KR101188172B1 - Electrochemical biosensor and method of fabricating the same - Google Patents

Abstract

The present invention relates to an electrochemical biosensor using a diamond thin film electrode of a graphene sheet on which metal nanoparticles are deposited, and a method of manufacturing the same. Biosensor according to the invention, the substrate; An electrochemical biosensor comprising a working electrode for generating an enzyme reaction and a counter electrode for detecting a difference in electrical output from the working electrode, wherein the working electrode comprises: an impurity doped diamond layer; A graphene layer formed on the impurity doped diamond layer; And a metal nanoparticle layer formed on the graphene layer.

Description

Electrochemical biosensor and method of fabricating the same

The present invention relates to a biosensor and a method of manufacturing the same, and more particularly, to an electrochemical biosensor and a method of manufacturing the electrode material improved to improve the sensitivity of the working electrode.

In the conventional biosensors, various biosensors are known that detect or analyze a target substance by fixing a specific substance to a substrate and binding a target substance that specifically binds to the substrate. In particular, among these biosensors, a variety of electrochemical biosensors are known for detecting electrical signals resulting from an electrochemical reaction between enzymes and substrates using enzymes, such as blood glucose levels and blood cholesterol levels. Makes it possible to measure them relatively precisely.

In particular, research on a biosensor for measuring blood glucose concentration has been extensively performed, which measures blood glucose concentration according to the following principle. That is, glucose oxidase (GOx), which is an enzyme capable of oxidizing glucose, is immobilized on the working electrode, and the glucose oxidase converts glucose in the measurement sample into gluconic acid and reduces itself. To glucose glucose oxidase. The reduced glucose oxidase again reacts with oxygen to generate hydrogen peroxide or to convert the electron transfer mediator to a reduced type, by measuring the oxidation current measured during the reoxidation of the hydrogen peroxide or the reduced electron transfer mediator. Blood glucose levels can be measured.

Existing biosensors have been developed with an emphasis on sensitivity improvement through electrode surface treatment. However, in some diseases such as cancer, even a single molecule substance is unavoidable. Therefore, for the accurate diagnosis of various diseases, the detection of trace concentrations and reproducibility, as well as high sensitivity, are important factors in the development of the sensor.

Research using a boron-doped diamond thin film as a sensor electrode has been attempted. In addition to biocompatibility, diamond has excellent chemical stability, thermal conductivity, robustness and optical properties. However, because diamond itself is an insulator that does not conduct electrical current, it exhibits the characteristics of a semiconductor by doping boron. The boron-doped diamond thin film electrode has high signal-to-noise ratio, long-term stability, high sensitivity, and good reproducibility, so it can detect a lower concentration than using an electrode such as gold or platinum. However, since the boron-doped diamond thin film is hydrophobic, it is not easy to be combined with a hydrophilic biomaterial and thus requires a surface treatment technique.

Most of the surface treatment methods studied so far are mostly chemicals, which are complicated, limited, and have a long processing time. For this reason, there are not many cases where boron-doped diamond thin films have been applied to biosensors. Therefore, there is a need for an easy and universal surface treatment technique.

In addition, the amount of biomaterials immobilized on the electrode surface is one of the important factors affecting the sensitivity of the biosensor, and the effective area to immobilize these materials is also an important factor in the development of high sensitivity sensors. In order to overcome these problems, development of nanomaterials having excellent physical properties and development of sensors using them are required.

SUMMARY OF THE INVENTION The present invention has been made in an effort to provide a biosensor capable of detecting high sensitivity and low concentrations of substances and having good reproducibility, and a method of manufacturing the same.

In order to solve the above technical problem, the biosensor according to the present invention, the substrate; An electrochemical biosensor comprising a working electrode for generating an enzyme reaction and a counter electrode for detecting a difference in electrical output from the working electrode, wherein the working electrode comprises: an impurity doped diamond layer; A graphene layer formed on the impurity doped diamond layer; And a metal nanoparticle layer formed on the graphene layer.

