Conductive fiber

Abstract

To provide a conductive polymer fiber excellent in conductivity, strength in a dry state and a wet state and flexibility, and a bioelectrode having the same.SOLUTION: In a conductive polymer fiber 40, a base material fiber 41 is impregnated with and / or subjected to adhesion of a conductive material 42 containing a conductive polymer. The conductive polymer fiber 40 is characterized in that the conductive material is arranged in the state of adhering tightly to the base material fiber, among the plurality of base material fibers.SELECTED DRAWING: Figure 6

Description

[Technical field]

[0001] The present invention relates to a conductive polymer fiber, a method and apparatus for manufacturing a conductive polymer fiber, a bioelectrode, a biosignal measuring device, an implantable electrode, and a biosignal measuring device. This application claims priority based on Japanese Patent Application No. 2011-251524 filed in Japan on November 17, 2011, Japanese Patent Application No. 2012-185343 filed in Japan on August 24, 2012, Japanese Patent Application No. 2012-189102 filed in Japan on August 29, 2012, and Japanese Patent Application No. 2012-212998 filed in Japan on September 26, 2012, the contents of which are incorporated herein by reference. [Background technology]

[0002] Conventionally, conductive fibers include those coated with metals such as copper, those woven with carbon or thin metal wires, and those made of conductive polymers formed into strings. These conductive fibers are widely used in bioelectrodes, biointerfaces, antistatic clothing, and the like. However, conventional conductive materials such as metals and carbon are hydrophobic and hard. For this reason, there is a problem that they are not suitable for applications that contact the body surface or internal tissues of living organisms, which are rich in moisture and flexible. For example, when placing a bioelectrode on the body surface, if the bioelectrode is made of a hard and hydrophobic material, it is difficult to make direct electrical contact with the body surface. For this reason, it is necessary to prepare and use a conductive paste (jelly) separately to electrically connect the bioelectrode and the body surface. In recent years, the development of conductive fibers has been promoted by extruding an aqueous solution of PEDOT-PSS {poly(3,4-ethylenedioxythiophene)-poly(styrenesulfonic acid)}, a conductive polymer with particularly excellent electrical conductivity and hydrophilicity, from a nozzle into a coagulating bath of acetone to form threads. The practical application of such fibers is also being considered (for example, see Non-Patent Document 1).

[0003] However, when the conductive fiber made of the above-mentioned PEDOT-PSS is used in a high humidity environment, the PEDOT-PSS absorbs moisture, causing a decrease in strength (especially tensile strength). In addition, the conductive fiber made of PEDOT-PSS expands when it absorbs moisture, and shrinks when it dries. This causes cracks to form inside the fiber or the fiber breaks, resulting in a problem that the conductivity of the fiber is easily reduced or lost. Clothing may get wet from rain or sweat during use. Bioelectrodes and biointerfaces are used in environments with high humidity. Therefore, in order to utilize PEDOT-PSS, which has excellent conductivity and hydrophilicity, for a wide range of applications, it was necessary to solve the above problems.

[0004] In addition to the above problem that the strength of the conductive fiber made of PEDOT-PSS is significantly reduced when it contains water, there are the following problems. That is, the fiber manufactured by the wet-spinning method described in Non-Patent Document 1 is a thin fiber with a diameter of about 10 microns. Therefore, there is a problem that it is difficult to handle and has insufficient strength even when dry. Furthermore, the fiber is highly rigid and feels rough to the touch. Therefore, there is also an issue that it does not provide the flexibility required for applications such as clothing.

[0005] On the other hand, body surface-mounted bioelectrodes are widely used for recording bioelectrical signals such as electroencephalograms, event-related potentials, evoked potentials, electromyograms, and electrocardiograms, and for electrical stimulation of living organisms. (Hereinafter, body surface-mounted bioelectrodes may be simply referred to as bioelectrodes.)

[0006] Conventionally widely used bioelectrodes are composed of a metal electrode plate and a gel or paste containing an electrolyte solution. The basic structure of these bioelectrodes is to fix the metal electrode plate and the skin surface by using (applying) a gel or paste between the metal electrode plate and the skin surface. Wearing the bioelectrode constantly seals a specific position on the skin surface. For this reason, particularly with long-term continuous use, discomfort or itching due to sweating may occur, and contact dermatitis or bacterial infection may also occur. There is a need to solve such problems in the conventional technology.

[0007] In countries where the population is aging, there are an increasing number of cases where biosignals such as electrocardiograms are monitored for long periods of time in medical institutions or at home. As various skin functions of the elderly are reduced, highly adhesive electrodes using conventional adhesive tapes are prone to causing skin irritation and discomfort such as itching. Furthermore, there are many cases where wearers who suffer from dementia or nocturnal delirium remove the bioelectrodes, and a solution to this problem is required.

[0008] In conventional bioelectrodes, which are prone to the above problems, a gel or paste containing an electrolyte solution is used between the skin and a metal electrode plate. When placing a bioelectrode on the skin surface via a gel or paste, it is necessary to increase the contact area of ​​the electrode. This is because the conductivity of the gel or paste is not high, and it is necessary to reduce the electrode resistance by increasing the contact area with the skin. However, the increase in the contact area of ​​the electrode is also the main cause of the above problems. Thus, existing bioelectrode configurations that rely on electrolyte gels or pastes are uncomfortable to wear and make it difficult to further miniaturize and densify the electrodes.

[0009] On the other hand, implantable bioelectrodes are necessary to receive electrical signals from within the body accurately and efficiently with an external device, and to transmit electrical signals from an external device to the body. Signals such as action potentials and synaptic potentials of nerve cells are particularly weak. Therefore, many signals are difficult to measure or input unless electrodes are placed very close to the cells. In addition to the nervous system, implantable bioelectrodes are widely used in cardiac pacemakers and cochlear implants. Development of implantable bioelectrodes for brain-machine interfaces and other future human interfaces is also underway.

[0010] The living body is rich in water and electrolytes and is flexible. In contrast, conventional bioelectrodes for implantation in the body are made of hard, hydrophobic conductive materials such as metal or carbon. For this reason, there are problems with mechanical and electrochemical compatibility between conventional bioelectrodes and living tissue. In particular, mechanical stress at the interface between the bioelectrode and the biological tissue can cause inflammation, resulting in tissue damage (invasiveness).

[0011] When electrodes are embedded in biological tissues, particularly in the nervous tissue of the brain and spinal cord, the following problems may occur. In other words, inflammation caused by microscopic damage to the nervous tissue gradually spreads, resulting in degeneration and loss of nerve cells around the electrode, making measurement and stimulation (signal input) difficult. In particular, when electrodes are permanently embedded in nervous tissue, loss of nerve cells and glial scarring may occur, resulting in a decrease in the efficiency of electrical stimulation and a deterioration or loss of the measured waveform. Furthermore, loss of nerve cells may cause nerve dysfunction. For these reasons, a solution is required. [Prior art documents] [Non-patent literature]

[0012] [Non-Patent Document 1] “Spinning and Characterization of Conducting Microfibers” Macromol. Rapid Commun. (2003) 24, pp261-264 Summary of the Invention [Problem to be solved by the invention]

[0013] In this invention, in order to overcome the problems of the conductive fibers made of PEDOT-PSS, a composite material is proposed in which PEDOT-PSS, a conductive polymer, is fixed to the inside and outside of fibers such as silk or fiber bundles (yarns). Such conductive fibers made of a PEDOT-PSS composite material are expected to be particularly useful as a material for bioelectrodes, since they have electrical conductivity, hydrophilicity, tensile strength, and water resistance.

[0014] Until now, commercially available PEDOT-PSS materials (e.g., Clevios P ​​Heareus) have been supplied as a PEDOT-PSS solution, and are generally used by immobilizing them on a substrate. Thus, methods for immobilizing PEDOT-PSS can be broadly classified into two types: chemical and electrochemical. The chemical method is a relatively simple immobilization method that can immobilize PEDOT-PSS on various substrates, but its conductivity and strength are not as high as those of the electrochemical method described below. The electrochemical method is a method in which PEDOT-PSS is electrically polymerized and immobilized on the electrode surface, and is superior in conductivity and strength to the chemical method, but since it is necessary to pass electricity between the substrate and the solution, it is necessary to use a conductive material for the substrate.

[0015] Here, when manufacturing a composite fiber of PEDOT-PSS and fiber bundle as described above, since the fiber bundle as the base material is an insulator (non-conductive), an electrochemical method cannot usually be used and a chemical fixing method must be used. Specifically, it can be manufactured by spraying an organic solvent such as acetone, ethanol, or methanol, or an electrolyte solution such as magnesium oxide solution, or by immersing the fiber bundle impregnated with PEDOT-PSS in the liquid.

[0016] However, conductive polymer fibers made of fiber bundles of PEDOT-PSS produced by the chemical fixation method have a conductivity of only about 40 to 50 MΩ / cm for No. 9 silk thread (fiber bundle diameter approximately 280 microns), and furthermore, the adhesive strength between the PEDOT-PSS and the fiber bundle is low. This resulted in the disadvantage that PEDOT-PSS was prone to peeling and its conductivity was easily reduced, so further improvement was necessary.

[0017] In addition to the above, other techniques for fabricating PEDOT-PSS into fibers include spinning methods (wet spinning and electrospinning). However, it is technically difficult to produce PEDOT-PSS blended with long fiber bundles (yarns) by spinning.

[0018] In addition, as a method for manufacturing a composite fiber of PEDOT-PSS and a fiber bundle, a method is also conceivable in which the fiber bundle is first given electrical conductivity by applying a metal coating thereto, or PEDOT-PSS is fixed to the fiber bundle by a chemical fixing method to give electrical conductivity, and this electrical conductivity is then used as an electrode to electrochemically fix PEDOT-PSS. However, this method requires a two-step fixation process, which may result in poor productivity and high costs.

[0019] The composite fiber of PEDOT-PSS and silk or other fiber bundles as described above is a conductive material with excellent biocompatibility and is expected to be applied to bioelectrodes. Although such composite fibers can be produced by the chemical fixation method described above, further improvements in conductivity and durability, as well as efficient production, are required.

[0020] In addition, a wide range of fields, including medicine, health promotion, information technology, and wearable computers, are in demand for surface-mounted bioelectrodes that can be used continuously for long periods of time. Not only are there demands for stable and reliable measurement of bioelectrical signals, but there is also a demand for comfortable wear. There is also a demand for implantable electrodes that are minimally invasive to biological tissue.

[0021] The inventors of the present invention focused on the fact that conductive polymers, typified by PEDOT-PSS, which have been developed recently, are highly hydrophilic and flexible, and believed that the stress on biological tissues when a conductive polymer is embedded in the body as an electrode could be reduced. As a result of extensive research, they completed the present invention. The first to fourth aspects of the present invention have been made in consideration of the above-mentioned problems of the conventional art and through various investigations by the present inventors. In the present invention, the conductive polymer fiber of the first aspect can be preferably produced by the device and method of the second aspect. Furthermore, the electrodes and devices of the third and fourth aspects can preferably use the conductive polymer fiber of the first aspect.

[0022] A first aspect of the present invention aims to provide a conductive polymer fiber having excellent electrical conductivity, strength in a dry state and in a wet state, and flexibility, and a bioelectrode including the conductive polymer fiber.

[0023] A second aspect of the present invention aims to provide a method and apparatus for producing conductive polymer fibers, which can impregnate or attach a conductor containing PEDOT-PSS as a conductive polymer to an insulating fiber (fiber bundle) and continuously polymerize and fix it electrochemically, and which can efficiently produce conductive polymer fibers that have high biocompatibility, good homogeneity, and excellent conductivity and durability.

[0024] A third aspect of the present invention aims to provide a bioelectrode that has stability and reliability when measuring bioelectrical signals and has a more comfortable fit than conventional bioelectrodes, and a biosignal measuring device equipped with such a bioelectrode.

[0025] A fourth aspect of the present invention aims to provide an implantable electrode capable of detecting weak electrical signals within a living body, having excellent biocompatibility and low invasiveness to biological tissue, and a biosignal measuring device equipped with such an implantable electrode. [Means for solving the problem]

[0026] The present invention relates to a bioelectrode using a conductive fiber containing a conductive polymer, the conductive fiber being formed by impregnating and / or attaching a conductor containing a conductive polymer to a base fiber, and the conductive polymer is hydrophilic. (First aspect) A first aspect related to the present invention provides the following conductive polymer fiber. I-(A): A conductive polymer fiber in which a conductor containing a conductive polymer is impregnated and / or attached to a base fiber, and the conductor is held and fixed between a plurality of the base fibers. I-(B): A conductive polymer fiber in which a conductor containing a conductive polymer is impregnated and / or attached to a base fiber, and the conductor is disposed between a plurality of the base fibers in close contact with the base fiber.

[0027] As a conductive polymer fiber related to the conductive polymer fiber of the first embodiment described above, a conductive polymer fiber having the following characteristics can be mentioned. I-(1): The conductive polymer fiber related to the first aspect described above is characterized in that a conductor containing a conductive polymer is impregnated and / or attached to a base fiber containing an animal fiber, and the conductive polymer is PEDOT-PSS. I-(2): In the above conductive polymer fiber, the conductor contains glycerol, polyethylene glycol-polypropylene glycol copolymer, ethylene glycol, sorbitol, sphingosine or phosphatidylcholine as an additive.

[0028] I-(3): In the conductive polymer fiber of I-(1) or (2), the conductor is coated around the base fiber.

[0029] I-(4): In the conductive polymer fiber of any one of I-(1) to I-(3), the conductor is impregnated into the base fiber.

[0030] I-(5): The conductive polymer fiber of any of I-(1) to (4) has the conductor impregnated into the base fiber, the base fiber is coated with metal or carbon, and the conductor is further coated on the coated metal or carbon.

[0031] I-(6): In the conductive polymer fiber of any of I-(1) to (5), the conductor is disposed between a plurality of the base fibers in close contact with the base fibers.

[0032] I-(7): In any one of the above I-(1) to (5), the conductive polymer fiber is further coated with an insulating layer around the conductive polymer fiber.

[0033] I-(8): A bioelectrode related to the present invention is characterized in that it comprises, as an electrode, the conductive polymer fiber described in any one of I-(1) to (7).

[0034] (Second Aspect) A second aspect related to the present invention provides the following method and apparatus for producing a conductive polymer fiber. II-(1): The method for producing the conductive polymer fiber described above is characterized by comprising a dipping step of dipping an insulating base fiber consisting of a thread-, string-, cloth- or ribbon-shaped fiber bundle into a conductor solution containing PEDOT-PSS {poly(3,4-ethylenedioxythiophene)-poly(styrenesulfonic acid)} as a conductive polymer, thereby impregnating and / or adhering the base fiber with the conductor, and a fixing step of electrochemically polymerizing and fixing the conductor impregnated and / or attached to the base fiber by running the base fiber between electrodes while pulling the base fiber vertically out of the conductor solution and passing an electric current through the base fiber. The above-mentioned manufacturing method may sequentially include the immersion step, the fixing step, and a drying step of blowing air to dry the base material fiber to which the conductor has been polymerized and fixed, and may further be characterized in that each of the immersion step, the fixing step, and the drying step is performed while controlling the atmospheric humidity. The electrode referred to here includes a single (unipolar) electrode as well as a configuration of multiple electrodes. It is also preferable that the second aspect II-(1) above has the following features.

[0035] II-(2): In the second embodiment of the method for producing conductive polymer fibers, not only a single electrode but also an array of a plurality of electrodes can be used in the fixing step. Specifically, with regard to the fixing process, multiple electrodes are used as the electrodes, and the multiple electrodes are comb-tooth electrodes having multiple comb teeth in the longitudinal direction of the base material fiber, the comb-tooth electrodes are arranged so as to sandwich the base material fiber from both radial sides of the base material fiber, and the multiple comb teeth are arranged so as to be alternately combined in the longitudinal direction of the base material fiber from both radial sides of the base material fiber, and the method may also include running the base material fiber while pressing and guiding the multiple comb teeth provided on the comb-tooth electrodes against the base material fiber from both radial sides to pass an electric current through the base material fiber.

[0036] II-(3): In the second aspect of the method for producing a conductive polymer fiber, in the fixing step, a plurality of electrodes are used as the electrodes, and the plurality of electrodes are arranged in the longitudinal direction of the base fiber, and are rotor electrodes arranged so as to sandwich the base fiber from both radial sides of the base fiber, the rotor electrode arranged on one side in the radial direction of the base fiber is roller-shaped, and the rotor electrode arranged on the other side is pulley-shaped, the rotor electrodes arranged on both sides of the base fiber are alternately arranged in the longitudinal direction of the base fiber, and it is preferable that the base fiber is made to run between the plurality of electrodes while being guided by a groove formed in the pulley-shaped rotor electrode while the roller-shaped rotor electrode is pressed against the base fiber to pass a current.

[0037] II-(4): In the manufacturing method of the conductive polymer fiber of the second aspect described above, in the fixing step, it is more preferable to adjust the amount of the conductor impregnated into and / or attached to the base fiber while passing electricity through the base fiber, by pressing the base fiber against the roller-shaped rotor electrode and guiding it in the groove of the pulley-shaped rotor electrode to adjust the arrangement shape of the fiber bundle.

[0038] The second aspect also provides the following manufacturing device for conductive polymer fibers. II-(5): The manufacturing apparatus for conductive polymer fibers of the second aspect described above is characterized by comprising: an immersion container in which a conductive solution containing PEDOT-PSS {poly(3,4-ethylenedioxythiophene)-poly(styrenesulfonic acid)} as a conductive polymer is contained, and an insulating base fiber consisting of a thread-like, string-like, cloth-like or ribbon-like fiber bundle is immersed in the conductive solution to impregnate and / or adhere the conductor to the base fiber; a winding section for vertically pulling up the base fiber from the conductive solution contained in the immersion container; an electrode for electrochemically polymerizing and fixing the conductor impregnated and / or adhered to the base fiber by passing an electric current through the base fiber pulled up vertically while running the base fiber; a drying section for drying the base fiber to which the conductor has been polymerized and fixed by blowing air; and a humidity control section for adjusting the atmospheric humidity in the vicinity of the base fiber.

[0039] The above device also preferably has the following features: II-(6): The electrode is composed of a plurality of electrodes, and the plurality of electrodes are comb-tooth electrodes having a plurality of comb teeth arranged in the longitudinal direction of the base material fiber, the comb-tooth electrodes are arranged so as to sandwich the base material fiber from both radial sides of the base material fiber, and the plurality of comb teeth are arranged so as to be alternately combined in the longitudinal direction of the base material fiber from both radial sides of the base material fiber, and it is also preferable to adopt a configuration in which the base material fiber is run while the plurality of comb teeth provided on the comb-tooth electrodes are pressed against and guided from both radial sides of the base material fiber to pass electricity through it.

[0040] II-(7): In the manufacturing apparatus for conductive polymer fibers of the second aspect described above, the electrode is preferably composed of a plurality of electrodes, the plurality of electrodes being arranged in the longitudinal direction of the base fiber, and being rotor electrodes arranged so as to sandwich the base fiber from both radial sides thereof, the rotor electrode arranged on one side of the radial direction of the base fiber being roller-shaped, and the rotor electrode arranged on the other side being pulley-shaped, the rotor electrodes arranged on both sides of the base fiber being alternately arranged in the longitudinal direction of the base fiber, and the base fiber being made to run between the plurality of electrodes while being guided by a groove formed in the pulley-shaped rotor electrode while the roller-shaped rotor electrode is pressed against the base fiber, to pass a current through the base fiber.

[0041] II-(8): In the manufacturing device for conductive polymer fibers of the second aspect described above, it is more preferable that the plurality of electrodes are configured such that the roller-shaped rotor electrode is pressed against the base fiber while being guided by the groove of the pulley-shaped rotor electrode to adjust the arrangement shape of the fiber bundle, thereby adjusting the amount of the conductor impregnated into and / or attached to the base fiber while passing electricity through the electrode.

[0042] (Third Aspect) A third aspect related to the present invention is the following bioelectrode. III-(1): A bioelectrode characterized by using a conductive composite fiber containing a conductive polymer. It is also preferable that the bioelectrode has the following features.

[0043] III-(2): The bioelectrode according to III-(1), characterized in that it is provided with a string-, band- or cloth-shaped contactor made of the conductive composite fiber.

[0044] III-(3): The contactor is a bioelectrode described in III-(2), characterized in that the contactor is formed by bundling a plurality of the conductive composite fibers or by winding the conductive composite fibers around a metal wire.

[0045] III-(4): The bioelectrode according to any one of III-(1) to III-(3), wherein the conductive composite fiber has adsorptivity or hydrophilicity to the skin.

[0046] III-(5): A bioelectrode according to any one of III-(2) to (4), characterized in that the contact is supported by a bow-shaped or hairpin-shaped frame.

[0047] III-(6) The bioelectrode according to any one of III-(2) to (4), wherein the contact is disposed on a surface of a sheet-like substrate.

[0048] III-(7): A bioelectrode according to III-(6), characterized in that an elastic holder is provided on the back surface of the sheet-like substrate, and the holder is arranged so as to be slidable along the back surface.

[0049] A third aspect related to the present invention includes the following biosignal measuring device. III-(8): A biosignal measuring device comprising the bioelectrode according to any one of III-(1) to III-(7).

[0050] (Fourth aspect) A fourth aspect related to the present invention is the following electrode. IV-(1): An implantable electrode comprising a conductive composite fiber containing a conductive polymer. It is also preferable for the implantable electrode in accordance with the present invention to have the following features:

[0051] IV-(2) The implantable electrode according to IV-(1) above, wherein the conductive composite fiber is formed into a rod-like or coil-like shape.

[0052] IV-(3): An implantable electrode according to IV-(1) or (2) above, characterized in that the conductive composite fiber is bonded to the tip of a needle.

[0053] IV-(4): The implantable electrode according to IV-(3), characterized in that the conductive composite fiber is adhered to the needle via a water-soluble adhesive material.

[0054] IV-(5): The implantable electrode according to any one of IV-(1) to (4), wherein the conductive composite fiber is in a dry shrinkage state.

[0055] IV-(6): An implantable electrode according to any one of IV-(1) to (5), characterized in that a metal, silicon or carbon electric wire is connected to the conductive composite fiber.

[0056] IV-(7): The implantable electrode according to any one of IV-(1) to IV-(6), characterized in that the conductive composite fiber formed into a rod or string shape is used as a core, and at least a portion of the core is coated with a water-resistant polymer, thereby forming a flow path for liquid to permeate from one end of the coated core to the other end.

[0057] IV-(8): The implantable electrode according to IV-(7), wherein the flow path contains a solution containing a drug.

[0058] IV-(9): A biosignal measuring device comprising the implantable electrode according to any one of IV-(1) to (8).

[0059] The conductive composite fibers described in IV-(1) to (8) may contain one or more of glycerol, sorbitol, ethylene glycol, squalane, silicone, mineral oil, and MPC (2-methacryloyloxyethyl phosphorylcholine). By containing these, the rate at which the conductive composite fiber absorbs water in biological tissue can be slowed down, and the swelling rate of the conductive composite fiber can be slowed down. As a result, the conductive composite fiber can be prevented from swelling during the operation of embedding the conductive composite fiber in biological tissue, and the mechanical strength of the conductive composite fiber can be prevented from decreasing.

[0060] The conductive composite fiber described in IV-(7) may be connected to a reservoir or chamber capable of containing a drug solution at the one end. By putting a solution containing one or more drugs, such as glycerol, sorbitol, mannitol, fructose, NGF (Nerve Growth Factor), BDNF (Brain-derived neurotrophic factor), SKF96365, Cilostazol, Gadolinium, NT3 (Neurotrophin-3), GSNO (S-Nitrosoglutathione), Magnesium, TRIM (1-(2-Trifluoromethylphenyl)imidazole), EGTA (ethylene glycol tetraacetic acid), or Ruthenium Red, into the reservoir or chamber, the solution can be permeated from the one end to the other end of the coated conductive composite fiber. That is, the solution can be released from the other end in the biological tissue, and the drug or the like contained in the solution can be administered locally around the conductive composite fiber.

[0061] One or more of the drugs may be impregnated or coated on the conductive composite fiber described in any one of IV-(1) to (8) above. In this case, the drug is gradually released from the conductive composite fiber placed in the biological tissue, and the drug can be administered locally around the conductive composite fiber.

[0062] A thread may be connected to the conductive composite fiber described in IV-(1), (2), (5), (6) or (7). Furthermore, a surgical needle may be attached to the thread. The thread is first introduced into a living body, and the thread connected to the thread can be smoothly introduced into the living tissue by pulling the thread. Effect of the Invention

[0063] The present invention can provide the following excellent effects. (Effect of the first aspect) According to a first aspect of the present invention, it is possible to provide a conductive polymer fiber having excellent conductivity, strength in a dry state and in a wet state, and flexibility, and a bioelectrode including the conductive polymer fiber. According to the conductive polymer fiber described in I-(1) of the present invention, a fiber can be obtained that has both the high strength and flexibility of the base fiber and the electrical conductivity and hydrophilicity of the conductive polymer PEDOT-PSS. According to the configuration of I-(2), the additive suppresses the water absorption of PEDOT-PSS and prevents a decrease in strength due to the wet state, resulting in a conductive polymer fiber with higher strength. According to the configuration of I-(3), the conductor covers the base fiber, thereby further increasing the conductivity of the conductive polymer fiber and making it easier to bring multiple conductive polymer fibers into contact with each other and establish electrical conductivity. According to the configuration of I-(4), since the conductor is impregnated into the base fiber, there is no risk of the conductor and the base fiber separating, resulting in a fiber with excellent long-term reliability. According to the configuration of I-(5), the metal or carbon is disposed inside the conductive polymer fiber and sandwiched between the surrounding conductors, resulting in a fiber with even higher conductivity. In addition, the metal or carbon is not exposed on the fiber surface, so corrosion and deterioration of the metal or carbon is prevented. According to the configuration of I-(6), the conductor covers the base fiber, thereby further increasing the conductivity of the conductive polymer fiber and making it easier to bring multiple conductive polymer fibers into contact with each other and establish electrical conductivity. According to the configuration of I-(7), the conductive polymer fiber is protected by the insulating layer, resulting in a fiber with excellent durability. The bioelectrode of I-(8) of the present invention is provided with conductive polymer fibers that are excellent in electrical conductivity, strength in both dry and wet states, and flexibility. This allows for a high degree of freedom in the location of placement on the body surface or inside the body of a living organism, excellent workability during placement, and allows for sufficient electrical measurements, as well as measurements over a relatively long period of time.

[0064] (Effect of the second aspect) According to the second aspect of the present invention, the following effects can be obtained. According to the second aspect of the present invention, the method for producing a conductive polymer fiber of the present invention employs a method in which, as described above, a base fiber impregnated with and / or attached to a conductor containing PEDOT-PSS is vertically pulled up from a conductor solution and run between one or more electrodes to pass electricity through it. This allows the process of electrochemically polymerizing and fixing the conductor to the base fiber to be performed continuously in a single step, improving productivity. Furthermore, by running the base fiber between the electrodes while pulling it up vertically, the conductor to be polymerized and fixed to the base fiber can be evenly dispersed and uneven distribution can be prevented. Therefore, it is possible to produce a conductive polymer fiber with high biocompatibility, good homogeneity, and excellent conductivity and durability with good productivity.

