WO2021200859A1

WO2021200859A1 - Electrode and biosensor using same

Info

Publication number: WO2021200859A1
Application number: PCT/JP2021/013381
Authority: WO
Inventors: ダニエルポポビッチ
Original assignee: 日東電工株式会社
Priority date: 2020-03-30
Filing date: 2021-03-29
Publication date: 2021-10-07
Also published as: JPWO2021200859A1

Abstract

The present invention provides an electrode constitution that enables noise reduction and detection of stable signal waveforms, and a biosensor using the same. This electrode, in which a polymer material is used, comprises a conductive layer containing the polymer material and one or more holes, said holes penetrating through the conductive layer in the thickness direction. The area of the electrode is 80 mm² or greater and the distance from an end of the conductive layer to the nearest hole and the distance between adjacent holes are 2.7 mm or longer.

Description

Electrodes and biosensors using them

The present invention relates to an electrode and a biosensor using the electrode.

When a dry electrode is used for a sensor that detects biological signals such as electrocardiography (ECG) waveforms, pulse waves, electroencephalograms, and myoelectric signals, the electrodes are exposed on the surface of the sensor and the electrodes are brought into direct contact with the skin to make the living body. Measure the potential. At this time, it is required that the electrodes are in stable contact with the skin. It is known that an electrode is placed on the surface of a biocompatible polymer substrate and attached to the skin to detect data (see, for example, Patent Document 1).

In the case of a dry electrode, the contact impedance is higher than that of a wet electrode using gel or the like, and efficient noise removal is a technical issue. Although noise can be removed to some extent by digital signal processing, it cannot be completely removed. This is because noise generated at the same frequency as the peak waveform in the signal cannot be removed by digital processing.

An object of the present invention is to provide an electrode configuration that reduces noise and detects a stable signal waveform, and a biosensor using the electrode configuration.

In one aspect of the present invention, an electrode made of a polymer material is
The conductive layer containing the polymer material and
One or more holes penetrating the conductive layer in the thickness direction,
Have,
The electrode has ^{an area of 80 mm 2} or more, and the distance from the end of the conductive layer to the nearest hole and the distance between adjacent holes are at least 2.7 mm.

With the above configuration, noise can be reduced and a stable signal waveform can be detected.

It is a schematic diagram which shows an example of the biological sensor to which the electrode of an embodiment is applied. It is a figure which shows the example of the hole formed in an electrode. It is a figure which shows the state at the time of use of a biological sensor. It is a schematic diagram of the setup which evaluates the influence of the electrode size on the biological waveform. It is a signal waveform measured by changing the electrode size. It is a signal waveform measured by changing the electrode size. It is a signal waveform measured by changing the electrode size. It is a signal waveform measured by changing the electrode size. It is a signal waveform measured by changing the electrode size. It is a signal waveform measured by changing the electrode size. It is a signal waveform measured by changing the electrode size. It is a signal waveform measured by changing the electrode size. It is a signal waveform measured by changing the electrode size. It is a signal waveform measured by changing the electrode size. This is an evaluation result of the influence of the electrode size on the ECG waveform. It is a schematic diagram of the sample which evaluates the number and size of a hole provided in an electrode. It is an ECG waveform measured with a sample having a hole diameter of 3 mm. It is an FFT spectrum of the ECG waveform of FIG. It is an ECG waveform measured with a sample having a hole diameter of 5 mm. It is an FFT spectrum of the ECG waveform of FIG. It is an ECG waveform measured with a sample having a hole diameter of 6 mm. It is an FFT spectrum of the ECG waveform of FIG. This is an evaluation result of the influence of the holes formed in the electrodes on the ECG waveform. Another evaluation result of the effect of the holes formed in the electrodes on the stability of the measurement. It is a figure which shows the example of the electrode shape.

The inventor includes (1) the area of the entire electrode and (2) the design of the holes formed in the electrode (number, size, and arrangement of holes) in order to suppress noise and obtain a stable biological signal. ), Found to optimize. The area of the electrode and the design of the holes formed determine the width of the conductive path created in the electrode during measurement. The width of the conductive path is defined as the distance from the edge of the electrode to the nearest hole and / or the distance between adjacent holes. When the width of the conductive path is appropriate, the signal detection sensitivity is good and an accurate signal waveform can be obtained. The number, size, and arrangement of holes formed in the electrode also affects the stability of contact between the electrode and the skin. This is because, as will be described later, when the biosensor is pressed against the skin, the adhesive layer enters the holes of the electrodes to improve the adhesion to the skin. When the adhesion between the electrode and the skin is improved, the contact impedance, that is, noise is suppressed.

