Stent-graft-mounted blood leakage detection sensor, manufacturing method thereof, and wireless blood leakage detection system using blood leakage detection sensor

Abstract

The present invention relates to a stent-graft-mounted blood leakage detection sensor, a manufacturing method thereof, and a wireless blood leakage detection system using the blood leakage detection sensor. According to an embodiment of the present invention, the a stent-graft-mounted blood leakage detection sensor is attached to the surface of the stent-graft, does not induce blood flow leakage or damage to the stent-graft, and allows a change in capacitance to occur in response to blood leakage so as to detect blood leakage occurring after stent-graft treatment, thereby enabling blood leakage (endoleak) to the outside of the stent to be monitored in daily life without medical imaging examination.

Description

Stent-graft attached blood leak detection sensor, method of manufacturing the same, and wireless blood leak detection system using the blood leak detection sensor

[0001] The present invention relates to a stent-graft attached blood leak detection sensor, a method for manufacturing the same, and a wireless blood leak detection system using the blood leak detection sensor. More specifically, the invention relates to a sensor attached to the surface of a stent-graft to detect endoleak occurring after stent-graft treatment, a method for manufacturing the sensor, and a wireless system using the sensor.

[0002]

[0003] Currently, surgery or procedures are being performed to treat aortic aneurysms, which are one of the major aortic diseases. Surgery is a traditional treatment method in which the aortic aneurysm is removed and replaced with an artificial blood vessel through open abdominal or thoracotomy under general anesthesia. The procedure involves inserting a stent-graft into the blood vessel through a tube and installing it at the location of the aortic aneurysm. Referring to the patent document in the prior art literature below and Fig. 1, the stent-graft consists of a cylindrical metal mesh covered with fabric; it prevents the aneurysm from expanding further by blocking blood flow into the aortic aneurysm and allowing blood to move only within the stent-graft. Therefore, while surgery removes the aortic aneurysm, the procedure prevents the aneurysm from expanding and rupturing. Since the procedure is performed under local anesthesia without skin incision, it carries a lower risk of complications than surgery and allows for a quick return to daily life; thus, stent insertion is primarily performed unless surgical treatment is unavoidable.

[0004] However, after stent implantation for an aortic aneurysm, there are cases where blood leaks out of the stent due to vascular aging or stent displacement. This blood leakage is called an endoleak. Among the five types of endoleaks, Type 2 endoleaks are the most common but usually resolve naturally, while for Type 1 endoleaks, which carry the highest risk, surgery is recommended immediately upon detection.

[0005] These endoleaks primarily occur within five years of stent implantation and are difficult to detect early because they do not show significant external characteristics until the aorta ruptures due to continued leakage; when they do manifest externally, the patient's mortality risk is extremely high. Currently, to detect them early, patients must visit the hospital periodically for CT or MRI scans. However, since the visit interval is often approximately six months, the situation may already be dangerous by the time the condition is discovered, and the examination process entails various inconveniences.

[0006] Therefore, there is an urgent need for measures to monitor and detect endoleaks occurring after stent implantation in real time at an early stage.

[0007]

[0008] The present invention aims to solve the problems of the aforementioned prior art. One aspect of the present invention is to provide a stent-graft attached blood leak detection sensor that is installed on the outer surface of a stent-graft, does not cause blood flow leakage, does not damage the stent-graft, and induces a change in capacitance corresponding to blood leakage to detect blood leakage within the blood vessel after insertion of the stent-graft, a method for manufacturing the same, and a wireless blood leak detection system capable of continuously monitoring blood leakage wirelessly outside the body based on an LC resonant circuit in which a change in resonant frequency is induced in response to a change in capacitance of the blood leak detection sensor.

[0009]

[0010] A blood leakage detection sensor attached to a stent-graft according to an embodiment of the present invention comprises: a base film surrounding the outer surface of a tubular stent-graft, having a continuous curved portion in a zigzag shape along the longitudinal direction; an adhesive film adhering one surface of the base film to the outer surface of the stent-graft; an encapsulation film formed in a shape corresponding to the base film, with one surface laminated to the other surface of the base film; and an electrode portion comprising a first conductive line formed in a shape corresponding to the zigzag shape of the base film, and a second conductive line formed in a shape corresponding to the first conductive line and arranged parallel to the first conductive line at a predetermined distance, and disposed between the base film and the encapsulation film, wherein the capacitance changes when blood comes into contact with the other surface of the encapsulation film; and detects the blood leaking from the blood vessel where the stent-graft is installed according to the change in capacitance.

[0011] In addition, in the stent-graft attached blood leakage detection sensor according to an embodiment of the present invention, the base film and the encapsulation film may each be formed of an insulating material.

[0012] In addition, in the stent-graft attached blood leakage detection sensor according to an embodiment of the present invention, the base film may be adhered to a region of the outer surface of the stent-graft that is in contact with the blood vessel.

[0013] In addition, in a stent-graft attached blood leakage detection sensor according to an embodiment of the present invention, the adhesive film may comprise: a porous heat-sealable sheet formed by arranging a plurality of fibers made of a first heat-sealable polymer; and a heat-sealable layer formed by coating a powder made of a second heat-sealable polymer on the surface of the heat-sealable sheet.

[0014] In addition, in a stent-graft attached blood leakage detection sensor according to an embodiment of the present invention, the first conductive line and the second conductive line may each have a width of 180 to 220 μm, a thickness of 0.5 μm or less, and a spacing of 20 to 40 μm.

[0015] In addition, in a stent-graft attached blood leakage detection sensor according to an embodiment of the present invention, the electrode portion may further include a third conductive line that is relatively shorter than the length of the second conductive line, is formed in a shape corresponding to a predetermined portion of the longitudinal direction region of the second conductive line, and is arranged parallel to the predetermined portion at a predetermined distance.

[0016] Meanwhile, a method for manufacturing a stent-graft attached blood leakage detection sensor according to an embodiment of the present invention comprises: (a) a step of manufacturing an adhesive film by coating a powder made of a second heat-adhesive polymer onto a porous heat-adhesive sheet formed by arranging a plurality of fibers made of a first heat-adhesive polymer; (b) a step of manufacturing an electrode structure on a substrate, the electrode structure comprising a base film having a continuous curved portion in a zigzag shape along the longitudinal direction, two or more conductive lines formed in a shape corresponding to the zigzag shape of the base film and arranged parallel to each other, and a encapsulation film formed in a shape corresponding to the base film and laminated on the base film with the conductive lines in between; and (c) a step of separating the electrode structure from the substrate and transferring it to one surface of the adhesive film.

[0017] In addition, in a method for manufacturing a stent-graft attached blood leakage detection sensor according to an embodiment of the present invention, the step (a) may include: a step of melting the first heat-sealable polymer; a step of manufacturing the heat-sealable sheet while spinning the melted first heat-sealable polymer into the form of the fiber; and a step of coating the surface of the manufactured heat-sealable sheet with a coating solution in which the powder is dissolved.

[0018] In addition, in a method for manufacturing a stent-graft attached blood leakage detection sensor according to an embodiment of the present invention, step (b) may include: forming a sacrificial layer on the substrate; forming the base film with a first insulating polymer on the sacrificial layer; forming the conductive line on the base film; and forming the encapsulation film with a second insulating polymer on the base film on which the conductive line is formed.

[0019] In addition, in a method for manufacturing a stent-graft attached blood leakage detection sensor according to an embodiment of the present invention, the step (c) may include: removing the sacrificial layer; attaching a first adhesive sheet to one surface of the electrode structure and separating it from the substrate; placing the other surface of the electrode structure attached to the first adhesive sheet on one surface of the adhesive film; placing a second adhesive sheet on the other surface of the adhesive film on which the electrode structure is placed to form a multilayer structure; heating the multilayer structure to transfer the electrode structure to the adhesive film; and immersing the heated multilayer structure in an organic solvent to remove the first adhesive sheet and the second adhesive sheet from the adhesive film.

[0020] In addition, the method for manufacturing a stent-graft-attached blood leakage detection sensor according to an embodiment of the present invention may further include the step of (d) arranging the electrode structure to surround the outer surface of a tubular stent-graft, and heating the adhesive film on which the electrode structure is transferred to adhere the electrode structure to the outer surface of the stent-graft.

[0021] In addition, in a method for manufacturing a stent-graft attached blood leak detection sensor according to an embodiment of the present invention, step (d) may include: a step of spraying ethanol to attach the adhesive film to the outer surface of the stent-graft; and a step of heating the attached adhesive film to attach the electrode structure to the outer surface of the stent-graft.

[0022] Meanwhile, a blood leakage detection wireless system according to an embodiment of the present invention comprises: a blood leakage detection sensor attached to the outer surface of a tubular stent-graft and acting as a variable capacitor whose capacitance changes in response to blood leaking from a blood vessel; a planar inductor attached to the outer surface of the stent-graft and electrically connected to the blood leakage detection sensor, forming an LC resonant circuit whose resonant frequency changes in response to the change in capacitance; and a wireless reader unit comprising a reader coil disposed outside the body and magnetically coupled with the planar inductor to wirelessly detect the change in resonant frequency.

[0023] In addition, in a blood leakage detection wireless system according to an embodiment of the present invention, the planar inductor may be formed in an elongated rectangular shape along the length direction of the stent-graft.

[0024] In addition, in a blood leak detection wireless system according to an embodiment of the present invention, the planar inductor may have a width of 20 to 40 mm and a length of 80 to 120 mm.

[0025] In addition, in a blood leakage detection wireless system according to an embodiment of the present invention, the planar inductor may be disposed in a semicircular region of the outer surface of the stent-graft.

[0026] In addition, in a blood leak detection wireless system according to an embodiment of the present invention, the reader coil is formed in a shape corresponding to the shape of the planar inductor, and its size may be 1.5 to 2.5 times that of the planar inductor.

[0027]

[0028] The features and advantages of the present invention will become more apparent from the following detailed description based on the accompanying drawings.

[0029] Prior to this, terms and words used in this specification and claims should not be interpreted in their ordinary and dictionary meanings, but should be interpreted in a meaning and concept consistent with the technical spirit of the invention, based on the principle that the inventor can appropriately define the concept of the terms to best describe his invention.

[0030]

[0031] According to the present invention, blood leakage (endoleak) outside the stent that may occur after stent-graft treatment (Endovascular Aneurysm Repair, EVAR) can be monitored in daily life without imaging examination, and a passive LC resonant circuit that can operate without a battery or a separate power supply is provided, so it can be used stably even after long-term implantation.

[0032] In addition, by integrating a large-area planar inductor into the outer wall of the stent and applying a reader coil larger than it to the outside, stable wireless signal transmission and detection are possible even in the aorta located deep inside the human body. Furthermore, since the resonance frequency changes distinctly in the MHz range depending on the area covered by blood over the electrode sensor, even minute amounts of blood leakage can be sensitively detected. Consequently, patients can check for the occurrence of endoleaks in real time using an external reader without visiting a hospital, thereby significantly reducing reliance on high-cost, high-risk imaging equipment such as CT or X-ray.

[0033] Furthermore, the present invention can be extended not only to the detection of aortic aneurysm endoleaks but also to the monitoring of various biomarkers such as pressure, body fluids, and blood flow, making it highly clinically useful.

[0034]

[0035] FIG. 1 is a perspective view showing a part of a conventional stent-graft.

[0036] FIG. 2 is a diagram showing a state in which a stent-graft attached blood leakage detection sensor according to an embodiment of the present invention is attached to the outer surface of a stent-graft.

[0037] FIG. 3 is a diagram showing the state in which a stent-graft attached blood leakage detection sensor according to an embodiment of the present invention is placed inside a blood vessel.

[0038] FIG. 4 is an exploded perspective view of a stent-graft attached blood leak detection sensor with the dotted box A of FIG. 2 disassembled.

[0039] FIG. 5 is a cross-sectional view of a stent-graft attached blood leakage detection sensor according to an embodiment of the present invention.

[0040] FIG. 6 is a diagram illustrating a stent-graft attached blood leakage detection sensor according to another embodiment of the present invention.

[0041] FIG. 7 is a diagram illustrating the electrode portion of a stent-graft attached blood leakage detection sensor according to another embodiment of the present invention.

