WO2024248179A1 - Electrode-type electromyogram sensor and manufacturing method therefor - Google Patents

Abstract

The present disclosure provides a method for manufacturing an electromyogram sensor, the method comprising the steps of: preparing a conductive ink by mixing 10-25 wt% of an elastic base material, 30-50 wt% of a conductive filler, and 30-50 wt% of a hydrophobic solvent; applying the conductive ink to a substrate and performing spin coating at a predetermined speed or less; sintering the conductive ink for a predetermined time or more to remove the hydrophobic solvent and generate a conductive film; and generating a plurality of throughholes in the conductive film by using a micro-patterning mold. Therefore, the electrode-type electromyogram sensor provides stretchability (100% or more) and reliable electrical conductivity despite stretching, and exhibits low skin contact resistance and moisture permeability and a stable signal-to-noise ratio (SNR) even when used for a long period of time, thereby increasing user convenience and providing improved quality.

Description

Electrode-type electromyography sensor and its manufacturing method

The present disclosure relates to an electrode type electromyography sensor and a method for manufacturing the same.

Systems for measuring bio-signals such as electromyography (EMG) are already available in a variety of products on the market. EMG is a method for measuring muscle fatigue, tension, and relaxation, and there are many reports of close relationships between changes in EMG signals according to the current subject's condition.

In order to manufacture an electromyography sensor, a separate human-friendly metal piece was attached to a specific substrate to manufacture an electromyography sensor. The electromyography sensor manufactured in this way was attached to the skin to measure biosignals. However, since the electromyography sensor includes a metal piece, it was inconvenient to attach it to the curved surface of the skin. In addition, when the skin and the electromyography sensor were attached for a long time, there was no ventilation between the metal piece of the electromyography sensor and the surface attached to the skin, so it was difficult to measure biosignals due to the sweat secreted at this time. In addition, since the electromyography sensor must be manufactured in a small size because the metal piece must be attached to the substrate, it was difficult to obtain biosignals from a specific, extremely small area.

Here, US patent US20130056689A1 discloses the application of a reducing agent that reduces the fatty acid of silver salt present on the surface of silver flakes, induces sintering, and forms a nano-unit conductive path to reduce resistance. Accordingly, the resistance of the conductive adhesive is reduced by applying such a reducing agent. However, the adhesive is mainly intended for an RF antenna, not an electromyography sensor. Accordingly, the application target is different, and no improvement steps are disclosed to secure breathability, adhesiveness, and flexibility.

Meanwhile, international patent WO2018111365A1 discloses an electrically conductive silicone adhesive (ECSA), a method for producing the same, a method for utilizing the same, and a device utilizing the same. ECSA is an active polymer silicone resin, and contains polydiorganosiloxane polymer, MQ organosiloxane resin, and silver (Ag) as a metal filler.

Such silicone adhesive does not disclose the alignment of the silver filler dispersed therein. When the silicone adhesive is utilized as an electromyography sensor electrode, if elasticity is required depending on the purpose, the conductivity may show a large difference depending on the alignment direction of the silver filler.

[Prior art literature]

US Patent US20130056689A1, Publication Date March 7, 2013

International Patent WO2018111365A1, Publication Date February 15, 2017

The present disclosure aims to improve the problems of existing medical electrode-type electromyography sensors having low elasticity and water vapor permeability and being thick, making them unsuitable for skin stretching and sweating of the wearer due to exercise.

The present disclosure proposes an electrode-type electromyography sensor having moisture permeability, thereby providing an electromyography sensor that can be reused for a long time.

The present disclosure seeks to provide an electrode-type electromyography sensor capable of maintaining conductivity and elasticity by firing the conductive filler in a state in which it is arranged to align in a specific direction.

In addition, the present disclosure provides a dry electromyography sensor by inducing alignment of fillers by mixing and then removing a solvent (133) to lower the viscosity of the polymer during the manufacturing step, and then sintering, while providing conductivity, retention, contact, and elasticity at levels similar to those of a wet sensor.

The present disclosure provides an electromyography sensor electrode comprising: an elastic matrix; and a plurality of conductive fillers dispersed within the elastic matrix and including a two-dimensional surface; wherein 70 to 80 w% of the plurality of conductive fillers are included with respect to the electromyography sensor electrode, and a horizontal plane of the electromyography sensor electrode and a two-dimensional surface of the conductive fillers are aligned to be parallel.

