Patent Application
20230414147
The present application claims priority to Korean Patent Application No. 10-2022-0076834, filed Jun. 23, 2022, the entire contents of which is incorporated herein for all purposes by this reference.
The present invention relates to a device for measuring biosignals and electrical stimulation, and more particularly, to a device for measuring electrical stimulation biosignals with low mechanical properties, excellent stress relief properties, and sustainable conductivity by incorporating a conductive composite including self-healing polymers and liquid metals.
In measuring biosignals and electrical stimulation, there is a demand for technologies related to flexible conductive materials.
Conventionally, a technology utilizing conductive materials including metals such as gold and silver has been used for measuring biosignals and electrical stimulation, but these conductive materials containing inorganic materials lose flexibility when the tissue is dynamic or deformed, leading to potential noise generation during biosignal measurement, which may result in the low signal-to-noise ratio, difficulties in accurate biosignal measurement, possible physical cracks in the conductive material itself, as well as physical cracks in the conductive materials themselves as well as in other structures in the device.
To address the issue of reduced flexibility in dynamic or deformable tissues, Thesis 1 (Sim, K. et al. An epicardial bioelectronic patch made from soft rubbery materials and capable of spatiotemporal mapping of electrophysiological activity. Nat. Electron. 3, 775-784 (2020)) introduced a technology in which an inherently flexible conductive composite material composed of one-dimensional silver (Ag) nanomaterials encapsulated in a gold (Au) shell and the rubbery polymer is applied to the epicardial mesh for a bidirectional electrical interface.
However, the technology introduced in thesis 1 has the drawback that fatigue caused by continuous cardiac contractions may lead to cracks in the polymer matrix. This issue may arise from the limited ability of the low-ductility polymer to efficiently release accumulated stresses within the rigid domains of the composite material.
Meanwhile, in the field of measuring biosignals and electrical stimulation, there is also a need for a technology allowing for uniform adhesion to tissue surfaces without the need for suturing.
However, conventional film-like polymers lack the necessary flexibility and conformability to wrap around the curved surfaces of tissues, making them unsuitable as films for implantable devices.
Additionally, conventional tissue adhesives require external factors such as temperature, ultraviolet (UV) light, or pressure, resulting in longer adhesion times and potential risks of tissue damage. Furthermore, in cases where tissues are not immobilized and exhibit non-periodic or periodic movement, the films or devices adhered to curved surfaces of the tissue may have a risk of detachment.
To address the mechanical adaptation and fixation issues related to dynamic tissues, thesis 2 (Lee, H. et al. Mussel-inspired surface chemistry for multifunctional coatings. Science 318, 426-430 (2007)) introduced an elastomer-based electronic device with an adhesive hydrogel layer incorporating polydopamine nanoparticles, inspired by mussels.
However, in thesis 1, improved adhesion strength and durability in dynamic tissues are required. Specifically, film-like elastomeric materials encounter challenges in conforming to the irregular curved surfaces of the epicardial tissue and maintaining integrity without cracks in films with nanometer thickness.
To address the above issues, the inventor of the present invention has developed a device for measuring biosignals and elastic electrical stimulation that exhibits high tissue adhesion strength and a conformable nature to cover curved surfaces comprehensively.
The present invention has been conceived to solve the above problems, and it is an object of the present invention to provide a device for measuring biosignals and electrical stimulation with the inclusion of a network fiber layer including hydrogen bondable self-healing polymers and a hydrogel coating layer with tissue-adhesive hydrogel coating on the fiber layer, the hydrogel being characterized by infiltrating into the fiber layer.
The technical objects of the present invention are not limited to the aforesaid, and other objects not described herein with can be clearly understood by those skilled in the art from the descriptions below.
In order to accomplish the above objects, an embodiment of the present invention provides a device for measuring biosignals and electrical stimulation that incorporates conductive composite including self-healing polymers and liquid metal.
In an embodiment of the present invention, the liquid metal may be a eutectic alloy.
In an embodiment of the present invention, the liquid metal may include one or more selected from the group consisting of eutectic gallium-indium alloy, eutectic gallium-tin alloy, eutectic gallium-indium-tin alloy, and gallium.
In an embodiment of the present invention, the liquid metal may be 82 to 88 wt % in content relative to the total weight of the conductive composite, denoted as 100 wt %.
In an embodiment of the present invention, the liquid metal may be dispersed within the conductive composite.
In an embodiment of the present invention, the conductive composite may be disposed on a network fiber layer that includes fibers made of the self-healing polymer.
In an embodiment of the present invention, the conductive composite may be coupled to the network fiber layer.
In an embodiment of the present invention, the device may include a hydrogel coating layer coated on the network fiber layer and the conductive composite.
In an embodiment of the present invention, the hydrogel may be infiltrated into the fiber layer.
In an embodiment of the present invention, the network fiber layer may be porous.
In an embodiment of the present invention, the network fiber layer may include fibers with a diameter of 4.5 to 6.5 μm and pores with a diameter of 50 to 70 μm.
In an embodiment of the present invention, the hydrogel may penetrate less than 60% of the thickness of the network fiber layer from the contact surface with the network fiber layer.