[0065] In addition, according to the second aspect of the present invention, the conductive polymer fiber manufacturing apparatus employs a configuration including a winding section that vertically pulls up the base fiber impregnated and / or attached with a conductor including PEDOT-PSS from the conductor solution in the immersion container, and one or more electrodes that pass a current while the base fiber is running. As a result, as in the above, the process of electrochemically polymerizing and fixing the conductor to the base fiber can be continuously performed in a single step, thereby improving productivity. Furthermore, by providing an electrode that passes a current while pulling up the base fiber vertically, the conductor to be polymerized and fixed to the base fiber can be evenly dispersed and uneven distribution can be prevented. Therefore, a conductive polymer fiber having high biocompatibility, good homogeneity, and excellent conductivity and durability can be obtained with high productivity. More specifically, the following effects can be obtained. According to II-(1) to (8) of the second embodiment, the following effects can be obtained.

[0066] According to the method for producing a conductive polymer fiber having the configuration of II-(1), a method is adopted in which a base fiber impregnated with and / or attached to a conductor including PEDOT-PSS is vertically pulled up from a conductor solution while running between electrodes and an electric current is applied, thereby electrochemically polymerizing and fixing the conductor to the base fiber. Furthermore, according to this configuration, by pulling the base fiber vertically while running it between the electrodes, the conductor to be polymerized and fixed to the base fiber can be evenly dispersed and uneven distribution can be prevented, resulting in a conductive polymer fiber with excellent conductivity and durability.

[0067] In the case of II-(2), the process of electrochemically polymerizing and fixing the conductor to the base fiber can be carried out continuously in a single step, thereby improving productivity. In addition, according to this configuration, a comb-tooth electrode having multiple comb teeth arranged in the longitudinal direction of the base material fiber is used, and the multiple comb teeth are arranged so that they are alternately combined in the longitudinal direction of the base material fiber from both radial sides of the base material fiber.Therefore, the base material fiber repeatedly comes into contact with the multiple comb teeth to pass electricity through it, thereby increasing the efficiency of polymerization fixation of the conductor to the base material fiber.

[0068] According to the configuration of II-(4), a method is adopted in which the base fiber is pressed against the roller-shaped rotor electrode and guided by a groove formed in the pulley-shaped rotor electrode, and the base fiber is run between the electrodes to pass an electric current, thereby reducing friction caused by contact between the base fiber and the electrodes and preventing the conductor fixed to the base fiber from peeling off. In addition, by adjusting the arrangement shape of the fiber bundles and passing an electric current while adjusting the amount of conductor impregnated and / or attached to the base fiber, any amount of PEDOT-PSS can be held and polymerized and fixed.

[0069] According to the conductive polymer fiber manufacturing apparatus of II-(5), the apparatus is configured to include a winding section that vertically pulls up the base fiber impregnated and / or attached with a conductor including PEDOT-PSS from the conductor solution in the immersion container, and an electrode that applies current while running the base fiber, so that the conductor is electrochemically polymerized and fixed to the base fiber. Furthermore, according to this configuration, by providing an electrode that applies current while pulling up the base fiber vertically, the conductor to be polymerized and fixed to the base fiber can be evenly dispersed and uneven distribution can be prevented, so that a conductive polymer fiber with excellent conductivity and durability can be manufactured. In the conductive polymer fiber manufacturing apparatus of the present invention, as for the electrode, not only a configuration using a single electrode but also a configuration in which multiple electrodes are arranged can be adopted. In this case, the electrochemical polymerization and fixation of the conductor to the base fiber can be performed continuously in one step, so that productivity is improved.

[0070] According to the configuration of II-(6), comb-tooth electrodes are provided as the multiple electrodes, and the multiple comb teeth are arranged so that they are alternately combined in the longitudinal direction of the base fiber from both radial sides of the base fiber. Therefore, by repeatedly contacting the base fiber with the multiple comb teeth and passing electricity through it, it is possible to increase the efficiency of polymerization fixation of the conductor to the base fiber.

[0071] According to the configuration of II-(8), the multiple electrodes are composed of a roller-shaped rotor electrode arranged on one side in the radial direction of the base fiber and a pulley-shaped rotor electrode arranged on the other side, so that the base fiber is run between the multiple electrodes while being guided by the groove formed in the pulley-shaped rotor electrode while pressing the roller-shaped rotor electrode against the base fiber, thereby reducing friction caused by contact between the base fiber and the electrodes and preventing the conductor fixed to the base fiber from peeling off. In addition, by adjusting the arrangement shape of the thread-like fiber bundles with the above-mentioned rotor electrode, it is possible to hold and polymerize and fix any amount of PEDOT-PSS by applying electricity while adjusting the amount of conductor impregnated and / or attached to the base fiber.

[0072] (Effect of the third aspect) According to the third aspect of the present invention, the following effects can be obtained. The body surface-mounted bioelectrode of the third aspect of the present invention includes a composite fiber made of a conductive polymer and a flexible fiber material. This improves the ease of installation on the body surface, and allows the electrode to be made smaller and the contact area with the skin to be reduced. In addition, since the bioelectrode of the present invention is made of a flexible fiber material, it causes less irritation to the skin when worn and is less likely to cause discomfort during wearing. Furthermore, since the composite fiber constituting the bioelectrode of the present invention has appropriate adhesiveness to the skin, it is not necessary to seal the skin and the electrode with a highly adhesive gel or tape, as in the case of conventional bioelectrodes. In other words, the bioelectrode of the present invention can reduce the burden on the subject (wearer) and provide a comfortable fit. More specifically, the following effects can be obtained.

[0073] The bioelectrode and device according to the third aspect can provide the following effects. The III-(1) bioelectrode uses conductive composite fibers that are conductive, flexible, and have excellent tensile strength. Compared with conventional bioelectrodes, this electrode causes less discomfort and damage to the wearer when worn, has low electrode resistance per unit area, enables precise transmission of biosignals, and is suitable for miniaturization and weight reduction. In III-(2), the shape of the contact is suitable for contact with the skin or the surface of a living body, so that more precise transmission of biological signals is possible. In addition, discomfort and damage to the wearer are further reduced. In III-(3), the bioelectrode is provided with a contact having the above-mentioned configuration, so that the structural strength of the contact and the bioelectrode is increased, and the electrode resistance can be reduced. In III-(4), the conductive composite fiber has an adhesive property or hydrophilicity to the skin, so that the bioelectrode can be installed independently on the skin surface or the surface of the living body, and it is possible to more easily transmit a highly accurate biosignal with reduced noise. In addition, the bioelectrode can be installed without using a paste or gel having adhesiveness and conductivity, which was necessary for installing a conventional bioelectrode on the skin surface.

[0074] According to the bioelectrode of III-(5), it is easier to insert the bioelectrode between the hair and bring the electrode surface into contact with the scalp. According to the bioelectrode of III-(6), the contact can be brought into stable contact with a wide area of ​​the skin surface or the surface of the living body. According to the bioelectrode of III-(7), the elasticity of the holder allows the contactor arranged on the surface of the sheet-like substrate to be pressed against the skin surface, so that the contactor can be stably installed. Furthermore, while maintaining this stable state, the holder slides on the back surface of the sheet-like substrate, and moves independently of the sheet-like substrate, so that the relative position between the holder and the skin surface can be changed. Therefore, even if the wearer of the bioelectrode moves his / her body, the biosignal can be stably transmitted. In III-(8), the biosignal measuring device may be, for example, an electrocardiogram measuring device, a heart rate monitor, an electroencephalogram measuring device, etc. In the biosignal measuring device, the bioelectrode may have not only a function of receiving a signal from the surface of the living body, but also a function of transmitting an electric signal (electrical stimulation).

[0075] The sheet-like substrates described in III-(6) and (7) and the sheet-like substrates that may be included in the bioelectrodes constituting the biosignal measuring device described in III-(8) are preferably provided with openings, which improves the air permeability between the sheet-like substrate and the skin, thereby reducing the sweating of the skin.

[0076] (Effects of the fourth aspect) According to the fourth aspect of the present invention, the following effects can be obtained. According to the fourth aspect of the implantable electrode of the present invention, the conductive composite fiber constituting the electrode is flexible and has excellent biocompatibility, so that the invasiveness to the biological tissue in which it is implanted can be reduced. In addition, the conductive polymer contained in the conductive composite fiber can detect weak electrical signals in the living body, so that highly accurate signal transmission and reception can be performed between an external device and the electrode implantation part. Furthermore, since the mechanical strength is increased by the fiber material that constitutes the conductive composite fiber, the electrode will not be damaged by external forces when implanted in the body, and it has excellent durability after implantation in the body. Since the implantable electrode provided in the biosignal measuring device of the present invention is less invasive to biological tissue, highly accurate signal transmission and reception can be performed between an external measuring device and the electrode implantation portion without impairing the inherent functions of the biological tissue. More specifically, the following effects can be obtained.

[0077] According to the electrode and device of the fourth aspect, the following effects can be obtained. According to the implantable electrode of IV-(1), the conductive composite fiber constituting the electrode has excellent flexibility and biocompatibility, so that the invasiveness to the biological tissue in which it is implanted can be reduced. In addition, the conductive polymer contained in the conductive composite fiber can transmit weak electrical signals in the living body, so that highly accurate signal transmission and reception can be performed between an external device and the electrode implantation part. Furthermore, since the mechanical strength is increased by the fiber material that constitutes the conductive composite fiber, the electrode will not be damaged by external forces when implanted in the body, and it has excellent durability after implantation in the body. According to IV-(2), the conductive composite fiber formed into a rod shape can be inserted like a needle when embedded in the body, thereby further reducing invasiveness to biological tissue. The conductive composite fiber formed into a coil shape (spiral shape) hardly shifts position within the embedded biological tissue, and can transmit and receive signals with high accuracy between an external device and the electrode embedded part. According to IV-(3), the conductive composite fiber can be installed in biological tissue easily and with low invasiveness by the simple operation of inserting (sticking) the needle having the conductive composite fiber attached to the tip end into biological tissue. Also, by the simple operation of removing the needle while leaving the installed conductive composite fiber inside the living body, the needle, which is not necessary for transmitting and receiving signals, can be removed outside the living body easily and with low invasiveness, completing the installation of the conductive composite fiber. According to IV-(4), when the needle is inserted into the biological tissue, the adhesive material dissolves by absorbing moisture such as body fluids, and the adhesion between the needle and the conductive composite fiber can be easily released. After the adhesion is released, the needle can be removed from the living body, and the conductive composite fiber can be left in the biological tissue.

[0078] According to IV-(5), the conductive fiber in the dry and shrunk state is flexible but has a relatively high mechanical strength, so that the conductive fiber can be prevented from being damaged (broken) when it is placed (inserted) into biological tissue. In addition, the physical volume of the conductive fiber in the dry and shrunk state is relatively small, so that the invasiveness when it is placed (inserted) into biological tissue can be reduced. According to IV-(6), the conductive composite fiber embedded in biological tissue can be electrically connected to an external device by the electric wire. According to IV-(7), by connecting one end of the core to a reservoir in which a drug solution or the like is stored or to a chamber equipped with a tube connector, and placing the other end of the core at a predetermined position within biological tissue, the drug solution or the like can permeate (penetrate) the conductive composite fiber that constitutes the core, and the drug solution can be transported into the biological tissue where the other end is located. According to IV-(8), the drug (medicine) can be administered to the site where the electrode is embedded through the flow path. The drug is preferably a drug having a pharmacological action of suppressing or promoting a biological reaction. Examples of the drug include a drug that reduces damage to biological tissue, a drug that promotes the repair of biological tissue, a drug that promotes the growth of biological tissue, and the like. According to IV-(9), since the implantable electrode is less invasive to biological tissue, signals can be transmitted and received with high accuracy between an externally installed measuring device and the implanted electrode portion without impairing the inherent functions of the biological tissue.

[0079] The base material fiber described in the present invention includes not only single-filament fibers, but also, for example, fibers obtained by twisting a plurality of base material fibers together to form a twisted thread of a desired thickness. The shape of the base material fiber in the present invention is not limited to the above-mentioned thread-like shape, but also includes, for example, a string-like shape, a cloth-like shape, a ribbon-like shape, and the like. [Brief description of the drawings]

[0080] [Figure 1] FIG. 2 is a schematic diagram showing a longitudinal cross section of an example of a conductive polymer fiber of the present invention. [Diagram 2] FIG. 2 is a schematic diagram showing a longitudinal cross section of an example of a conductive polymer fiber of the present invention. [Diagram 3] FIG. 2 is a schematic diagram showing a cross section in a direction perpendicular to the longitudinal direction in an example of a conductive polymer fiber of the present invention. [Figure 4] FIG. 2 is a schematic diagram showing a cross section in a direction perpendicular to the longitudinal direction in an example of a conductive polymer fiber of the present invention. [Diagram 5]FIG. 2 is a schematic diagram showing a cross section in a direction perpendicular to the longitudinal direction in an example of a conductive polymer fiber of the present invention. [Figure 6] FIG. 2 is a schematic diagram showing a cross section in a direction perpendicular to the longitudinal direction in an example of a conductive polymer fiber of the present invention. [Figure 7] FIG. 2 is a schematic diagram showing a cross section in a direction perpendicular to the longitudinal direction in an example of a conductive polymer fiber of the present invention. [Figure 8] FIG. 2 is a schematic diagram showing a cross section in a direction perpendicular to the longitudinal direction in an example of a conductive polymer fiber of the present invention. [Figure 9A] 1 is a graph showing the results of Comparative Example 1-1. [Figure 9B] 1 is a graph showing the results of Example 1-1. [Figure 10] 1 is a graph showing the results of Example 1-2. [Figure 11A] FIG. 1 is a schematic diagram showing an example of the present invention, illustrating a state in which a bioelectrode is placed on a body surface. [Figure 11B] This is a measured human electrocardiogram (leads I, II, and III from top to bottom). [Figure 12A] 13 is a photograph of the fabricated thread-shaped implantable bioelectrode observed with a stereomicroscope. [Figure 12B] These are the measured action potentials of a rat's sciatic nerve (from top to bottom: at rest, during muscle contraction, and during muscle relaxation).

[0081] [Figure 13] FIG. 1 is a schematic diagram illustrating an embodiment of a method and apparatus for producing conductive polymer fibers according to the present invention, and is a schematic diagram illustrating an example of the configuration of an apparatus for producing conductive polymer fibers. [Figure 14] FIG. 2 is an enlarged view of a main part, showing an example of an embodiment of the conductive polymer fiber manufacturing method and manufacturing apparatus of the present invention, in which comb-shaped electrodes are arranged on both sides of the base material fiber. [Figure 15A] FIG. 1 is an enlarged view of a main part that shows an example of an embodiment of a method and apparatus for producing conductive polymer fibers of the present invention, and shows an example of a roller-shaped rotor electrode. [Figure 15B] FIG. 13 is a diagram showing an example of a pulley-shaped rotor electrode. [Figure 15C] FIG. 13 is a diagram showing an example of a state in which roller-shaped rotor electrodes and pulley-shaped rotor electrodes are alternately combined in the longitudinal direction of the base fiber. [Figure 16] FIG. 2 is a schematic diagram showing another embodiment of the method and apparatus for producing conductive polymer fibers of the present invention, illustrating the case where a single electrode is used. [Figure 17A] FIG. 2 is a diagram for explaining an embodiment of a method and apparatus for producing conductive polymer fibers according to the present invention, and is a photograph showing conductive polymer fibers obtained by passing current through a rotor electrode according to the present invention in Example 2-1. [Figure 17B] FIG. 2 is a photograph showing conductive polymer fibers obtained by passing a current through a comb-shaped electrode in Example 2-2. [Figure 18A] FIG. 2 is a diagram for explaining an embodiment of a method and apparatus for producing a conductive polymer fiber according to the present invention, and is a photograph showing the state of a conductive polymer fiber obtained in Example 2-1 by electrochemical polymerization and fixing using the method and apparatus according to the present invention after a water resistance test. [Figure 18B] FIG. 1 is a photograph showing the state of the conductive polymer fiber obtained by conventional chemical fixation in Comparative Example 2-2 after a water resistance test.

[0082] [Figure 19A] This is a schematic diagram of a comb-shaped electrode according to a first embodiment of the third aspect of the present invention, and is a side view of a comb-shaped bioelectrode 310 showing how a contact 311 is fixed like a string to a bow-shaped first frame 312. [Figure 19B] FIG. 2 is a perspective view of the bioelectrode 310. [Figure 19C] 1 is a perspective view showing a state in which the bioelectrode 310 is inserted into a gap in hair H and fixed at a position where it contacts the scalp S. FIG. [Figure 19D] 1 is a photograph showing the area where the bioelectrode 310 contacts the skin. [Figure 20A]FIG. 13 is a schematic diagram of a hairpin-shaped electrode according to a first embodiment of the third aspect of the present invention, and is a side view of a bioelectrode 320 showing how two contacts 321 are fixed in parallel to a hairpin-shaped hair clip like strings via fixing posts 322. [Figure 20B] FIG. 2 is a perspective view of the bioelectrode 320. [Figure 20C] 13 is a perspective view showing a state in which the hairpin-shaped bioelectrode 320 is clamping the hair H near the scalp S and the contactor 21 is in contact with the scalp S. FIG. [Figure 20D] The top photo shows the top surface of the hairpin-shaped bioelectrode 320, and the bottom photo shows the bottom surface that comes into contact with the skin.

[0083] [Figure 21A] FIG. 2 is a schematic diagram showing an example of a cross-sectional structure of a contact, in which the left part shows a cross section taken along line XX of the right part. [Figure 21B] FIG. 11 is a schematic diagram showing a cross-sectional structure of a contact 321 according to a second embodiment of the third aspect, in which the left part shows a cross section taken along line XX of the right part. [Figure 21C] 19C is a photograph of a string-like contact 321 having the structure of FIG. 19B. [Figure 22A] 1 is a schematic diagram showing a comb-shaped bioelectrode 310 inserted between the hair and covered with a stretchable net-like holder N to hold the bioelectrode 310. FIG. [Figure 22B] FIG. 1 is a top view showing an example in which a bioelectrode 310 is installed on an elastic net N in accordance with the international 10-20 system. [Figure 23A] 1 is a graph showing human electroencephalograms (C3 C4 50 μV 400 msec / div) measured using hairpin-shaped EEG electrodes. [Figure 23B] 1 is a graph showing a human auditory brainstem response measured with a hairpin-shaped EEG electrode (0.2 μV, 1 ms / div 90 db, click sound, 1000 times averaging, showing the peak of IV-VII evoked potential).

[0084] [Figure 24A] 13 is a plan view showing the contact surface of an electrocardiogram electrode 330 with the skin. FIG. [Figure 24B] 13 is a side view showing the electrocardiogram electrode 330 placed on the skin surface S. FIG. [Figure 25A] 1 shows an example of a side view and a front view (skin contact surface, i.e., surface) of an electrode pad. [Figure 25B] 4 shows another example of a side view and a front view (skin contact surface, i.e., surface) of the electrode pad. [Figure 26A] 13 is a cross-sectional view showing the arrangement of the electrode pads 338 and the holder 335 relative to the trunk B. FIG. [Figure 26B] FIG. [Figure 27] 3 is a graph showing the recorded waveforms of electrocardiograms measured using the bioelectrode 1 of the present invention and conventional bioelectrodes 2 and 3 in Example 3-3, where the scale bar represents one second and 50 mV. [Figure 28] 1 is a graph showing the results of measuring the moisture content of the skin where bioelectrodes B and C according to the present invention or conventional bioelectrodes D and E are placed, and comparing the degree of sweating of the skin with each electrode.

[0085] [Figure 29A] FIG. 1 is a side view of a first embodiment of an implantable electrode according to the fourth aspect of the present invention. [Figure 29B] 29B is a side view showing how the conductive composite fiber bundle shown in FIG. 29A swells due to water absorption. FIG. [Figure 29C] FIG. 4 is a side view showing a state in which the needle is being removed from the swollen conductive composite fiber bundle. [Figure 30A] 11A and 11B are a side view and a bottom view showing an example of an electrode according to a fourth embodiment of the present invention, in which a guide needle is provided with one conductive composite fiber bundle. [Figure 30B] 11A and 11B are a side view and a bottom view showing an example of an electrode according to a fourth embodiment of the present invention, in which a guide needle is provided with two conductive composite fiber bundles. [Figure 30C] 11A and 11B are a side view and a bottom view showing an example of an electrode according to a fourth embodiment of the present invention, in which a guide needle is provided with four conductive composite fiber bundles. [Figure 31A] FIG. 11 is a side view of a second embodiment of the implantable electrode according to the fourth aspect of the present invention. [Figure 31B] 4 is a side view showing the conductive composite fiber bundle of the electrode swelling due to water absorption. FIG. [Figure 31C] 13 is a side view showing the state in which the needle is removed from the swollen conductive composite fiber bundle of the electrode. FIG.

[0086] [Figure 32A] FIG. 13 is a schematic diagram showing the introduction of a third embodiment of the implantable electrode according to the fourth aspect of the present invention into a nerve cord N'. [Figure 32B] FIG. 13 is a schematic diagram showing the introduction of a third embodiment of an implantable electrode into a nerve cord N'. [Figure 32C] FIG. 13 is a schematic diagram showing the introduction of a third embodiment of an implantable electrode into a nerve cord N'. [Fig. 32D] FIG. 13 is a schematic diagram showing the introduction of a third embodiment of an implantable electrode into a nerve cord N'. [Figure 33A] 13A and 13B are side and bottom views of a fourth embodiment of an implantable electrode with a flow path and a reservoir. [Figure 33B] 4 is a fourth embodiment with a chamber and a tube connector.

[0087] [Figure 34A] (A) Waveforms (Scale bar 250 ms 40 mV) of collective action potentials in the rat brain measured using the fourth embodiment of the implantable electrode (PEDOT-PSS) and a conventional metal electrode, respectively. [Figure 34B] This shows an immunostained image of glial cells (astrocytes) in the rat cerebral cortex using an anti-GFAP antibody, without any electrodes being implanted. [Figure 34C] 1 shows an immunostained image of glial cells (astrocytes) in the rat cerebral cortex using an anti-GFAP antibody, and also shows the position (dotted line) where an electrode according to a fourth embodiment of the present invention is embedded and the surrounding condition. [Fig. 34D]The figure shows an immunostained image of glial cells (astrocytes) in the rat cerebral cortex using an anti-GFAP antibody, and also shows the location where a conventional metal electrode was embedded (dotted line) and the condition of its surroundings. [Figure 35A] 2 is a waveform showing an action potential of the cerebral cortex (barrel cortex) of a rat recorded by the implantable electrode of the first embodiment of the present invention. [Figure 35B] 13 is a waveform showing a collective action potential of the sciatic nerve of a rat recorded by the implantable electrode of the third embodiment of the present invention (Scale bar 1 second 50 μV). [Figure 35C] 13 is a waveform showing a rat's electrocardiogram recorded by the implantable electrode of the third embodiment of the present invention (Scale bar 1 second 50 mV). [Diagram 36] 1 is a graph showing the kinetics of drug transport in a conductive composite fiber bundle coated with PDMS. DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

[0088] Hereinafter, preferred examples and embodiments of the present invention will be described with reference to the drawings for the first to fourth aspects of the present invention. However, the present invention is not limited to the following examples and embodiments. For example, the components and conditions of these preferred examples and embodiments may be appropriately combined. In addition, preferred examples may be used with each other between the various aspects. Furthermore, they may be combined with other components as long as there are no problems. Various changes can be made to the position, number, size, amount, and the like, without departing from the spirit of the present invention. For example, the examples that are preferred in the description of FIG. 1 can also be preferably used in other examples of this embodiment unless otherwise specified. Regarding the first aspect Hereinafter, an embodiment of the first aspect of the present invention will be described with reference to the drawings, but the present invention is not limited to such an embodiment. The first aspect of the present invention relates to a conductive polymer fiber and a bioelectrode. More specifically, the present invention relates to a conductive polymer fiber in which a conductive polymer is impregnated or attached to a base fiber, and a bioelectrode including the conductive polymer fiber. <First embodiment of the first aspect> The conductive polymer fiber 10 (first embodiment) of the present invention shown in FIG. 1 is a fiber in which a base fiber 11 is coated with a conductor 12 containing PEDOT-PSS {poly(3,4-ethylenedioxythiophene)-poly(styrenesulfonic acid)} as a conductive polymer. FIG. 1 is a longitudinal cross-sectional view of the conductive polymer fiber 10, and FIG. 3 is a cross-sectional view in a direction perpendicular to the longitudinal direction. The conductive polymer fiber 10 has the base fiber 11 as a core and is coated with the conductor 12 around it, so that the contact area between the two is large and the fibers are sufficiently bonded to each other to form a composite fiber. With this configuration, the conductor 12 is reinforced by the base fiber 11, so that the strength can be increased compared to a fiber consisting of only the conductor 12. In particular, the strength is excellent in a dry state and a wet state. In addition, the flexibility of the base fiber 11, which is the core, is imparted to the conductive polymer fiber 10.

[0089] The type of base fiber 11 is not particularly limited as long as it is made of a polymer. For example, synthetic fibers, vegetable fibers, animal fibers, etc. may be used. The material may be a single material or a mixture. Examples of the synthetic fibers include nylon, polyester, acrylic, aramid, polyurethane, carbon fiber, etc. Examples of the vegetable fibers include cotton, hemp, jute, etc. Examples of the animal fibers include silk, wool, collagen, and elastic fibers that constitute animal tissues.

[0090] Among the exemplified base fiber materials, animal fibers (protein-containing fibers) are preferred, which have excellent adhesion to the conductor 12, high strength in dry and wet states, and flexibility suitable for use in clothing, etc. Furthermore, silk fibers are more preferred, which are particularly excellent in adhesion and hydrophilicity to PEDOT-PSS, which will be described later. The base fiber is preferably silk alone. It is also preferably a mixture as necessary. In the case of a silk mixture, the silk content may be 0.1% or more and less than 100%, 1% or more and less than 95%, 3% or more and less than 90%, 10% or more and less than 80%, 30% or more and less than 70%, or 40% or more and less than 60%. It is also preferable to mix it with another material as appropriate depending on the purpose.

[0091] Examples of silk fibers that can be used as the base fiber 11 include natural silk fibers from silkworm moths, spiders, and bees, and artificial silk fibers made using genetic engineering technology. Silk contains a protein called fibroin, and is a fiber that is hydrophilic, biocompatible, and dyeable, and is used for clothing and surgical threads. It is one of the fibers that has been used by humans since ancient times. For this reason, it is suitable for use as the base fiber 11.

[0092] The silk fiber used for the base fiber 11 may be either unprocessed raw silk from which the gelatin component sericin has not been removed, or degummed yarn from which some or all of the sericin has been removed. Degummed yarn is more preferable from the viewpoint of increasing the adhesion to the conductor 12 and the fiber strength.

