JP6824527B2 - Devices for virus and bacteria detection biosensors, and biosensors - Google Patents

Description

The present invention relates to devices for virus and bacterium detection biosensors, and biosensors.

Infectious diseases are a social problem, and overcoming them requires prompt judgment as to whether or not they have been infected.
As a method of quickly monitoring whether or not a person is infected with a virus or a bacterium, there is a method using a biosensor. Biosensors are required to have higher sensitivity and the ability to easily monitor many types of viruses or bacteria.
Regarding the influenza virus monitoring technique, for example, a method of indirectly detecting an influenza virus by monitoring the health condition of an animal (Japanese Patent Laid-Open No. 2012-242250) has been reported. Further, as a method for detecting by means for capturing influenza virus, a detection method using an antibody against influenza virus (Japanese Patent Laid-Open No. 2014-038115, Japanese Patent Application Laid-Open No. 2007-261988, Japanese Patent Application Laid-Open No. 2014-095720, Japanese Patent Application Laid-Open No. 2013-174607, Bioses. Bioelectron. 2015, 65, 211-219) has also been reported.
On the other hand, as a means for detecting hemagglutinin of influenza virus, a detection method using a device in which an oligosaccharide is immobilized on a gate insulating film of a field-effect transistor (Anal. Chem. 2013, 85, 5641-5644) has also been reported. There is.

The use of oligosaccharides as a means for detecting viruses or bacteria has the advantage that it is expected that the target virus or bacterium can be selectively detected, and that oligosaccharides can be supplied at a lower cost than antibodies. It is also desired to develop a sensor capable of detecting viruses and bacteria other than the above-mentioned influenza.

Therefore, an object of the present invention is to provide a biosensor device and a biosensor that selectively and highly sensitively detect a virus or a bacterium.

Specific means for solving the problem include the following forms.
<1> A device for a biosensor having a conductive layer, an oligosaccharide that selectively captures a specific virus or bacterium, and an intermediate layer that binds the conductive layer and the oligosaccharide.
<2> The biosensor device according to <1>, wherein the intermediate layer is a self-assembled monolayer.
<3> The biosensor device according to <2>, wherein the self-assembled monolayer is formed of a thiol compound immobilized on the conductive layer with a thiol group.
<4> The biosensor device according to <3>, wherein the self-assembled monolayer contains a thiol compound having an oxylamino group or a hydrazide group.
<5> The biosensor device according to <1>, wherein the intermediate layer contains a conductive polymer.
<6> A group in which the conductive polymer is composed of a polythiophene-based conductive polymer, a polyphenylene-based conductive polymer, a polypyrrole-based conductive polymer, a polyaniline-based conductive polymer, a polyfluorene-based conductive polymer, and a polyindole-based conductive polymer. The biosensor device according to <5>, which is at least one selected from the above.
<7> The biosensor device according to <5>, wherein the conductive polymer is a polythiophene-based conductive polymer.
<8> The device for a biosensor according to any one of <5> to <7>, wherein the conductive polymer has an oxylamino group or a hydrazide group.
<9> The conductive polymer is selected from the group consisting of (A) at least one of an oxylamino group or a hydrazide group and (B) a thienyl group, a phenyl group, a pyrrol group, an aniryl group, a fluorene group and an indol group. The device for a biosensor according to <6>, which comprises a structural unit formed from a monomer containing at least one of the above.
<10> The biosensor device according to any one of <5> to <9>, wherein the conductive polymer is formed by an electrolytic polymerization method or an oxidative polymerization method.
<11> The biosensor device according to any one of <1> to <10>, wherein the oligosaccharide is an oligosaccharide that selectively captures a specific influenza virus.
<12> The biosensor device according to any one of <1> to <11>, wherein the conductive layer is a gate electrode of a field effect transistor or an electrode electrically connected to the gate electrode.
<13> A biosensor including the biosensor device according to any one of <1> to <12>.
<14> A plurality of biosensor devices according to any one of <1> to <13> are provided, and the oligosaccharides of the biosensor devices selectively select specific viruses or bacteria different from each other. The biosensor according to <13>, which is an oligosaccharide to be captured.
<15> The biosensor according to <13> or <14>, wherein the oligosaccharide of at least one biosensor device contains a sialic acid residue.
<16> The biosensor according to any one of <13> to <15>, which detects a virus or a bacterium based on any one change selected from an electric potential, a frequency, and a refractive index.

According to the present invention, it is possible to provide a device for a biosensor and a biosensor that selectively and highly sensitively detect a virus or a bacterium.

It is a schematic diagram which shows the structural example of the biosensor device and FET biosensor which concerns on 1st Embodiment of this invention. (A) is a schematic diagram showing a configuration example of an electrode on which an oligosaccharide-bound SAM is immobilized. It is a schematic diagram which shows the structural example of the biosensor device and FET biosensor which concerns on 1st Embodiment of this invention. (B) is a side schematic view showing a configuration example of a FET biosensor provided with a floating gate type electrode. It is a typical perspective view which shows the structural example of the biosensor which concerns on 1st Embodiment of this invention, and is the typical perspective view which shows the structural example of the FET biosensor provided with the extension gate type electrode. (A) is a schematic plane and an electric circuit diagram of a FET biosensor including a biosensor device according to the first embodiment of the present invention and an FET having an operational amplifier. (B) is an electric circuit diagram of an FET biosensor in which the device for a biosensor according to the first embodiment of the present invention is combined with an FET having an operational amplifier, a differential amplifier, an amplifier circuit and a phase compensation capacitor. It is a schematic diagram which shows the structural example of the QCM biosensor which concerns on 2nd Embodiment of this invention. It is a schematic diagram which shows the structural example of the SPR biosensor which concerns on 3rd Embodiment of this invention. It is a schematic diagram which showed an example of the manufacturing method of the device for a biosensor of this invention. It is a graph which shows the potential with respect to the measurement time of the influenza virus measured by the FET biosensor provided with the reference electrode. (A) is a graph showing the surface potential of influenza A virus measured by the FET biosensor equipped with the biosensor device having α2,3-sialyllactose used in Control 1 with respect to the measurement time. (B) is a graph showing the potential of influenza A virus measured by the FET biosensor equipped with the biosensor device having α2,6-sialyllactose used in Example 1 with respect to the measurement time. 3 is a graph showing the difference between the potentials obtained in Examples and Controls and the potentials obtained by measurement with a FET biosensor equipped with a reference electrode with respect to the concentration of influenza A virus, respectively. (A) It is a graph when the initial potential in Example 1 and control 1 is high. 3 is a graph showing the difference between the potentials obtained in Examples and Controls and the potentials obtained by measurement with a FET biosensor equipped with a reference electrode with respect to the concentration of influenza A virus, respectively. (B) It is a graph when the initial potential in Example 1 and control 1 is small. It is a figure which shows the change of the difference (ΔFrequency (Hz)) of the resonance frequency with respect to the measurement time of the influenza A virus measured by the QCM biosensor. It is a figure which shows the change of the capacitance and the change of the film thickness of the electrolytic polymerization process of the EDOTAO monomer at each ratio. (A)-(e) Cyclic voltanmogram: 10 mM monomer (EDOTAO relative to EDOT is 0 mol% (a), 25 mol% (b), 50 mol% (c), 75 mol% (d), 100 mol% (e). )) Aqueous solution, dopand is 100 mM NaClO ₄ . Counter electrode, Pt; Reference electrode is Ag / AgCl (in 3.3M KCl aq); Scan rate, 0.1V s ^-1 . (F): Change in capacitance with respect to change in cycle in the electrolytic polymerization process at each EDOTAO ratio. (G)-(l): EDOTAO monomer 0 mol% (g), 25 mol% (h), 50 mol% (i), 75 mol% (j) and 100 mol% (l) with respect to EDOT are electrolyzed on the Au electrode surface for 10 cycles. Change in film thickness during polymerization, dopanda is 100 mM NaClO ₄ . The displayed data is mean ± standard error; n = 3. It is a figure which shows the XPS high resolution spectrum. (A) N1s and (b) S2p: Monomers having 0 mol%, 25 mol%, and 50 mol% of EDOTAO with respect to EDOT were electropolymerized on the surface of the Au electrode, respectively. The dopant is 100 mM NaClO ₄ . Relationship between the supply amount of EDOTAO, the theoretical introduction rate of EDOTAO, and the actual introduction amount (c). The displayed data is mean ± standard error; n = 4. It is a figure which shows the low impedance surface of the conductive copolymer on the glass electrode. (A) Nyquist plot of conductive copolymer film, inset is Nyquist plot in high frequency domain. The measurement solution is a 5 mM ferricyanide / ferricyanide buffer solution (0.01 DPBS, pH 7.4), counter electrode, Pt; the reference electrode is Ag / AgCl in (in 3.3 M KCl aq); DC bias, 0.18 V; AC amplification, 50 mV; AC frequency domain, 10 kHz to 0.1 Hz. (B) The change in CPE (constant phase element) charge transfer resistance (Rct) with respect to the supply amount of EDOTAO was shown. Fitting was performed using the equivalent circuit of Randle's (insertion diagram). The displayed data is mean ± standard error; n = 3. It is a figure which shows the evaluation of hydrophilicity by a water drop contact angle meter. (A): Contact angle of the conductive polymer before the introduction of the sugar chain. It is a figure which shows the evaluation of hydrophilicity by a water drop contact angle meter. (B): Contact angle of the conductive polymer after the introduction of the sugar chain. It is a figure which shows the evaluation of hydrophilicity by a water drop contact angle meter. (C): Change in weight of the QCM sensor surface before and after the introduction of the sugar chain. The displayed data is mean ± standard error; n = 3. It is a figure which shows the specific detection of influenza virus (H1N1) by a QCM sensor. In (a), changes in weight derived from the molecular recognition reaction of 2,6-sialyllactose sugar chains or 2,3- (sugar chain-modified electrodes, respectively, at various influenza virus concentrations are expressed by frequency changes. It is a change with time when shown. The displayed data is the average value ± standard error; n = 3. It is a figure which shows the specific detection of influenza virus (H1N1) by a QCM sensor. (B) is the frequency change when each injection is in equilibrium in (a) (* p <0.01). The displayed data is mean ± standard error; n = 3. It is a figure which shows the specific detection of influenza virus (H1N1) by potential measurement. The concentration of influenza virus and the potential change derived from the molecular recognition reaction of the electrode modified with α2,6-sialyllactose sugar chain or α2,3-sialyllactose sugar chain were displayed. The displayed data is mean ± standard error; n = 3.

Hereinafter, the present invention will be specifically described. In the following description, "~" representing a numerical range represents a range including the numerical values described as the upper limit value and the lower limit value.

≪Device for biosensor≫
The device for a biosensor of the present invention has a conductive layer, an oligosaccharide that selectively captures a specific virus or bacterium, and an intermediate layer that binds the conductive layer and the oligosaccharide. Each configuration of the device of the present invention will be described in detail.
In addition, in this specification, a "biosensor" means an apparatus used for detecting a virus or a bacterium.

(Conductive layer)
The "conductive layer" refers to a layer having conductivity in which compounds constituting the intermediate layer are immobilized, and the conductive layer is bonded to oligosaccharides via a material forming the intermediate layer. The material forming the intermediate layer may be referred to as "intermediate layer forming material". The intermediate layer forming material may also have conductivity.
The conductive layer is formed of a conductive material, and as long as the above-mentioned intermediate layer forming material can be immobilized, the type and shape of the material are not particularly limited, and the virus or bacterium using the device of the present invention can be used. In consideration of the detection method and the like, it is preferable to use a metal and other conductive materials in a desired shape as appropriate.