When the electromyography sensor electrode is elastically deformed on the horizontal plane by the elastic substrate, electrical conductivity can be maintained by the alignment of the conductive filler.

The above conductive filler may be at least one of silver (Ag) flakes, gold (Au) flakes, and copper (Cu) flakes.

The above plurality of conductive fillers may have a longitudinal length of 2-10 μm.

The above EMG sensor electrodes may have a thickness of 30-90 um.

The above electromyography sensor electrode may have multiple holes formed on its surface.

The above plurality of holes can be formed in a cylindrical or conical shape.

The above plurality of pores may have a diameter of 50 to 400 um on the surface.

The above elastic material may be a silicone elastomer or polyurethane (PU).

Meanwhile, the present disclosure provides a method for manufacturing an electromyography sensor electrode, including the steps of preparing a conductive ink by mixing 10 to 25 wt% of an elastic matrix, 30 to 50 wt% of a conductive filler, and 30 to 50 wt% of a hydrophobic solvent; applying the conductive ink to a substrate and performing spin coating at a predetermined speed or lower; sintering the conductive ink for a predetermined time or longer to remove the hydrophobic solvent and generate a conductive film; and creating a plurality of holes in the conductive film using a micro-patterning mold.

The step of manufacturing the conductive ink may include stirring the conductive ink using a magnetic stirrer for 2 hours or longer at a predetermined speed or lower to induce the conductive filler in the conductive ink to be uniformly dispersed.

The step of performing the above spin coating can apply the conductive ink to the substrate without external stress and perform low-speed spin coating for a short period of time.

By the low-speed spin coating, the conductive filler in the conductive ink can be aligned on the substrate so that a two-dimensional plane is parallel to the plane of the substrate.

The step of producing the conductive film may include heating at a temperature range of 150 to 180 degrees for 30 minutes to 2 hours to evaporate the hydrophobic solvent and harden the elastic base material while sintering the conductive filler.

The step of creating the above-mentioned through hole may include a step of manufacturing the micro-patterning mold using 3D printing; and a step of rolling the conductive film on the micro-patterning mold to form the through hole.

The above conductive filler has a longitudinal length of 2-10 μm and may be at least one of silver (Ag) flakes, gold (Au) flakes, and copper (Cu) flakes.

The above EMG sensor electrodes may have a thickness of 30-90 um.

The above plurality of pores may be cylindrical or conical and have a diameter of 50 to 400 μm on the surface.

The above elastic material may be a silicone elastomer or polyurethane (PU).

The above hydrophobic solvent has a molecular weight of 60 g/mol or more and may be hexadecane or toluene.

Through the above solution, the electrode type electromyography sensor can increase user convenience and provide improved quality by showing elasticity (100% or more), low skin contact resistance (<50 kΩ at 100 Hz based on commercial electrode area), and stable signal-to-noise ratio (SNR) even after long-term use.

In addition, the electromyography sensor of the present disclosure has excellent breathability, so that it can maintain a comfortable wearing feeling even when sweating due to exercise.

In addition, since the manufacturing cost and unit price are very low, it can be distributed to households as a low-cost model, and by applying it to an integrated measuring device with a belt marker, etc., exercise data and muscle data can be analyzed together to create a more accurate injury risk assessment index.

FIG. 1A and FIG. 1B are cross-sectional views and top views of an electrode-type electromyography sensor according to the present disclosure.

Figure 2 is a conceptual diagram illustrating an application example in which the electrode-type electromyography sensor of Figures 1a and 1bb is attached to the epidermis of a subject.

FIG. 3 is a flowchart showing a method for manufacturing an electrode-type electromyography sensor according to the present disclosure of FIGS. 1a and 1b.

Figure 4 is a drawing showing the stress-free low-speed spin process of Figure 3.

Figures 5a and 5b are drawings showing the curing and support film removal processes.

FIG. 6a is a drawing showing a breathable patterning process according to the first embodiment, and FIG. 6b is a drawing showing a breathable patterning process according to the second embodiment.

Figures 7a and 7b are drawings showing the product manufacturing process.

FIG. 8a and FIG. 8b are cross-sectional photographs showing the breathability pattern of an electrode-type electromyography sensor manufactured by the manufacturing method of FIG. 3.

Fig. 9a is a cross-sectional photograph of a flexible electrode according to a comparative example, and Fig. 9b is a photograph of an electrode-type electromyography sensor of the present disclosure.