The biosensor according to the present invention may further include an enzyme reaction material layer including glucose oxidase on the metal nanoparticle layer, and the metal nanoparticle layer is preferably a platinum nanoparticle layer. Preferably, the graphene layer and the metal nanoparticle layer are graphene sheets on which metal nanoparticles are deposited.

In the biosensor manufacturing method according to the present invention, the method comprises: forming an impurity doped diamond layer; Forming a graphene layer on the impurity doped diamond layer; And forming a metal nanoparticle layer on the graphene layer to form a working electrode.

In particular, forming the impurity doped diamond layer comprises coating a diamond nanocrystalline seed on a substrate; And depositing a boron doped diamond layer on the substrate using B ₂ H ₆ , CH _4, and H ₂ gases. In this case, coating the diamond nanocrystal seed may include dispersing the diamond nanocrystals in an aqueous poly sodium 4-styrene sulfonate] solution; Preparing a silicon substrate thermally oxidized to the substrate; And immersing the substrate in an aqueous solution in which the diamond nanocrystals are dispersed, and then washing the substrate.

The forming of the graphene layer on the impurity doped diamond layer may include attaching a graphene sheet on the impurity doped diamond layer, and the forming of the metal nanoparticle layer on the graphene layer may include: Preferably, the impurity doped diamond layer having the pinned layer is immersed in an aqueous solution of the target metal to deposit nanoparticles of the metal on the graphene layer by an electrochemical method. In particular, depositing platinum nanoparticles on the graphene layer by an electrochemical method of immersing the impurity-doped diamond layer on which the graphene layer is formed in an aqueous solution containing K ₂ PtCl ₆ and applying a potential of -0.6V. desirable.

The biosensor according to the present invention has a large surface area of the working electrode, excellent electrical conductivity, and can increase the amount of immobilization of the enzyme material and increase the detection sensitivity. Accurate measurements can be obtained with only a small amount of raw material. In addition, according to the present invention, a highly sensitive electrochemical biosensor having a fast response time capable of detecting low concentrations can be manufactured with high reproducibility.

In particular, the effective electrode area of the diamond thin film electrode of the graphene sheet on which the platinum nanoparticles prepared according to the present invention is 1.463 cm ² , which is about 3.53 times higher than the effective area (0.415 cm ² ) of the electrode before the surface treatment. The sensitivity was 17.1 μA / mM, showing linearity over a relatively wide concentration range of 1.98 μM to 1.95 mM, and detection at a significantly lower concentration with a detection limit of 0.14 μM. When the platinum nanoparticles are dispersed on a boron-doped diamond thin film without a graphene sheet, the sensitivity is only 4.94 μA / mM, and thus the effect of the graphene layer on the biosensor according to the present invention can be confirmed.

In addition, glucose sensors using the conventional diamond electrode is a form in which the enzyme is immobilized after surface functionalization by chemical treatment only on the diamond electrode, the biosensor according to the present invention showed a sensitivity improvement of about 171 times.

That is, in the present invention, by synthesizing the graphene sheet on which the boron-doped diamond thin film and the platinum nanoparticles are deposited, the synergy effect improves the performance of the sensor such as an electrode area increase, sensitivity improvement, and low detection limit. Platinum nanoparticles increase the surface area and graphene has high electrical conductivity, which increases the sensitivity of the sensor and promotes direct electron transfer between the redox enzyme and the electrode surface. In addition, the diamond thin film has a low capacitive current, so the biosensor using the same can detect low concentrations of analytes.

1 is a plan view of a biosensor according to the present invention.
2 is a cross-sectional view of the working electrode of the biosensor according to the present invention.
3 is a schematic diagram of a process for making an improved working electrode according to the present invention.
4 are SEM images of platinum nanoparticles electrodeposited on a graphene layer at various potentials according to the experimental example of the present invention.
Figure 5 shows the relative CV response characteristics for the biosensor according to the present invention and the comparative example.
6 is a graph showing the relationship between I _p (peak current) and v ^1/2 (scan rate ^1/2 ) for the biosensor according to the present invention and the comparative example.
7 is a graph showing the correlation between the current change and the glucose concentration in the biosensor according to the present invention and the comparative example.