Such a conductive layer may be a simple substance of a metal, an alloy, or a material other than a metal such as glass carbon, a conductive polymer, or a conductive inorganic compound.
For example, in the case of metal, gold, silver, platinum, tantalum, aluminum, titanium, magnesium, zinc, silicon, tin and the like can be mentioned. Of these, a metal having a good affinity with the intermediate layer forming material is preferable. When, for example, alkanethiol is used as the intermediate layer forming material, it is preferably a metal having a good affinity for sulfur atoms, specifically, silver, platinum, or zinc, preferably gold. More preferably.
For example, in the case of conductive polymers, polythiophene-based conductive polymer, polyphenylene-based conductive polymer, polypyrrole-based conductive polymer, polyaniline-based conductive polymer, polyfluorene-based conductive polymer, polyindole-based conductive polymer and carbon-based Materials and the like can be mentioned. Specific examples thereof include polyethylene dioxythiophene, poly (p-phenylene), poly (p-phenylene vinylene), poly (p-phenylene sulfide), polythienylene vinylene, eumelanin, glass carbon, graphene and the like. Of these, a conductive polymer having a good affinity with the intermediate layer forming material is preferable. When, for example, a monomer for introducing an oligosaccharide is used as the intermediate layer forming material, it is preferable that it is a conductive polymer that is easily electrochemically adhered by electropolymerization or the like. It is preferably eumelanin or glassy carbon, more preferably polyethylenedioxythiophene.
For example, in the case of a conductive inorganic compound, indium tin oxide (ITO) and indium-gallium-zinc-oxygen (IGZO) can be mentioned.
Further, the shape of the conductive layer is not particularly limited as long as it is a thin film, and the shape in the top view may be circular, elliptical, rectangular, triangular, or polygonal. , May be irregular.

The conductive layer may have a part or all of the surface formed of at least one of a metal salt and a metal oxide.
When the device of the present invention is applied and a biosensor using a field effect transistor (hereinafter referred to as FET) described later is used as the biosensor, a part or all of the surface of the conductive layer is a metal salt. And at least one of the metal oxides is formed, and the coating portion is in contact with an aqueous solution containing a virus or a bacterium, so that the stability of the potential at the solid-liquid interface between the aqueous solution and the conductive layer can be improved.

As the metal salt, for example, the same material as the metal oxide described in JP2013-246060 can be used. Further, among these, since the aqueous solution containing a virus or a bacterium contains a large amount of chloride ions, the metal salt is preferably a metal chloride. Further, the metal constituting the metal salt preferably has a good affinity with an intermediate layer forming material, particularly a self-assembled monolayer (hereinafter referred to as "SAM"). When, for example, alcanthiol is used as the intermediate layer forming material, it is preferable that the metal constituting the metal salt is a metal having a good affinity with a sulfur atom, specifically, gold, silver, platinum, etc. Alternatively, it is preferably zinc, more preferably gold.

As the metal oxide, for example, the same material as the metal material described in JP2013-246060 can be used, and the preferred one is also the same.

As the conductive layer, glass carbon, ITO, and IGZO are also preferable. When a polythiophene-based conductive polymer is used as the intermediate layer forming material, for example, a roll-to-roll printing method or an electric field polymerization method can be applied, so that a thin film can be formed easily, inexpensively, and in large quantities. Can be done.

When the conductive layer has a flat plate shape, its thickness is preferably adjusted as appropriate, but is preferably 0.05 μm to 1 μm. When it is 0.05 μm or more, the conductivity as a device can be more sufficiently secured, and the strength as a device can be more sufficiently secured. Further, when it is 1 μm or less, the manufacturing cost can be suppressed.

Further, the conductive layer can be an electrode of the FET described later, but in this case, the conductive layer is preferably a gate electrode of the FET or an electrode connected to the gate electrode from the viewpoint of detection sensitivity of viruses or bacteria.

(Middle layer)
The intermediate layer forming material for forming the intermediate layer may be composed of any material as long as it is bonded to both the conductive layer and the oligosaccharide. As the intermediate layer forming material, the same substance as the substance forming the conductive layer may be used. Further, the intermediate layer may contain the same substance as the conductive layer in a part thereof.
SAM is preferable from the viewpoint of improving detection sensitivity and easy production. The SAM is a thin film having a structure having a certain order formed by the mechanism of the film material itself without applying fine control from the outside. This self-organization forms ordered structures and patterns in non-equilibrium situations. Sulfur-containing compounds such as alkanethiols and alcandisulfides are spontaneously adsorbed on substrates containing precious metals such as precious metals such as gold and silver, metal salts composed of precious metals, and oxides composed of precious metals. Gives a single molecule size thin film. Therefore, an intermediate layer can be easily formed on the electrode surface, and oligosaccharides can be immobilized at a high density through the intermediate layer, and the accuracy of virus or bacterium detection can be improved. ..

Examples of the material capable of forming the SAM include sulfur-containing compounds and silyl group-containing compounds.
Examples of the sulfur-containing compound include a thiol group-containing compound represented by an alkane thiol having an alkyl chain and a disulfide group-containing compound. The disulfide group-containing compound is a raw material for forming a thiol group.
Examples of the silyl group-containing compound include alkylalkoxysilanes having an alkyl chain.
The sulfur-containing compound and the silyl group-containing compound may be a sulfur-containing compound or a silyl group-containing compound having a polymer chain typified by polyethylene glycol (PEG) instead of the alkyl chain.
Among the above compounds, from the viewpoint of sensitivity as a biosensor, SAM is preferably a sulfur-containing compound immobilized on a conductive layer with a thiol group. Further, as the sulfur-containing compound, a thiol group-containing compound is more preferable, and a thiol group-containing compound having an alkyl chain is particularly preferable.

Further, the compound forming the SAM preferably has at least two or more different functional groups from the viewpoint of binding to both the conductive layer and the oligosaccharide. Further, from the viewpoint of easy bonding to the conductive layer, the compound forming the SAM is preferably a thiol group-containing compound and preferably has a functional group other than the thiol group. Having two or more different functional groups in this way is effective because it facilitates the formation of an oligosaccharide-bound SAM on the conductive layer.

Examples of the functional group include an oxylamino group, a hydrazide group, an amino group, a hydroxyl group, a carboxyl group, a carbonyl group, an azido group, an alkynyl group, an epoxy group and an isocyanate group, in addition to the above-mentioned thiol group. Further, the functional group other than the thiol group is preferably an oxylamino terminal or a hydrazide terminal because of the binding property of the oligosaccharide to the reducing terminal carbon. Since the end of the molecule forming SAM (the end on the side that binds to oligosaccharide) is an oxylamino group or a hydrazide end, it is necessary to impart a functional group for binding to SAM or the like on the oligosaccharide side. Therefore, it is also advantageous that the oligosaccharide can be used as it is for binding to SAM or the like.

Further, among the compounds used when forming SAM, for example, the above-mentioned thiol group-containing compound or disulfide group-containing compound may be one kind of compound having a functional group other than the thiol group or a functional group other than the disulfide group. Alternatively, it may be a plurality of types of compounds having a functional group other than the thiol group or a functional group other than the disulfide group.

When forming SAM, for example, the density of the number of molecules immobilized on the surface of the metal electrode (SAM density), the content of functional groups that bind oligosaccharides in the above compound, or the reaction conditions for binding oligosaccharides. By adjusting, the amount of oligosaccharides bound in a certain area on the SAM, that is, the distance between the oligosaccharides bound to the SAM can be adjusted. As a result, the distance between oligosaccharides on the SAM so as to be compatible with multiple recognition regions (oligosaccharide recognition sites) for oligosaccharides such as lectins possessed by bacteria and viruses or oligosaccharide recognition sites such as multiple lectins. Can be adjusted. Therefore, the biosensor device of the present invention dramatically improves the detection sensitivity when used as a biosensor by forming a multipoint bond between a lectin or the like possessed by a virus or a bacterium and a plurality of oligosaccharides on the device. It is possible to make it.

The SAM density (mol / nm ² ) can be determined by cyclic voltamentary (CV) measurement or the like.

Further, the compound having a disulfide group and further having a functional group other than the disulfide group dissociates the disulfide bond on the metal electrode which is the conductive layer, and two portions via the disulfide bond in the compound ( A and B of ASSB) can be fixed on a metal electrode by a thiol bond. As a result, two kinds of compounds having different functional groups can be arranged in the vicinity of each other on the metal electrode. When such a compound forms a SAM, it has a certain thickness (thickness of a single molecule) on the metal electrode, and two types of immobilization using each functional group are possible. For example, a molecule that captures a target and a molecule that suppresses non-specific adsorption can be immobilized on the same surface of the SAM. Therefore, such a SAM can be used to manufacture a biosensor device having high target recognition ability.

When the intermediate layer is formed on the conductive layer, for example, a solution in which the compound forming the SAM, which is the intermediate layer forming material, is dissolved is applied onto the conductive layer, and the solution dries until the SAM is formed. It may be held so as not to be formed, and washed after forming the SAM.
The oligosaccharide may be bonded to the intermediate layer forming material after forming the intermediate layer as described above. Further, after the oligosaccharide and the intermediate layer forming material are bonded in advance, this may be dissolved in a solvent and applied as a solution on the conductive layer.

As a method of applying the solution on the conductive layer, a method of applying the solution by application such as spin coating, air knife coating, bar coating, blade coating, slide coating, curtain coating, and further, a dropping method, a spray method, and a vapor deposition method. , Casting method, dipping method and the like. The intermediate layer forming material or the solution in which the intermediate layer forming material to which the oligosaccharide is bound is dissolved is randomly applied on a metal electrode or a conductive layer composed of a metal salt, a metal oxide, or the like. It may be given by arranging it in an array, or it may be given as a block in a domain shape. Further, it may be given in a pattern.

Water, alcohol, or a mixed solvent thereof is used as the solvent for dissolving the intermediate layer forming material or the intermediate layer forming material to which the oligosaccharide is bound. A water-insoluble organic solvent may be used to dissolve only the intermediate layer-forming material, but water, alcohol, dimethylformamide (DMF), and dimethyl sulfoxide are used to dissolve the intermediate layer-forming material to which oligosaccharides are bound. A water-soluble solvent such as oxide (DMSO), acetonitrile or acetone, or a mixed solvent thereof is used. Of these, water, ethanol, DMSO, or a mixed solvent thereof is preferable.

As the intermediate layer forming material, a conductive polymer is also preferable. The conductive polymer is not particularly limited as long as it can introduce oligosaccharides and can bind to the conductive layer to form an intermediate layer. Examples of the conductive polymer include polythiophene-based conductive polymer, polyphenylene-based conductive polymer, polypyrrole-based conductive polymer, polyaniline-based conductive polymer, polyfluorene-based conductive polymer, and polyindole-based conductive polymer. Specific examples thereof include polyethylene dioxythiophene, poly (p-phenylene), poly (p-phenylene vinylene), poly (p-phenylene sulfide), polythienylene vinylene, graphene and the like. The conductive polymer may be a homopolymer or a copolymer formed from two or more monomers. Further, the conductive polymer may contain additives used for enhancing conductivity, such as ethylene glycol (EG), dimethylformamide (DMF), and sulfolane, or additives used for other purposes. The conductive polymer used as the intermediate layer may be synthesized by a known method, or a commercially available product may be used. From the viewpoint of bondability with conductivity, it is preferable to form the conductive layer on the conductive layer by electrolytic polymerization or oxidative polymerization using an appropriate monomer.