[0042] FIG. 8 is a diagram illustrating the circuit connection relationship of the electrode portion of a stent-graft attached blood leakage detection sensor according to another embodiment of the present invention.

[0043] FIG. 9 is a configuration diagram of a blood leak detection wireless system according to an embodiment of the present invention.

[0044] FIG. 10 is a diagram illustrating wireless data transmission of a blood leak detection wireless system according to an embodiment of the present invention.

[0045] FIG. 11 is an image showing the connection relationship (a) between the blood leak detection sensor and the planar inductor and the reader coil (b) shown in FIG. 9.

[0046] Figure 12 is a drawing comparing the shape and size of the planar inductor and leader coil shown in Figure 11.

[0047] FIG. 13 is a flowchart of a method for manufacturing a stent-graft attached blood leakage detection sensor according to an embodiment of the present invention.

[0048] Figure 14 is a diagram illustrating the adhesive film manufacturing steps shown in Figure 13.

[0049] Figure 15 is a diagram illustrating the steps for manufacturing the electrode structure shown in Figure 13.

[0050] Figure 16 is a diagram illustrating the electrode structure transfer step shown in Figure 13.

[0051] Figure 17 is a diagram illustrating the electrode structure bonding step shown in Figure 13.

[0052] Figure 18 shows the results of evaluating the flexibility (a), adhesion (b), and absorbency (c) of the adhesive film prepared in Experimental Example 1 into the stent fabric.

[0053] Figure 19 shows the results of a compression expansion test for the blood leak detection sensor manufactured in Experimental Example 1.

[0054] Figure 20 is a graph showing the capacitance (a) according to the frequency of the applied voltage for each target substance of the blood leak detection sensor manufactured in Experimental Example 1, and the change in capacitance (b) according to blood leak.

[0055] FIG. 21 is a diagram illustrating a mechanism (a) for wirelessly detecting endoleaks using capacitance-based resonant frequency changes by inserting and mounting a capacitance-type blood leak detection sensor integrated into a stent of a blood leak detection wireless system according to Experimental Example 2, an arrangement (b) of a blood leak detection sensor for detecting the location and direction of blood leaks, a state of capacitance change due to endoleaks (left: low capacitance in normal state, right: increased capacitance due to blood leaks) (c), and a cross-sectional view of an electrode located between a blood vessel and a stent and an electric field distribution of a capacitance sensor (d).

[0056] FIG. 22 is a graph showing the change in capacitance according to the thickness of leaked blood (a), the change in capacitance according to the area of ​​blood covering the sensor (b), and the time response of the sensor according to the initial blood flow velocity (c) of the blood leak detection sensor according to Experimental Example 2.

[0057] Figure 23 shows the change in capacitance according to the blood leakage location of the blood leakage detection sensor according to Experimental Example 2 (when blood flow occurs at site 1: capacitance increases in both electric field (E-field) 1 and 2; when blood flow occurs at site 2: capacitance increases only in electric field (E-field) 2), the result of a stress test in which stent compression and expansion were repeated 1,000 times (b), and the daily change in capacitance after immersing the sensor in phosphate-buffered saline (PBS) for one month (c), respectively.

[0058] FIG. 24 shows an image of the in vitro experimental setup using a porcine aorta of the blood leak detection sensor according to Experimental Example 2 (a), the change in capacitance when blood, albumin, and PBS flow between the stent and the porcine aorta under experimental conditions (b), and the change in capacitance within the normal physiological blood pressure range (c), respectively.

[0059] FIG. 25 shows an image (a) of a blood leak detection sensor and a heat-sealed film according to Experimental Example 2, a bending strength (b) of the adhesive, and an adhesion strength (c) of the adhesive, respectively.

[0060] FIG. 26 shows an SEM image (a) of the inner surface of a fully cured stent fabric after coating or attaching the adhesive according to Experimental Example 2, a state in which the adhesive is applied to the outer surface of the stent (b), a blood droplet contact angle (c) on the inner fabric of the stent after the adhesive is applied, and the results of Calcein staining and cell viability analysis (d to e), respectively.

[0061] FIG. 27 shows the results of a numerical simulation (a) for an aortic aneurysm (left) when the electrode-attached stent according to Experimental Example 2 is not inserted and for an endo-leak when the electrode-attached stent is inserted, and the results of a numerical simulation and an analytical approach (b) for blood leakage rates according to the gap size induced by changes in electrode thickness.

[0062] FIG. 28 shows a blood leakage image (a) according to the thickness (5㎛ and 1,000㎛) of the blood leakage detection sensor of Experimental Example 2, the gap between the stent-graft and the blood vessel according to the thickness of the blood leakage detection sensor (the gap increases as the device thickness increases, b), and optical microscope images (c) of the top of the polyimide etched using O2 and the bottom of the polyimide etched using a mixture of O2 and CF4 gas during the etching process.

[0063] FIG. 29 shows an example of applying an LC-based blood leakage wireless monitoring system according to Experimental Example 2 to the human body (a), and images of a planar inductor, a reader coil, an electrode sensor with the planar inductor integrated therein, and a stent (b), respectively.

[0064] FIG. 30 shows the simulation configuration performed to optimize the dimensions of the planar inductor and the leader coil according to Experimental Example 2 (the leader coil (Tx) is designed to be twice the length of the planar inductor (Rx) so that measurements can be taken from a longer distance) (a), the change in resonance frequency according to the blood coverage of the blood leak detection sensor (b), and the change in resonance frequency measured when the distance between the planar inductor and the leader coil is 70 mm (c).

[0065]

[0066] The objects, specific advantages, and novel features of the present invention will become more apparent from the following detailed description and preferred embodiments in conjunction with the accompanying drawings. It should be noted that in assigning reference numbers to the components of each drawing in this specification, identical components are assigned the same number whenever possible, even if they are shown in different drawings. Furthermore, terms such as "first," "second," etc., are used to distinguish one component from another, and the components are not limited by these terms. In the following description of the present invention, detailed descriptions of related prior art that could unnecessarily obscure the essence of the invention are omitted.

[0067] Hereinafter, preferred embodiments of the present invention will be described in detail with reference to the attached drawings.

[0068]

[0069] FIG. 2 is a drawing showing a state in which a stent-graft attached blood leak detection sensor according to an embodiment of the present invention is attached to the outer surface of a stent-graft, and FIG. 3 is a drawing showing a state in which a stent-graft attached blood leak detection sensor according to an embodiment of the present invention is placed inside a blood vessel. FIG. 4 is an exploded perspective view of a stent-graft attached blood leak detection sensor with the dotted box A of FIG. 2 disassembled, and FIG. 5 is a cross-sectional view of a stent-graft attached blood leak detection sensor according to an embodiment of the present invention.

[0070] As illustrated in FIGS. 2 to 5, a stent-graft attached blood leakage detection sensor (100) according to an embodiment of the present invention comprises a base film (10) that surrounds the outer surface of a tubular stent-graft (1) and has a continuous zigzag curved portion along the longitudinal direction, an adhesive film (20) that adheres one surface of the base film (10) to the outer surface of the stent-graft (1), a encapsulation film (30) formed in a shape corresponding to the base film (10) and having one surface laminated to the other surface of the base film (10), a first conductive line (41) formed in a shape corresponding to the zigzag shape of the base film (10), and a second conductive line (42) formed in a shape corresponding to the first conductive line (41) and arranged parallel to the first conductive line (41) at a predetermined distance, and is disposed between the base film (10) and the encapsulation film (30), and the encapsulation film (30) It includes an electrode part (40) whose capacitance changes when blood comes into contact with the other surface, and detects blood leaking from the blood vessel where the stent-graft (1) is installed according to the change in capacitance.

[0071]

[0072] The present invention relates to a sensor attached to the surface of a stent-graft to detect blood leakage (endoleak) occurring after stent-graft treatment. Endoleak refers to the leakage of blood from a stent due to vascular aging or stent displacement following stent implantation for an aortic aneurysm. In particular, Type 1 endoleak carries a high risk and requires immediate surgery upon detection; however, since patients are currently only visited approximately every six months for CT or MRI examinations, early detection is difficult and the examination process is cumbersome. The present invention was devised to solve these problems.

[0073] Specifically, a stent-graft attached blood leakage detection sensor (100) according to an embodiment of the present invention comprises a base film (10), an adhesive film (20), an encapsulation film (30), and an electrode portion (40).

[0074]

[0075] The base film (10) is a member having a predetermined length and is provided with a continuous curved section in a zigzag shape along its length direction. That is, the curved section is formed in a shape similar to a sine wave overall. However, the curved section does not necessarily have to repeat a symmetrical shape at predetermined intervals like a sine wave; it is sufficient if it is formed in a wavy shape overall. When the base film (10) stretches or shrinks in the length direction due to an external force, this curved section unfolds and folds, preventing the base film (10) from breaking. Furthermore, when the external force is removed, the curved section has a restoring force, which can contribute to the base film (10) being restored to its original state. Additionally, the base film (10) is bonded through an adhesive film (20), and the bonding area can be increased due to the zigzag shape.

[0076] When this base film (10) is attached to the outer surface of the tubular stent-graft (1), it surrounds the outer surface of the stent-graft (1) along the arc direction. The stent-graft (1) is a cylindrical metal mesh structure covered with fabric and is formed in a tubular shape (see FIG. 1). When the stent-graft (1) is installed in a blood vessel, it is inserted into the blood vessel in a folded state and unfolded at the installation location, at which time a portion of its outer surface adheres to the inner wall of the blood vessel. Here, by having a curved portion of the base film (10), the base film (10) is prevented from breaking when the stent-graft (1) unfolds. The base film (10) is placed in a portion of the outer surface of the stent-graft (1) that contacts the blood vessel. Generally, the contact area between the stent-graft (1) and the blood vessel is 2 cm or more, and the base film is placed in that area.

[0077] The base film (10) supports the electrode portion (40) and the encapsulation film (30). That is, the electrode portion (40) and the encapsulation film (30) are sequentially laminated on the base film (10). Here, the base film (10) is formed of an insulating material that does not conduct electricity. For example, polyimide (PI) may be used, but is not necessarily limited thereto.

[0078]

[0079] The adhesive film (20) is an adhesive member that bonds one side of the base film (10) to the outer surface of the stent-graft (1). This adhesive film (20) can be heat-bonded and has adhesive properties as it melts. At this time, it must not seep into the fabric of the stent-graft (1). Also, since the base film (10) is bonded while the stent-graft (1) is unfolded and the stent-graft (1) is inserted into the blood vessel while folded and then unfolded again, the base film (10) contracts and stretches during this process, so the adhesive film (20) must be formed of a material that can maintain flexibility even after it has been melted and hardened.

[0080] In one embodiment, the adhesive film (20) may include a heat-bonded sheet and a heat-bonded layer.

[0081] A heat-sealable sheet can be formed by arranging a plurality of fibers composed of a first heat-sealable polymer. The first heat-sealable polymer may be nylon, polyester, etc., but is not necessarily limited thereto; there are no specific restrictions as long as it is a polymer that can be melted and spun into a fiber form and possesses adhesive properties when cured after heat melting. Since the fibers of this first heat-sealable polymer material are woven to form a heat-sealable sheet, the heat-sealable sheet may have a plurality of pores. Here, it is preferable that the fibers of the first heat-sealable polymer material be spun into a nonwoven structure, but is not necessarily limited thereto.

[0082] The heat-sealable layer is a layer formed by coating a powder composed of a second heat-sealable polymer onto the surface of a heat-sealable sheet. The powder may include polyester powder, polyamide (PA) powder, polyurethane hot melt adhesive powder, ethylene vinyl acetate hot melt adhesive (EVA) powder, etc., but is not necessarily limited thereto; there are no specific restrictions as long as the powder is made of a polymer material that is a thermoplastic polymer capable of being adhesive and powderable. By coating such powder onto the heat-sealable sheet to form a layer, the porosity of the heat-sealable sheet can be reduced and the adhesiveness improved.

[0083]

[0084] The encapsulation film (30) is a member formed in a shape corresponding to that of the base film (10). That is, the encapsulation film (30) also has a predetermined length and is provided with a zigzag curved portion along its length direction. One side of the encapsulation film (30) of this shape is laminated to the other side of the base film (10). Therefore, just like the base film (10), it can flexibly stretch and contract when the stent-graft (1) is unfolded and folded.

