Composite adhesive structure comprising porous structural layer and hydrogel

Abstract

The present application relates to a composite adhesive structure comprising a porous structural layer and a hydrogel, and a medical device comprising same and, particularly, to an implantable device comprising the composite adhesive structure comprising a porous structural layer and a hydrogel.

Description

Composite adhesive structure comprising a porous structural layer and a hydrogel

[0001] The present invention relates to a composite adhesive structure including a porous structural layer and a hydrogel, and a medical device including the same.

[0002] Implantable devices are medical devices inserted into the body for the purpose of diagnosing or treating diseases. However, the dynamics of body movement during physiological activity can cause the implanted device to move from its initial position. The present invention focuses on the movement of stents, which serve to maintain the patency of blood vessels and other body passages, including the esophagus, bronchial tubes, biliary tract, and ureters. After insertion, non-vascular contraction and relaxation, such as peristalsis, generate pressure differences within the vessel. This highlights the mechanical and biological mismatch between soft, moist human tissue and externally inserted materials, such as metals, oxides, and polymers used in stents, leading to migration problems. This migration can reduce therapeutic efficacy, potentially damage surrounding tissue, and increase the risk of adverse events. To prevent stent migration, various methods, including clips, stent fixation, barbs, flips, and shape modifications, are used to securely secure the stent. Although effective, these methods must be tailored to the individual needs of each patient, may cause damage or scarring of surrounding tissues, and are associated with pain and a foreign body sensation after insertion.

[0003] Likewise, traditional invasive methods such as sutures and staples continue to be widely used in the areas of hemorrhage control and wound closure. However, recent research has advanced sutureless technologies that leverage the properties of robust hydrogels and dry polymer films. Biocompatible hydrogel adhesives possess the remarkable ability to covalently bond to specific tissues and achieve significant levels of adhesion. Due to their high moisture content, softness, and stretchability, hydrogels are well-suited to mimic the mechanical properties of soft biological tissues, making them particularly advantageous in situations where traditional sutures and staples may be less effective.

[0004] Recent research has focused on strategies for effectively bonding hydrogels to substrates. These include surface bridging / initiation and hydrogel painting techniques. In particular, surface bridging utilizes bridging molecules such as silanes or silane coupling agents (e.g., TMSPMA and APTES) to establish durable bonds between hydrogels and substrates. However, this approach can be cumbersome, requiring optimization of both the substrate and the hydrogel. Despite the development of hydrogels with impressive physical properties, the weak bond between hydrogels and solid materials often poses a significant challenge, hindering their integration and functionality in devices and systems.

[0005] [Prior Art Literature]

[0006] [Patent Document]

[0007] Republic of Korea Patent No. 10-1845115.

[0008] The present invention provides a composite adhesive structure including a porous structural layer and a hydrogel, and a medical device including the same.

[0009] However, the problems that the present invention seeks to solve are not limited to the problems mentioned above, and other problems not mentioned will be clearly understood by those skilled in the art from the description below.

[0010] The first aspect of the present invention provides a composite adhesive structure comprising: a substrate; a porous structural layer formed on the substrate; and a hydrogel layer formed on the porous structural layer and including a hydrogel.

[0011] A second aspect of the present invention provides a medical device comprising a composite adhesive structure according to the first aspect.

[0012] A third aspect of the present invention provides a method for manufacturing a composite adhesive structure, comprising the steps of: (a) forming a porous structure layer on a substrate; and (b) forming a hydrogel layer on the porous structure layer.

[0013] The composite adhesive structure according to the embodiments of the present invention has an effect of strengthening the adhesiveness to living tissue and preventing movement when applied to a living body due to the strong adhesiveness between the porous structural layer and the hydrogel.

[0014] Figure 1 is a schematic diagram showing, in one embodiment of the present invention, (a) the cause of movement of an esophageal stent; (b) a conventional fixation method for preventing movement of an esophageal stent; (c) the appearance of an HFLAT-stent in which a hydrogel is coated on a conventional FLAT-stent and the appearance of peeling off after swelling due to weak adhesive force and the weak adhesive force; (d) the introduction of RINC on the surface of the stent to increase the adhesive force between the stent and the hydrogel and the adhesive principle thereof and the appearance of not peeling off even after swelling due to the strong adhesive force.

[0015] Figure 2 shows, in one embodiment of the present invention, (a) surface schematic diagrams of FLAT, HFLAT, RINC, and HiRINC; (b and c) surface and cross-sectional SEM images of RINC (b) and HiRINC (c); and (d) comparison of Fourier transform infrared spectra of hydroxyl groups of HFLAT and HiRINC.

[0016] Figure 3 shows, in one embodiment of the present invention, (a) force versus displacement curves of 90 degree peel tests of HFLAT and HiRINC; (b) comparison of interfacial strength, adhesion strength, and work of adhesion of HFLAT and HiRINC; (c and d) plots of the ratio of the final distance traveled and the maximum cycle during cyclic lap shear tests of HFLAT (c) and HiRINC (d).

[0017] FIG. 4 shows photographs of each step after a cross-cut test in the as-prepared (a) and swelling (b) state of HFLAT and HiRINC in one embodiment of the present invention.

[0018] FIG. 5 is a photograph showing, in one embodiment of the present invention, (a) the adhesive strength between the tissue and the hydrogel of HFLAT and HiRINC; (b and c) lifting of each tissue by HFLAT (b) and HiRINC (c).

[0019] Figure 6 shows, in one embodiment of the present invention, (a and b) EDS mapping images (left), low magnification (middle) and high magnification (right) SEM images of RINC-stent (a) and HiRINC-stent (b); (c and d) optical microscope images of HFLAT-NiTi wire (c) and HiRINC-NiTi wire (d) in dry (left) and swollen (right) states; (e and f) optical microscope images of HFLAT-stent (e) and HiRINC-stent (f) in swollen states.

[0020] Figure 7 shows experimental images (a) and maximum force values ​​(b) for in vivo lap shear tests in one embodiment of the present invention (a and b).

[0021] Figure 8 shows, in one embodiment of the present invention, (a) an overview of the randomization operation and follow-up of rats; (b) the extent of migration of stents installed in the esophagus of rats (control, top), HFLAT-stent (middle), and HiRINC-stent (bottom) tracked and compared through X-ray imaging for 4 weeks.

[0022] Figure 9 is a lap shear test curve for comparing the adhesion work and bond strength between (a and b) HFLAT (a) and HiRINC (b) in one embodiment of the present invention; and (c and d) SEM images of the separation at the interface between the hydrogel and RINC during the lap shear test of HiRINC.

[0023] FIG. 10 shows the results of 90 degree peel tests for various commercial tapes of FLAT, FLAT_Amine, silica thin film, porous silica nanoparticles, and RINC in one embodiment of the present invention.

[0024] Figure 11 shows the swelling dynamics results of HFLAT and HiRINC in one embodiment of the present invention, (a to d) showing graphs showing changes in weight (a), area (b), and thickness (c) over time, and (d) showing a graph comprehensively showing the swelling ratio according to these.

[0025] Figure 12 shows SEM images of (a to c) FLAT-stent (a), HFLAT-stent (b), and HiRINC-stent (c) after in vivo lap shear testing in one embodiment of the present invention.

[0026] Figure 13 shows SEM images according to the pore diameter and thickness of RINC (a, b, d, and e), and adhesion of HiRINC according to the pore diameter and thickness of RINC in one embodiment of the present invention (c and f).

[0027] Figure 14 shows, in one embodiment of the present invention, (a) the appearance of FLAT-PTFE (bare PTFE); (b) the appearance of a RINC-stent in which a porous structural layer (3D nanostructure) is synthesized and fused with RINC-PTFE; (c and d) comparison of SEM images of the surfaces of FLAT-PTFE (c) and RINC-PTFE (d).

[0028] FIG. 15 shows the results of a mechanical adhesive strength analysis according to a lap shear test of a polydopamine (PDA)-based hydrogel ((a) polydopamine content 0.12 w / w%, (b) polydopamine content 0.25 w / w%) in one embodiment of the present invention.

[0029] Figure 16 shows the peel resistance analysis results according to a peel test of a PDA-based hydrogel in one embodiment of the present invention.

[0030] Figure 17 shows the results of (a) lap shear test and (b) peel test analysis of a chitosan-based hydrogel in one embodiment of the present invention.

[0031] Figure 18 shows the results of comparing the adhesive strength according to the swelling time of a PVA-AAc-based hydrogel in one embodiment of the present invention.