Among the conductive polymers, polyethylenedioxythiophene, which is one of the polythiophene-based conductive polymers, has high electrical conductivity, is chemically stable, has hole injectability and doping properties, and is a rare element or. It is preferably used as an intermediate layer forming material from the viewpoint of not containing toxic elements and having high flexibility and elasticity. Further, since the polythiophene-based conductive polymer can be used as a dispersion liquid of water or an organic solvent, it is suitable for coating and processing. For example, it can be incorporated into a fiber to make a clothing-type wearable sensor, and the versatility of the biosensor can be enhanced.

The monomer used when forming the conductive polymer as an intermediate layer is soluble in water or an organic solvent, has compatibility with an inorganic anion or a polyanion as a dopant, and is electrolyzed by water or an organic solvent. It is preferable that the monomer has a low oxidation-reduction potential that does not cause decomposition. Examples of such a monomer include a thiophene-based monomer.
Examples of the thiophene-based monomer include 3,4-ethylenedioxythiophene, 3,4-dimethyloxythiophene, and 5,5''''-bis (2'''''', 2''''''-dicyanovinyl). -2,2': 5', 2'': 5'', 2''': 5''', 2''''-quinquethiophene (DCV5T), 5,5''''-bis (trideca) Fluorohexyl) -2,2': 5', 2'': 5'', 2'''-quaterthiophene, 2,5-bis (trimethylstanyl) -thieno [3,2-b] thiophene, 2,5-bis (trimethylstanyl) thiophene, 2,2'-bitieno [3,2-b] thiophene, 2,3'-bithiophene, 2-bromo-5-dodecylthiophene, 2-bromo-3- ( 2-Ethylhexyl) thiophene, 5-bromo-5'-hexyl-2,2'-bithiophene, 3,4-diaminothiophene dihydrochloride, 3,3'-dibromo-5,5'-bis (trimethylsilyl) -2, 2'-Bithiophene, 2,5-dibromo-3-cyclohexylthiophene, 2,5-dibromo-3-decylthiophene, 5,5'-dibromoo-4,4'-didodecyl-2,2'-bithiophene, 2 , 5-Dibromo-3,4-dihexylthiophene, 2,5-dibromo-3,4-dinitrothiophene, 5,5'-dibromo-4,4'-ditetradecyl-2,2'-bithiophene, 2,6- Dibromodithieno [3,2-b: 2', 3'-d] thiophene, 2,5-dibromo-3,4-ethylenedioxythiophene, 2,5-dibromo-3- (2-ethylhexyl) thiophene, 2, 5-dibromo-3-octylthiophene, 2,5-dibromothieno [3,2-b] thiophene, 3,4- (2', 2'-diethylpropylene) dioxythiophene, 5,5'-dihexyl-2, 2'-Bithiophene, 3,4-dihydroxythiophene-2,5-dicarboxylic acid diethyl ester, 3,4- (2,2-dimethylpropylenedioxy) thiophene, 3,4-ethylenedithiothiophene, 3- (2-) Ethylhexyl) thiophene, 5-perfluorohexyl-2,2'-bithiophene, 3-phenylthiophene, 3,4-propylenedioxythiophene, 3,4-propylenedioxythiophene-2,5-dicarboxylic acid, 3,3 ', 5,5'-Tetrabromo-2,2'-Bithiophene, Thieno [3,2-b] Thiophene-2-carboxyalde Hyd, terthiophene [3,2-b] thiophene-2,5-diboronic acid bis (pinacol ester), terthiophene [2,3-b] thiophene-2,5-dicarboxyaldehyde, thieno [3,2-b] Terthiophene-2,5-dicarboxyaldehyde, thiophen-2-boronic acid pinacol ester, thiophene-2,5-diboronic acid bis (pinacol) ester, 3,4-thiophene dicarboxylic acid, 5-hexyl-2-thiopheneborone Acid pinacol ester, dithieno [3,2-b: 2', 3'-d] thiophene, 3,3 "'-zidodecyl-2,2': 5', 2": 5 ", 2"'-quater Thiophene, 5,5'-di (4-biphenylyl) -2,2'-bithiophene, 5,5''-dibromo-2,2': 5', 2''-terthiophene, 2,5-dibromo -3-Dodecylthiophene, 3,3'-dibromo-2,2'-bithiophene, 5,5'-dibromo-2,2'-bithiophene, 2,5-dibromo-3-butylthiophene, 2,5-dibromo -3-hexylthiophene, 3,3 "'-dihexyl-2,2': 5', 2": 5 ", 2"'-quaterthiophene, 5'''''-dihexyl-2,2' : 5', 2'': 5'', 2''': 5''', 2'''': 5'''', 2'''''-Sexithiophene, 3,4-dihexyl Thiophene, 3,4-dimethoxythiophene, 5,5'-diiodo-2,2'-bithiophene, α-sexithiophene, terthiophene [2,3-b] terthiophene, terthiophene [3,2-b] thiophene, 2 , 2': 5', 2 "-terthiophene, 2,2': 5', 2" -terthiophene-5-boronic acid pinacol ester, 3-decylthiophene, 3-dodecylthiophene, 5,5'-bis (Tritrimethylstannyl) -2,2'-Bithiophene, 2,2'-Bithiophene, 2,2'-Bithiophene-5,5'-Diboronic acid bis (pinacole) ester, 2,2'-Bithiophene-5-boron Acid pinacol ester, 5-fluoro-2,3-thiophene dicarboxyaldehyde, 2-bromo-3-octylthiophene, 2-bromo-3-dodecylthiophene, 2-bromo-3- (bromomethyl) thiophene, 2-bromo- 3-Hexylthiophene, 5-bromo-2-hexylthiophene, 3-bromo-4-methylthiophene, 3-hexylthiophene-2-boronic acid pinacol ester, 5-hexy Examples thereof include ru-2,2'-bithiophene, 5'-hexyl-2,2'-bithiophene-5-boronic acid pinacol ester and the like.
Of these, 3,4-ethylenedioxythiophene is particularly preferable.

As the monomer forming the conductive polymer, it is preferable to use at least one kind of monomer having a functional group capable of binding to an oligosaccharide. Examples of the functional group include a thiol group, an oxylamino group, a hydrazide group, an amino group, a hydroxyl group, a carboxyl group, a carbonyl group, an azido group, an alkynyl group, an epoxy group and an isocyanate group. From the viewpoint of binding of the oligosaccharide to the reducing terminal carbon, it is preferable to have an oxylamino group or a hydrazide group. Since the conductive polymer has an oxylamino group or a hydrazide group, it is not necessary to impart a functional group for bonding with the conductive polymer to the oligosaccharide side, so that the oligosaccharide is directly used as the intermediate layer. Can be used for binding. As the monomer having a functional group capable of binding to an oligosaccharide, one type may be used, or a plurality of types may be used to form a copolymer.

As the monomer forming the conductive polymer, a monomer having a functional group (A) capable of binding to an oligosaccharide and a thienyl group, a phenyl group, anilyl group, a fluorene group, an indol group or a pyrrole group (B) (hereinafter referred to as a monomer). , A monomer for introducing oligosaccharides) is particularly preferable.
The oligosaccharide-introducing monomer can be obtained by introducing a functional group capable of binding to an oligosaccharide into the monomer described as the monomer forming the conductive polymer. Methods of introducing a functional group capable of binding an oligosaccharide into a monomer are known to those skilled in the art. Alternatively, a commercially available product may be used.
Preferred oligosaccharide introduction monomers are 3,4-ethylenedioxythiophene-ONH ₂ or 3,4-ethylenedioxythiophene-ONHNH ₂ , with 3,4-ethylenedioxythiophene-ONH ₂ (EDOTAO) in particular. preferable.

The conductive polymer may be a copolymer of an oligosaccharide-introducing monomer and at least one other monomer. The other monomer is not particularly limited, but is a monomer having a structure obtained by removing a functional group capable of binding to an oligosaccharide from the oligosaccharide introduction monomer (that is, an oligosaccharide introduction monomer before the functional group is introduced). Is preferable. By using such a monomer as another monomer, a copolymer having high conductivity can be obtained without disturbing the conjugated system of the polymer main chain. As the other monomer, one type may be used, or a plurality of types may be used.

Particularly preferable copolymers of the oligosaccharide-introducing monomer and other monomers include a copolymer of 3,4-ethylenedioxythiophene (EDOT) and EDOT-ONH ₂ (EDOTAO).

When a copolymer is used as the conductive polymer, the polymerization ratio of the copolymer is not particularly limited. For example, when a copolymer of one kind of oligosaccharide introducing monomer and one kind of other monomer is used, it is preferably in the range of 9: 1 to 1: 9 from the viewpoint of conductivity and virus recognition. , 8: 2 to 6: 4, more preferably.
For example, when a copolymer of EDOT and EDOTAO is used, the polymerization ratio of the monomers is not particularly limited, but from the viewpoint of conductivity and virus recognition, EDOT: EDOTAO has a molar ratio of 9: 1 to 1: 9. It is preferably present, and more preferably 8: 2 to 6: 4.

The conductive polymer used as the intermediate layer forming material is a conductive layer obtained by electrolytic polymerization from the viewpoints of being able to control properties such as conductivity by electrically controlling the polymerization reaction and being able to form a uniform thin film. It is preferably formed on top. In electrolytic polymerization, for example, a monomer is mixed with a solvent such as an aqueous solution of sodium perchlorate or an aqueous solution of polystyrene sulfonic acid (PSS), and a working electrode having a conductive layer on at least a part of the electrode surface and a counter electrode are brought into contact with the monomer mixture. It can be carried out by applying a potential equal to or higher than the oxidation potential of the monomer to the working electrode. A reference electrode may be further used.
The conditions of electrolytic polymerization such as the electrode to be used, the voltage / current to be applied, and the energization time can be appropriately set by those skilled in the art. As the electrode, for example, a glassy carbon working electrode, a platinum counter electrode, an Ag / AgCl reference electrode, or the like can be preferably used. The electrical conditions can be, for example, a cyclic potential of −0.6 to +1.1 to −0.6V (vs Ag / AgCl reference electrode) and a scanning speed of 0.1V / sec. There is no limitation.
The electrolytic polymerization may be carried out after the monomer and the oligosaccharide are bonded, but it is desirable to bond the oligosaccharide to the intermediate layer after the electrolytic polymerization. This is to prevent oligosaccharides from being redoxed by electropolymerization.

The intermediate layer may be added in addition to being further formed on the conductive layer.
For example, as a method of forming a conductive polymer on the conductive layer as an intermediate layer, a conductive polymer synthesized by another method such as an oxidative polymerization method is applied by a spin coating method or a roll-to-roll method. , Etc. (Yan H. Jo T., Okuzaki H, Highly Conductive and Transparent Poly (3,4-ethylenedioxythiophene) / Poly (4-styrenesulfonate) (PEDOT / PSS) Thin Films, Polymer Journal, 2009, 41 (12), 1028-9; Santos GH, Gavim AX, Silva RF, Rodrigues PC, Kamikawachi RC, de Deus JF, Macedo AG, Roll-to-roll processed PEDOT: PSS thin films: application in flexible electrochromic devices, Journal of Materials Science: Materials in Electronics, 2016, Volume 27, Issue 10, pp 11072-9).
When the same material is used as the intermediate layer forming material and the conductive layer forming material, the conductive layer and the intermediate layer may be integrally formed. Even if the boundary between the intermediate layer and the conductive layer is not clearly provided, it is included in the scope of the present invention as long as the integrally formed one has a function as the intermediate layer and the conductive layer of the present invention.