[0085] When the stent-graft (1) is installed in the blood vessel, the encapsulation film (30) is placed at the outermost edge, so that in the event of blood leakage from the blood vessel, the leaked blood comes into contact with the other side of the encapsulation film (30).

[0086] An electrode portion (40) is disposed between the encapsulation film (30) and the base film (10), and the encapsulation film (30) can be formed of an insulating material so that the electrode portion (40) and the blood vessel do not conduct electricity. For example, polyimide (PI) can be used, similar to the base film (10), but is not necessarily limited thereto, and the same material as the base film (10) does not have to be used.

[0087]

[0088] The electrode portion (40) includes a first conductive line (41) and a second conductive line (42), and accumulates an electric charge between the first conductive line (41) and the second conductive line (42). Here, the capacitance has a characteristic that changes when blood comes into contact with the other side of the encapsulation film (30).

[0089] The first conductive line (41) is a conductive member formed in a shape corresponding to the zigzag shape of the base film (10). Likewise, the second conductive line (42) is also a conductive member formed in a shape corresponding to the first conductive line (41). Thus, the first conductive line (41) and the second conductive line (42) are formed in a zigzag shape identical to the base film (10) and the encapsulation film (30) and are placed between the base film (10) and the encapsulation film (30). Therefore, since the first conductive line (41) and the second conductive line (42) also have flexibility, they do not easily break when stretched and contracted, and can be easily reduced.

[0090] The first conductive line (41) and the second conductive line (42) are spaced apart at a predetermined distance, and when power is applied between the first conductive line (41) and the second conductive line (42), they have capacitance. Here, the first conductive line (41) can be used as a voltage application line, and the second conductive line (42) can be used as a ground line. Therefore, when power is applied through the first conductive line (41), an electric field is formed, and electric field lines are formed through the second conductive line (42), so that the first conductive line (41) and the second conductive line (42) can be seen as being coupled by parasitic capacitance. Here, when blood comes into contact with the other side of the encapsulation film (30), a portion of the electric field lines moves into the blood, causing a change in capacitance. Depending on this change in capacitance, blood leakage from the blood vessel where the stent-graft (1) is installed can be detected. Here, in order to form an electric field between the first conduction line (41) and the second conduction line (42) and to check the change in capacitance, the first conduction line (41) and the second conduction line (42) are electrically connected to a separate circuit.

[0091] The first conductive line (41) and the second conductive line (42) may include gold (Au), iridium (Ir), platinum (Pt), etc., and their respective widths (L, FIG. 5) may be 180 to 220 μm, and the spacing between them (G, FIG. 5) may be 20 to 40 μm, but are not necessarily limited thereto. In addition, the curved shape of the first conductive line (41) and the second conductive line (42) may be formed, for example, in a shape where an arc with a radius of 0.8 to 1.2 mm and an angle of 120° to 180° is repeated while being inverted vertically. In addition, the total length may be 50 to 60 mm, but may vary depending on the type of stent-graft (1) to be attached and the attachment location. Here, a pair of first conductive lines (41) and second conductive lines (42) disposed on the base film (10) may surround the entire outer surface of the stent-graft (1) as a whole, that is, one first conductive line (41) and a second conductive line (42) may surround the entire outer surface of the stent-graft (1), or may surround the stent-graft (1) continuously in multiple pairs along the circumferential direction. The case of surrounding in multiple pairs refers to a case where multiple first conductive lines (41) are arranged continuously along the circumferential direction of the stent-graft (1), and multiple second conductive lines (42) are arranged continuously to correspond one-to-one with the multiple first conductive lines (41). In this case, the blood leakage detection range is partitioned along the circumferential direction of the stent-graft (1), so the leakage location can be subdivided and detected.

[0092] In addition, to induce a change in capacitance due to blood, it is desirable to apply an alternating voltage having a frequency in the range of 50 to 100 KHz. Here, the height of the electric field may be up to 200 to 230 μm, but is not necessarily limited thereto.

[0093]

[0094] Since the stent-graft attached blood leakage detection sensor (100) according to the present invention is placed in the area of ​​the stent-graft (1) that is in contact with a blood vessel, it must be formed as thin as possible so that blood leakage does not occur due to the sensor. To this end, the thickness of the base film (10) and the encapsulation film (30) may be 5 μm or less, the thickness (T, FIG. 5) of the first conductive line (41) and the second conductive line (42) may be 0.5 μm or less, and the thickness of the adhesive film (20) may be 50 μm or less, but is not necessarily limited thereto.

[0095] Additionally, the stent-graft attached blood leak detection sensor (100) according to the present invention may be disposed on the outer surface of a stent-graft (1) at least one at a predetermined interval along the longitudinal direction of the stent-graft (1). When a plurality of stent-graft attached blood leak detection sensors (100) according to the present invention are spaced apart in parallel, the blood leak detection area is partitioned along the longitudinal direction, so blood leak can be detected in more subdivided sections.

[0096]

[0097] FIG. 6 is a diagram illustrating a stent-graft attached blood leak detection sensor according to another embodiment of the present invention, FIG. 7 is a diagram illustrating an electrode portion of a stent-graft attached blood leak detection sensor according to another embodiment of the present invention, and FIG. 8 is a diagram illustrating a circuit connection relationship of an electrode portion of a stent-graft attached blood leak detection sensor according to another embodiment of the present invention.

[0098] As illustrated in FIGS. 6 to 8, a stent-graft attached blood leakage detection sensor (100) according to another embodiment of the present invention comprises the aforementioned base film (10), adhesive film (20), encapsulation film (30), and electrode portion (40), wherein the electrode portion (40) additionally comprises a third conductive line (43) in addition to a first conductive line (41) and a second conductive line (42).

[0099] The third conductive line (43) is a conductive member that is shorter in length than the second conductive line (42). The third conductive line (43) is formed in a shape corresponding to one region (refer to region a1 in FIG. 7) of the longitudinal region (refer to region a in FIG. 7) of the zigzag-shaped second conductive line (42). The third conductive line (43) formed in this way is arranged to be spaced apart from the second conductive line (42) by a predetermined distance. Here, the third conductive line (43) is used as a voltage application line, and when voltage is supplied from a separately provided circuit, it is coupled with the second conductive line (42), which is a ground line, by parasitic capacitance, and causes a change in capacitance when blood comes into contact with the other side of the encapsulation film (30). In this case, when the third conductive line (43) is arranged parallel to one area of ​​the second conductive line (42) (refer to area a1 in FIG. 7), the blood leakage detection area is distinguished from the area where the third conductive line (43) is not arranged (refer to area a2 in FIG. 7). Therefore, by arranging the third conductive line (43), the blood leakage detection area along the circumferential direction of the stent-graft (1) can be partitioned.

[0100] Overall, when a plurality of stent-graft attached blood leak detection sensors (100) according to another embodiment of the present invention are spaced apart along the longitudinal direction of the stent-graft (1), the blood leak detection area can be partitioned along the longitudinal direction. Additionally, a plurality of pairs of first conductive lines (41) and second conductive lines (42) are arranged continuously along the arc direction of the stent-graft (1), so that the blood leak detection area can be partitioned along the arc direction. Furthermore, by additionally arranging a third conductive line (43), the blood leak detection area can be distinguished along the arc direction of the stent-graft (1) into an area where the third conductive line (43) is arranged and an area where it is not arranged.

[0101]

[0102] In summary, the stent-graft attached blood leak detection sensor according to the present invention allows for the monitoring of blood leakage outside the stent that may occur after Endovascular Aneurysm Repair (EVAR) in daily life, without the need for radiological examinations. Visiting a hospital for MRI or CT scans is physically and financially cumbersome and requires frequent intervals; thus, by the time an endoleak is detected, the situation is already serious, and early detection can contribute to reducing mortality. Furthermore, the sensor according to the present invention has a thin thickness that does not affect the movement of the aorta or blood flow, and because it is attached to the stent and inserted together, it can be installed non-invasively. Because it is formed in a zigzag shape, the stent and the sensor can be loaded together into a catheter, and it remains undamaged despite repeated compression and expansion. The sensor according to the present invention operates based on the detection of capacitance by genetic materials such as blood. Compared to conventional sensors that determine leakage simply by generating a closed circuit, sensors installed within the aorta must be in close contact with the blood vessel; therefore, detecting capacitance, which varies in value depending on the substance, is a suitable principle as it allows for the differentiation of blood within the vessel. Furthermore, the electrode section is designed in a long, linear shape, enabling installation that surrounds the stent and detecting blood leakage that may occur anywhere within the cylindrical structure. By varying the length of the electrode to divide the blood leakage detection area, the location of the leak can be identified through a signal that defines the position of the leak. The design of the thin electrode section allows for securing the necessary space for adhesion in the narrow gap between the stent and the graft, which is inconsistent due to the nature of stents manufactured to fit the patient's aortic aneurysm structure. Additionally, the thin width of the electrode section results in a low electric field height, ensuring it is not affected by arterial blood inside the stent and enabling differentiation between the minimal blood present between the stent and the vessel and the leaked blood.Endoleaks, which leak only in certain areas, are suitable for use in capacitance change detection that is highly affected by area.

[0103]

[0104] Hereinafter, a blood leak detection wireless system according to an embodiment of the present invention is described. FIG. 9 is a configuration diagram of a blood leak detection wireless system according to an embodiment of the present invention, and FIG. 10 is a diagram illustrating wireless data transmission of a blood leak detection wireless system according to an embodiment of the present invention. FIG. 11 is an image showing the connection relationship (a) between the blood leak detection sensor and the planar inductor and the reader coil (b) shown in FIG. 9, and FIG. 12 is a diagram comparing the shape and size of the planar inductor and the reader coil shown in FIG. 11.

[0105] As illustrated in FIGS. 9 and 10, a wireless blood leak detection system according to an embodiment of the present invention includes a blood leak detection sensor (100) attached to the outer surface of a tubular stent-graft (1) and acting as a variable capacitor whose capacitance changes in response to blood leaking from a blood vessel, a planar inductor (200) attached to the outer surface of the stent-graft (1) and electrically connected to the blood leak detection sensor (200) to form an LC resonant circuit whose resonant frequency changes in response to a change in capacitance, and a reader coil (310) disposed outside the body and magnetically coupled with the planar inductor (200), and a wireless reader unit (300) that wirelessly detects a change in resonant frequency.

[0106]

[0107] As previously described, the blood leak detection sensor (10) is a sensor attached to the outer surface of the tubular stent-graft (1) to detect blood leakage. This blood leak detection sensor (10) acts as a variable capacitor whose capacitance changes in response to blood leaking from a blood vessel. As the blood leak detection sensor (10) has been previously described, a detailed explanation thereof is omitted.

[0108]

[0109] Referring to FIGS. 11 and 12, the planar inductor (200) is an inductor that is electrically connected to the blood leak detection sensor (100) to form an LC resonant circuit. This planar inductor (200) is attached to the outer surface of the stent-graft (1). As described above, when the blood leak detection sensor (100) comes into contact with blood, the capacitance of the blood leak detection sensor (100) changes. That is, the blood leak detection sensor (100) operates as a variable capacitor whose capacitance changes according to contact with blood. The blood leak detection sensor (100) operating as a variable capacitor has its capacitance changed according to blood contact and the area covered by the blood.

[0110] A planar inductor (200) is electrically connected to a blood leak detection sensor (100) that operates as a variable capacitor to form an LC resonant circuit. Therefore, when blood comes into contact with the blood leak detection sensor (100), the capacitance of the blood leak detection sensor (100) changes, and a change in the resonant frequency is induced in response to the change in capacitance. Since the blood leak detection sensor (100) operates as a variable capacitor whose capacitance changes according to the blood covering, the resonant frequency can be changed sensitively. A passive sensor is formed when an LC resonant circuit is formed by the blood leak detection sensor (100) and the planar inductor (200). Accordingly, the sensor can operate stably for a long period after implantation without a battery or a separate power supply.

