JP2020185336A - Electrode, manufacturing method for electrode, and biological signal measuring apparatus - Google Patents

Abstract

To provide an electrode high in adhesion between an insulation base material and a conductive polymer layer, and a manufacturing method for the electrode.SOLUTION: An electrode 1 has an insulation base material 10, a conductive intermediate layer 12 laminated on at least one surface of the insulation base material 10, and a conductive polymer layer 13 laminated on the surface opposite the insulation base material of the conductive intermediate layer 12. The conductive intermediate layer 12 has higher compatibility with the insulation base material 10 than the conductive polymer layer 13. The manufacturing method for the electrode has: a conductive intermediate layer forming step for forming a conductive intermediate layer 12 at least on one surface of the insulation base material 10; and a conductive polymer layer forming step for forming the conductive polymer layer 13 on the surface opposite the insulation base material of the conductive intermediate layer 12 by a plating method.SELECTED DRAWING: Figure 1

Description

The present invention relates to an electrode, a method for manufacturing the electrode, and a biological signal measuring device using the electrode.

In recent years, with the improvement of pacemakers, brain-machine interfaces (BMI), and measurement and control technology for electrical signals derived from muscle tissue, there is a demand for in-vivo implantable electrodes that can be used in a long-term and structurally stable environment even in a water-rich environment. Is increasing. Conventionally, in vivo electrical measurement has been widely performed using electrodes made of metal materials such as gold, silver, and indium tin oxide (ITO).

However, when an electrode made of a metal material is embedded in a living tissue, it is caused by the high rigidity, low hydrophilicity, and low growth of cells on the material surface of the metal material. As a result, long-term implantation for several months or longer may cause divergence from the living tissue in contact, making it difficult to obtain a signal. Therefore, it is necessary to replace the new electrode each time and insert the electrode, which causes problems such as damage to the inserted portion and mental load on the individual. Further, since the metal material has low light transmission, it is not possible to observe changes over time such as changes in the shape and mode of the living tissue in contact with the metal material, and the possibility of in-situ observation is strongly required. Furthermore, as the rarity of metal materials increases, the increase in manufacturing unit price has become a problem, and the development of alternative conductive materials having high biocompatibility and flexibility and low rarity has been actively carried out. ing.

Therefore, in recent years, it has been studied to use a conductive polymer as an electrode material for an implantable electrode in a living body. It is difficult to form a conductive polymer alone into an electrode shape. For this reason, in an implantable electrode in a living body using a conductive polymer as an electrode material, a conductive polymer layer is generally formed on the surface of an insulating base material. For example, Patent Document 1 describes that a conductive polymer fiber obtained by impregnating and / or adhering a conductor containing PEDOT-PSS to a base fiber is used as an in-vivo embedded electrode.

Republished WO 2013/073763

It is preferable that the implantable electrode in a living body has high adhesion so that the insulating base material and the conductive polymer layer do not peel off even if the electrode is distorted by the movement of the living body at the time of implanting the electrode. However, in the conventional electrode in which the conductive polymer layer is directly formed on the surface of the insulating base material, the insulating base material and the conductive polymer layer may be peeled off when excessive strain is applied. there were.

The present invention has been made in view of the above circumstances, and an object of the present invention is to provide an electrode having high adhesion between an insulating base material and a conductive polymer layer, and a method for manufacturing the electrode. The present invention also provides a biological signal measuring device using an electrode having high adhesion between a base material and a conductive polymer layer.

As a result of diligent research to solve the above problems, the present inventors have made an affinity between the insulating base material and the conductive polymer layer for the insulating base material rather than the conductive polymer layer. The present invention has been completed by finding that the adhesion between the insulating base material and the conductive polymer layer can be improved by providing the conductive intermediate layer having a high conductivity.

The electrode according to one aspect of the present invention includes an insulating base material, a conductive intermediate layer laminated on at least one surface of the insulating base material, and the insulating base material side of the conductive intermediate layer. Has a conductive polymer layer laminated on the surface on the opposite side, and the conductive intermediate layer is characterized in that it has a higher affinity for the insulating base material than the conductive polymer layer.

In the electrode according to one aspect of the present invention, the insulating base material is a flexible resin film, and the conductive intermediate layer has a structure showing affinity for the molecular structure of the flexible resin film. It may be configured to have.
Further, in the electrode according to one aspect of the present invention, the flexible resin film may contain a resin having a six-membered ring, and the conductive intermediate layer may have a six-membered ring structure.
Further, in the electrode according to one aspect of the present invention, the flexible resin film may contain a polyparaxylene resin, and the conductive intermediate layer may have a graphene structure.
Furthermore, in the electrode according to one aspect of the present invention, the conductive polymer layer is composed of poly (3,4-ethylenedioxythiophene), poly (3,4-ethylenedioxythiophene) -poly (4-). Styrene sulfonic acid), polythiophene polymer, polybithiophene polymer, polyisothiophene polymer, polydodecylthiophene polymer, polyisonite thiophene polymer, poly-3-hexylthiophene polymer, polyacene A configuration including one or more conductive polymers selected from the group consisting of a polymer, a polyparaphenylene polymer, a polyaniline polymer, a polydiacetylene polymer, a polypyrrole polymer, and a polyaniline polymer. It may have been.

