KR20220055433A - Strain gauge sensor and mehtod for manufacturing thereof - Google Patents

Abstract

According to one embodiment, the present invention includes a substrate having a frame and an operating unit protruding from the frame; a piezoresistive sensor including a signal line located in the operating unit, and an electrode pad formed on one end of the signal line and located in the frame; and a pillar including a first pillar located in the frame and a second pillar located in the operating unit while having the piezoresistive sensor therebetween. According to the present invention, 3D artificial myocardial tissues can be cultivated and real-time contraction measurement is possible at the same time.

Description

Shrinkage force measuring sensor and manufacturing method thereof

The present invention relates to a sensor for measuring contractile force, and more particularly, to a sensor for measuring contractile force of artificial myocardial tissue, and a manufacturing method thereof.

Currently, in order to conduct a study on myocardial tissue, a model for implementing the tissue in an in vitro environment has been proposed. Microphysiological systems (MPS), also known as organs-on-chips, are based on human cells and reproduce their structure and function, so they It is suitable for testing reactivity or toxicity. MPS made based on myocardial tissue contracts similarly to real myocardial tissue, and this contractile force is used as an important factor in evaluating the maturity of artificial myocardial tissue.

A number of measurement systems have been proposed to measure the contractile force of the artificial myocardial tissue.

Force transducers are usually used to measure the mechanical properties of various body tissues, such as bones, muscles, and tendons. To measure a small force (tens of μN in the case of artificial myocardial tissue), an electromagnetic linear variable differential transformer is used to measure the acceleration, and then the force can be measured through Newton's second law of motion. However, in this process, the process of directly connecting the already mature artificial myocardial tissue to the force transducer is essential, and there is a limit that damage to the artificial myocardial tissue may occur during this process.

In the case of an optical readout system, after culturing artificial myocardial tissue between a platform composed of a flexible post and a rigid post, the flexible post deforms when the artificial myocardial tissue contracts is measured using an optical camera. Then, the contractile force is calculated using Hooke's law. Since the strain is measured through an optical camera, the contractile force can be measured without damaging the 3D artificial myocardial tissue, but there is a limitation in that a large amount of data is required for real-time analysis.

Accordingly, the present invention is to solve the above-mentioned limitations of the existing systems, and to provide a contractile force measuring sensor and manufacturing method capable of simultaneously culturing a three-dimensional artificial myocardial tissue and measuring the contractile force in real time.

A piezoresistive sensor including a frame, a substrate having an operation part protruding from the frame, a signal line positioned in the operation part, an electrode pad formed at one end of the signal line and positioned in the frame, and the piezoresistive sensor according to an embodiment of the present invention It includes a column including a first column positioned on the frame, and a second column positioned on the operating unit.

The signal line may be embedded in the substrate, and the operation unit may be located in a cutout formed in the frame.

The first pillar and one end are connected and may further include a skeleton located in the frame.

Surrounding the cutout, it may further include an upper frame and a lower frame attached to the upper and lower portions of the substrate to form a cell culture space, respectively.

The electrode pad may be located outside the cell culture space.

The lower frame may have a wider width than the upper frame.

After the object for measuring the contractile force is formed to connect between the first pillar and the second pillar, the operation unit may be bent when the object is contracted.

The subject may include cardiomyocytes.

The signal line may be formed in a zigzag shape having a length in a direction in which the object is contracted.

The substrate, the piezoresistive sensor, and the pillar may be formed by a 3D printer.

A method of manufacturing a contractile force measuring sensor according to another embodiment of the present invention comprises the steps of forming a boundary line on a sliding glass, forming a first pillar and a second pillar spaced apart from each other in a space surrounded by the boundary line, filling the space surrounded by the boundary line Forming a substrate, forming a piezoresistive sensor on the substrate, forming and curing the piezoresistive sensor, and separating the substrate from the sliding glass; , the substrate, and the piezoresistive sensor are formed with a 3D printer.

After the step of forming the boundary line, a skeleton may be formed in the space, and a first column connected to one end of the skeleton may be formed.

The first pillar and the skeleton may be made of a material having greater rigidity than the second pillar.

The curing may be performed in a vacuum oven at 60° C. for 24 hours.

In the forming of the piezoresistive sensor, the piezoresistive sensor may be embedded in the substrate by injecting a material for the piezoresistive sensor into the substrate.