[0111] The planar inductor (200) not only forms an LC resonant circuit with the blood leak detection sensor (100) but also enables wireless communication by forming a strong magnetic coupling with the reader coil (310). Since the planar inductor (200) is attached to the stent-graft (1) and implanted into the body's blood vessels, it is necessary to optimize its shape and size for effective wireless communication with the outside. Accordingly, the planar inductor (200) can be formed as an elongated rectangular shape along the length of the stent-graft (1). Here, the width (W) of the planar inductor (200) may be 20 to 40 mm, and the length may be 80 to 120 mm. However, the width (W) and length (L) of the planar inductor (200) are not necessarily limited to these values.

[0112] Additionally, the planar inductor (200) may be placed in a semicircular area of ​​the outer surface of the stent-graft (1). The placement of the planar inductor (200) on the stent-graft (1) is designed so that the magnetic field is not canceled out, and is devised for effective wireless communication.

[0113]

[0114] The wireless reader unit (300) wirelessly detects changes in resonance frequency caused by the blood leak detection sensor (100) and the planar inductor (200) from outside the body. Here, the wireless reader unit (300) includes a reader coil (310). The reader coil (310) is placed outside the body and is magnetically coupled with the planar inductor (200). This reader coil (310) is placed outside the body and performs wireless communication with the aforementioned LC resonance circuit. Here, since the blood leak detection sensor (100) and the planar inductor (200) constituting the LC resonance circuit are integrated into the stent-graft (1) and inserted into the blood vessel, it is necessary to form a strong magnetic coupling with the planar inductor (200) placed at a depth of 60 to 70 mm or more. To form such magnetic coupling, the shape and size of the reader coil (310) are important. Accordingly, the leader coil (310) can be formed with a shape corresponding to the planar inductor (200), that is, the same shape as when the planar inductor (200) is rectangular. At this time, the size of the leader coil (310) can be expanded to 1.5 to 2.5 times the size of the planar inductor (200). In this case, stable wireless communication without signal attenuation becomes possible.

[0115] When blood covers the blood leak detection sensor (100), the resonant frequency shifts distinctly in MHz units depending on the covering area, so the wireless reader (300) can detect minute blood leaks.

[0116] The wireless reader unit (300) may further include an analyzer (320) for detecting blood leakage. The analyzer (320) is a device capable of measuring S-parameters. For example, the analyzer (320) may be a network analyzer. This analyzer (320) generates an RF signal and transmits it to the reader coil (310), and the LC resonant circuit can generate a resonant response through magnetic coupling with the reader coil (310). By the reader coil (310) receiving this resonant response and transmitting it to the analyzer (320), the analyzer (320) can stably measure the change in resonant frequency, i.e., the S11 response, and based on this, monitor blood leakage over a long period.

[0117]

[0118] In summary, according to the blood leak detection wireless system of the embodiment of the present invention, a passive LC resonant circuit capable of operating without a battery or a separate power supply is provided, so that it can be used stably for a long period even after implantation of a stent-graft (1) incorporating a blood leak detection sensor (100). In addition, by integrating a large-area planar inductor (200) into the outer wall of the stent and applying a reader coil (310) larger than that to the outside, stable wireless signal transmission and detection are possible even in the aorta located deep inside the human body. Furthermore, since the resonant frequency changes distinctly in MHz units depending on the area covered by blood over the blood leak detection sensor (100), even minute amounts of blood leakage can be sensitively detected. Accordingly, patients can check for the occurrence of endo-leakage in real time using an external reader without visiting a hospital, thereby significantly reducing dependence on high-cost, high-risk imaging equipment such as CT or X-ray. Furthermore, the present invention can be extended not only to the detection of aortic aneurysm endoleaks but also to the monitoring of various biomarkers such as pressure, body fluids, and blood flow, making it highly clinically useful.

[0119]

[0120] Meanwhile, when detecting a blood leak using the aforementioned blood leak detection system, a stent-graft (1) equipped with a blood leak detection sensor (100) can be placed through a blood vessel at the location of the aortic aneurysm to detect the blood leak.

[0121] Here, the Type 1 endoleak can be continuously monitored using the aforementioned blood leak detection system.

[0122]

[0123] The following describes a method for manufacturing a stent-graft-attached blood leak detection sensor according to an embodiment of the present invention. As the stent-graft-attached blood leak detection sensor has been described previously, details that overlap with it will be omitted or described only briefly.

[0124] FIG. 13 is a flowchart of a method for manufacturing a blood leak detection sensor according to an embodiment of the present invention, FIG. 14 is a diagram illustrating the adhesive film manufacturing step illustrated in FIG. 13, and FIG. 15 is a diagram illustrating the electrode structure manufacturing step illustrated in FIG. 13. Additionally, FIG. 16 is a diagram illustrating the electrode structure transfer step illustrated in FIG. 13, and FIG. 17 is a diagram illustrating the electrode structure adhesion step illustrated in FIG. 13.

[0125] As illustrated in FIG. 13, a method for manufacturing a stent-graft attached blood leakage detection sensor according to an embodiment of the present invention comprises the steps of: manufacturing an adhesive film by coating a powder made of a second heat-sealable polymer onto a porous heat-sealable sheet formed by arranging a plurality of fibers made of a first heat-sealable polymer (S100); manufacturing an electrode structure on a substrate, comprising a base film having a continuous curved portion in a zigzag shape along the longitudinal direction, two or more conductive lines formed in a shape corresponding to the zigzag shape of the base film and arranged parallel to each other, and a encapsulation film formed in a shape corresponding to the base film and laminated on the base film with the conductive lines in between (S200); and separating the electrode structure from the substrate and transferring it to one side of the adhesive film (S300).

[0126] As described above, the stent-graft attached blood leakage detection sensor manufactured according to the present invention has a structure in which an adhesive film, a base film, an electrode portion, and an encapsulation film are sequentially laminated, and the electrode portion may additionally include a third conductive line in addition to a first conductive line and a second conductive line.

[0127] Specifically, a method for manufacturing a stent-graft attached blood leak detection sensor includes an adhesive film manufacturing step (S100), an electrode structure manufacturing step (S200), and an electrode structure transfer step (S300).

[0128] The adhesive film manufacturing step (S100) is a process for manufacturing an adhesive film by forming a heat-seal layer on a porous heat-sealable sheet. The porous heat-sealable sheet is formed by arranging a plurality of fibers composed of a first heat-sealable polymer, and the adhesive film can be manufactured by coating a powder composed of a second heat-sealable polymer on its surface.

[0129] With reference to FIG. 14, a heat-sealable sheet can be manufactured by melting a first heat-sealable polymer and spinning the melted first heat-sealable polymer into a fiber form. The first heat-sealable polymer may be nylon, polyester, etc., but is not necessarily limited thereto. Here, it is preferable that the fibers of the first heat-sealable polymer material be spun into a nonwoven structure, but is not necessarily limited thereto.

[0130] Next, a coating solution in which powder is dissolved is coated onto the surface of the manufactured heat-bonded sheet to form a heat-bonded layer. The powder may include polyester powder, polyamide (PA) powder, polyurethane hot melt adhesive powder, ethylene vinyl acetate hot melt adhesive (EVA) powder, etc., but is not necessarily limited thereto. By coating the heat-bonded sheet with such powder to form a layer, the porosity of the heat-bonded sheet can be reduced and the adhesiveness improved.

[0131] The electrode structure manufacturing step (S200) is a process for forming a structure (hereinafter referred to as an 'electrode structure') in which a base film, an electrode portion, and an encapsulation film are sequentially laminated. Such an electrode structure can be formed on a substrate and transferred to an adhesive film.

[0132] For example, as shown in FIG. 15, a sacrificial layer may be formed on a substrate, a base film may be formed on the sacrificial layer using a first insulating polymer, conductive lines may be arranged on the base film, and then an encapsulation film may be formed using a second insulating polymer. The sacrificial layer may use materials such as PMMA, but is not necessarily limited thereto; there are no specific restrictions as long as it can be removed by physical or chemical means without affecting the electrode structure. Here, the conductive lines may be composed of three lines spaced apart from each other, with the central conductive line serving as a ground line and the other two lines serving as voltage application lines. In this case, one voltage application line and the ground line may be made long, while the other voltage application line may be made short, thereby allowing the blood leakage detection area to be distinguished. These conductive lines may be formed through methods such as metal deposition, lithography, or etching. Here, the base film and the encapsulation film may use polyimide (PI), but are not necessarily limited thereto. In addition, the conductive line may include gold (Au), iridium (Ir), platinum (Pt), etc., but is not necessarily limited to these.

[0133] The electrode structure transfer step (S300) is a process of transferring an electrode structure manufactured on a substrate to one side of an adhesive film.

[0134] Referring to FIG. 16, for example, the sacrificial layer on the substrate is first removed. In the case of a PMMA sacrificial layer, it can be removed by immersing it in an acetone solution and heating it. Next, the first adhesive sheet is attached to one side of the electrode structure to separate it from the substrate. The first adhesive sheet is a material such as a tape having an adhesive surface that can be removed by an organic solvent. By attaching this first adhesive sheet to a encapsulation film placed on the outermost edge of the electrode structure, the electrode structure can be separated from the substrate from which the sacrificial layer has been removed. Then, the other side of the electrode structure attached to the first adhesive sheet is placed on one side of the adhesive film, and the second adhesive sheet is placed on the other side of the adhesive film on which the electrode structure is placed to form a multilayer structure. Here, the multilayer structure is a structure in which the first adhesive sheet, the electrode structure, the adhesive film, and the second adhesive sheet are laminated. Here, the second adhesive sheet is also a material such as a tape having an adhesive surface that can be removed by an organic solvent. When the multilayer structure is formed in this way, the multilayer structure is heated to transfer the electrode structure onto the adhesive film. Since the adhesive film is made of a heat-sealable polymer, the electrode structure can be adhered to the adhesive film by applying heat. Then, the heated multilayer structure is immersed in an organic solvent. By doing so, the first adhesive sheet and the second adhesive sheet are removed from the adhesive film, and the electrode structure can be transferred onto the adhesive film.

[0135] Meanwhile, when the electrode structure is transferred to the adhesive film, an electrode structure bonding step (S400) can be performed. The electrode structure bonding step (S400) is a process of bonding the electrode structure to the outer surface of the stent-graft. Here, the electrode structure is positioned to surround the outer surface of the tubular stent-graft, and the adhesive film on which the electrode structure is transferred is heated.

[0136] Referring to FIG. 17, for example, ethanol is first sprayed to attach the other side of an adhesive film, on which an electrode structure is transferred to one side, to the outer surface of a stent-graft. Then, the adhesive film is heated to melt, so that the electrode structure can be attached to the outer surface of the stent-graft.

[0137]

[0138] [Explanation of the symbol]

[0139] 1: Stent-graft 10: Base film

[0140] 20: Adhesive film 30: Encapsulation film

[0141] 40: Electrode section 41: First conduction line

[0142] 42: 2nd Challenge Line 43: 3rd Challenge Line

[0143]

[0144] The present invention will be explained in detail below through experimental examples.

[0145]

[0146] Experimental Example 1

[0147] 1.1. Manufacturing of Adhesive Films

[0148] A heat-bonded fabric was prepared by melting polyester and spinning it into a nonwoven structure. Then, as shown in Fig. 10, polyamide (PA) powder was added to acetone at a ratio of 1:2 w% and heated to 120 to 180°C for 20 to 30 minutes to dissolve it. The dissolved powder was spin-coated onto the polyester heat-bonded fabric at 2000 rpm for 30 seconds. After spin-coating, a heat-bonded layer coated as a single layer was obtained, filling the empty spaces of the heat-bonded fabric.

[0149]

[0150] 1.2. Fabrication of Electrode Structure

[0151] Referring to Fig. 15, an electrode structure was fabricated on a glass substrate with a PMMA sacrificial layer in between. Polyimide (PI) was used for the base film and the encapsulation film, and gold (Au) was used as the material for the three conductive lines. The thickness of the base film and the encapsulation film is 2 μm each, and the thickness of the conductive lines is 0.2 μm. The width of the conductive lines is 200 μm, and the spacing between the conductive lines is 30 μm. The conductive lines were formed by bending into a repeating arc shape with a radius of 1 mm and an angle of 120°.

[0152] A base film was first fabricated by melting and curing PI, conductive lines were arranged on the base film by metal deposition, and then an encapsulation film was laminated on the base film by melting and curing PI.

