blood sample having a glucose concentration of 325 and 450 mg / dL was measured using the biosensor of the present invention at a scan rate of 0.1 Vs ^-1 . The bolt profile shows characteristic hydrogen uptake / desorption curves according to the Cu / Ni oxidation and reduction peaks (labeled), confirming the presence of PVA / PAA-GOD coated nanofibers on the Cu / Ni electrode. Glucose was oxidized by electrochemical deformation of ruthenium chloride. The ruthenium chloride was continuously regenerated by the electrodes under oxidizing conditions. The magnitude of the reaction current depends on the glucose concentration. Reaction currents were obtained from a CV result (approximately +0.065 V) performed at total-blood glucose concentration after the blood sample was injected into the PVA / PAA-GOD coated nanofibers. PVA / PAA-GOD coated nanofibers on Cu / Ni / PET films formed biosensor electrodes with very low electrical resistivity. Since the carbon electrode has a higher electrical resistance (approximately 80 Ω to 1 kΩ ± 5 to 10%) than the Cu / Ni electrode (approximately 0.01 Ω) and requires a high applied voltage (approximately 0.2 to 0.7 V) The enzyme reaction requires a longer reaction time to reach a steady state. In contrast, the PVA / PAA-GOD coated nanofiber Cu / Ni electrode reached a steady state within 3 seconds.

Example 9 Influence of Hematocrit on PVA / PAA-GOD Coated Nanofibers [

In highly sensitive glucose biosensors, separation of plasma-containing glucose based on size-exclusion principle is efficient. In human blood, erythrocytes have a diameter of 6 to 9 μm and a thickness of 1.8 to 2.8 μm, and leukocytes have a diameter of 6 to 10 μm or more. Thus, a structure with a cut-off size of about 5 μm can tolerate glucose in plasma, instead of blocking most leukocytes and red blood cells. The nanofibers of the present invention can control the size of voids by controlling the condition of electrospinning, so that they are not affected by blood cells and have high sensitivity to glucose.

Regarding the hematocrit effect, the biosensor of the present invention measured changes in glucose concentration (140, 230 and 309 mg / dL) according to the hematocrit concentration (35, 42, 50, 60%).

FIG. 12 is a graph showing a glucose concentration measured according to a hematocrit level measured with a conventional biosensor, and FIG. 13 is a graph showing a glucose concentration measured with a biosensor according to an embodiment of the present invention. Fig. 12 and 13 are graphs showing the results of measurement of the glucose concentration in the range of 37.1 mg / dL (2.06 mmol / L) to 544.7 mg / dL (30.24 mmol / L) at the level of hematocrit of 35% to 60%.

Referring to FIG. 12, as the hematocrit ratio is increased, the current of the glucose biosensor is reduced because erythrocytes and leukocytes are attached to the electrode surface. According to the Cottrell curling ceremony,

I = (nFAD ^1/2 C) / (tt) ^1/2

F = Faraday constant (charge of one mole of electrons = 96,485 Coulombs / mol), A = area of electrode (cm) ² ), D = diffusion coefficient, C = reaction concentration (mol / cm ³ ), t = time (seconds)).

However, referring to FIG. 13, since the biosensor current of the present invention tested under the same conditions was not influenced by the hematocrit level, the glucose concentration changed within an error range of 5%. This is because current is stably maintained because red and white blood cells are completely blocked by the fiber.

While the present invention has been particularly shown and described with reference to exemplary embodiments thereof, it is to be understood that the invention is not limited to the disclosed exemplary embodiments. For example, if the techniques described are performed in a different order than the described methods, and / or if the described components are combined or combined in other ways than the described methods, or are replaced or substituted by other components or equivalents Appropriate results can be achieved.

Therefore, other implementations, other embodiments, and equivalents to the claims are also within the scope of the following claims.

Claims (13)

A substrate comprising a plurality of electrodes; And
And a porous matrix layer for hematocrit filtration consisting of nanofibers coated with glucose oxidase,
Wherein the porous matrix layer has a pore diameter of 2.5 to 5.0 占 퐉 and a thickness of 10 to 30 占 퐉.
The method according to claim 1,
The nanofiber is made of PVA (polyvinylalcohol) / PAA (polyacrylamide)
Wherein the molar ratio of the polyvinyl alcohol (PVA) to the polyacrylamide (PAA) is from 2: 3 to 3: 2.
The method according to claim 1,
Wherein the diameter of the nanofiber coated with glucose oxidase is 400 to 600 nm.
delete delete The method according to claim 1,
Wherein the biosensor is such that the glucose concentration measured in an environment of a hematocrit level of 35% to 60% is constantly measured within an error range of 5%.
The method according to claim 6,
Wherein a glucose concentration measured constantly within an error range of 5% in the biosensor is 37.1 mg / dL to 544.7 mg / dL.
The method according to claim 1,
Wherein the biosensor is configured such that the glucose concentration is measured in 1 to 3 seconds after the sample is introduced into the porous matrix.
Preparing a substrate comprising a plurality of electrodes;
Preparing a hydrophilic polymer solution;
Preparing a porous matrix for filtering hematocrit comprising nanofibers by electrospinning the hydrophilic polymer solution on the substrate;
Heat treating the porous matrix at a temperature of 60 to 200 DEG C for 10 to 120 minutes; And
And coating the porous matrix with glucose oxidase,
Wherein the porous matrix layer has a pore diameter of 2.5 to 5.0 占 퐉 and a thickness of 10 to 30 占 퐉.
10. The method of claim 9,
The hydrophilic polymer is PVA (polyvinylalcohol) and PAA (polyacrylamide)
Wherein the molar ratio of PVA (polyvinylalcohol) to PAA (polyacrylamide) is 2: 3 to 3: 2.
10. The method of claim 9,
The step of electrospinning the hydrophilic polymer solution on the substrate comprises:
Wherein the concentration of the hydrophilic polymer solution is 8 to 20 wt%
Wherein the voltage of the electrospinning is 1 to 80 kV,
The spinning speed is 5 to 30 ml / hr, and
And a radiation distance is 10 to 30 cm.
10. The method of claim 9,
Wherein coating the porous matrix of nanofibers with glucose oxidase comprises:
Screen inserting method, foam impregnation method, spray method, or dispensing method.
delete

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