In an embodiment of the present invention, the hydrogel may be a catechol-bonded polymer.
In an embodiment of the present invention, the hydrogel may undergo hydrogen bonding or hydrophobic interactions with tissues.
In an embodiment of the present invention, the self-healing polymer may include a polymer main chain, a first structural unit containing —HN—C(═O)—NH— capable of forming strong hydrogen bonds, and a second structural unit containing —HN—C(═O)—NH— capable of forming weak hydrogen bonds.
In an embodiment of the present invention, the polymer main chain may include at least one selected from polysiloxane and polydialkylsiloxane (where alkyl is C1 to C6), such as polydimethylsiloxane, polyethylene oxide (PEO), polypropylene oxide (PPO), polybutylene oxide (PBO), perfluoropolyether (PFPE), polyolefin, poly(ethylene-co-1-butylene), polybutadiene, hydrogenated polybutadiene, poly(ethylene oxide)-poly(propylene oxide) copolymer, poly(hydroxyalkanoate), styrene-butadiene copolymer (SB), styrene-butadiene-styrene copolymer (SBS), styrene-ethylene-butylene-styrene copolymer (SEBS), ethylene propylene diene rubber (EPDR), acrylic rubber, polychloroprene rubber, polyurethane, fluoro-rubber, butyl rubber, or silicone rubber.
In an embodiment of the present invention, the first structural unit may be represented by Formula 1:
where Ar is a substituted or unsubstituted arylene group of C6 to C30 or a planar heteroarylene group of C3 to C30.
In an embodiment of the present invention, the second structural unit may be represented by Formula 2-1:
where a is an integer in the range of 5 to 20.
In an embodiment of the present invention, the second structural unit may be represented by Formula 2-2:
where Cy is a substituted or unsubstituted cyclic alkylene group of C5 to C30, and b is an integer of 1 or 3.
In an embodiment of the present invention, the biosignal may be a biosignal of human or animal tissue.
In an embodiment of the present invention, the biosignal may be a biosignal of the heart tissue in human or animal.
In order to accomplish the above object, an embodiment of the present invention provides a method for manufacturing an elastic device for measuring biosignals and electrical stimulation, the method including fabricating a network fiber layer by electrospinning self-healing polymer, arranging a conductive composite including a self-healing polymer and liquid metal onto the network fiber layer, and forming a hydrogel coating layer coated with hydrogel on the network fiber layer and the conductive composite.
In an embodiment of the present invention, the biosignal may be a biosignal of human or animal tissue.
In an embodiment of the present invention, the biosignal may be a biosignal of the heart tissue in human or animal.
In order to accomplish the above objects, an embodiment of the present invention provides a device for drug delivery or implantable in a body, the device including the device for measuring a biosignal of electrical stimulation, which is manufactured by the method for manufacturing a device for measuring biosignals and electrical stimulation.
Specifically, E-SHN and the EGaIn composites alone were used as control samples.
Specifically, in the absence of Alg-CA, the EGaIn composite lost charge injection characteristics, and after several stimulations, the stimulation artifact (gray asterisk) and heart rhythm (arrow) were completely separated. The EGaIn composite with Alg-CA demonstrated stable myocardial capture (arrow) following the stimulation artifact.
Hereinafter, the present invention will be described with reference to the accompanying drawings. However, the present invention may be embodied in many different forms and is not limited to the embodiments described herein. In order to clearly describe the present invention, parts irrelevant to the description may be omitted in the drawings, and similar reference numerals may be used for similar components throughout the specification.
Throughout the specification, when a part is said to be “connected (coupled, contacted, or combined)” with another part, this is not only “directly connected”, but also “indirectly connected” with another member in between. Also, when a part is said to “comprise” a certain component, this means that other components may be further included instead of excluding other components unless specifically stated otherwise.
The terminology used herein is for the purpose of describing particular embodiments only and is not intended to be limiting of the invention. As used herein, the singular forms are intended to include the plural forms as well, unless the context clearly indicates otherwise. It will be further understood that the terms “comprises” or “has,” when used in this specification, specify the presence of a stated feature, number, step, operation, component, element, or a combination thereof, but they do not preclude the presence or addition of one or more other features, numbers, steps, operations, components, elements, or combinations thereof.
Hereinafter, embodiments of the present invention will be described in detail with reference to the accompanying drawings.
In order to solve the above problems, an embodiment of the present invention provides a device for measuring biosignals and electrical stimulation that incorporates conductive composite including self-healing polymers and liquid metal.
With reference to
In an embodiment of the present invention, the liquid metal may be an eutectic alloy. By including the eutectic alloy, the liquid metal within may remain in a liquid state within the body.
In an embodiment of the present invention, the liquid metal may include one or more selected from the group consisting of eutectic gallium-indium alloy, eutectic gallium-tin alloy, eutectic gallium-indium-tin alloy, and gallium. By including the metal, the conductive composite may be maintained as a liquid metal within the body, and the use of gallium-based eutectic alloys or gallium in the liquid metal is suitable for the electrodes and wiring of the biosignal measuring device due to their biocompatibility.