The method for manufacturing an electrode according to one aspect of the present invention includes a step of forming a conductive intermediate layer on at least one surface of an insulating base material and a step of forming the conductive intermediate layer on the insulating base material side. On the surface opposite to the above, a conductive polymer layer forming step of forming a conductive polymer layer by a plating method is provided.

In the method for manufacturing an electrode according to one aspect of the present invention, in the conductive intermediate layer forming step, the conductive intermediate layer formed on the surface of the transfer substrate is transferred to at least one surface of the insulating base material. It may be configured to form a conductive intermediate layer.

The biological signal measuring device according to one aspect of the present invention has the electrode according to one aspect of the present invention described above.

According to the present invention, it is possible to provide an electrode having high adhesion between a base material and a conductive polymer layer, high durability, and a high-definition electrode and a method for manufacturing the electrode. Further, according to the present invention, it is possible to provide a biological signal measuring device using an electrode having high adhesion between a base material and a conductive polymer layer and high durability.

It is sectional drawing of the electrode which concerns on one Embodiment of this invention. It is sectional drawing explaining the manufacturing method of the electrode which concerns on one Embodiment of this invention. It is a graph which shows the relationship between the amount of electric charge which flowed per unit area when the PEDOT-PSS layer was formed by the plating method in Example 1, and the thickness of the formed PEDOT-PSS layer. It is a graph which shows the light transmittance of the electrode produced in Example 1. 6 is a graph showing the relationship between the contact time and the contact angle between the PEDOT-PSS layer of the electrode produced in Example 1 and the droplet. It is a graph which shows the relationship between the voltage value and the current value when the voltage is applied in the plane direction of the electrode produced in Example 1. It is a measurement result of cyclic voltammetry of the electrode produced in Example 1. It is a perspective view of the conical electrode produced in Example 1. FIG. It is an immunostaining image of the brain tissue of a rat into which a conical electrode was inserted. 3 are SEM images before and after laminating the PEDOT-PSS layer of the electrode produced in Example 2. It is a phase difference image before and after laminating the PEDOT-PSS layer of the electrode produced in Example 2.

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

<Electrode>
Hereinafter, an electrode according to an embodiment of the present invention will be described.
FIG. 1 is a cross-sectional view of an electrode according to an embodiment of the present invention.

As shown in FIG. 1, the electrode 1 is composed of an insulating base material 10, a conductive intermediate layer 12 laminated on one surface (upper surface in FIG. 1) of the insulating base material 10, and a conductive intermediate layer 12. It has a conductive polymer layer 13 laminated on the surface opposite to the insulating base material 10 side. The conductive intermediate layer 12 is made of a material having a higher affinity for the insulating base material 10 than the conductive polymer layer 13.

The insulating base material 10 is a flexible resin film 11. Although the flexible resin film 11 varies depending on the application, it is preferable that the flexible resin film 11 has enough flexibility to process the electrode 1 into an arbitrary shape such as a cylinder or a cone as a whole. The material of the flexible resin film 11 is not particularly limited, and can be appropriately selected in consideration of the combination with the material of the conductive intermediate layer 12 as described below.

The conductive intermediate layer 12 acts as a bonding layer (linker) for bonding the flexible resin film 11 and the conductive polymer layer 13 so as not to be peeled off from each other. The conductive intermediate layer 12 preferably has a structure that exhibits an affinity for the molecular structure of the flexible resin film 11. The structure showing affinity is, for example, a structure that improves the adhesion between the flexible resin film 11 and the conductive intermediate layer 12 by intermolecular interaction with respect to the molecular structure of the flexible resin film 11. is there. Examples of intermolecular interactions include van der Waals forces, dispersion forces (π-π interactions), electrostatic forces, and hydrogen bonds.

For example, when the flexible resin film 11 contains a resin having a six-membered ring, the conductive intermediate layer 12 preferably has a six-membered ring structure. The six-membered ring contained in the flexible resin film 11 is preferably a conjugated double-bonded six-membered ring. The six-membered ring structure of the conductive intermediate layer 12 is preferably a graphene structure. In this case, since the flexible resin film 11 and the conductive intermediate layer 12 are rich in electrons delocalized by the π-electron system of the six-membered ring, the π-π interaction is induced at the interface. The adhesion between the flexible resin film 11 and the conductive intermediate layer 12 is improved. The π-π interaction has stronger adhesion than van der Waals forces and hydrogen bonds. Therefore, it is possible to form a laminate of the flexible resin film 11 and the conductive intermediate layer 12 which are difficult to peel off not only in the air but also in water.

