JP2022088296A - High-sensitivity electrochemical biosensor - Google Patents

Abstract

To manufacture a biosensor capable of detecting only change in electrical characteristics due to biological response induced on a surface without influence of change in electrical characteristics of a measurement environment, that is to say, manufacture a biosensor capable of both eliminating background noise and immobilizing biomolecules serving as probes under an optimal condition to detect change in electrical characteristics due to capturing of target molecules with high sensitivity.SOLUTION: Chain-structured molecules in which a hydrophobic region, a functional region, probe molecules are linked in this order are provided on the surface of a substrate. The hydrophobic region is located on the surface side of the substrate, the substrate is a conductive substrate, and the hydrophobic region is composed of a water-insoluble polymer.SELECTED DRAWING: None

Description

The present invention relates to a biosensor that detects changes in electrical signals associated with the capture of biomolecules, and a substrate used thereof.

A biosensor is a device for detecting a specific target molecule in a biological sample by utilizing a specific biochemical reaction for the purpose of detecting a disease or measuring a health condition. Most biomolecules are polyelectrolytes that have dissociative groups such as carboxy groups, amino groups or phosphate groups in their molecules. Therefore, the electrochemical biosensing technique using a field effect transistor for detecting an electric signal is effective in terms of increasing sensitivity, miniaturization, integration, and the like of a biosensor.

In electrochemical biosensing, unexpected redox reactions due to water, lysates, non-target substances, etc. and fluctuations / biases in the measurement environment (ion concentration, etc.) may cause background noise, which may hinder higher sensitivity. .. In recent years, it has become clear that an electrode element having a hydrophobic polymer brush structure on its surface has an almost constant output regardless of the ion concentration in the measurement environment, and is a promising candidate as a reference electrode element that realizes high measurement sensitivity. It is one of. On the other hand, in order to detect the attachment / detachment of a specific biomolecule with high sensitivity, it is necessary to specifically interact between the biomolecule that is the probe immobilized on the sensor surface and the target biomolecule in the biosample. There is.

In order to induce a specific specific interaction in a biological sample containing a large amount of stenosis, all non-specific interactions should be eliminated as completely as possible, and the state of the biomolecule used as a probe, that is, It is necessary to optimize and immobilize the density, higher-order structure, surrounding environment, etc. On the other hand, it is very difficult to control such an advanced surface structure and surface characteristics on the surface of a hydrophobic polymer which is supposed to be used as a reference electrode element. In addition, blocking agents such as skim milk, albumin, and TritonX are often used for the purpose of suppressing non-specific adsorption, but the selection of the blocking agent largely depends on the measurement system, so it is optimal for each measurement. There are problems such as the need to examine the blocking agent, the possibility of peeling due to physical adsorption, and the inability to specify the adsorption density and adsorption film thickness.

Japanese Unexamined Patent Publication No. 2019-045417 International Publication No. 2017/1673715 Pamphlet International Publication No. 2016/035752 Pamphlet Japanese Unexamined Patent Publication No. 2004-226381

An object of the present invention is to create a biosensor that is not affected by changes in electrical characteristics of a measurement environment and can detect only changes in electrical characteristics due to a biological response induced on the surface.

In other words, the present invention is a biosensor that can detect changes in electrical characteristics due to capture of target molecules with high sensitivity by achieving both elimination of background noise and immobilization of biomolecules as probes under optimal conditions. The challenge is to create.

In the present invention, a conductive substrate having a surface modified with a two-layer polymer brush structure consisting of hydrophobic and functional segments can be preferably used. In general, small molecule substances such as water molecules and ions dissolved in the measurement system are considered to induce changes in the electrical characteristics of the measurement environment. Hydrophobic polymer segments are expected to suppress the diffusion of such small molecule substances into conductive substrates. In addition, the functional polymer segment is expected to have effects such as elimination of non-specific interactions, fixation of biomolecules by covalent bonds, and optimization of the environment for immobilization of biomolecules. These can be realized by introducing a functional group according to the desired function. For example, as shown in FIG. 1, such an effect can be realized by having a structure in which the hydrophobic polymer brush layer is arranged on the electrode side and the hydrophilic polymer brush layer is arranged on the bioenvironment side. It is possible.