[0153]

[0154] 1.3. Electrode Structure Transfer

[0155] With reference to Fig. 16, the substrate formed with the electrode structure fabricated in Experimental Example 1.2 was immersed in an acetone solution and heated at 150°C for about 1 hour to remove the sacrificial layer. Then, the electrode structure was peeled off using a tape having an adhesive surface that can be removed by acetone. Next, the lower surface of the base film of the electrode structure was placed on the adhesive film fabricated in Experimental Example 1.1, and another tape having an adhesive surface that can be removed by acetone was placed on the encapsulation film of the electrode structure. The tape / adhesive film / tape layers were stacked, and the three layers were bonded by applying heat of 100 to 150°C and hot air of 150 l / m using a blow soldering iron. The three bonded layers were immersed in an acetone solution at room temperature, and the electrode structure was transferred to the adhesive film by removing the tape.

[0156]

[0157] 1.4. Attachment of Electrode Structure

[0158] With reference to Fig. 17, an adhesive film having the electrode structure prepared in Experimental Example 1.3 transferred onto it was placed on the outer surface of a stent-graft, and ethanol was sprayed to create surface tension so that the adhesive film would adhere to the surface of the stent. Heat of 100 to 150°C and hot air of 150 l / m were applied using a blow soldering iron to melt the adhesive film and attach the electrode structure.

[0159]

[0160] 1.5. Performance Evaluation of Adhesive Films

[0161] In the case of conventional liquid adhesives, they can be absorbed into the fabric of the stent-graft and seep into the arterial blood passing through, potentially affecting it. Furthermore, after curing, they can harden, which may affect the movement of the stent or cause damage to the fabric. Since the sensor according to the present invention is attached to a commercially available stent and inserted into the blood vessel, when the sensor is attached to the stent, it must be installed folded inside the catheter and then stably unfolded, and there should be no rigid parts due to the irregular movement of the aorta. Therefore, the adhesive used to attach the electrode must maintain flexibility even after curing and allow a thin electrode of approximately 5 μm to be transferred onto the surface of the stent without loss. Accordingly, the flexibility, adhesion, and absorption into the stent fabric of the adhesive film prepared in Experimental Example 1.1 were evaluated, and the results are shown in FIG. 18. FIG. 18 shows the results of evaluating the flexibility (a), adhesion (b), and absorption into the stent fabric (c) of the adhesive film prepared according to Experimental Example 1.

[0162] After heat-melting and curing the adhesive film prepared in Experimental Example 1.1, a bending strain test was performed with another adhesive, and it was found that it maintained excellent flexibility even after curing (see Fig. 18 (a)). In addition, a pulling strain test was also performed, and the results confirmed that the adhesive strength increased as the heat-bonding powder was spin-coated onto the heat-bonding fabric (see Fig. 18 (b)). Furthermore, the pores of the heat-bonding fabric were reduced, and it was not absorbed into the stent fabric even in the heat-melted state (see Fig. 18 (c)).

[0163] In addition, a compression expansion test was performed more than 1,000 times on a stent-graft to which the electrode structure fabricated in Experimental Example 1.4 was attached, and the results are shown in Fig. 19. Fig. 19 shows the results of the compression expansion test on the blood leak detection sensor fabricated in Experimental Example 1. With reference to Fig. 19, it can be seen that the adhesive film maintains its adhesive strength even after multiple compression expansions of the stent, and maintains its flexibility even after curing, so it does not hinder the movement of the stent.

[0164]

[0165] 1.6. Evaluation of Blood Leak Detection Performance

[0166] When an alternating current voltage is applied to a conductive line, a change in capacitance occurs depending on the permittivity and frequency of the material. To distinguish between blood vessels and blood, blood, blood vessels, proteins, lipids, and a buffer solution (pbs) were used as comparison groups, and the change in capacitance according to frequency for each material was evaluated, and the results are shown in FIG. 20. FIG. 20 is a graph showing the capacitance (a) according to the frequency of the applied voltage for each target material of the blood leak detection sensor manufactured in Experimental Example 1, and the change in capacitance due to blood leakage (b).

[0167] Referring to FIG. 20, the change in capacitance by material in the frequency range of 50 to 100 KHz is clearly distinguished, and thus, blood leakage can be detected in the present invention by applying an alternating voltage having a frequency in that range. The structure of a thin and long conductive line has a higher influence per unit area than the height of the material, which is suitable for the characteristics of Type 1 endoleaks where blood leaks from a part of the stent rather than the entire stent.

[0168]

[0169] Experimental Example 2

[0170] 2.1. Overview

[0171] This study proposes a wireless real-time blood leakage monitoring technology capable of continuously detecting Type 1 endoleaks. In this system, an ultrathin capacitive sensor array with a filamentary serpentine structure is embedded at the proximal attachment site of a conventional stent-graft and coupled to an inductive-capacitance (LC) resonant circuit. Positioned between the stent-graft and the vessel wall, the sensor provides precise and localized information regarding blood flow and blood leakage based on the capacitance principle that distinguishes between blood and the vessel wall.

[0172] Since the sensor is ultra-thin and flexible, it does not affect the mechanical performance of the stent and maintains structural integrity even after repeated crimping, folding, and deployment processes. Additionally, because the sensor is barely exposed to blood flow, it can detect blood leaks early while maintaining the key functions of Endovascular Aneurysm Repair (EVAR). When a blood leak occurs, the sensor's capacitance changes, which in turn alters the resonant frequency of the LC circuit. This frequency change can be wirelessly detected through inductive coupling with a reader located outside the patient's abdomen, enabling non-invasive and continuous monitoring of the endoleak status.

[0173] The clinical applicability of this technology has been demonstrated through computer simulations, experiments using blood samples, evaluation of electrode cytotoxicity, and porcine aorta (in vitro) model experiments. Accordingly, the system based on this technology is evaluated to be able to revolutionize the paradigm of endoleak monitoring after EVAR procedures by detecting endoleaks in real time, reduce the risk of aneurysm rupture, and improve patient prognosis.

[0174]

[0175] 2.2. Materials and Methods

[0176] 2.2.1. Fabrication of Thin-Film Capacitive Sensors

[0177] First, a sacrificial layer was formed by spin-casting polymethyl methacrylate (PMMA, 950 PMMA A 2, Kayaku Advanced Materials) onto a glass substrate (MATSUNAMI, thickness 1.2 ~ 1.5 mm) and curing it at 180°C for 3 minutes. Subsequently, a base film was formed by spin-coating polyimide (PI) (Poly(pyromellitic dianhydride-co-4,4'-oxydianiline), amic acid solution, Sigma-Aldrich), baking it at 150°C for 5 minutes, and then fully crosslinking it at 350°C for 1 hour under vacuum. Afterward, a gold (Au) layer with a thickness of 200 nm was deposited on the base film using an electron-beam evaporator. A gold electrode structure was formed through photolithography (AZ 5214E, AZ Electronic Materials) and wet metal etching processes. A second PI layer was deposited using the same method to form an upper encapsulation film covering the gold electrode. Subsequently, a copper mask was patterned using photolithography and Cu wet etching, and a base film and encapsulation film structure were formed using a reactive ion etcher (Samco Inc.). The PMMA sacrificial layer was removed by immersion in acetone (Samchun Pure Chemical) at 150°C for 1 hour.

[0178]

[0179] 2.2.2. Manufacture of Heat-Sealed Films

[0180] Dulbecco's phosphate-buffered saline (DPBS, Sigma-Aldrich) and tetrahydrofuran (THF, Biograde, 99.8%, Thermo Scientific) were mixed in a weight ratio of 1:3. Co-polyamide (Co-PA) fabric (SINGER Iron-On Fusing Web, Singer) was dissolved in this mixed solvent and heated at 130°C for 10 minutes. The resulting solution was spin-coated onto a glass substrate at 2000 rpm for 30 seconds, and then cooled to 20–25°C.

[0181]

[0182] 2.2.3. Separation and Transfer of Thin-Film Capacitive Sensor

[0183] After removing the PMMA sacrificial layer, the electrode was separated from the substrate using water-soluble tape (AQUASOL, Aquasol Corporation). Then, a heat-sealed film was attached to the bottom of the electrode by applying heat. Subsequently, the water-soluble tape was removed with deionized water. The electrode was temporarily fixed by applying isopropyl alcohol (IPA) to the stent fabric, and the bonding process was completed by melting the heat-sealed film at 110°C using a flow gun.

[0184]

[0185] 2.2.4. Adhesive properties of stent fabric: Flexibility, adhesive strength, and permeability

[0186] Polydimethylsiloxane (PDMS) (Sylgard 184, USA, base-to-curing agent ratio 10:1), Super Glue (Mxbon® 224902M), tissue liquid adhesive (Mastisol liquid adhesive), and skin spray adhesive (Skinister Medical Adhesive Spray) were used as control groups. The adhesive films developed in this study were temporarily attached to stents using IPA and then permanently bonded through heat treatment. PDMS was applied to the fabric via spin coating (2000 rpm, 30 seconds) and cured in an 80°C oven for 2 hours. Super Glue and the medical adhesive were similarly spin-coated (2000 rpm, 30 seconds) and cured at 30°C for 1 hour. After the adhesives were cured, the flexural strength and adhesive strength of the adhesives were measured using a Universal Testing Machine (AMETEK). Bending strength was measured by applying a compressive force of 50 mm / min to a double-layer stent fabric (20 mm × 60 mm) with adhesive attached. For the adhesive strength test, two stent fabrics (20 × 60 mm) were bonded with adhesive, and a 180° peel test was performed with a force of 5 N and a speed of 60 mm / min. To evaluate permeability, the adhesive was applied to the outer surface of the fabric, and the inner surface of the fabric was observed using a scanning electron microscope (SEM).

[0187]

[0188] 2.2.5. Preparation of Gelatin Gel for Phantom Production

[0189] A blood-mimicking gel with an electrical conductivity of 0.6 S / m and a blood vessel-mimicking gel with an electrical conductivity of 0.2 S / m were prepared according to the conductivity concentration formula. Gelatin powder (Herbnare) and sodium chloride (NaCl, purity ≥99%, Sigma-Aldrich) were dissolved in deionized water at 90°C for 30 minutes on a hot plate. Subsequently, the solution was placed in a glass mold with a depth of 2 mm and cooled to 20–25°C to form a gel.

[0190]

[0191] 2.2.6. Sensor Thickness Measurement and Repetitive Mechanical Deformation Test

[0192] The thickness of the sensor electrode, including the metal layer (conductive line, electrode part), PI base film, and encapsulation film, was measured using a surface profilometer (Alpha Step, Bruker). Compression and expansion stress tests on the sensor electrode were performed more than 1,000 times using a stent-crimping device.

[0193]

[0194] 2.2.7. Quantitative Analysis of Capacitance in Blood Vessels and Blood Phantoms

[0195] The capacitance of a coplanar capacitor was measured using a probe station in the frequency range of 0 to 200 kHz. Capacitance values ​​were measured for equal volumes of blood, blood vessels, phantom materials (blood vessels and blood, 10 × 15 × 2 mm), albumin (Officeahn), PBS (Hyunil Lab-Mate), and fat (Hanyang Sangsa). Performance testing of the sensor electrodes was performed at a frequency of 100 kHz with the electrodes covered with vascular gelatin to replicate the in vivo environment. The residual blood height in the blood phantom was controlled using molds with heights of 50 µm, 100 µm, and 200 µm. To change the blood flow velocity, the electrode substrate was tilted at 0°, 45°, and 90°, and 1 ml of blood was injected into a 2 mm wide PDMS (10:1) channel to achieve blood flow velocities of 72 µm / s, 172 µm / s, and 860 µm / s, respectively.

[0196]

[0197] 2.2.8. Thickness-dependent blood leakage analysis

[0198] A phantom simulating the elastic modulus of a blood vessel (0.5 MPa) was fabricated by mixing PDMS and Ecoflex in a 2:1 ratio. The mixture was poured into a 3 mm × 100 mm × 100 mm mold and cured in a vacuum oven at 70°C for 2 hours. To fabricate an electrode with a total thickness of 1000 µm, 20 layers of PI tape (Kapton®) with a thickness of 50 µm were laminated. Subsequently, electrodes of various thicknesses (size: 1.5 mm × 25 mm) were attached to the stent. The blood vessel phantom was fabricated in a cylindrical shape capable of enveloping the stent and the electrodes attached thereto.