Fig. 10a is a conceptual diagram of a stretch according to a comparative example, and Fig. 10b is a conceptual diagram of a stretch according to the present disclosure.

Fig. 11a is the front conductivity of the present disclosure and a comparative example, and Fig. 11b is the back conductivity.

Fig. 12a shows the change in resistance according to the degree of elasticity of the present disclosure and the comparative example, and Fig. 12b shows the contact impedance according to the frequency of the present disclosure and the comparative example.

Figure 13 is a photograph showing the expansion of the electrode-type electromyography sensor of the present disclosure.

Figure 14 compares the adhesion of the electrode-type electromyography sensor of the present disclosure with that of a comparative example.

The use of terms such as ‘first’, ‘second’, etc. before the components mentioned below is only intended to avoid confusion regarding the components referred to, and has nothing to do with the order, importance, or main-subordinate relationship between the components. For example, an invention can also be implemented that includes only the second component without the first component.

In the drawings, the thickness or size of each component is exaggerated, omitted, or schematically illustrated for convenience and clarity of explanation. In addition, the size and area of each component do not entirely reflect the actual size or area.

In addition, the angles and directions mentioned in the process of explaining the structure of the present disclosure are based on those described in the drawings. In the description of the structure in the specification, if the reference point and positional relationship for the angle are not clearly mentioned, refer to the relevant drawings.

FIG. 1A and FIG. 1B are cross-sectional views and top views of an electrode-type electromyography sensor according to the present disclosure.

Referring to FIGS. 1A and 1B, an electromyography sensor according to the present disclosure may include a substrate (110) and an electromyography electrode (130).

The above substrate (110) may be a flexible substrate (110), and may be manufactured and sold in various shapes. For example, it may be manufactured in the form of a sports tape, and in this case, it may be manufactured using fibers formed from a resin having elasticity, such as polyurethane.

When a plurality of electromyography electrodes (130) are placed on a substrate (110), the plurality of electromyography electrodes (130) can be placed spaced apart from each other, and can be connected to pads so that each can transmit a signal to an external controller (not shown) through a different wire (120).

In contrast, multiple electromyography electrodes (130) can be connected simultaneously and the pads can be connected in parallel to transmit signals to the controller through the same wire (120), and this can be selectively implemented depending on the function.

For connection of such pads and wires (120), a circuit portion can be formed between the substrate (110) and the electrode (130).

The above circuit part is described as being implemented in its entirety in Fig. 1a, but may be described in only a part.

The adhesion between the above-mentioned electromyography electrode (130) and the substrate (110) can be formed using a general adhesive (not shown) and is not particularly limited.

The electromyography electrodes (130) of the present disclosure are individually formed and separately attached to the substrate (110) using an adhesive.

The above electromyography electrode (130) includes an elastic material (131) and a conductive filler (132), as illustrated in FIG. 1A.

The above elastic material (131) may be a silicone elastomer such as Ecoflex ^TM , Silbione ^TM , polydimethylsiloxane, etc., but may alternatively be an elastically deformable polymer resin such as polyurethane (PU).

The above elastic material (131) may be contained in an amount of 20 to 30 w%, preferably about 25 w%, relative to the electromyography electrode (130).

The above elastic material (131) is elastically deformable and can easily come into contact with the curved surface of the skin, and has viscosity, so that skin adhesion can be improved.

A plurality of conductive fillers (132) are dispersed within the elastic substrate (131). The conductive fillers (132) are connected to the pads and wires (120) below, and when in contact with the skin, the amount of change in the electric charge generated according to the muscle activity of the skin is detected and transmitted to the circuit unit (120) below as a signal.

The above-described plurality of conductive fillers (132) each have a two-dimensional surface, and can satisfy a longitudinal length (d1) of 2-10 μm, preferably 4 to 10 μm. When the longitudinal length (d1) of the two-dimensional surface of the conductive filler (132) exceeds the upper limit of the above-described range, the mechanical properties including the Young's modulus deteriorate, and when it is less than the lower limit, the electrical properties including the electrical conductivity deteriorate.

At this time, the plurality of conductive fillers (132) may be arranged so that the two-dimensional surface is parallel to the bottom surface (110a), that is, the bottom surface (110a) of the electromyography electrode (130). That is, when the plurality of conductive fillers (132) have a two-dimensional surface, they are aligned in parallel, rather than standing at a predetermined angle with respect to the bottom surface (110a) within the elastic substrate (131).

