A method for preparing an charged biomolecule delivery system

Abstract

This invention discloses a universal delivery method for charged biomolecules, belonging to the field of biomolecule delivery technology. The method involves physically mixing charged biomolecules with a charged, large-π-conjugated guide molecule, driving direct self-assembly into a nanoscale delivery system under non-covalent forces. Polyethylene glycol is then used for surface cross-linking to stabilize the self-assembled structure and shield the surface charge, thereby improving the in vivo pharmacokinetic properties of the nanoparticles. This method enables efficient and precise delivery of charged biopharmaceutical molecules with good stability and biocompatibility. The assembly method in this invention will provide a new paradigm for developing easily fabricated and multifunctional nanomedicines and guide a wider range of biomolecule delivery methods.

Description

Technical Field

[0001] This invention belongs to the field of biomolecule delivery technology, specifically relating to a universal delivery method for charged biomolecules. Background Technology

[0002] Biomolecules (such as peptides, nucleic acids, and polysaccharides) possess high target selectivity, high activity, and good biocompatibility, showing promising prospects in the treatment of cancer and other diseases. However, most biomolecules exhibit drawbacks such as poor drug-likeness, short blood circulation time, susceptibility to enzymatic hydrolysis, non-specific tissue distribution, and low cellular uptake, hindering their application in disease treatment. Nanodrug delivery systems (NDDSs), such as liposomes, micelles, and polymer nanoparticles, have been widely used for biomolecule delivery. NDDSs offer high controllability, multifunctionality, and tumor selectivity, enabling the specific delivery of payloads to tumors through passive accumulation or active targeting, and allowing for spatiotemporal control of drug release to achieve multiple functions. To effectively treat diseases, it is necessary to increase the concentration of biomolecules within cells. However, traditional NDDSs typically have poor drug loading capacity (less than 15%), insufficient to deliver adequate therapeutic drugs to the lesion. Correspondingly, increasing the dosage of NDDSs can deliver more biomolecules, but this increase may introduce more potential adverse toxicities to the body. Meanwhile, the complex chemical synthesis of NDDSs also leads to unpredictable toxicity and high drug development costs. Therefore, increasing drug loading and reducing complex chemical synthesis are of great value in improving the delivery efficiency, therapeutic efficacy, and safety of biomolecules, as well as reducing development costs.

[0003] NDDSs encounter a series of biological barriers on their journey from entering the bloodstream to exerting their effects in tumor cells. Generally, due to some unavoidable contradictions, NDDSs struggle to overcome these barriers simultaneously. For example, positively charged NDDSs facilitate cellular uptake but are susceptible to opsonin-induced capture by macrophages, thus shortening blood circulation time; negatively charged NDDSs can mitigate opsonin effects to some extent, but they are difficult to endocytose and deliver drugs to target cells. Structurally stable NDDSs can effectively load drugs but may not release them efficiently; loosely structured NDDSs facilitate complete drug release but are detrimental to their stability in the bloodstream. To balance these intractable contradictions, researchers have developed environmentally responsive NDDSs. Chen et al. developed an ultrasensitive pH-triggered, charge- and size-convertible NDDS to improve gene delivery barriers. However, due to the high heterogeneity and complexity of tumor tissues, single-treatment strategies still face insufficient efficacy. In previous studies, our research group developed a dual-size / charge-convertible crosslinked NDDS that responds to the tumor microenvironment, exhibiting longer blood circulation time, deeper tissue penetration, and stronger intracellular internalization. However, these strategies involve complex chemical synthesis and preparation, and suffer from low drug loading capacity. Therefore, there is an urgent need to develop an easy-to-use and more comprehensive NDDS that addresses these issues, achieving higher drug loading capacity, overcoming diverse biological barriers, effectively delivering biomolecules, and synergistically working with other therapeutic approaches. Summary of the Invention

[0004] In order to overcome the shortcomings of the prior art, the present invention aims to provide a universal delivery method for charged biomolecules, so as to solve the technical problems of existing nanomedicine delivery systems, such as poor drug loading capacity, different biological barriers, and poor delivery of biomolecules.

[0005] To achieve the above objectives, the present invention employs the following technical solution:

[0006] This invention discloses a universal delivery method for charged biomolecules, comprising the following steps:

[0007] S1: The charged large π conjugated guiding molecule and the charged biomolecule are premixed in an organic solvent to obtain a premixed working solution;

[0008] S2: After mixing the premixed working solution and deionized water, the mixture is stirred and evaporated, and a nanoscale delivery system is obtained by self-assembly under the drive of non-covalent forces.

[0009] S3: Mix polyethylene glycol and a nanoscale delivery system, and perform a surface cross-linking reaction on the nanoscale delivery system to obtain a polyethylene glycol cross-linked biomolecule delivery system, thereby achieving universal delivery of charged biomolecules.

[0010] The charged biomolecules have the opposite charge type to the large π-conjugated guide molecules.

[0011] Furthermore, in S1, the molar ratio of the charged large π-conjugated guiding molecule to the charged biomolecule is (0.5~100):1.

[0012] Furthermore, the ratio of the charged biomolecule to the organic solvent is (0.1~50) mg / mL.

[0013] Furthermore, in S1, the charged large π-conjugated directing molecule is dihydroporphyrin, pheophytin, tetraaminophenylporphyrin, or tetracarboxyphenylporphyrin.

[0014] Furthermore, the organic solvent is tetrahydrofuran or methanol.

[0015] Furthermore, in S1, the charged biomolecules are polypeptides, nucleic acids, or polysaccharides.

[0016] Furthermore, in S1, the charged biomolecule is a polypeptide drug, small interfering RNA, messenger RNA, hyaluronic acid, or chitosan.