Therefore, in the present invention, the following aspects can be proposed.
[1] The bio is characterized in that a chain structure molecule in which a hydrophobic region, a functional region, and a probe molecule are linked in this order is provided on the surface of the substrate, and the hydrophobic region is located on the surface side of the substrate. sensor.
[2] The biosensor according to [1], wherein the substrate is a conductive substrate.
[3] The biosensor according to [1] or [2], wherein the hydrophobic region is composed of a water-insoluble polymer.
[4] The biosensor according to any one of [1] to [3], wherein the functional region is composed of a functional polymer.
[5] Any of [1] to [4], wherein the chain structure molecule is arranged in a brush shape on the surface of the conductive substrate, and the density thereof is 0.01 molecule / nm ² to 1 molecule / nm ² . The biosensor described in.
[6] The biosensor according to any one of [1] to [5], which comprises an electrochemical detection element and the substrate is an electrode which is a conductive substrate.
[7] The biosensor according to any one of [1] to [6], wherein the molecule constituting the functional region of the chain structure molecule has a functional group for binding a probe molecule that captures a specific biomolecule. ..
[8] The biosensor according to [7], wherein the functional group that binds the probe molecule that captures a specific biomolecule is a carboxy group formed by dissociation of a tarsier butyl group.
[9] The biosensor according to any one of [1] to [8], wherein the water-insoluble polymer layer composed of the hydrophobic region of the chain structure molecule has a thickness of 1 nm to 100 nm.
[10] The biosensor according to any one of [1] to [9], wherein the hydrophobic region of the chain structure molecule is immobilized on a conductive substrate by a direct and / or indirect covalent bond.
[11] The biosensor according to [10], wherein the hydrophobic region of the chain structure molecule is bonded to the conductive substrate via a polymerization initiating group immobilized on the conductive substrate.
[12] The biosensor according to [11], wherein the water-insoluble polymer layer composed of the hydrophobic region of the chain structure molecule has a thickness of 1 nm to 50 nm.
[13] The biosensor according to any one of [1] to [12], wherein the hydrophobic region of the chain structure molecule comprises a monomer unit having a tertiary butyl group in the side chain.
[14] The biosensor according to any one of [1] to [13], wherein the contact angle between the substrate and pure water is 60 ° or less.
[15] The biosensor according to any one of [1] to [14], wherein the probe molecule is a molecule selected from the group consisting of DNA oligonucleotides, RNA aptamers, and antibodies.

According to the present invention, it is possible to develop an action electrode element in which changes in the measurement environment do not become background noise due to the presence of the hydrophobic polymer segment, and further, by devising the functional polymer segment, various biomolecules by electrochemical biosensing can be developed. High-sensitivity measurement can be realized. Furthermore, this high sensitivity leads to sensing in a small area, which can greatly contribute to the miniaturization and integration of biosensors.

It is a figure which showed one aspect of the biosensor of this invention. It is a figure which showed one aspect of the signal detection mechanism by the biosensor of this invention. It is a figure which shows the chemical structure of the surface of a biosensor produced in Example 1. FIG. It is a high-sensitivity reflection type (RAS) Fourier transform infrared spectroscopy (IRRAS) chart of a substrate before and after acid treatment. (a) Before acid treatment, (b) After acid treatment. It is a figure which showed the relationship between the contact angle of pure water, and the acid treatment time on the surface of a polymer brush two-layer structure. It is a figure which shows the potential difference in the pH standard solution (pH 4.01, 6.86, 9.26) in each electrode in Example 2. FIG. It is a figure which shows the time change of the potential difference with the adsorption of bovine serum albumin (BSA) when the gold electrode modified by the polymer brush layer made with the different polymerization time was used as the working electrode in Example 3. FIG. It is a figure which shows the change of the potential difference with Bcl2 capture in each electrode in Example 4. FIG.