[0032] Figure 19 shows the results of evaluating the drug loading and release performance of a porous silica nanostructure in one embodiment of the present invention.

[0033] Figure 20 shows the results of evaluating the process applicability of hydrogel coating in one embodiment of the present invention.

[0034] Figure 21 shows, in one embodiment of the present invention, (a) the results of an adhesive strength evaluation according to swelling time of a PVA-AAc-based hydrogel and (b) an image after 24 hours of swelling.

[0035] Hereinafter, with reference to the attached drawings, implementation examples and embodiments of the present invention will be described in detail so that those skilled in the art can easily practice the present invention. However, the present invention may be implemented in various different forms and is not limited to the implementation examples and embodiments described herein. In addition, in the drawings, parts irrelevant to the description have been omitted to clearly explain the present invention, and similar parts have been designated with similar drawing reference numerals throughout the specification.

[0036] Throughout this specification, when it is said that an element is "on" another element, this includes not only cases where the element is in contact with the other element, but also cases where another element exists between the two elements.

[0037] Throughout this specification, whenever a part is said to "include" a component, this does not mean that it excludes other components, but rather that it may include other components, unless otherwise specifically stated.

[0038] The terms "about," "substantially," and the like used in this specification are used in a meaning that is at or close to the numerical value when manufacturing and material tolerances inherent in the meanings mentioned are presented, and are used to prevent unscrupulous infringers from unfairly exploiting the disclosure in which exact or absolute values ​​are mentioned to aid understanding of the present application.

[0039] The terms “step of ~” or “step of ~” as used throughout this specification do not mean “step for ~.”

[0040] Throughout this specification, the term "combination(s) thereof" included in the expressions in the Makushi format means one or more mixtures or combinations selected from the group consisting of the components described in the expressions in the Makushi format, and means including one or more selected from the group consisting of said components.

[0041] Throughout this specification, references to “A and / or B” mean “A or B, or A and B.”

[0042] Below, the implementation examples of the present invention are described in detail, but the present invention may not be limited thereto.

[0043] The first aspect of the present invention provides a composite adhesive structure comprising: a substrate; a porous structural layer formed on the substrate; and a hydrogel layer formed on the porous structural layer and including a hydrogel.

[0044] In one embodiment of the present invention, the material may be used without limitation as long as it is a material used as a medical material or a medical device. The medical material and the medical device may each be implantable into a living body.

[0045] In one embodiment of the present invention, the substrate may include, but is not limited to, one or more selected from an alloy, a biodegradable polymer, and a non-degradable polymer.

[0046] In one embodiment of the present invention, the alloy may include one or more selected from stainless steel, cobalt-chromium, nickel-titanium (Nitinol), and platinum-chromium, but may not be limited thereto.

[0047] In one embodiment of the present invention, the biodegradable polymer may include one or more selected from PLA (Polylactic acid), PGA (Polyglycolide), PCL (Polycaprolactone), PEG (Polyethyleneglycol), and PLGA (Polylactic-co-glycolicacid), but may not be limited thereto.

[0048] In one embodiment of the present invention, the non-degradable polymer may include at least one selected from PTFE (polytetrafluoroethylene), PET (polyethylene terephthalate), PU (polyurethane), PDMS (polydimethylsiloxane), and silicone membrane, but may not be limited thereto.

[0049] In one embodiment of the present invention, the material of the above-mentioned description may be selected in consideration of mechanical strength, biocompatibility, processability, surface treatment suitability, etc., and may be manufactured by using the materials listed above alone or in combination.

[0050] In one embodiment of the present invention, the substrate may be manufactured in the form of a stent, a tissue scaffold, a biofixation implant, or an adhesion-preventing film applicable to internal organs of the human body, such as the esophagus, biliary tract, large intestine, trachea, bronchi, and ureter.

[0051] In one embodiment of the present invention, the porous structure layer may include a substrate-derived porous structure layer formed by treating the surface of the substrate; or a coated porous structure layer formed on the surface of the substrate and composed of a material different from the material of the substrate.

[0052] In one embodiment of the present invention, the porous structure layer derived from the substrate may be formed by surface treating the substrate through a process such as plasma, chemical etching, or laser ablation, but may not be limited thereto.

[0053] In one embodiment of the present invention, the coated porous structure layer may include, but is not limited to, one or more selected from inorganic oxides including silica (SiO2), titania (TiO2), alumina (Al2O3), zirconia (ZrO2), or ceria (CeO2); and crystalline porous structures including metal-organic frameworks (MOFs) or zeolitic imidazolate frameworks (ZIFs).

[0054] In one embodiment of the present invention, the coated porous structure layer may be formed through a sol-gel process, a wet chemical process, a spin coating, or an electrophoretic process, but may not be limited thereto.

[0055] In one embodiment of the present invention, the porous structural layer may have a form in which the lowermost region in contact with the substrate is partially or completely connected.

[0056] In one embodiment of the present invention, the porous structure layer may be a structure in which each nanostructure composed of a nanorod, nanoparticle, nanowire, nanoplate, nanocylinder or nanocube is formed on the substrate and connected to each other.

[0057] In one embodiment of the present invention, the thickness of the porous structure layer is from about 5 nm to about 1000 nm, from about 5 nm to about 900 nm, from about 5 nm to about 800 nm, from about 5 nm to about 700 nm, from about 5 nm to about 600 nm, from about 5 nm to about 500 nm, from about 5 nm to about 400 nm, from about 5 nm to about 300 nm, from about 5 nm to about 200 nm, from about 5 nm to about 100 nm, from about 10 nm to about 1000 nm, from about 10 nm to about 900 nm, from about 10 nm to about 800 nm, from about 10 nm to about 700 nm, from about 10 nm to about 600 nm, from about 10 nm to about 500 nm, from about 10 nm to about 400 nm, from about 10 nm to about 300 nm, from about It may be, but is not limited to, 10 nm to about 200 nm, about 10 nm to about 100 nm, about 50 nm to about 900 nm, about 50 nm to about 800 nm, about 50 nm to about 700 nm, about 50 nm to about 600 nm, about 50 nm to about 500 nm, about 50 nm to about 400 nm, about 50 nm to about 300 nm, about 50 nm to about 200 nm, or about 50 nm to about 100 nm.

[0058] In one embodiment of the present invention, the porous structure layer may include pores, grooves, channels, or valleys therein. The pores may have a diameter of about 1 nm to about 500 nm, and a depth of more than 0 nm to about 1000 nm. The pores may have a hybrid form in which open pores, closed pores, channel-type, hierarchical structure, or interpenetrated network structure are mixed, and this may act as a major factor controlling the degree of mechanical interlocking by interacting with the penetration and fixation properties of the hydrogel composition. In addition, the pore structure may act as a storage space capable of carrying a drug therein.

[0059] In one embodiment of the present invention, the diameter of the pores of the porous structural layer may be, but is not limited to, about 1 nm to about 500 nm, about 1 nm to about 400 nm, about 1 nm to about 300 nm, about 1 nm to about 200 nm, about 1 nm to about 100 nm, about 1 nm to about 50 nm, about 10 nm to about 300 nm, about 10 nm to about 200 nm, about 10 nm to about 100 nm, or about 10 nm to about 50 nm.

[0060] In one embodiment of the present invention, the porous structural layer acts beyond a simple coating and serves as a functional platform for the expression of adhesive strength of the hydrogel, and can play a key role in maintaining the hydrogel's peeling resistance, durability, and positional stability in a wet environment. In particular, the pore density and the pore area ratio of the porous structural layer to the substrate surface can directly affect the hydrogel's application stability and long-term adhesion.

[0061] In one embodiment of the present invention, the porous structural layer can suppress the expansion or damage of the hydrogel layer due to external pressure by mechanically integrating the hydrogel layer and the substrate, and can provide structural stability and pressure resistance. Hydrogel is a soft material with a high water content, and has low tensile strength and fracture toughness, so that external pressure (e.g., pressure, expansion force, and peristalsis in the intestine, etc.) can cause interfacial peeling, tearing, or shrinkage damage. Therefore, there is a problem that it is difficult to sufficiently secure the mechanical durability required for a long-term retention structure with only a single-layer hydrogel coating. The porous structural layer can improve resistance to shear stress by inducing mechanical interlocking between the substrate and the hydrogel layer, and can effectively distribute the expansion force of the hydrogel or external pressure to the substrate, thereby improving the structural stability of the entire composite structure. If the above porous structural layer is absent or the interfacial adhesion is weak, stress may be concentrated within the hydrogel, causing local damage.