[0093] The diameter (thickness) of the base fiber 11 is not particularly limited and can be appropriately selected depending on the application. For example, the diameter ranges include 0.1 μm to 1 mm, 1 μm to 1 mm, and 1 μm to 0.5 mm. When used in clothing, bioelectrodes, biointerfaces, and the like, the diameter is preferably, for example, 1 μm to 100 μm. The length of the base fiber 11 is not particularly limited and can be appropriately selected depending on the application. For example, it can be 10 μm to 10 cm for an electrode to be embedded in a biological tissue, 1 mm to 50 cm for use in a biointerface on the body surface, and 1 cm to 100 m as a fiber material for weaving or knitting into clothing. It is not limited to these lengths and can be selected as needed.

[0094] The base fiber 11 (41) is not particularly limited and can be selected as necessary. For example, a plurality of base fibers may be twisted together to form a twisted yarn of a desired thickness (see the example shown in FIG. 6), or a blended yarn in which different types of base fibers are blended may be used. The shape of the base fiber is not limited to the thread-like shape described above, and for example, string-like, cloth-like, ribbon-like, and other base fiber shapes may be used. In addition, in order to improve the hydrophilicity of the base fiber 11, a base fiber that has been subjected to plasma treatment, pore treatment, or chemical coating may be used.

[0095] The conductor 12 contains a conductive polymer, and may consist of only a conductive polymer, or may contain other additives.

[0096] The conductive polymer used in the present invention is PEDOT-PSS {poly(3,4-ethylenedioxythiophene)-poly(styrenesulfonic acid)}, which has excellent conductivity and hydrophilicity. PEDOT-PSS is a conductive polymer that can be obtained by polymerizing the monomer 3,4-ethylenedioxythiophene in the presence of poly(4-styrenesulfonic acid). PSS functions as a dopant that imparts a negative charge to PEDOT. In the present invention, from the viewpoint of increasing the conductivity of the conductive polymer fiber, it is preferable that the conductive polymer contains a dopant.

[0097] The present inventors have found that the adhesion between PEDOT-PSS and protein-containing fibers such as silk is particularly excellent, and that the adhesive surfaces of the two do not easily peel off. Based on this finding, in the present invention, it is more preferable to use silk fibers as the base fiber 11 and to use PEDOT-PSS as the conductive polymer contained in the conductor 12.

[0098] Other examples of conductive polymers that can be used include polyaniline sulfonic acid and polypyrrole. The conductive polymer contained in the conductor 12 may be one type, or two or more types may be used in combination. The molecular weight of the conductive polymer used in the present invention is not particularly limited. For example, a molecular weight in the range of several thousand to several hundred thousand can be used. It may be selected as necessary. To give a specific example, the weight average molecular weight (Mw) in terms of polystyrene may be in the range of 1000 to 900000, 3000 to 450000, or 5000 to 50000. However, it is not limited to these ranges.

[0099] An example of a method for forming the conductor 12 is to coat a solvent containing a conductive polymer such as PEDOT-PSS and a dilution solvent on the base fiber 11 and then dry the solvent to form the conductor 12 made only of the conductive polymer. However, the conductor 12 may contain additives other than the conductive polymer.

[0100] Examples of the additives include glycerol, sorbitol, polyethylene glycol-polypropylene glycol copolymer, ethylene glycol, sphingosine, phosphatidylcholine, etc. The additives contained in the conductor 12 may be one type, or two or more types may be used in combination.

[0101] The additives in the above examples can be used to adjust the wettability of conductive polymer fibers such as PEDOT-PSS, or to impart flexibility to improve affinity with biological tissues (skin and tissues) when used as a biological electrode. Specific examples of the adjustment of the wettability include adjustment of water absorbency and prevention of excessive expansion and contraction during wetting and drying.

[0102] The use of a combination of PEDOT-PSS and the additive is preferable because it makes it easier to adjust the wettability of the conductor 12, and in particular makes it easier to prevent excessive expansion and contraction. One of the reasons for this is thought to be that by incorporating PEDOT-PSS, which has high water absorption properties, together with the additive in advance, there is less room for moisture to penetrate later. As additives used for adjusting the wettability of PEDOT-PSS and imparting flexibility, among the above examples, glycerol, sorbitol, polyethylene glycol, and polyethylene glycol-polypropylene glycol copolymer are particularly preferred.

[0103] The conductive polymer fiber 10 having the conductor 12 containing the additive and PEDOT-PSS does not absorb excessive water even when used in a high humidity environment, has high fiber strength, and is excellent in conductivity. In addition, since it also has excellent flexibility, the stiff feel (rigidity) of PEDOT-PSS is mitigated, and it has excellent contact and affinity with biological tissue, so that it is possible to configure a biological electrode that can measure biological signals with little noise.

[0104] The additives contained in the conductor 12 are not limited to the above examples, and may be, for example, surfactants, alcohols, natural polysaccharides, sugar alcohols, acrylic resins, known organic solvents such as dimethyl sulfoxide, and the like.

[0105] The surfactant may be a cationic surfactant, an anionic surfactant, or a nonionic surfactant, which may be used alone or in combination of two or more.

[0106] Examples of the cationic surfactant include quaternary alkyl ammonium salts and alkylpyridinium halides. Examples of the anionic surfactant include alkyl sulfates, alkyl benzene sulfonates, alkyl sulfosuccinates, and fatty acid salts. Examples of the nonionic surfactant include polyoxyethylene and polyoxyethylene alkyl ether.

[0107] As the alcohol, a wide variety of known monohydric and polyhydric alcohols can be used. These alcohols may be used alone or in combination of two or more.

[0108] Examples of monohydric alcohols include methanol, ethanol, propyl alcohol, isopropyl alcohol, butanol, etc. The carbon skeleton constituting these alcohols may be linear, branched, or cyclic. Examples of polyhydric alcohols include glycols such as ethylene glycol, linear polyhydric alcohols such as glycerin, cyclic polyhydric alcohols such as glucose and sucrose, and polymeric polyhydric alcohols such as polyethylene glycol, polyvinyl alcohol, polyethylene glycol, and polypropylene glycol copolymers.

[0109] Examples of natural polysaccharides include chitosan, chitin, glucose, and aminoglycan. Examples of sugar alcohols include sorbitol, xylitol, and erythritol. Examples of the acrylic resin include polyacrylic acid, polymethyl methacrylate, and polymethyl methacrylate resin.

[0110] The thickness h of the conductor 12 coated around the base fiber 11 is not particularly limited and can be selected arbitrarily. Any thickness is acceptable as long as the effect of the present application can be obtained. For example, the thickness may be 0.001 to 2 times the diameter L of the base fiber 11. The thickness h can be selected as necessary, and may be, for example, 0.01 to 1 times the diameter L of the base fiber 11, 0.001 to 0.1 times the thickness, 1 to 2 times the thickness, or 0.1 to 1 times the thickness. More specifically, when a silk fiber of 2 to 3 denier (D) silkworm is used as the core, that is, a silk fiber with a fiber diameter of about 10 to 15 microns, the thickness may be 0.01 microns to 10 microns. The diameter of the fiber can be measured by any method, but can also be confirmed, for example, by an electron microscope photograph. If necessary, the base fiber may not be completely covered by the conductor. In the present invention, the thickness may be the length of a line from the center of the base fiber to the surface that is covered by the conductor.

[0111] By covering the periphery of the base fiber 11 with the conductor 12, the conductivity of the conductive polymer fiber 10 is increased and it becomes easier to bring multiple conductive polymer fibers 10 into contact with each other to conduct electricity. Furthermore, when the thickness is within the above range, the conductive polymer fiber 10 can have a fiber with better conductivity without impairing its flexibility. Within the above range, the thicker the fiber, the higher the conductivity of the fiber. In other words, the conductivity or electrical resistance of the conductive polymer fiber 1 can be adjusted by adjusting the thickness of the conductor 12.

[0112] <Preparation of conductive polymer fibers (1a)> As in the conductive polymer fiber 10 shown in FIG. 1, examples of a method for attaching or covering the surface of a base fiber 11 with a conductor 12 include the following methods. First, an aqueous solution containing a conductive polymer (for example, a commercially available PEDOT-PSS solution (Heraeus: CLEVIOS P)) is applied to the surface of the base fiber 11 in a solution bath. After this, or after the solution is uniformly applied to the surface of the base fiber 11 using a roller or a brush, a part of the water contained in the solution is dried and removed. Next, an organic solvent such as acetone, methanol, ethanol, or a fixing solution such as a magnesium chloride solution is applied to gel the conductive polymer such as PEDOT-PSS. This can be exemplified as a method (hereinafter sometimes referred to as preparation method 1a) in which a conductor 12 containing a conductive polymer such as PEDOT-PSS is fixed to the surface of the base fiber 11. An example of the aqueous solution is an aqueous solution containing a conductive polymer such as PEDOT-PSS at a concentration of 0.1 to 50 (v / v)%. This concentration can be selected as necessary. For example, the concentration may be 1 to 30%, 30 to 50%, or 0.5 to 15%. The aqueous solution can contain the additive as necessary. Furthermore, in the present invention, in addition to the example of CLEVIOS P, any solution containing PEDOT-PSS can be used as the aqueous solution containing a conductive polymer.

[0113] <Method of adding additives> Examples of methods for incorporating an additive into the conductor 12 include a method in which the conductor 12 applied to the base fiber 11 by the preparation method 1a is dried, and then the resulting conductive polymer fiber 10 is used to apply the additive to its surface, or a method in which the conductive polymer fiber 10 is immersed in a solution containing the additive for a predetermined time, and then the excess additive solution remaining on the surface is removed. As another method, a method in which a mixed solution in which an additive is mixed in a solution containing a conductive polymer to be applied to the surface of the base fiber 11 is used to apply or immerse the conductive polymer and the additive together.

[0114] An example of the mixed liquid is an aqueous solution containing a conductive polymer such as PEDOT-PSS at a concentration of 0.1 to 50 (v / v) % and an additive such as glycerol at a concentration of 0.1 to 50 (v / v) %. The concentration of the additive in the conductor 12 in the present invention is not particularly limited and may be, for example, 0.1 to 50 wt %. This concentration can be selected as necessary and may be, for example, 0.1 to 20 wt %, 20 to 50 wt %, or 0.1 to 5 wt %.

[0115] Second embodiment of the first aspect The conductive polymer fiber 20 (second embodiment) of the present invention shown in Fig. 2 is a fiber in which a base fiber 21 is impregnated with a conductor 22 containing a conductive polymer. Fig. 2 is a longitudinal cross-sectional view of the conductive polymer fiber 20, and Fig. 4 is a cross-sectional view in a direction perpendicular to the longitudinal direction. The conductor 22 permeates the interior of the base fiber 21, forming a composite fiber in which the two are integrated. With this configuration, there is no risk of the conductor 22 falling off the base fiber 21. In addition, the conductor 22 is reinforced by the base fiber 21, making it possible to increase the strength compared to a fiber consisting only of the conductor 22. In addition, it also has the flexibility of the base fiber 21. In the present invention, the entire internal space of the base fiber may be filled with the conductor, but some spaces may be unfilled. It is preferable that the conductor reaches the center of the base fiber, but if necessary, some parts may not reach the center.

[0116] The materials constituting the base fiber 21 and the conductor 22 may be the same as those constituting the base fiber 11 and the conductor 12 described in the first embodiment. As in the first embodiment, it is preferable that the conductor 22 contains the additive.

[0117] <How to prepare conductive polymer fibers (1b)> An example of a method for immersing a conductor 22 inside a base fiber 21, as in the conductive polymer fiber 20 shown in Figure 2, is to immerse the base fiber 21 in an aqueous solution containing a conductive polymer (e.g., a commercially available PEDOT-PSS solution (Heraeus CLEVIOS P)) in a solution bath for a predetermined period of time, then dry and remove some of the water contained in the solution, and then apply an organic solvent such as acetone, methanol, ethanol, etc. or a fixing solution such as a magnesium chloride solution to gel the PEDOT-PSS, thereby fixing the conductor 22 containing PEDOT-PSS to the surface of the base fiber 21 (hereinafter sometimes referred to as preparation method 1b). The aqueous solution may contain the additives as required.

[0118] Examples of methods for promoting the penetration of the solution containing a conductive polymer into the base fiber 21 include a method of immersing the base fiber 21 after adjusting the pH of the solution, a method of applying mechanical operations such as tension or compression to the base fiber 21 during immersion, a method of heating the solution during immersion, a method of applying a treatment such as reduced pressure or pressure during immersion, etc. Specifically, when base fiber 21 such as silk is immersed in a PEDOT-PSS solution, it is preferable to adjust the pH of the solution to 1 to 6.

[0119] <Third embodiment of the first aspect> A conductive polymer fiber 30 (third embodiment) of the present invention shown in Fig. 5 is formed by impregnating a base fiber 31 with a conductor 32 containing a conductive polymer, coating the base fiber 31 with a metal 33, and further coating the coated metal or carbon 33 with a conductor 34. Fig. 5 is a cross-sectional view of the conductive polymer fiber 30 in a direction perpendicular to the longitudinal direction. Hereinafter, unless otherwise specified, "metal or carbon" will be referred to as metals.

[0120] The third embodiment has the advantages of both the first and second embodiments. In addition, the coated metals 33 themselves contribute to improving the conductivity of the conductive polymer fiber 30. Since the metals 33 are sandwiched between the conductors 32 and 34, the metals 33 are not exposed on the fiber surface. This prevents the metals 33 from corroding or deteriorating. Note that, if necessary, a portion of the metals 33 may be exposed on the fiber surface.

[0121] The materials constituting the base fiber 31 and the conductors 32, 34 can be the materials constituting the base fiber and the conductors described in the first and second embodiments. As in the first and second embodiments, it is preferable that the conductors 32, 34 contain the additive. The materials constituting the conductors 32 and 34 may be the same or different.

[0122] The type of metal 33 is not particularly limited, and examples include titanium, gold, silver, copper, carbon, etc. Among these metals, gold is preferable because it has excellent corrosion resistance, electrical conductivity, and ductility. The carbon is preferably one containing carbon atoms as a main raw material, and examples of the carbon include carbon materials containing carbon black, glassy carbon, graphene, carbon nanotubes, fullerene, etc. The carbon content in these carbon materials is preferably 80 to 100 mass%, more preferably 90 to 100 mass%, and further preferably 95 to 100 mass%. The metals 33 may be one type of metal used alone, or two or more types of metals used in combination.

[0123] The thickness of the metals 33 (metal layer or carbon layer) coated around the base fiber 31 is not particularly limited, and can be changed appropriately depending on the type of metal. For example, the range can be 0.1 nm to 1 mm. For example, when gold is used, the thickness can be 1 nm to 2 μm. The metal layer 33 can be formed by a known film formation method such as a sputtering method or an electroless plating method. Also, the carbon layer can be formed by a known film formation method such as carbon vapor deposition.

[0124] <Preparation of conductive polymer fibers (2a)> The conductive polymer fibers 30 can be produced, for example, by the following method. First, the conductive polymer fiber 20 obtained by the preparation method 1b is coated with a metal 33 by a known film forming method. The fiber obtained here is immersed in an aqueous solution containing a conductive polymer (for example, a commercially available PEDOT-PSS solution (Heraeus CLEVIOS P)), and the metal 33 is used as an electrode, and a direct current voltage of +0.5V to 20V is applied, thereby producing a conductive polymer fiber 30 in which a conductive polymer such as PEDOT-PSS is electrochemically fixed to the surface of the metal 33. Hereinafter, this method may be referred to as preparation method 2a. Here, a method of forming a metal layer around the conductive polymer fiber 20 has been exemplified, but a method of simply forming a metal layer on a base fiber and similarly electrically fixing the conductive polymer around the metal layer may also be adopted.

[0125] By adding ethylenedioxythiophene (EDOT) to the solution, conductive polymer fibers 30 with even better conductivity can be obtained. The amount of ethylenedioxythiophene can be selected arbitrarily. For example, a 0.1 w / v % solution of ethylenedioxythiophene (Heraeus CLEVIOS M V2) may be added to the solution.

[0126] <How to prepare conductive polymer fibers (2b)> A method of electrochemically fixing the conductive polymer without forming the metal 33 can also be exemplified. That is, the conductive polymer fiber 20 obtained by the preparation method 1b already has conductivity. By utilizing this conductivity, the conductive polymer fiber 20 obtained by the preparation method 1b is added to a solution containing a conductive polymer (for example, a commercially available PEDOT-PSS solution (Heraeus CLEVIOS P)), and a direct current voltage of +0.5V to 20V is applied in this solution, thereby producing a conductive polymer fiber in which a conductive polymer such as PEDOT-PSS is electrochemically fixed to the surface around the conductive polymer fiber 20. Hereinafter, this method may be referred to as the preparation method 2b. The additive may be contained in the aqueous solution as necessary.

[0127] Although the configuration of the conductive polymer fiber obtained by preparation method 2b is not illustrated, the conductive polymer fiber has a configuration in which the metal 33 is removed from the conductive polymer fiber 30 shown in FIG.

[0128] <Fourth embodiment> The conductive polymer fiber 40 (fourth embodiment) of the present invention shown in FIG. 6 is formed by disposing a conductor 42 containing a conductive polymer between a plurality of base fibers 41 in close contact with the base fiber 41. FIG. 6 is a cross-sectional view of the conductive polymer fiber 40 in a direction perpendicular to the longitudinal direction. The number of base fibers can be selected arbitrarily and is an integer of 2 or more. For example, it may be 2, 3, 4, 5, 6, 7, or 8. It may be a number in the range of 1 to 1000 or a number in the range of 1 to 30. The conductive polymer fiber 40 may be a higher-order structure such as a twisted string, woven fabric, or nonwoven fabric formed by twisting or knitting a plurality of base fibers 41. As shown in the example of FIG. 6, a conductor 42 containing PEDOT-PSS, which is a conductive polymer, is disposed between a plurality of base fibers 41 in close contact with the base fiber 41, and the conductive polymer fiber 40 can be configured as a higher-order structure such as a twisted string, woven fabric, or nonwoven fabric by twisting or knitting a plurality of base fibers 41. The conductor 42 serves to bond the multiple base fiber 41 together, thereby increasing the strength of the higher-order structure. Furthermore, a relatively large amount of the conductor 42 can be disposed between the multiple base fiber 41, resulting in a conductive polymer fiber with superior conductivity. Note that the amount compared here is the amount of conductor disposed on the surface of a single base fiber.

[0129] The method for producing the conductive polymer fibers 40 is not particularly limited, and for example, a method in which the higher-order structure is immersed in a solution containing a conductive polymer and then dried can be mentioned. The fiber spacing between the multiple base fibers 41 can be selected arbitrarily. For example, the fiber spacing between the base fibers 41 may be about 0.01 to 3 times the diameter of the base fiber. For example, when base fibers 41 with a diameter of 10 μm to 15 μm are used, the fiber spacing can be 0.01 μm to 50 μm. With a fiber spacing in this range, the conductors 42 can be sufficiently distributed between the fibers.

[0130] The materials constituting the base fiber 41 and the conductor 42 may be the same as those constituting the base fiber and the conductor described in the first embodiment. As in the first embodiment, it is preferable that the conductor 42 contains the additive.

[0131] Fifth embodiment of the first aspect A conductive polymer fiber 50 (fifth embodiment) of the present invention shown in Fig. 7 is formed by arranging a conductor 54 containing a conductive polymer between a plurality of base fibers 51 impregnated with a conductor 52 containing a conductive polymer in close contact with the base fiber 51. Fig. 7 is a cross-sectional view of the conductive polymer fiber 50 in a direction perpendicular to the longitudinal direction.

[0132] The configuration of the fifth embodiment is similar to that of the fourth embodiment, except that the inside of the base fiber 51 is impregnated with the conductor 52. In this embodiment, the conductor 52 further improves the electrical conductivity.

[0133] The material constituting the conductor 52 may be the same as or different from the material constituting the conductor 54. The manufacturing method of the fifth embodiment can be the same as or different from the manufacturing methods of the first to fourth embodiments.

[0134] Sixth embodiment of the first aspect A conductive polymer fiber 60 (sixth embodiment) of the present invention shown in Fig. 8 is configured such that a base fiber 61 is coated with a conductor 62 containing a conductive polymer, and the periphery of the conductor 62 is coated with an insulating layer 63. Fig. 8 is a cross-sectional view of the conductive polymer fiber 60 in a direction perpendicular to the longitudinal direction. The fiber has excellent durability because the base fiber 61 and the conductor 62 are protected by the insulating layer 63. If necessary, a part of the insulating layer 63 can be removed to expose a part of the conductor 62 on the fiber surface.

[0135] Known insulating materials can be used as the material for the insulating layer 63. From the viewpoint of biocompatibility and flexibility, polytetrafluoroethylene (PTFE) and silicone resin (silicone rubber) are preferable. The thickness of the insulating layer 63 is not particularly limited. It can be selected arbitrarily and can be in the range of, for example, 0.1 μm to 3 mm, 0.1 μm to 2 mm, 1 μm to 2000 μm, or 10 μm to 500 μm. In addition, the base fiber 61 and the conductor 62 can be covered with the insulating layer 63 by a known resin coating method.

[0136] <Bioelectrode> The conductive polymer fiber according to the present invention has sufficient strength, conductivity and flexibility even under high humidity conditions, and is therefore suitable for use in not only bioelectrodes and biointerfaces but also clothing. By bundling a plurality of conductive polymer fibers according to the present invention into a thread or string, sufficient conductivity for measuring biosignals can be obtained. Since the conductive polymer PEDOT-PSS is arranged in the fiber, electrical continuity can be obtained immediately by contacting the fiber with a measurement object. Therefore, by contacting the fiber (thread) with a measurement object, or ligating, wrapping, sewing, folding, etc., it is possible to stably record biosignals for a long period of time. When a bioelectrode is made by using the conductive polymer fiber according to the present invention as an electrode, the fibers can be bound together into a thread and then knotted, knitted, sewn, or bound together to provide a variety of bioelectrodes in various shapes, such as cloth, belt, strap, etc. Furthermore, by combining the conductive polymer fibers and forming them into a nonwoven fabric, etc., a patch-shaped (cloth-shaped) bioelectrode can also be made.

[0137] Regarding the second aspect A second aspect of the present invention relates to a method and apparatus for producing conductive polymer fibers, and in particular to a method and apparatus for producing conductive polymer fibers, in which a conductor containing a conductive polymer is impregnated into or adhered to insulating fibers (fiber bundles). Hereinafter, the conductive polymer fiber manufacturing method and manufacturing apparatus according to the second aspect of the present invention will be described with reference mainly to Figs. 13 to 18 as appropriate, but the second aspect of the present invention is not limited to the following embodiments. Here, Figs. 13 to 16 are schematic diagrams showing the conductive polymer fiber manufacturing apparatus described in this embodiment. Figs. 1, 3, 6, etc. are schematic diagrams showing examples of conductive polymer fibers obtained by the manufacturing method and manufacturing apparatus of this embodiment.

[0138] [Conductive polymer fiber] The manufacturing method and manufacturing apparatus according to the second aspect of the present invention can preferably form the conductive polymer fiber described in the first aspect. The preferred conditions described in the first aspect can also be used here. For example, the conductive polymer fiber shown in Figs. 1, 3 and 6 can be easily formed. In this aspect, it is more preferable to use PEDOT-PSS as the conductive polymer contained in the conductor 12, but the base fiber used in this embodiment is not limited to silk fiber, and other general fiber materials can be used without any restrictions. It is only necessary that PEDOT-PSS is contained as the conductive polymer.

[0139] In this embodiment, when producing the conductive polymer fiber 10 by the production method and production apparatus described below, first, the base fiber 11 is immersed in a solution of the conductor 12 to impregnate and / or adhere the conductor 12 to the base fiber 11. In this case, the solution of the conductor 12 contains a dilution solvent in addition to the conductive polymer PEDOT-PSS, and may further contain additives other than the conductive polymer as necessary.

[0140] [Method and apparatus for producing conductive polymer fibers] Hereinafter, one embodiment of a method and apparatus for producing conductive polymer fibers according to the present invention will be described in detail mainly with reference to FIGS.

[0141] "Manufacturing equipment" First, the manufacturing apparatus used in this embodiment will be described in detail. A manufacturing apparatus (hereinafter, sometimes abbreviated as manufacturing apparatus) 210 for conductive polymer fibers shown in Fig. 13 includes an immersion container 205. The immersion container 205 is an immersion container that contains a conductor solution 204 containing PEDOT-PSS {poly(3,4-ethylenedioxythiophene)-poly(styrenesulfonic acid)} as a conductive polymer, and is used for impregnating and / or attaching a conductor (see also conductor 212 shown in Figs. 17A, 17B, etc.) to the base fiber 211 by immersing an insulating base fiber 211 made of a thread-like, string-like, cloth-like, or ribbon-like fiber bundle in the conductor solution 204. The manufacturing apparatus 210 also includes a winding section 209 for vertically pulling up the base fiber 211 from the conductor solution 204 contained in an immersion container 205, and a plurality of electrodes 202, 203 for electrochemically polymerizing and fixing the conductor 212 impregnated in and / or attached to the base fiber 211 by passing an electric current through the base fiber 211 being pulled up vertically while the base fiber 211 is traveling. The manufacturing apparatus 210 described in this embodiment is generally configured to include a drying section 208 for drying the base fiber 211 to which the conductor 212 has been polymerized and fixed by blowing air thereto, and a chamber (including a humidity control section) 207 capable of adjusting the atmospheric humidity in the vicinity of the base fiber 211.

[0142] The immersion container 205 is a container in which the conductor solution 204 containing PEDOT-PSS as the conductive polymer as described above is accommodated, and a conventionally known container may be used. A spool 206 is accommodated in the immersion container 205 so as to be immersed in the conductor solution 204. An insulating base fiber 211 made of a thread-like, string-like, cloth-like, or ribbon-like fiber bundle wound around the spool 206 is immersed in the conductor solution 204, whereby the base fiber 211 is impregnated with and / or adhered to the conductor 212.

[0143] In this embodiment, the thread spool 206 may also be a conventionally known one. For example, a bobbin-like shape is used, which can be rotated by a motor or the like to wind the base fiber 211. The base fiber 211 before the treatment of impregnating and / or attaching the conductor 212 is wound around the bobbin 204, and as the treated base fiber 211, i.e., the conductive polymer fiber 201, is wound by a winding unit 209 described below, the base fiber 211 wound around the bobbin 204 is unwound. The shape of the bobbin 204 is not limited to the above-mentioned bobbin, and when a cloth-like or string-like base fiber is used, for example, a winding shaft suitable for the shape of the base fiber may be used.

[0144] The winding unit 209 pulls up the base fiber 211 vertically at a constant speed from the conductive solution 204 contained in the immersion container 205, and winds and rolls up the base fiber 211. As with the above-mentioned spool 204, a rotatable device shaped like a roll bobbin or the like can be used for this purpose as well. In addition, the winding unit 209 is configured to be able to adjust the amount of the conductive material 212 electrochemically polymerized and fixed to the base fiber 211 by adjusting the running speed of the base fiber 211, as will be described in detail later.