[0017] Furthermore, in S2, the premixed working solution and deionized water are mixed by stirring at room temperature; the volume ratio of the premixed working solution to deionized water is 10%~50%.

[0018] Furthermore, in S3, the surface of the nanoscale delivery system is coated with amino or carboxyl groups;

[0019] When the surface of the nanoscale delivery system has amino groups, the polyethylene glycol has aldehyde, hydrazide or carboxyl groups. In this case, the nanoscale delivery system and polyethylene glycol are mixed at an equivalent ratio of amino to aldehyde, hydrazide or carboxyl groups of 1: (1~2).

[0020] When the surface of the nanoscale delivery system has carboxyl groups, the polyethylene glycol has amino or hydroxyl groups. In this case, the nanoscale delivery system and polyethylene glycol are mixed at an equivalent ratio of carboxyl groups to amino or hydroxyl groups of 1:(1~2).

[0021] Furthermore, in S3, the surface crosslinking reaction is carried out at room temperature for 48 hours.

[0022] Compared with the prior art, the present invention has the following beneficial effects:

[0023] This invention discloses a universal delivery method for charged biomolecules. It primarily involves physically mixing charged biomolecules with a charged large-π conjugated guide molecule, and then directly self-assembling them under non-covalent forces to form a nanoscale delivery system (NDDSs for delivering charged biomolecules). If the charged biomolecules are positively charged, the self-assembly guide molecule (the charged large-π conjugated guide molecule) must be a chemical molecule with the opposite charge. This method eliminates the need for chemical synthesis, greatly simplifying the preparation of NDDSs, and achieves high drug loading capacity. Experimental results demonstrate that nanostructures can be successfully assembled using a similar method. This indicates that the assembly method proposed in this invention can broadly guide the improvement of drug delivery through supramolecular assembly of biopharmaceuticals. The aforementioned nanoscale delivery system is surface-crosslinked with polyethylene glycol to stabilize its self-assembled structure and shield surface charges, thereby improving the in vivo pharmacokinetic properties of the nanoparticles. According to relevant experimental results, the polyethylene glycol-crosslinked biomolecular delivery system of this invention can significantly increase drug loading and deliver drugs efficiently and accurately, without premature leakage under physiological conditions, and is less susceptible to opsonin effects, exhibiting good stability and biocompatibility. The assembly method in this invention will provide a new paradigm for developing easily manufactured and multifunctional nanomedicines and guide a wider range of biomolecular delivery methods. Attached Figure Description

[0024] Figure 1 The structural characterization diagrams of the dKLA and Ce6 molecules in Example 1 are shown.

[0025] Among them: a-dKLA; b-Ce6 molecules;

[0026] Figure 2 The polydispersity index and potential statistics of the supramolecular assembled nanoparticles dKLA-Ce6 NPs obtained in Examples 1 to 4 are shown in the figure.

[0027] Wherein: c - polydispersity index statistical chart; d - potential statistical chart;

[0028] Figure 3 TEM image of the supramolecular assembled nanoparticles dKLA-Ce6 NPs obtained in Example 1;

[0029] Figure 4 This is a schematic diagram of the process of obtaining supramolecular assembled nanoparticles in Example 1;

[0030] Wherein: f - the process of dynamically forming stable nanostructures; g - a schematic diagram of hydrogen bonding;

[0031] Figure 5 SASA analysis of the supramolecular assembled nanoparticles obtained in Example 1;

[0032] Where: h - co-assembled structure; i - number of hydrogen bonds; j - hydrogen bond interaction;

[0033] Figure 6 This is a schematic diagram of the characterization of the DLS (Delivery System for Charged Biomolecules) obtained in Example 5.

[0034] Where: a-Size chart; b-PDI chart;

[0035] Figure 7 The diagram shows the size and PDI distribution of the charged biomolecule delivery system obtained in Example 6.

[0036] Figure 8 The diagram shows the size and PDI distribution of the charged biomolecule delivery system obtained in Example 7.

[0037] Figure 9 The diagram shows the size and PDI distribution of the charged biomolecule delivery system obtained in Example 8.

[0038] Figure 10 The image shows the addition of different concentrations of PEG. 2k Schematic diagram of the performance of DC NPs crosslinked with PEG after -2CHO;

[0039] Wherein: a- Schematic diagram of particle size and PDI; b- Schematic diagram of nanoparticle charge; c- TEM morphology image of PDC NPs; d- Stability of PDC NPs; e- Stability of PDC NPs; f- Schematic diagram of drug release under different physiological pH conditions;

[0040] Figure 11 This is a schematic diagram illustrating the size and surface charge characterization after incubation at pH 6.5 for 24 hours.

[0041] Wherein: a- Dimensional diagram; b- TEM micrograph; c- Surface charge; d- Intrinsic fluorescence assessment;

[0042] Figure 12 The pharmacokinetic properties of three formulations: Ce6, DC NPs, and PDC NPs;

[0043] Among them: a-blood clearance effect; b-near-infrared fluorescence (NIRF) imaging measurement; c-quantitative fluorescence; d and e-in vitro distribution of PDCNPs relative to Ce6 and DC NPs; f and g-frozen section results of tumor tissue;

[0044] Figure 13 This is a schematic diagram illustrating the good anti-tumor effect of PDC NPs in vitro.