The fact that the intermediate layer forming material is immobilized on the conductive layer means that, for example, the surface of the metal electrode is subjected to X-ray photoelectron spectroscopy such as XPS (X-ray Photoelectron Spectroscopy) and ESCA (Electron Spectroscopy for Chemical Analysis). It can be confirmed by performing an analysis. For example, when alkanethiol is used as the SAM, a peak derived from a sulfur atom is detected when the alkanethiol is immobilized on a metal electrode.
Further, it can be confirmed that the oligosaccharide is bound to, for example, SAM by labeling the biomolecule with a fluorescent molecule and performing fluorescence detection, for example.

(oligosaccharide)
The oligosaccharide is not particularly limited as long as it binds to the intermediate layer and can selectively capture a specific virus or bacterium. Since the type of oligosaccharide, that is, the sequence of sugar residues constituting the oligosaccharide and the number of sugar residues are unique to the target virus or bacterium, the affinity is higher depending on the virus or bacterium to be detected. A highly potent oligosaccharide is appropriately selected.
As described above, it is preferable that the oligosaccharide has a high detection sensitivity to bacteria and viruses of the biosensor.

Further, the length from the position in contact with the conductive layer of the intermediate layer to the end of the oligosaccharide (hereinafter, may be referred to as the probe layer thickness) when the intermediate layer and the oligosaccharide are bound impairs the effect of the present invention. Unless otherwise limited, it is preferably 0.1 nm to 100 nm from the viewpoint of detection sensitivity when used as a biosensor. Further, it is more preferably 1 nm to 10 nm. The thickness of the intermediate layer can be adjusted by selecting, for example, the length of the main chain of the compound used for SAM, and the length of the oligosaccharide can be adjusted by the number of sugar residues of the oligosaccharide described above. ..
Further, when the conductive polymer is formed by electrolytic polymerization, the film thickness of the conductive polymer can be controlled by the amount of electricity in the electrolytic polymerization, the composition of the monomer used, and the like. Those skilled in the art can appropriately determine the monomer composition and the amount of electricity for obtaining an appropriate film thickness.

Further, the oligosaccharide may be naturally occurring or non-existent, and may be a partially modified oligosaccharide.

Examples of such oligosaccharides include N-linked glycoprotein sugar chains, O-linked glycoprotein sugar chains, polysaccharides, cyclodextrin and the like. Among them, the oligosaccharide for virus detection is preferably an oligosaccharide containing sialic acid. Examples of oligosaccharides containing sialic acid include α2,6-sialyllactose (Neu5Ac (α2,6) Gal (β1,4) Glc) that captures type A influenza virus, and α2,3 that captures bird influenza virus. -Sialyl lactose (Neu5Ac (α2,3) Gal (β1,4) Glc), sialyl 2,6-galactose (Neu5,9Ac ₂ ((α2,6) Gal), which captures human coronavirus, and bovine coronavirus Sialyl 2,3-galactose Neu5,9Ac ₂ ((α2,3) Gal), sialic acid residue that captures mouse hepatitis virus (Neu4,5Ac ₂ ), ganglioside that captures adenovirus (GD1a) (Neu4Acα2,3Galβ1, Examples thereof include 3GalNAcβ1,4 (Neu4Acα2,3) Galβ1,4Glc-OH).
Among these, oligosaccharides that capture influenza virus can selectively capture specific influenza viruses that infect specific species such as humans and birds, and the device using the oligosaccharide was used as a biosensor. In some cases, it can be detected with high sensitivity.

Here, Gal, Neu, Glc, and GalNAc represent the types of sugar residues, Gal is a galactose residue, Neu is an N-acetylneuraminic acid residue which is a sialic acid residue, Glc is a glucose residue, and GalNAc is N. -Represents an acetylgalactosamine residue. In addition, the notation between each sugar residue indicates the mode of binding and the binding position. For example, in the case of Neu4Acα2,3Glc, it means that the 2-position of Neu4Ac and the 3-position of Glc are glycosidic-bonded at α. Further, Neu4,5Ac ₂ indicates that an acetyl group is bonded to the 4- and 5-positions of the N-acetylneuraminic acid residue.

The above oligosaccharides may be prepared from natural products by a known method, or may be chemically or enzymatically prepared by a known method. Alternatively, a commercially available product may be prepared as it is, or may be chemically or enzymatically induced. Examples of commercially available products include α2,3-sialyllactose and α2,6-sialyllactose.

When the oligosaccharide and the intermediate layer forming material are bound, the position of the binding in the oligosaccharide is not particularly limited as long as the effect of the present invention is obtained, and which part of the intermediate layer forming material and the sugar residue constituting the oligosaccharide is present. May be combined with. However, from the viewpoint of ease of bonding with the intermediate layer forming material, it is preferable that the carbon at the terminal having a reducible hemiacetal structure of the oligosaccharide is bonded to the intermediate layer forming material. Here, when the intermediate layer forming material has an oxylamino terminal, the hemiacetal at the reducing end of the oligosaccharide easily becomes an aldehyde depending on the reducing conditions, so that the aldehyde is an oxime on the surface of the intermediate layer forming material. By reacting with the amino group, a stable oxime structure can be formed.
In addition, the oxylamino group is more likely to bind to oligosaccharides than other functional groups, and this binding produces an oxime that is stable in aqueous solution. Therefore, as an intermediate layer forming material, a compound having an oxylamino group and another functional group, or a compound having an oxylamino group and a plurality of types of compounds having another functional group are mixed and used. Since it is possible to bind only the oligosaccharide and the oxylamino group so as not to bind the oligosaccharide to other functional groups, a substituent other than the oligosaccharide is attached to the other functional group of the molecule immobilized on the conductive layer. Can be introduced.

The binding of the oligosaccharide to the intermediate layer is not particularly limited as long as the effect of the present invention is exhibited, but for example, the functional group constituting the SAM is made into an oligo by using the "Glycoblotting method" described in International Publication No. 2004058687. Sugar can be introduced. In this case, for example, the reaction conditions for reacting the oligosaccharide with the oxylamino group in the intermediate layer are preferably 50 ° C. to 70 ° C. for 140 minutes to 240 minutes. Further, in using the Glyco blotting method, a commercially available kit, for example, a kit manufactured by Sumitomo Bakelite Co., Ltd. (Blot Glyco) can be used.

The virus or bacterium to be tested by the biosensor of the present invention is not particularly limited as long as it is a virus or bacterium that can be captured by oligosaccharides. Such bacteria include Mycobacterium tuberculosis, Mycobacterium tuberculosis, Bordetella pertussis, Regionella, Pseudomonas aeruginosa, various pathogenic Escherichia coli, Welsh, Clostridium tetani, Diffcil, Helicobacter pylori, Shigella, and Meningococcus. Bacteria that are pathogenic and may be captured by oligosaccharides. In addition, some lactic acid bacteria (including bifidobacteria) and non-pathogenic bacteria that may be captured by oligosaccharides can be tested by the biosensor of the present invention.
In addition, as viruses, viruses that may be captured by oligosaccharides such as influenza virus, parainfluenza virus, norovirus, adenovirus, dengue virus, herpes virus, corona virus, rhinovirus, and mouse hepatitis virus (MHV) are , Can be the subject of inspection of the biosensor of the present invention.

In addition, the concentration of bacteria used in the biosensor can be expressed by a known method, for example, the number of bacteria in a certain volume, and the number of bacteria can be determined by turbidity or colony forming unit (cfu). As a result, the presence and concentration of bacteria captured by oligosaccharides in the sample can be calculated from the magnitude of the electric potential or the like detected by the biosensor.
Since viruses have different infectivity to living organisms depending on the type and type, the relationship between the physical amount of virus (the number of viruses per unit volume) and the infectivity is not constant. Therefore, the virus titer, which is an index showing infectivity, is used. The amount of virus contained in the sample can be grasped by measuring the magnitude of the electric potential or the like with a biosensor for the sample and comparing the potential with the calibration curve of the potential at the titer calculated in advance. .. For example, in the case of influenza virus, the virus contained in the sample is quantified from the viewpoint of titer by comparing the magnitude of the potential at the titer obtained by the hemagglutination assay and the magnitude of the potential of the sample. Can be judged.

≪Biosensor≫
The biosensor of the present invention includes the above-mentioned biosensor device. The biosensor may include a biosensor device in any form as long as the oligosaccharide of the device can detect the virus or bacterium by capturing the specific virus or bacterium. For example, for the purpose of detecting one type of virus or bacterium, the biosensor can be supported by having one device for the biosensor. Further, if the biosensor is used to detect a plurality of types of viruses or bacteria, the biosensor is provided with a plurality of biosensor devices, and the oligosaccharides of the biosensor devices are specific to each other. It may be an oligosaccharide that selectively captures viruses or bacteria.
As described above, the type of oligosaccharide in this case is preferably appropriately selected depending on the type of target virus or bacterium. By using a biosensor device in which different types of oligosaccharides are arranged, different types of viruses or bacteria can be selectively and sensitively detected.

Further, the biosensor may include a plurality of types of devices having different densities (mol / mm ² ) of oligosaccharides immobilized on the conductive layer of the biosensor device. By providing biosensor devices having different densities of oligosaccharides, it is possible to improve the selectivity and detection sensitivity to a plurality of viruses or bacteria.
It may also be a biosensor equipped with biosensor devices arranged at different densities for each of the different types of oligosaccharides.

In addition, the biosensor provided with the above-mentioned biosensor device may detect a virus or a bacterium based on any change obtained from the device as long as it can detect a specific virus or bacterium. From the viewpoint of selectively and sensitively detecting a virus or a bacterium, the virus or a bacterium can be detected based on any one change selected from the potential, frequency and refractive index. Examples of the change in potential include changes obtained by FET. Examples of the change in frequency include a change obtained by a quartz crystal microbalance (hereinafter referred to as QCM). In addition, examples of the change in the refractive index include changes obtained by surface plasmon resonance (hereinafter referred to as SPR).
Of these, it is more preferable to detect a virus or a bacterium based on a change in electric potential. Further, the change in potential is preferably a change obtained by FET.

Hereinafter, a biosensor that detects a virus or a bacterium based on a change in potential obtained by an FET is referred to as an FET biosensor, and a biosensor that detects a virus or a bacterium based on a change in frequency obtained by a QCM is referred to as a biosensor. , A QCM biosensor, and a biosensor that detects a virus or a bacterium based on a change in the refractive index obtained by SPR is referred to as an SPR biosensor.

Further, examples of the electrode form of the FET biosensor include a floating gate type and an extension gate type. Among them, the extension gate type electrode is excellent in operability in detection. An FET biosensor having an extension gate type electrode may be referred to as an extension gate type FET biosensor.

Hereinafter, embodiments of the present invention will be described in more detail with reference to the accompanying drawings.

(First Embodiment)
<Device for biosensor>
FIG. 1A is a schematic diagram showing a configuration example of a biosensor device used in an FET biosensor. As shown in FIG. 1A, the biosensor device 10 forms an oxime (−ON = C−) on the gold electrode 12 constituting the conductive layer by binding with the oligosaccharide 16. The alkylthiol compound and the alkylthiol compound having a functional group 18 not bonded to the oligosaccharide are each fixed with a thiol group to form the intermediate layer SAM14. Further, the carbon at the reducing end of glucose of α2,6-sialyllactose, which is an oligosaccharide, is bonded to the alkylthiol.
The SAM 14 may be arranged on the entire surface of one side of the conductive layer, or may be arranged only in a certain area of the surface. The intermediate layer may be something other than SAM.