[0062] In one embodiment of the present invention, the coated porous structure layer, particularly, the silica-based coated porous structure layer, may have a hydroxyl group (-OH) on the surface, and various functional groups may be introduced based on the hydroxyl group. For example, a reactive functional group such as an amine group (-NH2), a carboxyl group (-COOH), an aldehyde group (-CHO), or a thiol group (-SH) may be introduced to the surface using a silane coupling agent. In addition, the reactive functional group of the coated porous structure layer may be bonded to a functional group (e.g., a catechol group, a carboxyl group, an amine group, etc.) in the hydrogel composition through a covalent bond or electrostatic interaction. Accordingly, the coated porous structure layer and the hydrogel layer may have excellent fixing strength based on chemical coupling, and may improve peeling resistance and adhesive durability of the hydrogel layer.

[0063] In one embodiment of the present invention, the silica-based coating-type porous structure layer may be formed by a manufacturing method including: (a1) adding an ionic surfactant and an auxiliary spacer to a reaction tank containing a substrate and stirring and heating; and (a2) adding a silica precursor solution to the reaction tank and stirring and heating.

[0064] In one embodiment of the present invention, the ionic surfactant may be a cationic surfactant or an anionic surfactant, and the auxiliary spacer may have a counter ionicity to the ionic surfactant.

[0065] In one embodiment of the present invention, in the step (a1), the ionic surfactant and the auxiliary spacer can self-assemble to form a micelle structure on the surface of the substrate.

[0066] In one embodiment of the present invention, in the step (a2), the silica precursor may be laminated on the surface of the substrate to ultimately form the porous structure layer.

[0067] In one embodiment of the present invention, the silica precursor solution may include at least one selected from tetramethyl orthosilicate (TMOS), tetraethyl orthosilicate (TEOS), tetrapropyl orthosilicate (TPOS), tetrabutyl orthosilicate (TBOS), tetramethoxyvinylsilane (TMVS), tetrakis(2-hydroxyethyl) orthosilicate (THEOS), and sodium metasilicate, but may not be limited thereto.

[0068] In one embodiment of the present invention, steps (a1) and (a2) may each be performed at a temperature range of about 40°C to about 95°C, but may not be limited thereto.

[0069] In one embodiment of the present invention, the ionic surfactant may be a cationic surfactant such as cetyltrimethylammonium chloride (CTAC), cetyltrimethylammonium bromide (CTAB), dodecyltrimethylammonium bromide (DTAB), or tetradecyltrimethylammonium bromide (TTAB), or an anionic surfactant such as sodium dodecyl sulfate (SDS) or sodium dodecyl benzene sulfonate (SDBS). It is more preferable to use a cationic surfactant for forming a porous structure layer having large pores. When a cationic surfactant is used, the auxiliary spacer may be an anionic counterion such as sodium salicylate, sodium benzoate, sodium cinnamate, salicylic acid, benzoic acid, cinnamic acid, sodium methoxybenzoate, methoxybenzoic acid, sodium vinylbenzoate, vinylbenzoic acid, or sodium p-toluenesulfonate.

[0070] In one embodiment of the present invention, the hydrogel may be formed by penetrating into the pores of the porous structural layer.

[0071] In one embodiment of the present invention, the hydrogel is formed by penetrating into the pores of the porous structure layer, so that mechanical interlocking can be formed between the porous structure layer and the hydrogel layer.

[0072] In one embodiment of the present invention, the porous structure layer and the hydrogel layer may form layers that are physically overlapped with each other.

[0073] In one embodiment of the present invention, the porous structural layer may include a silanol group on the surface.

[0074] In one embodiment of the present invention, the porous structure layer and the hydrogel layer may be chemically bonded.

[0075] In one embodiment of the present invention, the porous structure layer may be surface-modified using a silane compound having a hydroxy group, an amine group, an aldehyde group, a carboxyl group, a mercapto group, or an acrylic group. The silane compound may be, as a non-limiting example, 3-aminopropyl triethoxysilane (APTES), 3-mercaptopropyl trimethoxysilane (MPTMS), or 3-(trimethoxysilyl)propyl methacrylate (TMSPMA).

[0076] In one embodiment of the present invention, the hydrogel composition used to form the hydrogel layer may include a hydrophilic polymer; and one or more selected from a coupling agent, a crosslinking agent, an initiator, and a functional additive.

[0077] In one embodiment of the present invention, the hydrogel may include one or more hydrophilic polymers selected from poly(vinyl alcohol; PVA); alginate (Alg), polyethylene glycol (PEG), chitosan, gelatin, polyacrylic acid (PAAc), polyacrylamide (PAM), poly(PNIPAM), agar, poly(2-acrylamido-2-methylpropanesulfonic acid) (PAMPS), polyvinylpyrrolidone, polyethylene oxide, methyl vinyl ether maleic anhydride copolymer, isobutylene maleic anhydride copolymer, methacrylic acid butyl acrylate copolymer, methoxyethylene maleic anhydride copolymer, polyurethane, hyaluronic acid, alginic acid, dopamine, cellulose, collagen, and derivatives thereof, but may not be limited thereto. In one embodiment of the present invention, the hydrophilic polymer may be used alone or in combination to control the hydration, adhesiveness, and biocompatibility of the hydrogel.

[0078] In one embodiment of the present invention, the coupling agent includes a coupling group for forming a covalent bond between the hydrogel and a tissue or substrate, and the coupling group may be, as a non-limiting example, N-hydroxysuccinimide ester, N-hydroxysulfosuccinimide ester, aldehyde, imidoester, epoxide, isocyanate, or catechol.

[0079] In one embodiment of the present invention, the crosslinking agent induces chemical or physical three-dimensional bonding between polymer chains in the hydrogel, and may be, as non-limiting examples, N,N'-methylenebisacrylamide (NMBA or BIS), polyethylene glycol diacrylate (PEGDA), polyethylene glycol dimethacrylate, terephthalaldehyde (TPD), or dihydroxybenzaldehyde.

[0080] In one embodiment of the present invention, the initiator is one that induces polymerization or curing of the hydrogel by operating through heat, light, or an oxidation-reduction reaction, and may be, as non-limiting examples, ammonium persulfate (APS), potassium persulfate (KPS), AIBN (2,2'-Azobisisobutyronitrile), benzoyl peroxide, Irgacure series (e.g., Irgacure 2959), camphorquinone, Eosin Y, or α-ketoglutaric acid.

[0081] In one embodiment of the present invention, the functional additive may include a pH adjuster (e.g., NaOH), an ion assistant (e.g., NaCl, CaCl2), a metal ion chelator (EDTA), a viscosity adjuster, or a curing catalyst (e.g., triethanolamine (TEMED)).

[0082] In one embodiment of the present invention, the hydrogel composition is applied to the surface of a substrate in a liquid state, penetrates into the porous structural layer, and then hardens through heat or light stimulation, etc., thereby achieving mechanical fixation (interlocking) and a combination of hydrogen bonding and / or covalent bonding to ensure excellent adhesiveness and durability. The hydrogel composition can be precisely designed according to the content ratio of each component, reaction conditions, and hardening mechanism, and can simultaneously secure functional stability and structural adhesiveness even in various bio-insertion environments.

[0083] In one embodiment of the present invention, the hydrogel composition can simultaneously induce mechanical interlocking, chemical bonding, and physical bonding with the porous structural layer, thereby implementing a composite adhesive mechanism with excellent adhesive strength, durability, and biocompatibility.

[0084] The above mechanical interlocking may be achieved by the hydrogel composition penetrating into the nano- or micro-sized pores, channels, grooves, or uneven structures of the porous structural layer and then curing to fix the hydrogel composition. In this case, the hydrogel interlocks with the internal structures of the pores, thereby increasing resistance to shear and peeling stresses, and maintaining fixation even in environments where it is exposed to body fluids for long periods of time.

[0085] The above chemical bond can be induced by hydrogen bonding, coordination bonding, or covalent bonding between a hydroxyl group (-OH) or an amine group (-NH2) on the surface of the substrate and a functional group (e.g., a carboxyl group, a catechol group, an aldehyde group, or an NHS ester) within the hydrogel. For example, when the hydrogel is a dopamine-based hydrogel, the catechol group of the dopamine derivative can provide strong adhesive force based on its high affinity with metal oxide and silica-based surfaces, and can also implement an adaptive function that can control the adhesive force according to changes in external pH or ion concentration.

[0086] The above physical bonding includes the hydration property of the hydrogel, hydrogen bonding, and electrostatic interactions, and the ionic bonding and hydrophilic interaction between the polymer (e.g., chitosan, PAA, or PEG) and the substrate can contribute to improving the initial adhesive stability and peeling resistance.