In the following, in order to suppress noise and obtain a stable signal waveform, the area of the electrode and the optimum design of the holes formed in the electrode will be examined.

FIG. 1 is a schematic view of a biological sensor 100 to which the electrode 10 of the embodiment is applied. FIG. 1A is a top view and FIG. 1B is a bottom view. The surface on which the biosensor 100 is arranged is defined as the XY plane, and the thickness direction of the biosensor 100 is defined as the Z direction. The biosensor 100 has a package 101 that includes a front surface 102 and a back surface 103. The front surface 102 is the outermost surface of the top cover of the package 101, and the back surface 103 is the attachment surface for attaching the biosensor 100 to the living body.

The electronic component 150 is housed in the space 104 inside the package 101 of the biosensor 100. The electronic component 150 includes a microprocessor, a memory, a battery, and the like mounted on a circuit board. Programmable integrated circuits and logic devices may be implemented as needed.

The electrode 10 is exposed on the back surface 103 of the package 101. The electrode 10 is connected to the electronic component 150 by the wiring 160. The electrode 10 functions as a probe and comes into contact with the skin during measurement to detect a biological signal. The biological signal detected by the electrode 10 is processed by the electronic component 150 and recorded in the memory for a certain period of time.

In the example of FIG. 1, biometric information is acquired in a single channel using a pair of electrodes 10, but the present invention is not limited to this example. Two differential electrode pairs and one ground electrode may be used, or two or more pairs of electrodes may be used to acquire biometric information in a multi-channel manner. In either case, the front surface of the electrode 10 is exposed on the back surface 103 of the package 101. A wearable sensor is realized by attaching the biosensor 100 to the living body on the back surface 103 where the electrode 10 is exposed.

The electrode 10 is made of a polymer material. The polymer material is superior in flexibility, oxidation resistance, etc. as compared with the metal material, and is suitable for direct contact with the skin. The electrode 10 can be formed of, for example, a conductive composition containing a conductive polymer and a binder resin.

As the conductive polymer, polythiophene, polyacetylene, polypyrrole, polyaniline, polyphenylene vinylene, one of these, or a combination of two or more thereof can be used. As an example, polythiophene compounds, especially polystyrene sulfonic acid (poly4-styrene sulfonate; PEDOT-PSS doped with PSS) is used.

The content of the conductive polymer is preferably 0.20 parts by mass to 20 parts by mass, and more preferably 2.5 parts by mass to 15 parts by mass with respect to 100 parts by mass of the conductive composition. , 3.0 parts by mass to 12 parts by mass is more preferable. When the content of the conductive polymer is in the range of 0.20 parts by mass to 20 parts by mass with respect to the conductive composition, high conductivity, toughness, and flexibility can be imparted to the conductive composition. ..

The binder resin may be a water-soluble polymer or a water-insoluble polymer, but in the embodiment, the water-soluble polymer is used from the viewpoint of compatibility with other components contained in the conductive composition. The water-soluble polymer contains a polymer (hydrophilic polymer) that is completely insoluble in water and has hydrophilicity. As the water-soluble polymer, a hydroxyl group-containing polymer or the like can be used. As the hydroxyl group-containing polymer, saccharides such as agarose, polyvinyl alcohol (PVA), modified polyvinyl alcohol, or a copolymer of acrylic acid and sodium acrylate can be used. These may be used alone or in combination of two or more. Among these, polyvinyl alcohol or modified polyvinyl alcohol is preferable, and modified polyvinyl alcohol is more preferable.

The content of the binder resin is preferably 5 parts by mass to 140 parts by mass, more preferably 10 parts by mass to 100 parts by mass, and 20 parts by mass to 70 parts by mass with respect to 100 parts by mass of the conductive composition. It is more preferably parts by mass. When the content of the binder resin is in the range of 5 parts by mass to 140 parts by mass with respect to the conductive composition, excellent conductivity, toughness, and flexibility can be imparted to the conductive composition.

