Drug delivery carrier with blood-brain barrier penetration technology and method for preparing the same

Abstract

The present invention relates to a drug delivery system capable of crossing the blood-brain barrier and a method for manufacturing the same. More specifically, it relates to a stimulus-responsive drug delivery system that conditionally induces the passage of an active drug component through the blood-brain barrier in response to stimulation of the microenvironment surrounding the blood-brain barrier, and a method for manufacturing the same.

Description

Technology Field

[0001] The present invention relates to a drug delivery system capable of crossing the blood-brain barrier and a method for manufacturing the same. More specifically, it relates to a stimulus-responsive drug delivery system that conditionally induces the passage of an active drug component through the blood-brain barrier in response to stimulation of the microenvironment surrounding the blood-brain barrier, and a method for manufacturing the same. Background Technology

[0002] The blood-brain barrier is a complex biological barrier composed of cerebral capillary endothelial cells, astrocytes, peristaltic cells, and basement membranes, which performs the function of strictly controlling the transport of substances between the bloodstream and the brain parenchyma. The intercellular gaps of blood-brain barrier endothelial cells are sealed by tight junction proteins such as claudin-5, occludin, and ZO-1, and this tight junction structure effectively blocks the intercellular transport of water-soluble substances with a molecular weight of approximately 400 Da or more. Consequently, most active ingredients of drugs developed for the treatment of brain diseases are difficult to reach effective concentrations in brain tissue after systemic administration, and the blood-brain barrier acts as a critical technical obstacle in the drug treatment of central nervous system diseases such as glioblastoma, Alzheimer's disease, and Parkinson's disease.

[0003] Approaches to overcoming the blood-brain barrier can be broadly categorized into physical and biochemical methods. Physical methods studied include the temporary opening of tight junctions by focused ultrasound, endothelial cell contraction by osmotic shock, and direct intracerebral injection using microneedles. Biochemical methods reported include strategies to induce receptor-mediated transcytosis by conjugating ligands targeting receptors expressed on blood-brain barrier endothelial cells, such as transferrin receptors, insulin receptors, and LRP-1 receptors, to the surface of nanoparticles; strategies for adsorption-mediated transcytosis using cell-permeable peptides; and strategies using nanoparticles mimicking high-density lipoproteins.

[0004] Nanoparticle-based drug delivery systems have been designed to adopt a core-shell structure to encapsulate the active drug component within the core, prevent drug degradation during blood circulation through the shell, and induce passage through the blood-brain barrier by placing target ligands on the surface. Mesoporous silica, polymer nanoparticles, liposomes, and dendrimers have been studied as core materials for these delivery systems, while polyethylene glycol, chitosan, and polylactic acid-glycolic acid copolymers have been used as shell materials.

[0005] However, conventional nanoparticle-based drug delivery strategies have primarily focused on optimizing the means by which the carrier crosses the blood-brain barrier. When target ligands are constantly exposed on the surface of the carrier, non-specific interactions with endothelial cells at non-target sites occur during blood circulation, which tends to shorten the carrier's circulation time in the blood or accumulate in reticuloendothelial organs such as the liver and spleen. Additionally, if the active drug component is released prematurely before the carrier reaches the blood-brain barrier at the disease site, the concentration of the active drug reaching brain tissue is reduced, leading to fluctuations in therapeutic efficiency.

[0006] Meanwhile, disease states such as glioblastoma or neuroinflammation are accompanied by biochemical changes in the microenvironment surrounding the blood-brain barrier that are distinct from normal conditions. In disease sites, a decrease in pH due to anaerobic metabolism, an increase in reactive oxygen species concentration due to oxidative stress, and the overexpression of matrix metalloproteinases involved in tissue remodeling may be observed. These changes in the microenvironment imply that the blood-brain barrier in disease sites is under different physicochemical conditions compared to that in normal sites. Nevertheless, research on technical approaches in which the behavior of drug delivery systems is conditionally switched according to the state of the blood-brain barrier to selectively induce the passage of active ingredients has not yet been sufficiently advanced. Accordingly, there is a need for a technology that can improve the permeability of active ingredients across the blood-brain barrier while reducing unnecessary brain exposure, based on a condition-based perspective regarding the state of the blood-brain barrier under which the passage of active ingredients is induced. The problem to be solved

[0007] The present invention aims to provide a drug delivery system capable of conditionally inducing the passage of an active drug component through the blood-brain barrier in response to stimulation of the microenvironment surrounding the blood-brain barrier.

[0008] In addition, the present invention aims to provide a core-shell structured drug delivery system capable of stably protecting the active drug component during blood circulation while reducing the resistance of the active drug component to pass through the blood-brain barrier by converting the structure of the stimulus-responsive shell and releasing the mask of the responsive binding structure under stimulation of the microenvironment surrounding the blood-brain barrier.

[0009] In addition, the present invention aims to provide a drug delivery system in which a blood-brain barrier crossing step and a drug release step are linked so that an active ingredient of a drug can be delivered to a target site within brain tissue.

[0010] In addition, the present invention aims to provide a method for manufacturing a drug delivery system that can reduce unnecessary brain exposure and reduce variability in drug delivery efficiency.

[0011] However, the problems that the present invention aims to solve are not limited to those described above, and may include all problems that a person skilled in the art can recognize from the description in this specification. means of solving the problem

[0012] According to one aspect of the present invention, a drug delivery system capable of penetrating the blood-brain barrier may be provided, comprising: a core containing a drug active ingredient; a stimulus-responsive shell formed on the outer side of the core; and a responsive binding structure disposed on the outer surface of the stimulus-responsive shell; wherein the stimulus-responsive shell undergoes a structural transition from a shielded state to a relaxed state upon stimulation of the microenvironment surrounding the blood-brain barrier, and the responsive binding structure comprises a masking linker and a target ligand coupled to the masking linker.

[0013] For example, a drug delivery system capable of crossing the blood-brain barrier may be provided, wherein the stimulation of the microenvironment surrounding the blood-brain barrier is one or more of a change in pH, a change in reactive oxygen species concentration, and a change in enzyme concentration.

[0014] For example, the masking linker may be an enzymatically degradable peptide linker, and a blood-brain barrier penetration technology drug delivery system may be provided.

[0015] For example, a blood-brain barrier penetration drug delivery system may be provided, wherein the core comprises porous inorganic nanoparticles and the active drug component is supported within the pores of the porous inorganic nanoparticles.

[0016] For example, the porous inorganic nanoparticles mentioned above may be mesoporous silica nanoparticles, and a blood-brain barrier penetration technology drug delivery system may be provided.

[0017] For example, a drug delivery system capable of penetrating the blood-brain barrier may be provided, wherein the pore diameter of the mesoporous silica nanoparticles is 1 to 20 nm.

[0018] For example, a blood-brain barrier penetration technology drug delivery system may be provided, wherein the pore diameter is 2 to 10 nm.

[0019] For example, a blood-brain barrier penetration technology drug delivery system may be provided, wherein the pore diameter is 3 to 6 nm.

[0020] For example, a drug delivery system capable of penetrating the blood-brain barrier may be provided, wherein the particle size of the mesoporous silica nanoparticles is 30 to 500 nm.

[0021] For example, a blood-brain barrier penetration technology drug delivery system may be provided, wherein the particle size is 50 to 300 nm.

[0022] For example, a blood-brain barrier penetration technology drug delivery system may be provided, wherein the particle size is 80 to 200 nm.

[0023] For example, the stimulus-responsive shell may be provided as a blood-brain barrier penetration technology drug delivery system comprising a block copolymer including a pH-responsive segment and a hydrophilic segment.

[0024] For example, a blood-brain barrier crossing technology drug delivery system may be provided, wherein the block copolymer is poly(â-aminoester)-block-polyethylene glycol.

[0025] For example, a blood-brain barrier penetration drug delivery system may be provided, which corresponds to the structural transition in the pH range of 6.0 to 7.0 of the microenvironment surrounding the blood-brain barrier in the above pH change.

[0026] For example, a blood-brain barrier penetration technology drug delivery system may be provided, wherein the pH is in the range of 6.5 to 6.8.

[0027] For example, the above enzymatically degradable peptide linker may be a substrate metalloproteinase degradable peptide linker, and a blood-brain barrier penetration technology drug delivery system may be provided.

[0028] For example, the above substrate metalloproteinase-degradable peptide linker may be provided as a blood-brain barrier-crossing drug delivery system comprising a peptide sequence that is degraded by one or more of MMP-2 and MMP-9.

[0029] For example, the target ligand may be a blood-brain barrier-crossing drug delivery system, which is a transferrin receptor-binding peptide.

[0030] As an example, a blood-brain barrier crossing technology drug delivery system may be provided, further comprising a metal-organic framework nanovalve layer disposed at the mesopore entrance of the mesoporous silica nanoparticles.

[0031] For example, the metal-organic framework may be a zeolite imidazolate framework containing zinc ions and 2-methylimidazole, and a blood-brain barrier crossing technology drug delivery system may be provided.

[0032] For example, the metal-organic framework nanovalve layer may be provided as a blood-brain barrier penetration drug delivery system that decomposes in the pH range of 6.0 to 7.0.

[0033] For example, the metal-organic framework nanovalve layer may be provided as a blood-brain barrier penetration drug delivery system that decomposes in the pH range of 6.5 to 6.8.

[0034] For example, a blood-brain barrier penetration drug delivery system may be provided, further comprising a phenylboronic acid residue disposed on the outer surface of the stimulus-responsive shell.

[0035] For example, a blood-brain barrier-crossing drug delivery system may be provided, which further includes a catechol group in the stimulus-responsive shell.

[0036] For example, a blood-brain barrier penetration drug delivery system may be provided, which further comprises trehalose co-encapsulated with the drug active ingredient within the mesopores of the mesoporous silica nanoparticles, wherein the trehalose forms a glassy matrix within the mesopores.

[0037] For example, a blood-brain barrier penetration drug delivery system may be provided, wherein the content of the trehalose is 0.5 to 20 times by weight relative to the weight of the active ingredient of the drug.

[0038] For example, a blood-brain barrier penetration drug delivery system may be provided, wherein the content of the trehalose is 2 to 10 times by weight relative to the weight of the active ingredient of the drug.

[0039] For example, a blood-brain barrier penetration drug delivery system may be provided, wherein the content of the trehalose is 3 to 7 times by weight relative to the weight of the active ingredient of the drug.

[0040] For example, a blood-brain barrier penetration drug delivery system may be provided, further comprising a catalyst layer containing Prussian blue nanoparticles on the outermost layer of the stimulus-responsive shell.

[0041] For example, a blood-brain barrier penetration drug delivery system may be provided, wherein the content of the Prussian blue nanoparticles is 0.5 to 15 weight percent relative to the total weight of the drug delivery system.

[0042] For example, a blood-brain barrier penetration technology drug delivery system may be provided, wherein the content of the above-mentioned Prussian blue nanoparticles is 2 to 8 weight percent relative to the total weight of the drug delivery system.

[0043] According to one aspect of the present invention, a method for manufacturing a drug delivery system capable of penetrating the blood-brain barrier is provided, comprising: a core forming step of forming a core; a drug encapsulating step of encapsulating a drug active ingredient in the core; a shell forming step of forming a stimulus-responsive shell on the outer side of the core encapsulated with the drug active ingredient; and a surface functionalization step of forming a responsive binding structure including a masking linker and a target ligand on the outer surface of the stimulus-responsive shell, wherein the stimulus-responsive shell undergoes a structural transition from a shielded state to a relaxed state upon stimulation of the microenvironment surrounding the blood-brain barrier.

[0044] For example, a method for manufacturing a blood-brain barrier-crossing drug delivery system may be provided, wherein, in the shell-forming step, the stimulus-responsive shell comprises a block copolymer including a pH-responsive segment and a hydrophilic segment.

[0045] For example, in the surface functionalization step, a method for manufacturing a blood-brain barrier penetration drug delivery system may be provided, wherein the masking linker is an enzymatically degradable peptide linker.

[0046] For example, a method for manufacturing a blood-brain barrier-crossing drug delivery system may be provided, wherein the core formation step comprises the step of synthesizing mesoporous silica nanoparticles by the sol-gel method.

[0047] For example, a method for manufacturing a blood-brain barrier-crossing drug delivery system may be provided, wherein the shell forming step comprises the step of coating the outer surface of the core with poly(â-aminoester)-block-polyethylene glycol.

[0048] For example, a method for manufacturing a blood-brain barrier-crossing drug delivery system may be provided, further comprising a nanovalve layer forming step prior to the shell forming step, wherein zinc ions and 2-methylimidazole are reacted at the mesopore entrances of the mesoporous silica nanoparticles to form a nanovalve layer comprising a zeolite imidazolate framework.

[0049] For example, a method for manufacturing a blood-brain barrier-crossing drug delivery system may be provided, wherein the surface functionalization step comprises the step of introducing a phenylboronic acid residue to the outer surface of the stimulus-responsive shell and the step of introducing a catechol group to the stimulus-responsive shell.