[0199] To reproduce the endoleak phenomenon, 0.1 ml of blood was injected into the space between the stent and the vascular phantom using a syringe. For the blood pressure environment, a vinyl tube was inserted into the stent, and an internal pressure of 80 to 120 mmHg was applied at 60 to 100 beats / min (bpm) using an air pump.

[0200]

[0201] 2.2.9. Long-term Stability Assessment

[0202] After completely immersing a stent with a sensor electrode attached in a PBS solution at 37°C, the capacitance was measured every 3 days for up to one month. Before measuring the capacitance, the stent was removed from the PBS solution and dried at approximately 20–25°C for about 30 minutes. The measuring tip of the probe station was connected to the contact pad of a short electrode pair with a length of 13.5 mm, and after obtaining capacitance data in air, a blood phantom (10 mm × 5 mm) was placed on the electrode pair and the capacitance for blood was measured.

[0203]

[0204] 2.2.10. Toxicity Assessment of Electrode Sensors and Adhesives

[0205] Human umbilical vein endothelial cells (HUVECs) were cultured in EGM-2 medium (LONZA) at 5% CO₂ and 37°C. Conditioned media containing substances that could be eluted from PI and Co-PA films were prepared according to a standardized protocol (ISO 10993-12). Specifically, a film with an area of ​​600 mm² was placed in a 24-well plate, and the conditioned media was obtained after incubating in 1 ml of culture medium at 37°C for 24 hours. Cytotoxicity evaluation was performed by culturing HUVECs in the conditioned media for one day, and cells treated with normal EGM-2 medium were set as a control.

[0206] For live cell staining, cells were treated with DPBS containing calcein-AM (1:1000) for 10 minutes (using Invitrogen’s live / dead viability kit) and observed using a fluorescence microscope (Eclipse Ti2, Nikon; Tokyo, Japan). For the Thiazolyl Blue Tetrazolium (MTT) assay, cells were cultured with MTT solution at 37°C for 1 hour, and the resulting MTT formazan was extracted using DMSO (Sigma-Aldrich). Absorbance was measured at 550 nm using a microplate reader (Varioskan LUX, Thermo Scientific; Waltham, MA, USA).

[0207]

[0208] 2.2.11. Ex vivo experiment configuration

[0209] A sensor-integrated stent was inserted into the porcine aorta using a catheter (Medtronic). To replicate physiological blood pressure conditions, a vinyl tube was inserted into the stent, and an air pump was used to pressurize the stent to a range of 80–120 mmHg under conditions of a heart rate of 80–90 bpm. A syringe (Jibengao, 1 ml) was placed between the blood vessel and the stent, and 0.5 ml each of blood, protein, PBS, and lipids were sequentially injected at 37°C.

[0210]

[0211] 2.2.12. HFSS Simulation for Size Optimization of Planar Inductors and Lead Coils

[0212] A three-dimensional electromagnetic finite element analysis of the receiver (Rx) and transmitter (Tx) within the radio sensing system was performed using the commercial software ANSYS HFSS. Through this analysis, electromagnetic coupling, scattering coefficient (S11), and resonant frequency were determined. The Rx coil was modeled as a planar multi-winding (N = 3) rectangular inductance coil with dimensions set to L = 100 mm and W = 40 mm. The Tx coil was modeled with two representative designs: 1) a single-winding rectangular wire antenna (L = 110 mm, W = 50 mm), and 2) a single-winding rectangular wire antenna (L = 200 mm, W = 120 mm). Design optimization was performed to maximize the area coverage between the Rx and Tx radio systems and to enable sensing and detection at a maximum distance of 100 mm. The Rx antenna was designed to resonate in the 13.3 to 13.4 MHz range equipped with an external capacitor (C = 117 pF). An adaptive mesh (tetrahedral elements) and a spherical radiation boundary (radius 1000 mm) were applied for computational accuracy and fusion. The nonlinear relationship between the S11 parameter and the distance between the transmitter and receiver at resonance was calculated in the range of 60 to 100 mm. In the simulation, electrical conductivity (σ) was used as the Rx and Tx coil materials. Cu ) is 5.96 × 10 7 Copper (Cu) with an S / m value was used, and all simulations were performed in an air environment.

[0213]

[0214] 2.2.13. Wireless Monitoring Setup and Phantom Production

[0215] The sensor inductor, which forms an LC resonant circuit by connecting the electrode pair and the ACF film, was designed with dimensions of 30 mm × 100 mm, 13 turns, and a winding spacing and width of 400 μm. A rectangular copper coil (width 2.6 mm, diameter 60 mm × 200 mm) was connected to a network analyzer via an SMA connector (SubMiniature version A connector). S11 parameters were measured using the network analyzer at a sampling rate of 30 kHz within a frequency range of 1 to 20 MHz. A dielectric muscle phantom with a thickness of 70 mm was placed between the sensor inductor and the coil. The phantom was fabricated by pouring a 20% gelatin solution into a mold measuring 100 mm × 100 mm × 15 mm and allowing it to solidify at 4°C for 24 hours. After creating a circular hole with a diameter of 30 mm on the side of the solidified gelatin, the structure was stacked to a height of 150 mm.

[0216]

[0217] 2.3. Evaluation

[0218] 2.3.1. System Design, Functions, and Basic Concepts

[0219] FIG. 21 is a diagram illustrating a mechanism (a) for wirelessly detecting endoleaks using capacitance-based resonant frequency changes by inserting and mounting a capacitance-type blood leak detection sensor integrated into a stent of a blood leak detection wireless system according to Experimental Example 2, an arrangement (b) of a blood leak detection sensor for detecting the location and direction of blood leaks, a state of capacitance change due to endoleaks (left: low capacitance in normal state, right: increased capacitance due to blood leaks) (c), and a cross-sectional view of an electrode located between a blood vessel and a stent and an electric field distribution of a capacitance sensor (d).

[0220] Figure 21 (a) illustrates a medical application of the present blood leak detection wireless system. Here, the system enables non-invasive and wireless monitoring of potential blood leaks via an external reader. The sensor electrode specifically detects Type-Ia endoleaks occurring at the top of the proximal portion of the stent due to vascular degeneration or stent migration. Type-Ia endoleaks pose the highest risk among all endoleak types due to the high-pressure environment within the aneurysm, the absence of the possibility of natural resolution, and the need for immediate surgical intervention. In stent-graft procedures for aneurysm treatment, the proximal portion of the stent must be positioned to overlap the normal aortic region by at least 20 mm. This is because Type-Ia endoleaks typically occur in this region. As shown in Figure 21 (b), a circular capacitive sensor is arranged at the top of an EVAR stent (proximal average diameter 28 mm, length 150 mm) for blood leak detection. To enhance coupling efficiency with an external reader, a large induction coil is connected along the longitudinal direction of the stent. As shown in Figures 21 (c) and (d), proximal blood leakage causes a change in the sensor's capacitance, which alters the resonant frequency of the induction coil. This change in resonant frequency is detected through an external coil system placed near the abdominal skin, thereby enabling continuous wireless monitoring of blood leakage. The sensor consists of gold electrodes coated with biocompatible polyimide (PI). Since the sensor has a very thin and flexible structure with a total thickness of 5 μm and is attached to the stent, no additional surgical procedure is required during aortic stent insertion. The coplanar capacitor electrode surrounding the stent forms an electric field with four pairs of electrodes to detect blood leakage in both temporal and spatial terms.Depending on the location of the blood leak, capacitance changes occur in a single or both zones of the sensor, and the leak location can be localized by detecting these changes upon contact with blood through an array of four pairs of electrodes surrounding the proximal portion. The serpentine structure expands the capacitance area and increases the adhesion surface area, thereby enhancing adhesion stability. Furthermore, these structural features minimize the effective Young's modulus of the structure, enabling relatively free folding and deployment during catheter-based implantation.

[0221] These devices are strongly bonded to the fabric material of the stent-graft via a specially designed biocompatible adhesive. The adhesive was prepared by dissolving Co-PA material in a mixed solution of tetrahydrofuran (THF) and phosphate-buffered saline (PBS) and then forming it into a film. A micro-thick flexible electrode is attached to the stent fabric through the adhesive film. The adhesive film was manufactured by spin-casting the adhesive onto a silicon wafer and then separating it. Ethanol promotes initial adhesion between the adhesive film and the stent. The synthesized adhesive film exhibited excellent mechanical and adhesive properties and maintained structural integrity and functional stability even in dynamic vascular environments. The mechanical performance of the thin and flexible adhesive film ensures that the sensor does not impair the self-expanding function of the stent. Specifically, when deploying the stent using a commercial catheter, the sensor remained stably attached to the stent fabric without any damage or detachment, despite repeated compression and expansion deformations. Consequently, the combination of a thin chipless sensor and a flexible adhesive film becomes a key element that enables stable contact with blood vessels and continuous endoleak detection.

[0222] 2.3.2. Evaluation of the Sensor's Blood Leak Detection Performance

[0223] The coplanar capacitive sensor applied to the stent consists of two electrodes, an anode and a cathode, and measures the change in relative permittivity of the medium in contact with the sensor. As a result of analysis performed on air, blood, blood vessels, fat, albumin, etc., all samples showed a linear change in capacitance in the 50 to 200 kHz range. In particular, at 100 kHz, blood (19.8 pF) showed the highest capacitance, followed by albumin (16.3 pF), blood vessels (13.8 pF), and fat (1.7 pF).

[0224] In addition, blood and blood vessel phantom models were constructed and compared with actual samples to analyze the detection performance of the sensor. As a result, since the capacitance values ​​of the blood and blood vessel phantoms fall within the entire frequency range of 0 to 200 kHz, just like the actual subjects, the phantoms are suitable for quantitative experimental evaluation.

[0225] FIG. 22 is a graph showing the change in capacitance according to the thickness of leaked blood (a), the change in capacitance according to the area of ​​blood covering the sensor (b), and the time response of the sensor according to the initial blood flow velocity (c) of the blood leak detection sensor according to Experimental Example 2.

[0226] Figure 22(a) shows the sensitivity according to the blood contact area of ​​the sensor. In particular, the capacitance increased linearly as the blood contact area increased. It can be observed that the capacitance increases as the blood thickness increases to 0, 50, 100, 150, and 200 μm. Based on the coplanar capacitor principle, the sensor can detect blood thickness up to 150.25 μm because the electric field propagates in a parabolic shape from the anode to the cathode through the target medium. This limitation can be overcome by increasing the distance between the anode and cathode; however, such adjustment is unnecessary when blood leakage exceeds 150 μm, as this already implies total leakage. Additionally, increasing the distance between the electrodes may reduce sensitivity due to the influence of arterial blood flowing along the inside of the stent-graft. The sensor is not affected by blood flow inside the stent.

[0227] Figure 22 (b) shows that capacitance increases as the lateral range of blood penetrating between the blood vessel and the electrode due to blood leakage (i.e., blood leakage width) increases. Additionally, the sensor can measure the initial leakage rate by monitoring the change in capacitance over time.

[0228] Figure 22 (c) shows the change in capacitance from the start of leakage to the equilibrium stage when blood leakage occurs at three blood flow velocities (288, 688, 3440 pL / s) while maintaining a fixed blood volume of 1000 mm³ on top of the sensor. The resulting capacitance-time graph shows that the time to increase capacitance is shortened as the blood flow velocity increases, which can provide useful information for evaluating leakage urgency based on blood flow velocity.

[0229]

[0230] Figure 23 shows the change in capacitance according to the blood leakage location of the blood leakage detection sensor according to Experimental Example 2 (when blood flow occurs at site 1: capacitance increases in both electric field (E-field) 1 and 2; when blood flow occurs at site 2: capacitance increases only in electric field (E-field) 2), the result of a stress test in which stent compression and expansion were repeated 1,000 times (b), and the daily change in capacitance after immersing the sensor in phosphate-buffered saline (PBS) for one month (c), respectively.

[0231] Figure 23 (a) shows that the location of a leak can be detected by analyzing the capacitance value generated from the electric field occurring in the electrode line. Depending on the location of the blood leak, a change in capacitance occurs in a single zone or both zones of the sensor, thereby allowing the location of the leak to be identified.