The conductive composition may contain at least one of a cross-linking agent and a plasticizer. Crosslinkers and plasticizers impart toughness and flexibility to the conductive composition. "Toughness" refers to the property of having both strength and elongation. Flexibility refers to the property of being able to suppress the occurrence of damage such as breakage at the bent portion even when the electrode 10 which is a cured product of the conductive composition is bent.

In addition to cross-linking agents or plasticizers, surfactants, softeners, stabilizers, leveling agents, antioxidants, hydrolysis inhibitors, leavening agents, thickeners, colorings, as needed, in conductive compositions Agents, fillers and the like may be included in appropriate proportions.

The electrode 10 has the above-mentioned conductive layer 11 of the polymer material and one or more holes 15 penetrating the conductive layer 11. As the material of the electrode 10, in addition to the above-mentioned conductive polymer, a polymer mixed with conductive nanopowder such as carbon, Ag, Ni, or a polymer mixed with carbon nanotubes (CNT) I and silver (Ag) nanotubes. , Carbon nanobud (CNB) and the like may be mixed. At least one of the size and shape of the electrode 10, the number of holes 15, the size, and the arrangement of the electrode 10 is optimal so that the electrode 10 can stably contact the skin for a certain period of time and detect a biological signal with high sensitivity and low noise. Designed in range. The optimum range of these parameters will be described later.

FIG. 2 shows an example of a hole 15 formed in the electrode 10. The shape of the hole 15 does not matter as long as the electrode 10 has an area of a certain value or more and the minimum distance between the hole 15 and the hole 15 is a certain value or more. It may be a circular hole 15 as shown in FIG. 2A, or an elliptical, oval, or oval hole 15 as shown in FIG. 2B. It may be a polygonal hole 15 as shown in FIG. 2C. The polygon is not limited to a hexagon, and may be another polygon such as a quadrangle, a pentagon, or an octagon.

The diameter of the hole 15 formed in the electrode 10 is set to 3 mm or more and 8 mm or less, as will be described later. This is because within this range, the adhesive layer adheres to the skin through the holes 15 and the contact between the electrode 10 and the skin can be maintained. When the biosensor 100 is used during sweating or for a long period of time (for example, one week), the diameter of the hole 15 is preferably 5 mm or more and 8 mm or less. From the viewpoint of widening the width of the conductive path formed on the electrode 10 at the time of measurement, it is desirable that the diameter of the hole is 5 mm or more and 6 mm or less.

The size and shape of the holes 15 formed in the electrode 10 do not necessarily have to be the same. If the diameter in the above range or the total opening area of the holes 15 is secured in a certain range, the shapes shown in FIGS. 2A to 2C and other shapes are mixed. May be good. In the case of a polygon, the diameter shall be the distance from one vertex to the opposite vertex or side. In the case of an ellipse or an oval, the diameter is the average of the minor and major. The thickness of the electrode 10 can be appropriately designed as long as the holes 15 can be formed, have sufficient strength, and are easy to handle. As an example, the thickness is 0.1 μm to 100 μm.

FIG. 3 shows a state when the biosensor 100 is used. The back surface 103 of the biosensor 100 is pressed against and attached to the skin 20 of the person to be measured. The biosensor 100 may have a base material 110 that holds the electronic component 150 as part of the package 101. As an example, the electronic component 150 may be mounted on the surface 110a side of the base material 110 via an insulating layer, and a wiring 160 extending from the electronic component 150 may be formed.

A pressure-sensitive adhesive layer 25 is provided on the back surface 110b of the base material 110, and the electrode 10 is held by the pressure-sensitive adhesive layer 25. The pressure-sensitive adhesive layer 25 and the electrode 10 may be protected by a release paper or the like when the biosensor 100 is not in use.

When the back surface 103 of the package 101 of the biosensor 100 is pressed against the skin 20 during use, the biosensor 100 is fixed to the skin 20 by the pressure-sensitive adhesive layer 25. The electrode 10 exposed on the back surface 103 is also directly pressed against the skin 20. The electrode 10 is provided with a hole 15 whose size or diameter is designed within a predetermined range. When the biosensor 100 is pressed against the skin, the pressure-sensitive adhesive layer 25 adheres to the skin 20 through the holes 15. By adhering the pressure-sensitive adhesive layer 25 to the skin 20 through the holes 15, it is possible to prevent the electrode 10 from floating from the skin 20 and peeling off. By stabilizing the contact between the electrode 10 and the skin, it is possible to reduce noise caused by the contact impedance and stably detect the biological signal.