The above-described EMG sensor undergoes elongation and expansion, i.e., elastic deformation, and the elastic deformation progresses by increasing or decreasing the bottom surface (110a) of the EMG sensor, i.e., the plane formed by the X-axis and the Y-axis. By aligning the conductive filler (132) as described above, the overlap between the two-dimensional surfaces of the conductive filler (132) can be maintained despite the elastic deformation of the EMG sensor.

Accordingly, even if elastic deformation occurs due to alignment of the conductive filler (132), the electrical conductivity of the electromyography sensor can be maintained.

To this end, the conductive filler (132) may be contained in an amount of about 70 to 80 w%, preferably about 75 w%, relative to the entire electromyography electrode (130).

The above conductive filler (132) may be, but is not limited to, silver (Ag) flakes, gold (Au) flakes, or copper (Cu) flakes.

The electromyography electrode (130) in which the conductive filler (132) is uniformly aligned and distributed may have a thickness (h1) of 30-90 um, preferably 40 to 80 um, but is not limited thereto.

The above electromyography electrode (130) has a plurality of holes (135) formed on the upper surface (130a).

The upper surface (130a) above is a surface that comes into contact with the user's skin, and the perforations (135) can be arranged uniformly as shown in Fig. 1b, but are not limited thereto.

The above-mentioned hole (135) may be formed in a circular shape so that the diameter (d2) remains constant from the surface to the bottom surface, but alternatively, it may be formed in a cone shape in which the diameter (d2) decreases as it goes toward the bottom surface.

The maximum value of the diameter (d2) of the hole (135) on the above surface may be 50 to 400 um, but is not limited thereto.

The above-mentioned perforations (135) may be arranged on the upper surface (130a) in a matrix as shown in Fig. 1b, but may be formed so that the even and odd rows are misaligned.

In this way, when the electromyography electrode (130) is attached to the user's skin as shown in FIG. 2, the upper surface (130a) is attached to the user's skin and detects and transmits charge movement according to muscle activity of the user's skin.

Since a plurality of conductive fillers (132) that form the electrical conductivity of the above-mentioned electromyography electrode (130) are arranged parallel to the bottom surface (110a) of the above-mentioned electromyography electrode (130), even if the bottom surface of the above-mentioned electromyography sensor, i.e. the plane formed by the X-axis and the Y-axis, is increased or decreased, the overlap between the two-dimensional surfaces of the conductive fillers (132) can be maintained.

In addition, moisture generated by the user's skin sweating can pass through the multiple openings (135) of the upper surface (130a) and evaporate to the outside, thereby preventing the electromyography electrode (130) from falling off due to moisture.

In addition, the EMG sensor has more than 100% elasticity, can maintain low skin contact resistance, and can maintain a stable signal-to-noise ratio (SNR) even after long-term use, thereby ensuring user convenience and quality.

The process for manufacturing such an electromyography sensor is described below with reference to FIGS. 3 to 7.

FIG. 3 is a flowchart showing a method for manufacturing an electrode (130) type electromyography sensor according to the present disclosure of FIGS. 1a and 1b, FIG. 4 is a drawing showing a stress-free low-speed spin process of FIG. 3, FIGS. 5a and 5b are drawings showing a curing and support film removal process, FIG. 6a is a drawing showing a breathable patterning process according to the first embodiment, FIG. 6b is a drawing showing a breathable patterning process according to the second embodiment, and FIGS. 7a and 7b are drawings showing a product manufacturing process.

The electromyography sensor of the present disclosure first produces conductive ink for forming electromyography electrodes (130) that constitute the electromyography sensor.

First, the conductive ink includes a conductive filler (132), an elastic matrix (131), and a hydrophobic solvent (133).

The above elastic material (131) may be a silicone elastomer such as Ecoflex ^TM , Silbione ^TM , polydimethylsiloxane, etc., but may alternatively be an elastically deformable polymer resin such as polyurethane (PU).

The elastic substrate (131) may be 10 to 25 w% of the entire conductive ink relative to the electromyography electrode (130).

The above elastic material (131) is elastically deformable and can easily come into contact with the curved surface of the skin, and has viscosity, so that skin adhesion can be improved.

The above-described plurality of conductive fillers (132) each have a two-dimensional surface, and the longitudinal length (d1) can satisfy 2 to 10 μm, preferably 4 to 10 μm. When the length (d1) of the conductive filler (132) exceeds the upper limit of the above-described range, the mechanical properties including the Young's modulus deteriorate, and when it is less than the lower limit, the electrical properties including the electrical conductivity deteriorate.