Further, the thickness of the layer of the gold electrode 12 is preferably 0.05 μm to 1 μm from the viewpoint of detection sensitivity in the case of an FET biosensor.
Further, the probe layer thickness of the biosensor device used for the FET biosensor is preferably 1 nm to 10 nm from the viewpoint of detection sensitivity. Further, it is more preferably 1 nm to 5 nm.

The reason why high-sensitivity measurement can be performed when a sensor using the above biosensor device is used can be considered as follows. In ordinary lectins, the affinity (dissociation constant) with sugar chains is on the order of ^10-2 to ^10-5 , and many such sugar chain recognition sites are present on viruses and bacteria. Therefore, since the plurality of oligosaccharides on the surface of the sensor and the sugar chain recognition site of the lectin interact at multiple points, the virus or the like is strongly adsorbed on the electrode surface of the sensor, which causes an electric potential. It is presumed that it appears as a change and the detection sensitivity is effectively improved.

<FET biosensor>
The FET biosensor is a biosensor that detects a virus or a bacterium based on a potential that changes depending on the electric charge of the virus or the bacterium. Specifically, the biosensor surface generated when the virus or the bacterium is captured by an oligosaccharide. The above weak potential change can be measured.
For detection of a virus or bacterium using an FET biosensor, an electrode on which an oligosaccharide capable of capturing the target virus or bacterium is immobilized (hereinafter, may be referred to as a probe-fixed electrode) is used as a virus or bacterium. By immersing in a contained solution (hereinafter, may be referred to as a sample solution), the oligosaccharide captures a virus or a bacterium, and the electric signal of the captured virus or the bacterium is captured by a change in potential.

The electric potential method that captures the electrical signal of the captured virus or bacterium by the change of electric potential is the solid / liquid interface between the electrode (solid) and the solution (liquid) that is generated by the oligosaccharide capturing the charged virus or bacterium. It detects changes in potential (changes in surface charge density) at the solid-liquid interface.

In the detection of a virus or a bacterium, a current method that captures an electric signal of the captured virus or a bacterium by a change in an electric current may be applied.
In the current method, a method that utilizes a redox reaction on an electrode is usually adopted. However, the presence of an oxidizing or reducing substance (eg, ascorbic acid) in a solution containing a virus or bacterium may cause an electric current based on the oxidation or reduction to flow, which may hinder the detection of the virus or bacterium. In addition, the electrode reaction proceeds on the electrode with the current measurement. Since the electrode reaction is an irreversible and non-equilibrium reaction, corrosion of the electrode, generation of gas, etc. occur, the stability of current measurement is impaired, and the detection accuracy may deteriorate especially in the case of repeated measurement.
Therefore, it is preferable that the electrical signal of the virus or bacterium captured by the oligosaccharide is captured by the change in electric potential.

The FET biosensor is a field effect transistor (FET) in which the conductive layer of the biosensor device of the present invention is applied as a gate electrode, and a known field effect transistor configuration can be used for configurations other than the gate electrode. A configuration example of the FET biosensor will be described with reference to FIG. 1 (B).
FIG. 1B is a side schematic view showing a configuration example of a FET biosensor provided with a floating gate type electrode. The floating gate type FET biosensor may be referred to as a floating gate type FET biosensor.
As shown in FIG. 1 (B), in the floating gate type FET biosensor 20, the source 22S and the drain 24D are arranged apart from each other in the silicon substrate 32, and the silicon substrate 32 and the layer 28 formed of silicon oxide are formed. The layer 26 made of silicon nitride and the above-mentioned biosensor device 10 are laminated in this order as a gate electrode. In this case, the layer 28 formed of silicon oxide and the layer 26 formed of silicon nitride are layers that function as a gate insulating film.
In a transistor, a floating electrode is generally formed on a channel region between a drain and a source, and a potential is detected by the floating electrode. Therefore, in the floating gate type FET biosensor 20, the channel region between the drain 24D and the source 22S is located in the region vertically below the biosensor device 10.

With the layer structure shown in FIG. 1B, the biosensor can selectively and sensitively detect the target virus or bacterium.

FIG. 2 is a schematic diagram for explaining the FET biosensor 40 provided with the biosensor device 10 represented by FIG. 1A as a probe fixed electrode of the extension gate type FET biosensor. The probe fixed electrode portion 10a has the same configuration as the biosensor device 10.
In FIG. 2, the extension gate type FET biosensor 40 includes a silicon substrate 46, a drain 44D and a source 42S provided apart from each other in a part of the silicon substrate 46, and a silicon substrate 46, a drain 44D, and a source 42S. It includes a gate insulating film 48 provided above and a gate insulating film 47. Further, inside the gate insulating film 48, a gate electrode portion 13 made of conductive polycrystalline silicon is provided as the gate electrode 15. Further, the probe fixed electrode portion 10a is connected to the gate electrode portion 13 to form the gate electrode 15.
By providing the gate electrode portion 13 connected to the probe fixed electrode portion 10a in this way, the channel region between the drain 44D and the source 42S is set to the region vertically below the probe fixed electrode like the above floating electrode. It is not necessary to position the gate electrode portion 13 connected to the probe fixed electrode portion 10a, and the gate electrode portion 13 can be freely formed regardless of the position. Further, since the gate electrode portion 13 is made of a conductive material, the shape between the portion other than the probe fixed electrode portion 10a and the probe fixed electrode portion 10a in the extended gate type FET biosensor is free. Can be set to. Further, in FIG. 2, the gate electrode portion 13 is formed by using polycrystalline silicon, but if the material forming the gate electrode portion 13 has conductivity, a desired shape or the like is taken into consideration. Therefore, it is preferable that the material is appropriately selected.

In the floating gate type FET biosensor and the extended gate type FET biosensor, the gate insulating films 28 and 48 are silicon oxide (SiO ₂ ), silicon nitride (SiN), aluminum oxide (Al ₂ O ₃ ), and tantalum oxide (Ta ₂ ). Materials such as O ₅ ) are used alone or in combination, usually in order to maintain good transistor operation, on top of silicon oxide (SiO ₂ ) silicon dioxide (SiN), aluminum oxide (Al ₂ O ₃ ) or tantalum oxide (Al ₂ O ₃ ). It has a two-layer structure in which Ta ₂ O ₅ ) is laminated.
Polysilicon (Poly-Si) is desirable for the gate electrode portion 13, and the gate electrode portion 13 has good consistency with a so-called self-alignment process in which a source 42D and a drain 44D are formed by ion implantation through a polysilicon gate.
For the electrode 12, a noble metal such as gold, platinum, silver or palladium, or a conductive substance such as a carbon material can be used.
Specific examples of the carbon material include conductive carbon materials such as glassy carbon, graphite, graphene, and carbon nanotubes.

In the extended gate type FET biosensor, the gate electrode portion 13 may be used as wiring. In this case, a material having low resistance and good workability such as etching is preferable, and the material is polysilicon (Poly-Si), aluminum, or the like. Molybdenum or the like can be used.
Further, the gate electrode portion 13 and the electrode 12 connected to the gate electrode portion 13 can be formed of the same material such as gold.

Further, as the FET biosensor used for detecting a virus or a bacterium, the above-mentioned configuration of the FET biosensor itself may be used, but as shown in FIG. 3A, for example, the above-mentioned extended gate type FET. An operational amplifier (op amp) may be configured by an integrated circuit using the biosensor 10b as a part of the input stage to form the FET biosensor 50.
Further, as shown in FIG. 3B, the FET biosensor has the functions shown in FIG. 3A, and further includes an extension gate type transistor 63, a phase compensation capacitor 62, and a differential amplifier 64. A highly functional FET biosensor 60 may be provided with an amplifier circuit 66, an output buffer source follower 68, or the like.

The highly functional FET biosensor as described above can detect a weak potential difference caused by the presence or absence of virus or bacterium capture with higher sensitivity. Further, in the FET biosensor as shown in FIG. 3, the biosensor device (probe fixing electrode 10b) on which the oligosaccharide is immobilized can be attached to and detached from the transistor, and can be replaced only by the probe fixing electrode 10b. Therefore, it is excellent in economy and usability as a biosensor.

Further, since it is necessary to detect viruses or bacteria by the above-mentioned FET biosensor in a state where the salt concentration is low, for example, the sample solution is appropriately diluted and the probe fixing electrode is immersed, or the sample solution is placed on the electrode. It can be done by dropping it on. In this case, the concentration of virus or bacterium contained in the sample solution before dilution is important for detection. For example, in the case of influenza virus, the virus titer can be detected under the condition of 1/100 HAU to 10 HAU. It is desirable to have. Since the biosensor of the present invention can be rapidly detected at a lower virus concentration than the conventional immunochromatography, it can be said that it is a rapid high-sensitivity detection method that replaces the conventional method. HAU represents the titer of the virus. The titer of the virus can be determined by a known blood agglutination assay.

(Second Embodiment)
<QCM biosensor>
The QCM biosensor is configured to include at least a probe fixed electrode which is a device for a biosensor, a crystal oscillator, and an oscillation circuit for detecting the vibration of the crystal oscillator, and is based on the quartz crystal microbalance method. A biosensor that detects viruses or bacteria.
Here, the crystal oscillator microbalance method uses a laminated body in which electrodes are attached to both sides of a thin plate-shaped crystal, and when an AC electric field is applied to each electrode, the crystal oscillator has a constant frequency (resonance frequency). This is a method of measuring the mass of a substance attached to an electrode by utilizing the vibrating property.
The QCM biosensor measures the mass change of the attached virus or bacterium by utilizing the property that the resonance frequency decreases according to the mass of the attached virus or bacterium when the virus or bacterium adheres to the electrode adjacent to the crystal oscillator. By doing so, the attached virus or bacterium can be detected.
A configuration example of the QCM biosensor will be described with reference to FIG.

In FIG. 4, a QCM biosensor in which a gold electrode 112b, a crystal oscillator (quartz plate) 118, a gold electrode 112a, and a SAM 114 as an intermediate layer are laminated in this order, and an oligosaccharide 116 is bound to the SAM 114. 100 is shown. An oscillation circuit (not shown) is connected to the electrode 112a and the electrode 112b by a conducting wire. As described above, the QCM biosensor 100 includes a device for the QCM biosensor.
The QCM biosensor device is the same as the FET biosensor device 10 described above.
The electrode 112a and the electrode 112b are not particularly distinguished, and either of them may be an electrode (probe fixing electrode) to which the oligosaccharide 116 is bound, and both the electrode 112a and the electrode 112b are the probe fixing electrodes. May be good.

In the measurement with the QCM biosensor, the probe fixed electrode is immersed in a sample in which a virus or a bacterium is present in a solution, the sample is dropped on the surface of the probe fixed electrode, or the QCM biosensor is immersed in the sample. Just do it.

The presence of viruses or bacteria selectively bound to oligosaccharides on the surface of the probe fixed electrode is quantitatively detected with high sensitivity by the change in resonance frequency (Hz) based on the vibration of the crystal oscillator obtained by the above measurement. be able to.

(Third Embodiment)
<SPR biosensor>
The SPR biosensor includes at least a probe fixed electrode, a light irradiation device, and a light receiving device that receives reflected light and detects the reflectance, and by changing the refractive index of the electrode using surface plasmon resonance. , A sensor that detects viruses or bacteria.
A configuration example of the SPR biosensor will be described with reference to FIG.