<Evaluation analysis for optimizing the size of the electrode 10>
FIG. 4 is a schematic diagram of a setup for evaluating the effect of electrode size on the biological signal waveform. FIG. 4A is a side view of the setup, and FIG. 4B is a bottom view. The electrode sample 10S is held by the pressure-sensitive adhesive tape 250, and a part of the contact surface 111 of the electrode sample 10S with the skin is covered with the insulating layer 22. The effective area of the electrode sample 10S is changed by changing the area of the region covered by the insulating layer 22.

The electrode sample 10S is connected to the ECG monitor with a conductive hook 21. The electrode sample 10S is formed to a thickness of 25 μm using the above-mentioned PEDOT-PSS. One side of the rectangular electrode sample 10S is fixed to the length L, and the width W of the adjacent side is made variable. In the evaluation experiment, the length L is fixed at 20 mm (L = 20 mm), and the width W is changed in the range of 0 mm to 14 mm. In this setup, no holes are formed in the electrode sample 10S in order to estimate the optimum range of electrode area.

5A-5J show signal waveforms acquired using electrode samples 10S of various sizes. Through FIGS. 5A to 5J, (a) is an ECG waveform and (b) is a fast Fourier transform (FFT) spectrum thereof. The horizontal axis of the ECG waveform is time (seconds), and the vertical axis is electric potential. The horizontal axis of the FFT spectrum is frequency, and the vertical axis is magnitude. The ECG waveform alone cannot evaluate the hidden noise generated at the same frequency as the peak. Therefore, the state of noise is observed in the FFT spectrum.

In FIG. 5A, the width W of the electrode sample 10S is set to 0 mm, and the area of the electrode is zero. Since the entire electrode sample 10S is covered with the insulating layer 22 and is not in contact with the living body, the ECG waveform cannot be detected, and only noise is detected. Even in the FFT spectrum, only noise is present and no periodic signal peak appears.

In FIG. 5B, the width W of the electrode sample 10S is set to 1 mm, and the area of the electrode is 1 × 20 = 20 mm ² . The electrode sample 10S is in contact with the skin only in a very narrow area. Although the ECG waveform including the PQRST wave has been obtained, the position of the waveform is shifted up and down, and the influence of respiration is remarkable. In addition, the baseline has fluctuated, showing the effects of body movements. The baseline is represented by a line connecting the beginning of the P wave to the beginning of the next P wave. Furthermore, the U wave, which is a small wave that should appear after the T wave, is not observed. The U wave is used to determine diseases such as hypokalemia (when the peak direction is the same as the T wave), hypertension, and cardiomyopathy (when the peak direction is opposite to the T wave), and the detection of the U wave is ECG. It is one of the important elements of diagnosis. Looking at the FFT spectrum, some periodic peaks appear, but it is buried in the noise immediately after the measurement and biometric information cannot be obtained. The range of 0 to 0.5 Hz of the FFT is generally associated with baseline variability, and the paceline is also unstable when there is a lot of noise in this region.

In FIG. 5C, the width W of the electrode sample 10S is set to 2 mm, and the area of the electrode is 2 × 20 = 40 mm ² . At this time as well, the size of the electrode sample 10S is insufficient. Compared with FIG. 5B, the vertical shift of the ECG waveform (effect of respiration) is small, but the baseline is not stable and a clear U wave is not observed. In the FFT spectrum, some periodic signals (SIG) are observed, but the noise (N) is large with respect to the signal (SIG), and the signal-to-noise ratio (SNR) is low. Too much.

In FIG. 5D, the width W of the electrode sample 10S is set to 3 mm, and the area of the electrode is 3 × 20 = 60 mm ² . Similar to FIG. 5C, the effect of respiration is less than in FIG. 5B, but the baseline is not stable and no clear U wave is observed. Although some periodic signals (SIG) are observed in the FFT spectrum, the SNR is small and sufficiently reliable biometric information cannot be obtained.

In FIG. 5E, the width W of the electrode sample 10S is set to 4 mm, and the area of the electrode is 4 × 20 = 80 mm ² . In FIG. 5E, the baseline has begun to stabilize and the effects of respiration are lessened. Also, a small U wave is observed after the T wave. A periodic signal (SIG) is observed in the FFT spectrum, and the SNR is improved compared to FIG. 5D. If the electrode size is 80 mm ² , the ECG waveform can be detected stably.