[0145] In this embodiment, the base fiber 211 is vertically pulled up from the conductive solution 204 by the winding section 209, so that the amount of conductive material 212 impregnated into and / or attached to the base fiber 211 is constant, and therefore electrochemical polymerization fixation by passing current is uniform.

[0146] 13, in the manufacturing apparatus 210, a plurality of electrodes 202, 203 are alternately provided, which pass electricity through the base fiber 211 pulled up vertically by the winding section 209 while it is traveling, thereby electrochemically polymerizing and fixing the conductor 212 impregnated in and / or attached to the base fiber 211. The plurality of electrodes 202, 203 are provided in the longitudinal direction of the base fiber 211, and are arranged on both sides of the base fiber 211 in the lateral diameter direction, so that the base fiber 211 is sandwiched between the plurality of electrodes 202, 203. A constant current or a constant voltage is applied to the plurality of electrodes 202, 203 from a DC stabilized power supply (not shown). As long as the effects of the present invention can be obtained, the electrodes 202, 203 do not have to be arranged vertically. It is preferable that the electrodes are not in contact with each other and are spaced apart, but the distance can be selected arbitrarily. The number of electrodes can also be selected arbitrarily, and there may be one or more combinations of anode and cathode. For example, the number of combinations may be in the range of 1 to 10, 2 to 8, or 3 to 5.

[0147] In the example shown in Fig. 13, the multiple electrodes 202 are anodes, and the multiple electrodes 203 are cathodes. As a result, in Fig. 13, the base material fiber 211 impregnated with and / or attached to the conductor 212 travels vertically and passes through each electrode in the order of "positive electrode (+)" - "negative electrode (-)" - "positive electrode (+)" - "negative electrode (-)". In this way, the positive and negative electrodes are applied alternately.

[0148] The multiple electrodes 202, 203 are made of, for example, a conductive metal material or a carbon material, and while in contact with the vertically running base fiber 211, conduct electricity in the longitudinal direction of the base fiber 211. By conducting electricity between the electrodes to the running base fiber 211 in this manner, the PEDOT-PSS contained in the conductor 212 impregnated and / or attached to the base fiber 211 is polymerized and electrochemically polymerized and fixed.

[0149] As the multiple electrodes 202, 203, known electrodes having various shapes that have been conventionally used in the field of electrodes, such as metal rods and metal plates with smooth surfaces, can be used without any restrictions. In particular, when a comb-shaped electrode having multiple comb teeth (electrodes) is used, it is possible to further increase the efficiency of electrochemical polymerization fixation of the conductor 212 to the base fiber 211.

[0150] FIG. 14 shows an enlarged view of a main part of an example in which the multiple electrodes (reference numerals 202, 203) shown in FIG. 13 are composed of comb-tooth electrodes 221, 231 having multiple comb teeth. The number of comb teeth may be selected arbitrarily. The comb-tooth electrodes 221, 231 shown in FIG. 14 have multiple comb teeth 221a, 231a arranged in the longitudinal direction of the base fiber 211, and are arranged so as to sandwich the base fiber 11 from both radial sides of the base fiber 211, and the multiple comb teeth 221a, 231a are arranged so as to be alternately combined in the longitudinal direction of the base fiber 211 from both radial sides of the base fiber 11. The comb-tooth electrodes 221, 231 press and guide the multiple comb teeth 221a, 231a from both radial sides of the base fiber 211, and run the base fiber 211 in the direction of the arrow in FIG. 14 to pass electricity. At this time, a constant current or a constant voltage is supplied to the comb-shaped electrodes 221 and 231 by connecting a DC stabilized power supply (not shown) to terminals 221b and 231b.

[0151] As described above, the comb-tooth electrodes 221, 231 have anodes (comb teeth 221a) and cathodes (comb teeth 231a) arranged alternately along the running direction of the base fiber 211, and by coming into contact with the base fiber 211, electricity can be repeatedly passed through the base fiber 211 in a short period of time to polymerize and fix the PEDOT-PSS.

[0152] In this embodiment, first, the base fiber 211 is run in the vertical direction, thereby distributing the conductors 212 to be fixed evenly inside and outside the fiber bundle made of the base fiber 211, and preventing uneven distribution. Furthermore, comb-shaped electrodes 221, 231 are used as multiple electrodes, and comb teeth (anode) 221a and comb teeth (cathode) 231a are arranged alternately and continuously, so that the conductors 12 can be polymerized and fixed multiple times to the base fiber (fiber bundle) 211 in one run.

[0153] In addition, since the multiple comb teeth (electrodes) 221a, 231a are connected in parallel, the combined resistance between the electrodes is reduced, and the applied voltage can be set low. In general, when the applied voltage to each electrode is set high, problems such as electrolysis of water and deterioration due to heating of the polymer tend to occur, so it is preferable to set the applied voltage as low as possible within a range in which polymerization and fixation are possible. In this way, when the applied voltage to the electrodes is set low, it is necessary to efficiently pass the current required for polymer polymerization, so from this viewpoint as well, it is preferable to use the comb-shaped electrodes 221, 231 configured as described above. In this case, the applied voltage to the comb-shaped electrodes 221, 231 can be selected arbitrarily. For example, it can be set to a range of 0.1 to 18 (V).

[0154] In this embodiment, the configuration using the comb-shaped electrodes 221, 231 that are multipolar electrodes having a plurality of comb teeth 221a, 231a utilizes the characteristic that PEDOT-PSS (conductor 212) that has already been electrochemically polymerized does not usually decompose even if a current is passed in the reverse direction after polymerization. Therefore, in the process in which the base fiber 211 runs between the comb-shaped electrodes 221, 231, polymerization and fixation are performed when approaching the anode (positive electrode), and electrochemical polymerization is repeated by running between the comb-shaped electrodes 221, 231 that are multipolar electrodes, and the conductor 212 containing PEDOT-PSS is superposed on the base fiber 211.

[0155] The size of the comb-tooth electrodes 221, 231 is not particularly limited, but the inter-tooth distance (inter-electrode distance) of the multiple comb teeth 221a (231a) is preferably in the range of 1 to 50 mm, and from the viewpoint of the processing efficiency of electrochemical polymerization fixation, it is preferably about 10 mm.

[0156] In addition, the multiple electrodes (reference numerals 202, 203) shown in FIG. 13 are not limited to the comb-shaped electrodes 221, 231 as described above. For example, as the multiple electrodes, as shown in the detailed enlarged views of the main parts in FIGS. 15A to 15C, a plurality of rotor electrodes 222, 232 may be arranged in the longitudinal direction of the base fiber 211 and arranged so as to sandwich the base fiber 211 from both sides in the radial direction of the base fiber 211. In the example shown in FIGS. 15A to 15C, the rotor electrode 232 arranged on one side in the radial direction of the base fiber 211 is roller-shaped, and the rotor electrode 222 arranged on the other side is pulley-shaped. These rotor electrodes 222, 232 are arranged alternately in the longitudinal direction of the base fiber 211.

[0157] As shown in Figure 15C, when rotor electrodes 222, 232 are used as multiple electrodes, the roller-shaped rotor electrode 232 is pressed against the base fiber 211, and the base fiber 211 is guided by the groove portion 222b formed in the pulley-shaped rotor electrode 222 while running between the multiple electrodes 222, 232 to pass electricity.

[0158] 15A, the roller-shaped rotor electrode 232 is a negative electrode (-) in this embodiment, and is configured by assembling a roller 232a to a metal shaft portion 232c. The metal shaft portion 232c is connected to the negative side of a DC stabilized power supply (not shown), and therefore a conductive metal material is used for this metal shaft portion 232c. The roller 232a is also a small metal roller made of, for example, stainless steel material and incorporating a ball bearing for rotation.

[0159] 15C, the base material fiber 211 running in the vertical direction comes into contact with the outer circumferential surface 232b of the roller-shaped rotor electrode 232. For this reason, the size of the roller 232a can be selected as necessary, but taking into consideration the running speed of the base material fiber 211, for example, a diameter of about 6 mm and a width of about 3 mm can be used.

[0160] 15B, the pulley-shaped rotor electrode 222 is an anode (+) in this embodiment, and is configured by assembling a pulley 222a having a groove 222b formed on its outer circumferential surface to a metal shaft portion 222c. The metal shaft portion 222c is connected to the (+) side of a DC stabilized power supply (not shown), and therefore, a conductive metal material is used for this metal shaft portion 222c, as in the case of the roller-shaped rotor electrode 232. The pulley 222a is also a small metal pulley made of, for example, stainless steel and incorporating a ball bearing for rotation.

[0161] 15C, the pulley-shaped rotor electrode 222 is guided by the vertically running base fiber 211 in contact with groove 222b formed on the outer circumferential surface of the pulley-shaped rotor electrode 222. For this reason, the size of the pulley 222a can be selected as necessary, as in the case of the roller-shaped rotor electrode 232, but taking into consideration the running speed of the base fiber 11, for example, a pulley with a diameter of about 8 mm and a width of about 4 mm can be used.

[0162] Also, as shown in FIG. 15C, in this embodiment, the pulley-shaped rotor electrode 222 and the roller-shaped rotor electrode 232 are arranged alternately along the vertical running direction of the base material fiber 211, i.e., the anodes (+) and cathodes (-) are arranged alternately (see also symbols 202 and 203 in FIGS. 15A to 15C).

[0163] Then, the base fiber (fiber bundle) 211 impregnated with and / or attached with the conductor 212 comes into contact with a pulley-shaped rotor electrode (anode) 222 made of metal and incorporating a bearing, and a roller-shaped rotor electrode (cathode) 32, which are fixed to metal shaft portions 222c and 232c, respectively. By supplying a constant current or low voltage from a stabilized DC power supply (not shown) to each of these rotor electrodes 222 and 232, the conductor 212 containing PEDOT-PSS is electrochemically polymerized and fixed to the base fiber 211.

[0164] When using rotor electrodes 222, 232 as described above, the amount of electricity required for polymerizing and fixing conductor 212 containing PEDOT-PSS can be selected arbitrarily. For example, when silk thread (No. 9: silk thread: manufactured by Fujix Co., Ltd.) having a diameter of approximately 280 μm is used as the base fiber, the amount of electricity is 0.1 to 6 mC per 10 mm, and particularly, good polymerizing and fixing can be achieved at about 3 mC.

[0165] The rotor electrodes 222, 232 as described above rotate together with the traveling base fiber 211. This reduces friction due to contact, and prevents the conductor 212 containing PEDOT-PSS electrochemically fixed to the base fiber 211 from peeling off due to friction. That is, it is possible to prevent the surface destruction of the polymer in the conductive polymer fiber 201 from occurring.

[0166] In addition, in the manufacturing apparatus 210 of this embodiment, the arrangement shape of the thread-like fiber bundles is adjusted by pressing the roller-shaped rotor electrode 232 against the base fiber 211 and guiding it with the groove 222b of the pulley-shaped rotor electrode 222. This makes it possible to adopt a configuration that allows current to be applied while adjusting the amount of conductor 212 impregnated into and / or attached to the base fiber 211. In this case, for example, the shapes and arrangement of the pulley-shaped rotor electrode 222 and the roller-shaped rotor electrode 232, as well as the tension, running speed, and rotation state of the base fiber (fiber bundle) 201 are adjusted. This makes it possible to adjust the opening and bundling function, fiber spacing, and fiber bundle arrangement (shape), etc. of the base fiber (fiber bundle).

[0167] By adjusting the gaps between the base fibers by making the above-mentioned adjustments, it becomes possible to hold and fix any amount of conductor 212 (42) between the base fibers, for example, as in the example shown in Fig. 6. In particular, by setting the shape of base fiber (fiber bundle) 211 using pulley-shaped rotor electrode 422 and roller-shaped rotor electrode 232, it becomes possible to adjust to various shapes as described above.

[0168] Here, the set shape of the fiber bundle may be, for example, with or without twisting of the fiber bundle, the cross-sectional shape of the fiber bundle (flat, circular, elliptical, rectangular, etc.), and the twisting of the fiber bundle (reversing the twist to straighten the fiber bundle, or adding further twist), and the shape may be set by appropriately selecting and adjusting. In this embodiment, by setting the shape of the fiber bundle as described above, a composite fiber bundle of conductive polymer fibers 201 set (molded) into a predetermined shape is obtained. When forming such a composite fiber bundle, as in the example shown in Figure 6, a conductor 212 (42) containing the conductive polymer PEDOT-PSS is arranged in close contact with multiple base fibers 211 (41), and the multiple base fibers 211 (41) are twisted or knitted to form a higher-order structure such as a twisted string, woven fabric, or nonwoven fabric.

[0169] As shown in FIG. 13, in addition to the above configuration, the manufacturing apparatus 210 of this embodiment can further include a drying section 208 that blows air toward the base fiber 211 to which the conductor 212 is polymerized and fixed, and a chamber (including a humidity control section) 207 that adjusts the atmospheric humidity in the vicinity of the base fiber 211.

[0170] The chamber 207 has a humidity and temperature control function (humidity control section) and maintains a high humidity level inside the chamber to maintain a constant concentration (PEDOT-PSS concentration) of the conductor 212. As the chamber 207, a thermo-hygrostat tank or the like that has been conventionally used in this field and is sized to accommodate the immersion container 205 and the multiple electrodes 202 and 203 can be used without any restrictions.

[0171] The drying section 208 blows low-humidity dry air toward the base fiber 211 (conductive polymer fiber 201) to which the conductor 212 is polymerized and fixed, thereby drying the same. For example, any conventionally known blowing and drying means composed of a motor, a fan, etc. can be used without any restrictions.

[0172] In addition, in this embodiment, by adding the above-mentioned chamber and adjusting the humidity near the multiple electrodes 202, 203, it is possible to adjust the moisture content of the PEDOT-PSS solution (conductor 212) impregnated into the base fiber (fiber bundle) 211. Furthermore, when the chamber is configured with three compartments and the humidity of each compartment is adjusted independently, the following settings (A) to (C) can be used. (A) Immersion container: In order to prevent evaporation of water from the conductive solution containing PEDOT-PSS and to keep the concentration of PEDOT-PSS constant, the humidity setting is adjusted to be in the range of 50 to 100%, for example. (B) Multiple electrodes... By setting the humidity to a range from high to low, for example, 99 to 10%, the moisture content of the conductive solution containing PEDOT-PSS with which the fibers are impregnated is adjusted. (C) Drying section: A function for blowing dry air has been added to promote the drying of the base fiber (conductive polymer fiber) to which the conductor is polymerized and fixed by circulating low-humidity dry air (for example, humidity setting in the range of 0 to 40%).

[0173] In addition to the above-mentioned components, the manufacturing apparatus of this embodiment may further include a container-shaped disinfecting and cleaning unit that performs fixation and sterilization of residual monomers using, for example, an ethanol or acetone bath, although this is not shown in the drawings. Furthermore, the disinfecting and cleaning unit may be provided with a configuration capable of removing residual monomers using a cleaning bath of physiological saline or the like.

[0174] Furthermore, the electrodes used in the present invention are not limited to a plurality of electrodes as shown in FIGS. 13 to 15, and may be, for example, a manufacturing apparatus 250 equipped with unipolar (single) electrodes 252, 253 as shown in FIG. This manufacturing apparatus 250 includes an immersion container 255 that contains a PEDOT-PSS solution 204, an electrode (negative electrode) 253 made of a metal plate or the like that is placed in the PEDOT-PSS solution inside the immersion container 255, an electrode (anode) 252 made of a metal rod or the like that is placed outside the immersion container 255 and in contact with the base fiber 211, a spool 256 that is placed in the immersion container 255 and around which the base fiber 211 is wound, a chamber 257 that contains the immersion container 255, the spool 256, and each of the electrodes 252, 253 and controls the humidity inside, a DC stabilized power supply 251 that supplies current to the electrodes 252, 253, a drying section 258 that blows air to dry the base fiber 201 (conductive polymer fiber 201), and a winding section 259 that winds up the completed conductive polymer fiber 201.

[0175] Even when using such a manufacturing apparatus 250 equipped with single (monopolar) electrodes 252, 253, the base fiber 211 can be pulled up vertically while running between the electrodes, thereby evenly dispersing the conductor 212 to be polymerized and fixed to the base fiber 211 and preventing uneven distribution, thereby making it possible to produce a conductive polymer fiber 201 with excellent conductivity and durability.

[0176] "Manufacturing method" The procedure for producing conductive polymer fibers 201 using the above-mentioned production apparatus 210 will be described below with reference to the same drawings (FIGS. 13 to 15) as those used to explain the above-mentioned production apparatus. The manufacturing method of the conductive polymer fiber 201 described in this embodiment includes the steps (1) to (3) shown below in order, and each of the steps (1) to (3) is carried out while controlling the atmospheric humidity. (1) An immersion process in which an insulating base fiber 211 consisting of a thread-like fiber bundle is immersed in a solution of a conductor containing PEDOT-PSS as a conductive polymer, thereby impregnating and / or attaching a conductor 212 to the base fiber 211. (2) A fixation process in which the base fiber 211 is pulled vertically out of a conductive solution while running between multiple electrodes 202, 203 and passing an electric current through it, thereby electrochemically polymerizing and fixing the conductor 212 impregnated and / or attached to the base fiber 211. (3) A drying step of blowing air to dry the base fiber 211 to which the conductor 212 has been polymerized and fixed.

[0177] (Soaking process) In the immersion process, as described above, the base fiber 211 is immersed in a solution of a conductor containing PEDOT-PSS as a conductive polymer, thereby impregnating and / or attaching the conductor 212 to the base fiber 211. Specifically, a conductor solution containing PEDOT-PSS, which is a conductive polymer, is placed in an immersion container 205 as shown in Fig. 13, and a base fiber (fiber bundle) 211 is immersed in this solution. As a result, a conductive conductor 212 is impregnated into and / or attached to the base fiber 211, and the base fiber 211 becomes conductive.

[0178] When preparing a solution of a conductor containing a conductive polymer as described above, for example, additives and the like can be added to a commercially available PEDOT-PSS solution (Heraeus: CLEVIOS P, etc.) as necessary. That is, a method can be applied in which a mixed solution is prepared by mixing an additive into a solution of a conductor containing PEDOT-PSS, and the conductive polymer and the additive are simultaneously applied to or immersed in the base fiber 211. Such a mixed solution can be selected arbitrarily. One example is an aqueous solution containing a conductive polymer such as PEDOT-PSS at a concentration of 0.1 to 50 (V / V) % and an additive such as glycerol at a concentration of 0.1 to 50 (V / V) %. The concentration of the additive in the conductor solution is not particularly limited, and can be in the range of, for example, 0.1 to 50 wt %.

[0179] (Fixed process) Next, in the fixing step, the base fiber 211 is pulled vertically out of the solution while being transported between the multiple electrodes 202, 203 and an electric current is applied thereto. This causes the conductor 212 impregnated in and / or attached to the base fiber 211 to be electrochemically polymerized and fixed.

[0180] Specifically, for example, the base fiber (fiber bundle) 211 impregnated with and / or attached to the conductor 212 is pulled up vertically from the solution and run by a winding unit 209 as shown in Fig. 13. By pulling the base fiber 211 up vertically and running in this manner, the conductor 212 is not centered due to gravity, and is evenly distributed and impregnated with and / or attached to the base fiber 211.

[0181] 13, a current is applied while the base fiber 211 is sandwiched from both radial sides of the base fiber 211. In this manner, by contacting the base fiber (fiber bundle) 211 impregnated and / or attached with the conductor 212 containing PEDOT-PSS with the electrodes 202, 203 and passing a current, the conductor 212 containing PEDOT-PSS attached inside and outside the fiber bundle is electrochemically polymerized and fixed, and a conductive polymer fiber 201, which is a composite fiber of the conductor 212 containing PEDOT-PSS and the base fiber (fiber bundle) 211, is obtained.

[0182] In the fixing step, the comb-tooth electrodes 221, 231 having the plurality of comb teeth 221a, 231a shown in Fig. 14 as described above can be used as the plurality of electrodes. In this case, the comb-tooth electrodes 221, 231 are arranged so as to sandwich the base fiber 211 from both sides in the radial direction of the base fiber 211, and the plurality of comb teeth 221a, 231a are arranged so as to be alternately combined in the longitudinal direction of the base fiber 211 from both sides in the radial direction of the base fiber 211. Then, a method can be adopted in which the base fiber 211 is made to run vertically while the plurality of comb teeth 221a, 231a of the comb-tooth electrodes 221, 231 are pressed against and guided from both sides in the radial direction against the base fiber (fiber bundle) 211, and a current is applied to the base fiber 211.

[0183] In addition, in the fixing process, it is also possible to use rotor electrodes 222, 232 as multiple electrodes, which are arranged in the longitudinal direction of the base fiber 211 and are arranged so as to sandwich the base fiber 211 from both radial sides of the base fiber 211, as shown in Figures 15A to 15C. That is, a roller-shaped rotor electrode 232 is arranged on one radial side of the base fiber 211, and a pulley-shaped rotor electrode 222 is arranged on the other radial side, and the rotor electrodes 222, 232 arranged on both sides of the base fiber 211 are arranged alternately in the longitudinal direction of the base fiber 211. A method can be employed in which the roller-shaped rotor electrode 232 is pressed against the base fiber (fiber bundle) 211, and the base fiber 211 is guided by a groove 222b formed in the pulley-shaped rotor electrode 222 while the base fiber 211 is made to run between the multiple electrodes to pass electricity.

[0184] Furthermore, in the fixing process, as described in the configuration of the above manufacturing apparatus, it is also possible to adopt a method in which the base fiber 211 is pressed against the roller-shaped rotor electrode 232 while being guided by the groove portion 222b of the pulley-shaped rotor electrode 222 to adjust the arrangement shape of the thread-like fiber bundle, thereby passing electricity while adjusting the amount of conductor 212 impregnated into and / or attached to the base fiber 211.

[0185] (drying process) Next, in the drying step, low-humidity dry air is blown toward the base fiber 211 to which the conductor 212 has been polymerized and fixed, that is, the conductive polymer fiber 201, to dry the conductive polymer fiber 201.

[0186] 13, for example, dry air is blown onto the base fiber (fiber bundle) 211 to which the conductor 212 has been electrochemically polymerized and fixed in the fixing step, using a drying unit 208 equipped with a humidity control means and a blowing means (not shown). This dries and removes the water (solvent) contained in the solution of the conductor 212 containing PEDOT-PSS.

[0187] In the manufacturing method of this embodiment, thereafter, it is preferable to wash the conductive polymer fibers 201 with an electrolyte solution such as physiological saline to remove unpolymerized PEDOT-PSS and the solvent. Furthermore, in this embodiment, it is preferable to clean and disinfect the conductive polymer fiber 201 using an ethanol solution, and then dry it.

[0188] Although detailed explanation will be omitted, in the manufacturing method of this embodiment, it is also possible to manufacture conductive polymer fibers 201 using a manufacturing apparatus 250 equipped with unipolar (single) electrodes 252, 253 as shown in FIG. 16.

[0189] According to the manufacturing method of the conductive polymer fiber 201 of the present invention as described above, the base fiber 11 impregnated with and / or attached with the conductor 212 containing PEDOT-PSS is vertically pulled up from the conductor solution while running between the multiple electrodes 202, 203 and energized. This allows the process of electrochemically polymerizing and fixing the conductor 212 to the base fiber 211 to be continuously performed in a single step, improving productivity. Furthermore, by running the base fiber 211 between the multiple electrodes 202, 203 while pulling it up vertically, the conductor 212 to be polymerized and fixed to the base fiber 211 can be evenly distributed and uneven distribution can be prevented. Therefore, it is possible to produce the conductive polymer fiber 201 with high biocompatibility, good homogeneity, and excellent conductivity and durability with good productivity.

[0190] In addition, the conductive polymer fiber manufacturing apparatus 210 according to the present invention employs a configuration including a winding section 209 that vertically pulls up the base fiber 211 impregnated and / or attached with the conductor 212 containing PEDOT-PSS from the conductor solution in the immersion container 5, and a plurality of electrodes 202, 203 that apply electricity while running the base fiber 211. This makes it possible to continuously perform the process of electrochemically polymerizing and fixing the conductor 212 to the base fiber 211 in a single step, thereby improving productivity. Furthermore, by providing a plurality of electrodes 202, 203 that apply electricity while pulling up the base fiber 211 vertically, the conductor 212 that is polymerized and fixed to the base fiber 211 can be evenly distributed and uneven distribution can be prevented. Therefore, the conductive polymer fiber 201 that has high biocompatibility, good homogeneity, and excellent conductivity and durability can be obtained with high productivity.

[0191] Regarding the third aspect The third aspect of the present invention relates to a bioelectrode and a biosignal measuring device. More specifically, the present invention relates to a body surface-mounted bioelectrode using a composite material of a conductive polymer and a fiber (hereinafter, conductive composite fiber), and a biosignal measuring device equipped with the bioelectrode. In this aspect, the fiber described in the first aspect of the present invention can be preferably used. Hereinafter, an embodiment of the third aspect of the present invention will be described with reference to the drawings, but the present invention is not limited to this embodiment.

[0192] Example of using bioelectrodes as electrodes for measuring brain waves In recent years, EEG measurement has been expanded beyond testing in medical institutions to include home EEG testing, remote medical care, health information, and ubiquitous healthcare systems. In addition to the medical field, it is also expected that this technology will be applied to psychological research using event-related potential measurements, engineering such as BCI (brain-computer interface), and the nursing and welfare fields.

[0193] When measuring brain waves, it is necessary to place electrodes aside the hair on the scalp. In conventional EEG measurements using bioelectrodes, in order to stabilize and fix the electrodes, the electrodes are fixed to the skin using adhesive, the electrodes are compressed and fixed from above using a head cap that covers the entire head, or the amount of paste or gel between the electrodes and the scalp is increased to prevent the electrodes from floating up. However, these measures are inconvenient to wear and place a heavy burden on the subject, which is particularly problematic when performing continuous EEG measurements over a long period of time. In addition, the appearance of the electrodes can cause some discomfort to the wearer or others, so the use of EEG has not yet become widespread beyond medical use.

[0194] The bioelectrode of the first embodiment of the third aspect of the present invention described below utilizes the conductivity of the conductive polymer, making it possible to miniaturize the electrode and reduce the contact area with the skin. Furthermore, this electrode, which is made of a soft fiber material, causes less irritation to the skin when worn and is less likely to cause discomfort while wearing it. In addition, there is no need to seal the skin with highly adhesive gel or tape, as in the case of conventional bioelectrodes. The bioelectrode of the first embodiment of the third aspect of the present invention is excellent in fit, can be used continuously, and does not look uncomfortable when worn, so it is preferably used, for example, in applications such as electroencephalography.

[0195] [First embodiment of the third aspect] The bioelectrode of the first embodiment includes a string-like contactor made of conductive composite fiber. Brain waves can be measured by placing the contactor in close contact with the scalp through the gaps in the hair (see Figs. 19A to 19D and Figs. 20A to 20D).