Hereinafter, with reference to the accompanying drawings will be described in detail a preferred embodiment of the biosensor according to the present invention and a method for manufacturing the same. However, the present invention is not limited to the embodiments disclosed below, but will be implemented in various forms, and only the present embodiments are intended to complete the disclosure of the present invention and to those skilled in the art to fully understand the scope of the invention. It is provided to inform you. Therefore, the shapes and the like of the elements in the drawings are exaggerated in order to emphasize a clearer description, and elements denoted by the same symbols in the drawings denote the same elements. In addition, where a layer is described as being "on" another layer or substrate, the layer may exist in direct contact with the other layer or substrate, or a third layer may be interposed therebetween. Can be.

1 is a plan view of a biosensor according to the present invention. Referring to FIG. 1, the biosensor 100 according to the present invention includes a substrate 10, a working electrode 20, a reference electrode 30, and a counter electrode 40.

The substrate 10 used in the present invention is a transparent inorganic substrate such as glass, quartz, Al ₂ O ₃ , SiC, PET (polyethylene terephthalate), PS (polystyrene), PI (polyimide), Transparent organic substrates such as PVC (polyvinyl chloride), PVP (poly vinyl pyrrolidone) and PE (polyethylene) A substrate for use can be used. In the case of forming an electrode or the like on the substrate 10 by a high temperature deposition method on the substrate 10, a semiconductor substrate that can withstand high temperatures is preferable.

The working electrode 20 performs an enzyme reaction with a substance to be detected. The counter electrode 40 detects a difference in electrical output from the working electrode 20. Reference electrode 30 may also be provided for electrochemical measurement. The reference electrode 30 may be a silver chloride electrode, a calomel electrode (SCE), or a mercury sulfate (I) electrode. When the cell is composed of the reference electrode 30 and the working electrode 20, or the reference electrode 30 and the counter electrode 40, and the electromotive force is measured, the potential of the reference electrode 30 is already known. The potential of is known.

The working electrode 20, the reference electrode 30, and the counter electrode 40 are connected to the electrode pads 27, 37, 47 and the conductive wires 25, 35, and 45, respectively. Since the electrode pads 27, 37, 47 are substantially integrated with the conductive wires 25, 35, and 45 to be extended portions of the conductive wires 25, 35, and 45, the materials and thicknesses thereof may be the same. The extension portions 25, 35, and 45 opposite the electrode pads 27, 37, and 47 may form part of the working electrode 20, the reference electrode 30, and the counter electrode 40, respectively. It is preferable to use noble metals such as palladium (Pd), platinum (Pt), gold (Au) and silver (Ag) as materials for the conductive wires 25, 35 and 45. This is because the electrochemical properties such as stability and reproducibility are excellent in the electrode surface region and properties that are difficult to oxidize are excellent. The conductive wires 25, 35, and 45 may be appropriately adjusted in consideration of adhesion to the substrate 10 and economical efficiency, and may be approximately 100 nm thick.

The present invention particularly improves the working electrode 20 material. 2 is a cross-sectional view of the working electrode 20. Referring to FIG. 2, the working electrode 20 includes an impurity doped diamond layer 12, a graphene layer 14 formed on the impurity doped diamond layer 12, and a metal nano formed on the graphene layer 14. A particle layer 16, and an enzyme reactant layer 18.

The impurity doped diamond layer 12 exhibits a high signal-to-noise ratio, long-term stability, high sensitivity, and good reproducibility compared to conventional metal electrodes such as gold and silver. In addition, since the impurity doped diamond layer 12 has a low capacitive current, the biosensor 100 using the same may detect a low concentration of analyte. The impurity doped diamond layer 12 is laminated to an appropriate thickness in consideration of conductivity, adhesion, and economics.

