JP7298881B2 - Myoelectric sensor and electrode member - Google Patents

Description

The present invention relates to an electrode member using a polymer material and a myoelectric sensor using the same.

Myoelectric prostheses are useful tools for assisting and supporting disabled persons who have lost a part of their limbs such as hands, feet, fingers, etc., and various researches and improvements have been made. The motion of the myoelectric prosthesis is controlled using an electromyography (EMG) signal, which is a type of biosignal. The EMG signal is a signal representing the level of muscle contraction, and in a myoelectric prosthesis that is expected to be worn for a long time, it is desirable to non-invasively measure or detect the EMG signal from the surface of the skin.

When a dry electrode is used as a sensor for detecting EMG signals, the metal of the electrode is exposed and the electrode is brought into contact with the skin to measure myoelectric potential. Dry electrodes have poor flexibility and low biocompatibility. In order to stably measure EMG signals from the skin surface, wet electrodes using conductive gel or paste are sometimes used. Further, an electrode sheet formed of a flexible conductive layer containing a carbon filler and a sensor using the electrode sheet have been proposed (see, for example, Patent Document 1).

JP 2016-27137 A

The use of wet electrodes using conductive gel or the like poses the problem of high running costs for consumables. Therefore, a hybrid electrode that uses a conductive polymer material with high biocompatibility as an electrode material and partially exposes the metal is conceivable, but the metal part in direct contact with the skin stably detects EMG signals is difficult. Even when the electrodes are made of polymer materials, the case parts other than the electrodes are generally made of highly synthetic materials, and the electrode material fixed to the case causes uneven pressure applied from the electrodes to the skin. There are problems such as loss of flexibility and leaving marks on the skin when used for a long time.

An object of the present invention is to provide an electrode material capable of non-invasively and stably measuring a myoelectric potential signal, and a myoelectric sensor using the same.

For the above purpose, layers of conductive polymers having different conductivities are laminated and used as electrodes.

In one aspect of the present invention, the myoelectric sensor comprises:
A first polymeric material layer containing a first blended amount of nanocarbon material and a second polymeric material layer containing a second blended amount of nanocarbon material less than the first blended amount are laminated. an electrode member;
a metal wire at least partially in contact with the first polymeric material layer;
have

In another aspect of the present invention, an electrode member for a myoelectric sensor comprises:
a first polymeric material layer containing a first compounded amount of nanocarbon material;
and a second polymer material layer containing a nanocarbon material in a second compounding amount less than the first compounding amount, and are continuously laminated,
The second polymeric material layer is the layer that contacts the skin of the subject.

In a good configuration example, the conductive material content of the polymeric material layer in direct skin contact is optimized.

With the above configuration, myoelectric potential signals can be measured non-invasively and stably. Moreover, since it can be used repeatedly, the running cost can be reduced.

It is a schematic diagram of a myoelectric sensor using the electrode member of the embodiment. It is an image of the produced myoelectric sensor. 4 is a diagram showing the relationship between the carbon black concentration (%) of the second polymeric material layer 12 and the maximum average myoelectric potential (V) obtained. FIG. It is myoelectric data at the time of non-pressurization (equally fixed), and is a diagram showing myoelectric potential at grip strength, myoelectric potential at rest, and SN ratio. It is myoelectric data at the time of pressurization (one-sided contact), and is a diagram showing myoelectric potential at grip strength, myoelectric potential at rest, and SN ratio. FIG. 10 is a diagram showing variations in myoelectric value when no pressure is applied and when pressure is applied; The electrical properties between the electrode and skin at each carbon black concentration are shown. FIG. 4 is a diagram showing variations or differences in electrical characteristics between non-pressurized and pressurized states; FIG. 10 is a diagram for evaluating the correlation between the robustness of myoelectric measurement and electrical characteristics; FIG. 3 is a schematic diagram of a configuration model for measuring the resistance of an electrode; FIG. 11 is a schematic diagram showing a resistance component of each model in FIG. 10; 11 is a diagram showing measurement results of each model in FIG. 10; FIG. A model for considering the relationship of resistance between layers based on the results of FIG. 12 .

FIG. 1 is a schematic diagram of a myoelectric sensor 20 using the electrode member 10 of the embodiment. The myoelectric sensor has an electrode member 10 made of a conductive polymer material and a metal wiring 14 electrically connected to the electrode member 10 to take out an electric signal. The metal wiring 14 is connected to an electric circuit chip 15, for example.