[0196] The bioelectrode 310 shown in FIG. 19A to FIG. 19D includes at least a string-like contactor 311 made of conductive composite fiber, a first frame 312, and a second frame 313 (connecting portion). Both ends of the contactor 311 are fixed to both ends of the arch-shaped first frame 312. The shape of the flexible contactor 311 is adjusted by the first frame. A tension may be applied to the contactor 311 by the first frame 312. The second frame 313, which is horizontally supported by the first frames 312, functions as a beam for fixing each of the first frames 312. In addition, the end of each contactor 311 is connected to a signal cable 314. An electric signal is transmitted and received between each contactor 311 and an electroencephalogram analysis device (not shown) connected to the end of the signal cable 314. The direction of the electric signal may be unidirectional or bidirectional.

[0197] The shape of the contact 311 is not particularly limited as long as it can be brought into contact with the scalp S, and may be any shape such as string, thread, band, cloth, net, etc. The size and length of the contact 11 are appropriately adjusted.

[0198] The shape, number, and size of the first frame 312 and the second frame 313 are not particularly limited. For example, a shape capable of applying an appropriate tension to the contact 311 or a shape capable of fixing the contact 311 so as not to loosen can be adopted. The material constituting the first frame 312 and the second frame 313 is not particularly limited as long as it is a material that does not disturb the electric signal in the contact 311, and a conventionally known resin material can be used. Examples of the number of the first frames 312 include a range of 1 to 20, a range of 2 to 8, and a range of 2 to 4. Examples of the number of the second frames 313 include a range of 1 to 6, a range of 1 to 3, and a range of 1 to 2. The shapes of the first frames 312 and the second frames 313 may be, for example, a partial loop shape or a plate shape. The thickness is preferably constant, but may be partially different. In the figure, one second frame 313 is arranged at a right angle to the four first frames 312. A plurality of second frames 313 may be arranged, and may be arranged diagonally as necessary. Furthermore, as long as there is no problem with electroencephalogram measurement (signal measurement), at least one of the first frame 312 and the second frame 313 may be made of metal. For example, by making the first frame 312 and the second frame 313 out of a metal material, the signal cable 314 may be connected to the first frame 312 or the second frame 313 without being connected to the contactor 311, and transmission and reception of electrical signals to and from the contactor 311 may be performed via the first frame 312 or the second frame 313.

[0199] The method of connecting the conductive composite fiber constituting the contact 311 and the signal cable 314 is not particularly limited as long as it is a method that allows electrical connection. For example, any of the following methods can be applied: crimping with a metal, wrapping or tying the conductive composite fiber around the signal cable 314, or adhesion with a conductive adhesive.

[0200] In the present invention, one or more string-like contacts 311 can be used per electrode. As illustrated in Figs. 19A to 19D, a comb-shaped bioelectrode 310 is configured using a first frame 312 and a second frame 313, thereby obtaining stable contact between the contacts 311 and the skin. In the comb-shaped bioelectrode 310 illustrated in Figs. 19A to 19D, a plurality of string-like contacts 311 are arranged in parallel. The comb-shaped bioelectrode 310, in which a plurality of contacts 311 are arranged in parallel, can be inserted between hair or at the hairline like a comb, and fixed by covering the comb with a net-shaped holder (Figs. 22A to 22B).

[0201] (Modification of the first embodiment of the third aspect) As a further miniaturized bioelectrode configuration, a hairpin-shaped bioelectrode 320 with two contacts 321 fixed to a hairpin-shaped hair clip (metal leaf spring) is shown in Figs. 20A to 20D. This hairpin-shaped bioelectrode 320 can be used by inserting it near the root of the hair. The hairpin-shaped hair clip can hold the hair. In the illustrated example, the two contacts 321 are fixed on the scalp S.

[0202] The hairpin-shaped bioelectrode 320 shown in Figs. 20A to 20D includes at least a string-like contactor 321, a third frame 322, and a fourth frame 323 made of conductive composite fiber. Both ends of the contactor 321 are fastened between two cylindrical third frames 322. The tension applied to the contactor 321 can be adjusted by adjusting the distance between the two third frames 322. The two third frames 322 are fixed to the tip and bent portion of the fourth frame 324 using a hairpin, respectively. In the example shown in Figs. 20A to 20D, two string-like contactors 321 are provided on the hairpin. In addition, the end of each contactor 321 is connected to a signal cable 324. An electric signal is transmitted and received between each contactor 321 and an electroencephalogram analysis device (not shown) connected to the end of the signal cable 324. The direction of the electric signal may be unidirectional or bidirectional.

[0203] The shape of the contactor 321 is not particularly limited as long as it can be brought into contact with the scalp S, and may be any shape such as string, thread, strip, cloth, net, etc. The size and length of the contactor 321 are appropriately adjusted.

[0204] The shape of the third frame 322 is not particularly limited, and may be, for example, a cylindrical shape, a polygonal prism such as a triangular prism or a square prism, or a spherical shape. In this configuration, the fourth frame 323 has a hairpin structure and functions as a hairpin, so that the fourth frame 323 can be fixed to the hair H. As a result, the contactor 321 can be easily brought into contact with the skin (scalp) S and fixed at a desired position.

[0205] The material constituting the third frame 322 and the fourth frame 323 is not particularly limited as long as it does not disturb the electrical signal in the contactor 321, and for example, a conventionally known resin material can be used. Furthermore, at least one of the third frame 322 and the fourth frame 323 may be made of metal as long as it does not interfere with the electroencephalogram measurement (signal measurement). In this modification, for example, the third frame 322 may be made of insulating resin, and the hairpin which is the fourth frame 323 may be made of metal. As long as it does not interfere with the electroencephalogram measurement, the contactor 321 may be electrically connected to the fourth frame 323 which is made of metal.

[0206] One or more string-like contacts 321 can be used per electrode. As illustrated in Figs. 20A to 20D, a hairpin-shaped 320 electrode is formed using a third frame 322 and a fourth frame 323, thereby providing stable contact between the contacts 321 and the skin. In the hairpin electrode 320 shown in Figs. 20A to 20D, a plurality of string-like contacts 321 are arranged in parallel. The hairpin electrode 320, in which a plurality of contacts 321 are arranged in parallel, can be fixed in place in the same way as hair is fixed with a hairpin.

[0207] The conductive composite fiber constituting the string-like contact 321 may be the same as that of the contact 311 described above.

[0208] (Conductive composite fiber) A composite fiber of a conductive polymer and a conventionally known fiber material can be applied as the conductive composite fiber constituting the string-like contacts 311, 321. The form (method) of the composite is not particularly limited, and for example, the conductive polymer may be coated on the surface of the string-like (yarn-like) fiber material, the conductive polymer may be impregnated into the string-like fiber material, or the string-like conductive polymer and the string-like fiber material may be twisted or spun together. The materials and conductive polymer fibers described in the first embodiment may be preferably used, and the apparatus and method described in the second embodiment may also be used.

[0209] The type of the conductive polymer is not particularly limited, and known conductive polymers can be used. Examples of the conductive polymer include hydrophilic conductive polymers such as PEDOT-S (poly(4-(2,3-dihydrothieno[3,4-b][1,4]dioxin-2yl-methoxy-1-butanesulfonic acid, potassium salt) in addition to the above-mentioned PEDOT-PSS. By using a composite fiber containing a hydrophilic conductive polymer as the material for the contacts 11 and 21, the contacts 11 and 21 themselves can be easily imparted with adhesiveness (adhesiveness) to the skin.

[0210] As the fiber material, for example, conventionally known fiber materials such as silk, cotton, hemp, rayon, and chemical fibers can be used. Among these, silk is preferable. When silk is used, the strength and hydrophilicity of the composite fiber can be further improved. In addition, when silk is used, the wearing comfort when touching the skin is better. The type of the conductive polymer to be combined with silk is not particularly limited, but the above-mentioned hydrophilic conductive polymer such as PEDOT-PSS or PEDOT-S is preferred.

[0211] As the conductive composite fiber constituting the contact of each embodiment of the present invention, a conductive polymer fiber, which will be described in more detail later, can be used.

[0212] (Contact Structure) As the structure of the string-like contacts constituting each embodiment of the present invention, two types of structures are exemplified. The first structure of the string-like contactor is a structure in which the conductive composite fiber bundle is used alone. An example of the first structure is the contactor 311 shown in Figs. 19A to 19D. The contactor 311 has an appropriate thickness and strength because it is made by weaving a thread (string) made of a bundle of conductive composite fibers. The first structure made only of conductive composite fibers is flexible, and is therefore suitable for applications in which flexibility and a comfortable fit are required for the bioelectrode. The conductive composite fiber is preferably in the form of a thread or string.

[0213] The second structure of the string-like contactor is a structure in which the conductive composite fiber bundle is combined with a metal cable or a metal thin wire. Examples of the second structure include the structures shown typically in Figures 21A and 21B. Contactor 321 in Figure 20 has the structure shown in Figure 21B. Since the second structure has enhanced conductivity due to the metal cable or metal thin wire, it is suitable for applications in which it is necessary to reduce the electrode resistance per contact area of ​​the bioelectrode.

[0214] In the second structure shown in FIG. 21A, a metal thin wire 321f is wound around a bundle of multiple core materials 321g, and a conductive composite fiber 321e is wound around the metal thin wire 321f. The material constituting the core material 321g may be conductive or insulating. In the illustrated example, the core material 311 is a core made of insulating fiber. The number and thickness of the core material 311 are appropriately adjusted. In the figure, the number of turns of the metal wire and the number of turns of the conductive composite fiber 321e are drawn to be approximately the same, but the relative relationship of the number of turns is not limited to this. For example, the number of turns of the metal wire 321f may be less than the number of turns of the conductive composite fiber. Also, the string-like contact may be partially covered with an insulating cover 321z as necessary. An example of the insulating cover is a cover made of silicone resin.

[0215] In the second structure shown in FIG. 21B, a conductive composite fiber 321a is wound around a bundle of a plurality of metal cables 321b. The type of metal constituting the metal cable is not particularly limited, and a copper wire having high conductivity is preferable. In the illustrated example, the metal cable 321b is a copper wire. The number and thickness of the metal cables 321b are not particularly limited, and are adjusted as appropriate. In some cases, the use of a large number of thin metal cables or thin metal wires can increase flexibility even with the same diameter, rather than using a small number of thick metal cables. If necessary, the string-like contactor may be partially covered with an insulating cover 321c. An example of the insulating cover is a cover made of silicone resin. FIG. 21C is a photograph of the string-like contactor 321 having the structure of FIG. 21B.

[0216] Comparing the first and second structures, in the first structure, electrical connection between contactor 311 and metallic (conductive) signal cable 314 is made at one point at the end of contactor 311, whereas in the second structure, electrical connection between contactor 321 and signal cable 324 is made over the entire contactor 321. Therefore, in the second structure, the distance between the skin and the metallic cable is shorter than in the first structure, and the electrode resistance is reduced.

[0217] The thickness of the contact is not particularly limited, but is preferably a thickness that provides structural strength that makes it difficult to break when it comes into contact with the skin. For example, a thickness of 0.1 mm to 5 mm makes it easy to provide structural strength that makes it difficult to break. Usually, the second structure having the metal wire tends to have stronger structural strength than the first structure. Other examples of the thickness range include a thickness of 0.1 mm to 3 mm and a thickness of 0.5 mm to 1 mm.

[0218] Since the contactor having the conductive composite fiber has adhesiveness (adsorption) to the skin, the bioelectrode of the first embodiment can be installed independently without using paste or adhesive for electrode installation. However, the bioelectrode may come off (peel off) from the skin surface when an external force such as the subject's movement or pulling of the signal cable is applied. To prevent this, a means for pressing the bioelectrode against the skin surface can be adopted. For example, a net-shaped holder (cap) N as shown in FIG. 22A can be exemplified as the said means.

[0219] (Electroencephalogram electrode holder; Elastic net cap) The bioelectrode of the first embodiment of the third aspect, for example, the comb-shaped electrode 310 shown in Fig. 22A, can be fixed by covering it from above with the elastic net cap N shown in Fig. 22A. The net N can be used as a holder that lightly presses the comb-shaped electrode 310 from above and stably holds it. The comb-shaped electrode 310 is inserted into the gaps between the hair, so it is difficult to lift up. For this reason, the comb-shaped electrode 310 does not need to be strongly pressed and fixed with a head cap or the like as in the conventional electrodes, and a stable fixation can be obtained by using a cover such as a low-tension elastic net. As the low-tension elastic net, for example, a commercially available net bandage (manufactured by Nihon Eizai Co., Ltd.) can be applied. With the bioelectrode of the first embodiment, a design that fits under the hair can be easily realized. Furthermore, when using the elastic net, it is also possible to pull the hair out of the net. Therefore, by using the bioelectrode of the first embodiment, both the wearing comfort and the appearance during wearing are improved. FIG. 22A shows a cap of the comb-shaped electrode 310 and the elastic net N. FIG. 22B shows an example (top view) in which the comb-shaped electrode 310 is arranged on the elastic lattice-like net N. In FIG. 22B, △ represents the nose, and the two ellipses represent the left and right ears, respectively. In the area divided by the lattice, the position indicated by the diagonal lines represents the location where the comb-shaped electrode 310 is arranged. In this configuration, by adjusting the spacing between the strings of the stretchable net N, the placement site of the bioelectrode 310 can be adapted to conform to the international 10-20 system.

[0220] Example of using bioelectrodes as electrodes for measuring electrocardiograms Conventionally, bioelectrodes for Holter electrocardiogram examinations and bioelectrodes for monitoring heart rate or myoelectric potential have been widely used. Electrodes for Holter electrocardiogram examinations are often used in a state where they are fixed to the skin using highly adhesive tape or adhesive pads. By fixing the bioelectrodes to the skin, noise generation is prevented. Also, in order to fix the electrodes for monitoring, which are often used continuously for long periods of time, to the skin, adhesive pads made of conductive gel are often used. The measurement data obtained by these electrodes is less likely to be mixed with artifacts such as noise, and the measurement waveform is highly stable.

[0221] However, there is a problem of attenuation of high frequency components of biosignals. This problem is thought to be caused by the effect of the capacitance of the electrolyte solution, which is used in conventional bioelectrodes because electrolyte paste or gel is used between the metal electrode plate and the skin. Therefore, electrolyte paste and electrolyte gel are one of the factors that make it difficult to analyze biosignals that contain high frequency components and to achieve high-speed communication between the living body and external devices in BCI, etc. In addition, in conventional bioelectrodes, the electrodes, which have high adhesiveness, are attached to the skin, which tends to cause sweating and discomfort to the subject (wearer). In addition, as a pretreatment to obtain the effectiveness of the adhesive, it is necessary to degrease the adhesive surface of the skin with alcohol cotton or the like. However, degreasing treatment with alcohol is highly irritating to the skin, which is one of the causes of itching and contact dermatitis, and improvements are required.

[0222] Thus, the problem of the influence on frequency characteristics and the sweating caused by using a metal electrode plate is not limited to electrocardiogram measurement, but is a problem that must be solved when transmitting electrical signals or electrical stimuli between a living body and an electrode, including the above-mentioned electroencephalogram measurement. Attempts to improve these problems have been made in the past, but have not been sufficient.

[0223] For example, in order to reduce the effect of electrolyte paste on frequency characteristics, attempts have been made to measure EEG by placing small sintered metal electrodes directly on the skin without using electrolyte paste. However, this method causes problems with the stability of the measured waveform. That is, in the method of placing metal electrodes directly on the skin, the resistance between the skin and the electrode is likely to fluctuate due to mechanical compliance and electrochemical mismatch between the metal electrodes and the skin. Furthermore, the measurement signal is likely to become unstable due to vibrations such as body movements and breathing of the living body, and noise is often mixed into the measurement signal. In addition, there are problems to be solved, such as the fact that hard metal electrodes tend to cause discomfort or discomfort when directly contacted with the skin.

[0224] In order to reduce the problems of rigid metal electrodes, textile electrodes using conductive fibers have been developed in recent years and are becoming popular mainly in the fields of sports and health. Textile electrodes are cloth-like bioelectrodes incorporating conductive fibers, and are used by pressing and fixing them to the skin using elastic bands or the like. In textile electrodes, no electrolyte paste is used, and the mainstream is the pasteless type, in which the electrode is directly in contact with the skin or the cloth that constitutes the electrode is moistened. In measurements using textile electrodes, relatively good biosignals can be obtained if the contact with the skin is maintained stably. However, if the contact with the skin becomes even slightly unstable, the resistance between the skin and the electrode fluctuates significantly, and the reliability of the measured waveform is easily reduced due to artifacts such as distortion of the recorded waveform and the inclusion of hum noise.

[0225] The bioelectrode of the second embodiment of the third aspect of the present invention described below utilizes the conductivity of the conductive polymer, as in the bioelectrode of the first embodiment, and therefore can be made smaller and the contact area with the skin can be reduced. Furthermore, this electrode made of a soft fiber material causes less irritation to the skin when worn and is less likely to cause discomfort during wearing. In addition, there is no need to seal the skin with highly adhesive gel or tape, as in the case of conventional bioelectrodes. The bioelectrode of the second embodiment of the present invention is excellent in fit, can be used continuously, and does not look uncomfortable when worn, so it is preferably used, for example, in medical or sports textile applications.

[0226] [Second embodiment of the third aspect] A bioelectrode 330 of a second embodiment is shown in Figures 24A to 24B. The bioelectrode 330 is provided with a contact portion (electrode surface) 332 in which a plurality of string-like contacts 331 made of conductive composite fibers are arranged in a plane, and a sheet-like substrate 333 that supports the contact portion 332. The combination of the contact portion 332 and the substrate 333 is called an electrode pad. A signal cable 334 electrically connected to each contact 331 is provided. Furthermore, a holder 335 made of a stretchable material is provided as a means for pressing the contact portion 332 of the electrode pad against the skin S.

[0227] The shape of the planar contact portion 332 and the substrate 333 in which a plurality of contacts 331 made of conductive composite fiber bundles are arranged is not particularly limited as long as the shape allows contact between the contact portion 332 and the skin to be obtained over a surface, and does not necessarily have to be flat. In other words, the contact portion 332 or the substrate 333 may form a curved surface, concave portion, or convex portion along the curved surface of the skin. The shape of the electrode pad does not need to be rigidly fixed, and may flexibly deform in accordance with the contact with the skin.

[0228] By using the sheet-like substrate 332, the flatness of the skin contact surface of the bioelectrode can be ensured, and stable adhesion to the skin can be promoted. The material, size, and shape of the substrate can be selected arbitrarily. For example, the substrate can be a 0.2 mm thick PVC (polyvinyl chloride) sheet or a flat silicone sheet (1 mm thick). The substrate material is not limited to these, and a flexible film-like (sheet-like) material that easily maintains the flatness of the substrate and has good adhesion to the skin is preferably used.

[0229] Contacts 331 made of conductive composite fibers are arranged and fixed on the substrate to form a contact portion 332 (electrode pad) with the skin. One side of the sheet-like substrate (the side on which contact portion 332 is provided) may be given a slight adhesive force to allow the substrate surface to adhere to the skin. The size of the sheet constituting substrate 333 is not particularly limited, and for example, in the case of a square electrode for an electrocardiogram, one side can be set to about 30 mm (for example, in the range of 5 mm to 75 mm).

[0230] 25A to 25B show a specific example of the bioelectrode 30. The upper part is a side view and the lower part is a front view. The holder 335 is omitted and not drawn. 25A to 25B, a plurality of contacts 331 are arranged in parallel in the horizontal direction of the paper, and both ends of each contact 331 are connected to a signal cable 334 arranged in the vertical direction of the paper. Both ends of the contact 331 penetrate the substrate 333 and are connected to the signal cable 334. In the illustrated example, the signal cable 334 is arranged on the surface (rear surface) of the substrate 333, which is the base material, opposite to the surface (front surface) on which the contacts 331 are arranged and fixed. With this configuration, the signal cable 334 can draw the contact 331 to the substrate surface. The surface on which the signal cable 334 is arranged may be the front surface of the substrate 333 instead of the back surface.

[0231] An opening 336 is provided in the substrate 333 at the overlapping position of the contact portion 332. The opening 336 functions as a ventilation hole (air vent). That is, when an electrode pad consisting of the contact portion 332 and the substrate 333 is pressed against the skin, steam or sweat from the skin can be discharged from the opening 336 to the outside of the electrode pad. The shape of the opening 336 is not particularly limited as long as it is a shape that allows gas to pass through the substrate 333, and may be any shape such as a circle or a rectangle. The contactor 331 arranged on the front surface of the substrate 333 may be exposed to the back surface of the substrate 333 through the opening 336.

[0232] The position where the opening 336 is arranged on the substrate 333 is not particularly limited, but it is preferable that the multiple openings 336 are arranged symmetrically with respect to the center of the substrate 333. In addition, it is preferable that the multiple openings 336 are provided at positions overlapping the contact portion 332. The total opening area of ​​the openings 336 provided on the substrate 333 is not particularly limited. It is preferable to avoid a large opening area that impairs the structural strength of the substrate 333, and usually, the opening area is preferably about 1 to 40% of the area of ​​the substrate 333. It is preferable to select from this range according to the purpose, and for example, it may be 1 to 20%, may be 20 to 40%, or may be 10 to 30%. Within the above range, the structural strength of the substrate 333 is sufficiently maintained, while improving the breathability between the substrate 333 and the skin, and reducing stuffiness of the skin. In addition, it is preferable that the total opening area of ​​the openings 36 provided at positions overlapping the contact portion 332 is about 2 to 60% of the area of ​​the contact portion 332. It is preferable to select from this range depending on the purpose, and it may be, for example, 2 to 40%, 40 to 60%, 10 to 30%, or 5 to 45%. When in the above range, the structural strength of the contact portion 332 is sufficiently maintained, while improving the breathability between the contact portion 332 and the skin, and reducing stuffiness of the skin.

[0233] 25B, a humidity control pad 337 may be provided on the surface (back surface) of substrate 333 opposite contact portion 332. Steam or sweat that has passed through opening 336 can be absorbed by humidity control pad 337. There are no particular limitations on the material of humidity control pad 337 as long as it is a water-absorbent material. Also, a humidity control cover 338 for covering or fixing the humidity control pad 337 may be provided.

[0234] Humidity conditioning pad 337 can be used not only to absorb sweat and the like from the skin, but also to supply moisture or a moisturizer such as glycerol to the skin or contacts by impregnating humidity conditioning pad 337 with moisture or a moisturizer such as glycerol in advance. In Figs. 25A to 25B, contactor 331 arranged on the front surface of substrate 333 may be exposed to the back surface of substrate 333 through opening 336, and contactor 331 may come into contact with humidity conditioning pad 337. This contact makes it possible to supply moisture and the like to contactor 331.

[0235] By providing the opening 336 in the substrate 333, it is possible to respond to various skin conditions, from situations where sweat is high and stuffy, such as in summer, for young people, and during exercise, to dry situations, such as in winter, for the elderly, and at rest. The opening 336 is not only provided for the purpose of opening the skin closed (covered) by the electrode pad and dissipating moisture, but may also be provided for the purpose of actively supplying moisture to the skin. In other words, by placing a water-absorbent pad (sponge, etc.) 337 on the opening 336, sweat can be removed and humidity can be adjusted. In a dry environment, the humidity-regulating pad 337 contains water, glycerol, and moisturizing ingredients, so that the pad 337 can supply the ingredients to the contact 331 and the skin. For the purpose of moisturizing during dry times, it is preferable to further cover the outside of the humidity-regulating pad 37 with a cover such as PVC. The conductive composite fiber that constitutes the contactor 331 has moderate moisture absorption properties, and moisture moves and diffuses due to capillary action caused by the fine fibers that constitute the contactor 331. Therefore, by providing the opening 336 and pad 337, it is possible to smoothly adjust the humidity around the contactor 332.

[0236] The size (area) of the opening 336 can be adjusted as appropriate according to the conditions of use, such as room temperature, humidity, movement, and the presence or absence of heat generation. The total area of ​​the opening or openings 336 provided in the substrate 333 is not particularly limited as long as the structural strength of the substrate 333 is adequately maintained, and can be adjusted, for example, within a range of 0.1 to 50% of the area of ​​the sheet-like substrate 333 constituting the electrode pad. It is preferable to select within this range according to the purpose, and may be, for example, 0.1 to 30%, 30 to 50%, 5 to 40%, 15 to 50%, or 0.1 to 5%, etc.

[0237] The conductive composite fiber constituting the contact 331 is similar to the conductive composite fiber of the first embodiment described above. The structure of the contact 331 is similar to the contact of the first embodiment described above.

[0238] The density of the contacts 331 in the contact portion 332 or the number of contacts 331 per unit area of ​​the contact portion 332, and the area of ​​the contact portion 332 are not particularly limited and may be appropriately adjusted according to the application. The density of the contacts 331 in the contact section 332 is, for example, when the contacts 331 (fiber bundles) having a diameter of 280 microns are arranged in parallel, usually about 30 pieces are used per 10 mm electrode width, but is not limited to this. For example, it can be adjusted in the range of 1 to 200 pieces. More specifically, for example, when a conductive composite fiber bundle (a contactor made of a composite fiber of PEDOT-PSS and silk fiber impregnated with glycerol) similar to that of Example 3-1 described below is arranged in parallel without gaps and fixed to a substrate for use as a bioelectrode for electrocardiogram measurement, the contact area with the skin (area of ​​contact portion 332) is 1 cm x 1 cm (100 mm 2 ) is usually set to 10 to 50,000 mm 2 In addition, when the bioelectrode for electrocardiogram measurement is used as a skin surface electrode for electrical stimulation, the contact area of ​​the electrode can be set to, for example, 10 to 50,000 mm 2 can be set to.

[0239] The method of arranging the contacts 331 in the contact section 332 is not particularly limited, and is appropriately adjusted according to the application. For example, in addition to arranging the multiple contacts 331 (conductive composite fiber bundles) in parallel with no gaps, a configuration in which the multiple contacts 331 are laid out in multiple layers, a configuration in which the multiple contacts 331 are woven or knitted to form a cloth-like structure, or a configuration in which the multiple contacts 331 are made of cloth and raised like a towel can be adopted. In a configuration in which the multiple contacts 331 are overlapped with each other, the contacts 331 are in contact with each other and electrically connected (conductivity is obtained), so no problems arise in use. In addition, a configuration in which the multiple contacts 331 in the contact section 332 are arranged sparsely by widening the intervals between the multiple contacts 331 may be adopted. In this configuration, the surface of the substrate 333 is exposed from the gaps between the contacts 331, and the exposed surface can directly contact the skin between the contacts 331. Therefore, by making the exposed surface of the substrate 333 adhesive, it is possible to adjust the adhesive strength of the electrode pad to the skin, the current density, and the contact area between the electrode contact portion 332 and the skin.

[0240] By impregnating the conductive composite fiber that constitutes the contact 331 with a moisturizing ingredient such as glycerol, and by having the moisture (sweat) from the skin on which the bioelectrode is placed be absorbed by the conductive composite fiber, the conductive composite fiber is maintained in a moderately moist (wet) state. When the conductive composite fiber has an appropriate amount of moisture, the conductive composite fiber becomes slightly sticky.