In FIG. 5F, the width W of the electrode sample 10S is set to 5 mm, and the area of the electrode is 5 × 20 = 100 mm ² . In FIG. 5E, the baseline is stable and the effect of respiration is small. After the T wave, the U wave is stably observed. A high SNR is obtained in the FFT spectrum.

In FIG. 5G, the width W of the electrode sample 10S is set to 7 mm, and the area of the electrode is 7 × 20 = 140 mm ² . In FIG. 5F, the baseline is stable and there is almost no effect of respiration. After the T wave, the U wave is clearly observed. A high SNR is obtained in the FFT spectrum.

In FIGS. 5H, 5I, and 5J, the width W of the electrode sample 10S is set to 10 mm, 12 mm, and 14 mm, respectively. The area of the electrodes is 10 × 20 = 200 mm ² , 12 × 20 = 240 mm ² , and 14 × 20 = 280 mm ² . In all of FIGS. 5H to 5J, the baseline is stable and there is almost no effect of respiration. After the T wave, a U wave is clearly observed, and a high SNR is obtained in the FFT spectrum.

FIG. 6 is a table summarizing the measurement results of FIGS. 5A to 5J. The leftmost column is the electrode size (W × Lmm ² ), the second column from the left is the difference between the maximum and minimum potentials of the ECG waveform (peak-to-valley), and the third column from the left is the baseline. Stability, rightmost column shows SNR.

When the area of the electrode is 3 × 20 mm ² or less, the baseline of the ECG waveform is uneasy and there is a lot of noise. It is considered that this is because the contact area between the electrode and the skin is insufficient. By setting the area of the electrode to 4 × 20 mm ² or more, the baseline of the ECG waveform is stabilized and the U wave is observed. In addition, the noise is reduced and a sufficient SNR can be obtained. From this, it is desirable that the area of the entire electrode used in the biosensor 100 is 80 mm ² or more. However, if the size of the electrode 10 is too large, it hinders the miniaturization of the sensor and the noise picked up becomes large. The size of the electrode 10 is preferably 80 mm ² or more and 280 mm ² or less, more preferably 80 mm ² or more and 200 mm ² or less.

In FIGS. 5A to 5J, the ECG waveform is measured in a resting state using the electrode sample 10S in which the hole 15 is not formed. That is, the ECG waveform is measured with only the outer circumference of the electrode sample 10S fixed to the skin with the pressure-sensitive adhesive tape 250. On the other hand, in daily life, there is body movement, and in the case of the biological sensor 100 that performs ECG measurement, twisting or bending of the body affects the adhesiveness between the electrode 10 and the skin. In order to measure the ECG signal for a certain period of time (for example, 24 hours or more) in daily life, the adhesion to the skin is maintained not only on the four sides of the electrode 10 but also in the inner region of the electrode 10. Is desirable. Below, the optimum hole design for bringing the pressure-sensitive adhesive layer 25 into contact with the skin through the holes 15 formed in the electrode 10 will be examined.

<Evaluation analysis for optimal design of hole 15>
FIG. 7 is a schematic view of an electrode sample 10S for evaluating the number and size of holes 15 formed in the electrode 10. Similar to FIG. 4, a part of the contact surface of the electrode sample 10S with the skin is covered with the insulating layer 22, and the ECG signal is measured via the hook 21. The area of the electrode sample 10S ^{is fixed to L × W mm 2} , and the number and size of the holes 15 formed in the conductive layer 11 are changed. The conductive layer 11 is formed of PEDOT-PSS, and the holes 15 are formed as circular holes.

Specifically, the area of the electrode sample 10S ^{is fixed to 20 × 14 mm 2} , the diameter of the holes 15 is changed in the range of 3 mm to 6 mm, and the number of holes 15 is changed in the range of 1 to 5. In FIG. 7A, the number of holes 15 is one. When the number of holes 15 is one, if the holes 15 are arranged in the center of the conductive layer 11, the adhesion between the skin and the electrode 10 is stable.