The electrode member 10 is formed by stacking a first polymeric material layer 11 and a second polymeric material layer 12 having different contents of conductive substances. The first polymeric material layer 11 is in contact with the metal wiring 14 and has a higher conductive substance content than the second polymeric material layer 12 . Assuming that the content of the conductive substance in the first polymeric material layer 11 is X% and the content of the conductive substance in the second polymeric material layer 12 is Y%, the relationship Y<X is established. Conductive substances are, for example, nano-carbon materials such as carbon black, carbon nanoparticles, graphite powder, carbon nanotubes.

The second polymeric material layer 12 is overlapped with the first polymeric material layer 11 so that the metal wiring 14 is not exposed, and the content of the nanocarbon material is less than that of the first polymeric material layer 11 . including. In use, the second polymeric material layer 12 is the contact layer that contacts the skin.

Polyurethane resin, epoxy resin, silicone resin, polyisopyrene, polybutadiene, and other elastomers can be used as the polymer material for the first polymer material layer 11 and the second polymer material layer 12 . By setting the content of the nano-carbon material contained in the second polymer material layer 12 to a content in a predetermined range that is less than the content of the nano-carbon material in the first polymer material layer 11, it will be described later. Thus, a large myoelectric potential can be stably measured regardless of the pressing state of the electrode member 10 against the skin.

As an example, the first polymeric material layer 11 is a silicone resin compounded with 4-6% carbon black (X=4-6%). The second polymeric material layer 12 is a silicone resin containing 2.0 to 2.1% or 2.6 to 2.7% carbon black (Y = 2.0 to 2.1% or 2.6-2.7%).

The metal wiring 14 is a gold (Au) wiring, a low-resistance conductive wire such as silver (Ag) or copper (Cu), or a conductive material such as gold (Au), platinum (Pt), rhodium (Rh), or the like. The wiring is plated with The metal wiring 14 is in contact with the first polymeric material layer 11 and inputs the EMG signal detected from the skin surface by the electrode member 10 to the electric circuit chip 15 .

By setting the blending amount of carbon black in the first polymer material layer 11 to 4% or more, the electrical resistance between the first polymer material layer 11 and the metal wiring 14 (that is, a good conductor such as Au, Cu, etc.) becomes close to zero. On the other hand, if the carbon black content exceeds 6%, the silicone will solidify, making it difficult to mold an electrode layer in which the carbon black is evenly dispersed.

The electrical circuit chip 15 filters and amplifies the EMG signal and outputs the amplified electrical signal to an external data logger. The data logger may be omitted and the amplified EMG signal directly supplied to a calculator such as a personal computer (PC) with built-in memory and A/D converter.

In the configuration example of FIG. 1, the electric circuit chip 15 is adhered to the electrode member 10 by the conductive nonwoven fabric 13, but the electric wiring 14 may be led out and connected to an external electric circuit. The conductive non-woven fabric 13 is made by combining a polymer non-woven fabric such as polyester or acrylic with a metal such as Ni, Cu or an alloy thereof.

The entire myoelectric sensor 20 may be sealed with silicone, except for the signal line (not shown) drawn out from the electric circuit chip 15 and the back surface of the second polymer material layer 12 (the surface in contact with the skin). . The sealing silicone functions as a casing for the electrode member 10 and the electric circuit chip 15 . By using a flexible silicone casing, the flexibility of the electrode member 10 can be maintained and it is possible to prevent marks from being left on the skin even after long-term use.

Such a myoelectric sensor 20 can be produced, for example, by the following procedure. First, a silicone resin containing 4 to 6% carbon black is poured into a mold, defoamed and baked to form the first polymeric material layer 11 . The silicone resin containing 4-6% carbon black is sufficiently stirred, and the carbon black is dispersed almost uniformly in the silicone.

A conductive nonwoven fabric 13 is adhered to one side (for example, the surface) of the baked first polymeric material layer 11 . From the opposite side (for example, the back surface) of the first polymeric material layer 11, a U-shaped metal wiring 14 is pierced to penetrate the first polymeric material layer 11 and the conductive nonwoven fabric 13, The bottom surface of the U-shaped metal wiring 14 is brought into contact with the first polymeric material layer 11 . At least part of the bottom surface of the metal wiring 14 may be embedded in the first polymeric material layer 11 . A second polymeric material layer 12 containing a different amount of carbon black is formed on the rear surface of the first polymeric material layer 11 while the metal wiring 14 is in contact with the first polymeric material layer 11 . The electric circuit chip 15 is adhered to the conductive nonwoven fabric 13 to electrically connect the metal wiring 14 and the electric circuit chip 15 . As a result, the myoelectric sensor can be produced by a simple method at low cost.