[0050] For example, a method for manufacturing a blood-brain barrier-crossing drug delivery system may be provided, wherein the drug encapsulation step comprises the step of co-encapsulating the drug active ingredient and trehalose within the mesopores of the mesoporous silica nanoparticles.

[0051] For example, a method for manufacturing a blood-brain barrier-crossing drug delivery system may be provided, wherein the trehalose is co-encapsulated at a weight ratio of 0.5 to 20 times the weight of the active ingredient of the drug.

[0052] For example, a method for manufacturing a blood-brain barrier-crossing drug delivery system may be provided, further comprising a vitrification step after the drug encapsulation step, wherein a vitrified matrix of the trehalose is formed within the mesopore by freeze-drying or vacuum drying.

[0053] For example, a method for manufacturing a blood-brain barrier-crossing drug delivery system may be provided, further comprising a catalyst layer forming step in which, after the surface functionalization step, a catalyst layer containing Prussian blue nanoparticles is formed on the outermost surface of the stimulus-responsive shell. Effects of the invention

[0054] According to the present invention, since the passage of a drug active ingredient through the blood-brain barrier can be conditionally induced in response to stimulation of the microenvironment surrounding the blood-brain barrier, the penetration rate of the drug active ingredient through the blood-brain barrier can be improved.

[0055] In addition, according to the present invention, under normal blood flow conditions, the active ingredient of the drug is protected by the stimulus-responsive shell and the target ligand is masked and maintained in an inactive state, so non-specific interactions at non-target sites are reduced and unnecessary brain exposure can be reduced.

[0056] In addition, according to the present invention, a structure in which the surface characteristics of a drug delivery vehicle are switched based on the state of the blood-brain barrier can be adopted, so the variability of drug delivery efficiency can be reduced.

[0057] In addition, according to the present invention, since the blood-brain barrier crossing step and the drug release step can be linked, the active ingredient of the drug can exert its effect at a target site within brain tissue, and the drug delivery efficiency for treating brain diseases can be improved.

[0058] However, the effects of the present invention are not limited to those mentioned above, and other unmentioned effects will be clearly understood by a person skilled in the art from the description below. Brief explanation of the drawing

[0059] FIG. 1 is a flowchart illustrating a method for manufacturing a drug delivery system capable of penetrating the blood-brain barrier according to one embodiment of the present invention. Specific details for implementing the invention

[0060] In one aspect of the present invention, a drug delivery system capable of penetrating the blood-brain barrier may comprise a core, a stimulus-responsive shell, and a reactive binding structure. The core may be a structure that encapsulates a drug active ingredient inside. The stimulus-responsive shell may be formed on the outer side of the core and surround the core. The reactive binding structure may be disposed on the outer surface of the stimulus-responsive shell. The drug delivery system may be configured so that the drug active ingredient encapsulated in the core passes through the blood-brain barrier and is delivered to brain tissue.

[0061] In one embodiment, the core may be a particulate structure having an internal space capable of accommodating a drug active ingredient. The drug active ingredient may be an active substance delivered to brain tissue to exhibit a pharmacological effect, and may be one or more of a low-molecular-weight compound, a peptide, a protein, a nucleic acid, and an antibody. The core may physically confine the drug active ingredient within or encapsulate it through chemical interaction, thereby reducing the degradation or inactivation of the drug active ingredient by blood enzymes or plasma proteins during blood circulation. The material of the core may be an inorganic material, an organic polymer material, or a composite material thereof. The core may have a porous structure, a hollow structure, or a matrix structure, and a suitable structure may be selected according to the molecular weight and polarity of the drug active ingredient.

[0062] For example, the particle size of the core may be in the range of 10 to 1000 nm, and in one embodiment, may be in the range of 30 to 500 nm. When the core particle size is within the above range, the drug delivery vehicle can secure an opportunity to contact blood-brain barrier endothelial cells while avoiding early capture by the reticuloendothelial system.

[0063] In one embodiment, the stimulus-responsive shell may structurally transition from a shielded state to a relaxed state upon stimulation of the microenvironment surrounding the blood-brain barrier. The shielded state may be a state in which the stimulus-responsive shell adheres to the outer surface of the core to isolate the active drug component inside the core from the external environment. The relaxed state may be a state in which the structure of the stimulus-responsive shell swells, dissolves, or undergoes phase separation, allowing the active drug component inside the core to be exposed or released to the outside. The stimulation of the microenvironment surrounding the blood-brain barrier is a physicochemical condition different from the normal blood flow environment and may occur around the blood-brain barrier at the disease site. Since the stimulus-responsive shell maintains a shielded state under normal blood flow conditions where the stimulation is not present, the unnecessary release of the active drug component at non-target sites or non-selective exposure to brain tissue may be reduced.

[0064] For example, the stimulus-responsive shell may be formed of a polymer material, and the polymer may be a stimulus-responsive polymer that changes solubility, charge state, or molecular chain orientation in response to stimulation of the microenvironment surrounding the blood-brain barrier. The thickness of the stimulus-responsive shell may be in the range of 5 to 100 nm, and in one embodiment, may be in the range of 10 to 50 nm.

[0065] In one embodiment, the reactive binding structure may include a masking linker and a target ligand bound to the masking linker. The masking linker may be a binding structure that connects the target ligand to the outer surface of the stimulus-responsive shell and may be degraded or cleaved by stimulation of the microenvironment surrounding the blood-brain barrier. Before the masking linker is degraded or cleaved, the target ligand may be sterically shielded by the masking linker and maintained in an inactive state. After the masking linker is degraded or cleaved, the target ligand is exposed and may bind to a specific receptor on the surface of the blood-brain barrier endothelial cells, and receptor-mediated transcytosis may be initiated by binding to the receptor, allowing the drug delivery vehicle to cross the blood-brain barrier.

[0066] For example, the masking linker may be a cleavable linker composed of a peptide bond, an ester bond, a disulfide bond, or a combination thereof. The target ligand may be one of a peptide, an antibody fragment, an aptamer, or a small molecule ligand that specifically binds to a receptor expressed on blood-brain barrier endothelial cells. The reactive binding structure may be arranged on the outer surface of the stimulus-responsive shell at a density of 1 to 10,000 / particle, and in one embodiment, at a density of 100 to 5,000 / particle.

[0067] In one embodiment, the drug delivery system may exist in an inactive state in which the active drug component is protected and non-target interactions are suppressed by the shielding state of the stimulus-responsive shell and the shielding of the target ligand by the masking linker during circulation in the bloodstream. When the drug delivery system reaches the microenvironment surrounding the blood-brain barrier, a structural transition to a relaxed state of the stimulus-responsive shell and the degradation of the masking linker proceed, allowing the active drug component to pass through the blood-brain barrier and be delivered to a target site within the brain tissue. After passing through the blood-brain barrier, the active drug component is released from the brain tissue and can exert a therapeutic effect at the target site. Accordingly, the penetration rate of the active drug component through the blood-brain barrier is improved, and the drug delivery efficiency for treating brain diseases can be improved.

[0068] In one aspect of the present invention, the stimulation of the microenvironment surrounding the blood-brain barrier may be one or more of a change in pH, a change in reactive oxygen species concentration, and a change in enzyme concentration. The microenvironment surrounding the blood-brain barrier may exhibit physicochemical conditions different from the normal blood flow environment due to neuroinflammation, brain tumor, cerebral ischemia, or neurodegenerative disease, and the change in pH, the change in reactive oxygen species concentration, and the change in enzyme concentration may be biological signals that have a significant deviation from the normal state in these disease states.

[0069] In one embodiment, the change in pH may refer to a pH condition lower than the normal blood flow pH. The pH of normal blood flow may be in the range of 7.35 to 7.45, and in the neuroinflammation or brain tumor microenvironment, the pH may be lowered due to anaerobic metabolism and lactic acid accumulation. The change in reactive oxygen species concentration may refer to a condition in which the concentration of reactive oxygen species is elevated compared to the normal blood flow environment, and the reactive oxygen species may include one or more of superoxide anions, hydrogen peroxide, and hydroxyl radicals. The change in enzyme concentration may refer to a condition in which the expression level of a specific protease around the blood-brain barrier is elevated compared to the normal state.

[0070] For example, the stimulus-responsive shell may undergo structural transformation in response to a single stimulus among the pH change, the reactive oxygen species concentration change, and the enzyme concentration change, or it may undergo structural transformation under conditions where two or more stimuli are present simultaneously. When the stimulus-responsive shell is configured to respond to two or more stimuli, non-specific activation by a single stimulus is reduced, thereby improving the selectivity of the drug's active ingredient for delivery into the brain.

[0071] In one aspect of the present invention, the masking linker may be an enzymatically degradable peptide linker. The enzymatically degradable peptide linker may include a sequence in which amino acid residues are linked by peptide bonds and may be selectively cleaved by a specific protease present in the microenvironment surrounding the blood-brain barrier.

[0072] In one embodiment, the amino acid sequence length of the enzymatically degradable peptide linker may be 4 to 20 residues, and a recognition site for the protease may be included within the sequence. The enzymatically degradable peptide linker may not be substantially cleaved in a normal blood flow environment because the concentration of the protease is low, and cleaving may be initiated under conditions where the concentration of the protease is elevated around the blood-brain barrier of the disease site. By cleavage, the target ligand may be exposed, enabling binding to a surface receptor of the blood-brain barrier endothelial cell.

[0073] For example, the above enzymatically degradable peptide linker may include a Pro-Leu-Gly-Leu-Ala-Gly sequence, a Gly-Pro-Leu-Gly-Val-Arg-Gly-Lys sequence, or a modified sequence thereof.

[0074] In one aspect of the present invention, the core may comprise porous inorganic nanoparticles, and the drug active ingredient may be supported within the pores of the porous inorganic nanoparticles. The porous inorganic nanoparticles may be nanoscale particles of an inorganic material having regular or irregular pores inside, and the pores may function as receiving spaces for the drug active ingredient.

[0075] In one embodiment, the porous inorganic nanoparticles may be silica-based, alumina-based, titania-based, or zeolite-based nanoparticles. The specific surface area of ​​the porous inorganic nanoparticles is 200 to 1500 m² 2 It may be in the range of / g, and in one embodiment, 500 to 1200 m 2 It may be in the range of / g. When the specific surface area is within the above range, the loading amount of the active drug component per unit mass can be secured. The pore volume of the porous inorganic nanoparticle is 0.3 to 2.0 cm². 3It may be in the range of 3 / g, and in one embodiment, 0.5 to 1.5 cm 3 It can be in the range of 3 / g.

[0076] For example, the drug active ingredient may be supported within the pores of the porous inorganic nanoparticles by physical adsorption, capillary condensation, or hydrogen bonding with silanol groups, and the support may be performed by dispersing the porous inorganic nanoparticles in a solution in which the drug active ingredient is dissolved and then removing the solvent.

[0077] In one aspect of the present invention, the porous inorganic nanoparticle may be a mesoporous silica nanoparticle, hereinafter MSN. The MSN may be a nanoparticle having regularly arranged mesopores within a framework composed of amorphous silica. The mesopores of the MSN may function as a loading space for the drug active ingredient, and the openings of the mesopores may function as release pathways for the drug active ingredient.

[0078] In one embodiment, the MSN may have a mesopore array structure of the MCM-41 type, SBA-15 type, or MCM-48 type. The MCM-41 type may have a cylindrical mesopore structure of a two-dimensional hexagonal array, and the SBA-15 type may have a larger pore diameter and thicker pore walls compared to the MCM-41 type. Silanol groups may be distributed on the surface of the MSN, and the silanol groups may be involved in interactions with the drug active ingredient or function as reaction sites for subsequent surface modification.

[0079] For example, the surface of the MSN may be modified with an aminosilane, a thiolsilane, or a carboxysilane, and electrostatic interactions or covalent bonds with the active ingredient of the drug may be formed by the surface modification.

[0080] In one aspect of the present invention, the pore diameter of the mesoporous silica nanoparticles may be 1 to 20 nm. The pore diameter is the internal diameter of the mesopores formed within the MSN framework and can be measured by the Barrett-Joyner-Halenda (hereinafter BJH) method of nitrogen adsorption-desorption analysis.

[0081] In one embodiment, if the pore diameter is less than 1 nm, the pore is narrow compared to the molecular size of the drug active ingredient, so the loading efficiency may be reduced, and if the pore diameter exceeds 20 nm, the retention capacity of the drug active ingredient within the pore is reduced, and premature release may occur during blood circulation.

[0082] For example, the pore diameter can be controlled by the type and concentration of the template used in the synthesis process, and when cetyltrimethylammonium bromide is used as a template, a pore diameter in the range of 2 to 4 nm can be formed, and when Pluronic P123 is used as a template, a pore diameter in the range of 5 to 10 nm can be formed.

[0083] In one aspect of the present invention, the pore diameter may be 2 to 10 nm. The range may be suitable for the drug active ingredient to enter and be loaded into the mesopore, considering the molecular size of low molecular weight drugs, peptides, and small nucleic acids.