[0241] The method of placing the electrode pad of the bioelectrode 330 on the skin surface is not particularly limited as long as it can stably fix the bioelectrode, and for example, the electrode pad can be attached to the skin independently by utilizing the adhesiveness of the conductive composite fiber described above or the adhesiveness of the sheet-like substrate 333. When the contact part 332 on which the contacts 331 are arranged is brought into contact with the body surface (skin surface), the contacts 331 are quickly attached to the skin surface, and electrical continuity is obtained between the contacts 331 and the skin surface, so that the biosignal can be acquired. The biosignal is sent to an external device such as a bioamplifier via a signal cable 334 (metal conductor) connected to the contacts 331.

[0242] In this way, the electrode pad can be fixed (placed) on the skin surface by itself by utilizing the adhesiveness of the electrode pad. In this fixing method, the electrode pad is attached to the skin by utilizing the weakly adhesive substrate 333 and the adhesiveness caused by the wetness of the conductive composite fiber of the contactor 331, so the fixing force to the skin is not strong. Therefore, the electrode pad may be displaced or fall off due to traction by the signal cable 334 or large body movements. Therefore, a holder 335 may be applied to press the electrode pad against the skin surface S in order to stably hold the electrode pad and prevent it from being displaced or falling off.

[0243] There is no particular limitation on the shape, size, and other configuration of the holder 335. For example, as shown in Figures 25A to 25B, a strip-shaped stretchable cloth (curtain) may be used to place the electrode pad 338 on the skin surface S of the body B, and the stretchable holder 335 may be wrapped around the waist of the body B from above the electrode pad 338. With this configuration, the electrode pad 338 will not easily fall off even if the body B is moved significantly, and the electrode pad 338 can be fixed more stably.

[0244] For example, as shown in Figs. 26A to 26B, the holder 335 can be installed inside the underwear (shirt) T. The electrode pad 338 and a part of the holder 335 are fixed inside the underwear T. The holder 335 and the electrode pad 338 are structurally independent, and the holder 335 is movable laterally over the electrode pad 338, that is, movable (shiftable) in a direction along the surface of the body B. Therefore, it is preferable that the electrode pad 338 and the holder 336 are arranged so that they can be separated from each other, and it is preferable that the electrode pad 338 and the holder 335 are not completely fixed. In this way, the electrode pad 338 and the holder 335 are structurally independent, the contact portion of the electrode pad 338 and the holder 335 is not fixed, and the holder 335 is slidable over the electrode pad 338 at the contact portion, thereby preventing the electrode pad 338 from falling off due to the misalignment between the body B and the underwear T, the loss of a biological signal, and the generation of noise due to the misalignment of the electrode. Also, if necessary, the electrode pad 338 can be removed or replaced by actively separating the holder 335 from the body B. The holder 335 not only serves to stably hold the electrode pad 338, but is also used to hold accessories of the bioelectrode (such as a cable, a connector, and an amplifier 339).

[0245] The fabric (material) constituting the holder 335 is not particularly limited. For example, it is preferable to use a stretchable fabric among cloth, sheet, mesh, rubber band, etc. Specifically, a strip-shaped two-way stretchable fabric, Lycra (registered trademark) (generic name: Spandex) (manufactured by Toray Industries, Inc.), can be sewn to the inside of an underwear shirt with a width (vertical length) of 15 cm in accordance with the height of the heart (see Figs. 25A and 25B). This holder is suitable as, for example, a holder for electrodes for a Holter electrocardiogram (for CC5 lead). In the case of CC5, a configuration can be adopted in which electrode pads 338 are placed on the left and right of the anterior chest and covered by the electrode holder 335.

[0246] The holder 335 is not limited to being placed inside the underwear for the upper body as described above, but may be placed, for example, wrapped in a band around the limbs, head, neck, or fingers depending on the application of the bioelectrode 330. The material of the holder 335 is not limited to the above-mentioned spandex, and various types of cloth, sheet, mesh, band, etc. can be used as long as they are elastic flat materials (fabrics).

[0247] Effects achieved by the bioelectrodes according to the first and second embodiments of the third aspect of the present invention Examples of effects obtained by the materials and structures of the bioelectrode of each embodiment are listed below.

[0248] When placing the bioelectrode of the present invention at the measurement site, there is no need to use a conductive gel (electrolyte gel) or a conductive paste (electrolyte paste). By not using a conductive gel or paste, the following effects (A) to (E) can be obtained. (A) Improved wearing comfort. There is little discomfort associated with attaching the electrodes to the skin. Because the gel or paste is not used, there is no need to seal the skin with a liquid or gel, and the electrodes can be placed in an open state to the outside air. In other words, measurements can be performed with a string-like electrode lightly touching the skin, or with a soft cloth-like electrode in contact with the skin. (B) Problems caused by electrolyte paste are avoided. There is no risk of leakage of the electrolyte solution, or of poor contact or noise occurring when the water in the gel or paste dries up. (C) The electrical properties of the electrode are improved. It is possible to reduce the resistance of the electrode per unit area compared to conventional bioelectrodes. This is advantageous for measuring weak signals such as electroencephalograms and evoked potentials. In addition, since the capacitance of the electrode according to the present invention is small, it has excellent high-frequency propagation characteristics and is advantageous for recording biosignals containing high-frequency components such as electroencephalograms and electromyograms. (D) The electrodes are highly convenient to use. Since the paste or gel is not used, there is no need to remove the paste or gel after measurement (after the test). For example, washing hair after EEG measurement, which was necessary when using conventional electrodes, can be omitted. (E) The electrodes can be made smaller and lighter. Since the electrode resistance per unit area is smaller than that of conventional electrodes, the electrode can be made smaller, lighter, and denser than conventional bioelectrodes.

[0249] By providing the bioelectrode according to the present invention with the conductive composite fiber, the following effects (F) to (K) can be obtained. (F) Improved installation stability. The electrodes can be stably placed using light pressure or weak adhesive materials. Unlike conventional bioelectrodes, there is no need for strong adhesives, bands, headgear, or other strong compression fixation. (G) A low-noise signal is obtained. Due to the properties of the conductive composite fiber, such as adhesiveness, flexibility, thinness and lightness, when the person wearing the electrode (subject) moves, unnecessary vibration of the electrode is reduced and noise is reduced. (H) A natural appearance is achieved. In particular, in the case of EEG electrodes, the electrodes are small and flat, and the design allows them to be hidden under the hair, making them unnoticeable even when worn. In other words, EEG measurements can be made at any time during daily life. (I) It can reduce sweating on the skin caused by wearing bioelectrodes for long periods of time. Generally, when an electrode is worn continuously for a long time, the skin tends to become sweaty due to sweating. However, when the bioelectrode according to the present invention uses a hydrophilic conductive composite fiber as the electrode material as described above and has a ventilation opening in the substrate, the skin can be further reduced from becoming sweaty during long-term use. (J) The range of applications of bioelectrodes can be expanded. The overall shape (basic shape) of the bioelectrode can be processed into a thin, flat shape (cloth-like) or linear shape (string-like). Because it is lighter, flatter, and more flexible than conventional electrodes, it is also possible to fabricate linear electrodes that are thinner than strings. In addition, the wearing sensation is comfortable. Due to these properties, the bioelectrode of the present invention can be used as a wearable electrode, expanding the range of its application. (K) It is possible to perform measurements that are as stable or more stable than conventional bioelectrodes. The bioelectrode of the present invention overcomes the drawbacks of conventional pasteless electrodes, such as noise contamination and instability of the measurement signal, even when it does not use electrolyte paste (pasteless). In other words, it provides a stability of the measurement signal that is equal to or greater than that of conventional medical bioelectrodes that use electrolyte paste.

[0250] Hereinafter, a conductive polymer fiber that can be used as the conductive composite fiber constituting the bioelectrode of the present invention will be described in detail, although the conductive composite fiber is not limited to this conductive polymer fiber.

[0251] Regarding the fourth aspect A fourth aspect of the present invention relates to an implantable electrode and a biosignal measuring device. More specifically, the present invention relates to an implantable bioelectrode using a composite material of a conductive polymer and a fiber (hereinafter, conductive composite fiber), and a biosignal measuring device including the bioelectrode. Hereinafter, an embodiment of the fourth aspect of the present invention will be described with reference to the drawings, but the present invention is not limited to such an embodiment.

[0252] As described above, PEDOT-PSS, a conventional conductive polymer, gels in biological tissues due to its high water absorption, and its mechanical strength is extremely reduced. For this reason, it is difficult to place PEDOT-PSS processed into a needle or rod shape alone in the body, as in conventional metal or carbon bioelectrodes. Even if PEDOT-PSS can be embedded in biological tissues, there is a problem that the connection part between the electrode made of PEDOT-PSS alone and the metal conductor (cable) connecting to an external device becomes weak due to PEDOT-PSS absorbing water, and is prone to breakage (disconnection).

[0253] In a fourth aspect of the present invention, a conductive composite fiber in which a conductive polymer is composited with fiber is used as an electrode to be implanted in the body, thereby solving the problems of the electrode itself disintegrating due to the conductive polymer absorbing water and the weakening of the connection between the conductive polymer and the metal conductor.

[0254] First embodiment of the fourth aspect The implantable electrode 410 of the first embodiment of the fourth aspect of the present invention shown in Figs. 29A to 29C includes a conductive composite fiber bundle 401 formed into a rod (needle) shape by bundling a plurality of conductive composite fibers containing a conductive polymer. A metal conductor 402 is wound around a part of the conductive composite fiber bundle 401 to form a connection part 403. The connection part 403 is covered with an insulating and water-resistant polymer 404 (resin). The conductive composite fiber bundle 401 has a higher mechanical strength than a conductor formed into a rod shape having an equivalent diameter by a single conductive polymer both before and after water absorption. This makes it possible to prevent the conductive composite fiber bundle 401 from being damaged when embedded in biological tissue, and from being disintegrated in biological tissue after embedding.

[0255] The conductive composite fiber bundle 401 is preferably in a dry state before use. The conductive composite fiber bundle 401 in the dry state has high mechanical strength and has a relatively small volume because it has shrunk compared to when wet. Therefore, by using the conductive composite fiber bundle 1 in a dry and shrunk state, it is possible to reduce the invasiveness when inserting it into biological tissue.

[0256] (Conductive composite fiber) As the conductive composite fiber constituting the conductive composite fiber bundle 401, a composite fiber of a conductive polymer and a conventionally known fiber material can be applied. The form (method) of the composite is not particularly limited, and for example, the conductive polymer may be coated on the surface of the thread-like (string-like) fiber material, the conductive polymer may be impregnated into the thread-like fiber material, or the thread-like conductive polymer and the thread-like fiber material may be twisted or spun together. The fibers described in the first aspect of the present invention can be preferably used.

[0257] The type of the conductive polymer is not particularly limited, and known conductive polymers can be used. Examples of such conductive polymers include hydrophilic conductive polymers such as PEDOT-S (poly(4-(2,3-dihydrothieno[3,4-b][1,4]dioxin-2yl-methoxy-1-butanesulfonic acid, potassium salt) in addition to the above-mentioned PEDOT-PSS. By using a composite fiber containing a hydrophilic conductive polymer as the material for the conductive composite fiber bundle 1, the conductive composite fiber bundle 1 itself can be easily imparted with adhesiveness (tackiness) to the needles 5.

[0258] As the fiber material, for example, conventionally known fiber materials such as silk, cotton, hemp, rayon, and chemical fibers can be used. Among these, silk is preferable. When silk is used, the strength and hydrophilicity of the composite fiber can be further improved. In addition, silk is preferable because it has almost no toxicity to biological tissue, is unlikely to cause inflammation due to an immune reaction, and has excellent compatibility with biological tissue. The type of the conductive polymer to be combined with silk is not particularly limited, but the above-mentioned hydrophilic conductive polymer such as PEDOT-PSS or PEDOT-S is preferred.

[0259] The length and thickness of the conductive composite fiber constituting the implantable electrode of the present invention are not particularly limited, and are appropriately adjusted according to the length and thickness of the fiber material to be composited. In addition, the length and thickness of the conductive composite fiber bundle formed by twisting or binding a plurality of conductive composite fibers are not particularly limited, and can be appropriately adjusted according to the purpose and use. For example, the thickness may be in the range of 0.01 μm to 5 mm, and the length may be in the range of 0.1 μm to 1 m. In another example, the thickness may be in the range of 0.1 μm to 1 mm, and the length may be in the range of 0.1 μm to 50 cm. As a specific example, the rod-shaped conductive composite fiber bundle 401 shown in Figs. 29A to 29C may have a thickness of 0.1 μm to 500 μm and a length of 1 μm to 10 mm. The coil-shaped conductive composite fiber bundle 401 shown in Figs. 31A to 31C may have a thickness of 10 μm to 500 μm and a length of 100 μm to 50 cm. Here, the thickness and length are the thickness and total length of the conductive composite fiber bundle 1 wound in a coil shape in a stretched state. The outer diameter of the coil wound in a coil shape shown in Figs. 301A to 301C may be, for example, 10 μm to 5 mm, and the length of the coil in the central axis direction may be, for example, 100 μm to 50 mm. The conductive composite fiber bundle 401 connected to a surgical thread shown in Figs. 32A to 32D may have a thickness of 0.1 μm to 500 μm and a length of 1 μm to 10 cm. Moreover, the conductive composite fiber bundle 401 constituting the core shown in FIG. 33B can have a thickness of, for example, 10 μm to 10 mm and a length of, for example, 10 μm to 50 cm.

[0260] As the conductive composite fiber constituting the implantable electrode of each embodiment of the present invention, a conductive polymer fiber, which will be described in more detail later, can be used.

[0261] (Adhesion between needle and conductive composite fiber bundle) The conductive composite fiber bundle 401 of the first embodiment of the fourth aspect is adhered to the tip of a needle 405 (guide needle). When the conductive composite fiber bundle 401 is moistened with water or alcohol, the conductive polymer on its surface becomes adhesive (tacky), and when it is dried again, it shrinks and hardens. By utilizing this property, the conductive composite fiber bundle 401 can be adhered (fixed) to the tip of a needle 405 (FIG. 29A). When an implantable electrode 410 having this configuration is inserted into biological tissue, the conductive composite fiber bundle 401 absorbs bodily fluids (extracellular fluid, cerebrospinal fluid, etc.) and swells (FIG. 29B). Furthermore, the adhesive strength (fixing strength) between the swollen conductive composite fiber bundle 401 and the needle 405 decreases, so that the needle 405 can be removed while leaving the conductive composite fiber bundle 1 in the biological tissue (FIG. 29C). The conductive composite fiber bundle 401 placed in the biological tissue is connected to an external device via an electric wire 402 (metal conductor 402), and signals (electrical signals or electrical stimuli) are sent and received.

[0262] The material of the needle 405 is not particularly limited, and examples thereof include metals such as gold, platinum, and copper, carbon, and resin (plastic).

[0263] The electric wire 402 is preferably a wire capable of electrically conducting the conductive composite fiber bundle and an external device. The constituent material of the electric wire 402 is not particularly limited, and for example, metal, silicon, carbon, etc. can be applied. The type of the metal is not particularly limited, and may be a metal used in a conventionally known electric wire. In order to prevent the electric wire 402 embedded in the biological tissue from picking up electrical noise and to function stably for a long period of time, the electric wire 402 is preferably coated with a polymer having insulating properties and water resistance. The type of the polymer is not particularly limited, and for example, a water-resistant polymer that coats the conductive composite fiber bundle of the fourth embodiment of the present invention described later can be applied. The thickness and length of the electric wire 402 are not particularly limited, and can be appropriately adjusted depending on the application.

[0264] The conductive composite fiber bundle 401 may be bonded to the needle 5 via a hydrophilic adhesive material (adhesive), in addition to the above-mentioned method utilizing the adhesiveness of the conductive polymer when wet. The adhesive material is not particularly limited, and is preferably a material that can bond the conductive composite fiber bundle 1 and the needle 5 together (demonstrates adhesiveness) in a dry state, and whose adhesive strength (adhesive strength) decreases when it absorbs water. Examples of the adhesive material include materials containing PEG (polyethylene glycol), PEDOT-PSS, polylactic acid, sorbitol, fibrin glue, starch glue, etc.

[0265] The type of PEG is not particularly limited, and for example, a relatively high molecular weight PEG that is solid at room temperature (e.g., about 20°C) to body temperature (e.g., about 40°C) and becomes liquid when heated can be used. PEG that has been heated and dissolved is applied to a needle, and then a conductive composite fiber bundle is brought into contact with the needle, and then the temperature is returned to room temperature, whereby the PEG solidifies and the needle and fiber bundle can be bonded together. When this is placed in an environment containing body fluids, such as in tissue, the PEG gradually dissolves, allowing the conductive composite fiber bundle to naturally separate from the needle.

[0266] In addition, as a method of adhering the conductive composite fiber bundle 401 to the needle 405, the conductive composite fiber bundle 1 and the needle 5 may be indirectly adhered via the adhesive material applied to the polymer 404 covering the connection portion 403.

[0267] (Protection of wiring parts) The connection part 403 is covered with a polymer 404. In biological tissue, the conductive composite fiber bundle 401 constituting the connection part 403 is covered with the polymer 404, so there is almost no swelling or loss of mechanical strength due to water absorption. In addition, since the mechanical strength of the conductive composite fiber bundle 401 is increased by combining it with a fiber material, the connection between the conductive composite fiber bundle 401 and the metal conductor 403 does not break (disconnect) even after water absorption, and the electrical connection can be sufficiently maintained.

[0268] The method of connecting the metal conductor 402 to the conductive composite fiber bundle 401 at the connection portion 403 is not particularly limited, and examples of the method include wrapping, tying, crimping, and adhesion with a conductive adhesive (silver paste, silver epoxy, etc.). The type of polymer 404 that covers the connection portion 403 is not particularly limited, and examples of the polymer 404 include silicone, PTFE (polytetrafluoroethylene), and PVC (polyvinyl chloride). By covering the connection portion 403 with the polymer 404, it is possible to prevent an electrical short circuit and protect the connection portion 3.

[0269] (Placement of electrodes inside the body) A method of placing the implantable electrode 410 of the first embodiment of the fourth aspect in a living tissue includes, for example, a method of inserting the needle 405 into the living body at high speed (in a short time) using a manipulator capable of high speed operation. The needle 405 leads the way to a predetermined position in the living body, and the conductive composite fiber bundle 401 attached to the needle 405 and the connected metal conductor 402 are both introduced to the predetermined position in the living body. This insertion is preferably completed at high speed and completed before the conductive composite fiber bundle 401 starts to swell in the living body. The insertion speed is not particularly limited, but can be, for example, about 100 to 1000 mm / sec. As a specific example, a conductive actuator capable of high speed operation can be used to insert the conductive composite fiber bundle 401 of the implantable electrode 410 at a speed of 10 to 20 msec so that the conductive composite fiber bundle 401 is placed at a depth of 2 mm below the cerebral cortex of the animal. Thereafter, the conductive composite fiber bundle 401 is caused to swell with body fluids, dissolving the adhesive material bonding the conductive composite fiber bundle 401 and the needle 405, and when the adhesive strength weakens, it is possible to remove only the needle 405. The conductive composite fiber bundle 401 placed in the biological tissue swells and comes into close contact with the surrounding biological tissue.

[0270] (Multiple electrode placement) The implantable electrode 10 of the first embodiment may have one conductive composite fiber bundle 401 at the tip of the needle 405 as shown in FIG. 30A, or may have multiple conductive composite fiber bundles 401 at the needle 405 as shown in FIGS. 30B and 30C.

[0271] In the configuration of FIG. 30B, the heights of the adhesive points of the two conductive composite fiber bundles 401 on the needle 405 are shifted (the positions are shifted in the length direction of the needle 405). With this configuration, when inserted into the body, each conductive composite fiber bundle 401 is placed at a different height (depth), and each can function as an independent electrode (2-channel electrode). In the configuration example of FIG. 30B, when viewed from the side and bottom, the conductive composite fiber bundles 401 are fixed on both sides of the needle 405, but two electrodes may be fixed to one side of the needle 405. In this case, when viewed from the bottom, the two conductive composite fiber bundles 401 appear to overlap in the depth direction (height direction). By fixing in this manner, it is possible to further reduce the invasiveness of the catheter to living tissue when the catheter is inserted into the body.

[0272] 30C, the height of the bonding points of the four conductive composite fiber bundles 401 on the needle 405 is set to be the same, and the four conductive composite fiber bundles 401 are arranged to surround the periphery of the needle 405. In this case, the four conductive composite fibers 401 (4-channel electrodes) can be placed in the biological tissue with the position where the needle 405 is inserted as the center.

[0273] Second embodiment of the fourth aspect An implantable electrode 420 according to a second embodiment of the fourth aspect of the present invention shown in Figures 31A to 31C is similar to the first embodiment, except that a conductive composite fiber bundle 401 is wound in a coil shape around the tip of a needle 405. In Figures 31A to 31C, the same components as those in the first embodiment of the fourth aspect are denoted by the same reference numerals.

[0274] The coil-shaped conductive composite fiber bundle 401 may be adhered to the tip of the needle 405, or may simply be wound around it. Since the coil-shaped conductive composite fiber bundle 401 is tightly wound around the tip of the needle 405, the conductive composite fiber bundle 401 is prevented from falling off the needle 405 when the tip of the needle 405 is inserted into the biological tissue (FIG. 31A). In addition, the outer diameter of the coil when dried and shrunk (the diameter of the circle that the coil describes) is small, so that the invasiveness to the biological tissue is reduced. The implantable electrode 420 inserted into the biological tissue swells as the conductive composite fiber bundle 1 absorbs bodily fluids, becoming a swollen fiber bundle 401' (FIG. 31B), that is, the coil spontaneously expands to increase its outer diameter, and the conductive composite fiber bundle 401 and the biological tissue come into close contact with each other. Because the adhesive strength between the conductive composite fiber bundle 401 and the needle 405 weakens due to water absorption, the needle 405 can be removed while leaving the conductive composite fiber bundle 401 in the biological tissue (FIG. 31C).

[0275] The second embodiment of the implantable electrode 420 having a coil-shaped conductive composite fiber bundle 401 is suitable for cases where the installation of the electrode (conductive composite fiber bundle 401) may cause atrophy or dead space to form in biological tissue, or where the cells or nerve fibers to be measured are scattered throughout the biological tissue.

[0276] Third embodiment of the fourth aspect 32A to 32D show an implantable electrode 430 according to a third embodiment of the fourth aspect of the present invention. A metal conductor 402 is connected to one end of a conductive composite fiber bundle 401, and a nylon monofilament thread 406 for surgical use is attached to the other end by the method described above. A curved needle 405 for surgical suture is tied to the nylon thread 406.

[0277] A method of placing the implantable electrode 430 in a nerve cord (bundle) is illustrated. A needle 405 is pierced (pierced) through the nerve cord N' in a manner similar to suturing the nerve cord N by a microsurgery surgical technique used for neurovascular suturing (FIG. 32A), and the nylon thread 406 is pulled up, so that the conductive composite fiber bundle 401 is introduced into the nerve cord N' by being pulled by the nylon thread 406 (FIG. 32B). Thereafter, the conductive composite fiber bundle 401 absorbs bodily fluids at a predetermined position in the nerve cord N, so that the nylon thread 406 can be peeled off from the conductive composite fiber bundle 401 and removed from the nerve cord N'. The conductive composite fiber bundle 401 placed in the nerve cord N' expands by absorbing water and adheres closely to the inside of the nerve cord N' (FIG. 32C). FIG. 32D shows the placement of multiple implantable electrodes on multiple nerve cords N within a nerve bundle.

[0278] Many peripheral nerves are mixed nerves of motor, sensory, and autonomic nerves, forming nerve cords (bundles). The nerve cords run three-dimensionally through the nerve bundles, and the distribution of the nerve bundle cords varies greatly from person to person, so it is difficult to identify the nerve cords by brain coordinates like the central nerves. However, clinically, the main nerve fibers can be identified by observation under a microscope and measurement of nerve activity. The implantable electrode of the third embodiment makes it possible to selectively record and stimulate signals of motor, sensory, and autonomic nerves by utilizing this clinical technique.

[0279] The implantable electrode of the third embodiment can be inserted not only by surgical technique, but also mechanically by an automatic anastomosis instrument, a micromanipulator, or the like.

[0280] (Adjustment of water absorption speed) It is possible to delay the rate at which the conductive composite fiber and the conductive composite fiber bundle constituting the implantable electrode of the present invention absorb body fluids in the body. The method for delaying the rate is to impregnate or apply at least one of glycerol, sorbitol, ethylene glycol, squalane, silicone, mineral oil, and MPC (2-methacryloyloxyethyl phosphorylcholine) to the conductive composite fiber (bundle) in advance. For example, by impregnating the conductive composite fiber bundle 401 of the third embodiment with glycerol in advance, even if the introduction of the electrode into the living body is difficult and the operation time is prolonged, it is possible to prevent the conductive composite fiber bundle 401 from swelling due to water absorption and the nylon thread 406 from peeling off during the operation. By maintaining the diameter of the conductive composite fiber bundle 401 small, it is possible to reduce the invasion of the living tissue during the introduction of the electrode.

[0281] Fourth embodiment of the fourth aspect 33A and 33B, an implantable electrode 440 according to a fourth embodiment of the present invention has a core made of a conductive composite fiber (bundle) 401 formed into a rod (needle) or string (cable) shape, at least a part of the core is covered with a water-resistant polymer 404, and a flow path is formed for liquid to permeate (transmit) from one end 1a (401a) of the core to the other end 1b. The one end 1a and the other end 1b are not covered with the polymer 404 and are exposed.

[0282] The "flow path" does not mean a hollow tube, but a configuration in which the water-resistant polymer 404 forms a tube and the conductive composite fiber bundle 1 is arranged in the tube. The conductive composite fiber bundle 401 has water absorption and substance permeability, so that liquid can permeate from one end 1a to the other end 1b and move spontaneously. The method of transporting liquid or substance through the flow path is not limited to spontaneous movement such as permeation, capillary action, and diffusion, but may also include a method of electrophoresing a substance by using one of the one end 1a and the other end 1b as a positive electrode and the other as a negative electrode, or a method of connecting a pump (e.g., an osmotic pump) to the one end 1a to pressurize the liquid and send the liquid. Either method allows stable drug transport and liquid delivery at a constant speed.

[0283] A reservoir 407 (FIG. 33A) or a chamber 408 (FIG. 33B) capable of containing a drug solution is connected to one end 1a of the core formed by the conductive composite fiber bundle 401. A pump for feeding the liquid may be connected to a tube connector 409 provided in the chamber 408. By storing a solution containing a drug in reservoir 407 or chamber 408, the solution can permeate the flow path and flow from one end 1a of the core to the other end 1b. Therefore, by placing the other end 1b at a desired position in the living tissue, the drug can be administered locally around the other end 1b.