FIG. 2 is an image of the myoelectric sensor 20 casing with the sealing resin 16 such as silicone as viewed from the back side (opposite side to the electric circuit chip 15). In this example, three electrode members 10-1 to 10-3 are arranged, and signal lines 17 are led out from each of the electrode members 10-1 to 10-3. The sensor as a whole is encapsulated by a sealing resin 16 with the signal line 17 and the rear surfaces of the electrode members 10-1 to 10-3 exposed (surfaces in contact with the skin). Since both the sealing resin 16 serving as the casing and the back surfaces of the electrode members 10-1 to 10-3 are made of a flexible polymer material layer, the user does not feel uncomfortable even if the myoelectric sensor 20 is kept in close contact with the skin for a long period of time. without giving

Electrode members 10-1 and 10-2 are a differential electrode pair. The central electrode 10-3 is a ground electrode that takes a reference potential. A pair of electrode members 10-1 and 10-2 are arranged along the direction of the muscle fibers of the site to be measured, and the potential difference between the two electrodes is detected as a surface myoelectric potential signal. When one electric circuit chip 15 is commonly used for the three electrode members 10-1 to 10-3, the electric circuit chip 15 has, for example, a differential amplifier circuit together with a filter circuit. The difference between 10-2 and the ground potential of the body surface may be amplified and output as an EMG signal.

Next, the optimum range of carbon black concentration contained in the second polymeric material layer 12 will be described.

FIG. 3 is a diagram showing the relationship between the carbon black concentration (%) of the second polymeric material layer 12 and the obtained maximum average myoelectric potential (V). The inventors found that by using the electrode member 10 having the laminated structure of FIG. It was confirmed by experiments that

<Demonstration experiment 1>
A plurality of electrode samples are prepared by changing the carbon black blending amount of the second polymer material layer 12, and the myoelectric potential is measured. As experimental conditions, Au wiring is used for the metal wiring 14, and the first polymer material layer 11 in contact with the metal wiring 14 is made of silicone resin. The concentration of carbon black blended in the first polymer material layer is fixed at 4%, and the carbon black concentration in the second polymer material layer 12 is varied from 1.7% to 4.0%. FIG. 3 shows the measurement results of myoelectric potential from 1.7% to 2.7% of them.

Eleven types of electrode samples (electrode pairs) were prepared by changing the carbon black concentration of the second polymeric material layer 12 from 1.7 to 2.7% in increments of 0.1%. Wipe the electrodes and the subject's skin with alcohol prior to measurement, and secure the electrode sample to the skin with cloth tape. The measurement point is just above the flexor muscle of the forearm skin. A pair of electrodes are placed along the flexor muscle at intervals of 2 cm, and the back surface of the second polymeric material layer 12 is brought into contact with the skin and fixed with a force of 5N. A wet electrode is applied to the elbow as a body ground.

A test subject (one person) performs a grip motion of 10 kg for 3 seconds with a grip dynamometer, and myoelectric potential is measured from the corresponding muscle at a sampling frequency of 2000 Hz. The root mean square (RMS) of 10 measured myoelectric data is calculated, and the maximum value of RMS is calculated as representative data from 200 data. A similar measurement is performed five times for each electrode sample, and the maximum average myoelectric potential value for five trials is obtained for one electrode sample.

In FIG. 3, data 1 to data 5 show data for five trials, and the thick solid line is the average value (Ave). The data of this subject shows a double peak in which the myoelectric potential increases when the carbon black concentration of the second polymeric material layer 12 is 2.1% and 2.6%. That is, the maximum average myoelectric potential shows a low value up to a carbon black concentration of 1.7 to 2.0%, whereas the maximum average myoelectric potential increases at a carbon black concentration of 2.1%, and then decreases slightly. , 2.6%.
<Demonstration experiment 2>
Next, in addition to electrode samples with carbon concentrations increased in increments of 0.1% from 1.7% to 2.7%, three carbon concentrations of 3.0%, 3.5%, and 4.0% Additional samples of the kind are used. Three electrode samples are made at each carbon black concentration. Myoelectric potential during grip strength and myoelectric potential at rest are measured for all three electrode samples prepared for each carbon black concentration, and the SN ratio is calculated. The myoelectricity during grip strength is the myoelectricity when a gripping motion of 10 kg is performed with a grip strength meter, and the myoelectricity at rest is the myoelectricity when the grip is released and the force is relaxed. The myoelectric potential measured in a resting state is a potential captured by the sensor even when no force is applied, and can be regarded as background noise. The SN ratio is represented by the ratio of myoelectric potential during grip strength to myoelectric potential at rest (more specifically, the square of the myoelectric potential value obtained by dividing the myoelectric potential value during grip strength by the myoelectric potential value at rest is used). .