[0084] In one embodiment, when the pore diameter is in the range of 2 to 5 nm, it may be suitable for loading low molecular weight drugs with a molecular weight of 1000 Da or less, and when the pore diameter is in the range of 5 to 10 nm, it may be suitable for loading peptide drugs with a molecular weight in the range of 1000 to 10000 Da.

[0085] In one aspect of the present invention, the pore diameter may be 3 to 6 nm. The range may be a range in which a balance is secured between the molecular size of the active ingredient of a drug used to treat brain diseases and the capillary retention force inside the mesopore.

[0086] In one embodiment, when the pore diameter is in the range of 3 to 6 nm, the loading rate of the drug active ingredient can be secured in the range of 5 to 40 weight% relative to the weight of the MSN, and the early leakage rate of the drug active ingredient under blood circulation conditions can be maintained at 10% or less based on 48 hours.

[0087] In one aspect of the present invention, the particle size of the mesoporous silica nanoparticles may be 30 to 500 nm. The particle size is the outer diameter of the MSN and can be measured by dynamic light scattering or transmission electron microscopy.

[0088] In one embodiment, if the particle size is less than 30 nm, early excretion from the body by renal filtration may occur, and if the particle size exceeds 500 nm, the capture rate by the reticuloendothelial system increases, which may shorten the blood circulation time. When the particle size is in the range of 30 to 500 nm, the blood circulation time of the drug delivery system is secured, and the opportunity for contact with blood-brain barrier endothelial cells can be maintained.

[0089] In one aspect of the present invention, the particle size may be 50 to 300 nm. Within this range, the accumulation rate of the drug delivery system in the cerebral blood vessels of the disease site may be improved by enhanced permeability and retention effects.

[0090] In one embodiment, when the particle size is in the range of 50 to 300 nm, the blood half-life of the drug delivery agent can be maintained in the range of 2 to 24 hours.

[0091] In one aspect of the present invention, the particle size may be 80 to 200 nm. The range may be a size range suitable for crossing the blood-brain barrier by receptor-mediated transcytosis.

[0092] In one embodiment, when the particle size is in the range of 80 to 200 nm, compatibility between the efficiency of encapsulation into transcytosis vesicles and the size that can be accommodated within the vesicles may be possible, and the efficiency of passing through the blood-brain barrier may be improved compared to cases where the particle size is outside the range.

[0093] In one aspect of the present invention, the stimulus-responsive shell may comprise a block copolymer comprising a pH-responsive segment and a hydrophilic segment. The block copolymer may be a polymer in which two or more chemically different polymer segments are connected by covalent bonds.

[0094] In one embodiment, the pH-responsive segment may be a polymer segment whose solubility or charge state changes as it is protonated or deprotonated within a specific pH range. The pH-responsive segment may exhibit hydrophobicity under normal blood flow pH conditions and adhere to the core surface in a condensed state, and may be converted to hydrophilicity by protonation and swell under reduced pH conditions of the microenvironment surrounding the blood-brain barrier. The hydrophilic segment may form a hydration layer on the outer surface of the drug delivery vehicle during blood circulation, thereby contributing to reducing non-specific adsorption of plasma proteins and avoiding capture by the reticuloendothelial system.

[0095] For example, the pH-responsive segment may be a polymer containing a tertiary amine group in a main chain or a side chain, and the pKa value of the tertiary amine group may be in the range of 6.0 to 7.2. The hydrophilic segment may be one of polyethylene glycol, polyvinylpyrrolidone, or polyoxazoline, and the molecular weight of the hydrophilic segment may be in the range of 1000 to 10000 Da.

[0096] In one aspect of the present invention, the block copolymer may be poly(â-aminoester)-block-polyethylene glycol, hereinafter PAE-b-PEG. The PAE-b-PEG may be an amphiphilic block copolymer in which a PAE segment as the pH-responsive segment and a PEG segment as the hydrophilic segment are connected by a covalent bond.

[0097] In one embodiment, the PAE segment may include an â-aminoester repeating unit within the main chain, and the tertiary amine group within the repeating unit may be protonated by a change in pH. The number average molecular weight of the PAE segment may be in the range of 3,000 to 30,000 Da, and in one embodiment, may be in the range of 5,000 to 15,000 Da. The number average molecular weight of the PEG segment may be in the range of 1,000 to 10,000 Da, and in one embodiment, may be in the range of 2,000 to 5,000 Da.

[0098] For example, the above PAE-b-PEG can be prepared by synthesizing the PAE segment by reacting a diacrylate monomer and a diamine monomer by Michael addition polymerization, and then conjugating it with NHS-activated PEG through terminal amine groups.

[0099] In one aspect of the present invention, the pH change may correspond to the structural transition in the pH range of 6.0 to 7.0 of the microenvironment surrounding the blood-brain barrier. The pH range may be an acidification range that may occur in the microenvironment surrounding the blood-brain barrier due to neuroinflammation, brain tumor, or cerebral ischemia.

[0100] In one embodiment, the stimulus-responsive shell can maintain a shielded state under conditions greater than pH 7.0, and when the pH range of 6.0 to 7.0 is entered, a structural transition to a relaxed state can be initiated by the protonation of the pH-responsive segment. The initiation pH for the structural transition can be determined by the pKa value of the pH-responsive segment, and the pKa value can be set by adjusting the type and ratio of monomers.

[0101] In one aspect of the present invention, the pH may be in the range of 6.5 to 6.8. The range may correspond to the pH range reported in the microenvironment surrounding glioblastoma or in sites of acute neuroinflammation.

[0102] In one embodiment, when the stimulus-responsive shell is designed to undergo structural transformation in the pH range of 6.5 to 6.8, the pH difference from the normal blood flow pH of 7.35 to 7.45 can be maintained in units of 0.55 to 0.95. Due to the pH difference, non-specific activation in normal blood flow can be suppressed while selective activation at the disease site can be ensured.

[0103] For example, if a monomer combination is selected such that the pKa value of the PAE segment is in the range of 6.5 to 6.8, the pKa range can be achieved by copolymerizing 1,6-hexanediol diacrylate and 4,4'-trimethylenedipiperidine.

[0104] In one aspect of the present invention, the enzymatically degradable peptide linker may be a substrate metalloproteinase, hereinafter MMP, or a degradable peptide linker. The MMP may be a zinc-dependent endopeptidase-type protease and may have activity for degrading extracellular matrix components.

[0105] In one embodiment, the MMP-degradable peptide linker may include an amino acid sequence that is recognized and cleaved by the active site of the MMP. The MMP may be overexpressed by blood-brain barrier endothelial cells or surrounding glial cells in the neuroinflammation or brain tumor microenvironment, and may exist in an inactive precursor form or be maintained at a low concentration in a normal cerebrovascular environment. Accordingly, the MMP-degradable peptide linker may be selectively cleaved at the disease site to expose the target ligand.

[0106] In one aspect of the present invention, the MMP-degradable peptide linker may comprise a peptide sequence that is degraded by one or more of MMP-2 and MMP-9. MMP-2 may also be referred to as gelatinase A, and MMP-9 may also be referred to as gelatinase B.

[0107] In one embodiment, the MMP-2 and MMP-9 may be expressed at concentrations 2 to 20 times higher than normal brain tissue in the glioblastoma microenvironment, and the activity of the MMP-2 and MMP-9 may be increased even around the blood-brain barrier in a neuroinflammatory state.

[0108] For example, the peptide sequences degraded by the MMP-2 and MMP-9 may include Pro-Leu-Gly-Leu-Ala-Gly, Gly-Pro-Leu-Gly-Ile-Ala-Gly-Gln, or Pro-Val-Gly-Leu-Ile-Gly, and the Gly-Leu or Gly-Ile bonds within the sequence may be the cleavage sites by the MMP.

[0109] In one aspect of the present invention, the target ligand may be a transferrin receptor binding peptide. The transferrin receptor may be a transmembrane glycoprotein expressed on the luminal surface of blood-brain barrier endothelial cells and may be a receptor involved in receptor-mediated transcytosis.

[0110] In one embodiment, the transferrin receptor binding peptide may be a peptide composed of 7 to 20 amino acid residues that bind to the extracellular domain of the transferrin receptor. When the transferrin receptor binding peptide binds to the transferrin receptor, the drug delivery vehicle can cross the blood-brain barrier through a transcytosis pathway in which it is encapsulated into an endothelial cell via a clathrin-coated vesicle and then released toward brain tissue.

[0111] For example, the transferrin receptor binding peptide may include the His-Ala-Ile-Tyr-Pro-Arg-His sequence, which is a T7 peptide, or a modified sequence thereof. The T7 peptide may have a dissociation constant at the micromolar level for the transferrin receptor, and since the binding site for the endogenous transferrin and the receptor are different, competitive inhibition with endogenous transferrin may be reduced.

[0112] In one aspect of the present invention, the drug delivery vehicle may further comprise a metal-organic framework (hereinafter MOF) nanovalve layer disposed at the mesopore entrance of the mesoporous silica nanoparticle. The MOF nanovalve layer may be a layer in which a porous crystal structure formed by the coordination bonding of a metal ion and an organic ligand blocks the entrance of the mesopore.

[0113] In one embodiment, the MOF nanovalve layer can be formed on the entrance surface of the mesopore by an in situ growth method, and the crystal size of the MOF can be set larger than the diameter of the mesopore to physically seal the mesopore. The MOF nanovalve layer can maintain structural stability under normal blood flow conditions to prevent the leakage of active drug components within the mesopore to the outside. When the coordination bond of the MOF is dissociated due to a change in pH of the microenvironment around the blood-brain barrier, the MOF nanovalve layer can decompose, the mesopore can be opened, and the active drug components within the mesopore can be released.

[0114] For example, the thickness of the MOF nanovalve layer may be in the range of 5 to 50 nm, and in one embodiment, may be in the range of 10 to 30 nm. When the thickness of the MOF nanovalve layer is within the above range, a balance between the blocking efficiency under normal conditions and the degradation rate under disease conditions can be ensured.

[0115] In one aspect of the present invention, the metal-organic framework may be a zeolite imidazolate framework comprising zinc ions and 2-methylimidazole, hereinafter ZIF-8. The ZIF-8 is Zn 2 It may be a MOF having a sodalite-type crystal structure formed by the coordination bonding of an ion with a 2-methylimidazolate ligand.

[0116] In one embodiment, the diameter of the pore opening within the crystal structure of the ZIF-8 may be 0.34 nm, and the pore opening may limit the permeation of small molecules. The ZIF-8 may be structurally stable under neutral to weakly alkaline conditions, and under acidic conditions, Zn is protonated by 2-methylimidazolate. 2 The coordination bond with the ion may dissociate, causing the structure to collapse. Zn due to the decomposition of the above ZIF-8 2Ions may be emitted, and the emitted Zn 2 Ions can interact with blood-brain barrier tight junction proteins claudin-5 and occludin, contributing to the reversible relaxation of intercellular pathways.

[0117] For example, the synthesis of the ZIF-8 can be carried out by dispersing the MSN in an aqueous zinc ion solution and an aqueous 2-methylimidazole solution, and then stirring the reaction at room temperature. The zinc ions may be supplied in the form of zinc nitrate hexahydrate, zinc acetate dihydrate, or zinc chloride. The molar ratio of the 2-methylimidazole to the zinc ions may be in the range of 4:1 to 70:1, and in one embodiment, may be in the range of 8:1 to 40:1.

[0118] In one aspect of the present invention, the MOF nanovalve layer may decompose in a pH range of 6.0 to 7.0. The pH range may overlap with the pH range in which structural conversion of the stimulus-responsive shell is induced, and accordingly, the relaxation of the stimulus-responsive shell and the decomposition of the MOF nanovalve layer may proceed together by the same pH stimulus.

[0119] In one embodiment, the degradation rate of the MOF nanovalve layer may depend on pH, and at pH 7.0, almost no degradation occurs, at pH 6.5, a degradation rate of 50% or more may be shown within 6 hours, and at pH 6.0, a degradation rate of 90% or more may be shown within 2 hours.

[0120] In one aspect of the present invention, the MOF nanovalve layer may decompose in a pH range of 6.5 to 6.8. This range may correspond to the pH range observed around the blood-brain barrier in glioblastoma or acute neuroinflammation sites.

[0121] In one embodiment, when the MOF nanovalve layer is set to decompose in the pH range of 6.5 to 6.8, coordination bond dissociation may be initiated in the said pH range based on the pKa value of 2-methylimidazolate constituting the ZIF-8. Zn released upon decomposition of the MOF nanovalve layer 2 The local concentration of the ion may be in the range of 10 to 500 μM, and reversible relaxation of tight junction proteins may be induced within said concentration range. In an environment of 7.2 to 7.4, which is the normal pH of the brain parenchyma after crossing the blood-brain barrier, the Zn 2 Ions can be cleared by intracellular zinc homeostasis mechanisms, and relaxed tight junctions can be spontaneously restored.

[0122] In one aspect of the present invention, the drug delivery vehicle may further comprise a phenylboronic acid, hereinafter PBA, residue disposed on the outer surface of the stimulus-responsive shell. The PBA residue may be an organic boronic acid residue having a structure in which a phenyl group and two hydroxyl groups are bonded to a boron atom.