In the present invention, the graphene layer 14 and the metal nanoparticle layer 16 are formed on the impurity doped diamond layer 12 in order to improve the surface treatment and the electrode area when only the diamond thin film is used as the electrode. Preferably, the graphene layer 14 and the metal nanoparticle layer 16 are graphene sheets on which metal nanoparticles are deposited.

Graphene is a honeycomb two-dimensional thin film made of a layer of carbon atoms. When carbon atoms are chemically bonded by sp2 hybrid orbits, carbon atoms form a carbon hexagonal network spread in two dimensions. Graphene is an aggregate of carbon atoms having this planar structure. Graphene is flexible, has very high electrical conductivity and has a large surface area. However, the surface treatment method is complicated and limited, making it difficult to use in biosensors.

In the present invention, the metal nanoparticle layer 16 is formed on the graphene layer 14 to improve the surface treatment method more easily and effectively. In general, the metal film itself is known to easily change the biomaterials to reduce the activity, but the metal nanoparticles used in the present invention does not change the easily adsorbed biomaterial. Therefore, a large amount of biomaterials may be immobilized on the surface of the metal nanoparticle layer 16 through adsorption. The metal nanoparticle layer 16 maximizes the electron transfer generated by the electrode area and the chemical reaction.

As described above, the biosensor 100 according to the present invention forms the graphene layer 14 and the metal nanoparticle layer 16 on the impurity doped diamond layer 12, thereby improving sensitivity and detecting the target material. Maximize the activity of enzymes that oxidize substances. This is believed to be the effect caused because the metal nanoparticle layer 16 more stably fixes the enzyme, and the graphene layer 14 increases the rate of electron transfer generated as a result of the enzyme reaction.

The enzyme reactant layer 18 contains an enzyme that reacts with a sample to be detected. The enzyme to be used may be variously selected according to a sample to be detected, and for example, it may be appropriately used according to a sample to detect glucose oxidase, lactic acid oxidase, alcohol oxidase, cholesterol oxidase and the like. In particular, glucose oxidase may be used when blood glucose or glucose is to be detected.

As such, the biosensor 100 according to the present invention is a high sensitivity biosensor capable of detecting fast response speed and low concentration.

Hereinafter, the present invention will be described in more detail with reference to the drawings and experimental examples for a method of manufacturing such a biosensor and specific experimental evaluation results.

3 is a schematic diagram of a process for making an improved working electrode according to the present invention.

First, the substrate 10 is prepared as shown in FIG. In the present experimental example, the substrate 10 was a silicon wafer thermally oxidized. A silicon wafer having a (100) plane was prepared to form a silicon oxide film (SiO ₂ ) having a thickness of about 1 mm by a thermal oxidation method.

Next, as shown in FIG. 3B, a boron-doped diamond thin film BDD is deposited on the substrate 10 as the impurity-doped diamond layer 12. To this end, first, diamond nanocrystals are coated on the substrate 10 in a high density to be used as a seed for film deposition. Then, a boron doped diamond thin film is grown thereon.

Specifically, it was performed as follows. First, diamond nanocrystals were dispersed in an aqueous solution of 10% (v / v) poly sodium 4-styrene sulfonate (MW: 70,000)] using an attrition mill at 1000 rpm for 5 hours. The diamond nanocrystals used had an average particle size of 3.2 nm.

This yields cationic diamond nanocrystals coated with anionic PSS chains. The substrate 10 prepared as shown in FIG. 3 (a) is immersed in 10% (v / v) PDDA [poly diallyldimethyl ammonium chloride (MW: 400,000-500,000)] aqueous solution for 30 minutes, and then washed with deionized water to obtain a substrate (10). ) A thin cationic polymer film can be coated over. In this process, most of the cationic polymer film is dissolved in the oxide film on the substrate 10 to be in direct electrostatic bonding with the substrate 10. Similarly, diamond nanocrystals coated with anionic PSS polymer adhere to the surface of the cationic substrate 10 at high density.