[0087] In one embodiment of the present invention, the hydrogel can perform various physiological functions in addition to its adhesive function when applied in vivo. Functional drugs such as anti-inflammatory agents, antibiotics, anticancer agents, or dyes can be loaded into the hydrogel composition through physical entrapment or chemical bonding, and the drugs can also be directly loaded within the porous structural layer. This dual loading strategy allows for controlled initial rapid drug release (burst release) and sustained release, enabling long-term delivery regardless of swelling, dehydration, or degradation of the hydrogel. The release rate and amount of the drug can be controlled depending on the swelling degree of the hydrogel, pore structure, interactions between polymers, and external stimuli (pH, temperature, ionic concentration, etc.).

[0088] In one embodiment of the present invention, the coating process for forming the hydrogel layer comprises the steps of applying, infiltrating, and curing a hydrogel composition. The hydrogel composition can be infiltrated into the pores of the porous structure layer by gravity, capillary action, or vacuum / pressure assistance. The application can be performed by a drop casting, dip-coating, spin-coating, or spray-coating process. After infiltration, the hydrogel composition can be gelled and structurally fixed by heat, ultraviolet (UV) light, a chemical initiator, or an oxidation-reduction reaction. The hydrogel network formed during this process can interlock within the pores to exhibit a mechanical fixing effect. The curing conditions (time, temperature, light intensity, etc.) can affect the hydrogel network density and drug release characteristics.

[0089] The above hydrogel coating process can be applied to various substrate shapes such as tubular, flat, or mesh-shaped, and can achieve uniform and robust coating even on complex shapes or polyhedral surfaces.

[0090] In one embodiment of the present invention, the hydrogel composition can be configured through a trigger-responsive detachment mechanism designed to enable detachment at a necessary time in a biological implant environment as well as maintain fixation strength. For example, the Schiff base reaction between aldehyde groups and amine groups is stable in acidic or neutral environments, but can be reversibly dissociated and inducible in basic conditions of pH 9 or higher, leading to detachment. Such conditions can be provided using solutions such as sodium bicarbonate or glutathione, and can be used for recovering medical devices that require removal after a procedure. In addition, components that respond to specific stimuli, such as ultraviolet light, reducing environments, or peroxidation conditions, can be introduced to selectively decompose the hydrogel network itself or weaken the adhesive interface. For example, hydrogels containing specific photoinitiators will induce detachment by destroying the cross-linked structure upon UV irradiation, and if disulfide bonds are included, network collapse occurs in a reducing environment. This detachment mechanism is ideally suited for applications requiring temporary fixation, coating removal after drug delivery, or removal without damaging living tissue.

[0091] In one embodiment of the present invention, the hydrogel may include a hydroxyl group, an amine group, an aldehyde group, a carboxyl group, and / or an acrylic group on the surface.

[0092] In one embodiment of the present invention, the silanol group, hydroxyl group, amine group, aldehyde group, carboxyl group, mercapto group, and / or acrylic group present on the surface of the porous structure layer and the surface-modified porous structure layer may form a covalent bond, hydrogen bond, electrostatic attraction, and / or van der Waals force with the hydroxyl group, amine group, aldehyde group, carboxyl group, and / or acrylic group present on the surface of the hydrogel.

[0093] In one embodiment of the present invention, hydrogen bonds, electrostatic attraction, and / or van der Waals forces may be formed between the silanol groups of the porous structural layer and the hydroxyl groups of the hydrogel.

[0094] In one embodiment of the present invention, the porous structural layer and the hydrogel may be chemically and physically bonded.

[0095] In one embodiment of the present invention, the thickness of the hydrogel may be from about 10 nm to about 5 mm, but may not be limited thereto. In one embodiment of the present invention, the thickness of the hydrogel may be, but is not limited to, about 10 nm to about 5 mm, about 10 nm to about 4 mm, about 10 nm to about 3 mm, about 10 nm to about 2 mm, about 10 nm to about 1 mm, about 10 nm to about 100 μm, about 10 nm to about 10 μm, about 10 nm to about 1 μm, about 100 nm to about 5 mm, about 100 nm to about 4 mm, about 100 nm to about 3 mm, about 100 nm to about 2 mm, about 100 nm to about 1 mm, about 100 nm to about 100 μm, about 100 nm to about 10 μm, or about 100 nm to about 1 μm.

[0096] In one embodiment of the present invention, the thickness of the hydrogel may be the sum of the thickness of the porous structure layer and the hydrogel physically overlapping each other (which may substantially mean the depth of the pores of the porous structure layer) and the thickness of the hydrogel that is not overlapping the porous structure layer.

[0097] In one embodiment of the present invention, the swelling ratio of the hydrogel of the medical material with respect to area and thickness may be reduced by at least about three times compared to a hydrogel formed on a flat substrate. In one embodiment of the present invention, due to the significantly reduced swelling ratio of the hydrogel of the medical material, wrinkling, buckling, and peeling of the hydrogel, which may occur due to compressive stress caused by swelling, may be prevented.

[0098] A second aspect of the present invention provides a medical device comprising a composite adhesive structure according to the first aspect.

[0099] Detailed descriptions of parts that overlap with the first aspect of the present application have been omitted, but the contents described in the first aspect of the present application may be equally applied even if the description is omitted in the second aspect of the present application.

[0100] In one embodiment of the present invention, the medical device may be an implantable device.

[0101] In one embodiment of the present invention, the medical device may be, but is not limited to, a non-vascular stent, a vascular stent, an artificial prosthesis, a catheter, a medical dressing, a film, a fiber, a mesh, a powder, a microsphere, a sheet, a sponge, a foam, a suture anchoring device, a conduit, a plate and a screw, a drug delivery device, or an anti-adhesion membrane. The medical device may be applied without limitation as long as it has a form to which the composite adhesive structure can be applied. In particular, a medical device including the composite adhesive structure can be applied to an organ of an animal, and can be utilized for a suitable purpose depending on the type of organ to which it is applied and the application environment.

[0102] In one embodiment of the present invention, the non-vascular stent may be an esophageal, bronchial, biliary, gastrointestinal, or ureteral stent, but may not be limited thereto.

[0103] In one embodiment of the present invention, the non-vascular stent may include a nitinol wire as a substrate, but may not be limited thereto.

[0104] A third aspect of the present invention provides a method for manufacturing a composite adhesive structure, comprising the steps of: (a) forming a porous structure layer on a substrate; and (b) forming a hydrogel layer on the porous structure layer.

[0105] Detailed explanations of parts that overlap with the first aspect of the present application have been omitted, but the contents described in the first aspect of the present application may be equally applied even if the explanation is omitted in the third aspect of the present application.

[0106] In one embodiment of the present invention, (a) may include surface-treating the substrate; or forming a coating-type porous structure layer on the surface of the substrate.

[0107] In one embodiment of the present invention, surface treatment of the substrate may be performed through a process of plasma, chemical etching, or laser ablation, but may not be limited thereto.

[0108] In one embodiment of the present invention, forming the coated porous structure layer may be performed through a sol-gel process, a wet chemical process, a spin coating, or an electrophoretic process, but may not be limited thereto.

[0109] In one embodiment of the present invention, the coated porous structure layer may be a silica-based coated porous structure layer.

[0110] In one embodiment of the present invention, the method for manufacturing the silica-based coating-type porous structure layer may include the steps of (a1) adding an ionic surfactant and an auxiliary spacer to a reaction tank containing a substrate and stirring and heating; and (a2) adding a silica precursor solution to the reaction tank and stirring and heating.

[0111] In one embodiment of the present invention, the hydrogel may include one or more hydrophilic polymers selected from poly(vinyl alcohol; PVA); alginate (Alg), polyethylene glycol (PEG), chitosan, gelatin, polyacrylic acid (PAAc), polyacrylamide (PAM), poly(PNIPAM), agar, poly(2-acrylamido-2-methylpropanesulfonic acid) (PAMPS), polyvinylpyrrolidone, polyethylene oxide, methyl vinyl ether maleic anhydride copolymer, isobutylene maleic anhydride copolymer, methacrylic acid butyl acrylate copolymer, methoxyethylene maleic anhydride copolymer, polyurethane, hyaluronic acid, alginic acid, dopamine, cellulose, collagen, and derivatives thereof, but may not be limited thereto.

[0112] In one embodiment of the present invention, the step (b) may be performed by a method such as dip-coating, cast-coating, drop-coating, spray-coating, or spin-coating, but may not be limited thereto.