[0123] In one embodiment, the PBA residue may form a reversible boronate ester bond with a diol compound. Sialic acid residues may be distributed in the glycocalyx of the blood-brain barrier endothelial cells, and the sialic acid residues may include a diol array within their structure. The PBA residues may form a dynamic boronate ester bond with the diol array of the sialic acid residues, thereby allowing the drug delivery system to dock to the surface of the blood-brain barrier endothelial cells. The dynamic boronate ester bond may be a reversible covalent bond in which the bond-dissociation equilibrium can shift depending on pH, competitive diol concentration, and redox conditions.

[0124] For example, the PBA residue may be a derivative of 3-aminophenylboronic acid, 4-carboxyphenylboronic acid, or 2-fluoro-4-aminomethylphenylboronic acid attached to the outer surface of the stimulus-responsive shell by amide or carbamate bonds. The surface density of the PBA residue is 0.1 to 5.0 nmol / cm² based on the outer surface area of ​​the stimulus-responsive shell. 2 The range may be 0.5 to 2.0 nmol / cm² in one embodiment. 2 It can be a range.

[0125] In one aspect of the present invention, the drug delivery system may further include a catechol group in the stimulus-responsive shell. The catechol group may be a functional group having a 1,2-dihydroxybenzene structure in which a hydroxyl group is attached to two adjacent positions of a benzene ring.

[0126] In one embodiment, the catechol group may be introduced into the side chain or terminal of the polymer constituting the stimulus-responsive shell. The catechol group may maintain a dihydroxy state under reducing conditions and may be converted to quinone under oxidizing conditions by reactive oxygen species. The quinone may form cross-links with nucleophilic amine groups via a Michael addition reaction or a Schiff base reaction. Under conditions where the concentration of reactive oxygen species is elevated in the microenvironment surrounding the blood-brain barrier, if the catechol group is oxidatively converted to quinone, the quinone may form cross-links with amine groups of proteins within transcytosis vesicles, thereby allowing the drug delivery vehicle to ride on the transcytosis transport pathway. On the brain tissue side of the blood-brain barrier, the quinone-amine bond may be dissociated by an intracellular reducing environment, and accordingly, the drug delivery vehicle may be released toward the brain tissue.

[0127] For example, the catechol group may be in a form in which a catechol portion of dopamine, 3,4-dihydroxy-L-phenylalanine, or caffeic acid is conjugated to the polymer. The surface density of the catechol group may be in the range of 0.5 to 3.0 molar ratio relative to the surface density of the PBA residue, and in one embodiment, may be in the range of 0.8 to 1.5 molar ratio.

[0128] In one aspect of the present invention, the drug delivery system may further comprise trehalose co-encapsulated with the drug active ingredient within the mesopores of the mesoporous silica nanoparticles. The trehalose may be a non-reducing disaccharide having the chemical name á-D-glucopyranosyl á-D-glucopyranoside. The trehalose may form a glassy matrix within the mesopores.

[0129] In one embodiment, the glassy matrix may be in a state where the trehalose exists in an amorphous solid state, and molecular mobility is extremely limited below the glass transition temperature. The glass transition temperature of the trehalose may be about 115°C, and the glassy matrix may maintain a stable amorphous solid state at the biological temperature of 37°C. The glassy matrix can disperse and immobilize molecules of the drug active ingredient within the amorphous solid within the mesopores, thereby maintaining the structural stability of the drug active ingredient. Stabilization by the glassy matrix can reduce thermal denaturation and chemical degradation during blood circulation for structurally unstable biomolecular drug active ingredients such as peptides, antibodies, and siRNA.

[0130] In one embodiment, the glassy matrix may be rapidly dissolved by the inflow of an aqueous medium when the mesopore is opened. When the trehalose glassy matrix is ​​rapidly dissolved within the closed nanoscale mesopore, a local increase in osmotic pressure may occur within the mesopore. The driving force caused by the increase in osmotic pressure may actively eject the drug active ingredient toward the outside of the mesopore. Accordingly, the release rate of the drug active ingredient is improved compared to release by passive diffusion, and the penetration depth of the drug active ingredient into brain tissue may be increased.

[0131] For example, the formation of the glassy matrix may be performed by impregnating a solution in which the drug active ingredient and the trehalose are dissolved in a solvent into the mesopores, and then removing the solvent by freeze-drying or vacuum drying. The primary drying temperature during freeze-drying may be in the range of -40 to -20°C, and the secondary drying temperature may be in the range of 20 to 40°C.

[0132] In one aspect of the present invention, the content of the trehalose may be 0.5 to 20 times by weight relative to the weight of the active ingredient of the drug.

[0133] In one embodiment, if the content of the trehalose is less than 0.5 times by weight, the formation of a glassy matrix within the mesopore is incomplete, and the stabilization effect of the active drug component may be reduced. If the content of the trehalose exceeds 20 times by weight, the space for carrying the active drug component within the mesopore is excessively occupied by the trehalose, and the amount of drug loaded may be reduced.

[0134] In one aspect of the present invention, the content of the trehalose may be 2 to 10 times by weight relative to the weight of the active ingredient of the drug.

[0135] In one embodiment, within the above range, the trehalose can substantially fill the interior of the mesopore to form a continuous glassy matrix, and the loading amount of the drug active ingredient can be maintained in the range of 3 to 30 weight% relative to the weight of the MSN.

[0136] In one aspect of the present invention, the content of the trehalose may be 3 to 7 times by weight relative to the weight of the active ingredient of the drug.

[0137] In one embodiment, an optimal balance between the density of the glassy matrix and the loading amount of the drug active ingredient can be achieved within the above range. When the trehalose content is in the range of 3 to 7 weight times, the osmotic driving force generated during the rapid dissolution of the glassy matrix can be secured to a level sufficient to actively eject the drug active ingredient, and at the same time, the loading amount of the drug active ingredient can be maintained in the range of 5 to 20 weight percent relative to the weight of the MSN.

[0138] For example, if the trehalose content is less than 3 times by weight, the osmotic driving force may not show a significant difference compared to release by passive diffusion, and if it exceeds 7 times by weight, securing a therapeutic effective dose may be limited due to a decrease in the loading amount of the active drug component.

[0139] In one aspect of the present invention, the drug delivery system may further comprise a catalyst layer containing Prussian blue nanoparticles on the outermost layer of the stimulus-responsive shell. The Prussian blue is Fe 3 4[Fe(CN) 3 6] 3 It may be a transition metal cyanide coordination polymer having the chemical formula of 3. The Prussian blue nanoparticles may be nanozymes having one or more multienzyme-mimicking activities among catalase-like activity, peroxidase-like activity, and superoxide dismutase-like activity.

[0140] In one embodiment, the catalyst layer may be a layer formed by distributing the Prussian blue nanoparticles on the outermost surface of the stimulus-responsive shell. The catalyst layer can catalytically decompose and eliminate reactive oxygen species surrounding the drug delivery vehicle. In normal blood flow environments and mild inflammatory environments, the reactive oxygen species scavenging capacity of the catalyst layer may exceed the surrounding reactive oxygen species concentration, so the reactive oxygen species reactive configuration inside the drug delivery vehicle may not be activated. If the reactive oxygen species concentration around the blood-brain barrier of the disease site exceeds the scavenging capacity of the catalyst layer, excess reactive oxygen species may penetrate the catalyst layer and reach the reactive oxygen species reactive configuration inside. Accordingly, the catalyst layer can perform a threshold gating function that converts a continuous reactive oxygen species concentration gradient into a binary on / off response.

[0141] In one embodiment, the Prussian blue nanoparticles can reduce oxidative stress by continuously scavenging residual reactive oxygen species in the brain parenchyma even after the drug delivery vehicle has passed through the blood-brain barrier.

[0142] For example, the particle size of the above Prussian blue nanoparticles may be in the range of 5 to 50 nm, and in one embodiment, may be in the range of 10 to 30 nm. The above Prussian blue nanoparticles may be made of the same material as the active ingredient of Radiogardase®, an FDA-approved drug, and may be a material whose biosafety has been confirmed.

[0143] In one aspect of the present invention, the content of the Prussian blue nanoparticles may be 0.5 to 15 weight percent relative to the total weight of the drug delivery system.

[0144] In one embodiment, if the content of the Prussian blue nanoparticles is less than 0.5 weight%, the scavenging capacity of the catalyst layer is insufficient, so the threshold gating function may not be expressed. If the content exceeds 15 weight%, the thickness of the catalyst layer is excessive, which increases the overall particle size of the drug delivery system and may physically hinder the structural transformation of the stimulus-responsive shell.

[0145] In one aspect of the present invention, the content of the Prussian blue nanoparticles may be 2 to 8 weight percent relative to the total weight of the drug delivery system.

[0146] In one embodiment, the scavenging capacity of the catalyst layer in the above range may be sufficient to scavenge reactive oxygen species in the range of 0.1 to 1.0 µM, which is the concentration of reactive oxygen species in a normal blood flow environment, and in the range of 1.0 to 10 µM, which is the concentration of reactive oxygen species in a mild inflammation environment, and in the range of 50 to 500 µM, which is the concentration of reactive oxygen species in a severe disease environment, the scavenging capacity may be saturated and excess reactive oxygen species may permeate inward. Accordingly, selectivity for disease severity may be ensured in the above content range.

[0147] For example, if the content of the above Prussian blue nanoparticles is less than 2 weight%, reactive oxygen species exceeding the threshold may reach the inside even in a mild inflammatory environment and non-specific activation may occur, and if it exceeds 8 weight%, the time to exceed the threshold may be excessively delayed even in a severe disease environment, which may reduce the timeliness of drug delivery.

[0148] FIG. 1 is a flowchart illustrating a method for manufacturing a drug delivery system capable of crossing the blood-brain barrier according to an embodiment of the present invention. Referring to FIG. 1, the manufacturing method may include a core formation step (S10), a drug encapsulation step (S20), a shell formation step (S30), and a surface functionalization step (S40). In the core formation step (S10), a core may be formed. In the drug encapsulation step (S20), a drug active ingredient may be encapsulated in the core. In the shell formation step (S30), a stimulus-responsive shell may be formed on the outer surface of the core containing the drug active ingredient. In the surface functionalization step (S40), a reactive binding structure may be formed on the outer surface of the stimulus-responsive shell. Each of the above steps may be performed sequentially according to a temporal order.

[0149] In one aspect of the present invention, a method for manufacturing a drug delivery system capable of crossing the blood-brain barrier may include a core forming step (S10), a drug encapsulation step (S20), a shell forming step (S30), and a surface functionalization step (S40).

[0150] In one embodiment, referring to FIG. 1, in the core forming step (S10), a core having an internal space capable of accommodating the drug active ingredient may be formed. The formation of the core may be performed by one of a synthesis method, a self-assembly method, or an emulsion method. In the drug encapsulation step (S20), the drug active ingredient may be encapsulated in the internal space of the core formed in the core forming step (S10). The encapsulation may be performed by one of a solution impregnation method, a vacuum inhalation method, or an ultrasonic dispersion method.

[0151] In one embodiment, in the shell forming step (S30), a stimulus-responsive shell may be formed on the outer side of the core containing the drug active ingredient. The stimulus-responsive shell may be formed of a polymer material that undergoes a structural transition from a shielded state to a relaxed state upon stimulation of the microenvironment surrounding the blood-brain barrier. The formation of the shell may be performed by one of the solvent evaporation method, the electrostatic adsorption method, or the grafting method.

[0152] In one embodiment, in the surface functionalization step (S40), a reactive binding structure including a masking linker and a target ligand may be formed on the outer surface of the stimulus-responsive shell. The formation of the reactive binding structure may be performed by covalently bonding one end of the masking linker to the outer surface of the stimulus-responsive shell and covalently bonding the target ligand to the other end of the masking linker. One of NHS-amine coupling, maleimide-thiol coupling, or click chemistry may be used for the bonding.

[0153] For example, the entire manufacturing process from the core formation step (S10) to the surface functionalization step (S40) may be performed under aqueous or organic solvent conditions, and a purification process by centrifugation, dialysis, or filtration may be interposed between each step.

[0154] In one aspect of the present invention, in the shell forming step (S30), the stimulus-responsive shell may comprise a block copolymer including a pH-responsive segment and a hydrophilic segment.

[0155] In one embodiment, the block copolymer may be coated on the outer surface of the core while dissolved in an organic solvent or an aqueous solvent. The coating may be performed by dispersing the core containing the drug active ingredient in the block copolymer solution, and then inducing self-assembly of the block copolymer by removing the solvent or adjusting the pH.

[0156] For example, the concentration of the block copolymer solution may be in the range of 0.1 to 50 mg / mL, and in one embodiment, may be in the range of 1 to 20 mg / mL. The temperature during coating may be in the range of 4 to 40℃, and the stirring time may be in the range of 1 to 24 hours.

[0157] In one aspect of the present invention, in the surface functionalization step (S40), the masking linker may be an enzymatically degradable peptide linker.