[0232] Figure 23 (b) shows that the capacitance of the electrode attached to the stent remains stable even after more than 1,000 cycles of stent crimping and expansion. This demonstrates the sensor's ability to continuously and stably detect leakage in the dynamic environment of the aorta, which maintains a hypertensive environment even after implantation.

[0233]

[0234] 2.3.3. Ex vivo sensor evaluation in a porcine aortic model

[0235] Figure 23 (c) shows the long-term stability of the sensor in an in vitro environment. The sensor was immersed in a phosphate-buffered saline (PBS) solution at 37°C, and the capacitance response was measured for one month; no significant changes were observed. The baseline capacitance value in air and the value upon blood detection were measured at 3-day intervals, and the results were consistent, indicating that the sensor can be used stably for a long period in vivo.

[0236]

[0237] FIG. 24 shows an image of the in vitro experimental setup using a porcine aorta of the blood leak detection sensor according to Experimental Example 2 (a), the change in capacitance when blood, albumin, and PBS flow between the stent and the porcine aorta under experimental conditions (b), and the change in capacitance within the normal physiological blood pressure range (c), respectively.

[0238] Figure 24(a) shows an in vitro experiment using a porcine aorta model, indicating that the sensor can be used for endoleak detection. In the experiment, a blood leakage environment between the porcine aorta and the top of the stent was reproduced on a benchtop. A stent equipped with a sensor was inserted into the porcine aorta using a catheter, and blood was injected between the aorta and the stent to simulate a Type 1 endoleak. To reproduce normal arterial blood pressure, a pressure of 80 to 120 mmHg at 80 to 90 bpm was repeatedly applied to the stent.

[0239] Figure 24(b) shows the change in capacitance during the injection of various substances under in vitro conditions. When blood, albumin, and PBS were sequentially injected between the blood vessel and the stent using a syringe, the capacitance values ​​for each substance remained consistent, and stable results were observed even with cross-injection. In terms of electrical response, the sensor exhibited the highest capacitance with blood, followed by albumin and PBS. Fluctuations in arterial pressure at a normal heart rate can cause subtle changes between the vessel wall and leaking blood. These changes are detected through changes in capacitance in a coplanar capacitor sensor in contact with the surrounding medium. The sensor can capture minute changes in capacitance within the physiological pressure range, thereby enabling simultaneous heart rate monitoring and the detection of potential blood leaks.

[0240] Figure 24 (c) is a graph comparing the capacitance change measured by the sensor (black line) and the intravascular pressure value reproduced by the pump (red line). A capacitance change of up to 6 pF was observed in the range of 80 to 120 mmHg, which shows that heart rate monitoring is possible simultaneously with minimal impact on sensing due to the movement of blood vessels caused by blood pressure.

[0241]

[0242] 2.3.4. Thermal adhesive for sensor attachment

[0243] EVAR stent-grafts are made of a waterproof, fabric-like material such as Gore-Tex, which has flexibility and excellent mechanical stability, and expands and adheres to the blood vessel. The adhesive used to attach the electrode to the graft surface within the blood vessel must have characteristics such as (1) mechanical flexibility similar to the stent fabric, (2) strong adhesion capable of withstanding long-term implantation, (3) hardening properties that do not penetrate the inner wall and do not obstruct blood flow, and (4) high biocompatibility that can minimize inflammatory responses.

[0244] These characteristics help mitigate functional damage to the electrode, reduce the risk of detachment, and minimize immune responses. Conventional adhesives are not suitable for attaching ultrathin devices to stent-graft materials. For example, biomedical adhesives, such as tissue adhesives, are flexible but biodegradable in the body, making them unsuitable for long-term implantation. Conversely, bone adhesives offer excellent long-term stability but possess high rigidity, making them unsuitable for vascular applications.

[0245] The thermal adhesive proposed in this study is designed to overcome the limitations of existing adhesives and function stably within the intravascular environment. This study developed an adhesive suitable for stents by utilizing a material that hardens upon heat while maintaining flexibility. This adhesive possesses mechanical flexibility, strong adhesion for long-term implantation, controlled curing to prevent penetration into the inner wall, and high biocompatibility.

[0246] FIG. 25 shows an image (a) of a blood leak detection sensor and a heat-sealed film according to Experimental Example 2, a bending strength (b) of the adhesive, and an adhesion strength (c) of the adhesive, respectively.

[0247] The newly developed adhesive possesses flexibility, strong adhesion, and biocompatibility, making it highly effective for attaching microscale thin-film sensors to graft materials (see Fig. 25(a)). Figs. 25(b) and (c) show the results of comparing the bending strength and adhesion strength of the newly developed thermal adhesive with various industry-standard adhesives (PDMS, medical liquid instant adhesive Ferndale Mastisol Adhesive 52348, tissue adhesive 3M Vetbond, and topical skin adhesive Skinister Topical Skin Adhesive). An ideal adhesive should possess both low bending strength and high adhesion strength simultaneously. Here, the newly developed thermal adhesive exhibits a low bending strength of 1.5 N, making it flexible, while maintaining a high adhesion strength of 10 N. In contrast, PDMS had a higher bending strength (4.5 N) but lower adhesion strength (6.1 N). Super glue exhibited the highest flexural strength (7.6 N) and adhesive strength (12.8 N). Tissue adhesive and skin adhesive showed higher flexural strength and lower adhesive strength than the newly developed adhesive (tissue adhesive: flexural strength = 6.0 N, adhesive strength = 2.2 N; skin adhesive: flexural strength = 3.2 N, adhesive strength = 2.4 N). These results indicate that the newly developed thermal adhesive possesses the optimal combination of characteristics suitable for intravascular implantation in hypertension.

[0248]

[0249] FIG. 26 shows an SEM image (a) of the inner surface of a fully cured stent fabric after coating or attaching the adhesive according to Experimental Example 2, a state in which the adhesive is applied to the outer surface of the stent (b), a blood droplet contact angle (c) on the inner fabric of the stent after the adhesive is applied, and the results of Calcein staining and cell viability analysis (d to e), respectively.

[0250] When fully curing the adhesive, it is important to ensure that it does not penetrate the surface of the stent-graft. This is because inward penetration can obstruct blood flow. To evaluate the degree of adhesive penetration, the adhesive was applied to the outside of the stent, and the interior was observed using SEM (see Figures 26 (a) and (b)). As a result, no traces of the developed adhesive penetrating the stent fabric were found, nor were any noticeable changes observed on the inner surface compared to an untreated stent. This is because the newly developed adhesive exists in the form of a film, and its properties can be controlled by adjusting its thickness. The thickness of the adhesive film used was 25 µm. This was optimized to ensure the material does not penetrate inward while maintaining the necessary adhesive properties. Additionally, a rapid curing time of 5 seconds at 110°C minimizes the risk of penetration into the fabric during the curing process.

[0251] On the other hand, SEM analysis observed that PDMS and cyanoacrylate (super glue) penetrated finely into some graft materials, which is interpreted as being due to these adhesives being in a low-viscosity liquid state with a long curing time. In addition, the waterproofing function of the stent is also very important, and the existing waterproofing function was maintained on the inner surface of the stent (see Fig. 26 (c)).

[0252] These results demonstrate that the newly developed adhesive meets the mechanical requirements for stably attaching electrodes to the outer surface of stent-grafts, while simultaneously possessing excellent flexibility, strong adhesion, and anti-penetration properties.

[0253] The biocompatibility evaluation of the polyimide and adhesive used in this study revealed that the materials did not exhibit significant cytotoxicity to endothelial cells. When HUVECs were cultured in conditioned media containing substances extractable from each material and phase-contrast and live-cell images were observed, cell density and sprouting were similar across all groups, including the control group (see Fig. 26(d)). Quantitative MTT analysis also showed no adverse effects on HUVECs from the conditioned media containing substances extractable from the polyimide film. Furthermore, cell viability was measured at 92% in the polyimide film conditioned media and 75% in the adhesive conditioned media (see Fig. 26(e)). The relatively low cell survival rate in the adhesive-treated group is presumed to be due to small molecule-induced cytotoxicity caused by the incomplete fine control of the adhesive in the manufacturing environment, but the degree of reduction was minor.

[0254]

[0255] 2.3.5. Evaluation of Blood Leakage According to Electrode Thickness

[0256] FIG. 27 shows the results of a numerical simulation (a) for an aortic aneurysm (left) when the electrode-attached stent according to Experimental Example 2 is not inserted and for an endo-leak when the electrode-attached stent is inserted, and the results of a numerical simulation and an analytical approach (b) for blood leakage rates according to the gap size induced by changes in electrode thickness.

[0257] Figures 27 (a) and (b) show the results of evaluating blood leakage according to electrode thickness using numerical analysis and analytical modeling based on the Carreau viscosity model. Simulations were performed at six different thicknesses (5 µm, 10 µm, 50 µm, 100 µm, 500 µm, and 1000 µm). As boundary conditions, a pulse frequency of 120 bpm and an average aortic pressure of 100 mmHg were applied to reflect the cardiac characteristics of a healthy person. A comparison of blood leakage rates at 5 µm and 1000 µm confirmed that electrode thickness has a significant effect on the blood leakage rate. At 5 µm, the blood leakage rate was 10 -3 Although it was less than mm / s, at 1000 μm, the velocity increased significantly to exceed 4 mm / s (see Fig. 27(a)). In particular, it was analyzed that as the thickness increased, a recirculation zone formed between the electrode inlet and the artery wall, which increased the residence time of flow particles and could raise the risk of endoleak. In addition, an analytical model was developed to verify the calculation results and optimize the device design, and the velocity predictions for each thickness from the analytical model were consistent with the numerical analysis results (see Fig. 27(b)).

[0258]

[0259] FIG. 28 shows a blood leakage image (a) according to the thickness (5㎛ and 1,000㎛) of the blood leakage detection sensor of Experimental Example 2, the gap between the stent-graft and the blood vessel according to the thickness of the blood leakage detection sensor (the gap increases as the device thickness increases, b), and optical microscope images (c) of the top of the polyimide etched using O2 and the bottom of the polyimide etched using a mixture of O2 and CF4 gas during the etching process.

[0260] Figure 28(a) shows an experiment conducted to verify the simulation results of Figure 27 in a real-world environment. To observe blood leakage caused by sensor thickness, stents equipped with sensors having thicknesses of 5 µm and 1000 µm, respectively, were fabricated and placed in a cylindrical phantom simulating the elastic modulus of vascular tissue (see Figure 28(b)). As a result, blood leakage occurred in the 1000 µm thick sensor due to its thickness, which is consistent with the simulation results. On the other hand, no blood leakage was induced in the 5 µm thick sensor. Here, the PI (polyimide) encapsulation layer insulating the electrode of the capacitive sensor accounts for approximately 93.6% of the total sensor thickness. Through this experiment, it was confirmed that the 5 µm thick sensor does not cause blood leakage. However, in the case of the vertically fabricated encapsulation layer, its edge was not suitable for completely sealing the electrode. Accordingly, a uniform 45° slope was applied to the edge of the encapsulation layer (see (c) in Fig. 28).

[0261]

[0262] 2.3.6. LC-based wireless monitoring for endoleak detection

[0263] Electrode sensors attached to stents can be applied to various wireless systems to wirelessly monitor endoleaks from the outside. However, due to the specific environment of insertion inside blood vessels, the thickness of the sensor must be taken into account. In the case of an LC-based wireless system that detects resonant frequencies based on an LC resonant circuit consisting of separate inductors and coils, no additional data transmission or charging device is required, so it can be utilized in implantable sensors.

[0264] FIG. 29 shows an example of applying an LC-based blood leakage wireless monitoring system according to Experimental Example 2 to the human body (a), and images of a planar inductor, a reader coil, an electrode sensor with the planar inductor integrated therein, and a stent (b), respectively.

[0265] As shown in Fig. 29 (a), wireless data transmission to an external reader is difficult because the aorta is typically located at a depth of at least 60 mm from the surface of the abdominal skin. For effective communication at this depth, large inductive coils must be used in both the internal transmitter and the external receiver. Fortunately, since the stent-graft has a large surface area, large inductive coils can be attached to the outer wall of the graft. Therefore, capacitance data from the electrode sensor can be wirelessly transmitted to an external receiver using an LC resonant circuit.