[0113] In one embodiment of the present invention, the step (b) may further include curing the hydrogel by external environmental changes such as ultraviolet (UV), heat, and pH after forming the hydrogel.

[0114] Hereinafter, the present invention will be described in more detail using examples. However, the following examples are provided only to help understand the present invention, and the contents of the present invention are not limited to the following examples.

[0115] [Example]

[0116] 1. Materials and Methods

[0117] 1.1 Materials

[0118] The reagents for synthesizing robust interlocking nanoconnectors (RINCs) for robust interlocking (with adhesives such as hydrogels) are triethanolamine (TEA), sodium salicylate (NaSal), cetyltrimethylammonium chloride (CTAC), and tetraethyl orthosilicate (TEOS). For the preparation of hydrogel precursors, acrylic acid (AAc), poly(vinyl alcohol) (PVA; Mw = 146,000 to 186,000), poly(ethylene glycol) dimethacrylate (PEGDMA), acrylic acid N-hydroxysuccinimide ester (AAc-NHS ester), and α-ketoglutaric acid were used.

[0119] 1.2 RINC synthesis on various substrates

[0120] Before RINC synthesis on silicon wafers, nitinol wires, nitinol stents, and glass, they were cleaned by sonication in acetone and ethanol, respectively, rinsed with deionized water, and then oven-dried. Similarly, silicon membranes and PTFE films were sonicated in isopropyl alcohol (IPA), rinsed with deionized water, and then oven-dried. All cleaned substrates were treated with O2 plasma for 40 seconds immediately prior to RINC synthesis.

[0121] 1.3 Typical RINC synthesis on silicon wafers, nitinol wires, and nitinol stents

[0122] In a typical synthesis of RINC, 272 mg of TEA was dissolved in 95 mL of deionized water at 80°C with stirring, and various substrates (silicon wafer, nitinol wire, and nitinol stent) were added. NaSal (1136 mg), CTAC (5.56 mL), and TEOS (10 mL) were added to this solution with stirring, and the resulting RINC-coated substrate was rinsed several times with ethanol to remove residual reactants.

[0123] 1.4 Control of pore diameter and thickness of RINC on silicon wafers

[0124] The pore diameter and thickness of RINCs on silicon wafers were controlled by adjusting the duration of specific steps or the amount of reagents used throughout the general synthesis process. Specifically, the pore thickness was manipulated from 5 nm to 500 nm by varying the TEOS synthesis time from 3 minutes to 3 hours.

[0125] 1.5 Preparation of hydrogel precursor

[0126] 700 mg of PVA was mixed with 6.6 mL of deionized water at 90°C, and 3.3 mL of AAc, 20 mg of α-ketoglutarate, 4.55 uL of PEGDMA, and 30 mg of AAc-NHS ester were added. Finally, the mixture was thoroughly stirred and stored away from ultraviolet (UV) light.

[0127] 1.6 Application of hydrogel precursor on substrate

[0128] Hydrogel dip coating

[0129] Silicon wafer, nitinol wire, and nitinol stent at 13 mm min -1The hydrogel was coated by dip-coating at a speed of 100 s and cured in a UV chamber (365 nm, 15 W) for 30 minutes to produce a hydrogel-coated smooth and flat surface (HFLAT) and a hydrogel-impregnated robust interlocking nanoconnector (HiRINC) coating.

[0130] Cast coating of hydrogel

[0131] For lap shear testing in dry condition, PDMS molds were made of 15 × 15 mm 2 The silicon wafer was placed in the center, and the hole of the PDMS mold was filled with 350 μL of hydrogel precursor. After casting the hydrogel precursor, another 15 × 15 mm 2 Gently place the silicon wafer on the PDMS mold to form a 10 × 10 mm 2 After forming a square-shaped overlapping area, it was cured under a UV lamp (365 nm, 15 W power) for 1 hour.

[0132] Peel test samples are 15 × 35 mm 2 After casting the hydrogel precursor into the PDMS mold, 15 × 30 mm 2 Place the silicon wafer on the edge of the mold to form a 10 × 30 mm 2 A rectangular overlapping area was formed. After placing the silicon wafers, UV (365 nm, 15 W power) was exposed to 90 pre-cured peel test samples for 1 hour.

[0133] 1.7 Measurement of interface strength and adhesion for characterization

[0134] Interfacial strength and adhesion were measured using a mechanical testing system equipped with a digital force gauge (ESM303, Mark-10) and a force test stand (F305-EM, Mark-10). Interfacial strength and travel displacement were measured at 13 mm min -1 The strain rate was recorded at 50 mm min. The bond strength and displacement were -1 was recorded at the deformation rate of .

[0135] 1.8 Swelling time adhesion test

[0136] To compare the effect of nanostructures on swelling-induced lap shear strength reduction, lap shear test samples were swollen in deionized water for predetermined swelling times. Lap shear strength was measured after each swelling period, keeping all conditions consistent except for varying time intervals.

[0137] 1.9 In vitro lap shear test

[0138] To demonstrate strong adhesion between hydrogels and porcine tissue, the hydrogel surfaces of HFLAT and HiRINC were applied to the tissue with a gentle pressure of 1 kPa for 30 s prior to lap shear testing.

[0139] 1.10 Adhesive interface fatigue test

[0140] To evaluate the interface weakening threshold, lap shear tests were performed under cyclic loading (force-controlled mode) using the same experimental setup. In lap shear tests performed over multiple cycles, a cyclic lap shear strength F was applied to the hydrogel coating for N cycles (N ≤ 100), and the interfacial crack extension, denoted as t, was recorded.

[0141] 1.11 Cross cut test

[0142] The adhesion strength of the coating was evaluated using a blade on a hydrogel-coated substrate. Two sets of cross-shaped scratches were created on the coated substrate using the blade, three in one direction and three in the vertical direction, forming a cross-shaped pattern. The hydrogel-coated substrate was immersed in deionized water for 1 hour to induce swelling conditions. Finally, a piece of tape (Scotch tape, 3M) was applied to the scratched sample and steadily removed within 5 seconds, after which a force of 1 kPa was applied and the sample was pulled at a 180° angle.

[0143] 1.12 Anti-swelling test

[0144] After preparing the cured hydrogel (Φ = 10 mm, 0.5T) on a silicon wafer, the swelling performance of the hydrogel was tested by immersing the sample in deionized water at room temperature. The swelling ratio (SR) was calculated by the following formula: Swelling ratio (%) = (W t -W0) / W0× 100(%)

[0145] 1.13 In vivo stent adhesion test

[0146] For in vivo lap shear testing in a rat model, three groups of stents were tied with a string before introduction into the esophagus. After tying the string to a mechanical testing device, lap shear testing was performed using the method described above. The utilized stents were recovered, and images of the residual hydrogel coating in each stent group were captured using SEM.

[0147] 2. Results

[0148] 2.1 Adhesive properties

[0149] The present invention presents the results of an investigation of the adhesive properties of a nanofilm having a defect-free porous structural layer, called a "robust interlocking nanoconnector (RINC)." By utilizing the unique nanostructure of the "RINC," we enhanced adhesion to a hydrogel, resulting in a composite, the "hydrogel-impregnated robust interlocking nanoconnector (HiRINC)," which represents an innovative approach to improving hydrogel-substrate adhesion. A high-magnification scanning electron microscope (SEM) image (Fig. 2b) reveals the porous structure of the "RINC," detailing the consistent thickness of the nanofilm of approximately 150 nm, confirming the presence of pores important for hydrogel integration. Upon application of the hydrogel, the "RINC" is converted to "HiRINC." This process is performed via methods such as spin-coating, mold-casting, or dip-coating, and produces a transparent and smooth hydrogel layer, as shown in Figure 2c. The cross-sectional view highlights the extensive penetration of the hydrogel into the pores of the RINC, demonstrating direct contact and mechanical interlocking between the hydrogel and the porous silica nanofilm.