[0158] In one embodiment, the enzymatically degradable peptide linker may be synthesized in advance by a solid-phase peptide synthesis method and then conjugated to the outer surface of the stimulus-responsive shell in the surface functionalization step (S40). A reactive functional group for conjugation with the outer surface of the stimulus-responsive shell may be introduced at the N-terminus or C-terminus of the enzymatically degradable peptide linker, and the reactive functional group may be one of an amine group, a thiol group, an azide group, or an alkyne group.

[0159] In one aspect of the present invention, the core forming step (S10) may include a step of synthesizing mesoporous silica nanoparticles, hereinafter MSN, by a sol-gel method.

[0160] In one embodiment, the sol-gel method may be a synthesis method in which a silica precursor is hydrolyzed and condensed to form an amorphous silica framework, and a template is used to form regular mesopores. The silica precursor may be tetraethoxysilane, hereinafter TEOS, tetramethoxysilane, hereinafter TMOS, or a mixture thereof. The template may be a cationic surfactant, a nonionic block copolymer surfactant, or a mixture thereof.

[0161] For example, the synthesis of the MSN can be carried out by dissolving the template in an alkaline aqueous solution, adding the silica precursor dropwise, and stirring the reaction at a temperature range of 40 to 80°C for 2 to 48 hours. After the reaction, the template can be removed by organic solvent extraction or calcination at a temperature range of 400 to 600°C.

[0162] In one aspect of the present invention, the shell forming step (S30) may include the step of coating the outer surface of the core with poly(â-aminoester)-block-polyethylene glycol, hereinafter PAE-b-PEG.

[0163] In one embodiment, the coating may be performed by dissolving the PAE-b-PEG in a buffer solution in the pH range of 7.0 to 8.0, dispersing the MSN containing the drug active ingredient, and inducing the PAE-b-PEG to self-assemble on the surface of the MSN by hydrophobic interactions of the PAE segment.

[0164] For example, the amount of PAE-b-PEG added may be in the range of 0.5 to 5.0 times by weight relative to the weight of MSN, and in one embodiment, may be in the range of 1.0 to 3.0 times by weight. The coating reaction may be carried out by stirring for 2 to 12 hours at a temperature in the range of 4 to 25°C.

[0165] In one aspect of the present invention, the manufacturing method may further include a nanovalve layer forming step prior to the shell forming step (S30), wherein a nanovalve layer comprising a zeolite imidazolate framework, hereinafter ZIF-8, is formed by reacting zinc ions and 2-methylimidazole at the mesopore entrances of the mesoporous silica nanoparticles.

[0166] In one embodiment, the nanovalve layer formation step may be performed after the drug encapsulation step (S20) and before the shell formation step (S30). In the nanovalve layer formation step, the MSN containing the drug active ingredient may be dispersed in an aqueous zinc ion solution, and subsequently, an aqueous 2-methylimidazole solution may be added to induce in situ crystal growth of ZIF-8 on the surface of the mesopore entrance of the MSN. By the in situ crystal growth, the mesopore entrance may be blocked by the ZIF-8 crystal, thereby forming a nanovalve layer.

[0167] For example, the concentration of the zinc ions may be in the range of 5 to 100 mM, and the concentration of the 2-methylimidazole may be in the range of 50 to 1000 mM. The in situ crystal growth reaction may be carried out at room temperature for 10 minutes to 6 hours, and in one embodiment, for 30 minutes to 2 hours.

[0168] In one aspect of the present invention, the surface functionalization step (S40) may include the step of introducing a phenylboronic acid, hereinafter PBA, residue to the outer surface of the stimulus-responsive shell, and the step of introducing a catechol group to the stimulus-responsive shell.

[0169] In one embodiment, in the step of introducing the PBA residue, a PBA derivative having an amino group or a carboxyl group may be conjugated to a reactive functional group present on the outer surface of the stimulus-responsive shell by an amide bond or an ester bond. In the step of introducing the catechol group, a catechol-containing compound may be conjugated to a side chain or terminal of the polymer constituting the stimulus-responsive shell by a covalent bond.

[0170] For example, the introduction of the PBA residue and the introduction of the catechol group may be performed sequentially, or may be performed in a single step using a bifunctional compound containing the PBA residue and the catechol group, respectively. The conjugation reaction may be performed for 1 to 12 hours in a buffer solution with a pH in the range of 6.5 to 8.0.

[0171] In one aspect of the present invention, the drug encapsulation step (S20) may include the step of co-encapsulating the drug active ingredient and trehalose within the mesopores of the mesoporous silica nanoparticles. The trehalose may be a non-reducing disaccharide having the chemical name á-D-glucopyranosyl á-D-glucopyranoside.

[0172] In one embodiment, the encapsulation may be performed by preparing a solution in which the drug active ingredient and the trehalose are co-dissolved in an aqueous solution or a buffer solution, dispersing the MSN in the solution, and then removing residual gas inside the mesopore under reduced pressure or vacuum conditions to induce the solution to be impregnated into the mesopore.

[0173] For example, the concentration of the active ingredient of the drug at the time of the above-mentioned empty encapsulation may be in the range of 0.1 to 50 mg / mL, and the concentration of the trehalose may be in the range of 1 to 200 mg / mL. The impregnation may be performed for 2 to 24 hours at a temperature in the range of 4 to 25℃.

[0174] In one aspect of the present invention, the trehalose may be co-encapsulated at a weight ratio of 0.5 to 20 times the weight of the drug active ingredient.

[0175] In one embodiment, the weight ratio of the trehalose and the drug active ingredient can be controlled by the ratio of the dissolution concentration of the trehalose and the drug active ingredient in the solution used for the empty encapsulation.

[0176] In one aspect of the present invention, the manufacturing method may further include a vitrification step after the drug encapsulation step (S20), wherein a vitrified matrix of the trehalose is formed within the mesopore by freeze-drying or vacuum drying.

[0177] In one embodiment, the vitrification step may be a step of introducing the MSN, which has been sealed, into a freeze dryer or a vacuum dryer to remove the solvent remaining in the mesopore, thereby inducing the trehalose to form a glassy matrix in an amorphous solid state. The vitrification step may be performed prior to the nanovalve layer formation step, so that the nanovalve layer can block the mesopore entrance after the glassy matrix is ​​formed within the mesopore.

[0178] For example, the freeze-drying may be carried out by freezing the MSN dispersion in the range of -80 to -40°C, then performing primary drying for 12 to 48 hours under reduced pressure conditions in the range of 0.01 to 1.0 mbar, and subsequently performing secondary drying for 6 to 24 hours at the range of 20 to 40°C. The reduced pressure drying may be carried out for 12 to 72 hours at a temperature in the range of 40 to 60°C under reduced pressure conditions in the range of 1 to 100 mbar.

[0179] In one aspect of the present invention, the manufacturing method may further include a catalyst layer forming step, wherein, after the surface functionalization step (S40), a catalyst layer comprising Prussian blue nanoparticles is formed on the outermost surface of the stimulus-responsive shell.

[0180] In one embodiment, in the catalyst layer formation step, the Prussian blue nanoparticles may be provided in the form of a pre-synthesized nanoparticle dispersion, and the surface-functionalized drug delivery vehicle may be dispersed in the Prussian blue nanoparticle dispersion to induce the Prussian blue nanoparticles to attach to the outermost surface of the stimulus-responsive shell by electrostatic adsorption or covalent bonding. The surface of the Prussian blue nanoparticles may be coated with one or more dispersion stabilizers selected from citric acid, polyvinylpyrrolidone, or polyethylene glycol.

[0181] For example, the above Prussian blue nanoparticles are K 3 3[Fe(CN) 3 6] Aqueous solution and FeCl 3 3. It can be synthesized by the co-precipitation method by mixing aqueous solutions. K during the above synthesis 3 3[Fe(CN) 3 6] The concentration may be in the range of 1 to 50 mM, and the FeCl 3 3. The concentration may be in the range of 1 to 50 mM. The stirring time in the catalyst layer formation step may be in the range of 1 to 12 hours, and the temperature may be in the range of 4 to 25℃.

[0182] In one aspect of the present invention, a pharmaceutical composition for treating brain diseases may comprise the drug delivery system and a pharmaceutically acceptable carrier. The active drug component encapsulated in the drug delivery system may be an active substance to be delivered to brain tissue for the treatment of brain diseases, and may be one or more of an anticancer agent, a neuroprotective agent, a treatment for neurodegenerative diseases, an anti-inflammatory agent, and a neurotransmitter modulator.

[0183] In one embodiment, the pharmaceutically acceptable carrier may be one or more of an isotonic aqueous solution, phosphate-buffered saline, Ringer's solution, or an aqueous dextrose solution. The pharmaceutical composition may be administered by intravenous injection, arterial injection, or intraventricular injection, and in one embodiment, may be administered by intravenous injection. The concentration of the drug delivery agent in the pharmaceutical composition may be in the range of 0.1 to 100 mg / mL, and in one embodiment, may be in the range of 1 to 50 mg / mL.

[0184] For example, the brain disease may be one or more of glioblastoma, brain metastatic cancer, Alzheimer's disease, Parkinson's disease, cerebral ischemia, cerebral hemorrhage, multiple sclerosis, and encephalitis. The anticancer agent may be one or more of doxorubicin, temozolomide, paclitaxel, or cisplatin, and the neuroprotective agent may be one or more of edaravone, memantine, or riluzole.

[0185] In one aspect of the present invention, a composition for improving skin condition through brain-skin axis regulation may comprise the drug delivery system and a pharmaceutically acceptable carrier, and the drug active ingredient may be a neuromodulator having brain-skin axis regulatory activity.

[0186] In one embodiment, the brain and skin originate from the same embryological ectoderm, and neuropeptides and neurotransmitters secreted from the brain may be involved in the proliferation, differentiation, melanin synthesis, collagen metabolism, and inflammatory response of skin cells via the neuro-endocrine-immune pathway. Through this brain-skin axis, neuromodulatory substances delivered to the brain may contribute to the condition of the skin. The neuromodulatory substances may be one or more of α-melanocyte-stimulating hormone, adrenocorticotropic hormone-releasing hormone regulatory peptide, brain-derived neurotrophic factor, oxytocin, or melatonin.

[0187] In one embodiment, the composition may be administered by intravenous injection, and the neuromodulator delivered to the brain by passing through the blood-brain barrier via the drug delivery system may contribute to skin anti-aging, skin barrier strengthening, regulation of melanin synthesis, or reduction of skin inflammation through the brain-skin axis pathway. The improvement of skin condition may be related to stress-induced skin aging, neurodermatitis, pigmentation caused by neuropeptide imbalance, or deterioration of skin barrier function caused by neuro-immune pathway disruption.

[0188] For example, the brain-derived neurotrophic factor mentioned above is involved in neuronal survival and plasticity in the brain, while simultaneously promoting the differentiation of skin keratinocytes via the brain-skin axis and contributing to maintaining homeostasis of skin nerve endings. The melatonin mentioned above is secreted from the pineal gland in the brain and reaches the skin through systemic circulation, and can contribute to antioxidant activity and the regulation of melanin synthesis.

[0189] In one aspect of the present invention, a health functional food composition for improving brain health may comprise the drug delivery system, a food-grade acceptable carrier, and an enteric coating layer formed on the outer surface of the drug delivery system. The drug active ingredient may be a functional ingredient for brain health.

[0190] In one embodiment, the enteric coating layer may be a layer formed of an enteric polymer that does not dissolve in the pH range of 1 to 3, which is gastric acid condition, but dissolves at pH 5.5 or higher, which is intestinal condition. The enteric polymer may be one or more of hydroxypropylmethylcellulose phthalate, cellulose acetate phthalate, Oidrazit L100, or Oidrazit S100. Due to the enteric coating layer, the drug delivery vehicle is not degraded in the gastrointestinal environment and can be absorbed into the bloodstream through the intestinal epithelium after the enteric coating layer dissolves in the small intestine, and subsequently reaches the blood-brain barrier through blood circulation, thereby allowing the brain health functional ingredient to be delivered to brain tissue by the conditional passage mechanism of the present invention.

[0191] In one embodiment, the brain health functional ingredient may be one or more of curcumin, resveratrol, phosphatidylserine, omega-3 fatty acids, ginkgolide, or ergothioneine. The food-grade acceptable carrier may be one or more of gelatin capsules, hydroxypropylmethylcellulose capsules, or excipients for tablets.

[0192] For example, the above health functional food composition may be provided in the form of a hard capsule or a soft capsule, and the content of the drug delivery agent per single intake may be in the range of 10 to 500 mg. The thickness of the enteric coating layer may be in the range of 10 to 100 ìm, and in one embodiment may be in the range of 20 to 50 ìm.