[0045] a-Animal experimental process; b-Tumor growth curve; c and d-Tumor size and weight; e-Mouse body weight growth;

[0046] Figure 14 A schematic diagram illustrating the drug tolerability and biosafety of PDC NPs+L combination therapy;

[0047] Wherein: a-Schematic diagram of WBC content; b-Schematic diagram of histopathological examination;

[0048] Figure 15 The diagram shows a schematic representation of the charged biomolecule delivery system and its preparation and universal delivery method disclosed in this invention. Detailed Implementation

[0049] To facilitate understanding of the features and effects of the present invention by those skilled in the art, the following provides a general description and definition of the terms and expressions used in the specification and claims. Unless otherwise indicated, all technical and scientific terms used herein have the ordinary meanings as understood by those skilled in the art regarding the present invention. In the event of conflict, the definitions in this specification shall prevail.

[0050] The theories or mechanisms described and disclosed herein, whether right or wrong, should not in any way limit the scope of the invention, that is, the contents of the invention can be implemented without being limited by any particular theory or mechanism.

[0051] All features, such as values, amounts, contents, and concentrations, described herein as numerical ranges or percentage ranges are provided for simplicity and convenience only. Accordingly, the description of numerical ranges or percentage ranges should be considered to include and specifically disclose all possible subranges and individual values ​​within the range (including integers and fractions).

[0052] In this document, unless otherwise specified, “include,” “including,” “contains,” “has” or similar terms cover the meanings of “consisting of” and “mainly consisting of,” for example, “A includes a” covers the meanings of “A includes a and other” and “A only includes a.”

[0053] For the sake of brevity, not all possible combinations of the technical features in each implementation scheme or embodiment are described herein. Therefore, as long as there is no contradiction in the combination of these technical features, the technical features in each implementation scheme or embodiment can be combined arbitrarily, and all possible combinations should be considered within the scope of this specification.

[0054] This invention provides a universal delivery method for charged biomolecules, which can broadly guide the development of nanoscale delivery systems for biomolecules such as peptides, nucleic acids, and polysaccharides.

[0055] The specific methods and steps are as follows: Figure 15As shown, X represents charged biomolecules, including polypeptide drugs with amino or carboxyl groups, such as apoptosis-promoting polypeptides (dKLA), transmembrane bee venom peptides (p5RHH), exenatide, liraglutide, teriparatide, octreotide, eptifibatide, etc.; nucleic acid drugs with phosphate groups, such as small interfering ribonucleic acid (siRNA), messenger ribonucleic acid (mRNA), and deoxyribonucleic acid (DNA); and polysaccharides with amino, carboxyl, or sulfonic acid groups, such as chitosan, hyaluronic acid, heparin, and chondroitin sulfate.

[0056] Y represents a guide molecule with a conjugated large π structure that carries the opposite charge to the biomolecule: such as charged porphyrin molecules, dihydroporphyrin e6 (chlorin e6, Ce6), protoporphyrin, pheophytin a, pyropheophytin a, 5,10,15,20-tetra(4-aminophenyl)porphyrin TAPP, and 5,10,15,20-tetra(4-carboxyphenyl)porphyrin TCPP.

[0057] Z represents the polyethylene glycol (PEG) crosslinking process: PEG 1000, PEG 2000, PEG 5000, PEG 10000, multi-arm PEG 1000, multi-arm PEG 2000, multi-arm PEG 5000, and multi-arm PEG 10000 all have active chemical groups at their ends. All ends of the PEG molecule, such as the two ends of linear PEG and the multiple ends of multi-arm PEG, must have chemically active groups that can react with X or Y to facilitate the crosslinking and reinforcement of charged biomolecule delivery systems by PEG. Linear PEG has two ends, four-arm PEG has four ends, eight-arm PEG has eight ends, and so on.

[0058] The main steps include:

[0059] By physically mixing charged large π-conjugated guide molecules, such as dihydroporphyrin (Chlorin e6, Ce6), tetraaminophenylporphyrin (TAPP), and tetracarboxyphenylporphyrin (TCPP), with charged biomolecules such as peptides (e.g., the pro-apoptotic peptide dKLA), nucleic acids (e.g., small interfering RNA (siRNA), messenger RNA (mRNA)), and polysaccharides (e.g., hyaluronic acid (HA), chitosan (CTS)), they directly self-assemble into nanoscale delivery systems (NDDSs for delivering charged biomolecules) under the drive of non-covalent forces. Furthermore, polyethylene glycol is used to surface-crosslink these NDDSs to stabilize their structure and endow the delivery system with charge-changing capabilities to overcome biological barriers, thus forming polyethylene glycol-crosslinked biomolecule delivery systems (PEG-crosslinked biomolecule NDDSs). If the charged biomolecules have a positive charge, then the self-assembly guide molecule must be a chemical molecule with an opposite charge. This method does not require chemical synthesis, which can greatly simplify the preparation of NDDSs and has a high drug loading capacity.

[0060] Preferably, if X is a negatively charged siRNA, mRNA, DNA, hyaluronic acid, heparin, etc., the Y-directing molecule can be selected to co-assemble with TAPP to form a charged biomolecule delivery system, and reinforced and cross-linked by polyethylene glycol with benzaldehyde at both ends.

[0061] Preferably, if X is a positively charged dKLA, p5RHH, chitosan, teriparatide, etc., the guiding molecule Y can be TCPP, Ce6, protoporphyrin, pheophytin, etc., and it is reinforced and crosslinked by polyethylene glycol with benzaldehyde at both ends.

[0062] Preferably, if X or Y has an amino group, the end of Z needs to have an aldehyde group, hydrazide group, carboxyl group, etc., to form a Schiff base for cross-linking; if X or Y has a carboxyl group, the end of Z needs to have an amino group, hydroxyl group, etc., that can react with the carboxyl group.

[0063] When the surface of the nanoscale delivery system has amino groups, the polyethylene glycol has aldehyde, hydrazide or carboxyl groups. In this case, the nanoscale delivery system and polyethylene glycol are mixed at an equivalent ratio of amino to aldehyde, hydrazide or carboxyl groups of 1: (1~2).