[0266] Figure 29(b) shows an example of an LC-based wireless system applied to a newly developed electrode sensor. In this system, the sensor acts as a variable capacitor. The passive sensor consists of an electrode sensor and a planar inductor. The electrode sensor was fabricated by depositing copper (Cu, thickness 200 nm) using an electron-beam evaporator, and the planar inductor (30 × 100 mm) was fabricated from a copper film patterned by a laser ablation system. The wireless reader consists of a network analyzer that measures S-parameters and a reader coil. The reader coil (60 × 200 mm) is made with the same shape as the planar inductor but twice the size, and uses copper wire to facilitate long-range electromagnetic coupling.

[0267] When the planar inductor attached to the stent is aligned with the leader coil, the resonant frequency peak appears most distinctly, allowing for easy detection of changes in the resonant frequency caused by blood leakage. To verify whether the detection performance of the electrode sensor is maintained even with this wireless system and planar inductor, blood was applied over various areas on the electrode sensor, and the resulting resonant frequencies were measured.

[0268]

[0269] FIG. 30 shows the simulation configuration performed to optimize the dimensions of the planar inductor and the leader coil according to Experimental Example 2 (the leader coil (Tx) is designed to be twice the length of the planar inductor (Rx) so that measurements can be taken from a longer distance) (a), the change in resonance frequency according to the blood coverage of the blood leak detection sensor (b), and the change in resonance frequency measured when the distance between the planar inductor and the leader coil is 70 mm (c).

[0270] Figure 30(a) shows a numerical simulation performed to optimize the shape and size of the inductor. The external inductor must be large enough to enable long-range magnetic coupling. This is because as the surface area of ​​the inductor increases, the magnetic field spreads over a wider area, improving the coupling coefficient with the reader coil. To prevent magnetic-field cancellation, the inductor was placed in a semicircular area surrounding the stent. As a result, the inductor was designed in a vertically elongated shape proportional to the length of the stent.

[0271] In addition, since the shape and size of the leader coil, as well as the size of the inductor, affect magnetic coupling, the size of the leader coil was considered as an important analytical variable in the simulation. As a result of the simulation, when the distance between the inductor and the leader coil was 100 mm and the leader coil was the same size as the inductor, the S11 peak was measured at -0.2 dB. On the other hand, when the leader coil was fabricated to be twice the size of the inductor, the S11 peak decreased to -1.8 dB. This is because the larger leader coil provided much stronger magnetic coupling. Based on these simulation results, a vertically elongated inductor and a leader coil twice the size of the inductor were finally selected for the wireless measurement system.

[0272] Figure 30(b) shows the change in resonant frequency according to the area of ​​the electrode covered with blood (0% to 100%). Initially, when only air was present on the electrode sensor (0%), the resonant frequency was 17.1 MHz. However, when 12.5% ​​of the electrode was covered with blood, the frequency decreased sharply to 13.2 MHz. As the area of ​​the electrode sensor covered with blood increased, the resonant frequency also decreased proportionally. It was found that even a slight increase in the area covered with blood caused the resonant frequency to drop significantly in the MHz range. This demonstrates that highly sensitive detection is possible even wirelessly. As a result of plotting a trend line from the measured data, it was found that the resonant frequency changes non-linearly depending on the area covered with blood, which is consistent with the theoretical relationship that frequency is inversely proportional to the square root of capacitance. In other words, it was confirmed that the experimental results match the theoretical prediction.

[0273] Because the stent is positioned deep within the aorta, a significant amount of biological tissue exists between the planar inductor and the leader coil. To replicate these conditions, the gap between the sensor inductor and the leader coil was set to 70 mm, and S-parameters were measured in the range of 1 MHz to 20 MHz using a network analyzer. Additionally, a phantom with the relative permittivity of muscle tissue was fabricated and placed between the inductor and the coil to simulate the conditions under which electromagnetic waves pass through or reflect through biological tissue. Figure 30(c) shows the S11 data measured when blood vessels and blood were sequentially placed on the electrode sensor under these conditions. A significant change in resonance frequency was observed, with the resonance frequency being approximately 9 MHz when the blood vessels were placed and approximately 6 MHz when the blood was placed. Notably, a distinct resonance peak was maintained despite the significant distance between the planar inductor and the leader coil, indicating that this sensor can reliably detect blood leakage even when implanted in the body. Using such a sensor-integrated stent allows patients to rapidly monitor changes in blood flow without visiting a hospital. In other words, patients with abdominal aortic aneurysms can significantly reduce the time and effort required to detect Type 1 endoleaks.

[0274]

[0275] 2.4. Conclusion

[0276] This study presents a capacitive electrode sensor for real-time endoleak monitoring to prevent aneurysm recurrence. The thin and flexible capacitive sensor detected the capacitance difference between blood and blood vessels and demonstrated a linear increase in capacitance even with residual blood volumes reaching a thickness of 100 μm. Furthermore, the sensor demonstrated high sensitivity and accuracy in measuring heart rate, blood flow location, and velocity within a blood pressure range of 80 to 120 mmHg. Experimental results showed that the proposed sensor, with a thickness of 5 μm, did not induce blood leakage. A strong bonding force was secured between the stent fabric and the electrode by using a film-type thermal adhesive. The proposed sensor was confirmed to maintain stable capacitance values ​​even after more than 1,000 compression and expansion cycles, exhibiting stable performance even in the high-pressure environment of the aorta. Therefore, this sensor demonstrated long-term stability for both the blood vessels and the sensor in vascular, soft tissue, and dynamic environments.

[0277]

[0278] Although the present invention has been described in detail through specific embodiments and experimental examples, this is for the purpose of specifically explaining the invention, and the invention is not limited thereto. It is evident that modifications or improvements can be made by those skilled in the art within the technical scope of the invention.

[0279] All simple variations or modifications of the present invention fall within the scope of the present invention, and the specific scope of protection of the present invention will be clarified by the appended claims.

[0280]

[0281] The present invention is recognized as having industrial applicability as a stent-graft attached blood leak detection sensor that is installed on the outer surface of a stent-graft, does not cause blood flow leakage and does not damage the stent-graft, and induces a change in capacitance corresponding to blood leakage to detect blood leakage within a blood vessel after stent-graft insertion, a method for manufacturing the same, and a wireless blood leak detection system using the blood leak detection sensor.

Claims

1. A base film having a continuous zigzag curve along the longitudinal direction and surrounding the outer surface of a tubular stent-graft; An adhesive film that bonds one side of the base film and the outer surface of the stent-graft; A capsulation film formed in a shape corresponding to the base film, wherein one side is laminated to the other side of the base film; and The electrode portion comprises a first conductive line formed in a shape corresponding to the zigzag shape of the base film, and a second conductive line formed in a shape corresponding to the first conductive line and arranged parallel to the first conductive line at a predetermined distance, disposed between the base film and the encapsulation film, and having a change in capacitance when blood comes into contact with the other surface of the encapsulation film. A stent-graft attached blood leak detection sensor that detects the blood leaking from the blood vessel where the stent-graft is installed, according to the change in the above-mentioned capacitance.

2. In Claim 1, Each of the above base film and the above encapsulation film is, A stent-graft attached blood leak detection sensor formed of an insulating material.

3. In Claim 1, The above base film is, A stent-graft attached blood leakage detection sensor that is adhered to a region of the outer surface of the stent-graft that contacts the blood vessel.

4. In Claim 1, The above adhesive film is, A porous heat-sealable sheet formed by arranging a plurality of fibers made of a first heat-sealable polymer; and A stent-graft attached blood leakage detection sensor comprising: a heat-sealable layer formed by coating a powder made of a second heat-sealable polymer onto the surface of the heat-sealable sheet.

5. In Claim 1, The above-mentioned first conductive line and the above-mentioned second conductive line are, Each has a width of 180 to 220 µm and a thickness of 0.5 µm or less, and A stent-graft attached blood leak detection sensor with a spacing of 20 to 40 µm.

6. In Claim 1, The above electrode part is, A stent-graft attached blood leakage detection sensor further comprising: a third conductive line which is relatively shorter than the length of the second conductive line, is formed in a shape corresponding to a predetermined area in the longitudinal direction of the second conductive line, and is arranged parallel to the predetermined area at a predetermined distance.

7. (a) A step of manufacturing an adhesive film by coating a powder made of a second heat-sealable polymer onto a porous heat-sealable sheet formed by arranging a plurality of fibers made of a first heat-sealable polymer; (b) a step of manufacturing an electrode structure on a substrate, comprising a base film having a continuous zigzag curved portion along the longitudinal direction, two or more conductive lines formed in a shape corresponding to the zigzag shape of the base film and arranged parallel to each other, and a encapsulation film formed in a shape corresponding to the base film and laminated on the base film with the conductive lines in between; and (c) a step of separating the electrode structure from the substrate and transferring it to one side of the adhesive film; a method for manufacturing a stent-graft attached blood leakage detection sensor.

8. In Claim 7, The above step (a) is, A step of melting the first heat-sealable polymer; A step of manufacturing the heat-sealable sheet by spinning the molten first heat-sealable polymer into the form of the fiber; and A method for manufacturing a stent-graft attached blood leakage detection sensor comprising the step of coating a coating solution in which the powder is dissolved on the surface of the manufactured heat-bonded sheet.

9. In Claim 7, The above step (b) is, A step of forming a sacrificial layer on the above substrate; A step of forming the base film with a first insulating polymer on the sacrificial layer; A step of forming the conductive line on the base film; and A method for manufacturing a stent-graft attached blood leakage detection sensor comprising the step of forming the encapsulation film with a second insulating polymer on the base film having the aforementioned conductive line formed thereon.

10. In Claim 9, The above step (c) is, Step of removing the above-mentioned sacrificial layer; A step of attaching a first adhesive sheet to one surface of the electrode structure and separating it from the substrate; A step of placing the other side of the electrode structure attached to the first adhesive sheet onto one side of the adhesive film; A step of forming a multilayer structure by placing a second adhesive sheet on the other side of the adhesive film on which the electrode structure is placed; A step of heating the above multilayer structure to transfer the electrode structure to the adhesive film; and A method for manufacturing a stent-graft attached blood leakage detection sensor comprising the step of immersing the heated multilayer structure in an organic solvent to remove the first adhesive sheet and the second adhesive sheet from the adhesive film.

11. In Claim 7, (d) a step of arranging the electrode structure to surround the outer surface of a tubular stent-graft, and heating the adhesive film on which the electrode structure is transferred to adhere the electrode structure to the outer surface of the stent-graft; further comprising a method for manufacturing a stent-graft-attached blood leakage detection sensor.

12. In Claim 11, The above step (d) is, A step of attaching the adhesive film to the outer surface of the stent-graft by spraying ethanol; and A method for manufacturing a stent-graft attached blood leakage detection sensor comprising the step of heating the attached adhesive film to adhere the electrode structure to the outer surface of the stent-graft.

13. A blood leakage detection sensor according to claim 1, which is attached to the outer surface of a tubular stent-graft and acts as a variable capacitor whose capacitance changes in response to blood leaking from a blood vessel; A planar inductor attached to the outer surface of the stent-graft and electrically connected to the blood leakage detection sensor, forming an LC resonant circuit in which the resonant frequency changes in response to a change in capacitance; and A wireless blood leak detection wireless system comprising: a wireless reader unit that wirelessly detects changes in resonance frequency, including a reader coil disposed outside the body and magnetically coupled with the planar inductor.

14. In Claim 13, The above planar inductor is, A blood leakage detection system formed in an elongated rectangular shape along the longitudinal direction of the above-mentioned stent-graft.

15. In Claim 14, The above planar inductor is, The width is 20 to 40 mm, and Blood leak detection system with a length of 80 to 120 mm.

16. In Claim 14, The above planar inductor is, A blood leakage detection system positioned in a semicircular area on the outer surface of the above-mentioned stent-graft.

17. In Claim 14, The above leader coil is, It is formed in a shape corresponding to the shape of the above-mentioned planar inductor, A blood leak detection system having a size 1.5 to 2.5 times that of the above-mentioned planar inductor.

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