When there is one hole 15, the diameter φ of the hole 15 is changed to 3 mm, 5 mm, and 6 mm. The minimum width Pmin of the conductive path formed in the electrode sample 10S at the time of measurement is the distance from the edge of the long side (L = 20 mm) of the conductive layer 11 to the contour of the hole 15. When the diameter φ of the hole 15 is 3 mm, the minimum width Pmin is 5.5 mm. When the diameter φ is 5 mm, the minimum width Pmin is 4.5 mm. When the diameter φ is 6 mm, the minimum width Pmin is 4 mm. As will be described later, the minimum width Pmin of the conductive path needs to be at least 2.7 mm. Therefore, when there is one hole 15, the diameter φ of the hole 15 may be 8 mm. In this case, the minimum width Pmin is 3 mm, and a sufficient conductive path is formed around the hole 15.

In (B) of FIG. 7, the number of holes 15 is two. When the number of holes 15 is two, it is desirable to evenly arrange the two holes 15 in the long axis direction of the electrode sample 10S from the viewpoint of adhesiveness. The diameter φ of the hole 15 is changed to 3 mm, 5 mm, and 6 mm. The minimum width Pmin of the conductive path formed in the electrode sample 10S at the time of measurement is the distance from the edge of the short side (W = 14 mm) of the conductive layer 11 to the contour of the hole 15 and / or between the adjacent holes 15. The distance. When the diameter φ is 3 mm, the minimum width Pmin is 14/3 mm (about 4.7 mm). When the diameter φ is 5 mm, the minimum width Pmin is 10/3 mm (about 3.3 mm). When the diameter φ is 6 mm, the minimum width Pmin is 8/3 mm (about 2.74 mm).

In (C) of FIG. 7, the number of holes 15 is four. When the four holes 15 are evenly arranged in the electrode sample 10S, two holes are arranged vertically and two horizontally. Since the two holes 15 are also arranged along the short side (W = 14 mm), the minimum width Pmin of the conductive path formed in the electrode sample 10S at the time of measurement is the short side (W = 14 mm) of the conductive layer 11. ) To the contour of the hole 15 and / or the distance between adjacent holes 15 in the minor axis direction. When the diameter φ is 3 mm, the minimum width Pmin is 8/3 mm (about 2.7 mm).

When the number of holes 15 is 4, it becomes difficult to form holes 15 having a diameter φ of 5 mm or more. This is because the adjacent holes 15 in the short axis direction are too close to each other, and it is not possible to secure a conductive path having a width sufficient for detecting a signal. The strength of the electrode sample 10S also decreases.

In (D) of FIG. 7, the number of holes 15 is five. When the five holes 15 are evenly arranged in the electrode sample 10S, one is arranged in the center and four are arranged in the periphery. The minimum width Pmin of the conductive path formed in the electrode sample 10S at the time of measurement is the distance from the edge of the short side (W = 14 mm) of the conductive layer 11 to the contour of the hole 15 and / or between the adjacent holes 15. The distance. When the diameter φ is 3 mm, the distance between the holes 15 adjacent to each other in the short side direction is 8/3 mm (about 2.7 mm). For the same reason as (C) in FIG. 7, it is difficult to form a hole 15 having a diameter φ of 5 mm or more.

When there are five holes 15, the distance between the diagonally adjacent holes 15 and the holes 15 is as narrow as Pmin in the short side direction. When viewed as a whole of the electrode sample 10S, (C) in FIG. 7 has a wider conductive path and stronger mechanical strength.

The ECG waveform and its FFT spectrum are measured using the electrode samples 10S shown in FIGS. 7A to 7D.

FIG. 8 is an ECG waveform when the diameter φ of the hole 15 is 3 mm. (A) of FIG. 8 shows the ECG waveform when the hole 15 is not formed as a reference. 8 (B), (C), (D), and (E) are ECG waveforms when the number of holes 15 having a diameter of 3 mm is changed to 1, 2, 4, and 5. be.

In FIGS. 8B and 8C, the baseline is slightly deviated. In (C), the signal potential itself is also low. This is because the number of holes 15 is small, and the contact area between the pressure-sensitive adhesive layer and the skin cannot be sufficiently secured. It is considered that the electrode sample 10S is lifted from the skin. In FIGS. 8D and 8E, the ECG waveform is stable because the total hole area increases.

FIG. 9 is an FFT spectrum of FIGS. 8A to 8E. In FIGS. 8 (D) and 8 (E), the SNR is good and the signal is observed with low noise. In FIGS. 8B and 8C, the SNR is decreased.