As the resin having a six-membered ring contained in the flexible resin film 11, for example, a polyparaxylene resin can be used. As the material of the conductive intermediate layer 12 having a graphene structure, for example, a graphene-based carbon material such as graphene, graphene oxide, carbon nanotubes, and graphite can be used. The graphene-based carbon material is preferably single-layer graphene or 2 to 30-layer multi-layer graphene.

The conductive polymer layer 13 is composed of poly (3,4-ethylenedioxythiophene), poly (3,4-ethylenedioxythiophene) -poly (4-styrene sulfonic acid), polythiophene-based polymer, and polybithiophene-based polymer. Polymer, polyisothiophene polymer, polydodecylthiophene polymer, polyisonitethiophene polymer, poly-3-hexylthiophene polymer, polyacene polymer, polyparaphenylene polymer, polyaniline polymer It is preferable to include a conductive polymer selected from the group consisting of a molecule, a polydiacetylene polymer, a polypyrrole polymer and a polyaniline polymer. One of these conductive polymers may be used alone, or two or more of these conductive polymers may be used in combination. All of the above conductive polymers have high biocompatibility. Therefore, by using the above-mentioned conductive polymer as the material of the conductive polymer layer 13, the biocompatibility of the electrode 1 is improved.

The conductive polymer layer 13 preferably contains poly (3,4-ethylenedioxythiophene) -poly (4-styrenesulfonic acid), and in particular, PEDOT-PSS [poly (3,4-ethylenedioxythiophene). ) -Poly (styrene sulfonic acid)] is preferably contained. PEDOT-PSS has a structure in which PEDOT, which is a π-conjugated conductive polymer, and PSS, which is a polymer electrolyte, are composited, and is excellent in hydrophilicity and biocompatibility.

By using a conductive polymer having low biotoxicity and high biocompatibility as the material of the conductive polymer layer 13, cells can be cultured on the surface of the electrode 1 or pierced into a living tissue as an in vivo implantable electrode. When it enters, it can be used as a conductive interface that does not show toxicity to biological samples such as cells and biological tissues. In particular, the electrode 1 using the above-mentioned conductive polymer has a feature that the structure and activity of the biological sample are not changed even when the electrode 1 is in contact with the biological sample for a long period of one year or more. For this reason, the duration of in vivo implantation, which was previously limited to a short period of several months, can be further extended, which makes it possible to stably measure data on living organisms over a long period of time. Become. The electrode 1 using the above conductive polymer is an electrode for in vivo implantation for repairing nerve tissue such as tantrum or spinal cord injury, a brain machine interface (BMI), or an electrical signal measurement electrode derived from muscle tissue. It can be applied as a wide range of substrate elements such as a multi-electrode array (MEA) for measuring electrical signals derived from cells, an element for a fine flexible battery, and an element for an actuator using a conductive polymer.

Further, as the thickness of the conductive polymer layer 13 becomes thicker, the unevenness of the surface becomes larger, and the area of the electrode surface (the portion in contact with air or the solvent inside) forming the electric double layer increases, and the portion thereof is charged. Is more likely to be retained (capacitor effect). The electrode 1 in which the conductive polymer layer 13 has a capacitor effect can obtain electrode characteristics capable of measuring a weak signal that tends to be buried in noise, for example, an in-vivo electrical signal derived from the brain. Therefore, this electrode 1 is useful as an in-vivo implantable electrode for a biological signal measuring device.

The electrode 1 preferably has a wavelength region in which the transmittance is 30% or more in the wavelength region of light of 300 to 1300 nm. By having such light transmission, for example, when observing with a microscope, it can be used in all kinds of optical microscopes regardless of whether it is an upright type or an inverted type, and further, it depends on the wavelength using a fluorescence microscope or a confocal microscope. It can also be applied to less sexual observations. The electrode 1 having such light transmittance not only enables in-situ observation of the state of cells and biological tissues to be observed, but also makes it possible to combine an optical system such as optogenetics with a biomaterial. It can also be applied to operation and analysis methods. The electrode 1 preferably has a transmittance of 30% or more (particularly 40% or more) in the wavelength region of light of 300 to 450 nm.

The flexible resin film 11 of the electrode 1 preferably has a thickness in the range of 100 nm or more and 900 μm or less. It is preferable that the thickness of the conductive intermediate layer 12 is within the range of the thickness of one atom or more and the thickness of 50 atoms or less. The thickness of the conductive polymer layer 13 is preferably in the range of 10 nm or more and 1.5 μm or less. The electrode 1 in which the thicknesses of the flexible resin film 11, the conductive intermediate layer 12 and the conductive polymer layer 13 are within the above ranges is usually excellent in flexibility, and for example, various shapes such as cylindrical and conical. It becomes easy to process into the shape of. In addition, it has excellent light transmission and facilitates in-situ observation with an optical microscope.