In an embodiment of the present invention, the liquid metal content may be 82 to 88 wt % relative to the total weight of the conductive composite, denoted as 100 wt %. Liquid metal content below 82 wt % may compromise the balance between elasticity and conductivity, while liquid metal content exceeding 88 wt % may interfere with the self-healing of the polymers through dynamic hydrogen bonding, potentially leading to liquid metal leakage.
In an embodiment of the present invention, the liquid metal may be dispersed within the conductive composite. By dispersing the liquid metal within the conductive composite, it is possible to effectively deliver electrical stimulation within the desired range of biological tissues and receive corresponding biological electrical signals.
In an embodiment of the present invention, the conductive composite may be positioned on a network fiber layer that includes fibers made of the self-healing polymer. The conductive composite attached onto the network fiber layer may be conformally mounted on curved surfaces of biological tissue.
In an embodiment of the present invention, the conductive composite may be coupled to the network fiber layer. The fibers and conductive composite share the same self-healing polymer composition, allowing fibers and conductive composite to bond together through mutual coupling when the conductive composite is arranged on the network fiber layer.
In an embodiment of the present invention, the device may include a hydrogel coating layer coated on the network fiber layer and the conductive composite.
With reference to
In an embodiment of the present invention, the hydrogel may be infiltrated into the fiber layer.
With reference to
The physically entangled structure, formed by the hydrogel wrapping around the fibers and interacting either through intermolecular interactions within the hydrogel or with the fibers, may provide a physically bonded structure with the fibers serving as a supporting framework.
Preferably, the hydrogel infiltrates the fiber layer, forming a physically entangled structure with a portion of the fiber layer, which desirable as it may help maintain the inherent elasticity of the fiber layer while providing strong tissue adhesion. More preferably, the hydrogel may be penetrated less than 60% of the thickness of the network fiber layer from the contact surface with the network fiber layer. When the hydrogel penetrates more than 60% of the network fiber layer thickness, it may be inefficient in terms of cost and slightly reduce the flexibility of the device itself.
Moreover, the device for measuring biosignals and electrical stimulation possesses viscosity and flexibility, allowing the device to relieve stress from external forces and adapt to the shape of curved human or animal tissues, ensuring a close contact without any gaps between the device and the tissue surface.
In an embodiment of the present invention, the network fiber layer may be porous. The porous network fiber layer may be manufactured through electrospinning, and the porous structure allows for easy penetration of the hydrogel. When a device containing a non-porous fiber layer is used to adhere to tissues, the device may detach from the tissue over time.
In an embodiment of the present invention, the network fiber layer may include fibers with a diameter of 4.5 to 6.5 μm and pores with a diameter of 50 to 70 μm. A fiber diameter smaller than 4.5 μm may result in weak strength and potential breakage, and a fiber diameter larger than 6.5 μm may lead to reduced elasticity.
In an embodiment of the present invention, the hydrogel may be a catechol-bonded polymer.
In an embodiment of the present invention, the hydrogel may exhibit hydrogen bonding or hydrophobic interactions with tissues. The hydrogel may adhere quickly to the tissue due to hydrogen bonding or hydrophobic interactions, and intermolecular hydrogen bonding within the hydrogel may provide elasticity. The hydrophobic interactions may involve interactions between the hydrophobic ring structure of catechol groups and hydrophobic amino acids in the tissue. The main chain of the hydrogel may be alginic acid, and the functional group may include a catechol derivative including polyphenols containing pyrogallol moiety.
In an embodiment of the present invention, the self-healing polymer may include a polymer main chain, a first structural unit containing —HN—C(═O)—NH— capable of forming strong hydrogen bonds, and a second structural unit containing —HN—C(═O)—NH— capable of forming weak hydrogen bonds.
With reference to
The polymer main chain of the self-healing polymer may be derived from a homopolymer, copolymer, or terpolymer that possesses flexibility, and the polymer may also be a crosslinked copolymer, block copolymer, or random copolymer, among others, without being particularly limited. For example, the polymer main chain may include at least one selected from polysiloxane and polydialkylsiloxane (where alkyl is C1 to C6), such as polydimethylsiloxane, polyethylene oxide (PEO), polypropylene oxide (PPO), polybutylene oxide (PBO), perfluoropolyether (PFPE), polyolefin, poly(ethylene-co-1-butylene), polybutadiene, hydrogenated polybutadiene, poly(ethylene oxide)-poly(propylene oxide) copolymer, poly(hydroxyalkanoate), styrene-butadiene copolymer (SB), styrene-butadiene-styrene copolymer (SBS), styrene-ethylene-butylene-styrene copolymer (SEBS), ethylene propylene diene rubber (EPDR), acrylic rubber, polychloroprene rubber, polyurethane, fluoro-rubber, butyl rubber, or silicone rubber. The polyolefin may be selected from polyethylene (PE), polypropylene (PP), polybutylene (PB), copolymers thereof, and mixtures thereof. The poly(ethylene oxide)-poly(propylene oxide) copolymer may be a block copolymer or a random copolymer.