[0064] When the surface of the nanoscale delivery system has carboxyl groups, the polyethylene glycol has amino or hydroxyl groups. In this case, the nanoscale delivery system and polyethylene glycol are mixed at an equivalent ratio of carboxyl groups to amino or hydroxyl groups of 1:(1~2).

[0065] The structural formula of protoporphyrin is shown below:

[0066] ;

[0067] The structural formula of pheophorbide a is shown below:

[0068] ;

[0069] The structural formula of pyropheophorbide a is shown below:

[0070] ;

[0071] The structural formula of dihydroporphyrin E6 (chlorin E6) is shown below:

[0072] ;

[0073] The structural formula of TAPP: 5,10,15,20-tetrakis-(4-aminophenyl)porphyrin is shown below:

[0074] ;

[0075] The structural formula of TCPP: 5,10,15,20-tetrakis-(4-carboxyphenyl)porphyrin is shown below:

[0076] ;

[0077] The structural formula of heparin is shown below:

[0078] ;

[0079] The structural formula of chitosan is shown below:

[0080] ;

[0081] The structural formula of hyaluronic acid is shown below:

[0082] ;

[0083] The structural formula of chondroitin sulfate is shown below:

[0084] ;

[0085] The structural formula of teriparatide is shown below:

[0086] ;

[0087] The structural formula of exenatide is shown below:

[0088] ;

[0089] The structural formula of liraglutide is shown below:

[0090] ;

[0091] The structural formula of octreotide is shown below:

[0092] ;

[0093] The structural formula of eptifibatide is shown below:

[0094] ;

[0095] When X or Y has an amino group, and Z has an acylhydrazine group, a hydrazide structure is formed:

[0096] ;

[0097] When X or Y has an amino group, Z has a carboxyl group, forming an amide bond structure;

[0098] ;

[0099] When X or Y has a carboxyl group, Z has an amino group, forming an amide bond structure;

[0100] ;

[0101] When X or Y has a carboxyl group, Z has a hydroxyl group, forming an ester bond structure;

[0102] .

[0103] The present invention will be further illustrated below with reference to specific embodiments. It should be understood that these embodiments are for illustrative purposes only and are not intended to limit the scope of the invention. Furthermore, it should be understood that after reading the teachings of this invention, those skilled in the art can make various alterations or modifications to the invention, and these equivalent forms also fall within the scope defined by the appended claims.

[0104] The following examples use instruments and equipment conventional in the art. Experimental methods in the following examples, unless otherwise specified, are generally performed under conventional conditions or as recommended by the manufacturer. All raw materials used in the following examples are conventional commercially available products with specifications conventional in the art. In this specification and the following examples, unless otherwise specified, "%" refers to weight percentage, "parts" refers to parts by weight, and "ratio" refers to weight proportion.

[0105] Example 1

[0106] A universal method for delivering charged biomolecules, comprising the following steps:

[0107] S1: Dihydroporphyrin (Ce6) and pro-apoptotic peptide (dKLA) are premixed to obtain a premixed working solution, wherein the molar ratio of pro-apoptotic peptide (dKLA) to dihydroporphyrin (Ce6) is 1:0.5; and the volume ratio of dKLA to tetrahydrofuran is 0.1 mg / mL.

[0108] S2: Add the premixed working solution to deionized water at room temperature and stir until the organic solvent evaporates completely. Driven by non-covalent forces, it directly self-assembles into a nanoscale delivery system (supramolecular assembled nanoparticles dKLA-Ce6 NPs). The volume ratio of the premixed working solution to deionized water is 20%.

[0109] S3: The DC NPs obtained in Example 1 were mixed with PEG in a 1:1 ratio of amino to aldehyde groups for surface crosslinking to stabilize their self-assembled structure. The surface crosslinking was carried out at room temperature for 48 hours to improve the in vivo circulation properties of the nanoparticles, resulting in a PEG (polyethylene glycol crosslinked biomolecular delivery system) with an aldehyde group at each end. The aldehyde group can interact with the amino group on the DC NPs to form a Schiff base, which can shield the positive charge of the DC NPs.

[0110] Example 2

[0111] Unlike Example 1, in this example, the molar ratio of the pro-apoptotic peptide (dKLA) to dihydroporphyrin (Ce6) is 1:1; the remaining preparation steps and parameters are the same as in Example 1, resulting in a nanoscale delivery system (supramolecular assembled nanoparticles dKLA-Ce6 NPs).

[0112] Example 3

[0113] Unlike Example 1, in this example, the molar ratio of the pro-apoptotic peptide (dKLA) to dihydroporphyrin (Ce6) is 1:2; the remaining preparation steps and parameters are the same as in Example 1, resulting in a nanoscale delivery system (supramolecular assembled nanoparticles dKLA-Ce6 NPs).

[0114] Example 4

[0115] Unlike Example 1, in this example, the molar ratio of the pro-apoptotic peptide (dKLA) to dihydroporphyrin (Ce6) is 1:3; the remaining preparation steps and parameters are the same as in Example 1, resulting in a nanoscale delivery system (supramolecular assembled nanoparticles dKLA-Ce6 NPs).

[0116] Example 5

[0117] Unlike Example 1, the charged biomolecule is negatively charged small interfering ribonucleic acid (siRNA), the charged large π-conjugated guiding molecule is tetraaminophenylporphyrin (TAPP), and the molar ratio of TAPP to siRNA is 10:1, 25:1, 50:1, and 100:1. The remaining preparation steps and parameters are the same as in Example 1, resulting in four nanoscale delivery systems.