[Chain structure molecule]
In the biosensor of the present invention, the chain structure molecule provided on the surface of the substrate is characterized in that the hydrophobic region, the functional region, and the probe molecule are linked in this order. The hydrophobic region of this chain structure molecule is located on the surface side of the substrate.

The hydrophobic region of this chain structure molecule is preferably composed of a water-insoluble polymer. This hydrophobic region can serve as an impermeable layer, i.e. a hydrophobic polymer brush layer, and contribute to the elimination of background noise in signal detection.

Further, in the biosensor of the present invention, the functional region of the chain structure molecule provided on the surface of the substrate is preferably composed of a functional polymer. In this functional region, it is possible to form a functional layer having a polymer layer composed of a monomer unit having a functional group for chemically immobilizing a biomolecule in a side chain.

In the biosensor of the present invention, if the functional layer formed by the chain structure molecule provided on the surface of the substrate has a structure arranged on the upper part of the impermeable layer, the method of connecting the functional layer is particularly limited. do not have to. That is, as a method for constructing this two-layer structure, after constructing the hydrophobic polymer brush layer, the surface thereof may be appropriately treated, a block type polymer brush structure may be used, or the hydrophobic polymer brush may be used. After building the layer, it may be physically coated with a suitable polymer.

Further, in the biosensor of the present invention, the impermeable layer blocks the movement of small molecule substances, and the functional layer blocks the adsorption of non-specific biopolymers, thereby causing background noise and polymer-derived noise ( It is possible to reduce both non-specific adsorption noise). Further, when the chain structure molecule is covalently bonded to the surface of the substrate, it is possible to prevent its peeling.

Further, the molecule constituting the functional region of the chain structure molecule preferably has a functional group for binding a probe molecule that captures a specific biomolecule, and a probe that captures this specific biomolecule. Typical functional groups that bind the molecule include a carboxyl group or an active ester thereof, an epoxy group, a tosyl group, an amino group, a thiol group, a bromoacetamide group, and the like. The carboxy group may be formed by, for example, dissociation of a tarshally butyl group.

Further, in the biosensor of the present invention, it is preferable that the chain-structured molecules are arranged in a brush shape on the surface of the conductive substrate and the density thereof is 0.01 molecule / nm ^2-1 molecule / nm ² . This makes it possible to contribute to the elimination of background noise in signal detection. Further, by controlling the density of the chain structure molecule, it is possible to set the optimum density of the chain structure molecule for each type of sample.

〔substrate〕
The substrate used in the biosensor of the present invention is preferably a conductive substrate. Examples of the material of the substrate include platinum and gold, but gold is preferable. Further, the substrate used in the biosensor of the present invention is preferably an electrode provided with an electrochemical detection element.

Further, in the substrate used in the biosensor of the present invention, it is preferable that the hydrophobic region of the chain structure molecule is bonded to the substrate via a polymerization initiating group immobilized on the substrate. Further, the water-insoluble polymer layer composed of the hydrophobic region of the chain structure molecule preferably has a thickness of 1 nm to 100 nm, and preferably has a thickness of 1 nm to 50 nm. More preferred.

Further, in the substrate used in the biosensor of the present invention, it is preferable that the hydrophobic region of the chain structure molecule is immobilized on the conductive substrate by a direct and / or indirect covalent bond.

Further, in the substrate used in the biosensor of the present invention, it is preferable that the hydrophobic region of the chain structure molecule is bonded to the conductive substrate via the polymerization initiating group immobilized on the conductive substrate. ..

The hydrophobic region of the chain structure molecule is preferably composed of a monomer unit having an alkyl group, an aromatic ring, a fluorinated alkyl group or the like in the side chain, and is preferably water-insoluble. Further, the contact angle between the substrate and pure water is preferably 60 ° or more.

Further, in the biosensor of the present invention, the type of the probe molecule is not particularly limited as long as it can supplement the target substance, but the probe molecule is a molecule selected from the group consisting of DNA oligonucleotides, RNA aptamers, and antibodies. It is preferable to have.

