Ph-controlled biodegradable polymeric particles for sustained release of positively charged species

Abstract

The disclosure relates to microparticles and nanoparticles comprising a porous polymer matrix comprising an uncapped polymer for sustained delivery of a net positively charged therapeutic agent. More particularly the disclosure relates to particles comprising PLGA or PLA which have a first state with relatively more interconnected pores at a first pH and a second state with relatively less interconnected pores at a second pH. Methods of making the particles and administering the particles are also provided.

Description

PH-CONTROLLED BIODEGRADABLE POLYMERIC PARTICLES FOR SUSTAINED RELEASE OF POSITIVELY CHARGED SPECIESSTATEMENT OF GOVERNMENT SUPPORT

[0001] This invention was made with government support under DK127817 awarded by the National Institutes of Health. The government has certain rights in the invention.FIELD

[0002] The present technology relates to pH-controlled biodegradable polymeric particles for sustained release of a positively charged species. More particularly, the disclosure relates to pH-controlled biodegradable polymeric particles for sustained, controlled release of a therapeutic agent, such as, a peptide or protein which has a net positive charge.BACKGROUND

[0003] Injectable, biodegradable polymeric particles, such as nanoparticles and / or microparticles, provide a means to deliver and control the release of therapeutic agents such as small molecule drugs, proteins, peptides, and antigens (which can be classified as species of proteins or peptides). Once injected, the biodegradable polymeric particles can release the therapeutic agent over the course of hours, days, or more extended periods such as weeks or months, thus eliminating the need for daily injections, and thereby improving patient acceptance and compliance as well as outcomes. Controlled release of a therapeutic agent can therefore beneficially reduce the number of injectable doses for the therapeutic agent in an immunization schedule.

[0004] Nevertheless, significant obstacles have been encountered with development of injectable, biodegradable polymeric particles for controlled release. Two significant concerns are the stability of therapeutic agents during encapsulation and high manufacturing costs. Methods for encapsulating therapeutic agents in biodegradable polymers can involve harsh processing conditions, including exposure to organic solvents, high temperatures, homogenization methods such as mixing, sonication, and high-speed agitation, and require aseptic processing. These methods alone or in combination can destabilize therapeutic agents, particularly proteins and peptides, as well as lower loading efficiencies for these therapeutic agents. Micronization of a therapeutic agent prior to encapsulation can further destabilize the therapeutic agent. Finally, conventional methods of encapsulation commonly involve using high concentrations of drug dissolved in water or acetic acid, which can either be destabilizing to the drug or not possible when drug concentrations are above the drug solubility.

[0005] PLGA, also referred to as poly(lactic-co-glycolic acid) and poly(lactide-co- glycolide), is the most commonly used polymer for the preparation of injectable, biodegradable polymeric particles. PLGA can be manufactured by polycondensation or ringopening polymerization. The co-polymer ratio of lactic (L) to glycolic (G) or lactide (L) to glycolide (G) can be controlled precisely from greater than 0 to 100%. PLGA can be formed as semi crystalline, or as fully amorphous. PLGA can be made with differing blockiness and block length of the L or G monomers and can be made with different end groups on the carboxyl groups (e.g., typically either aliphatic ester end-capped or free acid end-capped). PLGA can be made as a star polymer when polyalcohol compounds such as glucose are used to initiate ring opening polymerization.

[0006] Conventional microencapsulation approaches to manufacture PLGA microspheres include solvent evaporation, coacervation, and spray-drying. In each of these methods, the API is combined with PLGA dissolved in an organic solvent before forming microspheres. This combination creates several undesirable issues: (a) the peptide-loaded microspheres most often cannot be terminally sterilized, thus requiring expensive aseptic processing with organic solvents and numerous unit operations; (b) yields are often low, which is particularly problematic when the API is expensive; (c) products can include one or more residual organic solvents, which pose challenges to storage stability of the final products (van de Weert et. al. Pharmaceutical Research, 17 (2000) 1159-1167); (d) there is little opportunity to manipulate the polymer structure once the peptide-PLGA matrix is formed, limiting the ability to engineer release kinetics (Hines, D. J. et al. Grit. Rev. Therapeutic Drug Carr. Syst. 30 (2013) 257-276; Fu, Y. et al. Exp. Opin. Drug Del. 7 (2010) 429-444); and (e) mixing organic solvent / water mixtures in the presence of peptides, or other forms of micronization, particularly with higher-order structure, can be detrimental to drug stability (Schwendeman, S. P. et al. J. Control Release 190 (2014) 240-253).

[0007] The concept of remote loading in aqueous solution was initially shown by encapsulating large molecules including leuprolide and large proteins in porous aliphatic ester end capped PLGAs, where pores are closed spontaneously by passive healing of the polymer with elevated temperature (Reinhold et. al., Angewandte Chemie Int. Ed., 51 (2012) 10800-10803; US Patent 8,017,155). However, this technique generally requires a “trapping agent” such as aluminum hydroxide adjuvant or dextran sulfate inside the polymer pores to bind the therapeutic agent to overcome low encapsulation efficiency (Reinhold et. al., Angewandte Chemie Int. Ed., 51 (2012) 10800-10803). Additionally, incorporation of a water-insoluble base such as MgCOa is often required to raise internal pH of the particle or facilitate continuous drug release (Schwendeman, Recent Advances in the Stabilization of Proteins Encapsulated in Injectable PLGA Delivery Systems, 19 (2002) 26).

[0008] Previous studies on uncapped PLGA and net positively charged therapeutic agents demonstrated that the negatively charged acid end-group of PLGA (after deprotonation) and positively charged therapeutic agents can interact in aqueous solution (Sophocleous et. al. J. Controlled Release, 172, 662-670 (2013)). However, leuprolide which has a ~+1 charge at neutral pH was absorbed more readily than octreotide, which has a -+1.7 charge at neutral pH. Based on these results, the additional charge on octreotide was expected to cause the therapeutic agent to associate with more than one polymer chain in PLGA and thus decrease the number of binding sites available for the peptide and consequently reduce encapsulation efficiency and loading in PLGA particles, at least relative to a less highly charged therapeutic agent such as leuprolide.SUMMARY

[0009] One aspect of the present invention provides a method of preparing a biodegradable particle for sustained release of a therapeutic agent by providing one or more biodegradable particles comprising a polymer matrix, wherein the polymer matrix comprises an uncapped polymer, and incubating the one or more biodegradable particles with a therapeutic agent in a loading solution at a first pH between about 6.8 and about 8.0, then incubating the one or more biodegradable particles in a loading solution at a second pH between about 4.5 and about 6.5, wherein the uncapped polymer comprises free carboxyl groups such that the biodegradable particle includes free carboxyl groups, wherein the particle has an average particle diameter in the range of about 10 nm to about 100 pm, and wherein the therapeutic agent comprises a protein or peptide with a net positive charge in the loading solution at the first and second pH of the loading solution.

[0010] In another aspect, the invention provides a formulation for sustained release of biodegradable particles containing a therapeutic agent comprising a biodegradable particle comprising a polymer matrix, wherein the polymer matrix comprises an uncapped polymer, a therapeutic agent absorbed and encapsulated by the polymer matrix, and a solution with a pH of about 4.5 to about 6.5, wherein the biodegradable particle has an average particle diameter in the range of about 10 nm to about 100 pm, and wherein the therapeutic agent comprises a protein or peptide with a net positive charge in the solution.

[0011] Another aspect of the present invention provides a method of preparing a biodegradable particle for sustained release of a therapeutic agent by providing one or more biodegradable particles comprising a porous polymer matrix, wherein the polymer matrix has a first state with more interconnected pores and a second state with less interconnected pores, and incubating the one or more biodegradable particles with a therapeutic agent in a loading solution at a first pH between about 6.8 and about 8.0, thereby encapsulating thetherapeutic agent in the first state of the porous polymer matrix, then incubating the one or more biodegradable particles in a loading solution at a second pH between about 4.5 and about 6.5, thereby transitioning the polymer matrix from the first state to the second state, wherein the polymer matrix comprises an uncapped polymer comprising free carboxyl groups such that the biodegradable particle includes free carboxyl groups, wherein the therapeutic agent comprises a protein or peptide with a net positive charge in the loading solution at the first and second pH of the loading solution, and wherein the biodegradable particles have an initial burst release of the therapeutic agent of about 5% or less after 24 hours when the particles are in the second state.

[0012] In another aspect, the invention provides a formulation for sustained release of biodegradable particles containing a therapeutic agent, comprising a plurality of biodegradable particles, the biodegradable particles comprising a porous polymer matrix, wherein the porous polymer matrix comprises an uncapped polymer comprising PLGA or PLA, a therapeutic agent absorbed and encapsulated by the polymer matrix, and a media capable of providing a pH of about 4.5 to about 6.5 within an unstirred boundary layer of the biodegradable particles upon hydration, wherein the porous polymer matrix has a first state with more interconnected pores and a second state with less interconnected pores, wherein the biodegradable particles can transition from the first state to the second state when exposed to a pH between about 4.5 to about 6.5, wherein the biodegradable particles can transition from the second state to the first state when exposed to a pH between about 6.8 to about 8.0, wherein the biodegradable particles are dispersed in the media, and wherein the therapeutic agent comprises a protein or peptide with a net positive charge at the pH between about 4.5 to about 6.5.

[0013] Further aspects of the disclosure may become apparent to those skilled in the art from a review of the following detailed description, taken in conjunction with the examples and appended claims. While the invention is susceptible to embodiments in various forms, described herein are specific embodiments of the invention with the understanding that the disclosure is illustrative, and is not intended to limit the invention to specific embodiments described herein.BRIEF DESCRIPTION OF THE DRAWINGS

[0014] Figure 1 shows the effect of microparticle porosity on encapsulation efficiency and drug loading.

[0015] Figure 2 shows FT-IR spectra of (a) PLGA polymer, (b) microparticles remote loaded with setmelanotide, and (c) setmelanotide.

[0016] Figure 3 shows powder X-ray diffraction patterns of (a) PLGA polymer, (b) microparticles remote loaded with setmelanotide, and (c) setmelanotide.

[0017] Figure 4 shows the volume average diameter of blank microparticles and setmelanotide-loaded microparticles (Mean ± SEM, n = 3).

[0018] Figure 5 shows the particle size distribution (span) of blank microparticles and setmelanotide-loaded microparticles (Mean ± SEM, n = 3).

[0019] Figure 6 shows the final pH after remote loading of microparticles with setmelanotide in different loading solutions and different loading times.

[0020] Figure 7 shows the kinetics of encapsulation efficiency overlayed with the pH change during remote loading of PLGA microparticles with setmelanotide over 48 h (Mean ± SEM, n=2 or 3 for each time point).

[0021] Figure 8 shows the titration of 0.1 M HEPES buffer solution with acetic acid at room temperature.

[0022] Figure 9 shows the in vitro release kinetics in PBST at 37 °C of setmelanotide from remote-loaded PLGA microparticles prepared with different porosity (mean ± SEM, n=3).