[0118] Example 6

[0119] Unlike Example 1, the charged biomolecule is negatively charged messenger ribonucleic acid (mRNA), the charged large π-conjugated guide molecule is tetraaminophenylporphyrin (TAPP), and the molar ratio of TAPP to mRNA is 20:1. The remaining preparation steps and parameters are the same as in Example 1, resulting in a nanoscale delivery system.

[0120] Example 7

[0121] Unlike Example 1, the charged biomolecule is hyaluronic acid (HA), a sugar molecule with a negative charge, and the charged large π-conjugated guiding molecule is tetraaminophenylporphyrin (TAPP). The molar ratio of TAPP to HA is 1:1. The remaining preparation steps and parameters are the same as in Example 1, resulting in a nanoscale delivery system.

[0122] Example 8

[0123] Unlike Example 1, the charged biomolecule is chitosan (CTS), a positively charged carbohydrate molecule, and the charged large π-conjugated directing molecule is tetracarboxyphenylporphyrin (TCPP), which has the opposite charge. The molar ratio of TCPP to CTS is 1:1. The remaining preparation steps and parameters are the same as in Example 1, resulting in a nanoscale delivery system.

[0124] Example 9

[0125] Unlike Example 1, the organic solvent used was methanol, and the ratio of dKLA to methanol was 30 mg / mL; the volume ratio of the premixed working solution to deionized water was 45%, and the remaining preparation steps and parameters were the same as in Example 1, resulting in a nanoscale delivery system.

[0126] Example 10

[0127] Unlike Example 1, the organic solvent used was methanol, and the ratio of dKLA to methanol was 50 mg / mL; the volume ratio of the premixed working solution to deionized water was 50%, and the remaining preparation steps and parameters were the same as in Example 1, resulting in a nanoscale delivery system.

[0128] Figure 1 The structural characterization diagrams of the dKLA and Ce6 molecules in Example 1 are shown below. Figure 1 As can be seen in a, the pro-apoptotic peptide dKLA has 6 amino groups and carries a strong positive charge; from Figure 1 As can be seen from b, the Ce6 molecule has a large π-conjugated structure and three electronegative carboxyl groups (see appendix). Figure 1 (b) It can co-assemble with dKLA to form nanostructures (DC NPs) through non-covalent interactions such as electrostatic interactions, hydrogen bonds and hydrophobic forces.

[0129] Figure 2 The polydispersity index and potential statistics of the supramolecular assembled nanoparticles dKLA-Ce6 NPs obtained in Examples 1 to 4 are shown in the figure. Figure 2 As shown in c, different proportions of dKLA and Ce6 can form nanoparticles with a hydrodynamic diameter distribution of about 100 nm and a polydispersity index (PDI) of about 0.2, indicating that DC NPs have good dispersibility. Figure 2 As shown in Figure d, the Zeta potential results indicate that all DCNPs have a positive surface charge, and the potential tends to decrease when more Ce6 is introduced into the system. Considering the high drug loading and amino-mediated surface modification, we ultimately determined that a final mixing ratio of dKLA to Ce6 of 1:2 is optimal in the DCNP system.

[0130] The morphology of DCNPs was characterized by TEM, such as Figure 3 As shown, DC NPs have a uniform morphology and a regular spherical structure. The actual size of the nanoparticles is about 80 nm. They are well dispersed and do not aggregate.

[0131] To better understand the assembly mechanism of the dKLA-Ce6 co-assemblies from a theoretical perspective, a 500 ns molecular dynamics simulation was performed on the self-assembly system. Using hydrogen bonds as a characterization of the co-assembly process, we calculated the number of hydrogen bonds, which not only demonstrates the stability of the entire simulation but also assesses the potential for the Ce6 / dKLA composite to form nanoparticles through hydrogen bond formation. As shown in Table 1, the Ce6 / dKLA composite exhibits a high average number of hydrogen bonds, indicating that this composite has sufficient potential to form a more compact co-assembly structure. Figure 4 f visually illustrates the process of dynamic formation of stable nanostructures from an initial random configuration (0 ns). Further analysis of binding interactions shows that hydrogen bonds mainly occur between molecules (N–H··O=C) and within peptides (…). Figure 4 Specifically, the hydrogen atom on the amino group of the lysine residue can potentially form a hydrogen bond with the oxygen atom on the Ce6 molecule. Furthermore, π-π interactions and van der Waals interactions are observed between the large planar rings, contributing to the internal stable arrangement. It should be noted that Ce6, in addition to its normal hydrogen bond donor and acceptor, also possesses an ethynyl group. Studies have shown that the ethynyl moiety has multiple functions. For example, its π-cloud can mimic aromatic groups, facilitating the formation of π-π interactions. Analysis of the solvent-accessible surface area (SASA) indicates that hydrophobic interactions contribute to the co-assembly of Ce6 and dKLA molecules. Figure 5 h). Figure 5 i shows the number of hydrogen bonds generated between the co-assembled structure and the solvent water, and the reduced number of hydrogen bonds further confirms the hydrophobic interaction. Furthermore, the increased number of intermolecular hydrogen bonds between Ce6 and dKLA indicates that hydrogen bonding interactions also play a significant role in the self-assembly of DCNPs. Figure 5 Therefore, hydrogen bonding, hydrophobic interaction, and π-π stacking interaction provide important driving forces for the self-assembly of the two drugs, enabling them to spontaneously form dense nanoparticles with 100% drug loading through supramolecular self-assembly.