The contents of the present invention will be described with reference to the following examples, but the scope of the present invention is not construed as being limited to these examples.

(Example 1)
FIG. 3 shows an example of the chemical structure of the biosensor surface produced in the present invention.
First, atom transfer radical polymerization (Cr: 10 nm, Au: 200 nm) and gold sensor substrate for Quartz crystal microbalance with dispersion (QCM-D) measurement using the method shown below. The polymerization initiation group of ATRP) was fixed. Before use, each substrate was UV / O3 treated for ₁₀ minutes to clean the surface. Immediately after the treatment, the silicon substrate was put into a mixed solution of ethanol and water (93% ethanol, containing ammonia at a concentration of 1.4 mol / L) of (2-bromo-2-methyl) propionyloxypropyltrimethoxysilane adjusted to a concentration of 20 mmol / L. The gold substrate was immersed in an ethanol solution of α-bromobutyrate-11-undecanethiol adjusted to a concentration of 3.0 mmol / L, respectively. The substrate was allowed to stand overnight at room temperature, the substrate was removed from the solution, washed with ethanol, and then dried with a nitrogen blow to recover. Next, the poly (tert-butylryl (tBMA)) (PtBMA) brush structure shown in the upper part of FIG. 3 was constructed on the surface of each substrate by using the surface-initiated ATRP (SI-ATRP) method. In a eggplant-shaped flask, tBMA, copper (I) chloride (CuCl), 4,4'-dinonyl-2,2'-dipyridyl (dN-Bpy), ethyl 2-bromoisobutyrate (EBIB), tBMA / EBIB / CuCl / It was weighed at a ratio of dN-Bpy = 2000 / 1.0 / 2.5 / 5.0 (by molar fraction), dissolved while bubbling with argon, added to the polymer tube on which the prepared initiating group-immobilized substrate was installed, and sealed. Polymerization was promoted by shaking the polymerization tube at a speed of about 150 rpm in an oil bath maintained at 70 ° C. After a lapse of a predetermined time, the polymerization tube was quickly cooled to room temperature to stop the polymerization, and the substrate was taken out from the polymerization solution. The removed substrate was thoroughly washed with tetrahydrofuran (THF), dried with a nitrogen blow, and recovered. When necessary, the polymerization time is described after the abbreviation, such as PtBMA-1h. Monomers are obtained by diluting the recovered polymerization solution with THF to prepare a solution of known concentration and using this solution for Gel permeation chromatography (GPC) measurements to obtain information on the molecular weight of the graft polymer chains on the substrate. Information on conversion and molecular weight was obtained. In GPC measurement, THF was used as the eluent, KF-404 HQ was used as the column, the flow rate was 0.3 mL / min, and the standard sample was poly (methyl acrylic). As a result of measuring the film thickness of the polymer brush layer at the time of drying by spectroscopic ellipsometry, the film thickness of the polymer brush layer could be controlled in the range of 10 to 40 nm depending on the polymerization time. The graft density was 0.36 chains / nm ² . The molecular weight distribution of the polymer synthesized in the solution decreased from 1.50 to 1.15 with increasing polymerization time.
The PtBMA brush substrate was brought into contact with concentrated hydrochloric acid water at room temperature to modify the surface. In the IRRAS chart of the PtBMA brush substrate before and after the acid treatment for 120 minutes shown in FIG. 4, absorption peaks of 1368, 1393 and 2979 cm ^-1 derived from the methyl group were detected on both substrates, but on the hydroxy group. The resulting wide absorption peak around 3200 cm ^-1 was detected only on the acid-treated substrate. From this result, it was found that the hydroxy group was generated by the acid treatment. That is, a substrate (PtBMA-COOH substrate) in which the tBMA unit near the surface was changed to a methacrylic acid unit (MA) could be produced.
From FIG. 5, which shows the relationship between the contact angle of pure water with respect to the PtBMA-COOH brush substrate and the acid treatment time, it was found that the contact angle of pure water gradually decreased with the lapse of the acid treatment time. From FIGS. 4 and 5, it was shown that a polymer brush two-layer structure substrate in which the hydrophobic polymer brush layer was arranged on the substrate side and the hydrophilic polymer brush layer was arranged on the surface side could be produced.