[0284] The type of the drug is not particularly limited, but it is preferably a drug having a pharmacological action of suppressing or promoting a biological reaction. Examples of the drug include drugs that reduce damage to biological tissues, drugs that promote the repair of biological tissues, and drugs that grow biological tissues. Specific examples of the drug include soluble drugs such as glycerol, sorbitol, mannitol, fructose, BDNF (Brain-derived neurotrophic factor), NGF (Nerve Growth Factor), NT3 (Neurotrophin-3), GSNO (S-Nitrosoglutathione), SKF96365, Cilostazol, TRIM (1-(2-Trifluoromethylphenyl)imidazole), Gadolinium, Magnesium, EGTA (ethylene glycol tetraacetic acid), and Ruthenium Red. A liquid in which one or more of these drugs are dissolved is stored in the reservoir 407 or the chamber 408.

[0285] The type of water-resistant polymer 404 that coats the core is not particularly limited as long as it is a polymer that can form a coating layer around the core (form a water seal on the outer surface of the core), and for example, a polymer (resin) that is used in the field of medical devices such as conventionally known catheters can be applied. It is preferable that the water-resistant polymer 404 also has insulating properties to prevent the conductive composite fiber 401 that constitutes the core from being electrically short-circuited with the surroundings. The thickness of the coating layer formed by the water-resistant polymer 404 is not particularly limited. For example, it can be 0.1 μm to 5 mm.

[0286] Specific examples of the water-resistant polymer 404 include silicone, PTFE (polytetrafluoroethylene), PVC (polyvinyl chloride), ABS (acrylonitrile butadiene styrene), ANS (acrylonitrile styrene), PEN (polyethylene napthalate), PBT (polybutylene terephthalate), polycarbonate, PEI (polyetherimide), PES (polyether sulfone), PET (polyethylene terephthalate), polyamide, aromatic polyamide, polyester, polyether block amide copolymer, polymethylmethacrylate, polyurethane, and EVA (ethylene vinyl acetate). acetate, ethylene vinyl alcohol, polyethylene, latex rubber, PTFE, FEP, PFA, polypropylene, polysiloxane, ionomer, SAN (styrene acrylonitrile), nylon, and thermoplastic elastomers.

[0287] The drug may be pre-impregnated or pre-applied to the conductive composite fiber constituting the implantable electrode of the present invention. In this case, too, the drug is gradually released from the conductive composite fiber placed in the biological tissue, and the drug can be administered locally around the conductive composite fiber.

[0288] In a state where the other end 1b of the core is embedded in the biological tissue, the one end 1a and the reservoir 407 or the chamber 408 may be embedded in the biological tissue or may be placed outside the biological tissue. When the one end 1a is embedded in the biological tissue, it is desirable that the reservoir 407 or the chamber 408 connected to the one end 1a has a small volume as much as possible. For example, a capsule-shaped reservoir 407 may be mentioned. In addition, when the chamber 408 is placed in the biological tissue, it may be connected to the outside of the biological tissue via a tube connector 409 provided in the chamber 408. From the viewpoint of reducing invasiveness to the biological tissue, it is preferable that the one end 1a of the core and the reservoir 7 or the chamber 408 are placed outside the biological tissue. The core covered with the water-resistant polymer 4 can be produced in a desired length (for example, 100 μm to 10 cm) and thickness (for example, 10 μm to 5 mm) depending on the application.

[0289] The size and material of the reservoir 407 and the chamber 408 are not particularly limited and can be appropriately changed depending on the purpose and mode of use. For example, a bag or a box (case) made of a plastic such as silicone resin can be used as the reservoir 407 or the chamber 408.

[0290] In the fourth embodiment, it is not necessary for the entire section from one end 1a to the other end 1b of the conductive composite fiber bundle 401 to be covered with the water-resistant polymer 404. Some sections may not be covered. It is preferable that the section in which the conductive composite fiber bundle 401 functions as a liquid transport path is covered with the water-resistant polymer 404. In the fourth embodiment shown in Figures 33A and 33B, the section from the vicinity of the other end 1b to the reservoir 407 or chamber 408 in which the one end 1a is provided is covered with the water-resistant polymer 404.

[0291] There is no particular limitation on the method of connecting the reservoir 407 and the chamber 408 to the one end 1a. An example of a connection method is to place the reservoir 407 or the chamber 408 in a liquid storage section with the one end 1a exposed, and adhere the water-resistant polymer 404 covering the central side of the one end 1a to the outer wall of the reservoir 407 or the chamber 408 with a known adhesive or the like.

[0292] (Application of drug delivery function) Conventionally, when electrodes are implanted in the nervous system (central nervous system tissue), there is a problem that limited damage caused by the invasiveness of the implantation process can expand and cause permanent damage over an area larger than the size of the electrode, and a solution to this problem is needed.

[0293] The drug transport function (drug delivery function) of the implantable electrode 440 of the fourth embodiment of the present invention is applicable to administration of drugs to reduce damage caused by implantation of the electrode, and is particularly effective in reducing damage caused by implantation in nerve tissue. By administering a drug having the effect of reducing damage to the central nervous tissue, such as GSNO (S-Nitrosoglutathione), from the other end 1b of the core, damage caused by the implantable electrode 440 to the central nervous tissue can be significantly reduced. As a result, transmission and reception of signals between the electrode and the nerve tissue can be performed stably and with high accuracy for a longer period of time than in the past. This will be described in detail with reference to data in Example 4-4 (FIGS. 34A to 34D) described later.

[0294] When using a conventional metal or carbon bioelectrode in combination with a drug such as GSNO, it is necessary to separately install a tube for drug transport (for example, a very fine hollow needle such as a microcapillary) at the installation site of the bioelectrode. Alternatively, it is necessary to form a bundle structure (bundle structure) in which the very fine hollow needle is placed along the bioelectrode and stored (tied) in a single sheath (tube). In such a bundle structure, the drug release hole (release port) of the drug transport path and the electrode are installed separately (built into the sheath), which not only makes the configuration of the structure embedded in the biological tissue complicated, but also makes it difficult to distribute the drug uniformly on the interface between the electrode and the biological tissue.

[0295] On the other hand, the fourth embodiment of the present invention has a simple structure because the conductive composite fiber bundle 401 itself, which is the electrode, also serves as a drug transport path. Furthermore, because the drug seeps out from the surface of the electrode, it is possible to administer the drug uniformly to the interface between the electrode and the cell tissue, i.e., the area where damage is most likely to occur. In addition, the release rate of the drug can be adjusted by adding the aforementioned additives such as glycerol to the conductive composite fiber bundle 401.

[0296] The drug transport function (drug delivery function) of the implantable electrode 440 of the fourth embodiment of the fourth aspect of the present invention is not limited to the use of drug administration to reduce damage to biological tissues, and can be used for various uses to stimulate or utilize the physiological functions of cells and biological tissues of a living body, such as selective binding of nerve fibers by neurotrophic factors (selective formation of nerve wiring), recording of electrical signals accompanying selective stimulation of nerve fibers, etc. In addition, the liquid flowing through the flow path of the implantable electrode 440 is not limited to a drug solution, and the composition and function of the liquid are not limited as long as it is a liquid that can permeate and move through the conductive composite fiber bundle 401.

[0297] Example of Effects Obtained by the Fourth Aspect of the Present Invention The implantable electrode of the present invention provides the following advantages, for example. 1. Bioelectrodes made of flexible conductive composite fibers can be placed inside biological tissue. 2. The connection between the conductive composite fiber and the electric wire (signal cable) can be maintained stably inside the body. 3. Drugs can be transported at a constant speed to the area where the electrode and biological tissue come into contact. 4. Damage to biological tissues (especially brain and nerve tissues) caused by the placement (implantation) of electrodes can be reduced. 5. It becomes possible to stably record biological signals for a long period of time. 6. Electrodes can be placed three-dimensionally to match the three-dimensional structure of neural tissue. EXAMPLES

[0298] Example of the first aspect The first aspect of the present invention will now be described in more detail with reference to examples, but the first aspect of the present invention is not limited to the following examples. <Evaluation of tensile strength> [Comparative Example 1-1] A PEDOT-PSS film (cross-sectional area 0.03 mm2) was prepared by drying and concentrating Heraeus CLEVIOS P solution (Heraeus) and applying it uniformly onto a flat plate, allowing it to dry naturally, and then fixing it with ethanol. 2 The tensile strength of a sample (3 cm long, 1.5 mm thick) in a dry state and in a wet state (absorbed pure water until saturated) was investigated. The results are shown in FIG. 9A. It is clear from the graph that the tensile strength of the PEDOT-PSS linear bodies in the wet state (right) is drastically reduced to about 10% of the tensile strength in the dry state (left).

[0299] [Example 1-1] The tensile strength of the raw silk thread (No. 9 silk thread; Fujix Co., Ltd., twisted thread of 18 strands of 21D denier silk fiber, thread diameter approximately 280 μm, length 20 cm) in a dry and wet state (absorbed pure water until saturated) is investigated. The results are shown in Figure 9B. In addition, after immersing in 20cc of the CLEVIOS P solution for 1 hour, electrochemical fixation was performed by passing a current of 3mC per 1cm using a comb-shaped multi-point electrode, and ethanol was used as an organic solvent to dry out the water in the solution by blowing dry air to remove 60% of the water. The results of investigating the tensile strength of a conductive polymer fiber bundle (hereinafter sometimes referred to as PEDOT-PSS silk fiber bundle 1) with a diameter of about 280μm, which is made of PEDOT-PSS and the silk thread (silk thread No. 9) obtained by the above-mentioned preparation method 2b, in a dry state and a wet state (a state in which pure water was absorbed until it was saturated) are shown in Figure 9B. The vertical axis of the graph indicates the maximum tension (CN: centinewton), and the error bars indicate the standard deviation of 10 samples. The tensile strength test was performed in accordance with the JIS L 1013 standard using a constant-speed extension tester (Orientec Co., Ltd., model RTC-1210A). The fiber gripping distance was 20 cm, the tensile speed was 20 cm / min, and the maximum strength was calculated from the average of the measured values ​​obtained 10 times.

[0300] The specific measured values ​​(CN) in the graph were as follows: Raw silk dry; mean value = 1350.4, standard deviation = 8.11 Raw silk wet; mean value = 1082.9, standard deviation = 12.28 · PEDOT-PSS silk dry; mean value = 1238.8, standard deviation = 16.93 · PEDOT-PSS silk wet; mean value = 1031.4, standard deviation = 24.45

[0301] From the graph, when comparing the raw silk thread in the dry and wet states with the PEDOT-PSS silk fiber bundle 1 in the dry and wet states, no clear difference in strength was observed between the two. That is, the tensile strength of the PEDOT-PSS silk fiber bundle 1 in the wet state was 83% of the strength when dry, and the tensile strength of the raw silk thread in the wet state was 80% of the strength when dry. From this, it is clear that the conductive polymer fiber according to the present invention has excellent strength equivalent to the raw silk thread in both the dry and wet states, and is less likely to break or crack, so its conductivity is less likely to decrease. Furthermore, regarding the strength difference between the dry and wet states, the PEDOT-PSS silk fiber bundle 1 (which was reduced by 17% (207 CN) based on the strength when dry) has a smaller strength difference than the raw silk thread (which was reduced by 20% (268 CN) based on the strength when dry), so it can be seen that the PEDOT-PSS silk fiber bundle 1 has a smaller change in strength due to wetting and has stable strength characteristics. Furthermore, from the results of Example 1-1 and Comparative Example 1-1, it is clear that the tensile strength of the conductive polymer fiber (in a dry state) of the present invention is approximately 10 times higher than that of a conductive fiber (in a dry state) consisting only of PEDOT-PSS.

[0302] <Water resistance evaluation> [Example 1-2] Regarding the conductive polymer fiber (PEDOT-PSS silk fiber bundle 1) produced in the same manner as in Example 1, sample A impregnated with glycerol and sample B not impregnated with glycerol were prepared. Each sample A and B was immersed in pure water and shaken 10 times at 3 Hz with a horizontal amplitude of 5 cm, and then naturally dried. This washing process was repeated three times, and the change in resistance of each sample A and B was recorded. The resistance was calculated from the amount of current at a DC 5V load using a DC stabilized power supply (PAB18-5.5; Kikusui Electronics Co., Ltd.) and a digital multimeter (VOAC7511; Iwasaki Electric Co., Ltd.). The resistance was measured for samples in a dry state (containing no moisture). The results are shown in Figure 10. The vertical axis of the graph shows the resistance (MΩ / mm) per mm of length for PEDOT-PSS silk fiber bundle 1 (dry state) with a fiber diameter of approximately 280 microns.

[0303] From the graph, in the case of the PEDOT-PSS silk fiber bundle 1 without added glycerol (sample B; the plots of "◇" connected by a solid line in the graph), the resistance value increased and the conductivity decreased with repeated washing treatment. On the other hand, in the PEDOT-PSS silk fiber bundle 1 with added glycerol (sample A; the plots of "△" connected by a two-dot chain line in the graph), no increase in the resistance value was observed and it is clear that the conductivity was maintained. In other words, the water resistance can be improved by impregnating the conductive polymer fiber according to the present invention with an additive such as glycerol.

[0304] <Evaluation of Bioelectrodes (1)> [Examples 1-3] Using a conductive polymer fiber (hereinafter sometimes referred to as PEDOT-PSS silk fiber bundle 2) with a diameter of about 280 μm and a length of 300 mm, consisting of PEDOT-PSS obtained by the above-mentioned preparation method 2b and the above-mentioned silk thread (silk thread No. 9), a string-shaped body surface type bioelectrode was fabricated by providing a rubber band 4 and a metal conductor 5 on a fixing string 3, and further winding the PEDOT-PSS silk fiber bundle 2 in a coil shape as an electrode, as shown in FIG. 11A. This bioelectrode was placed on the surface of a human body 6. An example of the results of measuring a human electrocardiogram using this bioelectrode is shown in FIG. 11B. When measuring a human electrocardiogram, the PEDOT-PSS silk fiber bundle 2, which is an electrode constituting the bioelectrode, could be brought into contact with the skin and the measurement could be performed without using a paste (jelly) containing electrolytes, etc. In other words, it is clear that the bioelectrode including the PEDOT-PSS silk fiber bundle 2 according to the present invention has excellent strength, flexibility and conductivity and can be attached in close contact with the body surface. The bioelectrodes were placed on the skin (body surface) of the right upper limb, left upper limb, and left lower limb. Each bioelectrode was connected to an electrocardiograph (polygraph, AP1124; manufactured by TEAC) to record the human electrocardiogram at rest using the bipolar limb lead method (set sensitivity 2000 μV / mm, time scale 1 second, leads I, II, III).

[0305] <Evaluation of Bioelectrodes (2)> [Examples 1-4] The conductive polymer fiber, which is made of PEDOT-PSS and the silk thread (silk thread No. 9) obtained by the above-mentioned preparation method 1a and has a diameter of about 280 μm and a length of 1.5 mm, was covered with silicone resin to insulate a part of it. Specifically, the length of the exposed part (non-insulated part) was about 500 μm, and the length of the insulating coating part was about 1000 μm. The electrode resistance of the obtained conductive polymer fiber (hereinafter sometimes referred to as PEDOT-PSS silk fiber bundle 3) was about 500 kΩ, and this was connected to a metal thin wire (Xwire, manufactured by Tanaka Kikinzoku Kogyo Co., Ltd.) to form a thread-shaped embedded bioelectrode. A photograph of it observed with a stereomicroscope is shown on the left in Figure 12A.

[0306] Next, the fabricated bioelectrode was inserted under a microscope just below the epineurium of the sciatic nerve of a rat, and ligated and fixed using microsurgery thread (S&T 10-0) (right diagram in Figure 12A). After surgery, the metal conductor was connected to a preamplifier, and the action potential (collective action potential) of the sciatic nerve was recorded using a biosignal recording device (AP1024, TEAC). An example of the measurement result is shown in Figure 12B. The measurement conditions were a sensitivity setting of 2000 μV / mm, a time scale of 1 second, and from the top, the time was resting, muscle contraction, and muscle relaxation.

[0307] The implantable bioelectrode comprising the PEDOT-PSS silk fiber bundle 3 according to the present invention is thread-like, and can be sewn into tissue by surgery. Therefore, compared with conventional large metal electrodes that are less flexible, the bioelectrode according to the present invention has a higher degree of freedom in the placement location, can be fixed in a stable state, and is highly durable and capable of long-term recording because only the minimum necessary parts are exposed and the rest is covered.

[0308] <Evaluation of electrical conductivity> [Examples 1-5] Using the silk thread (No. 9 silk thread), conductive polymer fibers with PEDOT-PSS disposed inside and on the outer periphery of the silk thread were produced by the above-mentioned production method 2b. Specifically, the following samples were prepared: Sample C, which was electrochemically coated once with PEDOT-PSS on the outer periphery and dried; Sample D, which was further impregnated with glycerol on Sample C; Sample E, which was electrochemically coated again with PEDOT-PSS (a total of two coats); and Sample F, which was further impregnated with glycerol on Sample E. The electrical conductivity of each of the samples C, D, E, and F in a dry state (containing no moisture) was measured using the resistance measurement method described in Example 1-2. The results of measuring each resistance value using the same method as in Example 1-2 are shown in Table 1. From the results obtained, it is clear that in order to improve the conductivity and resistance value, a thicker conductor is preferable (two coats are preferable rather than one coat) and the addition of glycerol is preferable.

[0309] [Table 1]

[0310] Example of the second aspect The second aspect of the present invention will be described in more detail below with reference to examples, but the second aspect of the present invention is not limited to these examples in any way.

[0311] [Example 2-1] In this example, silk fibers (No. 9 silk thread: diameter about 280 μm) were prepared as substrate fibers, and a manufacturing device equipped with the rotor electrode of the present invention as shown in Figure 13 and Figures 15A-B was used to produce conductive polymer fibers in which a conductor containing PEDOT-PSS was polymerized and fixed on the outer periphery of the silk fibers and inside the fiber bundle made of the silk fibers. At this time, two types of conductor solutions were prepared: one without additives and one with glycerol added, and electrochemical polymerization and fixation were performed on each.

[0312] As shown in Figures 13A-13B, a pulley-shaped rotor electrode 222 and a roller-shaped rotor electrode 232 were used as rotor electrodes, and these rotor electrodes 222, 232 were arranged alternately so as to sandwich the silk fibers from both sides in the radial direction of the silk fibers. In this case, the pulley-shaped rotor electrode 222 had a pulley 222a with a diameter of 8 mm and a width of 4 mm. The roller-shaped rotor electrode 232 had a roller 232a with a diameter of 6 mm and a width of 3 mm.

[0313] Specifically, as explained in the embodiment of the present invention above, silk fibers were immersed in a conductive solution containing PEDOT-PSS contained in an immersion container, and then vertically pulled up by a winding section. At this time, a comb-shaped electrode as shown in FIG. 14 was used, and the silk fibers were alternately sandwiched from both sides by multiple comb teeth, and electricity was applied while the silk fibers were pulled up. At this time, a DC power supply of 20 μA and 18 V was supplied to the comb-shaped electrode using a DC stabilized power supply (manufactured by Kikusui Electronics Co., Ltd.: PAB18-5.5), and an electric quantity of 3 to 6 mC (electric flux density: 5.85 to 9.95×10) was used to polymerize and fix the silk fibers at a length of 10 mm in the longitudinal direction. 4 Cm 2 To ensure that the required voltage was maintained, a digital multimeter (VOAC7511, manufactured by Iwasaki Electric Co., Ltd.) was used to monitor and adjust the current and voltage.

[0314] The fiber resistance and conductivity of the obtained conductive polymer fiber were measured at a fiber length of 10 mm with a resistance measuring device "DM2561 (manufactured by NF Circuit Design Block Co., Ltd.)" at a direct current of 350 mA. A Nanoclip manufactured by Stack Electronics Co., Ltd. was used as a fiber gripping jig. The measurements were performed in a dry state (a state that did not contain moisture), and the results are shown in Table 2 below. In Table 2, Example 1 shows Example 2-1, and Comparative Example 1 shows Comparative Example 1-1.

[0315] The conductive polymer fibers thus obtained were observed using a stereomicroscope to visually confirm the state of coating with the conductor containing PEDOT-PSS, and a photograph of this observation is shown in FIG. 17A. In addition, the water resistance of the obtained conductive polymer fiber after immersion in physiological saline (0.9% NaCl solution: 20°C) for one month was evaluated by photographing it with a stereomicroscope (using a Leica SZ), and the photograph is shown in Figure 18A.

[0316] [Table 2]

[0317] [Example 2-2] In this example, conductive polymer fibers were produced under the same conditions and procedures as in Example 2-1, except that electricity was applied using comb-shaped electrodes 221, 231 as shown in Fig. 14 in the production apparatus shown in Fig. 13. In this case, the comb-shaped electrodes 221, 231 used had a plurality of comb teeth 221a, 231a with an inter-tooth distance (inter-electrode distance) of 10 mm.

[0318] The conductive polymer fibers thus obtained were then observed using a stereomicroscope to visually confirm the state of coating with the conductor containing PEDOT-PSS, and a photograph of this observation is shown in FIG. 17B.

[0319] [Comparative Example 2-1] In the comparative example, a conductive polymer fiber was produced in which a conductor containing PEDOT-PSS was fixed to the outer periphery of the silk fiber and inside the fiber bundle made of the silk fiber under the same conditions and procedures as in Example 2-1 above, except that the conductor was fixed to the silk fiber (base fiber) by a conventional chemical fixing method. Then, the fiber resistance and conductivity of the obtained conductive polymer fiber were measured by the same method as above, and the results are shown in Table 2.

[0320] Moreover, the water resistance was evaluated in the same manner as in Example 2-1 above, and a photograph taken as a stereomicroscope image is shown in FIG. 18B.

[0321] [Evaluation Results] As shown in the results in Table 2, it was revealed that the conductive polymer fiber of Example 1, in which a conductor containing PEDOT-PSS was electrochemically polymerized and fixed to silk fibers (base fiber) using the manufacturing apparatus of the present invention and the manufacturing method defined in the present invention, had lower fiber resistance and excellent conductivity, regardless of the presence or absence of additives, compared to the conductive polymer fiber of the comparative example produced by a chemical fixation method using a conventional manufacturing apparatus. Furthermore, as shown in the photograph of Figure 17A, in the conductive polymer fiber obtained in Example 2-1, the surface of the silk fiber and even the inside of the fiber bundle are uniformly coated with a conductor comprising PEDOT-PSS, and it can be seen that the conductor is fixed without exposing the silk fiber. Furthermore, as shown in the photograph of Figure 18A, it was confirmed that the conductive polymer fiber obtained in Example 2-1 maintained a state in which the surface of the silk fiber and even the inside of the fiber bundle were coated with a conductor (black color of the silk fiber surface) even after a one-month water resistance test.

[0322] On the other hand, it was revealed that the conductive polymer fiber produced using the conventional chemical fixation method had a higher fiber resistance and a lower conductivity than the conductive polymer fiber of Example 2-1, as shown in Table 2. Furthermore, as shown in the photograph of Figure 18B, after one month of water resistance testing, the conductive polymer fiber obtained in Comparative Example 2-1 was found to have exposed silk fibers (white to gray color on the silk fiber surface), and most of the conductors had been peeled off and lost.

[0323] As shown in the photograph of Fig. 17B, it was confirmed that the conductive polymer fiber obtained using the comb-shaped electrode in Example 2-2 had a part of the silk fiber surface exposed, compared to the conductive polymer fiber of Example 1 produced using the rotor electrode. This is because the conductor peeled off due to contact with some of the comb teeth (metal rod electrode) when electricity was passed through the comb-shaped electrode while it was pulled up vertically. However, since the coverage rate of the silk fiber surface is higher than that of the conductive polymer fiber produced in the comparative example, it is considered to be superior to the conventional product in terms of both fiber resistance and conductivity.

[0324] Example of the third aspect The third aspect of the present invention will now be described in more detail with reference to examples, but the third aspect of the present invention is not limited to the following examples. (Example 3-1; Interdigitated electrode for electroencephalogram) Silk fiber bundles (manufactured by Fujix Co., Ltd., tire size 9, fiber diameter approximately 280 μm) before composite fiber formation were immersed in a solution of PEDOT-PSS (Clevios P, manufactured by Heraeus, Germany) with 0.1% EDOT (manufactured by Heraeus, Germany). Electric current was then passed through the silk fiber bundle, and PEDOT-PSS was electrochemically fixed on the surface and inside of the silk fiber bundle to produce a conductive composite fiber of silk fiber bundle and PEDOT-PSS. Four of these conductive composite fiber bundles were bundled together and fixed in a shape stretched over a comb-shaped polystyrene arch frame (four places, a total of 16 bundles), to obtain a comb-shaped bioelectrode 10 shown in Figures 19A to 19D. A signal cable for an electroencephalogram measuring device (manufactured by Nihon Kohden Corporation) was used as the signal cable 314 to be joined to the contact 311 made of conductive composite fiber. 1 cm of the coating of this signal cable was stripped off to expose the copper wire, which was then wrapped with conductive composite fiber and ligated. The joint between the conductive composite fiber and the signal cable was insulated with an ethylene vinyl alcohol adhesive. At this time, the joint was fixed to the end of the frame together with the signal cable. Before using the bioelectrode in EEG measurement, the contact 311 (conductive composite fiber) was impregnated with glycerol. By impregnating the conductive composite fiber with glycerol, the conductivity and water resistance of the conductive composite fiber are improved, and the flexibility of the fiber is improved, so that good contact between the contact 311 and the scalp can be obtained, enabling stable measurement of EEG. The bioelectrode 310 produced in this embodiment is 12 mm wide, 35 mm long, 6 mm thick, and 2 mm thick at the comb tip, and weighs 1.1 g (electrode part only, not including cable weight), making it thin and lightweight. Furthermore, because of its comb-shaped shape, the bioelectrode 310 can be worn hidden under the hair.

[0325] (Example 3-2; Hairpin-shaped EEG electrode) The same conductive composite fiber as in Example 1 was used. A hairpin-shaped hair clip with a length of 3.5 cm was used as the frame. The hairpin was made of steel, and its surface was painted with a urethane resin. 3 cm of the coating of the signal cable for measuring brain waves (manufactured by Nihon Kohden Co., Ltd.) was stripped, and the conductive composite fiber was wrapped twice around the exposed copper wire to form a contact with a thickness of about 1 mm (FIG. 21C). The two contacts were fixed to both ends of the U-shaped frame of the hairpin via an ethylene vinyl support, and a hairpin-shaped bioelectrode 320 shown in FIGS. 20A to 20D was obtained. The bioelectrode 320 produced in this example has a length of 35 mm, a width of 2-5 mm, and a height of 3 mm, and weighs 0.5 g (electrode part only, not including the weight of the cable). When measuring brain waves using the bioelectrode 320, the electrode itself can hold the hair, so it can be fixed independently, and a holder such as an elastic net may or may not be used.