[0132] Table 1. Dynamic changes of hydrogen bonds in the last 100 ns of the simulation process

[0133]

[0134] Based on the results of Example 1, we applied this assembly rule to the assembly of more biomolecule delivery systems and characterized the nanoparticles in Examples 5 to 8 using DLS. Figure 6The figure shows a schematic diagram of the characterization of the charged biomolecule delivery system DLS obtained in Example 5. As can be seen from the figure, tetraaminophenylporphyrin (TAPP) and siRNA can be co-assembled into nanoparticles under different molar ratios. The particle size of the nanoparticles increases with the increase of the TAPP ratio, but the corresponding PDI is around 0.2, indicating that the nanoparticle size has good dispersibility and uniformity.

[0135] Figure 7 The figure shows the size distribution of the charged biomolecule delivery system obtained in Example 6. As can be seen from the figure, tetraaminophenylporphyrin (TAPP) and mRNA can be co-assembled to form 129 nm nanoparticles under the condition of nitrogen-phosphorus ratio of 2.5:1, and the PDI is about 0.1, indicating that the nanoparticles have good dispersibility and uniformity.

[0136] Figure 8 The figure shows the size distribution of the charged biomolecule delivery system obtained in Example 7. As can be seen from the figure, the formed nanoparticles have a particle size of about 120 nm and a PDI of 0.09.

[0137] Figure 9 The figure shows the size and PDI distribution of the charged biomolecule delivery system obtained in Example 8. As can be seen from the figure, the formed nanoparticles have a particle size of about 37 nm and a PDI of 0.28.

[0138] The biomolecular delivery systems assembled in Examples 1-8 were cross-linked with polyethylene glycol to stabilize their self-assembled structure and shield surface charges, thereby improving the in vivo pharmacokinetic properties of the nanoparticles. Simultaneously, a weakly acid-sensitive Schiff base was introduced to endow the nanoparticles with charge-reversal properties responsive to the tumor microenvironment, thus promoting the accumulation of nanoparticles at the tumor site and achieving better anti-tumor effects.

[0139] Taking Example 1 as an example, by adding different concentrations of PEG 2k -2CHO was used to optimize PEG-crosslinked DC NPs (PDCNPs); where dKLA and PEG 2k The molar ratios of -2CHO were 1:0.1, 1:0.2, 1:0.5, and 1:1; subsequently, the performance of the formed polyethylene glycol cross-linked biomolecule delivery system was characterized, such as... Figure 10 As shown; from Figure 10 a can be seen from the introduction of PEG 2k After -2CHO, the particle size of PDC NPs increased from 100 nm to 120-130 nm, while PDI remained at ~0.2, indicating that PEG crosslinking slightly increases the size of nanoparticles, but does not affect particle size uniformity. Figure 10 b shows that when dKLA and PEG2k When the molar ratio of -2CHO is 1:0.1 and 1:0.2, the nanoparticles still possess a positive surface charge. However, when dKLA and PEG... 2k When the ratio of -2CHO increased to 1:0.5 and 1:1, the surface charge decreased to negative, indicating that the positively charged amino groups were completely shielded, and PEG crosslinking was successfully formed on the surface of DC NPs. Considering that ideal NDDSs should have high drug loading and minimal excipients, dKLA with PEG... 2k The molar ratio of -2CHO was finally determined to be 1:0.5. The morphology of PDC NPs was observed using TEM, as shown... Figure 10 As shown in Figure c, PDC NPs possess a spherical structure similar to DC NPs, and PEG crosslinking did not affect the morphology of DC NPs. Further calculations revealed that the total drug loading of PDC NPs was 68.3%, with Ce6 and dKLA accounting for 33.6% and 34.7%, respectively. The total drug loading of PDC NPs significantly exceeded that of traditional NDDSs (~15%), indicating that our charged biomolecule delivery system can deliver drugs more efficiently.

[0140] Furthermore, we monitored the stability of DC NPs and PDC NPs using DLS. Figure 10 d and Figure 10 e). DC NPs do not maintain long-term stability in aqueous solution, and their particle size increases slightly; in serum solution, the particle size of DC NPs increases more significantly due to protein adsorption. In contrast, PDC NPs exhibit better stability; even after one week of incubation in serum solution, the particle size did not fluctuate significantly, and the PDI remained within a reasonable range without significant change, indicating that PEG crosslinking significantly improves the stability of nanoparticles. pH changes affect the ionization of charged groups, disrupting the intermolecular interactions of self-assembled molecules and further leading to drug release from charged biomolecule delivery systems. Therefore, we evaluated the drug release properties of PDC NPs in systems at pH 7.4 and pH 5.0 (simulating intracellular lysosomal pH), respectively. Figure 10 As shown in f, PDC NPs are relatively stable at physiological pH, with minimal drug release. In contrast, an acidic environment at pH 5.0 can trigger drug release from PDC NPs. This result indicates that PDC NPs can remain stable at pH 7.4, preventing premature drug leakage under physiological conditions, but controlled drug release can occur within the tumor cell microenvironment.

[0141] Because this invention employs reversible chemical bonds to crosslink the delivery system with PEG, the crosslinked delivery system undergoes changes within the tumor microenvironment, resulting in more efficient drug delivery. Take the Schiff base used in PDC NPs as an example. Schiff bases are sensitive to weak acids; the PEG crosslinking of PDC NPs is stripped in a TME (acidic pH 6.5-6.8), re-exposing the amino groups in dKLA and achieving a "negative to positive" charge reversal at the tumor site. We incubated PDC NPs at pH 6.5 for 24 hours and characterized their size and surface charge. Figure 11 As shown in Figure a, the particle size of PDC NPs incubated for 24 hours in a pH 6.5 environment becomes slightly larger than that of DC NPs and PDC NPs at pH 7.4. (TEM micrograph) Figure 11 (b) shows that the morphology of PDC NPs swells when treated with pH 6.5, consistent with the DLS results. Notably, after treatment with pH 6.5, the surface charge of PDC NPs reverses from negative to positive. Figure 11 c) indicates that weakly acidic pH can separate the PEG layer and reverse the surface charge of PDC NPs, which helps promote the internalization of PDC NPs into cells.