(Example 2)
FIG. 6 shows the potential difference between solutions of various pH when a gold electrode having different polymer brush layers is used as a working electrode. Specifically, an Ag / AgCl reference electrode was placed in the solution, and the potential difference between the working electrode and the reference electrode was measured using a source / measure unit.
If there is a pH-responsive functional group, there is inherently a linear relationship between pH and the potential difference.
・ Gold electrode modified with PtBMA brush layer → pH-independent potential difference compared to Bare Au ・ Gold electrode modified with PtBMA-COOH brush layer → Linear relationship between pH and potential difference ・ Not yet Modification, COOH-SAM modification or bovine serum albumin (BSA) coated gold electrode → No correlation between pH and potential difference From the above, pH responsiveness to the surface whose charge state is constant by the hydrophobic polymer brush layer. It was clarified that the function as a pH sensor was imparted due to the introduction of the carboxy group of.
It is considered that the background noise due to the change in pH, which is one of the measurement environments, could be reduced by introducing the PtBMA brush layer as an impermeable layer.
BSA is a protein often used as a blocking agent in biosensing studies. The results of the BSA-coated gold electrodes showed that blocking with BSA could not reduce background noise. The superiority of background noise reduction by introducing an impermeable layer was shown.

(Example 3)
FIG. 7 shows the time change of the potential difference due to the adsorption of bovine serum albumin (BSA) when a gold electrode modified with a polymer brush layer prepared at different polymerization times is used as a working electrode. A source / measure unit was used to measure the potential difference, and the potential difference between the Ag / AgCl reference electrode and the working electrode placed in the solution was measured in real time. As a specific method of measurement, a phosphate buffer solution was contacted in advance, and a BSA solution was added thereto so that the final concentration was 1.0 mg / mL.
As shown in FIG. 7, even on the electrode surface covered with the hydrophobic polymer brush layer, the change in the potential difference due to the adsorption of the protein could be detected as in the case of the unmodified gold electrode.
Therefore, it is shown that it is possible to detect a bioreaction on this electrode.
Further, from FIG. 7, when the film thickness of the PtBMA brush layer was large, the change value of the potential difference became small. The characteristics of the outermost surface do not depend on the film thickness.
That is, in the polymer brush layer having a large film thickness, the adsorption surface of BSA became far from the gold surface, and the sensitivity decreased.
Considering this together with the results shown in FIG. 6, it was found that in the range implemented this time, both the impermeable effect and the bioreaction detection can be achieved on the surface having a film thickness of about 15 nm.