[0193] Example 1

[0194] Mesoporous silica nanoparticles were synthesized by the sol-gel method. 1.0 g of cetyltrimethylammonium bromide was dissolved in 480 mL of distilled water, the pH was adjusted to 11.0 using an aqueous sodium hydroxide solution, and the temperature was raised to 80°C. 5.0 mL of tetraethoxysilane was added dropwise at a rate of 0.25 mL / min, and the reaction was stirred at 80°C for 2 hours. The reaction product was recovered by centrifugation, and the template was removed by repeating reflux extraction twice in a mixed solution of ethanol and hydrochloric acid. The pore diameter of the obtained mesoporous silica nanoparticles was confirmed to be 4.5 nm by the BJH method of nitrogen adsorption-desorption analysis; the particle size determined by dynamic light scattering was 150 nm, and the BET specific surface area was 980 m². 2 / g was.

[0195] 100 mg of the above mesoporous silica nanoparticles were dispersed in 10 mL of phosphate-buffered saline solution containing 10 mg of doxorubicin hydrochloride, and the drug was loaded by stirring at 4°C for 12 hours. Unloaded drug was removed by centrifugation, and the drug loading rate was measured by an ultraviolet-visible spectrophotometer and was found to be 8.2 wt% relative to the weight of the mesoporous silica nanoparticles.

[0196] Poly(â-aminoester)-block-polyethylene glycol was synthesized by Michael addition polymerization. Poly(â-aminoester) segments were obtained by reacting 1,6-hexanediol diacrylate and 4,4'-trimethylenedipipiridine in a 1.2:1 molar ratio in dimethylformamide at 50°C for 24 hours, and amine-terminated polyethylene glycol was conjugated to the terminal acrylate groups. The number average molecular weight of the poly(â-aminoester) segments of the obtained block copolymer was 8500 Da, and the number average molecular weight of the polyethylene glycol segments was 3000 Da.

[0197] 200 mg of the above block copolymer was dissolved in 20 mL of pH 7.4 phosphate-buffered saline, then 100 mg of the above drug-loaded mesoporous silica nanoparticles were dispersed, and a stimulus-responsive shell was formed by stirring at 4°C for 6 hours.

[0198] A peptide with the sequence Gly-Pro-Leu-Gly-Ile-Ala-Gly-Gln, prepared by a solid-phase peptide synthesis method as a substrate metalloproteinase-degradable peptide linker, was conjugated to the amine groups on the outer surface of the stimulus-responsive shell using an NHS-maleimide heterofunctional crosslinking agent, and a T7 peptide sequence His-Ala-Ile-Tyr-Pro-Arg-His was conjugated to the C-terminus of the peptide linker to form a reactive bond structure. The reaction was carried out for 4 hours at room temperature in phosphate-buffered saline at pH 7.2, and unreacted materials were removed by dialysis.

[0199] Example 2

[0200] In Example 1, a nanovalve layer was additionally formed after the completion of drug loading and prior to shell formation. 100 mg of drug-loaded mesoporous silica nanoparticles were dispersed in 20 mL of a 25 mM aqueous solution of zinc nitrate hexahydrate, then 20 mL of a 500 mM aqueous solution of 2-methylimidazole was added dropwise and stirred at room temperature for 1 hour to grow a zeolite imidazolate framework in situ at the mesopore entrance. Unreacted materials were removed by centrifugation. Subsequent shell formation and surface functionalization were performed in the same manner as in Example 1.

[0201] Example 3

[0202] After the surface functionalization of Example 2 was completed, the introduction of phenylboronic acid residues and catechol groups was further performed. 15 mg of 3-aminophenylboronic acid was reacted for 4 hours with the carboxyl groups on the outer surface of the stimulus-responsive shell in 10 mL of pH 7.0 2-(N-morpholino)ethanesulfonic acid buffer solution in the presence of 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide and N-hydroxysuccinimide to conjugate phenylboronic acid residues. Subsequently, 10 mg of dopamine hydrochloride was reacted for 3 hours with the polyethylene glycol segment ends of the stimulus-responsive shell in 10 mL of pH 8.0 Tris buffer solution to introduce catechol groups. After each reaction, unreacted materials were removed by dialysis.

[0203] Examples 4-1 to 4-5

[0204] In the manufacturing method of Example 3, trehalose co-encapsulation and vitrification were additionally performed during the drug encapsulation step, and five types of drug delivery systems with varying trehalose content were prepared. Specifically, 100 mg of mesoporous silica nanoparticles were dispersed in a solution in which doxorubicin hydrochloride and trehalose were co-dissolved in phosphate-buffered saline, and after removing residual gas inside the mesopores by treating under reduced pressure conditions of 0.05 mbar for 30 minutes, co-encapsulation was performed by stirring at 4°C for 12 hours. After the completion of co-encapsulation, the mesoporous silica nanoparticles were recovered by centrifugation, frozen at -60°C, and then subjected to primary drying at 0.05 mbar for 24 hours and secondary drying at 30°C for 12 hours to form a vitrified matrix. The subsequent formation of the nanovalve layer, shell formation, surface functionalization, and introduction of phenylboronic acid residues and catechol groups were performed in the same manner as in Example 3.

[0205] The trehalose content of Examples 4-1 to 4-5 was set as follows. For Example 4-1, it was 0.5 times by weight relative to the weight of the active ingredient, for Example 4-2, 2 times by weight, for Example 4-3, 5 times by weight, for Example 4-4, 7 times by weight, and for Example 4-5, 10 times by weight.

[0206] Examples 5-1 to 5-5

[0207] Based on the drug delivery system prepared in Example 4-3, a catalyst layer containing Prussian blue nanoparticles was additionally formed, and five types of drug delivery systems with varying Prussian blue nanoparticle content were prepared. First, K 33 [Fe(CN) 36 ] 10 mM aqueous solution and FeCl 33 A Prussian blue nanoparticle dispersion was synthesized by mixing equal volumes of 10 mM aqueous solutions and stirring at room temperature for 30 minutes in the presence of polyvinylpyrrolidone. The particle size of the Prussian blue nanoparticles, measured by dynamic light scattering, was 18 nm. The drug delivery system of Example 4-3 was dispersed in the Prussian blue nanoparticle dispersion and stirred at 4°C for 6 hours to form a catalyst layer on the outermost surface of the stimulus-responsive shell. Unattached Prussian blue nanoparticles were removed by centrifugation.

[0208] The content of Prussian blue nanoparticles in Examples 5-1 to 5-5 was set as follows relative to the total weight of the drug delivery system by adjusting the amount of the Prussian blue nanoparticle dispersion added. Example 5-1 was 0.5 wt%, Example 5-2 was 2 wt%, Example 5-3 was 5 wt%, Example 5-4 was 8 wt%, and Example 5-5 was 15 wt%.

[0209] Examples 6-1 to 6-3

[0210] In the preparation method of Example 1, three types of drug delivery systems with varying pore diameters were prepared by changing the synthesis conditions of the mesoporous silica nanoparticles. The particle size was maintained at 150 nm for all, and the pore diameter was controlled by changing the type and concentration of the template. In Example 6-1, mesoporous silica nanoparticles with a pore diameter of 2 nm were synthesized using 0.5 g of cetyltrimethylammonium bromide. In Example 6-2, mesoporous silica nanoparticles with a pore diameter of 4.5 nm were synthesized under the same conditions as in Example 1. In Example 6-3, mesoporous silica nanoparticles with a pore diameter of 10 nm were synthesized by using 2.0 g of Pluronic P123 as a template and reacting in 60 mL of a 2 M aqueous hydrochloric acid solution at 40°C. Subsequent drug loading, shell formation, and surface functionalization were performed in the same manner as in Example 1.

[0211] Examples 7-1 to 7-3

[0212] In the preparation method of Example 1, three types of drug delivery systems with varying particle sizes were prepared by changing the synthesis conditions of the mesoporous silica nanoparticles. The pore diameter was maintained at 4.5 nm for all, and the particle size was controlled by changing the drop rate and stirring speed of the silica precursor. In Example 7-1, mesoporous silica nanoparticles with a particle size of 50 nm were obtained by increasing the tetraethoxysilane drop rate to 0.5 mL / min and setting the stirring speed to 900 rpm. In Example 7-2, mesoporous silica nanoparticles with a particle size of 150 nm were obtained under the same conditions as in Example 1. In Example 7-3, mesoporous silica nanoparticles with a particle size of 300 nm were obtained by lowering the tetraethoxysilane drop rate to 0.1 mL / min and setting the stirring speed to 300 rpm. Subsequent drug loading, shell formation, and surface functionalization were performed in the same manner as in Example 1.

[0213] Example 8

[0214] As the drug delivery system of Example 5-3, the entire composition is integrated. Specifically, mesoporous silica nanoparticles with a pore diameter of 4.5 nm and a particle size of 150 nm were used as a core, and doxorubicin hydrochloride and trehalose were co-encapsulated at 5 times the weight of the drug active ingredient, followed by vitrification, a zeolite imidazolate framework nanovalve layer was formed, a poly(â-aminoester)-block-polyethylene glycol shell was coated, a substrate metalloproteinase-degradable peptide linker and a T7 peptide were conjugated, phenylboronic acid residues and catechol groups were introduced, and a Prussian blue nanoparticle catalyst layer was formed at 5 wt% of the total weight of the drug delivery system. The specific conditions for each step are as described in Examples 1 to 5-3.

[0215] Example 9

[0216] 50 mg of the drug delivery system prepared in Example 8 was dispersed in 10 mL of physiological saline to prepare a pharmaceutical composition for treating brain diseases with a drug delivery system concentration of 5 mg / mL. The active drug component encapsulated in the drug delivery system was doxorubicin hydrochloride.

[0217] Example 10

[0218] In the preparation method of Example 1, the active ingredient was changed from doxorubicin hydrochloride to recombinant brain-derived neurotrophic factor protein, and subsequently, in the same manner as in Example 4-3, 5 times by weight of trehalose was encapsulated, vitrified, nanovalve layer formed, shell formed, surface functionalization, introduction of phenylboronic acid residues and catechol groups, and Prussian blue catalyst layer formed to prepare a fully integrated drug delivery system. 50 mg of the above drug delivery system was dispersed in 10 mL of physiological saline to prepare a composition for improving skin condition through brain-skin axis regulation.

[0219] Example 11

[0220] In the manufacturing method of Example 1, the active ingredient was changed from doxorubicin hydrochloride to curcumin, and subsequently, in the same manner as in Example 4-3, 5 times by weight of trehalose was encapsulated, vitrification, nanovalve layer formation, shell formation, surface functionalization, introduction of phenylboronic acid residues and catechol groups, and Prussian blue catalyst layer formation were performed to prepare a fully integrated drug delivery system. After dispersing the drug delivery system in an aqueous solution of Oidrazit L100, an enteric coating layer with a thickness of 30 µm was formed on the outer surface of the drug delivery system by a fluid bed coating method. After the completion of the enteric coating, 200 mg of the drug delivery system was filled into hydroxypropylmethylcellulose capsules to prepare a health functional food composition for improving brain health.

[0221] Comparative Example 1

[0222] In the manufacturing method of Example 1, only the synthesis of mesoporous silica nanoparticles and drug loading were performed, and shell formation, surface functionalization, and all subsequent additional processes were not performed. That is, the mesoporous silica nanoparticles loaded with drug alone were used as the sample of Comparative Example 1.

[0223] Comparative Example 2

[0224] In the preparation method of Example 1, a polyethylene glycol-only shell was applied instead of a stimulus-responsive shell. A shell was formed by grafting methoxypolyethylene glycol-silane with a molecular weight of 5000 Da onto the surface of mesoporous silica nanoparticles. Subsequent surface functionalization was performed in the same manner as in Example 1. The polyethylene glycol-only shell is a non-reactive shell that does not undergo structural transformation due to changes in pH.

[0225] Comparative Example 3

[0226] In the manufacturing method of Example 1, the T7 peptide was directly conjugated to the outer surface of the stimulus-responsive shell without using a masking linker. That is, a drug delivery system was manufactured in which the target ligand is constantly exposed.

[0227] Comparative Examples 4-1 and 4-2

[0228] The same manufacturing method as in Examples 4-1 to 4-5 was followed, but the trehalose content was set outside the numerical range. Comparative Example 4-1 was set to 0.3 times by weight relative to the weight of the active ingredient, and Comparative Example 4-2 was set to 25 times by weight.

[0229] Comparative Examples 5-1 and 5-2

[0230] The same manufacturing method as in Examples 5-1 to 5-5 was followed, but the Prussian blue nanoparticle content was set outside the numerical range. Comparative Example 5-1 was 0.2 wt% relative to the total weight of the drug delivery system, and Comparative Example 5-2 was 20 wt%.

[0231] Comparative Examples 6-1 and 6-2

[0232] The same manufacturing method as in Examples 6-1 to 6-3 was followed, but the pore diameter was set outside the numerical range. Comparative Example 6-1 used microporous silica with a pore diameter of 0.8 nm as non-porous silica nanoparticles synthesized by the Stöber method without using a template. Comparative Example 6-2 synthesized mesoporous silica nanoparticles with a pore diameter of 25 nm using polystyrene-block-polyethylene oxide as a template along with a swelling agent. The particle size was maintained at 150 nm for all cases.