As described above, after the diamond nanocrystals are coated on the substrate 10 to form a high density seed layer, the doped impurities are grown by growing a boron doped diamond thin film using a hot-filament chemical vapor deposition (HFCVD) method. Formed diamond layer 12 is formed. The flow rates of the B ₂ H ₆ , CH _4, and H ₂ gases that are the source gases were 0.05 sccm, 1 sccm, and 100 sccm, respectively. The deposition was carried out for 5 hours at 90 torr pressure and 800 ℃ temperature conditions. The as-grown thickness of the deposited film was about 1 mm.

Next, the graphene layer 14 is formed on the impurity doped diamond layer 12 as shown in FIG. In the present invention, a method of attaching a sheet-like graphene thin film on the impurity doped diamond layer 12 is adopted.

The sheet-like graphene thin film was obtained by the following method. First, a nickel layer having a thickness of 300 nm or less was formed on a silicon substrate on which a silicon oxide film was formed using an electron-beam evaporator. The nickel layer serves as a catalyst for the precipitation of graphene from the gaseous carbon source.

Then, the substrate temperature was raised to 1000 ° C. in an argon atmosphere quartz tube. Methane (CH ₄ ) was used as the gaseous carbon source. A mixed gas of CH ₄ , H ₂ , and Ar was fed at a ratio of 50 sccm: 65 sccm: 600 sccm. It was then quenched using an argon stream to room temperature (25 ° C.). The cooling rate is about 10 ° C. ⁻¹ . In this process, a graphene thin film was grown on the nickel layer. The graphene thin film was then separated from the silicon substrate. After immersing in 1 M FeCl ₃ solution for several minutes and rinsing with deionized water, the nickel layer melts and the graphene thin film is separated and suspended on the solution. When the sheet-like graphene thin film thus formed is transferred onto the impurity doped diamond layer 12, a relatively strong bond is formed between each other by an electrostatic force.

Next, as shown in FIG. 3D, the metal nanoparticle layer 16 is formed on the graphene layer 14. In the present invention, the substrate 10 formed up to the graphene layer 14 is subjected to an electrochemical method by applying a potential of −0.6 V for 720 seconds in a 3 mL solution containing 20 mM K ₂ PtCl ₆ and 0.5 MH ₂ SO ₄ . Platinum nanoparticles were deposited and deposited on the graphene layer 14.

In this way, the working electrode 20 of the biosensor 100 of FIG. 1 is formed. Next, the reference electrode 30, the counter electrode 40, the conductive wires 25, 35, and 45 and the electrode pads 27, 37, and 47 are formed to form the electrode structure shown in FIG. 1. When the working electrode 20 is immersed in a solution containing glucose oxidase (GOx), the glucose oxidase is directly adsorbed on the surface of the platinum nanoparticles. This forms up to the enzyme reaction material layer 18, followed by rinsing with phosphate buffer (PB) solution to remove residual monomer or weakly bound enzyme molecules. D-glucose and glucose oxidase (GOx; E.C. 1.1.3.4, 200 U / mg) were used in a mixture of 0.05 M phosphate buffer at pH 7.

(Experimental Example 1)

When forming the metal nanoparticle layer 16 on the graphene layer 14 as shown in Figure 3 (d), it was experimented by changing the potential. 4 are SEM images of platinum nanoparticles electrodeposited onto graphene layer 14 at various potentials.

Referring to (a)-(c) of FIG. 4, when the negative voltage is increased to reach -0.6V, the platinum nanoparticles can be found to spread more densely and broadly with increasing potential.

Applying a voltage greater than -0.6V significantly reduced the density, as shown in (d)-(f) of FIG. From these results, it is concluded that when platinum nanoparticles are deposited on the graphene layer 14, it is desirable to apply a potential of -0.6 V in a 3 mL solution containing 20 mM K ₂ PtCl ₆ and 0.5 MH ₂ SO ₄ . Got it.