For each of the three electrode samples, when a pair of electrodes were fixed to the skin with equal pressure (this is called "unpressurized") and when the pressure was biased to one side of the electrode pair. It is done both when it is being pressed (this is called "pressurization" or "one-sided contact"). The uneven contact state is a state in which 20 N is applied to the entire one electrode by a force tester.

FIG. 4 shows myoelectric data when no pressure is applied (fixed evenly), showing myoelectric potential (a) at grip strength, myoelectric potential (b) at rest, and SN ratio (c). Three error bars accompanying each bar indicate the standard error of triplicate samples.

FIG. 5 shows myoelectric data at the time of pressurization (one-sided contact), showing myoelectric potential at grip strength (a), myoelectric potential at rest (b), and SN ratio (c). The three error bars accompanying each bar indicate the error for each of the triplicate samples.

In FIGS. 4 and 5, outliers are removed from the analysis targets. For each carbon black concentration, calculate the mean and standard error of 15 data for a total of 5 trials of triplicate electrode samples. In each data, data outside the range of mean ± 2 standard deviations (2σ) are defined as outliers and excluded from analysis.

Referring to the myoelectric data at the time of non-pressurization (at the time of uniform fixation) in FIG. 4, the myoelectric data at the time of grip strength (a) shows a single peak characteristic with a peak at a carbon concentration of 2.1%. However, it is dangerous to perform statistical comparison only from the myoelectric data shown in FIG.

At rest (b) in FIG. 4, the subject's forearm with the electrodes has little movement and no noticeable features. It can be seen that a constant potential mixes in as noise regardless of the carbon black concentration of the second polymeric material layer 12 . In the SN ratio (c) of FIG. 4, the myoelectric value is high at 2.1%, 2.6%, and 4.0%. Of these, 4.0% have large manufacturing errors between electrodes. Even if the carbon black concentration is increased to 3.0% to 3.5%, a sufficiently high SN ratio cannot be obtained. because it is amplified.

Turning to FIG. 5, during pressurization (one-sided contact), the myoelectric value at the time of grip strength (a) tends to increase, but there is a large variation within the electrodes, and there is also a large variation between electrodes. . In order to stably measure myoelectricity regardless of the fixation state of the electrodes, it is desirable that the variation between pressure (partial contact) and no pressure (equal fixation) is small.

FIG. 6 is a diagram showing the amount of variation when no pressure is applied and when pressure is applied. (a) shows myoelectric difference, (b) shows SN ratio difference, and (c) shows robustness. The left bar at each density of the difference (a) of the myoelectric potential is the difference in the myoelectric potential value at grip strength, and the right bar is the difference in the myoelectric potential value at rest. The difference in myoelectric potential between the even fixation state and the one-sided contact state shows a double peak with a minimal value at 2.0-2.1% and another minimal value at 2.6-2.7%. .

The difference (b) of the SN ratio also shows a double peak that takes a minimum value at 2.0 to 2.1% and again at 2.6 to 2.7%. In addition, it can be seen that the difference among the three electrode samples is also small in the range of carbon black concentration that takes the minimum value.

In order to obtain the final stability of the electromyograph speed, the value obtained by dividing the value of the SN ratio at grip strength when no pressure is applied by the amount of change in the SN ratio is defined as robustness and evaluated. Robustness shows a high value as the SN ratio at grip strength when no pressure is applied is large and the difference between the SN ratios when pressure is not applied and when pressure is not applied is small. From the result of robustness (c) in FIG. 6, peaks also appear at 2.0 to 2.1% and 2.6 to 2.7%. On the other hand, at a carbon black concentration of 3.5 to 4.0%, sufficient robustness is not ensured, and the measured EMG value varies greatly depending on how the electrode is pressed against the skin, making the measurement unstable. Become.