The self-healing polymer includes a first structural unit containing —HN—C(═O)—NH— moiety capable of forming strong hydrogen bonds connected to the polymer main chain and a second structural unit containing —HN—C(═O)—NH— moiety capable of forming weak hydrogen bonds. The first structural unit capable of forming strong hydrogen bonds may form hydrogen bonds with all carbonyl groups of —HN—C(═O)—NH— moieties in one chain with all amino groups of —HN—C(═O)—NH— moieties in another chain. Additionally, the second structural unit capable of forming weak hydrogen bonds may include at least one portion that cannot form hydrogen bonds with carbonyl groups of —HN—C(═O)—NH— moieties in another chain while having amino groups of —HN—C(═O)—NH— moieties in one chain.
The strong hydrogen bonds provide elasticity and mechanical strength, while the weak hydrogen bonds contribute to extensibility and energy dissipation performance.
The first structural unit includes an aromatic ring and may be represented by the following Formula 1.
In Formula 1, —Ar— represents a substituted or unsubstituted arylene group of C6 to C30 or a planar heteroarylene group of C3 to C30.
The arylene group may be a single aromatic ring, a condensed ring formed by the fusion of two or more aromatic rings, or a combination of two or more aromatic rings connected by functional groups selected from single bonds, fluorenyl groups, cycloalkylene groups with a carbon chain length between C1 and C10, —O—, —S—, —C(═O)—, —CH(OH)—, —S(═O)2-, —Si(CH3)2-, —(CH2)p- (where 1≤p≤10), —(CF2)q- (where 1≤q≤10), —C(CH3)2-, —C(CF3)2-, —C(═O)NH—, and combinations thereof.
In Formula 1, —Ar— may include an aromatic group selected from substituted or unsubstituted phenylene, substituted or unsubstituted naphthylene, substituted or unsubstituted anthrylene, substituted or unsubstituted phenanthrylene, substituted or unsubstituted pyrenylene, substituted or unsubstituted peylenylene, substituted or unsubstituted fluorenylene, substituted or unsubstituted pentalene, substituted or unsubstituted pyrazole, substituted or unsubstituted imidazole, substituted or unsubstituted thiadiazole, substituted or unsubstituted triazole, substituted or unsubstituted carbazole, substituted or unsubstituted pyridine, substituted or unsubstituted pyridazine, substituted or unsubstituted pyrimidine, substituted or unsubstituted pyrazine, substituted or unsubstituted triazine, substituted or unsubstituted indazole, substituted or unsubstituted indoline, substituted or unsubstituted benzimidazole, substituted or unsubstituted benzothiazole, substituted or unsubstituted thienothiophene, substituted or unsubstituted benzothiophene, substituted or unsubstituted benzofuran, substituted or unsubstituted isoquinoline, and substituted or unsubstituted purine, or an aromatic group formed by connecting two or more of the aforementioned aromatic groups using a linker.
Here, the linker may be selected from a single bond, a substituted or unsubstituted alkylene group of C1 to C30, and a substituted or unsubstituted alkylene group of C1 to C30 including at least one link selected from —O—, —NRa—, —C(═O)—, —OC(═O)—, —S(═O)2—, —Si(RaRb)2—, and —C(═O)NRd— (where Ra, Rb, Rc, and Rd are independently selected among hydrogen, alkyl groups of C1 to C10, alkenyl groups of C2 to C10, alkynyl groups of C2 to C10, aryl groups of C6 to C18, and heteroaryl groups of C2 to C18).
In Formula 1, —Ar— may be represented by the following Formula 1A or 1B.
In Formula 1A, R1 to R3 are independently hydrogen or alkyl groups of C1 to C6, and a, b, and c are integers representing the number of hydrogen atoms attached to the aromatic ring (e.g., they may be integers 0 to 4, 0 to 3, 0 to 2, or 1).
Ra and Rb are independently hydrogen, halogen, alkyl groups of C1 to C6, or fluoroalkyl groups of C1 to C6.
In Formula 1B, R1 to R3 are independently hydrogen or alkyl groups of C1 to C6, and a, b, and c are integers representing the number of hydrogen atoms attached to the aromatic ring (e.g., they may be integers 0 to 4, 0 to 3, 0 to 2, or 1).
Ra and Rb are independently hydrogen, halogen, alkyl groups of C1 to C6, or fluoroalkyl groups of C1 to C6.
In Formulas 1A and 1B, the bond (indicated by *) connected to —HN—C(═O)—NH— may be arranged to have linearity, promoting the linear nature of the self-healing polymer. The aromatic rings of different chains in Formulas 1A and 1B may be oriented in a way that facilitates stacking and favorable hydrogen bonding.
The structural units that confer linearity to the self-healing polymer may be represented by Formula 1AA or 1BB.
In Formula 1AA, R1 to R3 are independently hydrogen or alkyl groups of C1 to C6, and a, b, and c are integers representing the number of hydrogen atoms attached to the aromatic ring (e.g., these may be integers 0 to 4, 0 to 3, 0 to 2, or 1), Ra and Rb are independently hydrogen, halogen alkyl groups of C1 to C6, or fluoroalkyl groups of C1 to C6.