[0142] Charge reversal increases the internalization of PDC NPs: "Negative to positive" charge reversal favors cellular uptake of PDC NPs. Here, 4T1 cells were co-incubated with DC NPs, PDC NPs, and weakly acidic (pH 6.5) pretreated PDC NPs, and cellular uptake was assessed by measuring the intrinsic fluorescence of Ce6. Figure 11 As shown in Figure d, the fluorescence signal of the DC NPs group was stronger than that of the PDC NPs group, indicating that positively charged nanoparticles are more easily taken up by cells. In contrast, the intracellular fluorescence signal of the weakly acid-pretreated PDC NPs group was almost identical to that of the DC NPs group, suggesting that pH-triggered charge reversal significantly improves nanoparticle internalization. This result validates our hypothesis that negatively charged, surface-protected PDC NPs are less susceptible to opsonin effects and accumulate more at tumor sites; the reversed positive charge significantly improves cellular uptake, thereby enhancing drug delivery efficiency to tumor cells.

[0143] Pharmacokinetic Evaluation of PDC NPs in Mice: To demonstrate whether PEG crosslinking can prolong the blood circulation of nanoparticles, we compared the pharmacokinetic properties of the three formulations by administering a single tail vein injection of Ce6, DC NPs, and PDC NPs (10 mg / kg Ce6) to healthy Balb / c mice (n=3). Figure 12a. Free Ce6 was cleared from the blood more quickly, while serum circulation of DC NPs was not effectively improved. This is because positively charged nanoparticles are easily captured by macrophages due to opsonin activity and cleared from the bloodstream, whereas serum circulation of PDC NPs was significantly improved. The area under the curve (AUC) for free Ce6, DC NPs, and PDC NPs is shown in the figures. 0~ꝏ The concentrations of PDC NPs were 22.29±2.64 μg / mL / h, 24.75±1.98 μg / mL / h, and 39.48±1.29 μg / mL / h, respectively, and their AUCs were... 0~ꝏ It is about twice as high as that of free Ce6 and DC NPs. The half-lives (T5) of free Ce6, DC NPs, and PDC NPs are... 1 / 2 The blood circulation times of PDC NPs were 0.85±0.26 h, 1.3±0.19 h, and 1.5±0.2 h, respectively, indicating a longer blood circulation time. This suggests that our PEG crosslinking strategy can effectively improve the blood circulation of the charged biomolecule delivery system, facilitating greater accumulation of the delivery system in tumor tissue and achieving a stronger anti-tumor effect.

[0144] Tissue distribution of PDC NPs in tumor-bearing mice: The intratumoral accumulation and biodistribution of PDC NPs in 4T1 tumor-bearing mice were determined by near-infrared fluorescence (NIRF) imaging. Figure 12 b. Free Ce6 showed tumor accumulation 2 hours after administration, but with low signal intensity, and was subsequently cleared rapidly. DC NPs began to accumulate in the tumor 2 hours after administration, reaching their peak level at 6 hours, and then gradually subsided. In contrast, PDC NPs began to accumulate gradually at the tumor site from 4 hours after administration, reaching their peak signal intensity at the tumor site 8 hours after injection. The corresponding quantitative fluorescence results are as follows: Figure 12 As shown in c, the AUCs of Ce6, DC NPs and PDC NPs were 5370 ± 552, 12638 ± 1100 and 21937 ± 1410, respectively, indicating that the PEG-crosslinked charged biomolecule delivery system can accumulate more at the tumor site than the free molecules and non-crosslinked delivery system.

[0145] Based on in vivo imaging results, mice were sacrificed 6 hours after drug administration, and the in vitro distribution of PDC NPs relative to Ce6 and DC NPs was observed. Figure 12As shown in d and 12e, Ce6 is mainly distributed in the lungs, liver, and kidneys, with free molecules being rapidly excreted via the kidneys. Uncrosslinked nanoparticles accumulate highly in the liver, kidneys, and tumor tissues. DC NPs may prematurely release drugs during blood circulation, which are then excreted via the kidneys. In contrast, PDC NPs accumulate the most at the tumor site, followed by the liver, with less accumulation in the kidneys. This result demonstrates the importance of PEG crosslinking. Furthermore, frozen section results of tumor tissue show that PDC NPs penetrate the tumor site more deeply than free Ce6 and DC NPs. Figure 12 (f and 12g), indicating that PEG-crosslinked charged biomolecule delivery systems can effectively accumulate and penetrate tumor tissues.