The contractile force measuring sensor according to the present invention is capable of simultaneously measuring contractile force in real time according to the culture while culturing the three-dimensional artificial myocardial tissue.

In addition, since the contractile force measuring sensor according to the present invention is made of transparent PDMS, the maturation process can be observed under a microscope while culturing the three-dimensional artificial myocardial tissue.

In addition, the contractile force measuring sensor according to the present invention can be manufactured by 3D printing technology using a stacking method, thereby making it possible to repeatedly and precisely manufacture a platform having a relatively small (several hundred μm size) and complex shape.

In addition, it is possible to easily manufacture a contraction force measuring sensor composed of various materials by using multiple printing heads.

1 is a schematic diagram of a contractile force measuring sensor according to an embodiment of the present invention.
FIG. 2 is a photograph of a cross-section taken along line II-II of FIG. 1 .
3A and 3B are photographs for explaining the operation of the contractile force measuring sensor according to the presence or absence of a skeleton according to an embodiment of the present invention.
4 to 9 are diagrams schematically illustrating a method of manufacturing a contractile force measuring sensor according to an embodiment of the present invention.
10 is a view for explaining the operation of the contractile force measuring sensor according to an embodiment of the present invention.
11 is a photograph of a contractile force measuring sensor according to an embodiment of the present invention that is actually manufactured.
12 is a view for explaining a culture space of the apparatus for measuring contractility according to an embodiment of the present invention.
13 is a photograph illustrating self-alignment of myocardial cells inside the hydrogel part according to an embodiment of the present invention.
14 is a photograph of a change in contractility according to an embodiment of the present invention.
15 is a graph in which the contraction of artificial myocardial tissue according to time change is measured as a resistance change value through the contractile force measuring sensor according to an embodiment of the present invention, and a graph showing the contractile force corresponding to the measured resistance value in units of force am.
16 is an apparatus used for deriving a relationship between a contractile force value and a resistance change value of the contractile force measuring sensor according to an embodiment of the present invention.
17 is a graph illustrating a change in electrical resistance corresponding to a contractile force measured using the apparatus of FIG. 16 .

The present invention will be illustrated in more detail through the following examples, but the specific materials, structures and functions presented are merely described for the purpose of illustrating the exemplary embodiments of the present invention, and are not intended to limit the present invention to the following examples. . In addition, the embodiments according to the present invention may be implemented in various forms, and include all substitutes and equivalents included in the technical and design scope of the present invention.

Hereinafter, a contractile force measuring sensor and a manufacturing method thereof according to an embodiment of the present invention will be described with reference to the drawings.

1 is a schematic view of a contractile force measuring sensor according to an embodiment of the present invention, and FIG. 2 is a photograph taken when a cross-section is taken along line II-II of FIG. 1 .

As shown in FIG. 1 , the contractile force measuring sensor 100 according to an embodiment of the present invention is formed to protrude from the substrate 102 , the piezoresistive sensor 120 formed on the substrate 102 , and the substrate 102 . and a pillar 200 .

The substrate 102 is made of a flexible insulating material, and may be a polymer substrate that can be used in a culture medium, for example, polydimethylsiloxane (PDMS).

The substrate 102 includes a frame 11 including a cutout 3 having an inside cut out, and an operation section 13 protruding from the frame 11 toward the cutout 3 . Although the cutout 11 has a rectangular planar shape, it is not limited thereto, and any shape such as a circle or a polygon does not matter as long as it does not affect the movement of the operation unit 13 . The operation unit 13 may be flexibly bent even with a small force according to the growth of an artificial cardiomyocyte, which will be described later, and may be formed to have a width and a length capable of maintaining its shape even with repeated bending.

A second pillar 122 and a piezoresistive sensor 130 to be described later may be located in the operation unit 13 , and a first pillar 121 may be located in the frame 11 .