[0150] 3000 cm -1 3700 cm inland -1 The bonding within the hydrogel can be quantitatively evaluated through FTIR spectroscopy analysis (Fig. 2d) in the hydroxyl (OH) bonding region of the range. The present inventors have determined that the bonding within the hydrogel is 3200 cm -1 Intramolecular bonding within the hydrogel and 3400 cm -1Intermolecular hydrogel-silanol bonds were noted. A sample called 'hydrogel-coated surface (HFLAT)' consisting of hydrogel on a flat silicon wafer was observed at 3200 cm -1 showed a higher integral area (15.58) at 3400 cm, indicating strong intramolecular bonding within the hydrogel. In contrast, at 3400 cm -1 The integrated area was lower (4.56) in the hydrogel and untreated silicon wafer, indicating weaker intermolecular bonding between the hydrogel and the untreated silicon wafer, indicated by the 'smooth and flat surface (FLAT)'. This observation suggests limited interactions at the hydrogel-silicon interface. In contrast, the HiRINC sample showed a weaker interaction at 3200 cm -1 showed a slightly lower integral area (13.22) at 3400 cm, indicating a slight decrease in intramolecular bonding. However, at 3400 cm -1 A significant increase in the integral area (11.48) was observed, indicating that the intermolecular hydrogel-silanol bonds became stronger and thus the interaction with the porous silica nanostructures on the silicon wafer was enhanced.

[0151] To evaluate the adhesive properties of the samples, we focused on interfacial strength, bond strength, and work of adhesion (ΔE). These properties were evaluated using 90° peel tests and lap shear tests, as illustrated in Figures 9a and 9b. Figure 3a shows force versus displacement data from peel tests for two different surface types (HFLAT and HiRINC). The hydrogel-coated bare silicon wafer, referred to as HFLAT, consistently exhibited low force levels, indicating weak interaction with the flat surface. In contrast, HiRINC, which applied the hydrogel coating on a porous silica nanofilm, exhibited higher force levels, indicating strong adhesion due to its porous structure. The lap shear tests further elucidate the adhesive strength between the hydrogel and each surface type. For HFLAT, adhesive failure occurred at the point of maximum force, indicating detachment between the hydrogel and the substrate. In contrast, HiRINC achieved higher peak forces and exhibited stronger adhesion, with cohesive failure occurring within the hydrogel itself rather than at the interface, as illustrated in Figures 9c and d. Furthermore, Figures 9a and b contrast the stress behavior of HFLAT and HiRINC. HiRINC exhibited stronger adhesion with significantly greater hydrogel stretching between the substrates compared to HFLAT, suggesting that the nanostructured surface provides enhanced shear resistance. The work of adhesion calculated as the work of adhesion showed a larger area under the force-displacement curve for HiRINC than for HFLAT, indicating that more energy is required to peel the hydrogel from the nanostructured surface and thus the adhesion is stronger. The bar graph in Fig. 3b compares the interfacial strength, adhesion strength, and work of adhesion of HFLAT and HiRINC. In particular, the interfacial strengths of HiRINC and HFLAT were 7.6 N cm, respectively. -1 and 0.13 N cm -1 As such, the interfacial strength of HiRINC is more than 58 times that of HFLAT. This remarkable difference is also reflected in the bonding strength and work of adhesion, with the bonding strength of HiRINC (145 kPa) being 16 times greater than that of HFLAT (8.67 kPa), and the work of adhesion of HiRINC (1.63 kJ m -2 ) is HFLAT (0.025 kJ m -2 ) was shown to increase by approximately 650 times compared to the nanostructured surface, highlighting the excellent adhesion of the nanostructured surface.

[0152] As detailed in Figure 10, unlike a flat surface, the unique porous design appears to contribute to the strong adhesion of RINC by facilitating mechanical bonding. This hypothesis was further supported by 90° peel tests using various commercial tapes (Kapton, masking tape, Scotch, and double-sided tape), with RINC demonstrating greater interfacial strength in all tests. The present invention highlights the superior adhesive qualities of RINC, distinguishing it from other silica-based materials in adhesive technology applications.

[0153] In particular, the lap-shear fatigue strength of HFLAT and HiRINC was investigated to specifically evaluate the durability of this adhesion under cyclic stress. The inventors measured the extent to which the material moved from its original position under a constant force over several cycles of lap shear tests (Fig. 3c and d). While HFLAT easily desorbed the hydrogel under various levels of applied shear stress (Fig. 3c), HiRINC exhibited excellent durability and robustness, as the adhesion between the hydrogel and NANO remained strong and stable even under cyclic loading and high stress conditions (Fig. 3d).

[0154] 2.2 Durability and Stability Evaluation

[0155] Figures 4a and b show the results for HFLAT and HiRINC in their pristine and swollen states, respectively. Panel I shows the pristine state of the coating. Panel II shows the sample after cross-hatching and tape application, and panels III and IV show the substrate and tape, respectively, after peeling. For HFLAT, the hydrogel layer shows significant peeling, as evidenced by the visible removal in Figure 4a, and highlights the weak adhesion that remains consistent in the swollen state, causing peeling and making evaluation impossible (N / A). In contrast, HiRINC exhibits excellent adhesive properties by maintaining cohesion during tape removal in both the pristine and swollen states, highlighting the robust bond enhanced by the incorporation of the porous silica nanofilm.

[0156] Figures 11a to d detail the swelling dynamics of two surfaces, HFLAT and HiRINC, by tracking the changes in diameter (D0), thickness (T0), and weight (W0) over time. The summary in Figure 11d contrasts the swelling ratios between HFLAT and HiRINC. HFLAT exhibits a pronounced swelling ratio increase of up to 600% in area and thickness and up to 2000% in weight, leading to significant water uptake and delamination at the adhesive interface. In contrast, HiRINC exhibits a more controlled swelling response, maintaining a lower swelling ratio while increasing by 200% in area and thickness and 700% in weight.

[0157] To evaluate the tissue adhesion potential of HiRINC, quantitative adhesion strength tests were performed using ex vivo porcine stomach tissue (Figs. 5a to 5c). The results shown in Fig. 5a demonstrate the excellent adhesion performance of HiRINC, with an impressive adhesion strength of 65 kPa between the hydrogel and the tissue. In contrast, HFLAT, which includes hydrogel directly attached to the tissue, exhibited a significantly lower lap shear strength of 5 kPa, indicating adhesion failure at the surface. Figures 5b and 5c provide photographic evidence of actual adhesion tests on porcine tissue, further validating these results. The separation of the hydrogel layer from the substrate, indicated by the dotted circle, indicates immediate adhesion failure of HFLAT during the lifting process. In contrast, HiRINC maintains robust adhesion when force is applied, and the hydrogel remains attached to the substrate even when lifted from the tissue.

[0158] 2.3 Biomedical applications

[0159] Energy-dispersive X-ray spectroscopy (EDS) analysis confirmed the presence of silicon (Si) and carbon (C) elements corresponding to the porous silica nanofilm and the hydrogel layer, respectively. Figure 6a highlights the Si element distribution throughout the RINC-stent, while the SEM image shows the porous structure, and the C mapping shown together with the SEM image of the stent shown in Figure 6b confirms the presence of the hydrogel coating on the HiRINC-stent. Figures 6c and 6d illustrate the contrasting behavior of the hydrogel layer on the nitinol (NiTi) wire, showing both the HFLAT-NiTi wire and the HiRINC-NiTi wire. For the HFLAT-NiTi wire, the left image depicts the hydrogel layer in the dry state, whereas the right image shows significant swelling and subsequent delamination after immersion in water, indicating inadequate adhesion in the swollen state. In contrast, the HiRINC-NiTi wire maintained the integrity of the hydrogel layer in both dry and swollen states without any noticeable delamination. This contrasts with HFLAT, as HiRINC exhibited strong adhesion even in the swollen state, suggesting a role for the nanostructure in enhancing the bonding of the hydrogel and the NiTi wire. Furthermore, the HFLAT-stent exhibited significant hydrogel swelling, characterized by wrinkles, buckles, and delamination, as indicated by the dashed outline, suggesting instability when immersed in water (Fig. 6e). In contrast, the HiRINC-stent demonstrated effective hydrogel adhesion to the stent structure with minimal swelling (Fig. 6f). This indicates that the hydrogel is strongly and stably integrated into the stent even in the swollen state. These findings highlight the unique behavior of the HFLAT-stent and HiRINC-stent in water and emphasize the enhanced stability and integrity afforded by the nanostructured design.