Experimental Example 2

The biosensor 100 according to the present invention was manufactured by the above-described method, but the counter electrode 40 was made of platinum wire, and the reference electrode 30 was composed of a three-electrode system using SCE. When only the doped diamond electrode is used for comparison with the working electrode 20 of the biosensor 100 according to the present invention composed of doped diamond layer (BDD) / graphene / platinum (Pt) -nanoparticles (comparative example) 1_BDD) and the case where an electrode in which a graphene thin film was stacked on a doped diamond thin film (Comparative Example 2_BDD / graphene) were also tested.

5 shows relative CV response characteristics of three different electrodes of a biosensor according to the present invention and a comparative example. Three different electrodes were measured in 3M KCl solution containing 10 mM Fe (CN) ₆ ^{3- / 4-} .

In Comparative Example 1 using only a doped diamond thin film as an electrode, anionic peaks (current: 0.494 mA, potential: 0.395 V) and cationic peaks (current: -0.664 mA, potential: 0.114 V) were observed. The peak potential difference interval ΔE _p was about 0.281V.

In Comparative Example 2 using an electrode in which a graphene thin film was stacked on a doped diamond thin film, the anionic potential peak E _pa was 0.338 V (current: 0.716 mA), and the cationic potential peak E _pc was 0.177 V. (Current: -0.823 mA), the peak potential difference interval ΔE _p was about 0.161V.

Notably, in the case of the present invention, the oxidation peak current increased to ˜1.08 mA (E _pa 0.332 V) and the peak current decreased to −1.25 mA, ΔE _p was about 0.136 V. The magnitude of the electrochemical response of this electrode was the largest in the case of the present invention, followed by Comparative Example 2 and the smallest.

In the present invention, the peak current is increased because the effective surface area is increased. Well-defined and quasi-reversible redox peaks suggest direct electron transitions between electrodes and redox species. Significantly increased signal in the working electrode 20 of the biosensor 100 according to the present invention composed of a doped diamond layer (BDD) / graphene / platinum (Pt) -nanoparticles is a graphene thin film and platinum nanoparticles. This is because of the synergistic effect. From these results, it became clear that the graphene thin film and platinum nanoparticles accelerated the electron transition to increase the signal.

Experimental Example 3

The effective surface area of each electrode was measured using the Randles-Sevcik equation. According to this equation, the effective surface area A is proportional to I _p / v ^1/2 (Ip: peak current, v: scan rate). The biosensor according to the present invention and the biosensors of Comparative Examples 1 and 2 were measured in a continuous circulation potential with various scan rates in a 3 M KCl solution containing 10 mM K ₃ Fe (CN) ₆ .

6 shows a linear relationship between I _p and v ^1/2 for the electrode according to the invention. This shows that the reaction occurring at the electrode is almost reversible and that the mass transfer phenomenon in the bilayer region is largely controlled by diffusion.

Table 1 summarizes the results.

Compared with Comparative Example 1, Comparative Example 2 increased the surface area by 1.95 times, and in the case of the present invention, increased by 3.53 times.

The interfacial properties of the three electrodes were measured using impedance spectroscopy. In the case of Comparative Example 1, the electron transition resistance was measured as 750 Ω, which was very large compared to the present invention. Accordingly, it can be seen that as in the electrode of the present invention, when graphene and metal nanoparticles are present together, they promote electron transfer.

Experimental Example 4

In order to verify the usefulness of the graphene thin film in a 50 mM phosphate buffer solution while adding 0.6V to the SCE reference electrode for Comparative Example 3 and the present invention, which is the case of the electrode with platinum nanoparticles attached to the diamond thin film without the graphene layer, Prepared glucose samples were measured electrically.