[0023] Figure 10 shows the burst release of the low porosity microparticles remote loaded with leuprolide in solutions at pH values of 7.4, 6.5, and 5.5 over 2 h.

[0024] Figure 11 shows the burst release of the medium porosity microparticles remote loaded with leuprolide in solutions at pH values of 7.4, 6.5, and 5.5 over 2 h.

[0025] Figure 12 shows the burst release of the high porosity microparticles remote loaded with leuprolide in solutions at pH values of 7.4, 6.5, and 5.5 over 2 h.

[0026] Figure 13 shows the pharmacokinetics of setmelanotide after a single subcutaneous administration of setmelanotide remote-loaded PLGA microparticles to mice fed a high-fat-diet. Dose of setmelanotide was 40 mg / kg (mean ± SEM, n = 5 or 6 for each time point).

[0027] Figure 14 shows the pharmacokinetic profile of setmelanotide after intravenous administration at a dose of 15mg / kg with the best fitted line (mean ± SD, n=4).

[0028] Figure 15 shows the cumulative setmelanotide absorbed as estimated by deconvolution of the pharmacokinetic profile according to the Loo-Riegelman method (mean plasma levels used only) overlaid with the in vitro release profile of the high-porosity formulation of Figure 9 (time scaled by levy plot). The inset figure shows the linear correlation between in vitro release and in vivo absorption.

[0029] Figure 16 shows the effect of diet-induced obese mice administered a single subcutaneous injection of setmelanotide-loaded microparticles (day 0) compared to control, obese mice injected with drug-free microspheres until day 24 (mean ± SEM, n = 6).

[0030] Figure 17 shows the food intake on a high fat diet (HFD, 60%) of mice with setmelanotide-loaded microparticles compared with the drug-free microparticles control group. Each symbol represents a single mouse’s food intake data.

[0031] Figure 18 shows the improvement of body weight control over the vehicle-treated group (overlaid with the in vivo PK profile of setmelanotide-loaded microspheres of Figure 13). The value for each symbol was calculated by subtracting the average percentage weight change in setmelanotide-loaded microparticles group from the average percentage weight change in vehicle-treated controls (A control - A Setmelanotide).

[0032] Figure 19 shows the drug loading and encapsulation efficiency percentages of microparticle formulations remote loaded with setmelanotide in solutions with varying HEPES buffer concentrations.

[0033] Figure 20 shows the cumulative drug release from microparticle formulations remote loaded with setmelanotide in solutions with varying HEPES buffer concentrations vs. the pH drop as measured after loading.

[0034] Figure 21 shows drug loading percentages of microparticle formulations remote loaded with setmelanotide (n = 2, Mean ± SD).

[0035] Figure 22 shows the in vitro release kinetics in PBST at 37 °C of the setmelanotide remote-loaded PLGA microparticle formulations (n = 3, Mean ± SD).

[0036] Figure 23(a) shows SEM images of a specific setmelanotide remote-loaded PLGA microparticle formulation before remote loading (left) and after remote loading (right).

[0037] Figure 23(b) shows SEM images of a further specific setmelanotide remote-loaded PLGA microparticle formulation before remote loading (left) and after remote loading (right).

[0038] Figure 23(c) shows SEM images of an additional specific setmelanotide remote- loaded PLGA microparticle formulation before remote loading (left) and after remote loading (right).

[0039] Figure 24 shows the burst release of microparticles remote loaded with leuprolide after incubation in 15% hydrogel buffered at pH values of 7.4, 6.5, and 5.5 over 2 h (PBST pH 7.4 was used as control).

[0040] Figure 25 shows the burst release of microparticles remote loaded with leuprolide after incubation in 15% hydrogel buffered at pH values of 7.4, 6.5, and 5.5 over 4 h (PBST pH 7.4 was used as control).

[0041] Figure 26 shows the burst release of microparticles remote loaded with leuprolide after incubation in 18% hydrogel buffered at pH values of 7.4, 6.5, and 5.5 over 2 h (PBST pH 7.4 was used as control).

[0042] Figure 27 shows the burst release of microparticles remote loaded with leuprolide after incubation in 18% hydrogel buffered at pH values of 7.4, 6.5, and 5.5 over 4 h (PBST pH 7.4 was used as control).

[0043] Figure 28 shows the burst release of microspheres remote loaded with leuprolide after incubation in 15% hydrogel buffered at pH values of 7.4, 6.5, and 5.5 over 4 h. Formulations with hydrogel and free leuprolide were used as control (Free drug at Gel).DETAILED DESCRIPTION

[0044] The present invention discloses methods for encapsulating therapeutic agents, in biodegradable polymeric particles, such as microparticles and nanoparticles, as well as formulations for sustained release comprising such biodegradable polymeric particles. The disclosed methods and formulations advantageously provide control over the loading and / or release profile of the biodegradable polymeric particles by specifically controlling the pH of the environment surrounding the particles. By controlling the (bulk) pH of a loading solution at a first pH and / or the (bulk) pH of an incubation solution at a second pH as described herein, or the (bulk) pH of a formulation of the polymeric particles, or by providing the particles in a media capable of providing a pH of about 4.5 to about 6.5 within an unstirred boundary layer of the particles upon hydration, the biodegradable polymeric particles advantageously adopt a preferred state for either encapsulation or release, and even be caused to switch between a first state and a second state, respectively. The first state of the biodegradable particle has relatively more interconnected pores and channels than the second state of the particles. The relatively more interconnected pores and channels beneficially enables the first state of the biodegradable particles to enhance loading and encapsulation of therapeutic agents, preferably proteins or peptides with net positive charges, at the first pH. The second state of the particle has relatively less interconnected pores and channels than the first state of the particles. The relatively less interconnected pores and channels enable the second state of the particles to provide enhanced, sustained release of therapeutic agents, specifically by preventing an undesirable initial burst release of more than 5% of the therapeutic agent during the first 24 hours after administration I exposure to physiological conditions, as determined by the amount of therapeutic agentreleased from t = 0 hours to t = 24 hours in a phosphate-buffered saline buffer (pH = 7.4). Thus, the disclosed formulations advantageously provide precise control over the release profile of the loaded particles. Furthermore, by using the aqueous-based methods described herein, exposure of the therapeutic agent to harsh solvents like methylene chloride or to a micronization step can advantageously be avoided, thereby avoiding destabilization of therapeutic agents and lowering costs associated with manufacture. Loading of the biodegradable polymeric particles as disclosed herein also beneficially provides high encapsulation efficiency (e.g., >60%) as well as particles with high loading of the therapeutic agent (e.g., >5% w / w), and is particularly advantageous for use with molecules having net positive charges greater than +1 at pH values that are about neutral or greater (for example, at pH values between about 6.8 and 8.0 at which the loading step in the disclosed methods is conducted).

[0045] As used herein, the term “pH of a solution” refers to the pH of a bulk solution and can be measured using a conventional pH meter.

[0046] The term “about” is used according to its ordinary meaning, for example, to mean approximately or around. In one embodiment, the term “about” means ±10% of a stated value or range of values. In another embodiment, the term “about” means ±5% of a stated value or range of values. A value or range described in combination with the term “about” expressly includes the specific value and / or range as well (e.g., for a value described as “about 40,” “40” is also expressly contemplated).Composition of Polymeric Particles

[0047] Natural and synthetic polymers such as poly(lactide)s, poly(glycolide)s, poly(lactide-co-glycolide)s, poly(lactic acid)s, poly(glycolic acid)s, poly(lactic acid-co-glycolic acid)s, homopolymers, copolymers, and blends of these and other polymers may be used to form bidegradable polymer matrices. In the particles according to the disclosure, uncapped natural and synthetic polymers are used to provide the polymer matrix in the particles disclosed herein. The particles of the disclosure feature a polymer matrix, typically comprised of uncapped PLGA, PLA, or a combination thereof. In some embodiments, the polymer matrix comprises an uncapped PLGA with a lactic acid content in the range of 20% to 70%, 30% to 70%, 40% to 70%, or 50% to 70%, relative to the total amount of lactic acid and glycolic acid in the uncapped polymer.

[0048] PLGA and PLA based polymers feature carboxyl groups at the end of the polymer. As used herein, the term “uncapped polymer” refers to a polymer with “free” carboxyl groups at the end of the polymer. Thus, these free carboxyl groups may include acidic protons on the carboxyl groups under relatively acidic conditions, generally at about pH 5.5 or less, orbe present in substantially ionized form and thus be at least partially negatively charged at higher pH values. Therefore, the carboxyl groups of the uncapped polymer are such that the polymeric particles feature carboxyl groups which are negatively charged at the first pH between about 6.8 and 8.0 and thus are able to associate with positively charged moieties of a therapeutic agent and thereby facilitate absorption of the therapeutic agent into the polymer matrix when placed in an aqueous solution having a pH between about 6.8 and 8.0, at a temperature above room temperature, for example, about 30 °C, or about physiological temperature (37 °C), or about the polymer’s glass transition temperature (40 °C to 55 °C). Generally, the polymeric particles are placed in solutions about 37 °C such that after absorption of the therapeutic agent, the mobility of the polymer matrix is increased to facilitate the sealing of the particle surface when exposed to a solution at the second pH, between about 4.5 and 6.5. These carboxyl groups available to associate with the positively charged moieties of the therapeutic agent are herein referred to as “free carboxyl groups.” In various cases, the uncapped polymer has a weight average molecular weight in the range of about 2kDa to about 50 kDa, for example, about 3 kDa to about 45 kDa, about 5 kDa to about 40 kDa, about 7.5 kDa to about 35 kDa, and / or about 10 kDa to 20 kDa. In additional embodiments, the uncapped polymer has a molecular weight in the range of about 2 kDa to about 20 kDa, about 5 kDa to about 17 kDa, about 5 kDa to about 15 kDa, about 5 kDa to about 10 kDa, about 7 kDa to about 17 kDa, and / or about 8 kDa to about 12 kDa. The weight average molecular weight of the uncapped polymer can be determined using methods known in the art, particularly gel permeation chromatography. Thus, the uncapped polymer can have a lactic acid content in the range of 50% to 70%, relative to the total amount of lactic acid and glycolic acid in the polymer matrix and a molecular weight in the range of about 2 kDa to about 20 kDa, about 5 kDa to about 17 kDa, about 5 kDa to about 15 kDa, about 5 kDa to about 10 kDa, about 7 kDa to about 17 kDa, and / or about 8 kDa to about 12 kDa. In some embodiments, the uncapped polymer has a molecular weight in the range of about 5 kDa to about 10 kDa. In some embodiments, the uncapped polymer has a molecular weight less than 10 kDa.

[0049] Additionally, the particles disclosed herein generally are substantially free of “capped polymer” which refers to a polymer in which the carboxyl groups have been substituted or replaced with other functional groups (i.e. , not protons). Typically, in a capped polymer, the carboxyl groups at the end of the PLGA or PLA are replaced with hydrophobic groups, such as ester groups.