<Analysis of electrical characteristics>
FIG. 7 shows the electrical properties between the electrode and skin at each carbon black concentration. In FIG. 7, (a) is the impedance when no pressure is applied and when pressure is applied, (b) is the resistance when pressure is not applied and when pressure is applied, and (c) is the capacitance when pressure is not applied and when pressure is applied. For each electrode sample with different carbon black concentrations, triplicate samples were prepared, and the average value of the data obtained from five trials was taken.

In the measurement, after acquiring the data during pressurization (one-sided contact), the data during non-pressurization is acquired. The data in which the impedance (a) and the resistance (b) have high values over almost the entire carbon black concentration are the data at the time of non-pressurization (equally fixed).

Focusing on the average values of impedance and resistance, both when no pressure is applied and when pressure is applied, the carbon black concentration shows a minimum value of 2.6%. On the other hand, the capacitance in (c) shows large values at 2.0% and 2.6% in the non-pressurized characteristic, and shows a double peak as in the myoelectric characteristic.

FIG. 8 is a diagram showing fluctuations or differences in electrical characteristics between when no pressure is applied and when pressure is applied. (a) is the difference in impedance, (b) is the difference in resistance, and (c) is the difference in capacitance. As for the impedance difference (a), the carbon black concentration is 2.6%, and the difference between the non-pressurized state and the pressurized state is the smallest. As for the resistance difference (b), the difference between the non-pressurized state and the pressurized state is the smallest at the carbon black concentration of 2.2%, and the difference is relatively small even at 2.0% and 2.6%.

The difference in capacitance (c) is the smallest at a carbon black concentration of 1.9% when not pressurized and when pressurized, and the difference is small at 2.0% and 2.6 to 2.7%.

FIG. 9 is a diagram for evaluating the correlation between the robustness in (c) of FIG. 6 and the electrical characteristics. The horizontal axis of FIG. 9A indicates identification numbers of various electrical characteristics, and the vertical axis indicates the interlayer. (B) of FIG. 9 shows the electrical characteristics corresponding to the identification numbers of (A).

Identification numbers 1 to 3 indicate the impedance average, capacitance average, and resistance average when no pressure is applied (when uniformly fixed), respectively. Identification numbers 4 to 6 indicate the impedance average, capacitance average, and resistance average during pressurization (per piece), respectively. Identification numbers 7 to 9 indicate impedance changes, capacitance changes, and resistance changes when no pressure is applied and when pressure is applied, respectively.

It can be seen that the capacitance, resistance, and impedance between the skin affects the robustness of myoelectric properties, and changes in capacitance, resistance, and impedance correlate with the robustness of electromyography. As a result, in the above-mentioned demonstration experiment, double peaks were observed at carbon black concentrations of 2.0 to 2.1% and 2.6 to 2.7% with respect to robustness, and the resistance, impedance and capacitance in FIG. This can also be explained by the fact that a similar trend is observed in

As described above, by setting the carbon black concentration of the second polymer material layer 12 of the electrode member 10 to an appropriate range from the relationship between the impedance and/or capacitance between the myoelectric sensor 20 and the skin, the electrode member A stable EMG signal can be measured regardless of the fixing state of 10 to the skin.

Robustness in the fixed state (pressure state) of the electrode member 10 correlates with impedance and/or capacitance, and the robustness has two peaks as a function of carbon black concentration. It is presumed that this double peak characteristic appears even when the polymer material of the polymer material layers 11 and 12 and the type of nanocarbon material added are changed. For example, even when carbon powder is dispersed in polyurethane, polyisopyrene, or the like to form laminates with different concentrations, at least one of impedance and capacitance with the skin is presumed to exhibit two peaks.

A laminate is formed by stacking two polymer material layers with different amounts of conductive material, the layer with the smaller amount is used as the adhesive layer with the skin, and the layer with the higher amount is provided with metal wiring. By bringing the electrode into contact with the electrode and extracting the myoelectric potential, it is possible to stably measure the electromyographic potential regardless of the state of fixation of the electrode to the measurement site.

FIG. 10 is a schematic diagram of a model for measuring the resistance of the electrode itself. 7-9, the electrical characteristic values between the subject's skin and the sensor 20 were measured and discussed. From FIG. 10 onward, the electrical characteristics of the electrode itself are examined, and the usefulness of the electrode having a laminated structure is shown quantitatively.