FIG. 5 shows an SPR biosensor 200 in which an electrode (gold thin film) 212 and an intermediate layer 214 are laminated in this order on a prism 218, and an oligosaccharide 216 is further bound to the intermediate layer 214.
In the SPR biosensor 200, after incident through a prism 218 from a light irradiation device (laser oscillator) LV (not shown), light (hν) is irradiated to the gold thin film 212 at an incident angle θ, reflected by the gold thin film 212, and then the prism. Light is received by a light receiving device (laser detector) Ld (not shown) through 218. At this time, when the incident angle θ is changed, the angle (resonance angle) when the reflected light is most attenuated changes depending on the presence or absence of virus or bacteria adhering to the gold thin film 212. It is possible to confirm the adsorption of viruses or bacteria on the gold thin film 212. As the light receiving device, a CCD (Charge Coupled Device) camera or the like may be used. Further, a glass substrate may be arranged between the gold thin film 212 and the prism 218.

In addition, the evanescent wave that exudes onto the gold thin film when irradiated with laser light becomes less affected as the distance from the gold thin film increases, and does not affect changes in the resonance angle. Therefore, the thinner the probe layer, the better. The probe layer thickness of the probe fixed electrode of the SPR biosensor is preferably 1 nm to 10 nm from the viewpoint of detection sensitivity.

With such a configuration, the SPR biosensor can selectively detect a virus or a bacterium with high sensitivity.

Hereinafter, the present invention will be described in more detail with reference to Examples, but the present invention is not limited to the following Examples as long as the gist of the present invention is not exceeded. Unless otherwise specified, "%" and "part" are based on mass.

[Example 1]
<Biosensor>
(Manufacturing of device for FET biosensor)
FIG. 6 shows a schematic diagram illustrating a method for manufacturing a biosensor device 70 used for a FET biosensor, that is, a gold electrode on which oligosaccharides are immobilized, which is used in an example. First, as shown in FIG. 6, as an intermediate layer forming material, undecane disulfide having an oxylamino group (H ₂ N-O- (CH ₂ ) _11- S-S- (CH ₂ ) _11- OH, H ₂ N -O- (CH ₂ ) _11- S-S- (CH ₂ ) _11- NH ₂ ) (1 mmol / L) was synthesized with reference to the literature of Park et al. (Langmuir, 2008, Vol. 24, 6201-6207). went. The obtained intermediate layer forming material was dissolved in ethanol to prepare a SAM forming solution. Further, as the gold electrode 72 used as the conductive layer, a circular gold electrode (Au) plate having a diameter of 0.5 mm was prepared. The obtained SAM forming solution was applied to the gold electrode 72 to form a SAM on the gold electrode 72. The density of molecules having thiol groups on the surface of the gold electrode 72, measured in advance by CV, was shown to be 3.85 ± 0.81 mmol / nm ² .
Α2,6-Sialyl lactose (manufactured by Tokyo Kasei Kogyo Co., Ltd.) was applied to the gold electrode 72 on which SAM76a was immobilized as described above by a condensation reaction (60 ° C., 240 minutes) to form an oxylamino group of SAM76a. The oligosaccharide 74 is immobilized by binding to the reducing terminal carbon of the glucose residue of α2,6-sialyllactose and SAM76b via an oxime, that is, the following extended gate type. A biosensor device 70 used for an FET biosensor was manufactured.

(Extended gate type FET biosensor)
A gate electrode portion made of conductive polycrystalline silicon as shown in FIG. 2 was connected to the gold electrode 72 of the biosensor device 70 produced above to form an extended gate type electrode. This was mounted on an FET equipped with an operational amplifier or the like as shown in FIG. 3, an extended gate type FET biosensor 80 was manufactured, and the following measurements were performed.

(Extended gate type FET biosensor with reference electrode)
In Example 1, a product in which α2,6-sialyllactose was bound to a SAM formed on a gold electrode was prepared, and a silver / silver chloride electrode was incorporated into a biosensor as a reference electrode, and the same as in Example 1. The extended gate type FET biosensor was used as the extended gate type FET biosensor 82.

[Control 1]
In Example 1, a biosensor device was prepared and carried out in the same manner except that the α2,6-sialyll lactose used for producing the biosensor device used for the FET biosensor was changed to α2,3-sialyll lactose. An extension gate type FET biosensor 81 was manufactured in the same manner as in Example 1, and the following measurements were performed.

(Sample solution)
As a stock solution of the sample solution used in the examples and controls, a solution containing influenza A virus H1N1 (HA value 256) detoxified by 0.05% formalin treatment was prepared. In addition, the HA titer is an index showing the blood agglutination ability of the virus and is treated as an index of infectivity. The stock solution of the sample solution has a HA value of 256 (256 HAU) because the virus solution diluted up to 256 times showed blood agglutination ability.

<Measurement and results>
The extended gate FET biosensor 82 equipped with the reference electrode is first immersed in a buffer solution (0.01% phosphate buffered saline), and the buffer solution must be changed as appropriate until the potential drift line stabilizes. It was left still. Then, the FET biosensor 82 was immersed in the FET biosensor 82 at 10-minute intervals in order from the one with the lowest concentration within the range where the final virus concentration was 1/1280 to 4/5 HAU, and measurement was performed. The results are shown in FIG. In the potential measurement, for the drift line in the graph showing the relationship between the potential obtained in the buffer solution and the elapsed time of the virus before measurement, an approximate straight line that approximates the drift line to a straight line is created, and the virus When the solution to which is added is measured, how much the potential change is obtained from the approximate straight line is measured. Since the drift line is a gentle curve, it does not show a strict potential change, but the potential changes on the α2,6-sialyllactose surface and the α2,3-sialyllactose surface are relative to each other by the above measurement method. Can be compared to.

Next, as in the case of using the sensor 82 described above, as the first embodiment, the measurement result when the extended gate type FET biosensor 80 is used is shown in FIG. 8 (B) as the control 1 and the extended gate. The measurement results when the type FET biosensor 81 is used are shown in FIG. 8A. Each broken line is an approximate straight line obtained when the above-mentioned extended gate type FET biosensor 82 is used.

As shown in FIG. 8, the curves showing the relationship between the potential and the elapsed time of Control 1 and Example 1 both have a gentle slope from the approximate straight line and a large potential difference from the approximate straight line as the virus concentration is increased. You can see that. Further, in Example 1, it was shown that the potential difference from the approximate straight line was significantly larger than that of Control 1.

FIG. 9 shows a graph showing the relationship between the dilution ratio of the stock solution of the sample solution and the potential difference (ΔV) with respect to the potential difference obtained above. In addition, the potential difference obtained in Example 1 and the potential difference (ΔV) obtained in Control 1 were compared, and whether or not there was a significant difference was also added for each concentration as a p value determined by t-test. doing. In addition, each graph in FIG. 9 is obtained by performing the same experiment twice.

From the results of FIG. 9, it was shown that ΔV was significantly larger on the surface of α2,6-sialyllactose than on the surface of α2,3-sialyllactose in the range of virus concentration of 1/1280 to 4/5 HAU. This indicates that the virus was specifically adsorbed on the surface of α2,6-sialyllactose and the potential changed.
As described above, it was shown that the extended gate type FET biosensor is a highly sensitive sensor capable of discriminating between the human-infected type and the bird-infected type in the virus.

[Example 2]
(Biosensor device for QCM biosensor)
In the production of the biosensor device used for the QCM biosensor, "NaPiCOS" manufactured by Nihon Dempa Kogyo Co., Ltd. was used. As shown in FIG. 4, NaPiCOS has a structure in which a gold plate corresponding to two gold electrodes and a crystal are adhered to each other. Further, in the production of the biosensor device, the same device as in Example 1 was used except for the gold electrode and the crystal.
In the same manner as in Example 1, the SAM was fixed on the gold electrode and α2,6-sialyllactose was bound to the SAM to prepare a biosensor device used for the following QCM biosensor.
(QCM biosensor)
The biosensor device produced above is filled with a virus solution in a sample loop using a 250 μL syringe manufactured by Hamilton Co., Ltd., and the virus solution is appropriately flowed. This sensor utilizes the fact that the frequency generated by the resonance vibration of the crystal plate in the device is reduced by the weight increase due to the adsorption of influenza virus while the biosensor device is immersed in the solution at a constant flow velocity. It measures the amount of influenza adsorbed.

[Control 2]
After producing the biosensor device in Example 2 in the same manner except that the α2,6-sialyll lactose used for producing the biosensor device used for the QCM biosensor was changed to α2,3-sialyll lactoose. , A QCM biosensor was prepared in the same manner as in Example 2, and the following measurements were performed.

<Measurement and results>
First, 0.01-fold phosphate buffer was introduced into the QCM biosensor at a flow rate of 3.0 ml / h. After that, after confirming that the frequency was stable, a solution obtained by diluting the stock solution of the sample solution stepwise with 0.01-fold phosphate buffer was added to the flow path in order from the lowest concentration by 200 μL at 5 minute intervals. Introduced (from 1/32 HAU to 1 HAU). In addition, the measurements of Example 2 and Control 2 were carried out three times each. FIG. 10 shows the results of each of the three times.

From the results, it was shown that the amount of decrease in frequency (ΔFrequency (Hz)) in Example 2 was clearly larger than that in Control 2. Therefore, it was shown that a QCM biosensor having an electrode on which α2,6-sialyllactose is immobilized can detect influenza A selectively and with high sensitivity.
As described above, it was shown that the QCM biosensor is a highly sensitive sensor capable of discriminating between the human-infected type and the avian-infected type of influenza virus.

[Example 3]
Biosensor (monomer synthesis) using a polythiophene-based conductive polymer as an intermediate layer
The synthesis of the PEDOT-ONH ₂ (EDOTAO) monomer was carried out by the following five steps (Luo, SC; Ali, EM; Tansil, NC; Yu, HH; Gao, S .; Kantchev, EAB; Ying, JY. , Poly (3,4-ethylenedioxythiophene) (PEDOT) nanobiointerfaces: Thin, ultrasmooth, and functionalized PEDOT films with in vitro and in vivo biocompatibility. Langmuir 2008, 24 (15), 8071-8077).

-First step: Synthesis of EDOT-Cl 28 mL of dehydrated toluene was added to a 100 mL two-necked eggplant flask and degassed with stirring. Next, while substituting the entire reaction system with nitrogen, 3,4-dimethoxythiophene (1.36 g, 10.4 mmol), p-toluenesulfone (0.16 g, 0.9 mmol) and 3-chloro-1,2- Propanethiol (2.60 g, 23.1 mmol) was added and reacted at 90 ° C. for 24 hours. Then, 3-chloro-1,2-propanediol (2.60 g, 23.1 mmol) was added while substituting with nitrogen, and the mixture was reacted at 90 ° C. for 3 hours. Next, it was evaporated and purified by column chromatography (silica gel, dichloromethane / hexane = 2/8). A pale yellow oily reaction product, EDOT-Cl, was obtained. The yield was 56%.

-Second step: Synthesis of EDOT-N ₃ Add 30 mL of DMF, EDOT-Cl (1.18 g, 6.2 mmol) and NaN ₃ (0.48 g, 7.4 mmol) to a 100 mL two-necked eggplant flask and add 120. The reaction was carried out at ° C. for 90 minutes. The solution was added 15mL of _{H 2} O was allowed to cool, to carry out a liquid separation with ethyl acetate (3 × 15mL), the combined organic layers were washed with _{H 2 O (3 × 15mL)} , dried over anhydrous MgSO ₄ After filtration, the mixture was evaporated to obtain a brown oily reaction product EDOT-N ₃ . The yield was 81%.