7 is a graph showing the correlation between the change of current and glucose concentration. Referring to FIG. 7A, in Comparative Example 3, the linear section has a slope of 4.94 μA / mM (R = 0.916) and the glucose concentration is from 2.76 μM to 8.9 mM. On the other hand, as shown in Figure 7 (b) in the case of the present invention, the linear range is 1.98 μM to 1.95 mM and the sensitivity is 17.1 μA / mM (R = 0.969). The detection limit is 0.14 μM and the signal-to-noise ratio is 3.00. In general, the active redox center of glucose oxidase is deeply embedded in the protective protein shell, so that direct electron exchange with the electrode is very difficult. Graphene has a very high electrical conductivity and a large specific surface area. Because of these advantages, the graphene used in the working electrode in the present invention can promote electron transfer and enable direct electron transfer between glucose oxidase and the electrode.

As a result of the literature investigation, the sensitivity of the biosensor using the conventional electrodes was about 0.428 to 10 μA / mM, but in the case of the present invention, since it is 17.1 μA / mM, it is evaluated as a significant sensitivity improvement. In addition, the case of the present invention is evaluated as having a wider linear section and a lower detection limit than conventional electrodes.

In particular, in the case of the present invention, K _M ^app (Michaelis-Menten constant) was measured as low as 10.28 mM, which represents a high affinity between glucose and glucose oxidase immobilized on the electrode. Reproducibility was also verified.

While the present invention has been particularly shown and described with reference to exemplary embodiments thereof, it is clearly understood that the same is by way of illustration and example only and is not to be taken by way of limitation in the embodiment in which said invention is directed. It will be understood by those skilled in the art that various changes in form and detail may be made therein without departing from the scope of the appended claims.

Claims (10)

Board;
An electrochemical biosensor comprising a working electrode for generating an enzyme reaction and a counter electrode for detecting a difference in electrical output from the working electrode,
The working electrode includes an impurity doped diamond layer;
A graphene layer formed on the impurity doped diamond layer; And
Biosensor comprising a metal nanoparticle layer formed on the graphene layer. The biosensor of claim 1, further comprising an enzymatic reactive material layer comprising glucose oxidase on the metal nanoparticle layer. The biosensor of claim 1, wherein the metal nanoparticle layer is a platinum nanoparticle layer. The biosensor of claim 1, wherein the graphene layer and the metal nanoparticle layer are graphene sheets on which metal nanoparticles are deposited. Board;
An electrochemical biosensor manufacturing method comprising a working electrode in which an enzyme reaction occurs and a counter electrode for detecting a difference in electrical output from the working electrode,
Forming the working electrode,
Forming an impurity doped diamond layer;
Forming a graphene layer on the impurity doped diamond layer; And
Biosensor manufacturing method comprising the step of forming a metal nanoparticle layer on the graphene layer. The method of claim 5, wherein the forming the impurity doped diamond layer,
Coating the diamond nanocrystal seed on the substrate; And
And depositing a boron doped diamond layer on the substrate using B ₂ H ₆ , CH ₄ and H ₂ gases. The method of claim 6, wherein the coating of the diamond nanocrystalline seed,
Dispersing the diamond nanocrystals in an aqueous solution of poly sodium 4-styrene sulfonate [PSS];
Preparing a silicon substrate thermally oxidized to the substrate; And
And immersing the substrate in an aqueous solution in which the diamond nanocrystals are dispersed, and then washing the substrate. The method of claim 5, wherein the forming of the graphene layer on the impurity doped diamond layer,
Biosensor manufacturing method characterized in that to attach a graphene sheet on the impurity doped diamond layer. The method of claim 5, wherein the forming of the metal nanoparticle layer on the graphene layer,
And impregnating the impurity-doped diamond layer having the graphene layer in an aqueous solution of a target metal to deposit nanoparticles of the metal on the graphene layer by an electrochemical method. The method of claim 5, wherein the forming of the metal nanoparticle layer on the graphene layer,
Depositing platinum nanoparticles on the graphene layer by an electrochemical method in which the impurity-doped diamond layer having the graphene layer formed thereon is immersed in an aqueous solution containing K ₂ PtCl ₆ to apply a potential of −0.6 V. Biosensor manufacturing method.

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