In the electrode 1 of the present embodiment described above, the flexible resin film 11 and the conductive intermediate layer 12 have a high affinity due to the molecular structure, and the conductive intermediate layer 12 and the conductive polymer layer 13 are electrically connected. High affinity due to characteristics. Therefore, in the electrode 1 of the present embodiment, the flexible resin film 11 and the conductive polymer layer 13 are difficult to peel off even when strain is applied.

<Method of manufacturing electrodes>
Next, a method for manufacturing an electrode, which is an embodiment of the present invention, will be described.
FIG. 2 is a cross-sectional view illustrating a method for manufacturing an electrode according to an embodiment of the present invention.
The electrode manufacturing method of the present embodiment includes a substrate preparation step for film formation (FIG. 2 (a)), an insulating substrate forming step (FIG. 2 (b)), and a conductive intermediate layer forming step (FIG. 2 (FIG. 2)). c)), a conductive intermediate layer pattern forming step (FIGS. 2 (d) to (e)), a conductive polymer layer forming step (FIG. 2 (f)), and a peeling step (FIG. 2 (g)). Has.

In the film-forming substrate preparation step, the film-forming substrate 20 is prepared as shown in FIG. 2 (a). The film-forming substrate 20 has a release layer 21 formed on the surface (upper surface in FIG. 2A) on the side where the electrode 1 is formed. The peeling layer 21 is a layer that can be removed by a chemical or physical method, and is a layer for facilitating the peeling of the electrode 1 from the film forming substrate 20 by removing this layer after the electrode 1 is formed. Is.

It is desirable that the film-forming substrate 20 has a high surface flatness. Examples of the material of the film-forming substrate 20 include silicon, soda glass, quartz, magnesium oxide, and sapphire.

As the material of the release layer 21, for example, a metal dissolved by a predetermined etching solution or a physical gel capable of sol-gel transition can be used. As the physical gel, an alginate gel that can control the gel-sol transition depending on the presence or absence of calcium ions can be used. Further, as a gel that decomposes by changing light, heat, and pH, poly (N-isopropylacrylamide) (PNIPAM) or azobenzene-modified polymagel can be used.

In the film-forming substrate preparation step, as shown in FIG. 2B, the flexible resin film 11 is formed on the surface of the release layer 21 of the film-forming substrate 20. The method for forming the flexible resin film 11 is not particularly limited, and as a method for forming a resin film such as a chemical vapor deposition method (CVD method), a spin coating method, an inkjet printing method, a thermal vapor deposition method, a sputtering method, and an electrospray method. Various methods used can be used.

In the conductive intermediate layer forming step, as shown in FIG. 2C, the conductive intermediate layer 12 is formed on the surface of the flexible resin film 11. As a method for forming the conductive intermediate layer 12, a transfer method or a coating method can be used. The transfer method is a method in which a conductive intermediate layer is formed on the surface of a separately prepared transfer substrate, and the conductive intermediate layer formed on the surface of the transfer substrate is transferred to the surface of the flexible resin film 11. .. The transfer method is a useful method when single-layer or multi-layer graphene is used as the conductive intermediate layer 12. The coating method is a method in which a coating liquid in which powder of a material forming the conductive intermediate layer 12 is dispersed in a solvent is applied to the surface of the flexible resin film 11 and dried. The coating method is a useful method when graphene powder, graphene oxide powder, carbon nanotube powder, or graphite powder is used as the material of the conductive intermediate layer 12.

In the conductive intermediate layer pattern forming step, as shown in FIG. 2D, a physical mask 22 is formed on the surface of the conductive intermediate layer 12. The physical mask 22 can be formed by applying a photoresist to the surface of the conductive intermediate layer 12 and irradiating it with ultraviolet light by a lithography technique to pattern it. Next, the conductive intermediate layer 12 in the portion where the physical mask 22 is not formed is removed by etching. As the etching, an oxygen plasma etching method or the like can be used.

Next, as shown in FIG. 2E, the physical mask 22 is removed. As a result, the conductive intermediate layer 12 formed in a predetermined pattern is obtained.

In the conductive polymer layer forming step, as shown in FIG. 2 (f), the conductive polymer layer 13 is formed on the surface of the conductive intermediate layer 12. The conductive polymer layer 13 is formed by a plating method. The plating method is preferably an electroplating method. Specifically, the conductive polymer layer 13 is formed as follows. First, the conductive lamination substrate 20, the release layer 21, the flexible resin film 11, and the conductive intermediate layer 12 obtained in the conductive intermediate layer pattern forming step are laminated in this order. The body is immersed in a plating solution in which a conductive polymer is dissolved together with an auxiliary electrode. Next, a voltage is applied between the conductive laminate and the auxiliary electrode to deposit the conductive polymer on the surface of the conductive intermediate layer 12 to form the conductive polymer layer 13.