The piezoresistive sensor 130 includes a signal line 103 made of a conductive material and an electrode pad 104 formed at both ends of the signal line 103 , respectively. The signal line 103 may be positioned in the operation unit 13 , and the electrode pad 104 may be positioned in the frame 11 . In this case, the signal line 103 and the electrode pad 104 may be embedded in the operation unit 13 and the frame 11, respectively, but the present invention is not limited thereto, and the electrode pad 104 may be exposed to facilitate contact with an external device. can

The piezoresistive sensor 130 connects a measuring device to the electrode pad 104 to measure the change in resistance of the signal line 103 according to the bending movement of the operation unit 13. By disposing the long signal line 103 in a certain space, The three-dimensional artificial myocardial tissue may be formed in a zigzag shape to maximize the change in electrical resistance due to the contractile force. In this case, the signal line 103 may have a long zigzag shape in one direction, that is, a second direction X2 perpendicular to the first direction X1 connecting the two electrode pads 104 . The second direction X2 may be a direction in which the operation unit 13 protrudes and is bent.

The piezoresistive sensor 130 may be connected to form a half bridge or a Wheatstone bridge when measuring a change in electrical resistance.

The signal line 103 and the electrode pad 104 of the piezoresistive sensor 130 may include a conductive material, and may be Dow Corning, Sylgard 184 (PDMS) mixed with carbon black particles, but limited thereto. It is not limited thereto, and may include materials based on gold, silver, graphene, carbon, and the like.

The electrode pad 104 is a circle having a larger diameter than the width of the signal line 103 (or a rectangle having a larger width) in order to stably contact the electrical pin of the measuring instrument when measuring the change in electrical resistance. ) can be formed.

Since the contractile force measuring sensor according to an embodiment of the present invention can simultaneously measure the contractile force while immersed in the culture solution while culturing the artificial tissue, the piezoresistive sensor 130, which is a conductive material, is not exposed to the culture solution as shown in FIG. Likewise, it may be embedded in the substrate 10 .

The pole 120 includes a first pole 121 and a second pole 122, and the first pole 121 and the second pole 122 are respectively positioned on opposite sides with the piezoresistive sensor 130 interposed therebetween. and the first pillar 121 is located on the frame 11 and the second pillar 122 is located on the operation unit 13 . A skeleton 123 may be further formed on the substrate 10 , and the first pillar 121 may be positioned at one end of the skeleton 123 , and may be connected to and fixed to the skeleton 123 .

The skeleton 123 includes a central axis 21 and a plurality of branches 23 protruding to both sides from the central axis 21 . The skeleton 123 allows the first pillar 121 to be fixed without moving during the operation of the operation unit 13 .

3A and 3B are photographs for explaining the operation of the contractile force measuring sensor according to the presence or absence of a skeleton according to an embodiment of the present invention.

As shown in FIG. 3A , when the skeleton 123 is not formed, the first pillar 121 is lifted by the contractile force during cell culture, and the contractile force is not properly transmitted to the operation unit 13 , so that it may not be bent. If the moving part is not bent, there is no change in resistance, so it is impossible to measure the resistance change according to the contractile force properly.

Therefore, by connecting and fixing the first pillar 121 to the skeleton 123 as shown in FIG. 3B , the first pillar 121 is stably supported without lifting, so that the operation unit 13 is easily bent by the contractile force. make it possible In this case, the first pillar 121 and the skeleton 123 may be made of a material having greater rigidity than the second pillar 122 in order to maintain a constant shape when the substrate is supported.

The ends of the pillars 121 and 122 may be formed flat, but are not limited thereto, and irregularities (not shown) are formed on the upper surface of the ends in order to increase the adhesion of the hydrogel unit 300 to be described later, or the upper The width may be formed to be larger than the lower width.

The pillar 120 and the skeleton 123 serve as a support so that the three-dimensional artificial myocardial tissue can be stably cultured. For example, PDMS capable of 3D printing was used, but in addition, 3D printing such as PCL, PEVA, etc. is possible and biocompatible materials with high mechanical properties can be used.

Hereinafter, a method of manufacturing a contractile force measuring sensor according to an embodiment of the present invention will be described with reference to the drawings.

4 to 9 are diagrams schematically illustrating a method of manufacturing a contractile force measuring sensor according to an embodiment of the present invention.

First, referring to FIG. 4 , a slide glass (not shown) is prepared, and a boundary line 101 for maintaining the substrate shape on the slide glass is formed with a 3D printer. The boundary line 101 forms an outline of a substrate shape to be formed.