Third step: Synthesis of EDOT-NH _{2 In} a 100 mL eggplant flask, 10 mL THF, 10 mL 2 mol / L NaOH, EDOT-N ₃ (0.99 g, 5.1 mmol) and triphenylphosphine (1. 60 g (6.1 mmol) was added, and the mixture was reacted at 50 ° C. for 3 hours. When the mixed solution reached room temperature, it was evaporated to remove THF. The pH of the solution was then adjusted to less than 2 with 1 M aqueous HCl and separated with dichloromethane (3 x 15 mL). The organic layer was discarded and the pH of the aqueous layer was adjusted to be greater than 12 using 1 mol / L NaOH. The liquid was separated using dichloromethane (3 × 15 mL), the organic layers were collectively dried over anhydrous Then ₄ and filtered. Finally, it was evaporated to obtain a pale yellow oily EDOT-NH ₂ . The yield was 75%.

4th step: Synthesis of EDOT-ONH-Boc EDOT-NH ₂ (0.65 g, 3.8 mmol), 2-((tert-butoxycarbonyl) aminooxy) acetic acid (0.80 g, 4) in a 100 mL eggplant flask. .2mmol), and stirred for 10 minutes at room temperature of _{H 2} O in ethanol was added 7ml and 13 mL. Then, DMT-MM (1.57 g, 5.7 mmol) was added and the reaction was carried out for another 3 hours. After the mixture was evaporated, 15 mL of saturated Na ₂ CO ₃ was added and the mixture was separated by diethyl ether (3 x 15 mL). The organic layers were collectively washed with a 1 mol / L HCl aqueous solution, dried over anhydrous Na ₂ SO ₄ , and filtered. Finally, it was evaporated to obtain a transparent oily EDOT-ONH-Boc. The yield was 70%.

Fifth step: Synthesis of EDOT-ONH _{2 To} a 100 mL eggplant flask, add EDOT-ONH-Boc (0.46 g, 1.33 mmol), 15 mL of MeOH and 5 mL of H ₂ O, and add 10 mL of EDOT-ONH ₂ with stirring. Concentrated hydrochloric acid was slowly added dropwise, and the mixture was reacted at room temperature for 1 hour. After completion of the reaction, and evaporated in addition of _{H 2} O 20 mL. Separate with diethyl ether (3 x 15 mL), combine aqueous layers, prepare until pH is greater than 12, separate with diethyl ether (3 x 15 mL), combine organic layers, anhydrous Na ₂ SO _It was dried in ₄ and filtered. Finally, it was evaporated to obtain a transparent oily EDOT-ONH ₂ . The yield was 62%. The overall yield was 16%.

(Electropolymerization of EDOT-ONH ₂ (EDOTAO) and EDOT and evaluation of electrolytic polymer film)
Electropolymerization of EDOTAO and EDOT (3,4-ethylenedioxythiophene) monomer was carried out, and a copolymer was randomly attached to the surface of the glass carbon electrode in the aqueous solution.

As the working electrode, a commercially available glass carbon electrode (inner diameter 3 mm, manufactured by CH Instruments (TX, USA)) was used. Before use, the electrode surface was carefully polished with aluminum powder (particle size 0.05 μm (Baikouski International (Charlotte, NC, USA)) and washed with miliQ water. EDOTAO and EDOT had a molar ratio of EDOTAO: EDOT of 0. : 100 (hereinafter sometimes referred to as 0 mol% EDOTAO), 25:75 (hereinafter sometimes referred to as 25 mol% EDOTAO), 50:50 (hereinafter sometimes referred to as 50 mol% EDOTAO), 75:25 (Hereinafter, it may be referred to as 75 mol% EDOTAO), or 1 mL of an aqueous solution of sodium perchlorate (100 mmol / L) contained at a concentration of 10 mmol / L at 100: 0 (hereinafter, may be referred to as 100 mol% EDOTAO). In addition to the glass cell, three electrodes, a glass carbon action electrode, a platinum counter electrode, and an Ag / AgCl (in 3.3 mol / L KCl _aq ) reference electrode, were attached to the glass cell, and a potential stat (Versa STAT, manufactured by Powderon Applied Research) was attached. ) Was applied, and a cyclic potential (-0.6 ^to +1.1 ^to -0.6V vs Ag / AgCl at 0.1V / sec) was applied and swept continuously for 10 cycles. Necessary for forming a polymer film. It was observed that the oxidation potential (forward scan) was around + 1 V (vs Ag / AgCl) (FIGS. 11A, (a) to (e)). The increase in current density with the increase in the number of cyclics was polymerized. It represents the film growth process. The unit interfacial capacitance (C, mF / cm ² ) for each cycle is determined by the difference between the forward and reverse current densities (ΔI, mA) near + 0.2 V, and is the cyclic number. It increased with the increase (Fig. 11A, (f)).
C = ΔI / νA equation (1)

ν (V / sec) represents the scan rate and A (cm ² ) represents the surface area of the electrode. In the case of 0 mol% EDOTAO and 5 mol% EDOTA2, the significant increase in capacitance indicates that the electrode surface conductive polymer film was formed. On the other hand, in the case of 50 mol% EDOTA to 100 mol% EDOTA, no increase in capacitance was observed. A high content of EDOTAO suggests that the conductivity of the electrolytic polymerized film is reduced.

The film thickness of the polymerized film can be controlled by the amount of electricity polymerized, and varies depending on the composition of the electrolytically polymerized monomer. Therefore, the film thickness and surface roughness of the poly (EDOTAO-co-EDOT) electrolytic polymerized film fixed on the surface of the gold substrate were measured using a step meter (DekTak 150, manufactured by ULVAC). As the proportion of EDOTAO monomer increased, the film thickness became thinner and the surface roughness became smaller (FIGS. 11B, (g) to (i)). It is considered that the side chain of the EDOT skeleton of the EDOTAO monomer sterically hinders the formation of π-conjugated system, makes it difficult to polymerize, and reduces the conductivity.

The elemental composition of the surface of the electrolytic polymer film was observed by X-ray photoelectron spectroscopy (XPS) (AXIS-HSi165; 408, manufactured by Shimadzu-Kratos (Kyoto, Japan)). The light source is MgKα ray, acceleration voltage 15 kV, incident angle 90 °. The curve fitting of the high resolution spectrum was analyzed based on the Gaussian function. Mean and standard deviation were obtained from 4 measurement points. The composition of the monomer was changed on the surface of the gold substrate (0 mol% to 50 mol% EDOTAO), and 10-cycle electrolytic polymerization was performed. When EDOTAO was 0 mol%, the peak of N1s was not observed, but as the ratio of EDOTAO increased (25 mol% EDOTAO, 50 mol% EDOTAO), the atomic ratio of nitrogen increased (FIG. 12, (a)). When the molar ratios of EDOT and EDOTAO on the surface of the copolymer film are m and n, respectively, the following relationship holds:
N = 2n, S = m + n equation (2)
The structure of the conductive copolymer polymer is shown below.

The introduction rate of the EDOTAO composition on the surface of the electrolytic polymer film can be shown by the formula (3).
n / (m + n) = N / 2S equation (3)
N and S represent the actually measured element ratios of nitrogen and sulfur. It was found that the ratio of the EDOTAO composition of the conductive copolymer polymer was proportional to the supply amount of the EDOTAO monomer during electrolytic polymerization (FIG. 12, (c)).

The low impedance interface at the solid / liquid interface is one of the important properties as a material for biosensing and bioelectronics applications of conductive polymers. Therefore, electrochemical impedance measurement (EIS) was performed. PGSTAT 302 potentiostat (manufactured by Metrohm Autolab) is a glass carbon working electrode to which a conductive polymer is fixed, a three-electrode cell composed of an Ag / AgCl (in 3.3 mol / L KCl _aq ) reference electrode and a Pt counter electrode. Connected with. The measurement solution used was DPBS (0.01xDPBS pH 7.4, manufactured by Dulbecco), which contained 5 mM ferricyanide / ferricyanide and was diluted 100-fold with water.

Frequency range 0.1 kHz to 10 kHz (10 measurement points per digit), 50 mV AC potential, DC bias + 0.2 V vs Ag / AgCl. The charge transfer resistance was obtained by fitting the Nyquist plot to Randle's equivalent circuit model. Mean and standard error were obtained from four proprietary measurements. The electrochemical impedance of a poly (EDOTAO-co-EDOT) film electropolymerized on a glass carbon electrode was measured. The semi-circular part on the high frequency side of the Nyquist plot belongs to the charge transfer limit process of the redox reaction [Fe (CN) ₆ ³ / Fe (CN) ₆ ^4- ], and the linear part on the low frequency side is it. Belongs to the diffusion limit process of (Fig. 13A). The Nyquist plot of the 0 mol% EDPTAO (ie, PEDOT) polymer film responded linearly and behaved like a capacitor. From the fitted data based on the equivalent circuit of Randle, charge transfer resistance with increasing proportion of EDOTAO _{(R ct)} that increase was observed (Figure 13B). It was shown that the behavior of CPE (constant phase element) with respect to the change in the ratio of EDOTAO tended to be close to the interfacial capacitance (FIG. 11A) obtained from CV. It is considered that the higher the EDOTAO composition, the more steric hindrance the oxyamine side chain to the thiophene skeleton, the lower the film conductivity, and therefore the higher the impedance of the copolymerized film. Since the ratio of the oxyamine group of the copolymer film and the electron / ionic conductivity of the film cancel each other out, a 25 mol% EDOTAO copolymer was used in the following experiments from a comprehensive viewpoint.

Glycoblotting reaction on a copolymer film electrolytically polymerized on the surface of a glass carbon electrode (Hideshima, S .; Hinou, H .; Ebihara, D .; Sato, R .; Kuroiwa, S .; Nakanishi, T .; Osaka, T., Attomolar detection of influenza A virus hemagglutinin human H1 and avian H5 using glycan-blotted field effect transistor biosensor. Anal Chem 2013, 85 (12), 5641-4.) Using sugar chains (2,6) -Sialyllactose, 2,3-sialyllactose) was introduced.

0 mol% EDOTAO or 25 mol% EDOTAO was electropolymerized on the surface of the gold substrate after washing, and then washed with MilliQ water. The sugar chains of 2,6-sialyllactose or 2,3-sialyllactose in an aqueous acetic acid solution were prepared so as to have a final concentration of 0.1 mmol / L (pH 5.3), respectively. Then, a glycoblotting reaction was carried out at 60 ° C. for 12 hours.

The wettability of the membrane surface before and after the introduction of the sugar chain was evaluated using a contact angle meter (DM-501, manufactured by Kyowa Interface Science Co., Ltd.) (FIGS. 14A and 14B). The water droplet contact angles of the 0 mol% EDOTAO and the 25 mol% EDOTAO electrolytic polymerized membranes before the introduction of the sugar chain were 68.8 ± 0.2 ° and 63.8 ± 0.2 °, respectively (FIG. 14A). The water droplet contact angles of the 0 mol% EDOTAO and 25 mol% EDOTAO electrolytic polymerized membranes after the introduction of the sugar chain were 40.5 ± 0.5 ° and 4 °>, respectively, and the values of the contact angles decreased after the introduction of the sugar chain ( FIG. 14B).