In the peeling step, as shown in FIG. 2 (g), the flexible resin film 11, the conductive intermediate layer 12, and the conductive polymer layer 13 are arranged in this order by dissolving and removing the peeling layer 21. The electrodes 1 laminated in 1 are peeled off from the film-forming substrate 20. The obtained electrode 1 is heat-treated, if necessary. For example, when PEDOT-PSS is used as the material of the conductive polymer layer 13, it is preferable to heat-treat it at a temperature of 80 ° C. for 20 minutes.

According to the electrode manufacturing method of the present embodiment, the conductive polymer layer 13 is formed on the surface of the conductive intermediate layer 12 by using the plating method, so that the conductive intermediate layer 12 and the conductive polymer layer 13 are formed. It is possible to manufacture an electrode 1 having high adhesion to and high resistance to strain.

Further, according to the electrode manufacturing method of the present embodiment, the shape and area of the conductive intermediate layer 12 can be arbitrarily adjusted by using the photolithography method. Therefore, it is possible to produce a pattern of the conductive polymer layer 13 having an arbitrary shape. In the present embodiment, a photolithography method is used as a method for forming the pattern of the conductive intermediate layer 12, but the pattern forming method for the conductive intermediate layer 12 is not limited to this. As a method for forming the pattern of the conductive intermediate layer 12, for example, an electron beam lithography method, a dry etching method, or an inkjet method can be used. However, in the photolithography method, by shortening the wavelength of the emitted light, it is possible to produce a pattern having a finer structure than the ejection type patterning method such as the inkjet method. Therefore, by using the photolithography method, it is possible to form the conductive polymer layer 13 having a finer pattern.

<Biomedical signal measuring device>
The biological signal measuring device according to the embodiment of the present invention has the above-mentioned electrode 1. As described above, the electrode 1 has high biocompatibility and is therefore used as an implantable electrode in the living body. The biological signal measuring device preferably includes an external device having a function of analyzing or recording the biological signal measured by the electrode 1. The biological signal measuring device of the present embodiment can be used as, for example, a measuring device for electroencephalogram, electromyography, electrocardiography, and the like.

Hereinafter, the present invention will be described in more detail with reference to Examples.

[Example 1]
A glass substrate was prepared as the film-forming substrate (length: 32 mm, width: 24 mm, thickness: 100 μm). An alginate gel layer was formed as a release layer on the surface of the glass substrate. The alginate gel layer was formed by forming a sodium alginate thin film on the surface of a glass substrate by a spin coating method, and then immersing the glass substrate in a solution containing a large amount of calcium to convert sodium alginate into calcium alginate.
Next, a polyparaxylene resin film was formed on the alginate gel layer as an insulating base material. The polyparaxylene resin film was formed by growing 1.6 g of a paraxylene dima on a substrate using a CVD method. The film thickness of the formed polyparaxylene resin film was 1 μm.

Next, a graphene single layer was formed as a conductive intermediate layer on the surface of the polyparaxylene resin film. The graphene single layer was formed by preparing a copper foil with a graphene single layer having a single layer of graphene and transferring the graphene single layer to the surface of a polyparaxylene resin film. Specifically, after applying a polymer thin film (polymethyl methacrylate (PMMA)) to the surface of the graphene single layer of the graphene single layer copper foil, the copper foil of the graphene single layer copper foil is coated with ferric chloride. The graphene monolayer was recovered by dissolving with a solution. After repeating washing of the graphene single layer on the water surface, the polymer thin film coated with the graphene single layer is layered so as to be in contact with the surface of the polyparaxylene resin film, and the graphene single layer is transferred to the surface of the polyparaxylene resin film. did. The graphene single-layered copper foil was produced by forming a single-layer graphene on the surface of the copper foil by using a CVD method.
In this way, a conductive laminate in which an alginate gel layer, a polyparaxylene resin film, and a graphene single layer were laminated in this order was obtained on the surface of the glass substrate.

Next, a PEDOT-PSS layer was formed as a conductive polymer layer on the surface of the graphene single layer of the obtained conductive laminate by a plating method. Specifically, a silver paste was applied to the end portion of the graphite single layer of the conductive laminate, and a conductive portion was formed to dry and take conduction. Next, a portion of the conductive laminate in which the conductive portion was not formed was immersed in a plating tank in which a PEDOT-PSS solution was stored as a plating solution. The conductive portion of the conductive laminate was connected to the working electrode (WE), and the auxiliary electrode (CE) and the reference electrode (RE) were immersed in the plating tank. A silver-silver chloride electrode was used as the reference electrode, and a platinum coil was used as the auxiliary electrode. Next, the working electrode was set to a voltage of + 0.8 V with reference to the reference electrode, and a PEDOT-PSS layer was formed as a conductive polymer layer by an electroplating method.