At this time, the plurality of conductive fillers (132) may be flakes of silver, gold, copper, etc., or graphene at 30-50 wt% with respect to the entire ink, but are not limited thereto.

In addition, the conductive filler (132) can be used by mixing two conductive fillers (132) having different lengths (d1).

Meanwhile, the conductive ink further contains 30-50 w% of a hydrophobic solvent (133) (S10).

The hydrophobic solvent (133) lowers the viscosity of the ink in the conductive ink to induce alignment of the conductive filler (132) and delays the curing time to facilitate processing.

The above hydrophobic solvent (133) is a nonpolar solvent (133) having a molecular weight of 60 g/mol or more, and preferably hexadecane or toluene.

If the hydrophobic solvent (133) is less than the above weight ratio, the conductive filler (132) is not aligned, the curing time is short, and the time required for long-term uniform stirring of the ink cannot be secured, and if it exceeds the upper limit, the elasticity of the manufactured electrode (130) may be reduced.

In this way, the conductive filler (132), elastic base material (131), and hydrophobic solvent (133) for the conductive ink are mixed according to the weight ratio, and then stirred at low speed using a magnetic stirrer for 2 hours or more (S20).

By this type of low-speed, long-term stirring, the conductive filler (132), for example, silver flakes, in the ink (140) is maintained in a uniformly dispersed state.

Next, conductive ink (140) is applied on the substrate (20) as shown in Fig. 4 (S30).

The step of applying conductive ink (140) on the substrate (20) is performed by applying it to the substrate (20) without any specific external stress and without any external stress on the substrate (110) at least, and preferably performing low-speed spin coating at 100 rpm for about 60 seconds.

At this time, applying to the substrate (20) without a specific stress from the outside means that the coating is performed on the entire substrate (20) by pouring without a drop in the edge area or the center area and then rotating at a low speed for a short period of time, rather than applying to a wide area through a small area by applying pressure, such as a paste or injection method or a spray method such as a spray method.

When spin coating is performed at a low speed for a short time with external stress minimized in this way, the two-dimensional plane of the conductive filler (132) is aligned parallel to the plane of the substrate (20).

That is, the conductive filler (132) is laid down parallel to the plane of the substrate (20) by the uniform distribution and dispersion of the hydrophobic solvent (133) without being affected by the centrifugal force caused by the spin.

Next, as shown in Fig. 5a, the elastic base material (131) is hardened while sintering the conductive filler (132) by heating in a temperature range of 150 to 180 degrees for 30 minutes to 2 hours to evaporate all of the hydrophobic solvent (133) (S40).

After curing as shown in Fig. 5b, the lower substrate (20) is removed to manufacture a film for the electromyography sensor electrode (130) as a single layer.

Meanwhile, in order to form a breathable hole (135) on the surface of the electromyography sensor electrode (130) as shown in Fig. 1b, a micro-patterning mold (50a, 50b) as shown in Figs. 6a and 6b can be applied (S50).

The above micro patterning mold (50a) can be manufactured using a three-dimensional printer, and is formed with a plurality of needles (51, 52) formed.

The above plurality of needles (51, 52) may be cylindrical needles (51) as shown in Fig. 6a, or may be conical needles (52) as shown in Fig. 6b.

The film for the above-described electromyography sensor electrode (130) can form a circle pattern of FIG. 6a or a dot pattern of FIG. 6b on the surface, i.e., the upper surface (130a), by pressing or rolling it over the micro-patterning mold (50a, 50b).

The above-mentioned permeable hole (135) can maintain a shape in which the diameter (d2) decreases from the upper surface (130a) to the lower surface or in a conical shape when forming a circular pattern or a dot pattern. The maximum value of the diameter (d2) on the upper surface (130a) may be 50 to 400 um, but is not limited thereto.

The above-mentioned perforations (135) may be arranged on the upper surface (130a) in a matrix as shown in Fig. 1b, but may be formed so that the even and odd rows are misaligned.

In this way, the electromyography sensor electrode (130) with the permeable hole (135) formed can be cut to a desired size and attached to a wearable device to be commercialized (S60).

For example, as shown in Fig. 7a, an electromyography sensor electrode (130) is attached to a main body or substrate (110) using an adhesive.