[0050] As used herein, the term “substantially free” means that the compositions and / or particles according to the disclosure contain insignificant amounts of the indicated component. For example, the particles according to the disclosure may contain less than 5weight percent, less 2 wt.%, less than 1 wt.%, or less than 0.10 wt.% of the indicated component, based on the entire weight of the composition or particle.

[0051] Beneficially, the particles may be substantially free of a “trapping agent” such as aluminum hydroxide adjuvant or dextran sulfate and / or of a water-insoluble base such as magnesium carbonate. Such trapping agents and water-insoluble bases can be disadvantageous, for example, by creating additional complexity in formulations comprising the particles and / or potentially increasing inflammation at the injection site.Therapeutic Agents

[0052] Disclosed herein are methods for loading biodegradable polymeric particles with net positively charged therapeutic agents, as well as formulations for sustained release of biodegradable particles containing such therapeutic agents. As used herein, the term “therapeutic agent” refers to a protein, peptide, or small molecule drug. In many cases, the therapeutic agent is a protein or a peptide, particularly a positively charged protein or a positively charged peptide which is at least partially positively charged at the first pH between about 6.8 and 8.0. In preferred embodiments, the therapeutic agent is an antigen. In preferred embodiments, the therapeutic agent is a self-antigen.

[0053] The biodegradable polymeric particles have carboxyl groups on their surface and throughout the pores and channels of the particle and these free carboxyl groups which are at least partially negatively charged at pH values greater than 5 can associate with the positively charged moieties of the therapeutic agent, such that adsorption into the polymer matrix advantageously occurs in an aqueous loading solution at a pH between about 6.8 and 8.0. In general, such net positively charged therapeutic agents are preferred because the positively charged moieties of these net positively charged therapeutic agents can associate with free carboxyl groups from the ends of the uncapped polymer throughout the particle and / or on the surface of the biodegradable polymeric particle, thereby facilitating encapsulation and loading of the therapeutic agent.

[0054] Thus, the association between the positively charged moieties on the net positively charged therapeutic agent and the carboxyl groups of the ends polymer advantageously facilitates initial absorption of the therapeutic agent into the biodegradable polymeric particle in an aqueous loading solution at a pH between about 6.8 and 8.0. However, as noted above, it was expected that therapeutic agents with a net positive charge greater than about +1 would decrease the amount of the carboxyl groups on the surface and / or throughout the biodegradable polymeric particle, upon binding with the polymer. Thus, it was expected that this interaction of the multivalent cationic agent with more than one polymer chain would effectively decrease the number of binding sites available for further therapeutic agent bindingand consequently reduce encapsulation efficiency and loading. However, it was surprisingly found that controlling the pH of the loading solution as disclosed herein advantageously enhanced loading (i.e., therapeutic agent content) as well as encapsulation efficiency. In this respect, while positively charged moieties can improve association of the therapeutic agent to the polymer, as the net charge of the therapeutic agent becomes more positive than +1 , for example, such as +1.5, +2, or greater, because additional binding sites on the particle are believed to be involved, loading and encapsulation efficiency were generally expected to decrease, as mentioned above. However, despite requiring more binding sites on the particle, loading and encapsulation were surprisingly increased in the particles according to the invention when one or more of the particles are incubated with a therapeutic agent in a loading solution at a first pH between about 6.8 and about 8.0 and the net charge of the therapeutic agent is greater than +1 , for example, +1.5, +2, or greater.

[0055] Generally, the therapeutic agents of the disclosure have a net positive charge at about neutral pH (and thus at the first pH between about 6.8 and 8.0). In preferred embodiments, the net positively charged therapeutic agent features one or more positively charged moieties at neutral pH such that the net charge of the therapeutic agent is greater than or equal to about +1 at about neutral pH (and thus at the first pH between about 6.8 and 8.0). For example, the therapeutic agent can have a net charge greater than or equal to about +1.5, a net charge greater than or equal to about +1.7, and / or a net charge greater than or equal to about +2, at about neutral pH (and thus at the first pH between about 6.8 and 8.0).

[0056] In methods of the disclosure, therapeutic agents with net positive charges greater than about +1.5 or +2 are preferred. Without being bound by theory, therapeutic agents with net positive charges greater than about +1.5 or +2 can associate with the free carboxyl groups of the polymer more effectively than therapeutic agents with lesser net positive charges. Without intending to be bound by theory, therapeutic agents with net positive charges greater than about +1.5 or +2 are adsorbed into the polymer matrix more effectively relative to therapeutic agents with net positive charges of +1 because of a chelation effect resulting from association of the therapeutic agent with multiple negatively charged free carboxyl groups, thereby enhancing loading and encapsulation efficiency of the particles.

[0057] In embodiments, the net positively charged therapeutic agent is a protein or peptide containing more than one positively charged moiety at about neutral pH. In embodiments, the positively charged moieties are positively charged amino acid residues at neutral pH. For example, the peptides MOG 38-50, NRPA7, setmelanotide, and leuprolide all feature amino acid residues that are positively charged at neutral pH (pH = 7). In some cases, the positively charged amino acid residue includes one or more positively charged amino acid residue chosen from one or more in the group of positively charged lysine,positively charged arginine, and positively charged histidine residues. In some cases, the positively charged moiety is chosen from one or more in the group of guanidinium, ammonium, and imidazolium.

[0058] Generally, the therapeutic agents of the disclosure can be loaded into particles based on the size of the particle and the target of the therapeutic agent. The size of the particles can be selected based on the physiology of the subject of treatment and the target of the therapeutic agent. The nanoparticles of the disclosure can advantageously disperse throughout the body to provide sustained release of therapeutic agents to places in the body that microparticles cannot disperse.

[0059] In some cases, the therapeutic agent is MOG 38-50, a self-antigen peptide having the sequence GWYRSPFSRWHL (SEQ ID NO:1). In some cases, the therapeutic agent is a self-antigen peptide having the sequence NRPA7, KYNKANAFL (SEQ ID NO: 2). As used herein, the term “antigen” refers to a molecule capable of generating an immune response from a subject. As used herein, the term “self-antigen” refers to a protein or peptide which does not act as an antigen in a healthy subject, but is capable of generating an immune response in a subject with an autoimmune condition or disease. As used herein, the term “autoimmune condition or disease” refers to a disease or disorder that interferes with the proper functioning of the immune system, particularly when the immune cells in a subject attack its own healthy cells. It can be chronic pathology triggered by the loss of immunological tolerance to peptides, which can cause systemic or organ specific damage. In some instance, autoimmune response is mediated by autoreactive T and B lymphocytes responsible for the production of soluble mediators (e.g., cytokines, nitric oxide, etc.) and autoantibodies. Infections can be a cause of the autoimmune disease or disorder.

[0060] In some embodiments, an autoimmune disease or disorder can include but are not limited to Achalasia, Addison’s disease, Adult Still's disease, Agammaglobulinemia, Alopecia areata, Amyloidosis, Ankylosing spondylitis, Anti-GBM / Anti-TBM nephritism Antiphospholipid syndrome, Autoimmune angioedema, Autoimmune dysautonomia, Autoimmune encephalomyelitis, Autoimmune hepatitis, Autoimmune inner ear disease (Al ED), Autoimmune myocarditis, Autoimmune oophoritis, Autoimmune orchitis, Autoimmune pancreatitis, Autoimmune retinopathy, Autoimmune urticarial, Axonal & neuronal neuropathy (AMAN), Balo disease, Behcet’s disease, Benign mucosal pemphigoid, Bullous pemphigoid, Castleman disease (CD), Celiac disease, Chagas disease, Chronic inflammatory demyelinating polyneuropathy (Cl DP), Chronic recurrent multifocal osteomyelitis (CRMO), Churg-Strauss Syndrome (CSS) or Eosinophilic Granulomatosis (EGPA), Cicatricial pemphigoid, Cogan’s syndrome, Cold agglutinin disease, Congenital heart block, Coxsackie myocarditis, CREST syndrome, Crohn’s disease, Dermatitis herpetiformis, Dermatomyositis,Devic’s disease (neuromyelitis optica), Discoid lupus, Dressier’s syndrome, Endometriosis, Eosinophilic esophagitis (EoE), Eosinophilic fasciitis, Erythema nodosum, Essential mixed cryoglobulinemia, Evans syndrome, Fibromyalgia, Fibrosing alveolitis, Giant cell arteritis (temporal arteritis), Giant, cell myocarditis, Glomerulonephritis, Goodpasture’s syndrome, Granulomatosis with Polyangiitis, Graves’ disease, Guillain-Barre syndrome, Hashimoto’s thyroiditis, Hemolytic anemia, Henoch-Schonlein purpura (HSP), Herpes gestationis or pemphigoid gestationis (PG), Hidradenitis Suppurativa (HS) (Acne Inversa), Hypogammaglobulinemia, IgA Nephropathy, lgG4-related sclerosing disease, Immune thrombocytopenic purpura (ITP), Inclusion body myositis (IBM), Interstitial cystitis (IC), Juvenile arthritis, Juvenile diabetes (Type 1 diabetes), Juvenile myositis (JM), Kawasaki disease, Lambert-Eaton syndrome, Leukocytoclastic vasculitis, Lichen planus, Lichen sclerosus, Ligneous conjunctivitis, Linear IgA disease (LAD), Lupus, Lyme disease chronic, Meniere’s disease, Microscopic polyangiitis (MPA), Mixed connective tissue disease (MCTD), Mooren’s ulcer, Mucha-Habermann disease, Multifocal Motor Neuropathy (MMN) or MMNCB, Multiple sclerosis (MS), Myasthenia gravis, Myositis, Narcolepsy, Neonatal Lupus, Neuromyelitis optica, Neutropenia, Ocular cicatricial pemphigoid, Optic neuritis, Palindromic rheumatism (PR), PANDAS, Paraneoplastic cerebellar degeneration (PCD), Paroxysmal nocturnal hemoglobinuria (PNH), Parry Romberg syndrome, Pars planitis (peripheral uveitis), Parsonage-Turner syndrome, Pemphigus, Peripheral neuropathy, Perivenous encephalomyelitis, Pernicious anemia (PA), POEMS syndrome, Polyarteritis nodosa, Polyglandular syndromes type I, II, III, Polymyalgia rheumatic, Polymyositis, Postmyocardial infarction syndrome, Postpericardiotomy syndrome, Primary biliary cirrhosis, Primary sclerosing cholangitis, Progesterone dermatitis, Psoriasis, Psoriatic arthritis, Pure red cell aplasia (PROA), Pyoderma gangrenosum, Raynaud’s phenomenon, Reactive Arthritis, Reflex sympathetic dystrophy, Relapsing polychondritis, Restless legs syndrome (RLS), Retroperitoneal fibrosis,, Rheumatic fever, Rheumatoid arthritis, Sarcoidosis, Schmidt syndrome, Scleritis, Scleroderma, Sjogren’s syndrome, Sperm & testicular autoimmunity, Stiff person syndrome (SPS), Subacute bacterial endocarditis (SBE), Susac’s syndrome, Sympathetic ophthalmia (SO), Takayasu’s arteritis, Temporal arteritis / Giant cell arteritis, Thrombocytopenic purpura (TTP), Tolosa-Hunt syndrome (THS), Transverse myelitis, Type 1 diabetes, Ulcerative colitis (UC), Undifferentiated connective tissue disease (UCTD), Uveitis, Vasculitis, Vitiligo, and Vogt-Koyanagi-Harada Disease..