From FIGS. 8 and 9, when the holes 15 having a diameter of 3 mm are provided, the number of holes 15 is increased on condition that the minimum width Pmin of the conductive path generated in the electrode 10 is maintained at a certain value or more. It is desirable to increase the total hole area.

FIG. 10 is an ECG waveform when the diameter φ of the hole 15 is 5 mm. (A) of FIG. 10 shows the ECG waveform when the hole 15 is not formed as a reference. FIG. 10B is an ECG waveform when the number of holes 15 is one, and FIG. 10C is an ECG waveform when the number of holes 15 is two. In both (B) and (C) of FIG. 10, the baseline is stable and the waveform positions are aligned. It is considered that this is because the contact area between the pressure-sensitive adhesive layer and the skin is secured, and the electrode sample 10S is stably held on the skin.

FIG. 11 is an FFT spectrum of FIGS. 10A to 10C. The SNR is good in both (B) and (C) of FIG. 10, and the SNR is particularly good when the number of holes 15 is two. From this result, it is expected that a stable ECG signal can be obtained with low noise even when the diameter of the hole 15 is set to 8 mm, for example. This is because a wide contact area is secured in the hole 15, and a sufficient conductive path is secured in the conductive layer 11.

FIG. 12 is an ECG waveform when the diameter φ of the hole 15 is 6 mm. (A) of FIG. 12 shows the ECG waveform when the hole 15 is not formed as a reference. FIG. 12B is an ECG waveform when the number of holes 15 is one, and FIG. 12C is an ECG waveform when the number of holes 15 is two. In FIG. 12B, the baseline is stable and the waveform positions are aligned, but in FIG. 12C, the baseline fluctuates and the signal waveform is also small. It is considered that this is because the total area occupied by the holes 15 is too large and the signal detection area becomes small.

FIG. 13 is an FFT spectrum of FIGS. 12A to 12C. The SNR in FIG. 13 (B) is good, but the SNR in FIG. 13 (C) is low. Therefore, the SNR is particularly good when the number of holes 15 is two.

FIG. 14 is an evaluation result of the influence of the holes 15 formed in the electrode 10 on the ECG waveform. The leftmost column shows the number of holes 15, the second column from the left shows the diameter of the holes 15, the third column from the left shows the effect of breathing on the ECG waveform, and the rightmost column shows the effect of body movement on the ECG signal. ..

Changes in the number and size of holes are not related to the effect of respiration on the ECG waveform. Regarding the fluctuation of the baseline, the baseline is stabilized by increasing the diameter of the hole 15.

In the sample of FIG. 7, when three holes having a diameter of 3 mm are arranged vertically, the minimum width Pmin of the conductive path is 2.8 mm. When one hole with a diameter of 8 mm is formed, the minimum width Pmin of the conductive path is 3.0 mm, and a wider conductive path than when two holes 15 with a diameter of 6 mm are arranged (Pmin = 2.7 mm) can be obtained. .. When eight holes having a diameter of 2 mm are arranged evenly, the minimum width Pmin of the conductive path is 2.7 mm. When combined with the evaluation of FIG. 14, the diameter range of the hole 15 is larger than 2 mm, preferably 8 mm or less, more preferably 3 mm or more and 8 mm or less, and further preferably 5 mm or more and 6 mm or less. Since the hole 15 is not limited to a circle, the area of the hole, for example 4 mm ² or more, 50.5 mm ² or less, more preferably, 7 mm ² or more and 50.5 mm ² or less.

FIG. 15 is another evaluation result of the influence of the holes formed in the electrodes on the stability of measurement. In FIG. 4, the biosensor 100 was attached and the ECG waveform was measured all day long in a normal state. When the wearable biosensor 100 is realized, it may sweat depending on the season, and depending on the measurement, the sensor may be continuously attached to the skin for a certain period of time.

In FIG. 15, the measurement state is evaluated by changing the monitoring time and the skin condition using the sample of FIG. 7. The leftmost column is the diameter of the hole 15, and the second column from the left is the minimum width (Pmin) of the conductive path. In the rightmost column, the monitoring time is set to 1 day and 7 days, and the measurement result is evaluated when the skin condition is normal and when sweating, respectively. The double circle indicates that the measurement result is good, that is, a stable ECG waveform is obtained with low noise. The cross mark indicates a case where noise is high and a good ECG waveform cannot be obtained. The triangular mark indicates a case where noise is mixed, but an ECG waveform can be obtained relatively stably.