It is considered that the contact angle on the surface of the 25 mol% EDOTAO polymerized film was remarkably reduced because the sugar chain was introduced by the glycoblotting reaction and the surface of the polymerized film became superhydrophilic. In the case of 0 mol% EDOTAO (ie, PEDOT), the cause of the decrease in the contact angle is considered to be the non-specific adsorption of sugar chains.

In order to quantitatively evaluate the glicoblotting reaction, it was evaluated using a QCM sensor.
The QCM was measured at 25 ° C. using a twin QCM sensor (30 MHz, 1 Hz (ΔF) = 35 pg / cm ² ) and a QCM analyzer (manufactured by Nippon Denshi Kogyo Co., Ltd. (Tokyo, Japan)). After electrolytic polymerization of 0 mol% EDOTAO or 25 mol% EDOTAO on the QCM sensor, the QCM sensor was immersed in milliQ water for 12 hours in order to remove the dopant sodium perchlorate from the polymer film, and then the measurement was performed. Then, the sugar chain was introduced under the same conditions as above, washed with milliQ water, and then measured (FIG. 14C). In the case of the QCM sensor in which 25 mol% EDOTAO was electrolytically polymerized after the introduction of the sugar chain, a decrease in frequency due to the adhesion of the sugar chain was confirmed, and it was found that 9.4 μg / cm ^{2 was} attached. In the case of 0 mol% EDOTAO, it is presumed that the increase in frequency is due to the escape of the dopant NaClO ₄ remaining in the conductive polymer in a small amount into the solution.

(Specific detection of influenza virus)
-Real-time detection of influenza virus by QCM sensor For the influenza virus sample, inactivated human influenza A virus (H1N1, PR8, strain) was used. The degree of infection is 256 HA unit (HAU, chicken hemagglutination experiment). The QCM was measured using a twin QCM sensor (30 MHz) and a QCM analyzer (manufactured by Nippon Denshi Kogyo (Tokyo, Japan)). For the QCM sensor for measuring influenza virus, 25 mol% EDOTAO is electrolyzed on the surface of the QCM sensor for 10 cycles, and an aqueous acetate solution (pH 5.3) of 0.1 mM sugar chain (2,6-sialyllactose or 2,3-sialyllactose) is used. ) And 60 ° C. for 12 hours to prepare. Measurement conditions: DPBS (0.01 x DPBS pH 7.4, manufactured by Dalbecco) diluted 100 times with water, influenza virus sample volume 200 μL, influenza virus concentration 0.125, flow rate 50 μL / min, It was set to 0.25, 0.5, 1.0 and 2.0 HAU. Measurement temperature 25 ± 0.02 ° C. Injection was performed at 7 minute intervals. That is, the first 0.125 HAU influenza virus sample was injected, 7 minutes later, the 0.25 HAU sample was injected, and the samples up to 2.0 HAU were sequentially injected every 7 minutes. Mean and standard deviation were obtained from four equilibrium measurement data.

FIG. 15A shows the change over time when the change in weight due to the molecular recognition reaction of the 2,6-sialyllactose sugar chain or the electrode modified with the sugar chain was shown as a frequency when the concentration of influenza virus was changed. Shown. FIG. 15B is a plot of the frequency at which each injection is in equilibrium with respect to the influenza virus concentration.
The QCM sensor in which 2,3-sialyllactose was immobilized on poly (EDOTAO-co-EDOT) did not show a remarkable response until 2HAU influenza virus was injected, but on poly (EDOTAO-co-EDOT) 2, The QCM sensor immobilized with 6-sialyll lactose responded significantly from the first injection of the influenza virus sample (Fig. 15A).
From FIG. 15B, it was shown that the higher the influenza virus concentration, the greater the decrease in frequency due to the sugar chain capturing the influenza virus, and the concentration-dependent change in frequency. In addition, even in the case of 0.125 HAU, which has the lowest influenza virus concentration, the amount of influenza virus adhered was p <0.01, and there was a high significant difference. It was confirmed that the influenza virus selectively binds to a QCM sensor having 2,6-sialyllactose.

-Detection of influenza virus by potential measurement Using a potential measuring device, the cumulative dose response of the binding reaction between influenza virus and sugar chain was measured. Potential measurements were performed using 10-channel circular Au microelectrodes (500 μm in diameter) (Goda, T .; Higashi, D .; Matsumoto, A .; Hoshi, T .; Sawaguchi, T .; Miyahara, Y. , Dual aptamer-immobilized surfaces for improved affinity through multiple target binding in potentiometric thrombin biosensing. Biosensors and Bioelectronics 2015, 73, 174-180.). In order to detect influenza virus (H1N1), 25 mol% EDOTAO was electropolymerized on the surface of a 10-channel circular Au microelectrode for 10 cycles in the same manner as described. Then, a 2,6-sialyllactose sugar chain or a 2,3-sialyllactose sugar chain was introduced by a glycoblotting reaction, respectively.

When the concentration of influenza virus is 0.015, 0.031, 0.062, 0.12, 0.25, 0.50, 1.0 HAU, the potential response of the modified Au electrode is high-imput impedance. Measurements were made using an electrometer (Keythley 6517B, manufactured by Keythley) with respect to the Ag / AgCl (in 3.3 mol / L KCl _aq ) reference electrode without DC bias. First, a modified 10-channel circular Au microelectrode (500 μm in diameter) was immersed in DPBS (0.01xDPBS pH 7.4, manufactured by Dalbecco) diluted 100-fold with water for 150 minutes to stabilize the Au electrode surface potential. .. Then, the concentration of influenza virus was introduced from low concentration to high concentration at intervals of 10 minutes from 0.15 to 1.0 HAU, and the change in surface potential was measured (FIG. 16).

Since sialic acid at the end of the sugar chain has a negative charge and binds to the hemagglutinin recognition site (having a positive charge) of influenza virus (H1N1), the influenza virus concentration increases and the 2,6-sialyllactose sugar chain or 2 , The interfacial charges of the electrodes into which the 3-sialyllactose sugar chains were introduced tended to increase in both cases. However, in the electrode into which the 2,6-sialyllactose sugar chain was introduced, the interfacial potential was significantly increased as compared with the electrode into which the 2,3-sialyllactose sugar chain was introduced. This indicates that the influenza virus having more positive charges adhered to the electrode surface into which the 2,6-sialyllactose sugar chain was introduced. In a 0.01xDPBS solution, the electric double layer of the solution was about 7.4 nm. Since the length of the sugar chain is 2 nm or less, it can be detected in the electric double layer. The sugar chain has a negative charge and a positive net charge. Huang, Q .; Opitz, R .; Knapp, EW; Herrmann, A., Protonation and stability of the globular domain of influenza virus hemagglutinin. Biophysical Journal 2002, 82 (2), 1050-1058.), It is considered that the density of the plus net charge in the electric double layer became high and the potential increased.

We have developed a label-free detection method for influenza virus using an electrode modified with a conductive polymer in which a sugar chain that specifically binds to influenza virus is introduced at the end. A ligand molecule such as a sugar chain that specifically reacts with hemagglutinin, which is a protein on the surface of the virus, was introduced into PEDOT. After several synthetic steps, an oxylamine (-O-NH ₂ ) group was introduced as a substituent into the side chain of EDOT as a monomer, and the EDOT derivative (EDOTAO) was successfully synthesized. The synthesized monomer was evaluated by NMR and ESI-MS to confirm the synthesis of the target product. Next, in order to eliminate the possibility that the ligand molecule is affected by the electrolytic polymerization, EDOTAO was first introduced into the electrode substrate, and then a coupling reaction was carried out to introduce the ligand molecule. Using EDOT and EDOTAO, a copolymer was covalently introduced into the electrode surface by an electrolytic polymerization method. Glass carbon (GC) was used as a model electrode material as a base material. NaClO ₄ , which is a protonate, was used as a dopant species during electropolymerization. A sugar chain having a reducing end was introduced by a covalent bond by a glycoblotting reaction using an oxylamine group as a linker. 2,6-sialyllactose was used as a sugar chain specific for hemagglutinin of H1 influenza virus, and 2,3-sialyllactose was used as a sugar chain specific for hemagglutinin of H5 type. The prepared surface was evaluated by cyclic voltammetry (CV), AC impedance measurement (EIS), SEM, XPS, step measurement, and QCM. Particularly good results were obtained when a monomer having a composition of 25 mol% EDOTAO was electrolytically polymerized on the electrode. Influenza A (H1N1) virus was detected using the above electrodes. QCM measurement confirmed the specific binding of the virus to the surface of the 2,6-sialyllactose-introduced electrode. Furthermore, label-free electrical detection of viruses was realized by potential measurement. The sensitivity is hundreds of times higher than that of the conventional hemagglutination test, enabling rapid and real-time high-sensitivity specific detection of influenza virus.

The disclosure of Japanese patent application 2016-008084, filed January 19, 2016, is incorporated herein by reference in its entirety.
Further, the present invention is not limited to the description of the embodiments and examples of the above invention. Various modifications are also included in the present invention as long as those skilled in the art can easily conceive without departing from the description of the scope of claims.
All documents, patent applications, and technical standards described herein are to the same extent as if the individual documents, patent applications, and technical standards were specifically and individually stated to be incorporated by reference. Incorporated and incorporated herein.

<Explanation of symbols>
10 Biosensor device, 10a probe fixed electrode, 12 electrode (conductive layer), 13 gate electrode, 14 intermediate layer, 15 gate electrode, 16 oligosaccharide, 20 biosensor, 22S source, 24D drain, 28 gate insulating film , 32 Silicon substrate, 47 Gate insulating film

Claims (10)

A conductive layer, possess the oligosaccharides selectively capture a specific virus or bacteria, an intermediate layer coupling the oligosaccharide and the conductive layer, wherein the intermediate layer is,
(A) A self-assembled monolayer, the self-assembled monolayer is formed of a thiol compound immobilized on the conductive layer with a thiol group, and the thiol compound is formed via a disulfide bond2. Is it formed from a compound with a type of functional group; or
(B) Containing a conductive polymer, the conductive polymer contains a copolymer of 3,4-ethylenedioxythiophene (EDOT) and EDOT-ONH2 (EDOTAO).
Device for biosensor. The device for a biosensor according to claim 1 , wherein the self-assembled monolayer contains a thiol compound having an oxylamino group or a hydrazide group. The conductive polymer
(A) At least one of the oxylamino group or the hydrazide group and
(B) At least one selected from the group consisting of a thienyl group, a phenyl group, a pyrrole group, an aniline group, a fluorene group, and an indole group, and
The biosensor device according to claim 1 or 2 , comprising a structural unit formed from a monomer containing. The conductive polymer is formed by electrolytic polymerization or oxidative polymerization method, a device for biosensor of any one of claims 1 to 3. The biosensor device according to any one of claims 1 to 4 , wherein the oligosaccharide is an oligosaccharide that selectively captures a specific influenza virus. The biosensor device according to any one of claims 1 to 5 , wherein the conductive layer is a gate electrode of a field effect transistor or an electrode electrically connected to the gate electrode. A biosensor comprising the biosensor device according to any one of claims 1 to 6 . An oligo comprising a plurality of biosensor devices according to any one of claims 1 to 6, and the oligosaccharides of the biosensor devices selectively capture specific viruses or bacteria different from each other. The biosensor according to claim 7 , which is a sugar. The biosensor according to claim 7 or 8 , wherein the oligosaccharide of at least one biosensor device contains a sialic acid residue. The biosensor according to any one of claims 7 to 9 , which detects a virus or a bacterium based on any one change selected from an electric potential, a frequency, and a refractive index.