[0061] Peptides such as MOG 38-50 and NRPA7 have the potential to provide treatment strategies for autoimmune diseases and conditions that specifically work to reduce T cell recognition of peptides through tolerance. T cell tolerance is developed through sustained exposure to peptides that generate T cell responses. Typically, this requires regimentedinoculations to maintain exposure. Such regimented inoculations can advantageously be minimized using the particles and formulations disclosed herein, as the particles of the disclosure, when provided in the second state, can provide enhanced sustained release of peptides such as MOG 38-50 and NRPA7.

[0062] Generally, the molecular weight of the therapeutic agent can be determined by liquid chromatography with mass spectrometry. In various embodiments, the therapeutic agent has a molecular weight in the range of about 500 Da to about 5 kDa, about 1 kDa to about 4.5 kDa, about 1.5 kDa to about 4 kDa, and / or about 2 kDa to about 3.5 kDa.

[0063] In embodiments, the therapeutic agent is a peptide selected from setmelanotide, leuprolide, MOG 38-50, or NRPA7. In embodiments, the peptide is setmelanotide or leuprolide and the particle is a microparticle.

[0064] In another aspect, the biodegradable polymeric particles disclosed herein can be used to deliver fertilizers, biostimulants, and pesticides, particularly by providing microparticles, in the second state, loaded with such active agents, to an acidic media such as acidic soil. The pH of soil is important for plant growth because it can help promote the availability of nutrients to plants. For example, most edibles, grasses, and ornamentals should be cultivated in slightly acidic soils (pH 5.8 to 6.5) whereas azaleas, rhododendrons, blueberries, and conifers should be cultivated in relatively more acidic soils (pH 5.0 to 5.5). pH Controlled Biodegradable Polymeric Particles

[0065] As discussed above, the biodegradable polymeric particles have carboxyl groups on their surface and throughout the pores and channels of the particle. The protonation of these free carboxyl groups on the surface of the biodegradable polymeric particles can be controlled by providing a specific pH range of an aqueous solution surrounding the particle. At the first pH between about 6.8 and 8.0, for example at about pH = 7, the free carboxyl groups on the surface of the biodegradable polymeric particles are deprotonated and are negatively charged. Without being bound by theory, it is understood that these negatively charged carboxyl groups on the surface of the biodegradable polymeric particles repel one another which causes the particle to swell and increase the particle surface area, thereby forming particles in a first state with relatively more interconnected pores. Additionally, it is believed that the negatively charged carboxyl groups on the surface of the particle are generally incompatible with the bulk polymer matrix of the particles, resulting in a more porous structure.

[0066] The first state of the particles advantageously allows for high loading content and encapsulation efficiency of net positively charged therapeutic agents to be achieved by maintaining the first pH while loading the particles. In many embodiments the first pH can bebetween about 6.8 and about 8.0, for example, between about 7.0 and 7.8. Thus, the first pH can be about 6.8, about 6.9, about 7.0, about 7.1 , about 7.2, about 7.3, about 7.4, about 7.5, about 7.6, about 7.7, about 7.8, about 7.9, or about 8.0.

[0067] As the pH around the particles decreases, the negatively charged carboxylate groups on the surface of the biodegradable polymeric particles are protonated and become charge neutral. At a second pH, for example at about pH 5.5, without being bound by theory, it is believed that the charge neutral carboxyl groups on the surface of the biodegradable polymeric particles no longer repel one another. Additionally, these charge neutral carboxyl groups no longer inhibit the movement of the carboxyl groups on the surface of the particle between the surface and the bulk. As a result, the particle surface area decreases, thereby forming particles in the second state with relatively less interconnected pores. Additionally, without intending to be bound by theory, it is believed that the charge neutral carboxyl groups on the surface of the biodegradable polymeric particles are more compatible with the charge neutral bulk of the particle which enhances mobility of the polymer matrix. Therefore, the surface of the particles in the second state are more sealed than the surface of the particles in the first state. Further, the charge neutral carboxyl groups are believed to increase the resistance of the polymer matrix to mass transfer of the peptide i.e., release of the peptide.

[0068] Inducing the second state of the particles allows for loading and encapsulation of therapeutic agents to be truncated / stopped as desired. Advantageously, the second state of the particles can prevent an undesirable initial burst release of the therapeutic agent as well. The second pH can be between about 4.5 and about 6.5, for example, between 5.0 and 6.0. Thus, the second pH can be about 4.5, about 4.6, about 4.7, about 4.8, about 4.9, about 5.0, about 5.1, about 5.2, about 5.3, about 5.4, about 5.5, about 5.6, about 5.7, about 5.8, about5.9, about 6.0, about 6.1, about 6.2, about 6.3, about 6.4, or about 6.5.

[0069] Exposing particles in the first state to the second pH causes the particles to transition to the second state. Likewise, exposing particles in the second state to the first pH causes the particles to transition to the first state. Without being bound by theory, it is believed that a sufficient change in pH is required for the particles to effectively transition between the first state and the second state. The difference between the first pH and the second pH can be between about 1.0 to about 3.0 pH units, for example, between 1.5 and 2.5 pH units. Thus, the difference between the first pH and the second pH can be about 1.0, about 1.1, about 1.2, about 1.3, about 1.4, about 1.5, about 1.6, about 1.7, about 1.8, about1.9, about 2.0, about 2.1, about 2.2, about 2.3, about 2.4, about 2.5, about 2.6, about 2.7, about 2.8, about 2.9, or about 3.0. In embodiments, the difference between the first pH and the second pH is about 1.0 to about 3.0 pH units, about 1.2 to about 2.8 pH units, or about1 .5 to about 2.5 pH units. Generally, the transition of the particle between the first and second states and vice versa, can be confirmed by means well known in the art, including confocal microscopy, for example, by using a fluorescent dye such as 4,4-difluoro-4-bora- 3a,4a-diaza-s-indacene, or measuring airflow, for example, by using a porosimeter. Generally, the first state has an increased relative airflow compared to the second state of at least 20%, at least 30%, at least 40%, and / or at least 50%.

[0070] The particles in the first state display changes in morphology upon exposure to the second pH, accompanying the transition of the particles from the first state to the second state. Without being bound by theory, it is believed that self-healing of the polymer matrix occurs when the particle transitions to the second state. Self-healing generally occurs in polymers in the vicinity of their Tg or above and typically involves multiple physical phenomena such as: a) polymer chain interdiffusion, b) minimization of energetically unfavorable interfacial area, and c) transfer of energy stored in a defect. Further, remote loading of charged therapeutic agents has been previously shown to increase the Tg of the polymer by several degrees C. However, contrary to previous studies, the acid-terminated polymeric particles of the disclosure surprisingly and unexpectedly demonstrate self-healing when the particles are exposed to the second pH, rather than a temperature change. Accordingly, the methods of the disclosure advantageously allow the morphology and loading of polymeric particles to be controlled by providing the particles in solutions or media at the first and / or second pH values associated with the foregoing first and second states. In addition to being advantageous for enhancing release profiles, self-healing of the polymeric particles when provided in a solution at the second pH between about 4.5 to about 6.5 is surprising and unexpected, particularly because acid-terminated PLGAs have not been shown to heal effectively particularly relative to PLGAs prepared with an aliphatic ester endcap (Mazzara, J. M., Thouless, M. D. and Schwendeman, S. P., Healing kinetics of microneedle-formed pores in PLGA films, J. Controlled Release, 171 , 172-177 (2013)).Therefore, significant healing in the acid-terminated polymeric particles, as well as the ability to induce such healing by pH change, was surprising and unexpected.Formation of Biodegradable Polymeric Particles

[0071] The particles described herein are prepared via a method of either single or double oil-water or water-oil-water emulsion of uncapped polymers.

[0072] In some cases, polymeric particles are prepared using a single oil-water (O / W) emulsion of uncapped polymer. First, the uncapped polymer is dissolved in an organic solvent and an aqueous solution typically comprising polyvinyl alcohol (PVA) is added to create a water-in-oil (w / o) emulsion. The emulsion is then subjected to evaporation, and particles are formed and can be collected and stored frozen (e.g., -20 °C) until further use. For example,in accordance with the disclosure, a plurality of particles can be prepared by dissolving an uncapped polymer in an organic solvent, typically a polar aprotic solvent, such as CH2CI2,, thereby forming a solution, adding polyvinyl alcohol to the solution, thereby forming a polymer matrix, and removing the organic solvent, thereby forming the plurality of particles, wherein the polymer matrix comprises PLGA and / or PLA, wherein the uncapped polymer comprises free carboxyl groups such that the formed particle also includes free carboxyl groups.

[0073] Polymeric particles can also prepared using a double water-oil-water (W / O / W) emulsion of uncapped polymer. First, an aqueous solution of the uncapped polymer is dissolved in an organic solvent to create a water-in-oil (w / o) emulsion. Optionally, a porosigen and / or osmotic adjusting agent may be added to increase the porosity of the particles. Next, an aqueous solution typically comprising polyvinyl alcohol (PVA) is added to create the second emulsion. The emulsion is then subject to evaporation, and microparticles are formed and can be collected, sieved, lyophilized, and stored frozen (e.g., -20 °C) until further use.

[0074] One particle of the disclosure can have a particle diameter, and a plurality of particles can have an average particle diameter, ranging from about 10 nm to about 10 pm, for example at least about 20, 25, 30, 40, 45, 50, and / or 55 nm and / or up to about 1 , 2, 3, 4, 5, 6, 7, 8, 9 or 10 pm. In some cases, the particles or plurality of particles can have a particle diameter, or an average particle diameter, of about 30 nm to about 10 pm, about 100 nm to about 1 pm, or about 500 nm to about 900 nm. For example, particles of the disclosure can have an average particle diameter in the range of about 10 nm to about 10 pm, about 10 nm to about 1 pm, about 100 nm to about 10 pm, about 100 nm to about 1 pm, about 10 nm to about 100 nm, or about 100 nm to about 10 pm. Alternatively, one particle of the disclosure can have a particle diameter, and / or a plurality of particles can have an average particle diameter, ranging from about 10 pm to about 100 pm, for example at least about 15, 20, 25, 30, 35, 40, or 45 pm and / or up to about 40, 45, 50, 55, 60, 65, 70, 75, 80, 85, or 90 pm. In some cases, the particle can have a particle diameter, and a plurality of particles can have an average particle diameter, of about 10 pm to about 100 pm, about 15 pm to about 90 pm, or about 20 pm to about 80 pm. Particle sizes can be quantified by SEM images and / or by using a laser diffraction particle size analyzer (Master Sizer 2000, Malvern Instruments Ltd. Malvern, UK).