The 3D printer contains a metal syringe equipped with a metal nozzle with an inner diameter of 250 μm. After loading PEVA (polyethylene vinyl acetate) into a metal syringe, the temperature of the syringe is heated to 120°C. The heated PEVA is discharged onto the slide glass surface through a pneumatic pressure of 500 kPa. In this case, the printer head may be fed at a speed of 200 mm/min, and the boundary line 101 may be formed by stacking three layers with a height of 100 μm per layer. The height of the boundary line 101 may be the thickness of the substrate.

Although PEVA is used for the boundary line 101, 3D printing is possible, such as PCL (Polycaprolactone), and a biocompatible thermoplastic polymer may be used.

Thereafter, as shown in FIG. 5 , the skeleton 123 is formed in the space surrounded by the boundary line 101 , and the first pillar 121 is formed at one end of the skeleton 123 by a 3D printer. The skeleton includes a central axis 21 and a plurality of branches 23 protruding to both sides from the central axis 21 . A first pillar 121 may be formed at one end of the central axis 21 .

At this time, the 3D printer includes a plastic syringe equipped with a plastic tapered nozzle having an inner diameter of 250 μm.

The printer head of the 3D printer can be cooled to 4°C, and the mixture for the column is loaded into the cooled plastic syringe of the 3D printer. The mixture can be prepared by mixing PDMS and curing agent (Dow Corning) in a mass ratio of 10:1.

The mixture loaded in a plastic syringe is discharged onto the slide glass surface with a pneumatic pressure of 160 kPa. At this time, the printer head may be transferred at a speed of 200 mm/min and stacked eight times at a height of 200 μm per layer to form a column with a height of 3.6 mm.

Thereafter, as shown in FIG. 6 , the second pillar 122 spaced apart from the first pillar 121 is formed by a 3D printer. The second pillar 122 may be formed of a material in which a silicone elastomer and a curing agent are mixed.

Next, as shown in FIG. 7 , the space surrounded by the boundary line 101 is filled with a 3D printer to form the substrate 102 .

The substrate 102 is prepared by mixing PDMS and a curing agent (Dow Corning) in a mass ratio of 10:1. Then, it is loaded into a plastic syringe equipped with a plastic nozzle having an inner diameter of 250 μm, and the printer head is cooled to 4°C.

The cooled PDMS is discharged on the slide glass with a pneumatic pressure of 180 kPa. At this time, the print head is transferred at a speed of 120 mm/min.

Then, as shown in FIG. 8 , the piezoresistive sensor 130 is formed on the substrate 102 by a 3D printer.

Since the piezoresistive sensor 130 is formed before the substrate 102 is cured, the material for the piezoresistive sensor 130 is inserted into the substrate 102 so that the piezoresistive sensor 130 is embedded in the substrate 102 . (refer to FIG. 2). At this time, due to the difference in specific gravity between the material for the piezoresistive sensor 130 and the material for the substrate 102 , the material for the piezoresistive sensor 130 may be inserted into the substrate 102 or pushed by the discharge pressure discharged from the nozzle. A material for a piezoresistive sensor may be inserted into the substrate 102 .

The piezoresistive sensor 130 may include a conductive material, and may be PDMS in which carbon black particles are mixed, but is not limited thereto. Gold, silver, graphene , and may include a material based on carbon or the like.

The piezoresistive sensor 130 is formed by loading a conductive material mixture into a plastic syringe equipped with a plastic tapered nozzle having an inner diameter of 400 μm and then discharging it onto the substrate 102 with a pneumatic pressure of 700 kPa. In this case, after the printer head is transferred at a speed of 20 mm/min and the piezoresistive sensor 130 is formed, the electrode pad 104 may be formed.

The electrode pad 104 has a diameter larger than the width of the signal line 103 of the piezoresistive sensor 130 in order to stably contact an electrical pin when measuring the change in electrical resistance of the piezoresistive sensor 130 . The branches may be circular or polygonal having a large width.

The input resistance sensor 130 may be elongated in the direction in which the three-dimensional artificial myocardial tissue contracts, that is, the second direction X2, so that the change in electrical resistance can be maximized even with a small contractile force of the three-dimensional artificial myocardial tissue.

Thereafter, the curing process is performed in a vacuum oven at 60° C. to 80° C. for 18 hours to 24 hours.

Next, as shown in FIG. 9 , by separating the substrate 102 from the glass slide, the contraction force measuring sensor 100 is completed (see FIG. 1 ). At this time, the boundary line 101 may be separated together with the slide glass.