In Formula 1BB, R1 to R3 are independently hydrogen or alkyl groups of C1 to C6, and a, b, and c are integers representing the number of hydrogen atoms attached to the aromatic ring (e.g., they may be integers 0 to 4, 0 to 3, 0 to 2, or 1), Ra and Rb are independently hydrogen, halogen alkyl groups of C1 to C6, or fluoroalkyl groups of C1 to C6.
The second structural unit includes an aliphatic chain or an aromatic (non-aromatic) ring system and may be represented by Formula 2-1 or 2-2.
In Formula 2-1, the integer‘a’ is in the range of 5 to 20.
In Formula 2-2, ‘Cy’ represents a substituted or unsubstituted cyclic alkylene group of C5 to C30, and ‘b’ is an integer of 1 or 3. ‘Cy’ in Formula 2-2 may be represented by Formula 2-2A, 2-2B, or 2-2C.
In Formula 2-2A, R1 may be hydrogen or an alkyl group of C1 to C6, a represents the number of hydrogen atoms attached to the cyclohexyl ring, and L1 and L2 are independently a single bond or a substituted or unsubstituted alkylene group of C1 to C10.
In Formula 2-2B, R2 may be hydrogen or an alkyl group of C1 to C6, b represents the number of hydrogen atoms attached to the cyclohexyl ring, and L3 and L4 are independently a single bond or a substituted or unsubstituted alkylene group of C1 to C10.
In Formula 2-2C, R3 and R4 are independently hydrogen or alkyl groups of C1 to C6, c and d represent the number of hydrogen atoms attached to the cyclohexyl ring, Ra and Rb are independently hydrogen, halogen, alkyl groups of C1 to C6, or fluoroalkyl groups of C1 to C6, and L5 and L6 are independently a single bond or a substituted or unsubstituted alkylene group of C1 to C10.
As a specific example of Formula 2-2A, it is possible to consider the Formula 2-2AA, as a specific example of Formula 2-2B, it is possible to consider the Formula 2-2BB, and as a specific example of Formula 2-2C, it is possible to consider the Formula 2-2CC.
In Formula 2-2AA, 2-2BB, and 2-2CC, the hydrogen atoms present in each ring may be substituted with alkyl groups of C1 to C6.
When the structural unit is a moiety derived from 4,4′-methylenebis(phenylurea) (4,4′-first polydialkylsiloxane, and the polymer main chain is methylenebis(phenyl urea), MPU), and the second structural unit is a moiety derived from isophorone bisurea (IU), the self-healing polymer may include the repeating units of Formulas 3-1 and 3-2.
Here, n-1 in Formula 3-1 and in n-2 in Formula 3-2 are independently integers ranging from 10 to 200, e.g., from 20 to 150, 20 to 100, or 30 to 40.
In practical implementation, the relative mole ratio between the first structural unit, which provides strong hydrogen bonding, and the second structural unit, which provides weak hydrogen bonding, may be adjusted according to the desired properties of the composite. Increasing the content of the first structural unit enhances mechanical strength (Young's modulus and fracture energy), while increasing the content of the second structural unit improves self-healing rate and self-healing ability. For example, the first structural unit and the second structural unit may be included in the mole ratio of approximately 0.2:0.8 to 0.8:0.2, 0.3:0.7 to 0.7:0.3, 0.4:0.6 to 0.6:0.4, or 0.5:0.5.
The first structural unit and the second structural unit may be included in self-healing polymer in an amount of at least 0.01 mmol per mole of self-healing polymer, e.g., at least 0.02 mmol, at least 0.03 mmol, at least 0.04 mmol, or at least 0.05 mmol, and in an amount of up to 10 mmol, e.g., up to 9 mmol, up to 8 mmol, or up to 7 mmol, respectively. By controlling the composition within the aforementioned ranges, the elasticity, mechanical strength, and self-healing ability of the composite containing the self-healing polymer may be easily adjusted.
The self-healing polymer may have a number average molecular weight (Mn) in the range of approximately 10,000 or higher, e.g., around 11,000 or 12,000 or higher, and up to approximately 100,000 or lower, e.g., around 105,000, 110,000, or 115,000 or lower. When the number average molecular weight falls within the mentioned range, the self-healing polymer may exhibit excellent transparency and self-repairing characteristics. The number average molecular weight may be the polystyrene equivalent average molecular weight measured by gel permeation chromatography.
The polymer main chain may be derived from polymers with low glass transition temperatures. The glass transition temperature of the polymer may be around −40° C. or higher, e.g., around −30° C. or −20° C. or higher, and up to approximately 40° C. or lower, e.g., around 30° C. or 20° C. or lower. When the glass transition temperature falls within this range, the self-healing polymer may exhibit excellent transparency and self-restoration properties.