[0326] FIG. 23A shows human brain waves measured using the hairpin-shaped EEG electrodes of Example 3-2. The hairpin-shaped electrodes of Example 3-2 were placed at C3 and C4 as relevant electrodes. Silver-chloride plate electrodes (NE134A, manufactured by Nihon Kohden Corporation, for collodion electrodes) were placed as indifferent electrodes, fixed to both auricles (earlobes) with tape via absorbent cotton soaked in physiological saline. No pretreatment such as degreasing or exfoliation was performed on the skin where the electrodes were to be placed. FIG. 23A shows the waveform of the awake EEG of an adult male measured in a normal laboratory using MEB5504 manufactured by Nihon Kohden Corporation, with a low-cut filter of 1 Hz and a high-cut filter of 20 Hz. Here, the horizontal axis of the figure is 400 ms / div, and the vertical axis is 50 μV / div.

[0327] FIG. 23B shows the auditory brainstem response (evoked potential) of an adult male measured using the hairpin-shaped EEG electrode of Example 3-2. The measuring device (Nihon Kohden MEB5504) and the electrode placement are the same as those in the EEG measurement of FIG. 23A. A 90 db click sound was input from headphones to both ears, and 1000 arithmetic averages were taken using the standard settings for the auditory evoked potential. Here, the horizontal axis of the figure is 1 ms / div, and the vertical axis is 0.2 μV / div. The evoked potential waveforms measured under conditions of a low-pass filter of 1 Hz and a high-pass filter of 200 Hz show that the bioelectrode of the embodiment of the present invention can be used for evoked potential measurement.

[0328] (Example 3-3: Electrocardiogram Electrodes) In order to compare the stability and noise generation of the measured biosignals, the following three types of electrodes 1 to 3 were placed on the body surface of the same experimental animal (rat), the electrocardiogram was measured simultaneously, and the measured waveforms were compared. Electrode 1 (electrode according to the second embodiment of the present invention): A contactor having 30 fibers (length 12 mm) made of conductive composite fiber impregnated with glycerol, prepared in the same manner as in Example 3-1, was placed on the surface of the rat's body and fixed in place by the two methods described below. Electrode 2 (conventional type): a silver-silver chloride electrode coated with conductive gel (F120S, manufactured by Nihon Kohden Corporation) was placed on the surface of the rat's body and fixed in place using the two methods described below. Electrode 3 (conventional textile electrode): a commercially available electrode for sports heart rate monitors with silver-coated woven fabric (product name: Smart Fabric Sensor, WearLink+ strap electrode, manufactured by Polar) was placed on the surface of the rat's body and fixed in place using the two methods described below.

[0329] The skin of the rat's thorax and back was shaved and washed with disinfectant ethanol, after which the above three types of bioelectrodes (electrodes 1 to 3) were placed on the left and right thorax and back, respectively. In each area, electrodes 1 to 3 were placed as close as possible to each other. A medical bioelectrode (F-150S, Nihon Kohden Corporation) was placed on the thoracic and lumbar regions as an indifferent electrode (body earth). The signals obtained from each bioelectrode were analyzed using a measuring device (Polymate AP1124, TEAC Corporation).

[0330] Measurements were performed on the left and right thoracic and dorsal electrodes of the rat using two different fixing methods (elastic band or tape). The measurement results are shown in FIG. 27. First, when the electrode pads, each of which was constructed by placing a PVC sheet substrate on each electrode, were compressed and fixed with an elastic band, the signals obtained from the three types of electrodes were almost the same, and the signals could be recorded stably. Next, when the bands were removed and the electrodes were fixed with medical adhesive tape (Silkypore) (registered trademark), stable signals were recorded from electrodes 1 and 2 in both the state when the rat was at rest and when it was moving. However, the baseline of the signal from electrode 3 was unstable due to body movement, and hum noise was observed. From the above results, it is clear that the stability of the signal measured by electrode 1 according to the second embodiment of the present invention is similar to that of medical electrode 2 and is superior to that of textile electrode 3.

[0331] (Example 3-4: Regulation of skin moisture) The change in skin moisture content due to the sweating of the skin to which the bioelectrode was attached was measured using a skin moisture content measuring device (corneometer), and the skin moisture content was compared 6 hours after attachment of a conventional bioelectrode and the bioelectrode according to the second embodiment of the present invention. The skin of an adult male's forearm was used as the measurement site. The subjects performed desk work using a computer in an environment with a room temperature of 26 degrees and a humidity of 40%. The skin moisture content at each electrode attachment site was measured using a skin moisture content measuring device (TK59823, Courage + Khazaka Electronic, Germany) before and 6 hours after attachment of each electrode. The measurement results are shown in Figure 10.

[0332] The measurement results shown in FIG. 28 show the moisture content of the skin of the forearm before the electrodes were attached (result A) and 6 hours after the electrodes were attached (results B to E) (error bar 1 SD standard deviation n=10). Result B (+13.7%) is the result of using a bioelectrode according to the second embodiment of the present invention with a ventilation opening in the sheet substrate, result C (+15%) is the result of using the same bioelectrode as result B except that it does not have an opening, result D (+32.3%) is the result of using a conventional bioelectrode with an adhesive gel applied, and result E (+54.4%) is the result of using a conventional bioelectrode using a highly adhesive pad as the sheet substrate. The results in parentheses above show the increase rate of moisture content at the attachment site of each electrode, with the moisture content before the electrodes is set to 100%.

[0333] The specific configurations of the electrodes B to E corresponding to each result are as follows. Electrode B is a bioelectrode in the form shown in Fig. 25A, in which conductive composite fibers prepared in the same manner as in Example 3-1 were impregnated with glycerol, and 30 of the fibers were arranged in parallel on a 20 x 30 mm PVC sheet substrate to form a 7 x 12 mm contact. 2 The PVC sheet had two openings, and was fixed to the skin surface with an adhesive applied to the surface. Electrode C has the same structure as electrode B, except that a sheet-like substrate without an opening is used. Electrode D is a silver-silver chloride medical bioelectrode (F120S, 18x35mm, manufactured by Nihon Kohden Corporation) using a conductive adhesive gel. Electrode E is a silver-silver chloride medical bioelectrode (M150, Nihon Kohden Corporation, diameter 40 mm) with a highly adhesive foam pad. Bioelectrodes B to D were used in a state where they were independently fixed to the subject's forearm.

[0334] In the above results, the skin moisture content of the conventional electrode D using an adhesive gel increased by +32.3%, and the skin moisture content of the conventional electrode E using a highly adhesive foam pad increased by +54.4%. On the other hand, the skin moisture content of the electrode B (with openings) using conductive composite fiber increased by +13.7%, and the skin moisture content of the electrode C (without openings) increased by only +15.0% compared to the conventional types. This result shows that the electrode according to the second embodiment of the present invention is less likely to become stuffy compared to the conventional electrodes. Furthermore, the electrode B with openings in the sheet-like substrate showed a lower increase (increase) in skin moisture content than the electrode C without openings, and the effect of reducing humidity due to the openings was confirmed.

[0335] (Example 3-5: Comparison of electrical properties of bioelectrodes) Comparison of combined resistance between bioelectrodes and skin The following three types of bioelectrodes 4 to 6 were placed on the skin of a human forearm with an electrode interval of 5 cm, and the combined resistance of each bioelectrode and the skin was measured using a bioelectrode impedance meter (manufactured by Melon Technos Co., Ltd.) under the condition of 10 Hz and sine wave. The measurement results are shown in the table below as resistance ratios normalized by the electrode area, with the result of electrode 4 being set as "1". The contact area and impedance of each electrode are also shown. The above results show that the impedance per area of ​​the bioelectrode 4 according to the second embodiment of the present invention is the lowest. Note that the impedance of the sports bioelectrode 6 when the contact surface is dry was very high and could not be measured with the measuring device used.

[0336] The electrode 4 (the present electrode) is a bioelectrode in which 15 conductive composite fibers impregnated with glycerol, prepared in the same manner as in Example 1, are arranged in parallel on a PVC sheet-like substrate, and a 12 mm x 7 mm contact is fixed by the method of the second embodiment. The electrode 4 was placed on the surface of the skin of a human forearm and fixed with an elastic band. At this time, the contact area between the surface of the human forearm skin and the contact part formed by the contact was 84 mm 2 (7x12mm). Electrode 5 (conventional type): A silver-silver chloride electrode coated with conductive gel (Vitrode F 150S, manufactured by Nihon Kohden Corporation) was placed on the skin surface, covered with the sheet-like substrate used in electrode 4, and fixed with an elastic band. At this time, the contact area between the surface of the human forearm skin and electrode 5 was 630 mm 2 It was. Electrode 6 (conventional sports bioelectrode); a commercially available sports heart rate monitor electrode (Smart Fabric Sensor, WearLink+ strap electrode, manufactured by Polar) with a silver-coated nylon fiber fabric was placed on the skin surface and fixed with an elastic band. At this time, the contact area between the surface of the human forearm skin and electrode 6 was 600 mm 2 It was. The results of measuring the combined resistance of each bioelectrode and the skin are shown below.

[0337] [Table 3]

[0338] Frequency response In order to compare the frequency characteristics of the electrode 7 with conductive composite fiber and the conventional electrode 8 using an electrolyte solution, the frequency characteristics of both electrodes were measured using an autolab (PGSTAT, manufactured by Metrohm Autolab). The impedance of the electrode 7 according to the second embodiment of the third aspect of the present invention was shown to be lower than the impedance of the electrode 8 made of silk fiber impregnated with a sodium chloride electrolyte solution in the 10Hz-10KHz range.

[0339] The electrode 7 (main electrode) is a contact of conductive composite fiber with a length of 2 cm that is impregnated with glycerol and prepared in the same manner as in Example 3-1. The electrode 8 is an electrode (length 2 cm) obtained by impregnating the silk fiber (fiber diameter 280 microns) constituting the electrode 7 with a 0.9% sodium chloride electrolyte solution. The frequency characteristics of each electrode were measured as follows:

[0340] [Table 4]

[0341] Example of the fourth aspect The fourth aspect of the present invention will now be described in more detail with reference to examples, but the fourth aspect of the present invention is not limited to the following examples.

[0342] [Example 4-1] (Fabrication of implantable electrodes using conductive composite fibers) A silk fiber bundle (manufactured by Fujix Co., Ltd., tire size 9, fiber diameter approximately 280 μm) before being composite fiber was immersed in a solution of PEDOT-PSS (Clevios P, manufactured by Heraeus, Germany) with 0.1% EDOT (manufactured by Heraeus, Germany). Then, a current was passed through the silk fiber bundle using a comb-shaped electrode, and PEDOT-PSS was electrochemically fixed on the surface and inside of the silk fiber bundle, thereby obtaining a conductive composite fiber bundle of silk fiber bundle and PEDOT-PSS.

[0343] The coating at the tip of a polyimide-coated platinum iridium wire (diameter 30 microns) (manufactured by California Wire Company, USA) was removed, and the wire was ligated to the conductive composite fiber bundle (fiber diameter approximately 280 microns). The ligated portion and the surface of the conductive composite fiber bundle were covered with PDMS (Polydimethylsiloxane) (product name: Sylguard 184, manufactured by Dow Corning Toray Co., Ltd.). The coating of the tip of the conductive composite fiber bundle was peeled off by 500 to 2000 microns to expose the tip of the conductive composite fiber bundle. The PEDOT-PSS contained in the exposed tip was brought into contact with a stainless steel guide needle (diameter 100 microns, manufactured by Seirin Co., Ltd.), and further ethanol was applied to chemically fix (adhere) the tip, thereby producing an implantable electrode.

[0344] (Placement of implantable electrodes) SD rats were anesthetized with isoflurane, a window was made in the skull, and the dura was removed to expose the cerebral cortex. The implantable electrode was placed in the cerebral cortex using an electric actuator (RCD, IAI Co., Ltd.) fixed on a micromanipulator of a brain fixation device (SR-6R, Narishige Co., Ltd.). Specifically, the tip of the electrode was inserted into the left barrel cortex to a depth of 2 mm below the cortex in 0.01 to 0.02 seconds. The platinum iridium wire ligated to the conductive composite fiber bundle was connected to the head amplifier of a brain nerve signal measurement and recording analysis device (model number: RZ51, TDT, USA). A silver-silver chloride wire was placed on the cortex as a reference electrode, and a silver-silver chloride wire was placed under the skull as a body earth. The measured signals were recorded and analyzed using dedicated software (Open EX, open explorer TDT).

[0345] The composite material composed of PEDOT-PSS and silk fibers absorbs water slowly. For example, when the conductive composite fiber bundle in a dry state is immersed in 0.9% NaCl physiological saline, fiber expansion is clearly observed about 30 seconds after the start of immersion. If the conductive composite fiber bundle is inserted into biological tissue at high speed (in a short time (for example, within 1 second), the PEDOT-PSS will absorb water and expand, and the implantable electrode equipped with the conductive composite fiber bundle can be installed in the body before its strength decreases.

[0346] After the implantable electrode was inserted into the brain, the conductive composite fiber containing PEDOT-PSS that was placed in the tissue gradually absorbed bodily fluids (extracellular fluid or cerebrospinal fluid), expanded, and adhered closely to the surrounding tissue. As the conductive composite fiber absorbed more water, the adhesive portion between the conductive composite fiber and the guide needle peeled off, and the conductive composite fiber electrode was separated from the guide needle. The guide needle was then removed by a micromanipulator, and the conductive composite fiber bundle, which was the main body of the electrode, was left in the tissue.

[0347] (Recording of brain action potentials) Using the above insertion method, the implantable electrodes with a fiber diameter of 200 microns and a fiber length of 1 mm were placed at two positions in the left barrel cortex of the rat brain at a depth of 2 mm. The distance between the electrodes was 2 mm. Figure 35A shows the action potential of the rat's cerebral cortex (barrel cortex) recorded by the electrodes. The upper and lower graphs show the signals detected by the two placed electrodes. A burst-like collective action potential was recorded from the two placed electrodes by mechanical stimulation of the whiskers on the right side of the rat. In addition, synchronous collective potentials (↓: arrows) and asynchronous collective potentials (▼ marks) were observed in the waveforms of the two electrodes.

[0348] [Example 4-2] (Fabrication of implantable electrodes) A conductive composite fiber bundle (length 3 mm, wire diameter 50 microns) prepared in the same manner as in Example 4-1 was immersed in glycerol to impregnate the fibers with glycerol. A thread for insertion guide was attached to one end of the obtained conductive composite fiber bundle. A nylon monofilament suture thread with a curved needle for microsurgery (thickness: 10-0, manufactured by S&T) was used as the thread for insertion guide. A bare gold wire (X-wire, manufactured by Tanaka Kikinzoku Kogyo Co., Ltd.) with its insulating coating removed was wound and fixed to the other end of the conductive composite fiber bundle, and the fixed part was covered with PDMS (product name: Sylguard 184, manufactured by Dow Corning Toray Co., Ltd.).

[0349] (Recording of mass action potentials from the sciatic nerve) Wistar rats were anesthetized with isoflurane, and the skin of the left hind leg was incised to expose the left sciatic nerve. Under a microscope, a 10-0 guide thread was inserted into the adventitia of the sciatic nerve bundle. Next, the conductive composite fiber bundle connected to the guide thread was introduced into the sciatic nerve bundle by pulling the guide thread. Since the conductive composite fiber bundle was subjected to a delayed water absorption rate treatment, that is, the water absorption rate was slowed by impregnation with glycerol, it did not obviously expand during the surgical operation and was inserted into the tissue (under the epineurium). The conductive composite fiber bundle expanded and was fixed in the tissue 15 minutes after insertion. The collective action potential of the rat sciatic nerve (Scale bar 1 second 50 μV) measured after fixing the electrode is shown in Figure 35B.

[0350] [Example 4-3] (Electrocardiogram recording of rat) A rat's electrocardiogram was recorded using a conductive composite fiber bundle (length 20 mm, wire diameter 280 microns) prepared in the same manner as in Example 4-1. Under isoflurane anesthesia, the conductive composite fiber bundle was ligated to the subcutaneous tissue layer at three locations of the rat's right anterior chest, left anterior chest, and hypochondrium, to set up an implantable electrode in the subcutaneous tissue. The conductive composite fiber bundle constituting the electrode was connected to the signal cable of the preamplifier of a polygraph (AP1124, manufactured by TEAC Corporation) via a metal wire coated with an insulating and water-resistant polymer. The rat's electrocardiogram (bipolar lead) (scale bar 1 second 50 mV) recorded at a sampling frequency of 1 kHz is shown in Figure 35C.

[0351] [Example 4-4] (Drug delivery using implantable bioelectrodes) The same method as in Example 1 was used, but a relatively long conductive composite fiber bundle was prepared. A silicone bag containing a drug solution was connected as a reservoir to one end of the bundle. At this time, the outer surface of the conductive composite fiber bundle was covered (sealed) with PDMS (product name: Sylguard 184, manufactured by Dow Corning Toray Co., Ltd.) to form a drug transport path. This covering resulted in an implantable bioelectrode having the conductive composite fiber bundle at its core and a PDMS tube forming the outer shell of the transport path.

[0352] For the purpose of measuring the drug transport rate in the conductive composite fiber bundle constituting the core of the electrode, a test conductive composite fiber bundle was prepared and a drug transport test was carried out. First, the central part of a conductive composite fiber bundle (length 20 mm, fiber diameter 280 microns) prepared in the same manner as in Example 4-1 was covered with PDMS over a length of 5 mm, and one end of the conductive composite fiber bundle was immersed in a chamber containing 1 mL of physiological saline containing 100 μM of the fluorescent substance Lucifer Yellow, and the other end was placed in a dish containing 0.5 mL of normal (fluorescent substance-free) physiological saline. The water level of the chamber containing Lucifer Yellow was set to be 5 mm higher than the water level of the dish containing normal physiological saline. These were placed in a thermostatic chamber at 37 degrees, and the concentration of Lucifer Yellow contained in the physiological saline in the dish was measured on the 0th, 1st, 2nd, 3rd, 4th, and 7th days after placement. A fluorescence intensity measuring device (Multilabel Counter, ALVO SX1420, manufactured by PerkinElmer) was used for the measurement, and the measurement was performed by fluorescence measurement. The measurement results are shown in FIG. 36.

[0353] As Lucifer Yellow was transported from the chamber to the dish through the conductive composite fiber bundle, the concentration of Lucifer Yellow in the dish increased at a rate of 0.17 μM / day. This result indicates that Lucifer Yellow permeates (permeates) the conductive composite fiber at a constant rate (Figure 36, ▲ plot and dotted line).

[0354] [Example 4-5] A conductive composite fiber bundle with a central portion coated with PDMS was prepared in the same manner as in Example 4-4, except that the conductive composite fiber bundle was impregnated with glycerol before being coated with PDMS. Using this conductive composite fiber bundle, the drug transport rate was measured in the same manner as in Example 4-4, and the concentration of Lucifer Yellow in the dish increased at a rate of 6.7 μM / day (Figure 36, plot of ■ and solid line). This result indicates that the addition of glycerol to the conductive composite fiber increases the drug transport rate.

[0355] One of the reasons why the drug transport rate is improved by impregnating the conductive composite fiber bundle with glycerol is that when the conductive composite fiber bundle is coated with PDMS, glycerol prevents the PDMS from penetrating (seeping into) the inside of the conductive composite fiber bundle, thereby maintaining the state of the flow path formed by the conductive composite fiber bundle in a state suitable for drug transport.

[0356] [Examples 4-6] (Evaluation of the invasiveness of electrodes to central nervous tissue) The implantation of a bioelectrode into the tissue of the central nervous system causes permanent damage to the tissue of the central nervous system in an area larger than the size of the electrode, which impedes measurement, and this problem has been a problem for some time, and a solution to this problem is required. After inserting the implantable electrode prepared in Example 4-4 according to the present invention into the brain, a drug (GSNO: S-Nitrosoglutathione) that has the effect of reducing damage to the central nervous tissue was administered through the flow path (drug transport path) of the electrode, and whether the damage (invasiveness) caused to the central nervous tissue by the implantation of the electrode can be reduced was examined by animal experiments (Figures 34B to 34D).

[0357] The degree of damage to neural tissue caused by the insertion of electrodes into the rat brain was evaluated by immunohistochemical staining of glial cells (astrocytes) in the rat cerebral cortex and the degree of neural tissue loss. Immunohistochemical staining was performed as follows: 25-micron frozen sections were prepared from the cerebral cortex perfusion-fixed with 4% paraformaldehyde, and anti-GFAP antibody (MAB360 Chemicon) was bound at 1:1000 dilution at 4℃ overnight. The sections were then labeled with a secondary antibody (Alexa 568) and observed under a fluorescent microscope (BX51, Olympus Corporation).

[0358] When a conventional metal needle electrode was implanted in a rat brain and the cerebral cortex was observed one week later under a fluorescence microscope, a significant tissue defect (black area) was observed beyond the area where the metal needle electrode was implanted (dotted area), as shown in Figure 34D. GFAP-positive glial cells (astrocytes) had proliferated in the neural tissue (▲ in Figure 34D). Glial cells had proliferated densely in the area in contact with the electrode in particular, forming a glial scar (arrow in Figure 34D). Thus, seven days after the electrode was implanted, significant tissue defect and the formation of a glial scar (arrow) and a cluster of glial cells (▲) were observed.

[0359] Using the conventional metal needle electrode placed in this way, we measured the collective action potential of the rat's cerebral cortex. Measurements were taken on the 1st and 7th days after the electrode was implanted. The results are shown in the lower part of Figure 34A (Conventional) (Scale bar 250ms 40mV). The measured signal on the 1st day was good, but on the 7th day, the measured waveform was reduced and spikes were missing (arrows). In Figure 34A, Day 1 is the measurement record from the 1st day after the electrode was implanted, and Day 7 is the measurement record from the 7th day after the electrode was implanted. The arrows indicate the missing spikes.

[0360] On the other hand, the implantable electrode prepared in Example 4-4 was inserted into the brain of a rat, and an anti-inflammatory drug (GSNO) was administered at 15 μg / day per 250 g body weight via the drug transport pathway of the electrode. The rate of drug transport was adjusted by a small osmotic pump (Alzet, USA) connected to the drug transport channel. In the present example, in which an anti-inflammatory drug was administered, the tissue defect was smaller than that of the conventional electrode, and the tissue defect was limited to the electrode installation area (dotted area) (Figure 34C). In addition, the proliferation of glial cells in the nerve tissue was mild, and no clear glial scar was observed at the contact area with the electrode (Figure 34C). In Figure 34C, GSNO was administered around the electrode, and even 7 days after the electrode was implanted, the tissue defect was limited to the electrode installation area (dotted line), and the proliferation of glial cells in the tissue was also low. For comparison, FIG. 34B shows a fluorescent immunostained image of glial cells in a normal cerebral cortex without implanted electrodes.

[0361] Using the electrode of Example 4 according to the present invention thus installed, the collective action potential of the rat's cerebral cortex was measured. After the electrodes were implanted, the measurements were made on the first and seventh days. The results are shown in the upper part of FIG. 34A (PEDOT-PSS) (Scale bar 250 ms 40 mV). Good waveforms were observed in both the measurement signals on the first and seventh days.

[0362] The configurations and combinations thereof in the above-described embodiments are merely examples, and additions, omissions, substitutions, and other modifications of the configurations are possible without departing from the spirit of the present invention. Furthermore, the present invention is not limited to the embodiments, but is limited only by the scope of the claims. [Industrial Applicability]

[0363] The present invention can provide a conductive polymer fiber that is excellent in electrical conductivity, strength in a dry state and in a wet state, and flexibility, and a bioelectrode including the conductive polymer fiber. The present invention provides a method and an apparatus for producing conductive polymer fibers, which can impregnate or adhere a conductor containing PEDOT-PSS to a base fiber and electrochemically polymerize and fix the conductor continuously, thereby enabling the efficient production of conductive polymer fibers with excellent conductivity and durability. The bioelectrode according to the present invention can be widely used as a body surface-mounted bioelectrode that can be used continuously for long periods of time in a wide range of fields, including medicine, health promotion, information technology, wearable computers, etc. It is possible to provide a bioelectrode that is more comfortable to wear than conventional bioelectrodes and a biosignal measuring device equipped with the bioelectrode. The implantable electrode according to the present invention can be widely used as an implantable bioelectrode in a wide range of fields, including medicine, health promotion, information technology, wearable computers, etc. More specifically, it can be used for electrical stimulation treatments such as deep brain stimulation, implantable electrodes for recording neural activity, brain-machine interfaces, etc. It is possible to provide an implantable electrode that is capable of detecting weak electrical signals within a living body, has excellent biocompatibility, and is minimally invasive to biological tissue. [Explanation of symbols]

[0364] 2...Conductive polymer fiber 3…Fixing cord 4…Rubber band 5…Metal conductor wire 6. Human body surface 33...Metal or carbon 34...Conductor 54...Conductor 63...Insulating layer L: Diameter of the base fiber h…Thickness of the coated conductor 10, 20, 30, 40, 50, 60...Conductive polymer fiber 11, 21, 31, 41, 51, 61...Base fiber 12, 22, 32, 42, 52, 62...Conductors 210, 250...Conductive polymer fiber manufacturing equipment (manufacturing equipment) 202...multiple electrodes (anodes) 221... Interdigital electrode (multiple electrodes; anode) 221a...Comb tooth (electrode) 221b…Terminal 222...Rotor electrode (pulley-shaped electrode) 222a…pulley 222b…Groove 222c…Metal shaft part 203...Multiple electrodes (cathode) 231... Interdigital electrode (multiple electrodes; cathode) 231a...Comb tooth (electrode) 231b…Terminal 232...Rotor electrode (roller-shaped electrode) 232a…Roller 232b…Outer surface 232c…Metal shaft part 252...Single (monopolar) electrode (anode) 253...Single (monopolar) electrode (cathode) 204...Conductor solution 205, 255...Immersion container 206, 256…Thread spools 207, 257...Chamber (humidity control section) 208, 258...Drying section 209, 259…winding section

[0365] 310...Bioelectrode 311…Contact 312…First frame 313…Second frame 314…Signal cable H…Hair (head hair) S…Skin (scalp) 320...Bioelectrode 321...Contact 322…Third frame 323…Fourth frame 324…Signal cable 321a…Conductive composite fiber 321b…Metal wire rod 321c…Insulating coating material 321e…Conductive composite fiber 321f…Thin metal wire 321g…Core material 321z…Insulating coating material N: Low tension net 330...Bioelectrode 331...Contact 332…Contact part 333...Substrate (base material) 334…Signal cable 335…Holder 336…Opening 337...Moisture control pad 338...Electrode pad 339...Amplifier (external device) B…Body (trunk) T…Undergarments (shirt) 401...Conductive composite fiber bundle 402...Electric wire (metal conductor wire) 403…Connection section 404...Polymer 405...Needle (Guide needle) 406...Thread 407…Reservoir 408…Chamber 409…Tube connector 410, 420, 430, 440…Implantable electrodes N'... nerve cord

Claims

1. A conductive fiber characterized in that the base fiber is coated with carbon.

2. 2. The conductive fiber according to claim 1, wherein the carbon is carbon black.

3. The conductive fiber according to claim 1 or 2, characterized in that the base fiber is a straight or twisted fiber.

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