[0233] Comparative Examples 7-1 and 7-2

[0234] The same manufacturing method as in Examples 7-1 to 7-3 was followed, but the particle size was set outside the numerical range. In Comparative Example 7-1, the tetraethoxysilane dropwise addition rate was increased to 1.0 mL / min and the stirring speed was set to 1200 rpm to obtain mesoporous silica nanoparticles with a particle size of 15 nm. In Comparative Example 7-2, the tetraethoxysilane dropwise addition rate was lowered to 0.05 mL / min and the stirring speed was set to 200 rpm to obtain mesoporous silica nanoparticles with a particle size of 600 nm. The pore diameter was maintained at 4.5 nm in all cases.

[0235] Comparative Example 8

[0236] The recombinant brain-derived neurotrophic factor protein was not encapsulated in the drug delivery system of the present invention, and an aqueous solution dissolved in physiological saline at the same concentration was used as the sample of Comparative Example 8.

[0237] Comparative Example 9

[0238] Curcumin powder without encapsulating curcumin in the drug delivery system of the present invention and without applying an enteric coating was filled into hydroxypropylmethylcellulose capsules in equal weight, and this was used as the sample of Comparative Example 9.

[0239] Comparative Example 10

[0240] A drug delivery system containing curcumin was prepared in the same manner as in Example 11, but without forming an enteric coating layer, and the drug delivery system was filled into a hydroxypropylmethylcellulose capsule in the same weight as in Example 10, and this was used as the sample for Comparative Example 10.

[0241] Experimental Example 1: Analysis of Physicochemical Properties

[0242] The physicochemical properties of the drug delivery systems prepared in Examples 1, 2, 3, and 8 were analyzed. The hydrodynamic diameters measured by dynamic light scattering were 172 nm for Example 1, 178 nm for Example 2, 183 nm for Example 3, and 198 nm for Example 8. Transmission electron microscopy confirmed the regular mesopore structure of the mesoporous silica nanoparticle cores, and a nanovalve layer formed at the mesopore inlets was observed in samples from Example 2 onwards. The BET specific surface area determined by nitrogen adsorption-desorption analysis was 410 m² for Example 1. 2 / g, Example 2 was 285 m 2 / g, and the decrease in specific surface area in Example 2 is attributed to the blocking of mesopores by the zeolite imidazolate framework. The zeta potential under pH 7.4 conditions was -12.3 mV for Example 1, -8.7 mV for Example 3, and -6.2 mV for Example 8.

[0243] Experimental Example 2: Evaluation of pH Reactivity

[0244] The structural conversion behavior under pH conditions was evaluated for Examples 1, 2, 3, Comparative Example 1, and Comparative Example 2. Each sample was dispersed in phosphate-buffered saline adjusted to pH 7.4, 6.8, 6.5, and 6.0, and the rate of change in particle size was measured by dynamic light scattering after 2 hours at 37°C.

[0245] The particle size change rate of Example 1 was 1.2% at pH 7.4, 23.5% at pH 6.8, 41.8% at pH 6.5, and 52.3% at pH 6.0. The particle size change rate of Example 2 was 1.5% at pH 7.4, 28.1% at pH 6.8, 48.7% at pH 6.5, and 61.2% at pH 6.0. The particle size change rate of Example 3 was 1.8% at pH 7.4, 30.2% at pH 6.8, 51.3% at pH 6.5, and 64.5% at pH 6.0. Comparative Example 1 had a particle size change rate of within 2% under all pH conditions. Comparative Example 2 had a particle size change rate of within 3% under all pH conditions.

[0246] For Example 2, the decomposition rate of the zeolite imidazolate framework was further measured. The sample was dispersed in phosphate-buffered saline at each pH condition, and the decomposition rate was calculated by measuring the zinc ion concentration in the supernatant at different times using inductively coupled plasma mass spectrometry. The decomposition rate after 6 hours at pH 7.4 was 3.1%, after 6 hours at pH 6.8 was 48.7%, after 6 hours at pH 6.5 was 82.3%, and after 6 hours at pH 6.0 was 95.6%.

[0247] Experimental Example 3: Drug Loading Characteristics According to Variation in Pore Diameter and Particle Size

[0248] Drug loading rate and early leakage rate were measured for Examples 6-1 to 6-3 and Comparative Examples 6-1 to 6-2. Drug loading rate was determined by ultraviolet-visible spectrophotometer, and early leakage rate was calculated as the amount of drug released after 48 hours of dialysis at 37°C in pH 7.4 phosphate-buffered saline.

[0249] The drug loading rate of Comparative Example 6-1, with a pore diameter of 0.8 nm, was 0.9 wt%, and the early leakage rate was unmeasurable. The drug loading rate of Example 6-1, with a pore diameter of 2 nm, was 5.1 wt%, and the early leakage rate was 4.8%. The drug loading rate of Example 6-2, with a pore diameter of 4.5 nm, was 8.2 wt%, and the early leakage rate was 6.3%. The drug loading rate of Example 6-3, with a pore diameter of 10 nm, was 12.7 wt%, and the early leakage rate was 18.5%. The drug loading rate of Comparative Example 6-2, with a pore diameter of 25 nm, was 14.2 wt%, and the early leakage rate was 35.8%.

[0250] The cell endocytosis efficiency in bEnd.3 mouse cerebral vascular endothelial cells was measured for Examples 7-1 to 7-3 and Comparative Examples 7-1 to 7-2. After treating cells with each fluorescently labeled sample at a concentration of 100 g / mL for 4 hours, the proportion of fluorescently positive cells was measured by flow cytometry.

[0251] The cell encapsulation efficiency of Comparative Example 7-1, with a particle size of 15 nm, was 12.3%. The cell encapsulation efficiency of Example 7-1, with a particle size of 50 nm, was 38.7%. The cell encapsulation efficiency of Example 7-2, with a particle size of 150 nm, was 67.2%. The cell encapsulation efficiency of Example 7-3, with a particle size of 300 nm, was 41.5%. The cell encapsulation efficiency of Comparative Example 7-2, with a particle size of 600 nm, was 8.9%.

[0252] Experimental Example 4: Drug Stability and Release Characteristics According to Trehalose Content

[0253] Drug stability and release characteristics were evaluated for Examples 4-1 to 4-5, Comparative Example 4-1, and Comparative Example 4-2. Drug stability was assessed by storing each sample in pH 7.4 phosphate-buffered saline at 37°C for 72 hours, extracting the drug encapsulated within the mesopore, and measuring the residual activity rate by high-speed liquid chromatography. Release characteristics were assessed by switching each sample to pH 6.5 phosphate-buffered saline and measuring the release rate during the initial 30 minutes using an ultraviolet-visible spectrophotometer.

[0254] The trehalose content of Comparative Example 4-1 was 0.3 times by weight, the drug residual activity rate after 72 hours was 42.1%, and the initial 30-minute release rate was 11.3%. The trehalose content of Example 4-1 was 0.5 times by weight, the drug residual activity rate after 72 hours was 61.8%, and the initial 30-minute release rate was 15.7%. The trehalose content of Example 4-2 was 2 times by weight, the drug residual activity rate after 72 hours was 78.3%, and the initial 30-minute release rate was 28.4%. The trehalose content of Example 4-3 was 5 times by weight, the drug residual activity rate after 72 hours was 91.5%, and the initial 30-minute release rate was 52.8%. The trehalose content of Example 4-4 was 7 times by weight, and the drug residual activity rate after 72 hours was 93.2%, and the initial 30-minute release rate was 55.1%. The trehalose content of Example 4-5 was 10 times by weight, and the drug residual activity rate after 72 hours was 94.8%, and the initial 30-minute release rate was 48.6%. The trehalose content of Comparative Example 4-2 was 25 times by weight, and the drug residual activity rate after 72 hours was 95.1%, and the initial 30-minute release rate was 31.2%.

[0255] Meanwhile, for Example 3, which was not co-encapsulated with trehalose, measurements were taken under the same conditions, and the residual drug activity rate after 72 hours was 38.7%, and the initial 30-minute release rate was 8.5%.

[0256] According to the above results, the residual drug activity rate was significantly improved in the trehalose content range of 0.5 to 20 times by weight compared to the case without trehalose. In the trehalose content range of 3 to 7 times by weight, the initial 30-minute release rate exceeded 50%, which suggests that in this range, the osmotic driving force caused by the rapid dissolution of the glassy matrix exhibits a significant release acceleration effect compared to passive diffusion. In Comparative Example 4-1, both the drug stabilization and release acceleration effects were reduced due to incomplete formation of the glassy matrix. In Comparative Example 4-2, although the residual drug activity rate was high, the drug loading amount decreased to 1.8% by weight relative to the weight of the mesoporous silica nanoparticles, and the initial release rate actually decreased as drug release within the mesopores proceeded dispersively due to an excessive increase in osmotic pressure caused by the excess trehalose.

[0257] Experimental Example 5: Evaluation of Reactive Oxygen Species Threshold Gating

[0258] For Examples 5-1 to 5-5, Comparative Example 5-1, and Comparative Example 5-2, the catechol->quinone conversion rate under an active oxygen species concentration gradient was evaluated. Each sample was dispersed in phosphate-buffered saline at pH 6.5 and reacted at 37°C for 1 hour under conditions where the hydrogen peroxide concentration was set to 0.5, 5, 50, 200, and 500 ΩM. Then, the catechol->quinone conversion rate was calculated by measuring the absorbance at 395 nm, which is the quinone-specific absorbance.

[0259] The content of Prussian blue nanoparticles in Comparative Example 5-1 was 0.2 wt%, and the conversion rate was 18.5% at 0.5 μM hydrogen peroxide, 42.3% at 5 μM, 78.9% at 50 μM, 89.1% at 200 μM, and 93.2% at 500 μM. The content of Example 5-1 was 0.5 wt%, and the conversion rate was 8.3% at 0.5 μM hydrogen peroxide, 31.7% at 5 μM, 72.1% at 50 μM, 85.4% at 200 μM, and 91.8% at 500 μM. The content of Example 5-2 was 2 wt%, and the conversion rate was 2.1% at 0.5 µM hydrogen peroxide, 5.8% at 5 µM, 58.3% at 50 µM, 81.2% at 200 µM, and 90.5% at 500 µM. The content of Example 5-3 was 5 wt%, and the conversion rate was 0.8% at 0.5 µM hydrogen peroxide, 2.3% at 5 µM, 45.6% at 50 µM, 78.9% at 200 µM, and 88.7% at 500 µM. The content of Example 5-4 was 8 wt%, and the conversion rate was 0.3% at 0.5 μM hydrogen peroxide, 1.1% at 5 μM, 32.8% at 50 μM, 72.5% at 200 μM, and 85.1% at 500 μM. The content of Example 5-5 was 15 wt%, and the conversion rate was 0.1% at 0.5 μM hydrogen peroxide, 0.5% at 5 μM, 15.2% at 50 μM, 51.3% at 200 μM, and 72.8% at 500 μM. The content of Comparative Example 5-2 was 20 wt%, and the conversion rate was less than 0.1% at 0.5 oz M hydrogen peroxide, 0.2% at 5 oz M, 8.1% at 50 oz M, 32.7% at 200 oz M, and 55.3% at 500 oz M.

[0260] According to the above results, in the range of Prussian blue nanoparticle content of 2 to 8 weight%, the catechol->quinone conversion rate was suppressed to 6% or less at reactive oxygen species concentrations of 0.5 to 5 µM, which is the normal blood flow level, and was found to be 70% or more at 200 to 500 µM, which is the severe disease environment level. In Comparative Example 5-1, conversion proceeded even under low-concentration reactive oxygen species conditions, and non-specific activation was observed. In Comparative Example 5-2, the conversion rate was only 55.3% even under high-concentration reactive oxygen species conditions, resulting in reduced timeliness of drug delivery activation.

[0261] Experimental Example 6: Evaluation of In Vitro Blood-Brain Barrier Permeability

[0262] An in vitro blood-brain barrier model was constructed by culturing bEnd.3 mouse cerebral vascular endothelial cells on the porous polycarbonate membrane of a Transwell insert for 5 days. The transepidermal electrical impedance was 200 cm⁻¹ 2 Only the wells identified above were used in the experiment. The drug delivery systems prepared in Examples 1, 2, 3, 4-3, 8, Comparative Example 1, Comparative Example 2, and Comparative Example 3 were each introduced into the upper chamber at a concentration of 100 µg / mL. For the normal condition group, a pH 7.4 medium was used in the upper chamber, while for the disease simulation condition group, a pH 6.5 medium containing 100 µM hydrogen peroxide and 50 ng / mL recombinant MMP-9 was used. After incubation at 37°C for 4 hours, the drug concentration in the lower chamber was measured by fluorescence analysis, and the apparent transmittance was calculated.

[0263] The apparent transmittance coefficient under normal conditions for Example 1 is 0.8 x 10 6 cm / s, Example 2 is 0.9 x 10 6 cm / s, Example 3 is 1.1 x 10 6 cm / s, Example 4-3 is 1.2 x 10 6 cm / s, Example 8 is 1.0 x 10 6 cm / s, Comparative Example 1 is 0.3 x 10 6cm / s, Comparative Example 2 is 0.7 x 10 6 cm / s, Comparative Example 3 is 3.8 x 10 6 It was cm / s.