When the hole 15 having a diameter of 2 mm is arranged with the minimum width Pmin of the conductive path set to 2 mm, a stable ECG waveform can be obtained if the measurement is performed for one day in a normal state. However, good measurement results cannot be obtained in a sweaty state or in the measurement for 7 days.

When the hole 15 having a diameter of 3 mm is arranged with the minimum width Pmin of the conductive path set to 3 mm (for example, the arrangement of (C) in FIG. 7), it is stable if it is measured for one day in a normal state. The ECG waveform is obtained. However, in a sweaty state or in the measurement for 7 days, the measurement result becomes worse.

Holes with diameters of 2 mm and 3 mm, with one asterisk in the cross mark evaluation, peeled off at the interface between the pressure-sensitive adhesive layer and the skin 30 minutes after strong sweating, making measurement impossible. In the case of two asterisks, moisture accumulates at the interface between the dry electrode and the skin, and the contact between the pressure-sensitive adhesive and the skin is gradually impaired, indicating a large number of substances.

From the results of FIG. 15, the lower limit of the diameter of the hole 15 may be 2 mm for measurement in a normal state for about one day, but the diameter of the hole is 3 mm or more for measurement during sweating and long-term measurement. Is also desirable.

FIG. 16 shows an example of the planar shape of the electrode 10. The planar shape of the electrode 10 is not limited to a rectangle, and various shapes can be taken as long as holes 15 having an appropriate diameter are formed in the electrode 10 having an appropriate size. The electrode 10 is not limited to the rectangular, elliptical, and D-shaped electrodes shown in FIG. 16, and any shape applicable to the biosensor 100 such as a circle, a semicircle, an oval, and a polygon can be adopted.

The electrode 10 having the D-shaped or elliptical conductive layer 11 of FIG. 16 widens the area of the conductive layer 11 in accordance with the arc of the package 101 when both ends of the package 101 of the biosensor 100 are arcuate. Can be taken. In any shape, the area of the entire electrode 10 including the conductive layer 11 and the hole 15 is 80 mm ² or more, and the minimum width Pmin of the conductive path formed on the electrode 10 at the time of measurement is at least 2.7 mm. desirable.

By designing the electrode 10 in this way, the adhesion between the biological sensor 100 and the skin can be maintained, and the biological signal can be stably measured with low noise over a certain period of time.

This application is based on patent application No. 2020-059648 filed with the Japan Patent Office on March 30, 2020, and includes the entire contents thereof.

10 Electrode 10S Electrode sample 11 Conductive layer 15 Hole 20 Skin 25 Pressure-sensitive adhesive layer 100 Biosensor 101 Package 102 Front surface 103 Back surface 104 Space 110 Base material 150 Electronic components 160 Wiring 250 Pressure-sensitive adhesive tape Pmin Minimum width of conductive path

Japanese Unexamined Patent Publication No. 2012-10978

Claims

An electrode made of a polymer material
The conductive layer containing the polymer material and
One or more holes penetrating the conductive layer in the thickness direction,
Have,
The electrode has an area of 80 mm 2 or more and has an area of 80 mm 2 or more.
The distance from the end of the conductive layer to the nearest hole and the distance between adjacent holes is at least 2.7 mm.
electrode.
The electrode according to claim 1, wherein the diameter of the hole is larger than 2 mm and 8 mm or less.
The electrode according to claim 2, wherein the diameter of the hole is 3 mm or more and 8 mm or less.
Area of the holes, 7 mm 2 or more, the electrode according to any one of claims 1 to 3 is 50.5 mm 2 or less.
The electrode according to any one of claims 1 to 4, wherein the planar shape of the electrode includes a rectangle, a circle, an ellipse, an oval, a semicircle, and a D type.
The electrode according to any one of claims 1 to 5,
A pressure-sensitive adhesive layer arranged on the first surface of the electrode and
Electronic components connected to the electrodes
Biosensor with.
The biosensor has a surface to be attached to the living body and has a surface to be attached to the living body.
The biosensor according to claim 6, wherein the second surface of the electrode opposite to the first surface is exposed on the sticking surface.