[0160] In vivo adhesion strength was measured using a mechanical testing system equipped with a digital force gauge for fluoroscopic guidance and real-time quantitative evaluation (Fig. 7a). Control (n=3), HFLAT-stents (n=3), and HiRINC-stents (n=3) were inserted into the esophagus under fluoroscopic guidance. Before insertion, surgical sutures were attached to the tip of the stents, and the tip of the suture protruding out of the mouth was connected to a force gauge to measure the adhesion force. Side and top views are provided for clear visualization. Pulling the suture generated a force-displacement graph. Comparing the obtained smoothed maximum force values ​​(Fig. 7b), HiRINC showed the highest value, approximately 1000 kPa, compared to the control (uncoated) stent and the HFLAT-coated stent. These results suggest that the HiRINC-coated stent exhibits superior adhesion to esophageal tissue compared to the uncoated and HFLAT-coated stents. Figures 12a through 12c show a series of SEM images comparing the surface morphology of the bare stent (FLAT-stent), HFLAT-stent, and HiRINC-stent after in vivo lap shear testing. The SEM images of the HFLAT-stent show delamination of the hydrogel over a large area after testing. In contrast, the HiRINC-stent exhibits consistent hydrogel coating over most areas even after the mechanical stress of the lap shear test at low magnification. Higher magnification images still show nanostructures at the locations where the hydrogel detached. This integration suggests stronger adhesion of the hydrogel to the stent, potentially correlating with improved performance in a biomedical context.

[0161] 2.4 In Vivo Studies

[0162] During the 4-week follow-up period, stent migration occurred early in 5 (83.3%) rats in the control group. Except for one rat that initially exhibited complete stent migration, radiological examinations showed that the stent had partially migrated in 2 (33.3%) rats and had completely migrated in 2 (33.3%) rats at the last follow-up. More specifically, partial stent migration occurred in 4 rats at 3 or 14 days, and complete stent migration occurred in 2 of these rats at 4 weeks. In the HFLAT-stent group, partial migration occurred in 2 rats (2 / 6, 33.3%) at 4 weeks. In contrast, neither partial nor complete stent migration was observed in the HiRINC-stent group. The incidence of stent migration in the control group was significantly higher than that in the HiRINC-stent group (P = 0.014). The incidence rate in the HFLAT-stent group was higher than that in the HiRINC-stent group and lower than that in the control group, but there was no statistically significant difference (all P=0.439).

[0163] Figures 13a to 13f show that the pore diameter and thickness of the porous structure layer can be varied by controlling the reaction conditions, and thus the adhesive strength can be controlled. SEM images of RINCs with controlled pore diameter and thickness were observed from the surface (Figures 13a and d) and cross-section (Figures 13b and e), and the corresponding lap shear values ​​were measured (Figures 13c and f), demonstrating that the adhesive strength can be controlled through the reaction conditions. Figures 14a to d show a method for applying various nanostructures on the surface of a PTFE-coated nitinol stent. Figures 14a and c show bare PTFE surfaces with a noticeably uniform and smooth texture. Figures 14b and d show nanostructures after synthesis.

[0164] Figure 8a is a schematic diagram of the stent migration results after 4 weeks of stent, hydrogel-coated HFLAT-stent, and HiRINC-stent installation in the esophagus of rats, and Figure 8b is a comparative result through X-ray imaging for 4 weeks. In the control stent group, stent migration occurred in 5 rats (83.3%), and in the HFLAT group, partial migration occurred in 2 rats (33.3%) at 4 weeks. In contrast, no partial or complete stent migration was observed in the HiRINC-stent.

[0165] 3. Discussion

[0166] The adhesion enhancer RINC addresses these issues by bridging the gap between tissue and stents through hydrogel application, preventing migration and enabling noninvasive organ fixation within the rat esophagus. RINC exhibits mechanical interlocking after hydrogel impregnation, along with physical cross-linking via hydrogen bonding and van der Waals forces, ensuring robust adhesion to the hydrogel. After implantation, the hydrogel undergoes physical and covalent cross-linking via hydrogen and amide bonds, respectively, enhancing strong adhesion to biological tissue. This strategy differs from conventional mechanically-based stent fixation approaches. The present process involves coating a porous nanostructure onto a commercial medical substrate, followed by a simple hydrogel coating. This strategy ensures immediate bonding between the tissue and the hydrogel, enabling noninvasive and surgical-free fixation even after stent implantation. Furthermore, because there is no surgical suture, additional side effects that could lead to tissue inflammation can be prevented. Hydrogels not only enhance adhesion, but also provide mechanical elasticity, high biocompatibility, minimal fibroblast overgrowth, and an anti-inflammatory response.

[0167] Several considerations were taken into account in the design of HiRINC to facilitate clinical application. Initial efforts using clips, stent fixation, and barbs showed promising results in stent fixation. However, surgical complications related to tissue damage, the technically intensive manufacturing process, and, particularly, patient-specific requirements hindered widespread clinical application. HiRINC mitigates the risk of stent migration through perforated nanostructures that enhance the smooth surface and provide mechanical interlocking at the interface. Incorporating a flexible hydrogel into the deep pore channels of the nanostructures further enhances stability. This approach imparts enhanced adhesive stability and durability to HiRINC, effectively resisting challenges including delamination, shear stress, swelling, deswelling, and bending.

[0168] 4. Preparation of various hydrogels

[0169] 4.1 Preparation of polydopamine (PDA)-based hydrogel

[0170] Dopamine (30 mg) hydrochloride was added together with 22 mL of 1 M NaOH aqueous solution and stirred at 300 rpm for 20 minutes. Then, 10 g of acrylamide (AAm) was added and stirred for an additional 10 minutes. Ammonium persulfate (APS) (1000 mg) as an initiator, N,N-methylenebisacrylamide (24 mg) as a crosslinking agent, and triethanolamine (TEA) (20 μL) as a crosslinking catalyst were sequentially added and stirred further to homogenize the composition. The completed precursor composition was stored in a refrigerator and crosslinked under UV or temperature conditions just before use to form a hydrogel.

[0171] 4.2 Preparation of chitosan-based hydrogels

[0172] After adding 400 mg of chitosan and 16 mL of distilled water, 16 mL of acrylic acid (AAc) was added and stirred. Then, 240 mg of terephthalaldehyde (TPD) was added and stirred at room temperature. Then, 800 mg of polyethylene glycol (PEG) was added and stirred further. Then, 400 mg of ammonium persulfate (APS) and 200 mg of N,N-methylenebisacrylamide, a crosslinking agent, were sequentially added and stirred. The prepared hydrogel precursor solution was crosslinked under UV or temperature conditions to form a hydrogel.

[0173] 5. Application of various hydrogels and their evaluation

[0174] 5.1 Comparison of cross-cut adhesion performance of polydopamine (PDA)-based hydrogels with and without nanostructures

[0175] 15Х15 mm 2 Silica nanostructures were formed on a bare silicon (Si) wafer substrate using the sol-gel method, and a polydopamine (PDA)-based hydrogel was applied, followed by thermal curing at 70°C for 30 minutes. In addition, a PDA hydrogel was applied to the bare Si wafer substrate without forming silica nanospheres, and then thermal curing was performed under the same conditions.

[0176] A cross-cut test was performed. A blade was used to create a grid-like incision (crosshatch) at 5 mm intervals on the sample surface. Scotch tape was adhered to the sample surface under constant pressure for 10 seconds, and then peeled off in a 180° direction to evaluate the adhesion of the adhesive layer. As a result, the hydrogel on the nanostructured substrate remained firmly intact even at the cut grid area, with almost no signs of peeling observed. In contrast, on the bare substrate, the hydrogel was extensively peeled off along the cut area.

[0177] 5.2 Comparison of cross-cut adhesion performance of chitosan hydrogels with and without nanostructures

[0178] Chitosan-based hydrogels were applied to Si wafers with and without silica nanostructures, heat-cured at 75°C for 30 minutes, and then cross-cut tests were performed. Excellent hydrogel retention was observed on the cut surface on the silica nanostructured substrate, while clear peeling was observed along the cut line on the bare substrate.

[0179] 5.3 Mechanical adhesion analysis of PDA hydrogels using lap shear tests

[0180] After forming PDA hydrogels on bare and silica nanostructured Si wafers, respectively, lap shear tests were performed. For the lap shear test, two substrates coated with cured hydrogels were overlapped to an area of ​​1 cm × 1 cm, and a load was applied at a rate of 50 mm / min in the shear direction using a universal testing machine to measure the failure load and displacement. Even under conditions of relatively low PDA contents (0.12 w / w% and 0.25 w / w%), the maximum shear load on the silica nanostructured surface increased by approximately 2.8 times (0.12 w / w%) and approximately 6 times (0.25 w / w%), respectively, and the failure displacement was also extended. The higher the PDA content, the more enhanced the adhesion and energy absorption, and the mechanical reinforcing effect of the silica nanostructure was clearly observed (Figures 15a and b).