[0075] In some embodiments, the biodegradable polymeric particle is a nanoparticle. As used herein, the term “nanoparticle” refers to a solid or semi-solid particle having a diameter of less than about 2 pm. The nanoparticle of the disclosure can have a particle diameter, and a plurality of nanoparticles can have an average particle diameter ranging from about 10 nm to about 2 pm, for example, at least about 20, 25, 30, 40, 45, 50, and / or 55 nm and / or up to about 1 or 2 pm. For example, the nanoparticle can have a particle size, and a plurality of nanoparticles can have an average particle size, of about 30 nm to about 2 pm, about 100 nmto about 1 m, or about 500 nm to about 900 nm. The particle size can represent a weight-, number-, surface area-, or volume-average size for a particle size distribution of the nanoparticles. Nanoparticles having a spherical shape are referred to as nanospheres.

[0076] In embodiments, the biodegradable polymeric particle is a microparticle. As used herein, the term “microparticle” means a solid or semi-solid particle having a diameter of greater than about 10 pm and less than about 100 pm. The microparticle of the disclosure can have a particle diameter, and a plurality of microparticles can have an average particle size, ranging from about 10 pm to about 100 pm, for example at least about 15, 20, 25, 30, 35, 40, and / or 45 pm and / or up to about 40, 45, 50, 55, 60, 65, 70, 75, 80, 85, or 90 pm. In some cases, the microparticle can have a particle size, and a plurality of microparticles can have an average particle size, of about 10 pm to about 100 pm, about 15 pm to about 90 pm, or about 20 pm to about 80 pm. The particle size can represent a weight-, number-, surface area-, or volume-average size for a particle size distribution of the microspheres. Microparticles having a spherical shape are referred to as microspheres.

[0077] In general, nanoparticles are expected to have a larger density than microparticles. Accordingly, the nanoparticles of the disclosure are generally less porous than the microparticles of the disclosure which can be demonstrated by measuring the porosity of the particles. Further, nanoparticles are expected to have very short diffusion path lengths in the polymer matrix. As such, relatively non-porous nanoparticles are generally easier to load than porous microparticles.Methods of Loading Polymeric Particles

[0078] Generally, prior art methods involving loading particles comprising end-capped polymers and use various excipients such as trapping agents, and / or a metal salt (e.g., water-insoluble base(s)) to enhance loading and release profiles of therapeutic agents. To the contrary, the present method of the disclosure achieves comparable and / or even, surprisingly, significantly enhanced loading (therapeutic agent content), encapsulation efficiencies, and / or release profiles without significant amounts of end-capped polymers and / or without the need for including excipients in the polymer matrices such as trapping agent(s), and water-insoluble base(s), during loading. As a result, the particles may be “substantially free of excipients” such that the particle contains less than about 5, 4, 3, 2, 1 , 0.5, 0.1 , or 0.01 wt% of any trapping agent(s) or water-insoluble base(s), prior to being exposed to an aqueous solution.

[0079] The particles of the disclosure feature a polymer matrix made of at least some uncapped polymer such that the particle have free carboxyl groups which are available to associate with the therapeutic agent. The therapeutic agent is loaded into the preformedbiodegradable polymeric particles by incubating the particles in a loading solution of the net positively charged therapeutic agent at a first pH. At the first pH, the particle adopts the first state, characterized by relatively more interconnected pores and channels in the particle. At this first pH, loading and encapsulation of the therapeutic agent more efficiently occurs. After incubation in the loading solution at the first pH, the particles are incubated in a solution at the second pH, which may be provided by acid added to and / or generated in situ in the loading solution, thereby causing the particle to switch to the second state, which is characterized by relatively less interconnected pores and channels in the particle. At this second pH, loading and encapsulation of the therapeutic agent are inhibited, and self- healing of the polymer matrix occurs such that the particle advantageously demonstrates sustained release and avoids initial burst release.

[0080] The particles of the disclosure can adopt the first state when the particles are exposed to the first pH. The first pH is between about 6.8 and about 8.0. The first pH can be maintained during the incubation period when loading and encapsulation are performed. While in the first state, the pores and channels of the particles become more interconnected, i.e. , the particle “opens,” thereby increasing the particle surface area. At the first pH, the carboxyl groups are negatively charged, while the therapeutic agent is net positively charged. The electrostatic interactions between the negatively and positively charged groups facilitate loading because the positively charged moieties of the therapeutic agent can penetrate the pores and channels of the particle and associate with the polymer matrix. As a result, association of the negatively charged carboxyl groups and the net positively charged therapeutic agent can advantageously enhance loading and encapsulation.

[0081] The particles of the disclosure can adopt the second state when the particles are exposed to the second pH. The second pH is between about 4.5 and about 6.5. Without intending to be bound by theory, it is understood that incubating the loaded particles at the second pH prevents an undesirable initial burst release of the therapeutic agent by enhancing the self-healing properties of the polymeric matrix. While in the second state, the pores and channels of the particles become relatively less interconnected, i.e., the particle “closes,” thereby reducing the particle surface area. At the second pH, the carboxyl groups of the particle are charge neutral and the polymer chains no longer repel one another such that polymer healing can occur.

[0082] As the particles are incubated in the loading solution at the first pH, association of the positively charged moieties of the therapeutic agent with the negatively charged carboxyl groups releases protons (H+) into the loading solution, thereby acidifying the loading solution. In methods of the disclosure, the loading solution therefore can include a buffer to maintain the first pH and resist any significant pH change upon association of the therapeuticagent and the particle, or the loading solution can be formulated to generate the solution at the second pH in situ, i.e., upon association of the positively charged moieties and the negatively charged carboxyl groups. As the particles are exposed to the solution at the second pH, the carboxyl groups on the surface of the particles become charge neutral thereby inhibiting association of the positively charged moieties of the therapeutic agent with the free carboxyl groups of the particle and preventing further loading of the therapeutic agent. Because the pores and channels of the particles become relatively less interconnected in the second state, the loaded biodegradable polymeric particles demonstrate low initial burst upon subsequent additional exposure to a solution at the first pH, i.e., the polymeric particles advantageously release less than about 5% of the therapeutic agent in the first 24 hours after administration.

[0083] The therapeutic agent content or loading is quantified as: / Mass of therapeutic agent encapsulated in particlesX - x 100.\ Total mass of particles )

[0084] The percentage encapsulation efficiency of the therapeutic agent is calculated as: / Mass of therapeutic agent encapsulated in particles\ - x 100.\ Total mass of therapeutic agent in loading solution /

[0085] Advantageously, the particles of the disclosure are capable of high therapeutic agent loading. In many cases the loading wt% is in the range of about 2 wt.% to about 20 wt.%, about 4 wt.% to about 18 wt.%, about 6 wt.% to about 16 wt.%, about 8 wt.% to about 14 wt.%, or about 10 wt.% to about 12 wt.%, based on the entire weight of the particle. In preferred embodiments, the particle has a therapeutic content greater than or equal to 8%. Similarly, the methods described herein provide excellent encapsulation efficiency, for example at least about 95%, 90%, 92%, 95%, 98%, or 99% and / or up to about 95%, 98%, 99%, 99.5%, 99.9% or 100% efficiency. In embodiments, the encapsulation efficiency is about 40% to about 100%, about 50% to about 100%, or greater than or equal to 60%, for example, about 60% to about 100%, about 70% to about 100%, about 80% to about 100%, or about 90% to about 100%.

[0086] Generally, the loading solution can include an aqueous solvent, a buffer, a surfactant, and a therapeutic agent, preferably a net positively charged peptide at the first pH of the loading solution. In many embodiments, the aqueous solvent can be water or saline. In many embodiments, the surfactant is Tween® 80 (polyoxyethylene (80) sorbitan monooleate). In many embodiments, the buffer can include 2-[4-(2-hydroxyethyl)piperazin- 1-yl]ethanesulfonate (HEPES), or phosphate. The concentration of the buffer can generallyrange from about 0.01 M to about 0.5 M. In many embodiments, the solution at the second pH is formed in situ in the loading solution and the loading solution comprises a buffer. In embodiments where the solution at the second pH is generated in situ by upon association of the positively charged moieties and the negatively charged carboxyl groups, the concentration of the buffer can range from about 0.01 M to about 0.15 M or about 0.01 M to about 0.3 M. Thus, in embodiments in which the solution at the second pH is formed in situ in the loading solution, the loading solution comprises a buffer in a concentration of about 0.01 M to about 0.3 M. In such embodiments, the concentration of the buffer can be about 0.01 M, about 0.05 M, about 0.1 M, about 0.15 M, about 0.2 M, about 0.25 M, or about 0.3 M. In embodiments where the solution at the second pH is prepared by either adding an acid to the loading solution at the first pH or prepared separately, the concentration of the buffer can be greater than about 0.15 M, typically greater than about 0.3 M .

[0087] The therapeutic agent can have a net positive charge in the loading solution, for example, greater than about +1 , or greater than about +1.5, or greater than about +2. Typically, the therapeutic agent has a net positive charge greater than about +1 in the loading solution. More preferably, the therapeutic agent has a net positive charge greater than or equal to about +1.5 in the loading solution. For example, the therapeutic agent can have a net positive charge greater than or equal to about +2 in the loading solution.

[0088] The methods of the disclosure allow the pH of the loading solution to be adjusted to the first pH by addition of acid or base.

[0089] The methods of the disclosure include incubating the particles in a loading solution at a first pH between about 6.8 to about 8.0 and incubating the particles in an incubation solution at a second pH between about 4.5 to about 6.5. The solution at the second pH can be prepared separately, in situ by upon association of the positively charged moieties and the negatively charged carboxyl groups accompanied by a pH drop, or by adding an acid to the loading solution at the first pH. As used herein, the term “pH drop” can refer to the difference between the first pH and the second pH when the second pH is formed in situ from the loading solution at the first pH.

[0090] Any acid can be added to the loading solution at the first pH to arrive at the loading solution at the second pH. For example, phosphoric acid, trifluoroacetic acid, or acetic acid can be used. Of course, other acids may also be used. The separately prepared loading solution at the second pH can include the same components as the loading solution at the first pH.

[0091] The methods of the disclosure can further include collecting and drying the one or more particles, after incubating the one or more particles at the first pH and subsequentlyproviding the one or more particles in the solution at the second pH, thereby incubating the one or more particles at the second pH.

[0092] The methods of the disclosure can also include collecting and drying the one or more particles, after incubating the one or more particles at the second pH, and subsequently reconstituting the particles.Release Profile of Polymeric Particles

[0093] Loading the therapeutic agent into biodegradable polymeric particles and associating (e.g., binding or coupling) the therapeutic agent to the polymer and within the particle, as disclosed herein, provides a particle exhibiting high loading and encapsulation efficiency of the therapeutic agent, sustained release, and in particular the ability to provide a particularly desirable loading capability and then to switch the state of the particles so as to provide a particularly desirable release profile.