The multiple non-covalently hydrogen bonding interactions may lead to the crosslinking of the polymer main chains in the self-healing polymer, forming a supramolecular network (supramolecular elastomer network). That is, the chains of the self-healing polymer may self-assemble through continuous crosslinking via hydrogen bonding between —HN—C(═O)—NH— moieties. Such polymers crosslinked through hydrogen bonding exhibit faster self-healing rates and superior extensibility and self-recovery performance compared to structures crosslinked by covalent bonding. Accordingly, the composite material containing the self-healing polymer may exhibit high stretchability even under ambient conditions, allowing for rapid restoration of the network fiber layers incorporating the self-healing polymer during repetitive stretching, and it also possesses self-recovery characteristics, allowing for easy recombination and restoration of the original composite even upon damage. In one implementation, the composite material containing the self-healing polymer exhibits approximately 75% self-healing efficiency after 48 hours at room temperature (25° C.) and approximately 100% self-healing efficiency after 6 hours at 60° C.
The stretchability and self-recovery capabilities of the self-healing polymer are exemplified. The exemplified polymer has a polydimethylsiloxane (PDMS) backbone, with the first structural unit derived from 4,4′-methylenebis(phenyl urea) (MPU) and the second structural unit derived from isophorone bisurea (IU).
When external stress is applied, the weak hydrogen bonding may preferentially break before the strong hydrogen bonding, allowing for stretchability and strain energy dissipation.
The self-healing polymer forms a crosslinked matrix structure.
The tissue may be human or animal tissue, particularly rapidly pulsating cardiac tissue.
Hereinafter, a method for manufacturing an elastic device for measuring biosignals and electrical stimulation is described.
According to another embodiment of the present invention, a method for manufacturing an elastic device for measuring biosignals and electrical stimulation includes fabricating a network fiber layer by electrospinning self-healing polymer, arranging a conductive composite including a self-healing polymer and liquid metal onto the network fiber layer, and forming a hydrogel coating layer coated with hydrogel on the network fiber layer and the conductive composite.
Electrospinning the self-healing polymer creates pores in the fiber layer to fabricate a porous fiber layer, which allows for infiltration of hydrogel through these pores, leading to enhancement of the elasticity and tissue adhesion of the device used for measuring biosignals and electrical stimulation. In an embodiment of the present invention, the biosignals may be biosignals of human or animal tissues.
In an embodiment of the present invention, the biosignals may be biosignals of the heart tissue in human or animal.
Meanwhile, the device for measuring biosignals and electrical stimulation, manufactured by the method of the present invention, may be used as a device as an implantable device in the body.
Hereinafter, the above-described exemplary implementations are described in more detail through embodiments. However, the following embodiments are provided for illustrative purposes only and do not limit the scope of the invention.
The solution of H2N-PDMS-NH2 (100 g, Mn=6000, 1 eq, Gelest Inc.) was prepared by dissolving it in anhydrous CHCl3 (400 mL) and adding 10 mL of Et3N under argon at 0° C. After stirring for 1 hour, a mixture of 4,4′-methylenebis(phenyl isocyanate) (2.0 g, 0.4 eq.) and isophorone diisocyanate (2.7 g, 0.6 eq.) was added dropwise to the solution in CHCl3. The mixture was stirred for 1 hour while maintaining a temperature of 0° C. using an ice bath. The resulting solution was kept at room temperature and stirred for 4 days. After the reaction, MeOH (15 mL) was added to remove the remaining isocyanate, and the mixture was stirred for 30 minutes. The obtained solution was concentrated to half of its volume, and 60 mL of MeOH was poured onto the precipitate. A white precipitate-like viscous liquid was formed, and the mixture was left undisturbed for 30 minutes. The clear solution on the top was decanted into another container, and 100 mL of CHCl3 was added to dissolve the product. The solvent and a small amount of Et3N were removed by vacuum evaporation after repeating the dissolution-precipitation-decantation process three times. The final product (PDMS-MPU0.4-IU0.6 self-healing polymer (referred to as SHP), was obtained with a yield of 65 g (63%). The molecular weight of the obtained self-healing polymer was as follows (measured by GPC, Mw=103,400; Mn=65,000 (D=1.6)).
1H NMR (400 MHz, d5-THF): δ7.33 (d, J=8.0 Hz, 4H), 6.97 (d, J=8.0 Hz, 4H), 3.77 (s, 2H), 0.01 (b, 1325H). 13C NMR (400 MHz, CDCl3): δ158.78, 139.18, 137.31, 125.34.
Material Acquisition
The sodium alginate (intermediate viscosity), dopamine hydrochloride, N-hydroxysuccinimide (NHS), 2-(N-morpholino)ethanesulfonic acid (MES solution), 1-dodecanethiol, and octadecyltrimethoxysilane (OTMS, 376213) were obtained from Sigma-Aldrich (USA). 1-Ethyl-3-(3-dimethylaminopropyl)carbodiimide hydrochloride (EDC) was obtained from Tokyo Chemical Inc. (Japan). PDMS (Sylgard 184) was obtained from The Dow Chemical Company (USA). EGaIn (eutectic gallium-indium alloy) was purchased from Alfa Aesar (USA). Ag flakes (AgFs, DSF-500MWZ-S) were purchased from Daejoo Electronics (South Korea). Chloroform was purchased from Samchun Chemical (South Korea).