In FIG. 10, three models (a) to (c) are produced. In model (a), a silicone electrode 21 having a carbon concentration of 2% is placed on a conductive nonwoven fabric 13 and sandwiched between metals 31 and 32 . A measuring probe 33 is pressed against the metal 31 and the metal 32 to measure the resistance value. At that time, a weight of 100 g is placed on the uppermost portion in the stacking direction to apply a load.

Model (c) has the same laminated structure as model (a), but uses electrodes 22 made of silicone having a carbon concentration of 4%. A conductive nonwoven fabric 13 and an electrode 22 are sandwiched between metals 31 and 32, and a measurement probe 33 is pressed against them to measure the resistance value. A load is applied by placing a weight of 100 on top in the stacking direction.

In model (b), the laminated structure of the embodiment described above is used. A silicone electrode 22 with a carbon concentration of 4% and a silicone electrode 21 with a carbon concentration of 2% are laminated on the conductive nonwoven fabric 13 . In order to measure the resistance value, the structure is vertically symmetrical along the stacking direction. That is, two electrodes 22 with a carbon concentration of 4% sandwich an electrode 21 with a carbon concentration of 2%, and a measuring probe 33 is pressed against the outside of the electrodes 22 with a conductive nonwoven fabric 13 interposed therebetween.

In each of the models (a) to (c), measurement is performed for 300 seconds and data is acquired every 30 seconds. This process is treated as 1 trial, and a total of 4 trials are performed for each model.

In model (a) and model (c), measurement is performed using separately produced electrodes 21 or 22 for each treatment. In model (b), measurement is performed by changing the combination of the electrodes 21 and 22 for each operation. This includes an assessment of manufacturing tolerances. For data analysis, measurement data at 300 seconds, which is a steady state, is used.

FIG. 11 is a diagram showing respective resistance value components of models (a) to (c) in FIG. The resistance values measured in (a) of FIG. It is equivalent to the total value of the contact resistance between the silicone constituting the electrode 21 and the other metal.

Similarly, the resistance values measured in FIG. 3) Equivalent to the total value of the contact resistance between the silicone constituting the electrode 22 and the other metal.

The resistance values measured in (b) of FIG. 11 are (1) the contact resistance between the silicone constituting the electrode 22 and one metal, (2) the resistance of the main body of the electrode 22 containing 4% carbon, and (3) the carbon (4) the resistance of the body of the electrode 21 containing 2% carbon; (5) the silicone containing 2% carbon (electrode 21 ) and silicone containing 4% carbon (electrode 22), (6) the resistance of the body of the electrode 22 containing 4% carbon, and (7) the contact resistance between the silicone constituting the electrode 22 and one of the metals. Equivalent to sum.

In all of FIGS. 11(a)-(c), the contact resistance between the conductive nonwoven fabric 13 and the metal 31 and the contact resistance between the conductive nonwoven fabric 13 and the metal 32 are assumed to be zero.

FIG. 12 shows measurement results of models (a) to (c) in FIG. Each vertical bar indicates the mean value of resistance, and the error bars indicate the standard error. The asterisk indicates a significant difference (or significance probability p) at the 1% significance level as a result of one-way ANOVA. Since we are comparing the means of more than two samples here, we are doing a one-way ANOVA. The significance level is set at 1% as a criterion to show that the difference in statistics obtained from the three models (a)-(c) is not just an error.

As a result of the measurement, the model (a) with a carbon concentration of 2% had an average resistance value of 1883.6 Ω, whereas the model (b) in which a silicone with a carbon concentration of 2% and a silicone with a carbon concentration of 4% were laminated , the average resistance value is reduced to 924.4Ω. In model (c) with 4% carbon concentration, the average resistance is further reduced to 221.6Ω.

As a result of one-way analysis of variance, the significance probability p is smaller than the significance level of 1% (p<0.01), indicating that the difference in resistance values between the models is not just an error, but is significantly different. The resistance value of the structure sandwiched between the electrodes 22 with a carbon concentration of 4% (model (b)) is significantly smaller than the resistance value of the electrode 21 alone with a carbon concentration of 2% (model (a)), and the average value is approximately has been reduced by half.

From the experimental results of FIGS. 4 to 9, it has been found that it is desirable that the resistance value of the part in contact with the skin should have a somewhat large value in order to ensure robustness against pressure fluctuations (one-sided contact). The amount of carbon black added to the second polymeric material layer 12 that comes into contact with the skin does not need to be as high as possible. % to 2.1%) to give a certain resistance value. As a result, myoelectric signals are prevented from being buried in noise (myoelectric potential at rest), and a high S/N ratio and sensing stability are obtained.