[0181] 5.4 Analysis of peel resistance according to peel test of PDA hydrogel

[0182] After forming PDA-based hydrogels on the surfaces of bare Si wafers and silica nanostructured Si wafers, peel tests were performed. The peel resistance of the bonded hydrogel specimens with an area of ​​1.5 cm × 1 cm was evaluated by pulling them perpendicular to the shear direction at a rate of 13 mm / min. As a result, as shown in Figure 16, the peak peel load on the nanostructured substrate was approximately 13 times higher than that on the bare substrate, and a load-displacement curve was observed in which a high load was maintained throughout the entire peeling section and gradually decreased. This is attributed to the mechanical interlocking and enhanced interfacial bonding strength of the nanostructured surface.

[0183] 5.5 Lap shear test and peel test analysis of chitosan hydrogel

[0184] Chitosan-based hydrogels were formed on the surfaces of bare Si wafers and silica nanostructured Si wafers, respectively. Lap shear tests were performed on specimens measuring 1 cm × 1 cm at a rate of 50 mm / min, and peel tests measured the peel load on a 1.5 cm × 1 cm area at a rate of 13 mm / min. As shown in Figures 17a and b, the maximum shear load on the nanostructured substrate increased by approximately 6 times compared to the bare substrate, and the fracture displacement was also improved. In the peel test, the peak peel load increased by approximately 16 times or more, and a curve in which the average load continued at a high level was observed. This is attributed to the improved mechanical interlocking and interfacial bonding.

[0185] 5.6 Comparison of adhesive strength according to swelling time

[0186] The PVA-AAc-based hydrogel precursor was coated on bare Si wafers and porous silica nanostructured Si wafers, respectively, and fixed by UV curing. Each specimen was then swollen in distilled water for different periods (e.g., 1 h, 6 h, and 12 h) and then subjected to lap shear tests. The tests were performed at a speed of 50 mm / min over an overlapping area of ​​1 cm × 1 cm. As shown in Figure 18, the silica nanostructured substrate maintained high adhesion even after long-term swelling for more than 12 h, whereas the adhesion of the bare substrate rapidly decreased. This suggests that mechanical interlocking and interfacial bonding effects contribute to the enhanced interfacial stability even in an underwater environment.

[0187] 5.7 Visualization of drug loading within the porous structural layer

[0188] FITC, a fluorescent substance, was spin-coated on bare Si wafers and silica nanostructured Si wafers, respectively, at 9,000 rpm for 1 minute. Fluorescence was then observed via UV irradiation. While minimal or no fluorescence was observed on the bare substrate, a strong fluorescence signal was observed on the nanostructured substrate. This is believed to be due to the penetration and fixation of fluorescent molecules into the pores of the porous structural layer (porous silica nanostructure) (Fig. 19).

[0189] In addition, to evaluate the drug loading and release performance, the PDA-based hydrogel was hardened into a 0.5 cm disk shape and then loaded with the drug by swelling in 500 μL of a 10 mM Rhodamine B solution for 1 hour. The sample was immersed in distilled water for 24 hours and photographed. Initially, the dye was uniformly distributed within the gel, and over time, release into the external medium was observed, showing a shape similar to that of a hydrogel without dye loading. This result visually demonstrates that the hydrogel of the present invention can be utilized as a drug delivery platform (Fig. 19).

[0190] 5.8 Evaluation of the applicability of the spray coating process for thin-film hydrogel coatings

[0191] The hydrogel precursor solution used in 1.6 was sprayed onto a substrate formed with silica nanostructures (i.e., a silicon wafer on which RINCs were synthesized) using a spray method. Coating was performed under conditions of a spray distance of approximately 10 cm and a spray time of 5 seconds. As a result, a uniform, thin hydrogel film was formed across the entire substrate, and penetration into the silica nanostructures was confirmed (Fig. 20). By adjusting the spray coating conditions (pressure, distance, and spray time), precise control of the coating thickness and penetration depth was possible. This demonstrates that the hydrogel composition of the present invention is suitable for application to a thin film coating process.

[0192] 9) Evaluation of changes in hydrogel adhesive strength according to the desorption solution

[0193] A PVA-AAc-based hydrogel was formed on a Si wafer with silica nanostructures, and the adhesive strength was measured through a lap shear test. The same specimen was then immersed in a 0.5 M sodium bicarbonate (NaHCO3) aqueous solution to induce a desorption reaction of the hydrogel. It was confirmed that the adhesive strength of the hydrogel decreased within approximately 5 to 10 minutes. This indicates that the hydrogel composition of the present invention is desorbable under specific conditions (Fig. 21a and b).

[0194] The above description of the present invention is for illustrative purposes only, and those skilled in the art will readily appreciate that the present invention can be readily modified into other specific forms without altering the technical spirit or essential characteristics of the present invention. Therefore, the embodiments described above should be understood as illustrative in all respects and not restrictive. For example, each component described as a single entity may be implemented in a distributed manner, and similarly, components described as distributed may be implemented in a combined manner.

[0195] The scope of the present invention is indicated by the claims described below rather than the detailed description above, and all changes or modifications derived from the meaning and scope of the claims and their equivalent concepts should be interpreted as being included in the scope of the present invention.

Claims

1. Description; A porous structural layer formed on the above substrate; and A hydrogel layer formed on the porous structure layer and containing a hydrogel A composite adhesive structure comprising:

2. In paragraph 1, A composite adhesive structure, wherein the porous structure layer comprises a substrate-derived porous structure layer or a coating-type porous structure layer.

3. In paragraph 1, A composite adhesive structure, wherein the thickness of the porous structural layer is 5 nm to 1000 nm.

4. In paragraph 1, A composite adhesive structure, wherein the diameter of the pores of the porous structural layer is 1 nm to 500 nm.

5. In paragraph 1, A composite adhesive structure in which the hydrogel layer is formed by penetrating into the pores of the porous structural layer.

6. In paragraph 1, A composite adhesive structure in which the porous structural layer and the hydrogel layer are chemically and physically bonded.

7. In paragraph 1, The above hydrogel is a composite adhesive structure comprising at least one hydrophilic polymer selected from polyvinyl alcohol (PVA), alginate (Alg), polyethylene glycol (PEG), chitosan, gelatin, polyacrylic acid (PAAc), polyacrylamide (PAM), poly(PNIPAM), agar, poly(2-acrylamido-2-methylpropanesulfonic acid) (PAMPS), polyvinylpyrrolidone, polyethylene oxide, methyl vinyl ether maleic anhydride copolymer, isobutylene maleic anhydride copolymer, methacrylic acid butyl acrylate copolymer, methoxyethylene maleic anhydride copolymer, polyurethane, hyaluronic acid, alginic acid, dopamine, cellulose, collagen, and derivatives thereof.

8. In paragraph 1, A composite adhesive structure, wherein the thickness of the hydrogel layer is 10 nm to 5 mm.

9. In paragraph 1, A composite adhesive structure, wherein the above-mentioned substrate comprises at least one selected from an alloy, a biodegradable polymer, and a non-degradable polymer.

10. A medical device comprising a composite adhesive structure according to paragraph 1.

11. In paragraph 10, The above medical device is a medical device that is an implantable device.

12. In paragraph 10, The medical device is a non-vascular stent, a vascular stent, an artificial prosthesis, a catheter, a medical dressing, a film, a fiber, a mesh, a powder, a microsphere, a sheet, a sponge, a foam, a suture anchoring device, a conduit, a plate and a screw, a drug delivery device, or an anti-adhesion membrane. 13.(a) a step of forming a porous structure layer on the substrate; and (b) a step of forming a hydrogel layer on the porous structure layer; A method for manufacturing a composite adhesive structure, comprising:

14. In paragraph 13, A method for manufacturing a composite adhesive structure, wherein the step (a) comprises surface-treating the substrate; or forming a coating-type porous structure layer on the surface of the substrate.

15. In paragraph 13, A method for producing a composite adhesive structure, wherein the hydrogel comprises at least one hydrophilic polymer selected from poly(vinyl alcohol; PVA); alginate (Alg), polyethylene glycol (PEG), chitosan (Chitosan), gelatin, polyacrylic acid (PAAc), polyacrylamide (PAM), poly(PNIPAM), agar, poly(2-acrylamido-2-methylpropanesulfonic acid) (PAMPS), polyvinylpyrrolidone, polyethylene oxide, methyl vinyl ether maleic anhydride copolymer, isobutylene maleic anhydride copolymer, methacrylic acid butyl acrylate copolymer, methoxyethylene maleic anhydride copolymer, polyurethane, hyaluronic acid, alginic acid, dopamine, cellulose, collagen, and derivatives thereof.

16. In paragraph 13, A method for manufacturing a medical material, wherein the above step (b) is performed by a dip-coating, cast-coating, drop-coating, spray-coating or spin-coating method.

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