At this time, the main body (110) is a fiber formed from an elastic resin such as polyurethane, and can be manufactured and sold in the form of a sports tape.

A pattern (111) is printed in a band shape on the front surface facing the outside of the main body (110), and the electromyography sensor electrode (130) can be attached to a position corresponding to at least one specific part of the body on the inner surface that comes into contact with the body part, thereby manufacturing the body.

Meanwhile, as shown in Fig. 7b, it is possible to attach the wire-integrated electromyography electrode (130) by forming it on the medical tape body (110a).

The electromyography electrode (130) integrated into the medical tape body (110a) can transmit a detection signal to the pad (120a) at the end of the wire. In Fig. 7b, a magnetic cover (150) is separately attached to protect the electromyography sensor area, but is not limited thereto.

The electromyography sensor electrode (130) formed in this manner can secure electrical conductivity, reliability, and adhesiveness according to breathability and elasticity when applied to the wearable device.

Below, various characteristics are measured using an electromyography sensor electrode (130) manufactured through an experimental example of the present disclosure.

Experimental example

In the experimental example of the present disclosure, conductive ink (140) is prepared by mixing 43 wt% of silver paste, 14 wt% of Ecoflex ^TM , and 43 wt% of hexadecane.

The above silver flakes each have a two-dimensional surface, and those with a longitudinal length (d1) of 4-8 μm and those with a length of 10 μm were used in combination.

After stirring the conductive ink (140) at low speed using a magnetic stirrer for more than 2 hours, and maintaining the state in which the silver flakes in the ink (140) are uniformly dispersed by the low-speed long-term stirring, it is applied to a substrate (110) without external stress, and low-speed spin coating is performed at 100 rpm for about 60 seconds to form a thin film with a thickness of 60 μm (h1).

When spin coating is performed at a low speed for a short time with external stress minimized in this way, the two-dimensional plane of the silver flake is aligned parallel to the plane of the substrate (110).

Thereafter, the hexadecane was completely evaporated by heating in a temperature range of 150 to 180 degrees for 30 minutes to 2 hours, and the elastic matrix (131) was hardened while sintering the silver flakes to form a thin film with a thickness of 40 to 50 μm (h1).

Next, after removing the substrate (110), it was rolled on a micro patterning mold to form breathable holes (135) to form the electromyography sensor electrode (130) of FIGS. 8a to 8b according to the experimental example of the present disclosure.

FIGS. 8a and 8b are cross-sectional photographs showing the breathability pattern of an electrode (130)-type electromyography sensor manufactured using the manufacturing method of FIG. 3.

As shown in the enlarged photographs of FIGS. 8a and 8b, when observing photographs taken by an electron microscope of micro-scale breathable pores (135), it can be confirmed that the pores (135) are formed penetrating the electromyography sensor electrode (130).

Accordingly, moisture generated from the skin surface can be released to the outside through such a perforation (135), thereby improving water vapor permeability, as shown in the graph of Fig. 8c.

When measuring the water vapor permeability in the same area in the case of a circle pattern as in Fig. 8a and in the case of a dot pattern as in Fig. 6b, when the moisture permeability in the case of the entire surface being open is 30 g/hm ² , in the case of the present disclosure in which the through holes (135) are formed, the moisture permeability is 6 g/hm ² or more, preferably 10 g/hm ² or more, and particularly, in the case of the circle pattern of the experimental example, the moisture permeability is 12 g/hm ² or more.

This can provide meaningful penetration in cases where there is no pattern at all.

When moisture permeability is secured in this way, when attached to the skin and sweat is generated from the skin and moisture is generated, the moisture can be released to the outside to maintain adhesion.

Meanwhile, the electromyography sensor electrode (130) of the above experimental example has the following characteristics.

Fig. 9a is a cross-sectional photograph of a flexible electrode (130) according to the present disclosure, and Fig. 9b is a photograph of an electrode (130) type electromyography sensor of a comparative example. Fig. 10a is a conceptual diagram of a stretchable electrode according to the present disclosure, and Fig. 10b is a conceptual diagram of a stretchable electrode of a comparative example.

In the case of the electromyography sensor electrode (130) of the comparative example as shown in FIGS. 9b and 10b, the plane of the two-dimensional conductive filler (132) is not arranged parallel to the elongated plane. Therefore, the arrangement becomes more disordered as it elongates and contracts.