The liquid metal (eutectic gallium-indium, EGaIn) weighing 1 g was added to 5 mL of chloroform containing 0.5M 1-dodecanethiol, and the solution was uniformly dispersed using a tip sonicator (VCX750, SONICS & MATERIALS, INC., USA) with a power of 300 W for 5 minutes (10 seconds on, 5 seconds off pulses). After ultrasonic treatment, the solution was centrifuged at 4000 rpm for 5 minutes to remove the supernatant and precipitate the EGaIn. The precipitated EGaIn nano-/omicro-particles were washed several times with chloroform and then redispersed in the same solvent. Next, 1 mL of the redispersed EGaIn solution was mixed with 3 mL of chloroform containing dissolved SHP, and the mixture was stirred for 2 hours. All EGaIn and AgF composite solutions were cured at room temperature on OTMS-treated silicon wafers. After curing, a weak pressure was applied to induce conductivity pathways in the composite. Finally, the microstructure of the composite was observed using a field-emission scanning electron microscope (FE-SEM).
Except for using a PDMS solid film as the substrate, the manufacturing process followed the same method described in Manufacturing Example 1.
Except for dispersing AgF instead of EGaIn using a tip sonicator and omitting the process of separating and washing the solvent using a centrifuge, the manufacturing process followed the same method described in Manufacturing Example 1.
The SHP (1 g) was dissolved in 30 mL of CHCl3, and the SHP solution was prepared by stirring for 3 hours. Subsequently, the solution was drop-casted onto a Teflon dish with a diameter of 10 cm to produce a solid film. The network fiber layer (E-SHN) was fabricated by electrospinning the SHP solution.
Specifically, the SHP (1 g) was dissolved in 5 mL of CHCl3 and stirred for 3 hours. The SHP solution (3 mL) was filled into a syringe equipped with a 21-gauge metal needle and connected to a syringe pump (New Era-300 Pump System, USA). Aluminum foil was applied to a grounded flat collector (NanoNC, South Korea). For electrospinning, the solution was ejected at a rate of 1.4 mL/h−1 with a distance of 11 cm between the needle tip and the plate collector, at a voltage of 13 kV. Finally, 0.4 mL of the SHP solution was electrospun, and the resulting network fiber layer (E-SHN) had a thickness of 100 μm.
Alg-CA was prepared through the coupling reaction of 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide hydrochloride (EDC) and N-hydroxysuccinimide (NHS) at pH 4.5 to 4.8.
Specifically, alginate was dissolved in MES buffer (0.1M, adjusted to pH 4.5-4.8) to obtain a final concentration of 1% (w/v). The solution was then purged with nitrogen for 30 minutes. Dopamine hydrochloride, EDC, and NHS were added to the solution in an equimolar ratio with the carboxylic acid groups of the alginate backbone. The reaction was carried out at room temperature for 12 hours. To remove unreacted dopamine, the mixture was dialyzed in acidic distilled water (pH 5-5.5) for 12 hours and subsequently freeze-dried.
To coat the Alg-CA onto the substrate where the conductive composite is attached to the E-SHN, SHP was dissolved in distilled water at a concentration of 3% by weight. The substrates, measuring 1.5×1.5 cm2, were treated with oxygen plasma at 100 W for 6 minutes using a Femto Science plasma system (Korea). Immediately after plasma treatment, the Alg-CA solution was loaded onto the substrates and left at room temperature overnight for complete dehydration.
FE-SEM (JSM-7600F, JEOL) was used to observe the fine surface structure of the network fiber layer (E-SHN) and the EGaIn/SHP composite material and the cross-sectional structure of the network fiber layer (E-SHN) coated with Alg-CA.
To theoretically analyze the stress-relieving effect of the viscoelastic material, finite element analysis (FEA) was conducted using commercial software (ANSYS, Ansys Inc., Canonsburg, PA, USA) assuming that PDMS polymer and E-SHN are subjected to repetitive pressure on heart tissue. The analysis results were presented in
To investigate the tissue adhesion strength of the device for measuring biological signals of electrical stimulation of embodiment 1, porcine heart tissue (1×1 cm2 or 1×2 cm2) was affixed to a polymeric substrate (e.g., polyethylene terephthalate (PET) film) using commercial super glue. Additionally, the E-SHN, EGaIn/SHP composite material and solid film (1×1 cm2 or 1×2 cm2) attached to the substrate were compressed against the heart tissue. Shear stress or cyclic stress was applied to the terminals of the substrate using a 50N load cell at a speed of 10 mm/min−1. Adhesion strength (kPa) and elongation (mm) were monitored using Instron Bluehill software. The test results were presented in
For qualitative adhesion testing, heart tissue was dissected from 11-week-old Sprague Dawley rats. Alg-CA (1×3 cm2) was swiftly attached to the heart tissue and immediately pulled upward to evaluate the tissue adhesion of SAFIE and E-SHN. The same adhesion test was performed after rinsing with phosphate-buffered saline (PBS, 1×, pH 7.4). The test results were presented in
To measure the compatibility of E-SHN and solid films, models with E-SHN and solid film of thicknesses 100 μm or 200 μm were fabricated using a 3D printer and observed.