As described above, by using the contractile force measuring apparatus according to an embodiment of the present invention, it is possible to easily measure a change in electrical resistance according to a change in contractile force during culturing of artificial myocardial cells.

10 is a view for explaining the operation of the contractile force measuring sensor according to an embodiment of the present invention.

Referring to FIG. 10 , in the contractile force measuring sensor according to an embodiment of the present invention, a hydrogel part 300 including artificial myocardial cells is formed between the first pillar 121 and the second pillar 122 . The hydrogel part 300 may be immersed in a culture solution together with the contractility measuring sensor 100 and contracted while myocardial cells are cultured (refer to the red arrow). At this time, the operating part is bent by the contractile force, and when the contractile force measuring device according to an embodiment of the present invention is used, the contracting force is changed and the moving part is bent at the same time (see black and blue arrows) to change the electrical resistance, and the electrode pad 104 ) can measure the electrical resistance in real time through a measuring instrument connected to the

11 is a photograph of a contractile force measuring sensor according to an embodiment of the present invention that is actually manufactured.

11, the hydrogel part 300 may be formed by a 3D printer to connect between the pillars 121 and 122 of the contractile force measuring sensor 100 according to an embodiment of the present invention. The hydrogel unit 300 may include cardiomyocytes.

The hydrogel part 300 can be formed by 3D printing using bio-ink, and the bio-ink is a decellularized extracellular matrix solution, hydrogel, collagen (Collagen), fibrinogen (Fibrinogen), or a mixture thereof. can

Since the hydrogel part 300 can be suspended from the pillars 121 and 122 in the culture solution, a cap 400 is formed on the pillars 121 and 122 on which the hydrogel part 300 is formed, and the hydrogel part 300 ) is fixed so as not to be separated from the pillars (121, 122).

The hydrogel can be cultured in a cell culture solution together with a contractile force measuring sensor.

12 is a view for explaining a culture space of the apparatus for measuring contractility according to an embodiment of the present invention.

As shown in FIG. 12 , an upper frame 61 and a lower frame 62 , which form a space for containing the cell culture solution, are attached to the upper and lower portions of the substrate 102 on the upper and lower portions of the substrate 102 , respectively. The upper frame 61 and the lower frame 62 may be formed to surround the cutout 3 . The lower frame 62 may be wider than the upper frame 61 to support the substrate 102 , and may be formed to extend to a portion corresponding to the electrode pad 104 .

Accordingly, the upper frame 61 and the lower frame 62 form a culture space (S) that communicates through the cutout (3), the cell culture solution is filled in the culture space (S), and the operation unit (13) is submerged in cell culture. Since the operation unit 13 is embedded in the substrate 102 , it can be safely protected without being exposed to the cell culture solution.

The electrode pad 104 of the piezoresistive sensor is located outside the culture space S to connect with an external device, and may be exposed for connection with an external device. When the electrode pad 104 is formed to be embedded in the substrate, the upper substrate may be removed for connection with an external device, or a contact hole may be formed to connect the electrode pad 104 .

After a certain period of time after culturing, the cardiomyocytes included in the hydrogel unit 300 self-align in the longitudinal direction, that is, the second column connecting between the first column 121 and the second column 122 . Self-aligns in direction (X2).

13 is a photograph for explaining the self-alignment of myocardial cells inside the hydrogel part according to an embodiment of the present invention.

13 is a cell viability analysis method (live/dead imaging) and immunofluorescence imaging (immunofluorescence imaging) that was taken, cells were cultured and taken on days 0, 14, and 28, respectively, as time goes by, the cardiomyocytes are aligned in one direction. it can be confirmed that

After self-alignment, as the myocardial cells are cultured, the hydrogel part contracts and contractile force is generated, and accordingly, the moving part may be bent.

14 is a photograph of a contraction phenomenon according to an embodiment of the present invention.

As shown in FIG. 14A , when the hydrogel part is formed in the contractile force measuring sensor according to an embodiment of the present invention, it can be seen that the operating part is bent from the diastole to the systole as shown in FIG. 14B .

15 is a graph in which the contraction of artificial myocardial tissue according to time change is measured as a resistance change value through a contractile force measuring sensor according to an embodiment of the present invention, and a graph showing the contractile force corresponding to the measured resistance value in units of force am.