On the other hand, it is desirable that the resistance of the junction between the silicone electrode and the metal wiring 14 (or amplifier) is low. The experiments in FIGS. 10 to 12 show that the resistance value is lower when the electrode 21 with a carbon concentration of 4% is used than when the electrode 21 with a carbon concentration of 2% is used alone as an electrode member. . In addition, if the electrode 22 with a carbon concentration of 4% is used alone as an electrode member, noise (myoelectric potential at rest) is also amplified, reducing the S/N ratio and making it impossible to ensure robustness against uneven contact. As mentioned above. Thus, it is confirmed that the electrode formed of a polymer material has a laminated structure with different concentrations of carbon materials.

Next, a quantitative evaluation is performed. Since there are few conditions obtained for the number of types (variables) of resistance components shown in (a) to (c) of FIG. 11, it is difficult to uniquely obtain the value of each resistance component. One of the reasons for this is that a metal probe must be used to measure the resistance value of conductive silicone, and contact resistance between conductive silicone and metal occurs. Here, the relative values of the contact resistance between conductive silicones and the contact resistance between conductive silicone and metal are compared.

In order to eliminate common terms from models (a) to (c) in FIG.
(a) + (c) x 2 - (b) (1)
to calculate The reason for doubling model (c) is that model (b) contains two sets of the resistance of the C4% silicone itself and the contact resistance between the metal and the C4% silicone.

FIG. 13 shows the result of eliminating common terms in (a)+(c)×2−(b). Eliminating the common term,
2×[(C2%-M contact resistance) + (C4%-M contact resistance) - (C2%-C4% contact resistance)]
becomes. Here, "C2%" is silicone containing 2% carbon, "C4%" is silicone containing 4% carbon, and "M" is metal. In FIG. 13, the constant "2" is omitted.

Using the measurement results of FIG. 12, the resistance value of model (a) of 1883.6 Ω, the resistance value of model (c) of 221.6 Ω, and the resistance value of model (b) of 924.4 Ω are substituted into equation (1). Then, the values in FIG. 13 are
(1883.6+221.6×2-924.4)/2=701.2Ω (2)
becomes.

From the measurement results (total value of series resistance) in FIG. 12, the contact resistance between C2% silicone and metal in the first term of FIG. contact resistance is about 100Ω at maximum.

From equation (2), the total value of the model in FIG. 13 is about 700 Ω, so the sum of the first and second terms is in the range of 700 Ω to 1000 Ω, and the contact resistance of C2% silicone and C4% silicone in the third term. is assumed to be 0Ω to 300Ω.

From FIG. 12, assuming that the resistance value ratio between C2% silicone and C4% silicone is about 10:1, it is estimated that the contact resistance between C2% silicone and metal is 630Ω to 920Ω.

From the above discussion, it can be seen that the contact resistance value between conductive silicones is smaller than the contact resistance value between C2% silicone and metal. Further, as described above, the maximum contact resistance between C4% silicone and metal is about 100Ω, which is smaller than the contact resistance between C2% silicone and metal. This indicates the usefulness of laminating a conductive silicone with a high carbon concentration on a conductive silicone with a low carbon concentration.

If this is applied to the configuration of FIG. 1, the contact resistance between the first polymeric material layer 11 and the second polymeric material layer 12 is equal to and the sum of the contact resistances between the metal wiring 14 and the first polymeric material layer 11 . Also, the contact resistance between the first polymeric material layer 11 and the second polymeric material layer 12 is smaller than the contact resistance between the metal wiring 14 and the second polymeric material layer 12 .

In the configuration of FIG. 1, the metal wiring 14 does not necessarily have to be in contact with the second polymeric material layer 12, and may be embedded in the first polymeric material layer 11. FIG. In this case, the contact resistance between the first polymeric material layer 11 and the second polymeric material layer 12 is equal to or several times the contact resistance between the first polymeric material layer 11 and the metal wiring 14 . The difference is only about 100Ω. Therefore, the myoelectric signal sensed at a high S/N ratio using the relatively large resistance value of the second polymeric material layer 12 is efficiently transmitted through the first polymeric material layer 11 having a low contact resistance. can be amplified exponentially.

Finally, it is believed that increasing the thickness of the conductive silicone while maintaining the same contact surface with the metal will increase the resistance. If this linear or nonlinear increasing function can be estimated, the offset component can be identified as twice the contact resistance value between the conductive silicone and the metal. It is believed that this allows the values of all resistive elements to be identified.