[0094] In methods of the disclosure, the particles can be exposed to the first and second pH values to produce the first and second states in the particles, respectively. The first state is characterized by relatively more interconnected pores and channels of the polymer matrix with increased particle surface area exposed to an external solution. The therapeutic agent, particularly therapeutic agents having a net positive charge greater than or equal to +1 at neutral pH, or greater than about +1.5, or greater than about +2, such as positively charged peptides, is loaded and encapsulated more efficiently when the particle is in the first state, through association of positively charged moieties of the therapeutic agent with the negatively charged carboxyl groups of the particle as well as the increased surface area of the particle in the first state. The second state is characterized by the pores and channels of the polymer matrix being relatively less interconnected, with reduced particle surface area exposed to an external solution. When particles in the first state are exposed to the second pH, the negatively charged carboxyl groups of the particle become protonated, making the carboxyl groups charge neutral, causing the particle to become relatively more “closed” to an external solution and thereby inhibiting further association of the therapeutic agent with the particle. As a result, loading can be effectively ended by incubating the particles in a solution at the second pH. Without being bound by theory, it is believed that upon exposure to the second pH, the surface area of the particle decreases as the pores and channels of the particle become less interconnected, thereby trapping the therapeutic agent in the “closed” pores and channels of the particle in the second state. Additionally, exposing loaded particles in the second state to the first pH at physiological temperature (after the biodegradable particles are initially loaded with a therapeutic agent in the first state and transitioned to the second state) does not cause the pores to open again instantaneously but generally requires polymer degradation, polymer erosion, or osmotically mediated polymerpore breakage to move towards the previous first state having relatively more open and interconnected pores in order to allow the therapeutic agent to be released. Without being bound by theory, it is believed that this lag time directly after exposure of loaded particles in the second state to the first pH range provides the particles with low initial burst release, i.e. , the polymeric particles advantageously release less than about 5% of the therapeutic agent in the first 24 hours after administration.

[0095] In methods of the disclosure, loaded particles in the second state can advantageously be exposed to a solution in the first pH range to initiate release of the therapeutic agent from the particle. Thus, control of the pH in the vicinity of the particles can provide further control over the release profile of the therapeutic agent from the particles. The past presence of the second state of the particle enhances sustained continuous release of the therapeutic agent by trapping the therapeutic agent in the interconnected pores and channels of the particle. When the particles in the second state are exposed to a solution at the first pH, the particles at some point transition from the second state to the first state which advantageously promotes formation of interconnected pores and thereby promotes release of the therapeutic agent after a lag time, thereby still avoiding an initial burst release as described herein. On the other hand, when particles in the second state are provided in a media capable of providing a pH of about 4.5 to about 6.5 within an unstirred boundary layer of the particles upon hydration, even further enhanced sustained release of the therapeutic agent from the particles is expected by further reducing the initial burst release. It is noted that if the initial burst release is lower, more drug is remaining and faster release rates later on in the release period can be achieved.

[0096] The biodegradable polymeric particles of the disclosure are capable of continuous release of the therapeutic agent in vitro for > 21 days with low initial burst, i.e., the polymeric particles advantageously release less than about 5% of the therapeutic agent in the first 24 hours after administration, i.e., after exposure to a phosphate-buffered saline buffer (i.e., as determined by the amount of therapeutic agent released from t = 0 hours to t = 24 hours in a phosphate-buffered saline buffer (pH = 7.4)). Moreover, the biodegradable polymeric particles of the disclosure can provide sustained continuous release for at least 2 weeks, at least 3 weeks, for example, for about 4 weeks, about 5 weeks, about 6 weeks, about 7 weeks, about 8 weeks, or even longer.

[0097] In some cases, the particles of the disclosure can also advantageously avoid an undesirable initial burst of the drug from the particle. As used herein, “initial burst” or “initial burst release” refers to the amount of therapeutic agent released in the first 24 hours after administration I exposure to physiological conditions or release media. The particles of the disclosure can advantageously have an initial burst release of the therapeutic agent of about5% or less in the first 24 hours after exposure to a phosphate-buffered saline buffer (i.e., as determined by the amount of therapeutic agent released from t = 0 hours to t = 24 hours in a phosphate-buffered saline buffer (pH = 7.4)). For example, particles of the disclosure can have an initial burst release of the therapeutic agent in the range of about 0% to about 5%, about 1% to about 5%, about 1% to about 4%, about 1% to about 3%, about 1% to about 2%, in the first 24 hours after exposure to a phosphate-buffered saline buffer at a pH of about 7.4.

[0098] Additionally, the biodegradable polymeric particles of the disclosure can advantageously demonstrate a substantially zero-order release profile. A substantially zeroorder release profile refers to a release profile (i.e., a release rate) that is substantially constant over a period. As used herein, the term “substantially zero-order release profile” means that the rate of release of the therapeutic agent from the particle does not vary by more than about 100%, e.g., no more than about 35% over the lifetime of the particle. In one aspect, the term substantially zero-order release profile refers to a release profile in which about 10 wt.% to about 20 wt.% of the therapeutic agent (relative to the original amount of therapeutic agent in the particle) is released per week for at least 3 weeks, for at least 4 weeks, for at least five weeks, for at least six weeks, for at least seven weeks, and / or at least eight weeks.

[0099] The Loo-Riegelman method can be used to deconvolute the in vivo plasma profile of microspheres loaded with therapeutic agents. The percentage absorbed in vivo was estimated in the following way:Ab_ Cp+Ct+K10(AUC)t0A«, K10(AUC)“ where Cpand Ct are central compartment and peripheral compartment drug concentrations, respectively. Kto and AUC are the elimination micro rate constant and area under the plasma versus time curve, respectively. Aband Amare the cumulative amount absorbed at time t and at time infinity, respectively.

[0100] For the Loo-Riegelman method, estimates for distributive and elimination micro rate constants (Kto, K12, and K21) can be obtained by running a two-compartment model from Phoenix© WinNonlin® using separate single intravenous injection pharmacokinetic data. Pharmacokinetic parameters, maximum plasma concentration (Cmax), time to peak drug concentration (Tmax) and area under the plasma concentration-time curve (AUC), were estimated by Non-Compartmental Analysis (NCA) using Phoenix© WinNonlin® (Pharsight Corporation).Methods of Administration and Formulations

[0101] The polymeric particles of the disclosure can be used in injectable compositions.For example, the disclosure provides injectable formulations for parenteral administrationincluding the biodegradable polymeric particles of the disclosure. The formulations can further include a pharmaceutically acceptable excipient. Suitable routes for parenteral administration include intravenous, subcutaneous, intradermal, intramuscular, intraarticular, and intrathecal. In one aspect, subcutaneous administration is preferred. Advantageously, the biodegradable polymeric particles of the disclosure can provide a route for subcutaneously administering therapeutic agents such as peptides with controlled and sustained release thereof with delivery of the therapeutic agent occurs for a predetermined period, and can be stopped, for example, when a desired effect has been achieved.

[0102] The pharmaceutically acceptable excipient can include sterile water, saline, or a buffered solution. Additional excipients can include, but are not limited to carboxymethylcellulose sodium, D-mannitol, polysorbate, and combinations thereof, which can be added to help resuspension of the polymeric particles. The polymeric particles can advantageously be injected to a subject via incorporation into a microneedle.

[0103] In formulations of the disclosure, a media can be included that can provide a specified pH, typically at the second pH range between about 4.5 and about 6.5, within an unstirred boundary layer of the particles upon hydration. When the particles are hydrated, the flow of liquid in the vicinity of the therapeutic agent is restricted. The restricted liquid flow in the vicinity of the therapeutic agent produces a region in the particles where the diffusional movement of therapeutic agent exceeds movement due to convection. This results in a concentration gradient between the bulk solution and the particle surface known as the unstirred boundary layer. The formed unstirred boundary layer can reduce the apparent permeability of the particle, limiting the amount of therapeutic agent able to pass through the surface of the particle. Thus, the media can provide a pH of about 4.5 to about 6.5 within the unstirred boundary layer of the particles.

[0104] In embodiments, the media is in contact with, encapsulated, or accompanied by an anionic polysaccharide, a poorly water-soluble acid source, or a thermoreversible gel, as described herein, such that the media can provide a specified pH, typically at the second pH range between about 4.5 and about 6.5, within an unstirred boundary layer of the particles upon hydration, for example, by reconstituting the remote-loaded, biodegradable polymeric particles in an injection vehicle and / or interstitial fluid upon injection. In embodiments, the media that can provide a specified pH, typically at the second pH range between about 4.5 and about 6.5, within an unstirred boundary layer of the particles upon hydration, is an anionic polysaccharide, for example, an anionic polysaccharide such as dextran sulfate. The particles, whether in the first or second state, can be coated with the foregoing media or dispersed in such media.

[0105] In embodiments, the media that can provide a specified pH, typically at the second pH range between about 4.5 and about 6.5, within an unstirred boundary layer of the particles upon hydration can be a thermoreversible gel. As used herein, a thermoreversible get is a gel formed due to entanglement of polymer chains causing viscosity changes at a gelation temperature and capable of reverting back to a liquid state when at a temperature lower than the gelation temperature. The thermoreversible gel can be PEG-chitosan, chitosan-polyethylenimine, chitosan-arginine or glycol-chitosan-spermine, or other known thermoreversible gels. Since the gel is in the liquid state below the gelation temperature (and room temperature is below the gelation temperature), the thermoreversible gel can be included in an injectable formulation of particles of the disclosure. Upon injection, and exposure of the gel to a higher temperature above the gelation temperature, the viscosity of the thermoreversible gel can increase, preventing the movement of the particles.Advantageously, the thermoreversible gel formulation can be acidified to a desired pH between about 4.5 and about 6.5, thereby providing low initial burst and sustained release of therapeutic agents, without deleterious effects on the subject. The particles, whether in the first or second state, can be coated with a thermoreversible gel or dispersed in a thermoreversible gel. The thermoreversible gel may be a hydrogel, for example, a poloxamer.

[0106] As used herein, a poloxamer is a triblock copolymer with a central hydrophobic block of (polypropylene oxide)) flanked by two hydrophilic blocks of (poly(ethylene oxide)) and generally has the following structure of Formula (I):can range from 2 to 130 and y can range from 15-67. In various cases, x can range from 70 to 130. In many cases, x can range from 90-110. In various cases, x is 101. In various cases, y can range from 40 to 67. In many cases, y can range from 50 to 60. In various cases, y is 56. In various cases, x is 101 and y is 56. In many cases, the thermoreversible gel is a poloxamer and is present in the range of about 10 %(m / v) to about 30 % (m / v). In various cases, the poloxamer is present in the range of about 15 %(m / v) to about 20 %(m / v).