Referring to FIG. 15 , a change in the contractile force may be measured in real time through a change in an electrical resistance value while the time is changed to 0s, 5s, 10s, 15s, and 20s.

16 is a device used to derive the relationship between the value of the contractile force to the resistance change value of the contractile force measuring sensor according to an embodiment of the present invention, and FIG. 17 is an electrical corresponding to the contractile force measured using the device of FIG. It is a graph showing the resistance change value.

16, a suture 62 connected to a force transducer 60 instead of a hydrogel is connected to the first column and the second column of the contractile force measuring sensor according to an embodiment of the present invention. . And, when the contractile force measuring sensor 100 is positioned on the stage on which the motor is mounted and moves the contracting force measuring sensor in the contracting and relaxing directions, tension is applied to the suture connected to the first post and the second post, whereby the hydrogel is It bends and unfolds as when contracting and relaxing. At this time, the tension acting on the suture was matched to the contractile force generated during cell culture using Heidelgel.

Then, the tension applied to the suture was measured using a force transducer, and the change in resistance occurring at this time was measured.

Referring to FIG. 17 , it can be seen that the electrical resistance value changes according to the change in the contractile force, and it can be seen that there is a constant relationship like a quadratic curve. Therefore, the degree of culture can be easily determined by measuring the contractile force based on the change in electrical resistance measured during cell culture.

Although preferred embodiments of the present invention have been described above, the present invention is not limited thereto, and various modifications and variations are possible within the scope of the appended claims and the detailed description of the invention and the accompanying drawings. It goes without saying that it falls within the scope of the invention.

100: contraction force measuring sensor 102: substrate
103: piezoresistive sensor 123: skeleton
200: pillar 300: hydrogel part

Claims (16)

a frame, a substrate having an operation part protruding from the frame;
A piezoresistive sensor including a signal line positioned in the operation unit, an electrode pad formed at one end of the signal line and positioned in the frame;
A pole including a first pole positioned on the frame and a second pole positioned on the operation unit with the piezoresistive sensor interposed therebetween
A contractile force measuring sensor comprising a. In claim 1,
The signal line is a contractile force measuring sensor embedded in the substrate. In claim 1,
The operation unit is a contractile force measuring sensor located in the cutout formed in the frame. In claim 1,
The contractile force measuring sensor further comprising an upper frame and a lower frame surrounding the cutout, respectively attached to the upper and lower portions of the substrate to form a cell culture space. In claim 4,
The electrode pad is a contractile force measuring sensor located outside the cell culture space. In claim 4,
The lower frame is a contractile force measuring sensor having a wider width than the upper frame. In claim 1,
A skeleton having one end connected to the first pillar and positioned in the frame
Contraction force measuring sensor further comprising a. In claim 1,
A contractile force measuring sensor in which an object for measuring a contractile force is formed to be connected between the first pillar and the second post, and then the operation unit is bent when the object is contracted. In claim 8,
The object is a contractile force measuring sensor comprising myocardial cells. In claim 8,
The signal line is a contractile force measuring sensor having a long zigzag shape in a direction in which the object contracts. In claim 1,
The substrate, the piezoresistive sensor, and the pillar are a contractile force measuring sensor formed by a 3D printer. Forming a boundary line on the sliding glass,
Forming a first column and a second column spaced apart from each other in the space surrounded by the boundary line,
forming a substrate by filling the space surrounded by the boundary line;
forming a piezoresistive sensor on the substrate;
After forming the piezoresistive sensor, curing;
Separating the substrate from the sliding glass
including,
The boundary line, the first pillar, the second pillar, the piezoresistive sensor and the substrate are formed by a 3D printer. In claim 12,
After forming the boundary line,
A method of manufacturing a contractile force measuring sensor by forming a skeleton in the space and forming the first pillar connected to one end of the skeleton. In claim 13,
The method of manufacturing a contractile force measuring sensor, wherein the first pillar and the skeleton are made of a material having greater rigidity than the second pillar. In claim 12,
The curing step is
Method of manufacturing a contraction force measuring sensor proceeding for 18 hours to 24 hours in a vacuum oven of 60 ℃ to 80 ℃. In claim 12,
In the step of forming the piezoresistive sensor,
A method of manufacturing a contractile force measuring sensor in which the piezoresistive sensor is embedded in the substrate by injecting the material for the piezoresistive sensor into the substrate.

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