Meanwhile, in the experimental example of the present disclosure, as shown in FIGS. 9A and 10A, the plane of the two-dimensional conductive filler (132) is arranged mostly parallel to the elongated plane. As described above, this is implemented by dispersion by a hydrophobic solvent (133) and coating by ultra-low-speed rotation without external stress.

With the arrangement as described above, the arrangement is not disturbed according to the elongation and expansion of the electromyography electrode (130), and the upper and lower connections are continuously maintained.

Fig. 11a is the front conductivity of the present disclosure and a comparative example, and Fig. 11b is the back conductivity.

The electrical conductivity of the electromyography electrode (130) according to the experimental example of the present disclosure and the electrical conductivity of the comparative example are slightly different at the front and back, but their patterns are similar.

Figure 11a shows the front electrical conductivity, and Figure 11b shows the back electrical conductivity.

In general, it can be observed that the electrical conductivity of the front side is higher than that of the back side, whether the conductive filler (132) is aligned or not. At this time, the electrical conductivity is a comparison of the electrical conductivity of Fig. 9b, which is a comparative example in which the silver flakes are not aligned, and the electrical conductivity of Fig. 9a, which is an experimental example in which the silver flakes are aligned.

Referring to Figures 11a and 11b, it can be confirmed that the electrical conductivity of the experimental example is much higher on both the front and back skin-attached surfaces.

Figure 12a shows the change in resistance according to the degree of elasticity of the experimental example and comparative example of the present disclosure, and Figure 12b shows the contact impedance according to the frequency of the experimental example and comparative example of the present disclosure.

Referring to Figure 12a, it shows the change in resistance according to the degree of elasticity, that is, the ratio of the increased area to the initial area when the electromyography sensor electrode (130) is stretched in a plane.

It can be interpreted that the greater the change in resistance, the lower the electrical conductivity. Therefore, according to Fig. 12a, in the case of the experimental example where the silver flakes are aligned, it can be confirmed that the change in resistance is significantly small, so it can be interpreted that the change in electrical conductivity is small.

Meanwhile, Fig. 12b shows the values of Fig. 12a in the frequency domain.

At this time, Comparative Example 1 shows an electromyography sensor electrode (130) in a state where silver flakes are not aligned, as in the Comparative Example of Fig. 12a, and Comparative Example 2 shows an electromyography sensor electrode (130) of a commercially available Ag/AgCl wet electrode (130).

When examining FIG. 12b, the skin contact impedance of the experimental example of the present disclosure has a skin contact impedance at a level similar to that of a commercial wet electrode (130) when the electrode (130) area is 1.5 cm ² , and has a higher value than that of Comparative Example 1 including non-aligned silver flakes.

In addition, it has a signal-to-noise ratio similar to that of a commercial wet electrode (130), and accordingly, it can have a signal transmission efficiency similar to that of a commercial wet electrode (130).

Figure 13 is a photograph showing an extension of an electrode (130) type electromyography sensor according to an experimental example of the present disclosure.

As shown in Fig. 13, when the electromyography sensor electrode (130) of the experimental example is stretched, it has an elongation rate of 100% or more and can be stretched in any axis (x-axis or y-axis) on a plane. With such an elongation rate, attachment is possible along the curve of the skin.

Figure 14 compares the adhesion of the electrode (130) type electromyography sensor of the present disclosure with that of a comparative example.

In Fig. 14, the adhesion force of the electromyography sensor electrode (130) according to the experimental example of the present disclosure and the control example 3 was measured using a force gauge.

The measurement conditions were an electrode (130) attachment area of 2*2cm, skin humidity of 30%, and a 90-degree peeling test.

At this time, in the case of Control Example 3, it was confirmed that the electromyography sensor electrode (130) of the present disclosure had a skin adhesion strength of about 1.1 kPa, based on the commonly used medical tape having an adhesion strength of about 2.0 kPa.

In the case of this type of adhesive strength, although lower than that of general medical tape, it is an adhesive strength that induces position fixation when attached to a specific location, and can be stably attached for a long period of time, making it suitable for use as an electrode (130) of a wearable device.

In this way, the thin film-type conductive electrode (130) of the present disclosure is suitable for use in a wearable electrophysiological signal sensor, as only about 15% of resistance change is observed in the maximum elastic range of the skin, and the skin contact impedance based on a circular electrode (130) having a diameter of 2.5 cm, which is a commonly used electrode (130) area, is 50 kΩ or less (@100Hz), which is the skin contact impedance recommended for high-quality biosignals.