(Example 4)
FIG. 8 shows the time change of the potential difference when Bcl2 is brought into contact with the PtBMA-COOH brush layer on which the probe DNA is immobilized as the target DNA.
As probe DNA, 30 mer single-stranded DNA (5'-NH2-ATG AAG TAA GAG GAC AGG CAC CAC AGC CCC-3') with amino group modification at the 5'end is used as target DNA, 87 mer containing Bcl2 region. Single-stranded DNA (Bcl2, 5'-GCG GGC GGG CGG GCA GGC GGC GCG GAG GGG CGG GCG CGG GAG GAA GGG GGC GGG AGC GGG GCT GTG GTG CCT GTC CTC TTA CTT CAT -3') was used. Here, the underlined portion in the first half is the Bcl2 region, and the underlined portion in the second half is the complementary region with the probe DNA. The PtBMA-COOH brush layer was prepared under the conditions that the polymerization time for preparing the PtBMA brush layer was 1 hour and the concentrated hydrochloric acid water treatment for introducing the carboxy group was 60 minutes (thickness: 15 nm). .. Surface carboxy groups and amino groups of probe DNA are used to immobilize the probe DNA on the substrate using 1-ethyl-3- (3-dimethylamino propyl) carbodiimide, hydrochloride (WSC) and N-hydroxysuccinimide (NHS). The condensation reaction of The immobilization density of the probe DNA was quantified to be about 1.0 x 10 ^-2 DNA / nm ² by the energy-dissipating quartz crystal microbalance (QCM-D). As a negative control, COOH-SAM in which the above-mentioned probe DNA was pre-immobilized was used. In the experiment using COOH-SAM, DNA without complementary sequence was also used as the target DNA.
From FIG. 8, when Bcl2 came into contact with the probe DNA-immobilized COOH-SAM modified electrode, a negative potential difference change was detected (21 mV). On the other hand, when the probe DNA and the DNA having no complementary sequence were brought into contact with the same electrode, there was almost no shift in the potential difference (2 mV). From these results, it was found that the change in potential difference when Bcl2 was brought into contact was due to the fact that the negatively charged Bcl2 was captured on the electrode surface by hybridization with the immobilized probe DNA. Even in the electrode modified with the PtBMA-COOH brush layer on which the probe DNA was immobilized, a negative potential difference change of about 6 mV was detected, so the probe DNA immobilized on the electrode and Bcl2 in the solution were high. It was found that hybridization had progressed.
From the above, it was shown that the prepared PtBMA-COOH brush layer-modified gold electrode can detect the immobilization ability of probe molecules and specific biological reactions.
It was also shown that the PtBMA-COOH brush layer modified gold electrode can be used as an electrochemical biosensor.

The present invention can be expected to be used as a detection unit of a biosensing device to which a field effect transistor is applied. In particular, it is considered to be effective for sensing devices that require high sensitivity and array.
From the above, the present invention is extremely useful in industry.

Claims (15)

A biosensor comprising a chain structure molecule in which hydrophobic regions, functional regions, and probe molecules are linked in this order on the surface of a substrate, and the hydrophobic regions are located on the surface side of the substrate. The biosensor according to claim 1, wherein the substrate is a conductive substrate. The biosensor according to claim 1 or 2, wherein the hydrophobic region is composed of a water-insoluble polymer. The biosensor according to any one of claims 1 to 3, wherein the functional region is composed of a functional polymer. The invention according to any one of claims 1 to 4, wherein the chain structure molecule is arranged in a brush shape on the surface of the conductive substrate, and the density thereof is 0.01 molecule / nm ² to 1 molecule / nm ² . Biosensor. The biosensor according to any one of claims 1 to 5, wherein the biosensor comprises an electrochemical detection element and the substrate is an electrode which is a conductive substrate. The biosensor according to any one of claims 1 to 6, wherein the molecule constituting the functional region of the chain structure molecule has a functional group that binds a probe molecule that captures a specific biomolecule. The biosensor according to claim 7, wherein the functional group that binds the probe molecule that captures a specific biomolecule is a carboxy group formed by dissociation of a tarsier butyl group. The biosensor according to any one of claims 1 to 8, wherein the water-insoluble polymer layer composed of the hydrophobic region of the chain structure molecule has a thickness of 1 nm to 100 nm. The biosensor according to any one of claims 1 to 9, wherein the hydrophobic region of the chain structure molecule is immobilized on a conductive substrate by a direct and / or indirect covalent bond. The biosensor according to claim 10, wherein the hydrophobic region of the chain structure molecule is bonded to the conductive substrate via a polymerization initiating group immobilized on the conductive substrate. The biosensor according to claim 11, wherein the water-insoluble polymer layer composed of the hydrophobic region of the chain structure molecule has a thickness of 1 nm to 50 nm. The biosensor according to any one of claims 1 to 12, wherein the hydrophobic region of the chain structure molecule comprises a monomer unit having a tertiary butyl group in the side chain. The biosensor according to any one of claims 1 to 13, wherein the contact angle between the substrate and pure water is 60 ° or less. The biosensor according to any one of claims 1 to 14, wherein the probe molecule is a molecule selected from the group consisting of DNA oligonucleotides, RNA aptamers, and antibodies.

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