REFERENCE SIGNS LIST 10 electrode member 11 first polymeric material layer 12 second polymeric material layer 13 conductive nonwoven fabric 14 metal wiring 15 electric circuit chip 20 myoelectric sensor

Claims (13)

A first polymeric material layer containing a first blended amount of nanocarbon material and a second polymeric material layer containing a second blended amount of nanocarbon material less than the first blended amount are laminated. an electrode member;
a metal wire at least partially in contact with the first polymeric material layer;
has
The myoelectric sensor, wherein the second polymeric material layer is a layer that contacts the skin of the person to be measured, and has a resistance value that prevents myoelectric signals from being buried in noise. A first polymeric material layer containing a first blended amount of nanocarbon material and a second polymeric material layer containing a second blended amount of nanocarbon material less than the first blended amount are laminated. an electrode member;
a metal wire at least partially in contact with the first polymeric material layer;
has
The second polymeric material layer is a layer that contacts the skin of the subject,
The myoelectric sensor, wherein the electrode member has two peaks in capacitance between the electrode member and the skin as a function of the second compounding amount. 3. The myoelectric sensor according to claim 1 , wherein the metal wiring is not in contact with the second polymeric material layer. A first polymeric material layer containing a first blended amount of nanocarbon material and a second polymeric material layer containing a second blended amount of nanocarbon material less than the first blended amount are laminated. an electrode member;
a metal wire at least partially in contact with the first polymeric material layer;
and the metal wiring is in contact with the second polymeric material layer, the contact resistance between the first polymeric material layer and the second polymeric material layer is the second polymeric material layer A myoelectric sensor characterized by being smaller than the sum of the contact resistance between the layer and the metal wiring and the contact resistance between the first polymeric material layer and the metal wiring. A first polymeric material layer containing a first blended amount of nanocarbon material and a second polymeric material layer containing a second blended amount of nanocarbon material less than the first blended amount are laminated. an electrode member;
a metal wire at least partially in contact with the first polymeric material layer;
and the metal wiring is in contact with the second polymeric material layer, the contact resistance between the first polymeric material layer and the second polymeric material layer is the second polymeric material layer A myoelectric sensor characterized in that the contact resistance between the layer and the metal wiring is smaller than the contact resistance. A first polymeric material layer containing a first blended amount of nanocarbon material and a second polymeric material layer containing a second blended amount of nanocarbon material less than the first blended amount are laminated. an electrode member;
a metal wire at least partially in contact with the first polymeric material layer;
an electric circuit chip electrically connected to the metal wiring;
a conductive nonwoven fabric disposed between the first polymeric material layer and the electrical circuit chip ;
has
The myoelectric sensor, wherein the metal wiring penetrates the conductive nonwoven fabric from the first polymer material layer and is connected to the electric circuit chip. Claims 1 to 6 , wherein the nanocarbon material is carbon black, and the second compounding amount is either 2.0 to 2.1% or 2.6 to 2.7%. The myoelectric sensor according to any one of . 8. The myoelectric sensor according to claim 7 , wherein the first blending amount of the carbon black is 4 to 6%. The first polymer material layer and the second polymer material layer are selected from silicone resin, polyisopyrene, polybutadiene, and other elastomers according to any one of claims 1 to 8 . EMG sensor. An electrode member for a myoelectric sensor,
a first polymeric material layer containing a first compounded amount of nanocarbon material;
and a second polymer material layer containing a nanocarbon material in a second compounding amount less than the first compounding amount, and are continuously laminated,
The second polymeric material layer is a contact layer with the skin of the person being measured, and has a resistance value that prevents myoelectric signals from being buried in noise.
Electrode member. An electrode member for a myoelectric sensor,
a first polymeric material layer containing a first compounded amount of nanocarbon material;
and a second polymer material layer containing a nanocarbon material in a second compounding amount less than the first compounding amount, and are continuously laminated,
The second polymer material layer is a layer that contacts the skin of the person being measured,
An electrode member having two peaks in capacitance with said skin as a function of said second compounding amount. Claim 10 or 11, wherein the nanocarbon material is carbon black, and the second compounding amount is either 2.0 to 2.1% or 2.6 to 2.7%. The electrode member according to . 13. The electrode member according to claim 12 , wherein the first blending amount of the carbon black is 4 to 6%.

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