Next, the conductive laminate on which the PEDOT-PSS layer was formed was taken out from the plating tank, thoroughly washed with water, and then dried to form an alginate gel layer and a polyparaxylene resin film on the surface of the glass substrate. , A single layer of graphene and a PEDOT-PSS layer were laminated in this order to obtain a laminate. The voltage application time when forming the PEDOT-PSS layer was adjusted by the plating method to obtain a plurality of conductive laminates having PEDOT-PSS layers having different thicknesses. The obtained laminate is immersed in a chelating agent (ethylenediaminetetraacetic acid (EDTA) solution) to dissolve and remove the alginate gel layer, and the polyparaxylene resin film, the graphene monolayer, and the PEDOT-PSS layer are arranged in this order. The electrodes laminated in (1) were peeled off from the glass substrate.
The resulting electrode was heated at a temperature of 80 ° C. for 20 minutes.

The electrodes obtained as described above were evaluated for adhesion stability, PEDOT-PSS layer film thickness, light transmission, hydrophilicity, electrical characteristics, and biocompatibility by the following methods.

[Adhesion stability]
The entire obtained electrode was strained by buckling or twisting. The polyparaxylene resin film and the PEDOT-PSS layer are not peeled off or the PEDOT-PSS layer is not torn for any electrode having a different thickness of the PEDOT-PSS layer, and the adhesion stability against strain is excellent. It was confirmed that.
In addition, the electrodes were immersed in pure water for several days. When the electrode after immersion was taken out from pure water and observed, it was confirmed that the polyparaxylene resin film and the PEDOT-PSS layer were not peeled off and the adhesion stability in water was excellent.

[PEDOT-PSS layer film thickness]
The film thickness of the PEDOT-PSS layer was measured using a stylus profiler.
FIG. 3 is a graph showing the relationship between the amount of electric charge flowing per unit area when the PEDOT-PSS layer is formed by the plating method and the thickness of the formed PEDOT-PSS layer. In FIG. 3, the horizontal axis is the amount of charge flowing per unit area, and the vertical axis is the thickness of the PEDOT-PSS layer.
From the graph of FIG. 3, it can be seen that the amount of charge (voltage application time) flowing per unit area when the PEDOT-PSS layer is formed by the plating method has a high correlation with the thickness of the PEDOT-PSS layer. From this result, it can be seen that the PEDOT-PSS layer can be accurately managed by adjusting the amount of electric charge flowing per unit area when forming the PEDOT-PSS layer by using the plating method.

[Optical transparency]
The transmittance of light having a wavelength of 300 nm to 1300 nm was measured using a spectrophotometer. The result is shown in FIG.
FIG. 4 is a graph showing the light transmittance of the electrode produced in Example 1. In FIG. 4, the horizontal axis is the wavelength of light and the vertical axis is the transmittance. In FIG. 4, the polyparaxylene resin film is a polyparaxylene resin film alone, and the graphene / polyparaxylene laminate is a laminate in which a graphene single layer is transferred to the surface of the polyparaxylene resin film.

From the graph of FIG. 4, it can be seen that the polyparaxylene resin film and the graphene / polyparaxylene laminate show a transmittance of 80% or more with respect to light in the wavelength region of 300 nm to 1300 nm. In the case of the electrode in which the PEDOT-PSS layer is laminated, it was confirmed that the thickness of the PEDOT-PSS layer increases and the transmittance of light in the wavelength region of 600 nm or more decreases. On the other hand, in the wavelength region of 290 nm to 600 nm, the transmittance is maintained at 40% or more even when the thickness of the PEDOT-PSS layer is 1 μm (the amount of charge flowing per unit area: 666 μC / mm ² ). I understand.

[Hydrophilic]
In order to evaluate the hydrophilicity of the electrode, pure water was dropped on the surface of the PEDOT-PSS layer of the electrode, and the change over time in the contact angle between the PEDOT-PSS layer and the droplets formed on the surface was measured. The result is shown in FIG.
FIG. 5 is a graph showing the relationship between the contact time and the contact angle between the PEDOT-PSS layer of the electrode produced in Example 1 and the droplet. In FIG. 5, the horizontal axis is the contact time (seconds), and the vertical axis is the contact angle of the droplet. In FIG. 5, the hydrophilic glass plate is obtained by cleaning the surface of a commercially available glass plate and subjecting it to hydrophilic treatment. The hydrophobic glass plate is obtained by cleaning the surface of a commercially available glass plate and applying a water repellent treatment by applying a commercially available water repellent polymer. The polyparaxylene resin film is a 1 μm-thick polyparaxylene resin film deposited on the surface of a hydrophilic glass plate, and the graphene / polyparaxylene laminate is a single graphene on the surface of the polyparaxylene resin film. It is a laminated body in which layers are transferred. The PEDOT-PSS spin coating film is a PEDOT-PSS spin coating film formed on the surface of a hydrophilic glass plate by a spin coating method. After the heat treatment, the PEDOT-PSS spin coating film was heat-treated at a temperature of 80 ° C. for 20 minutes.