The tensile stress-strain curves of the solid film, E-SHN, and EGaIn/SHP composite material were recorded using a universal testing machine (UTM; Instron 34SC-1, USA). Samples with a width of 5 mm or 20 mm, thickness of 100 μm or 200 μm, and length of 10 mm were tested at a deformation rate of 20 mm/min−1 (50N load cell). All experiments were performed four times to calculate the Young's modulus and residual strain of each sample. Furthermore, all experiments were performed with a lengthwise notch of 10 mm in the center of each sample to calculate the fracture strength of each sample. The results of the experiments were presented in
First, the electrical resistance of the stretched EGaIn/SHP nano/micro composite material was measured using a digital multimeter (Keithley 2450 source meter, Tektronics). The sample (initial length 3 mm, width 3 mm, thickness 150 μm) was fixed onto a motorized stage (SMC-100, Jeil Optical Systems) using double-sided tape and elongated at a speed of 3 mm/min−1. The results of the experiment were presented in
Second, for cyclic testing, the resistance was measured at 0% and 50% strain, and the samples were subjected to 1,000 cycles at a speed of 20 mm/min−1. In all cases, EGaIn was used to form deformable electrical contacts between the composite and the multimeter wires. The results of the experiment were presented in
Third, the electrochemical impedance and cyclic voltammetry of the composite were monitored using a potentiostat (ZIVE sp1, ZIVE LAB) in the frequency range of 1 Hz to 1 kHz (potential range: −1.2 V to +1.2 V, scan rate: 50 mV/s−1). The sample (0.5×0.5 cm2 area) was immersed in PBS. The Ag/AgCl reference electrode and platinum counter electrode were used. The results of the experiment were presented in
The EGaIn/SHP composite, cut into a size of 0.5×1.5 cm2, was placed on SHP and PDMS substrates measuring 1×2.5 cm2 and, after approximately 10 minutes, subjected to a 100% strain using a manual stretcher from an initial length of approximately 5 mm to observe whether separation occurred between the composite and the substrates. Furthermore, after removing the sample from the stretcher, the substrate and composite were pulled in opposite directions using tweezers to confirm whether separation occurred. The results of the experiment were presented in
To assess the biocompatibility of the SAFIE device developed in this study, cell tests were conducted using mouse fibroblast cells (L929). Both E-SHN and EGaIn/SHP composite materials, the combined SAFIE device, were tested, and the number of living cells and dead cells were compared after a maximum of 72 hours to obtain the results. The results of the experiment were presented in
To conduct FTIR and XPS analysis, a sample containing only EGaIn particles, a sample with EGaIn-Alg-CA binding, and a sample containing only Alg-CA were prepared. First, the first sample was prepared by adding 36 mg of EGaIn to 3 ml of ultrapure water, and the second sample was prepared by adding 36 mg of EGaIn to 3 ml of 0.3 w/v % Alg-CA solution, and both samples were sonicated by a tip sonicator for 1 hour. The third sample was prepared using 0.3 w/v % Alg-CA solution. Each sample was dropped onto a 1×1 cm2 glass substrate, dried completely, and then subjected to analysis using an X-ray Photoelectron Spectrometer (XPS, ESCALAB250, Thermo, USA) and Fourier-Transform Infrared Spectrometer (FTIR, IFS-66/S, TENSOR27, Bruker, USA). The results of the experiment were presented in
To confirm the proper functioning of the device in vivo, the device was attached to the hearts of actual 11-week-old Sprague Dawley rats, and electrocadiogram was measured using a data acquisition system (Powerlab 8/35, AD Instruments, Australia), applying square pulses of 1.5 Vpp, 1 ms pulse width, and 6 Hz frequency using a pulse stimulator (model 2100, AM systems). Additionally, the actual device was implanted in the rats for 3 days, and ECG measurements were taken without anesthesia to compare the effects of Alg-CA coating. The rats were intravenously injected with a drug (diltiazem), and the induced arrhythmias were measured. The results of the above experiments are presented in
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A device for measuring for measuring biosignals and electrical stimulation according to an embodiment of the present invention is advantageous in terms of achievement of low mechanical properties, excellent stress relief properties, and sustainable conductivity by incorporating a conductive composite including self-healing polymers and liquid metals.
It should be understood that the advantages of the present invention are not limited to the aforesaid but include all advantages that can be inferred from the detailed description of the present invention or the configuration specified in the claims.
The above description of the present invention is for illustrative purposes only, and it will be understood by those skilled in the art that various modifications and changes may be made thereto without departing from the spirit and scope of the invention. Therefore, it should be understood that the embodiments described above are exemplary and not limited in all respects. For example, each component described as a single type may be implemented in a distributed manner, and similarly, components described as distributed may be implemented in a combined form.
The scope of the invention should be determined by the appended claims, and all changes or modifications derived from the meaning and scope of the claims and equivalent concepts thereof should be construed as being included in the scope of the present invention.
| Number | Date | Country | Kind |
|---|---|---|---|
| 10-2022-0076834 | Jun 2022 | KR | national |