[0107] Additionally, the media that can provide a specified pH, typically at the second pH range between about 4.5 and about 6.5, within an unstirred boundary layer of the particles upon hydration can comprise an acid source. Preferably, the acid source has low solubility in water, for example, the acid source has a solubility in water of less than 1 mg / mL or a solubility in the range of about 1 mg / mL to about 10 mg / mL. The poorly water-soluble acidsource can be formulated as a slowly dissolving particulate or granule with the PLGA / PLA particles to maintain a reduced pH for a brief period in the boundary layer of the microparticles. Alternatively, a poorly water-soluble acid could be applied as a coating of the PLGA / PLA particles to provide and maintain a pH around the particles low enough to prevent undesired burst release. The acid used and the coating thickness can be varied to control the release profile of the formulated particles. Suitable poorly water-soluble acids include benzoic acid, naphthoic acid, fatty acids, and other poorly water soluble organic acids acceptable in parenteral dosage forms. Furthermore, a biocompatible polyanionic excipient, such as dextran sulfate, heparin, or hyaluronic acid can be embedded into the surface of the PLGA / PLA microparticles to provide a lowering of the boundary layer pH by the Donnan effect.EXAMPLES

[0108] The following examples demonstrate the preparation and remote loading of microparticles with setmelanotide (Example I), preparation of microparticles with varying porosity and remote loading with setmelanotide (Example II), an evaluation of pH controlled remote loading of microparticles with setmelanotide (Example III), in vitro studies of pH controlled microparticles remote loaded with setmelanotide (Example IV), in vitro studies of pH controlled microparticles remote loaded with leuprolide (Example V), in vivo pharmacokinetic studies of pH controlled microparticles remote loaded with setmelanotide (Example VI), an in-vivo study of administration of pH controlled microparticles remote loaded with setmelanotide (Example VII), the preparation and remote loading of nanoparticles with MOG 38-50 and NRPA7 peptides (Example VIII), an evaluation of pH controlled remote loading of microparticles with setmelanotide under varying buffer concentrations and an in vitro study of their initial burst release profile (Example IX), the preparation of microparticles with varying PLGA molecular weights and remote loading thereof with setmelanotide (Example X), an in vitro study of pH controlled microparticles with varying PLGA molecular weight remote loaded with setmelanotide (Example XI), and an in vitro study of pH controlled microparticles remote loaded with leuprolide and suspended in pH Buffered Hydrogels (Example XII).Materials and Methods

[0109] PLGA (75:25 poly(lactic-co-glycolic acid prepared by polycondensation, 13kDa) was purchased from Wako Pure Chemical Industries, Ltd. (Japan). PLGA 50:50 (poly(lactide-co-glycolide prepared by ring opening polymerization, 7000-17000 kDa) was purchased from Evonik AG (Germany). Poly(vinyl alcohol) (PVA, MW 25000, 88% hydrolyzed) was purchased from Polysciences, lnc.(Warrington, PA, USA). D-mannitol was purchased from SPI Pharma™ (China). Sodium Carboxymethylcellulose (CMC) waspurchased from The Dow Chemical Company (Midland, Ml, USA). Trehalose, Tween® 80, and all other chemicals were purchased from Sigma-Aldrich, Inc. (St. Louis, MO, USA). HEPES, sodium acetate, and all solvents used were obtained from Thermo Fisher Scientific Inc. (Waltham, MA, USA). ddbhO was from Thermo Scientific™ Barnstead™ GenPure™. Setmelanotide acetate was obtained from RS Synthesis, LLC (Louisville, KY, USA).

[0110] Size Determination. The microsphere size and size distribution of blank and setmelanotide-loaded microspheres were measured using a Malvern Mastersizer 2000 (Malvern Panalytical Ltd, U.K.). Three measurements were performed with 40-50 mg dry microspheres suspended in the instrument sample dispersion unit at a stirring rate of 2500 rpm.

[0111] SEM Imaging. The morphology of microspheres was evaluated using TESCAN MIRA3 FEG SEM (TESCAN Brno, s.r.o., Kohoutovice, Czech Republic). Samples were mounted on brass stubs by double-sided carbon adhesive tape and then sputtered with gold for 90 seconds at 18 mA under vacuum. For cross-section images, a razor blade was used to fracture the microspheres on the stubs. Images were taken with a gun voltage of 10 kV and beam intensity of 8.

[0112] Energy-dispersive X-ray spectroscopy (EDS) with Focused Ion Beam-SEM (FIB- SEM). Thermo Fisher Helios G4 Plasma FIB UXe (Thermo Fisher, USA) was used for FIB cut and SEM imaging. Samples were prepared for SEM Imaging as described above. The sample stage was first tilted by 52 degrees. Then, a 30 kV plasma ion beam was used to obtain a cross-section on a selected microsphere. SEM images were taken with a tube voltage of 1 kV and tube current of 0.2 nA. The element identification of the cross-section was conducted using a coupled energy dispersive X-ray spectrometer (EDS). A tube voltage of 5 kV and tube current of 0.8 nA was used.

[0113] Powder X-Ray Diffraction (PXRD). Powder X-ray diffraction patterns of setmelanotide, PLGA polymer and remote-loaded microspheres were performed using Rigaku Miniflex 600 XRD (Tokyo, Japan) with Cu-Ka irradiation at a tube voltage of 40kV and tube current of 15mA. All samples were scanned from 26= 3 to 50° with 1 s scanning rate and 0.02° step intervals.

[0114] Fourier Transform Infrared Spectroscopy (FT-IR). FT-I R spectroscopy was conducted using Jasco FT / IR 4100 with diamond / ZnSe attenuated total reflectance (ATR) crystal (Tokyo, Japan). After scanning the background, Powder samples were pressed in the sampling accessory with best contact with the crystal base. All spectra were collected at 64 scans with a resolution of 1 cm-1and a scanning range of 4000-650 cm-1.

[0115] Ultra-Performance Liquid Chromatography (UPLC). LIPLC was conducted with a Waters Acquity system, using a UV absorbance detector set at 280 nm. Each sample was syringe filtered prior to analysis (Millex Syringe Filter, Durapore® 0.45 pm PVDF) using a ACQUITY UPLC® BEH C18 column (1.7 pm, 2.1*100 mm) and a mobile phase of acetonitrile containing 0.1% (v / v) trifluoroacetic acid (A) and water containing 0.1% (v / v) trifluoroacetic acid (B). Samples were collected using an isocratic elution gradient with 20% A: 80% B over 3.5 min at a flow rate of 0.4 mL / min. The sample injection volume was 10 pL.

[0116] Statistical and Regression Analysis. Statistical analyses and regressions were performed with the software Prism (Graphpad, San Diego, CA) to determine the significance of the difference. Repeated measures one-way ANOVA test with Tukey's multiple comparisons test was used to evaluate the statistical difference among three groups over three or more time points. Statistical comparison of the two groups was performed by an unpaired t test. Comparison of two groups over three or more time points was analyzed using repeated measures two-way ANOVA with Sidak's post hoc test. The value of P < 0.05 was accepted as statistically significant.Example I - Preparation & Remote Loading of Microparticles with Setmelanotide

[0117] PLGA microparticles were prepared by double water-oil-water (W / O / W) emulsion. The first emulsion was created by dissolving PLGA (50 / 50) (800 mg) in methylene chloride (1 mL) and homogenizing the solution (VirTis Tempest I.Q.2) at 10,000 rpm for 1 min with an ice bath to create a w / o emulsion. Next, 4 mL of a 5 wt.% polyvinyl alcohol (P A) solution was added to the emulsion. The mixture was vortexed (Scientific Industries Vortex Genie 2) for 1 minute at maximum speed to create the second emulsion. The w / o / w emulsion was hardened by adding the emulsion to a 100 mL stirring bath of 0.5 wt.% PVA and stirring for 3 hours until the solvent evaporated. After hardening, microparticles were washed extensively with diH2O and sieved for size in the range 20-63 m. Microparticles were lyophilized (Labconco FreeZone 2.5) and stored frozen (-20 °C) until further use.

[0118] Setmelanotide loading solution (20 mg / mL) was prepared by dissolving the peptide in a 0.1 M HEPES buffer and titrating the pH of the loading solution to pH 7.4 using 1 N sodium hydroxide (Fisher Scientific Accument® AE150 benchtop pH meter).

[0119] Pre-formed blank PLGA microparticles were incubated in the aforementioned setmelanotide loading solution. Specifically, 90 mg of blank microparticles were incubated in an aliquot of the titrated peptide solution (0.5 mL) at 37 °C and mixed for 24 h, using a rotational speed of 60 rpm. After incubation, microparticles were centrifuged (Eppendorf 5424R) for 10 minutes at 5,000 rpm and the supernatant was collected. Microparticles were next washed three times with 1 mL diH2O, and the supernatant was saved; centrifuging at5,000 rpm for 10 minutes between each wash. Loaded and washed microparticles were then lyophilized to remove excess water and stored at -20 °C until future use.Example II - Preparation of Microparticles with Varying Porosity & Remote Loading with Setmelanotide

[0120] Pre-formed blank PLGA microparticles with different porosities were prepared by adjusting the procedure described above. An aliquot of a porosigen solution was added to the CH2CI2 solution prior to homogenizing the solution and forming the first emulsion.

[0121] Specifically, three remote loaded PLGA microsphere formulations of high, medium, and low porosity, were prepared with 100 pL, 50 pL, and 0 pL, aliquots of a trehalose porosigen solution (500 mg / mL), respectively. Microparticles (180 mg / mL) were incubated in a solution of setmelanotide (20 mg / mL) in 0.1M HEPES buffer solution at pH 7.4 for 24 h at 37 °C.

[0122] The encapsulation efficiencies and drug loading of microspheres with high, medium, and low porosity were significantly different when incubating the drug-free microspheres with peptide solution at 37 °C for 24 h, i.e. , 6.3 ± 0.3%, 4.6 ± 0.2%, and 2.2 ± 0.2% w / w drug loading, and about 63%, about 46%, and about 22% encapsulation efficiency, respectively (Figure 1).

[0123] FT-IR spectroscopy was used to characterize the local environment of N-H stretching of setmelanotide before and after remote loading (Figure 2). The pure setmelanotide (shown as (c) in Figure 2) was found to exhibit structural features of N-H stretching at 3287.6 cm-1whereas the pure PLGA polymer (shown as (a) in Figure 2) had no absorption near this wavenumber. The remote loaded microspheres (shown as (b) in Figure 2) displayed a peak at 3348.5 cm-1, strongly supporting the conclusion of the ion pairing between cationic arginine residues and carboxylate anions, which causes the absorbance peak shift of N-H stretching to a higher frequency (Figure 2). However, the change was subtle considering the low weight ratio of setmelanotide to PLGA polymer (6:94).

[0124] The absence of a diffraction peak was also observed in powder XRD analysis of all samples: setmelanotide, PLGA polymer, and remote-loaded microspheres, indicating that setmelanotide was in an amorphous state before and after remote loading (Figure 3).

[0125] FIB-SEM in combination with EDS provided evidence for peptide distribution. A FIB-SEM cross-section of remote loaded high-porosity microspheres showed the nitrogen signal, which was associated with setmelanotide (not shown), was distributed evenly throughout the microspheres (not shown). The signals of carbon and oxygen were from both the polymer and the peptide.

[0126] SEM images of the surfaces of three different formulations before and after peptide remote loading demonstrate that the surface porosity of the microspheres decreased as the trehalose solution volume to the inner water phase was reduced (not shown). The inner porosity of the microspheres with 100 pL porosigen solution was slightly higher than the microspheres with 50 pL porosigen solution, albeit the surface morphology showed a more substantial difference. By contrast, the cross-sectional morphology of blank microspheres without addition of porosigen appeared quite non-porous with no sign of a network of larger pores.