As shown in the graph of FIG. 5, the contact angle of the hydrophilic glass plate with the droplets was 12 degrees, and the contact angle of the hydrophobic glass plate with the droplets was 88 degrees. The contact angle of the polyparaxylene resin film formed by the hydrophilic glass plate with the droplets was 80 degrees or more, which was equivalent to that of the hydrophobic glass plate. Further, when a graphene single layer was transferred to the surface of the polyparaxylene resin film, the hydrophilicity was slightly improved, and the contact angle with the droplet was 70 degrees. On the other hand, the contact angle of the PEDOT-PSS spin coating film (before treatment at the heating plant) was 65 degrees immediately after the dropping of pure water, but the contact angle gradually decreased to 40 degrees with the passage of time. Then it became unstable. The contact angle of the PEDOT-PSS spin coating film (after heat treatment) was 45 degrees immediately after dropping pure water, but it was observed that the contact angle gradually decreased due to swelling containing water. It was observed that the final value was 5 degrees or less. The contact angle of the electrode produced in Example 1 was 13 degrees immediately after the dropping of pure water, but then increased to 22 degrees, then gradually decreased, and became stable at around 5 degrees, PEDOT-. It was confirmed that the hydrophilicity was equivalent to that of the PSS spin-coated film.

[Electrical characteristics]
(1) Current-voltage A silver paste was applied and dried on the surface of the PEDOT-PSS layer of the electrode so as to have a gap of 5 mm to form two silver terminals. Next, a two-terminal probe was connected to the silver terminal, a voltage was applied between the silver terminals, and the current value when the voltage was applied in the plane direction of the electrode was measured. The result is shown in FIG.
As shown in FIG. 6, as compared with the graphene / polyparaxylene laminate (GR) in which a graphene single layer is transferred to the surface of the polyparaxylene resin film, the electrodes of the electrodes increase in thickness of the PEDOT-PSS layer. It was confirmed that the current value at the same voltage value tends to increase, that is, the resistance value tends to decrease.

(2) Cyclic voltammetry The cyclic voltammetry of the electrode was measured. The result is shown in FIG.
From the cyclic voltammetry curve shown in FIG. 7, it was confirmed that the electrode produced in Example 1 had a capacitor effect when the electrode potential was linearly swept and the response current was measured.

[Biocompatibility]
(1) Compatibility with nerve cells The surface of the PEDOT-PSS layer of the electrode prepared in Example 1 and the surface of the graphene layer of the graphene / polyparaxylene laminate in which a graphene single layer is transferred to the surface of the polyparaxylene resin film. In addition, rat fetal hippocampal primary cultured cells were seeded and cultured, and the cell shape and neurite outgrowth length were measured 2 weeks after seeding. As a result, it was confirmed that there was no significant difference in cell shape and neurite outgrowth between the surface of the PEDOT-PSS layer of the electrode and the surface of the graphene layer of the graphene / polyparaxylene laminate.

(2) Compatibility with myocardial cells The surface of the PEDOT-PSS layer of the electrode prepared in Example 1 and the surface of the graphene layer of the graphene / polyparaxylene laminate in which a graphene single layer was transferred to the surface of the polyparaxylene resin film. In addition, rat fetal myocardial primary cultured cells were seeded and cultured, and the cell shape and cell mass formation 2 weeks after seeding and the cycle of spontaneous beating of the cells were measured. As a result, it was confirmed that there was no significant difference in cell shape, cell mass formation, and spontaneous beating cycle of cells between the surface of the PEDOT-PSS layer of the electrode and the surface of the graphene layer of the graphene / polyparaxylene laminate. It was.

(3) Compatibility with rat brain cells A conical electrode was prepared using the electrode prepared in Example 1. FIG. 8 shows a perspective view of the produced conical electrode.
As shown in FIG. 8, the conical electrode 30 has a conical shape such that the electrode 1 (PEDOT-PSS layer thickness: about 1.5 μm) is on the outside and the conductive polymer layer 13 (PEDOT-PSS layer) is on the outside. It is wrapped around. The conical electrode 30 has a maximum diameter of 3 mm and a length of 5 mm.