[0146] Antitumor therapeutic effect: Based on the good antitumor effect of PDC NPs in vitro, we further investigated the antitumor efficacy of PDC NPs against a subcutaneous 4T1 tumor mouse model. The animal experimental process is attached. Figure 13 As shown in a, tumor-bearing mice were randomly divided into 7 groups (n=6). When the tumor size reached ~50 mm, the mice were divided into 7 groups. 3 Different formulations were used for treatment. All formulations were administered intravenously every three days, for a total of five doses. Based on biodistribution results, the laser treatment group received a 680 nm laser (0.5 W / cm²) 8 hours after administration. 2 Irradiation lasted 5 minutes. Mouse body weight and tumor size were monitored twice weekly. The seven groups were recorded as follows: PBS, a mixture of free Ce6-dKLA with or without laser treatment (Mix / Mix+L), DC NPs with and without laser treatment (DC NPs / DCNPs+L), and PDC NPs with and without laser treatment (PDC NPs / PDC NPs+L). Tumor growth curves are shown below. Figure 13 As shown in b, compared with the PBS group, tumor growth was moderately inhibited in the Mix, DC NPs, and PDC NPs groups, with the DC NPs and PDC NPs groups showing better therapeutic effects than the Mix group. All laser treatment groups showed better therapeutic effects than the non-laser treatment groups, with the PDC NPs+L group showing significantly better tumor inhibition than the Mix+L or DC NPs+L treatment groups, indicating that the better delivery efficiency of PDC NPs can broadly improve anti-tumor efficacy. Tumors were collected at the experimental endpoint for weighing and histopathological examination. The PDC NP+L group showed the greatest inhibition of tumor growth among all groups, exhibiting the smallest tumor size and weight (0.25g). Figure 13 c and Figure 13 d). Weight monitoring of mice in each treatment group as follows: Figure 13As shown in Figure e, compared to the PBS group, the weight gain of mice in the nanoparticle group was significantly affected. We further examined the tumor histopathology; the tumor tissue sections from the PDC NPs+L group showed the most porous nuclei and a significantly decreased nucleus-to-cytoplasm ratio, further demonstrating that PDC NPs combined with laser therapy can effectively inhibit tumor growth. The optimal antitumor effect of PDC NPs indicates that the PEG cross-linking strategy and charge reversal can broadly improve drug delivery efficiency, while the synergistic effect of phototherapy and pro-apoptotic peptides can significantly inhibit tumor growth.

[0147] Biosafety Evaluation: During the treatment trial, we evaluated the tolerability and biosafety of the PDC NPs+L combination therapy. Mice in the Mix treatment group experienced a slight decrease in body weight during treatment, which may be due to toxicity caused by the free peptides and photosensitizer. Mice in the nanoparticle group showed a similar trend of weight gain as the PBS group, indicating that the charged biomolecule delivery system can reduce the systemic toxicity of the drug (see appendix). Figure 13 e). Blood samples were collected from mice at the experimental endpoint for hematological and biochemical analysis. (e.g.) Figure 14 As shown in Figure a, the white blood cell count (WBC) of mice in all treatment groups was higher than that of healthy mice. Among the treatment groups, PDC NPs+L had the least impact on WBC, indicating that combined PDC NPs treatment can reduce systemic inflammation in tumor-bearing mice. Except for WBC, other hematological and biochemical parameters in all treatment groups fluctuated within the normal range compared to healthy mice. Further histopathological examination of major organs (heart, lungs, liver, spleen, and kidneys) was performed. The results showed that ( Figure 14 (b) In all treatment groups, the tissue cells of the mice were neatly arranged and the cellular structure was intact, with no obvious abnormalities or inflammatory infiltration even in the metastatic organs. These results indicate that the charged biomolecule delivery system has good biosafety and can serve as a potential photodynamic nanoparticle formulation.

[0148] The above content is only for illustrating the technical concept of the present invention and should not be construed as limiting the scope of protection of the present invention. Any modifications made to the technical solution based on the technical concept proposed in this invention shall fall within the scope of protection of the claims of this invention.

Claims

1. A method for preparing a charged biomolecule delivery system, characterized in that, Includes the following steps: S1: The charged large π conjugated directing molecule and the charged biomolecule are premixed in an organic solvent to obtain a premixed working solution; S2: After mixing the premixed working solution and deionized water, the mixture is stirred and evaporated, and a nanoscale delivery system is obtained by self-assembly under the drive of non-covalent forces. S3: Mix polyethylene glycol and a nanoscale delivery system, and perform a surface cross-linking reaction on the nanoscale delivery system to obtain a polyethylene glycol cross-linked biomolecule delivery system, thereby achieving the delivery of charged biomolecules; The charged biomolecules have the opposite charge type to the charged large π-conjugated guide molecules; In S1, the charged large π-conjugation guiding molecule is dihydroporphyrin, pheophytin, tetraaminophenylporphyrin, or tetracarboxyphenylporphyrin; the charged biomolecule is a polypeptide drug, small interfering RNA, messenger RNA, hyaluronic acid, or chitosan; the molar ratio of the charged large π-conjugation guiding molecule to the charged biomolecule is (0.5~100):1; the volume ratio of the charged biomolecule to the organic solvent is (0.1~50) mg / mL; The polypeptide drugs mentioned above are apoptosis-promoting polypeptides, transmembrane bee venom peptides, exenatide, liraglutide, teriparatide, octreotide, or eptifibatide. In S3, the surface of the nanoscale delivery system has amino or carboxyl groups; When the surface of the nanoscale delivery system has amino groups, the polyethylene glycol has aldehyde, hydrazide or carboxyl groups. In this case, the nanoscale delivery system and polyethylene glycol are mixed at an equivalent ratio of amino to aldehyde, hydrazide or carboxyl groups of 1: (1~2). When the surface of the nanoscale delivery system has carboxyl groups, the polyethylene glycol has amino or hydroxyl groups. In this case, the nanoscale delivery system and polyethylene glycol are mixed at an equivalent ratio of carboxyl groups to amino or hydroxyl groups of 1:(1~2).

2. The method for preparing a charged biomolecule delivery system according to claim 1, characterized in that, In S1, the organic solvent is tetrahydrofuran or methanol.

3. The method for preparing a charged biomolecule delivery system according to claim 1, characterized in that, In S2, the premixed working solution and deionized water are mixed by stirring at room temperature; the volume ratio of the premixed working solution to deionized water is 10%~50%.

4. The method for preparing a charged biomolecule delivery system according to claim 1, characterized in that, In S3, the surface crosslinking reaction is carried out at room temperature for 48 hours.

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