In addition, although the preferred embodiments of the present disclosure have been illustrated and described above, the present disclosure is not limited to the specific embodiments described above, and various modifications may be made by a person skilled in the art without departing from the gist of the present disclosure as claimed in the claims, and such modifications should not be individually understood from the technical idea or prospect of the present disclosure.

[Explanation of symbols]

110: substrate 120: wire

130: EMG sensor electrodes

Claims (20)

Elastic matrix; and A plurality of conductive fillers dispersed within the elastic matrix and comprising a two-dimensional surface; As an electromyography sensor electrode including: The above electromyography sensor electrodes include 70 to 80 w% of the above-described plurality of conductive fillers, An electromyography sensor electrode in which the horizontal plane of the electromyography sensor electrode and the two-dimensional plane of the conductive filler are aligned parallel to each other. In the first paragraph, An electromyography sensor electrode in which electrical conductivity is maintained by alignment of the conductive filler when the electromyography sensor electrode is elastically deformed on the horizontal plane by the elastic matrix. In the second paragraph, An electromyography sensor electrode wherein the conductive filler is at least one of silver (Ag) flakes, gold (Au) flakes, and copper (Cu) flakes. In the third paragraph, The above-mentioned plurality of conductive fillers are electromyography sensor electrodes having a longitudinal length of 2-10 μm. In paragraph 4, The above electromyography sensor electrode is an electromyography sensor electrode with a thickness of 30-90 um. In paragraph 5, The above electromyography sensor electrode is an electromyography sensor electrode having a plurality of holes formed on the surface. In Article 6, The above-mentioned plurality of holes are formed in a cylindrical or conical shape, and are electromyography sensor electrodes. In Article 7, The above-mentioned plurality of holes are electromyography sensor electrodes having a diameter of 50 to 400 um on the surface. In the second paragraph, The above elastic base material is an electromyography sensor electrode made of silicone elastomer or polyurethane (PU). A step of preparing a conductive ink by mixing 10 to 25 wt% of an elastic matrix, 30-50 wt% of a conductive filler, and 30-50 wt% of a hydrophobic solvent: A step of applying the conductive ink to a substrate and performing spin coating at a predetermined speed or lower; A step of sintering the conductive ink for a predetermined period of time to remove the hydrophobic solvent and generate a conductive film; and A step of creating a plurality of holes in the conductive film using a micro patterning mold. A method for manufacturing an electromyography sensor electrode including a . In Article 10, The step of manufacturing the above conductive ink is: A method for manufacturing an electromyography sensor electrode, wherein the conductive ink is stirred using a magnetic stirrer for 2 hours or more at a predetermined speed or lower to induce uniform dispersion of the conductive filler within the conductive ink. In Article 11, The step of performing the above spin coating is: A method for manufacturing an electromyography sensor electrode by applying the conductive ink to the substrate without external stress and performing low-speed spin coating for a short period of time. In Article 12, A method for manufacturing an electromyography sensor electrode, wherein the conductive filler in the conductive ink is aligned on the substrate by the low-speed spin coating so that a two-dimensional plane is parallel to the plane of the substrate. In Article 13, The step of producing the above conductive film is: A method for manufacturing an electromyography sensor electrode, comprising heating at a temperature range of 150 to 180 degrees for 30 minutes to 2 hours to evaporate the hydrophobic solvent and harden the elastic base material while sintering the conductive filler. In Article 14, The step of creating the above hole is: A step of manufacturing the above micro patterning mold using 3D printing; and A step of forming the through hole by rolling the conductive film on the micro patterning mold. A method for manufacturing an electromyography sensor electrode including a . In Article 10, A method for manufacturing an electromyography sensor electrode, wherein the conductive filler has a longitudinal length of 2-10 μm and is at least one of silver (Ag) flakes, gold (Au) flakes, and copper (Cu) flakes. In Article 10, The above electromyography sensor electrode is a method for manufacturing an electromyography sensor electrode having a thickness of 30-90 um. In Article 10, A method for manufacturing an electromyography sensor electrode, wherein the plurality of holes are cylindrical or conical and have a diameter of 50 to 400 um on the surface. In Article 10, A method for manufacturing an electromyography sensor electrode wherein the elastic base material is silicone elastomer or polyurethane (PU). In Article 10, A method for manufacturing an electromyography sensor electrode, wherein the hydrophobic solvent has a molecular weight of 60 g/mol or more and is hexadecane or toluene.

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