The prepared conical electrode was attached to the tip of a glass capillary and inserted into the rat brain. Two months after the insertion of the conical electrode, a rat brain section was obtained, and the obtained brain section was stained. Then, an immunostained image of the brain tissue of the rat in which the conical electrode was inserted was observed using an optical microscope. The immunostained image is shown in FIG.
From the immunostained image of FIG. 9, no dissociation was observed between the surface of the conical electrode and the brain tissue, and it was observed that brain cells were growing on the surface of the conical electrode.

[Example 2]
In the same manner as in Example 1, an alginate gel layer (release layer), a polyparaxylene resin film (insulating base material), and a graphene single layer (conductive intermediate layer) are placed on a glass substrate (deposition substrate). , Formed in this order. A positive photoresist was applied to the surface of the graphene single layer of the obtained conductive laminate, and a predetermined pattern of ultraviolet light was irradiated by a lithography technique to form a physical mask having a predetermined pattern shape. Then, using a plasma asher device, the portion where the physical mask was not formed was etched with oxygen plasma to remove the graphene simple substance layer and the polyparaxylene resin film in that portion. The physical mask was then removed.

Next, a PEDOT-PSS layer was formed as a conductive polymer layer on the surface of the graphene single layer of the etched conductive laminate by a plating method in the same manner as in Example 1, and after being sufficiently washed with water, It was dry. Then, the alginate gel layer of the obtained laminate was dissolved and removed, and the electrode in which the polyparaxylene resin film, the graphene single layer, and the PEDOT-PSS layer were laminated in this order was peeled off from the glass substrate. The resulting electrode was heated at a temperature of 80 ° C. for 20 minutes.

The surface of the obtained electrode on the PEDOT-PSS layer side was observed using an SEM (scanning electron microscope) and a retardation light microscope. FIG. 10 shows SEM images before and after laminating the PEDOT-PSS layer, and FIG. 11 shows phase difference images before and after laminating the PEDOT-PSS layer.
FIG. 10A is an SEM image before laminating the PEDOT-PSS layer (the surface of the graphene single layer of the etched conductive laminate), and FIG. 10B is an SEM image after laminating the PEDOT-PSS layer. It is an SEM image. FIG. 11A is a phase difference image before laminating the PEDOT-PSS layer (the surface of the graphene single layer of the etched conductive laminate), and FIG. 11B is a phase difference image after laminating the PEDOT-PSS layer. It is a phase difference image of.
As shown in FIGS. 10 and 11, it was confirmed that the PEDOT-PSS layer was formed in a predetermined pattern.

1 ... Electrode, 10 ... Insulating base material, 11 ... Flexible resin film, 12 ... Conductive intermediate layer, 13 ... Conductive polymer layer, 20 ... Deposition substrate, 21 ... Release layer, 22 ... Physical mask , 30 ... Conical electrode

Claims (8)

The conductive base material, the conductive intermediate layer laminated on at least one surface of the insulating base material, and the conductive intermediate layer laminated on the surface of the conductive intermediate layer opposite to the insulating base material side. An electrode having a polymer layer, wherein the conductive intermediate layer has a higher affinity for the insulating base material than the conductive polymer layer. The electrode according to claim 1, wherein the insulating base material is a flexible resin film, and the conductive intermediate layer has a structure having an affinity for the molecular structure of the flexible resin film. The electrode according to claim 2, wherein the flexible resin film contains a resin having a six-membered ring, and the conductive intermediate layer has a six-membered ring structure. The electrode according to claim 3, wherein the flexible resin film contains a polyparaxylene resin, and the conductive intermediate layer has a graphene structure. The conductive polymer layer is poly (3,4-ethylenedioxythiophene), poly (3,4-ethylenedioxythiophene) -poly (4-styrenesulfonic acid), polythiophene-based polymer, polybithiophene-based. Polymer, polyisothiophene polymer, polydodecylthiophene polymer, polyisonitethiophene polymer, poly-3-hexylthiophene polymer, polyacene polymer, polyparaphenylene polymer, polyaniline polymer The electrode according to any one of claims 1 to 4, which comprises one or more conductive polymers selected from the group consisting of a molecule, a polydiacetylene polymer, a polypyrrole polymer, and a polyaniline polymer. A conductive intermediate layer forming step of forming a conductive intermediate layer on at least one surface of an insulating base material,
A method for manufacturing an electrode, which comprises a conductive polymer layer forming step of forming a conductive polymer layer on the surface of the conductive intermediate layer opposite to the insulating base material side by a plating method. The electrode according to claim 6, wherein in the step of forming the conductive intermediate layer, the conductive intermediate layer formed on the surface of the transfer substrate is transferred to at least one surface of the insulating base material to form the conductive intermediate layer. Manufacturing method. A biological signal measuring device having the electrode according to any one of claims 1 to 5.