[0264] The apparent transmittance coefficient under disease simulation conditions was 8.7 x 10 for Example 1. 6 cm / s, Example 2 is 15.3 x 10 6 cm / s, Example 3 is 21.8 x 10 6 cm / s, Example 4-3 is 28.1 x 10 6 cm / s, Example 8 is 32.5 x 10 6 cm / s, Comparative Example 1 is 0.5 x 10 6 cm / s, Comparative Example 2 is 1.2 x 10 6 cm / s, Comparative Example 3 is 5.1 x 10 6 It was cm / s.

[0265] According to the above results, Examples 1 to 8 showed a significant increase in the apparent permeability under disease simulation conditions compared to the apparent permeability under normal conditions. Compared to Example 1, Example 2 showed an additional increase in permeability due to the opening of mesopores caused by the decomposition of the nanovalve layer and the relaxation of tight junctions caused by zinc ion release. Compared to Example 2, Example 3 showed an additional increase in permeability due to the docking of endothelial cell surfaces by phenylboronic acid residues and the riding of transcytosis by catechol-quinone conversion. Compared to Example 3, Examples 4-3 showed a permeability of 21.8 x 10⁻⁶ due to the addition of an osmotic driving force caused by the rapid dissolution of the trehalose glassy matrix. 6 28.1 x 10⁻⁶ at cm / s 6 It increased to cm / s, indicating that the acceleration of drug release by trehalose co-encapsulation independently contributes to the improvement of blood-brain barrier permeability. Example 8, which had an additional Prussian blue catalyst layer compared to Example 4-3, exhibited the highest permeability under disease simulation conditions, while the permeability under normal conditions was 1.0 x 10⁻⁶. 61.2 x 10 of Example 4-3 at cm / s 6 It was lower than cm / s, which is attributed to the additional suppression of non-specific activation under normal conditions by the scavenging of reactive oxygen species by the Prussian blue catalyst layer. Comparative Example 1 exhibited low permeability under both normal and disease conditions due to the absence of a shell and binding structure. Comparative Example 2 applied a non-reactive shell, so no significant increase in permeability was observed even under disease conditions. Comparative Example 3 showed non-selective permeability with a permeability of 3.8 x 10 cm / s under normal conditions due to continuous exposure to the target ligand, and the permeability under disease simulation conditions was lower compared to Example 3.

[0266] Experimental Example 7: Evaluation of Cell Viability and Tight Junction Recovery

[0267] bEnd.3 Cell viability and tight junction recovery were evaluated on mouse cerebral vascular endothelial cells. Cell viability was measured by the CCK-8 assay. Cell viability was measured after treating with the drug delivery systems prepared in Examples 2, 5-3, and 8 at concentrations of 25, 50, 100, and 200 g / mL for 24 hours.

[0268] The cell viability of Example 2 was 98.3% at 25 g / mL, 96.7% at 50 g / mL, 93.1% at 100 g / mL, and 88.5% at 200 g / mL. The cell viability of Example 5-3 was 97.8% at 25 g / mL, 95.2% at 50 g / mL, 91.8% at 100 g / mL, and 86.3% at 200 g / mL. The cell viability of Example 8 was 97.1% at 25 g / mL, 94.5% at 50 g / mL, 90.2% at 100 g / mL, and 84.7% at 200 g / mL.

[0269] Tight junction recovery was evaluated by the time-dependent change in trans-epidermal resistance in a transwell model. The drug delivery systems of Examples 2 and 8 were introduced into the upper chamber at a concentration of 100 g / mL under pH 6.5 conditions, and the trans-epidermal resistance was measured over time based on the point at which the medium was subsequently changed to pH 7.4.

[0270] The trans-epidermal resistance of Example 2 decreased to 61.2% after 4 hours of treatment under pH 6.5 conditions compared to immediately before the drug delivery system was administered. After changing the medium to pH 7.4, it recovered to 78.5% at 2 hours, 89.3% at 6 hours, 96.8% at 12 hours, and 99.1% at 24 hours. The trans-epidermal resistance of Example 8 decreased to 55.8% after 4 hours of treatment under pH 6.5 conditions. After changing the medium to pH 7.4, it recovered to 72.3% at 2 hours, 85.7% at 6 hours, 94.2% at 12 hours, and 98.5% at 24 hours.

[0271] According to the above results, the cell viability of Examples 2 and 8 was maintained at 90% or higher at a concentration of 100 g / mL. Although tight junctions were temporarily relaxed under pH 6.5 conditions, the transepidermal resistance recovered to 98% or higher of the pre-treatment level within 24 hours of returning to normal pH, confirming reversible relaxation and spontaneous recovery of tight junctions.

[0272] Experimental Example 8: Evaluation of In Vitro Blood-Brain Barrier Permeability of Pharmaceutical Composition

[0273] For Example 9 and Comparative Example 1, the apparent permeability coefficient under disease simulation conditions was measured using the same bEnd.3 Transwell blood-brain barrier model as in Experimental Example 6. The apparent permeability coefficient of Example 9 was 31.8 x 10⁻⁶. 6It was cm / s, and the apparent permeability coefficient of Comparative Example 1 was 0.5 x 10 cm / s. The above results confirmed that the drug delivery system included in the pharmaceutical composition of Example 9 exhibited blood-brain barrier permeability behavior similar to that of Example 8 of Experimental Example 6.

[0274] Experimental Example 9: Evaluation of Brain-Skin Axis Regulatory Activity

[0275] For Example 10 and Comparative Example 8, a co-culture system was established in which human keratinocytes, specifically HaCaT cells, were additionally cultured in the lower chamber of the bEnd.3 Transwell blood-brain barrier model and evaluated. Each sample was introduced into the upper chamber and cultured for 8 hours under disease simulation conditions of pH 6.5, hydrogen peroxide 100 µM, and recombinant MMP-9 50 ng / mL. Subsequently, the phosphorylation level of TrkB, a brain-derived neurotrophic factor receptor, in the HaCaT cells of the lower chamber was measured by Western blot.

[0276] In the case of Example 10, the level of TrkB phosphorylation in HaCaT cells in the lower chamber increased 3.8 times compared to the untreated group. In the case of Comparative Example 8, the level of TrkB phosphorylation was 1.2 times compared to the untreated group. In addition, when the secretion amount of procollagen type I from HaCaT cells was measured by enzyme-linked immunosorbent assay, Example 10 was 2.5 times higher than the untreated group, and Comparative Example 8 was 1.1 times higher than the untreated group. The above results indicate that the drug delivery system included in the composition of Example 10 passed through the blood-brain barrier to deliver brain-derived neurotrophic factor to the brain tissue, and then contributed to the activation of TrkB receptors and collagen synthesis in keratinocytes via the brain-skin axis pathway.

[0277] Experimental Example 10: Evaluation of Sequential Permeation of Simulated Gastrointestinal-Blood-Brain Barrier by a Health Functional Food Composition

[0278] For Example 11, Comparative Example 9, and Comparative Example 10, the blood-brain barrier permeability was evaluated sequentially through treatment with simulated gastric fluid followed by treatment with simulated intestinal fluid. Each sample was immersed in a pH 1.2 hydrochloric acid-pepsin solution, which is simulated gastric fluid, at 37°C for 2 hours. Afterward, the sample was retrieved and immersed in a pH 6.8 phosphate-buffered saline-pancreatin solution, which is simulated intestinal fluid, at 37°C for 4 hours. After treatment with the simulated intestinal fluid, the curcumin released in the supernatant was quantified by high-speed liquid chromatography, and the sample was introduced into the upper chamber of the bEnd.3 Transwell blood-brain barrier model to measure the apparent permeability coefficient under disease simulation conditions.

[0279] After 2 hours of treatment with simulated gastric fluid, the curcumin leakage rate was 2.3% for Example 11, was unmeasurable for Comparative Example 9, and 38.7% for Comparative Example 10. In Comparative Example 9, since the curcumin was in a powder state not encapsulated in the drug delivery system, most of the curcumin decomposed in the simulated gastric fluid, and the amount of curcumin recovered after treatment with simulated intestinal fluid was only 8.5% of the initial amount. In Comparative Example 10, the enteric coating layer was absent, causing the stimulus-responsive shell of the drug delivery system to be partially damaged in the simulated gastric fluid, resulting in early leakage.

[0280] As a result of introducing the drug delivery system recovered by simulating intestinal absorption after 4 hours of treatment with simulated serous fluid into a blood-brain barrier model, the apparent permeability coefficient under disease simulation conditions was 25.3 x 10⁻⁶ for Example 11. 6 cm / s, Comparative Example 9 is 0.2 x 10 6 cm / s, Comparative Example 10 is 8.7 x 10 6 It was cm / s. The above results indicate that the enteric coating layer of Example 11 protects the drug delivery system in an acidic environment and allows it to be released normally in the intestinal tract, thereby maintaining the pathway to cross the blood-brain barrier via blood circulation.

[0281] Although embodiments of the present invention have been described in detail above, the scope of the present invention is not limited thereto, and it will be obvious to those skilled in the art that various modifications and variations are possible within the scope of the technical concept of the present invention as described in the claims.

Claims

Claim 1 A core containing a drug active ingredient; a stimulus-responsive shell formed on the outer side of the core; and a reactive binding structure disposed on the outer surface of the stimulus-responsive shell; wherein the core comprises mesoporous silica nanoparticles having a pore diameter of 3 to 6 nm and a particle size of 80 to 200 nm, the drug active ingredient is supported within the pores of the mesoporous silica nanoparticles, and further comprises trehalose co-encapsulated together with the drug active ingredient within the mesopores of the mesoporous silica nanoparticles, wherein the trehalose forms a glassy matrix within the mesopores, and the content of the trehalose is 3 to 7 times by weight relative to the weight of the drug active ingredient, and further comprises a metal-organic framework nanovalve layer disposed at the entrance of the mesopores of the mesoporous silica nanoparticles, wherein the metal-organic framework is a zeolite imidazolate framework comprising zinc ions and 2-methylimidazole, and the stimulus-responsive shell comprises a block copolymer comprising a pH-responsive segment and a hydrophilic segment, and the A drug delivery system with blood-brain barrier penetration technology, wherein the block copolymer is poly(α-aminoester)-block-polyethylene glycol, further comprises a phenylboronic acid residue disposed on the outer surface of the stimulus-responsive shell and a catechol group introduced into the stimulus-responsive shell, further comprises a catalyst layer comprising Prussian blue nanoparticles on the outermost layer of the stimulus-responsive shell, wherein the content of the Prussian blue nanoparticles is 2 to 8 weight% relative to the total weight of the drug delivery system, wherein the stimulus-responsive shell undergoes a structural transition from a shielded state to a relaxed state upon stimulation of the microenvironment surrounding the blood-brain barrier, wherein the responsive binding structure comprises a masking linker and a target ligand bound to the masking linker, wherein the metal-organic framework nanovalve layer decomposes upon a pH change of the microenvironment surrounding the blood-brain barrier to release zinc ions, and wherein the released zinc ions interact with blood-brain barrier tight junction proteins to reversibly relax intercellular pathways. Claim 2 A drug delivery system capable of penetrating the blood-brain barrier according to claim 1, wherein the stimulation of the microenvironment surrounding the blood-brain barrier is one or more of a change in pH, a change in reactive oxygen species concentration, and a change in enzyme concentration. Claim 3 A method for manufacturing a blood-brain barrier penetration drug delivery system comprising: a core forming step of forming a core; a drug encapsulation step of encapsulating a drug active ingredient in the core; a nanovalve layer forming step of forming a nanovalve layer comprising a zeolite imidazolate framework by reacting zinc ions and 2-methylimidazole at the mesopore entrance of the core encapsulated with the drug active ingredient; a shell forming step of forming a stimulus-responsive shell on the outer side of the core encapsulated with the drug active ingredient; and a surface functionalization step of forming a reactive binding structure comprising a masking linker and a target ligand on the outer surface of the stimulus-responsive shell; wherein the core forming step comprises a step of synthesizing mesoporous silica nanoparticles by a sol-gel method, and wherein, in the shell forming step, the stimulus-responsive shell comprises a block copolymer comprising a pH-responsive segment and a hydrophilic segment, and the stimulus-responsive shell undergoes a structural transition from a shielded state to a relaxed state upon stimulation by the microenvironment surrounding the blood-brain barrier. Claim 4 delete Claim 5 A composition for improving skin condition through brain-skin axis regulation, comprising: a drug delivery system according to claim 1; and a pharmaceutically acceptable carrier; wherein the drug active ingredient is a neuromodulator having brain-skin axis regulatory activity. Claim 6 A health functional food composition for improving brain health, comprising: a drug delivery system according to claim 1; and a food-grade acceptable carrier; wherein the drug active ingredient is a functional ingredient for brain health, and an enteric coating layer for oral administration is formed on the outer surface of the drug delivery system.

Images