[0127] After remote loading peptide, some passive polymer healing, i.e. , rounding and / or disappearance of tiny holes, was observed, particularly visible in the medium porous formulation. The low-porosity microspheres showed some indentations in their surface after drug loading. The particle size and size distribution had negligible change before and after the incubation of blank microspheres and peptide solution (Figs. 7, 8).Example III - Study on pH Controlled Remote Loading of Microparticles with Setmelanotide

[0128] The microparticles were prepared with 100 pL of trehalose stock solution (500 mg / mL). The loading solution was titrated with 1 N NaOH to achieve the desired initial pH. The pH of the setmelanotide loading solution during remote loading of the high-porosity formulation was measured at different incubation times ranging from 0 h to 48 h. The pH of the loading solution was varied by either changing the concentration of the buffer present in the loading solution and / or titrating the loading solution to a given pH value.

[0129] Encapsulation efficiency, the ratio of actual encapsulated setmelanotide to total setmelanotide added, was determined by mass loss. Briefly, the supernatant was removed in 50 pL increments after centrifuging the remote loading suspension (6018*g for 5 min) at different incubation times. Then, the microspheres were washed by adding 1 mL diH2O and centrifuging again. 900 pL supernatant was removed and combined with the initial supernatant. Washing was repeated two more times to combine with the previous supernatant and obtain the final supernatant, which was analyzed by UPLC as described above.

[0130] The amount of setmelanotide encapsulated was calculated by subtracting the amount of setmelanotide in the supernatant from the total amount of setmelanotide added in the loading solution. The pH of the final supernatant was measured using a Fisher Scientific Accument® AE150 benchtop pH meter. The pH of the final supernatant and encapsulation efficiencies for each sample are presented in Table 1 and shown in Figure 9.Table 1. Remote loaded microparticle loading solution pH values and encapsulation efficiencies for the net positively charged peptide, setmelanotide.

[0131] The pH of the loading solution for Sample S1 in Table 1 gradually decreased from 7.4 to 4.7 from 0 h to 24h, finally stabilizing at pH~4.7 thereafter, as shown in Figure 7. Over the same time interval, the encapsulation efficiency determined by mass loss of peptide in solution increased rapidly over the first 4 hours and slowly reached a plateau by 16-24 h.

[0132] Shown in Figure 6, as incubation time increases from 24 hr to 48 hr, the pH of the loading solution drops for each formulation. Additionally, as the buffer concentration was decreased, the change in pH before and after loading was larger. As the buffer concentration decreases, the capacity of the buffer to resist pH changes is reduced. As a result, the pH of the loading solution with a lower concentration of buffer is less resistant to change.

[0133] The shape of the pH change curve in Figure 7 was similar to the titration curve of the HEPES buffer after adding acetic acid (Figure 8). In contrast, blank microspheres incubated under identical conditions without peptide displayed only a slight pH decrease from 7.4 to 6.8 after 24 h.

[0134] Hence, the pH drop after loading is consistent with the salt formation of PLGA- setmelanotide, which involves a proton exchange between the more acidic polymer end group proton and the acetate counterion of the peptide in aqueous solution. Without being bound by theory, it is believed that as the pH drops to below 6, carboxyl end groups of the PLGA gradually become protonated and the charge at the polymer surface becomes more positive, thus repelling the remaining peptide cations in the solution. The ionic interaction between setmelanotide and PLGA polymer is therefore likely inhibited at this stage, thereby preventing further loading. The reduced interfacial surface area and reduced interconnected pores also inhibit further loading at this lower pH.Example IV - In Vitro Study of pH Controlled Microparticles Remote Loaded with Setmelanotide

[0135] The three formulations were then tested in vitro to characterize both the initial burst release profile and the long-term release profile in PBS release medium for 8 weeks. As shown in Figure 9, all formulations showed triphasic release behavior with a small initial burst release, followed by a lag phase of around 1 week. Following this lag phase,setmelanotide release was slow and continuous for 3-4 weeks, and was complete by around day 42. The presence of differing levels of pores in the PLGA matrix had a negligible impact on the overall in vitro release profile, indicating that polymer porosity did not significantly alter the release mechanism.

[0136] Microspheres with high porosity often exhibit a high initial burst release (Kim, S. et al. J. Biomat. Sci., Polymer Ed. 30(18), 1725- 1743 (2019); Yoo, J. et al. ACS Biomater. Sci. Eng. 6(11), 6053-6062 (2020)), however the initial burst release phase of each of the three setmelanotide loaded microparticle formulations was surprisingly low. Without being bound by theory, the low initial burst release (4-7% in day 1 for all three formulations) of setmelanotide may be affected by an ion-pairing association of peptide with two polymer chains within the swollen polymer phase (corresponding to the first state of the particles). Setmelanotide which contains 2 arginine side chains and 1 histidine residue with no acidic side chains, is expected to possess approximately two positive charges at neutral pH. Remote loading of octreotide, another peptide with approximately two positive charges at neutral pH also displayed a lower initial burst. Previous studies of peptides with approximately one positive charge such as leuprolide did not show this lower initial burst (Giles, M. B. et al. Nat Common. 13(1), 3282 (2022)). Further, the sustained release profile that avoids an undesirable initial burst release was particularly surprising because microspheres manufactured by encapsulating microspheres with leuprolide using conventional methods in a similar acid-terminated PLGA 50 / 50 did not demonstrate reduced release after 24 hours. (Hirota, et al. Journal of Controlled Release 2016, 244, 302-313). Thus, remote loading of the active agent at the first pH in combination with incubation at the second pH appears to provide advantageous release profiles to the biodegradable polymeric particles described herein.

[0137] Initial burst release is often attributed to the dissolution of immediately available peptides near the surface or in the existing pores (Tomic, I. et al. European Journal of Pharmaceutics and Biopharmaceutics 125, 21-27 (2018)) although polymer pore healing can be a factor as well (Wang, J. et al. J. Controlled Release, 82(2), 289-307 (2002); Schutzman, R. et al. J. Controlled Release, S0168-3659 (2023)). Therefore, the release profile suggests that instead of peptide adsorbed on the surface of the microsphere or dwelling in the porous channel, setmelanotide was successfully absorbed into the polymer phase via remote loading technique, enabling sustained release for long-term administration of the therapeutic agent.

[0138] The external morphology of the loaded microparticles was studied during incubation in the pH 7. 4 release medium by SEM imaging (not shown). Before incubation (t = 0), the high-porosity formulation showed numerous pores on the surface. During the firsttwo weeks of incubation while the polymer began to degrade, the pores on the microsphere surface gradually disappeared, with the external appearance of the microsphere becoming smooth and compact. The significant alteration in microsphere morphology is correlated with the lag phase in the initial 14 days. Note that we unexpectedly found that pH drops in the release media are common when the release media has relatively low buffering capacity, and in this case the pH of the release media was unchecked. Next, as the incubation of microspheres progressed to day 28, the microsphere skeleton structure underwent bulk degradation, which is linked to the substantial release of setmelanotide after this lag phase. It is worth noting that the formulation with high porosity displayed a 3-week zero order release from day 14 to day 35. Within the same period, the release kinetics very slightly shifted from zero order to first order as the porosity of the formulation decreased, although with no overall significant difference (repeated measures one-way ANOVA, P > 0.05).

[0139] The effect of remote loading on microparticle morphology was also evaluated with fluorescence microscopy. Blank PLGA microparticles were prepared as described in Example I, using a 200 pL aliquot of trehalose stock solution. Blank microparticles (4.1 mg) were incubated with dipyrrometheneboron difluoride (BODIPY) (5 pg / mL) in 1 mL of PBS solution with Tween® 80 and varying concentrations of HEPES buffer at 4 °C for 3 hr. As the concentration of HEPES buffer was increased, the pH of the loading solution did not change as much. For the 0.1 M HEPES loading solution, the pH dropped from 7.4 to 4.9, while for the 0.2 M HEPES loading solution, the pH dropped from 7.4 to 6.2.

[0140] After incubation, the microparticles were irradiated with light, causing the loaded BODIPY molecules to fluoresce. SEM imaging of the fluorescent particles indicated the relative porosity of the microparticles in loading solutions at different pH values (not shown). Comparison of the blank microparticles, with microparticles loaded in 0.2 M HEPES buffer and a solution pH of 6.2 after loading suggests that the particles are morphologically similar. The fluorescence intensity is strong on the peripheral of the microparticles, while the center of the particle is less intense. In contrast, the microparticles loaded in the 0.1 M HEPES buffer solution, with a final pH of 4.9 after loading, do not display strong fluoresence, suggesting the microparticle pores are less interconnected with the surface.Example V - In Vitro Study of pH Controlled Microparticles Remote Loaded with Leuprolide

[0141] PLGA particles with a particle size of 20-63 pm were prepared using the procedure described above, including an aliquot (0, 50, or 100 pL) of trehalose solution (500 mg / mL). The volume of porosigen was added as an aliquot of a trehalose stock solution (500 mg / mL). Loading was performed with lyophilized PLGA microparticles (90 mg) and leuprolide acetate (10 mg) in 0.5 mL of a 0.1 M HEPES buffer solution (pH 7.4) at 37 °C on a rotator for24 h. After the incubation, microparticles were centrifuged at 8000 rpm for 5 min. The supernatant was discarded, and remaining PLGA microparticles were washed with water 3 times followed by lyophilization. The lyophilized microparticles were stored at -20 °C before use. Loading content were measured and are presented in Table 2. Also reported in Table 2 are particles sizes for a different batch of microparticles prepared under the same conditions which are expected to have the same particle size.Table 2. Remote loaded microparticle size and loading for the net positively charged peptide, leuprolide.

[0142] The results for leuprolide are different from the results of remote loading setmelanotide, where a significant difference in encapsulation efficiency was observed in the same type of PLGA microparticles prepared in the presence and absence of trehalose addition within the 24 h interval at 37 °C.

[0143] The initial burst release profile of the pH controlled microparticles remote loaded with leuprolide was evaluated using the analysis described in Example IV.

[0144] The effect of the pH of the release media on burst release was studied at pH 7.4, 6.5, and 5.5 with microparticles L1, L2, and L3 after 2 h. The cumulative release of each of L1-L3 at each pH after 2 h are presented in Figures 10, 11, and 12, respectively. For each microparticle, the burst release increased as the pH of the release media increased. For L1 , the low porosity microparticles, the release % was about 8% when the pH of the release media was 7.4, and not detectable when the pH of the release media was 5.5 or 6.4 (Figure 10). For L2, the medium porosity microparticles, the release % was less than 15% when the pH of the release media was raised to 7.4, and the release % was about 5% or lower when the pH of the release media was 6.5 or lower (Figure 11). For L3, the high porosity microparticles, the release % was about 15% when the pH of the release media was raised to 7.4, and the release % was less than 5% when the pH of the release media was 6.5 or lower (Figure 12).

[0145] After 4 hours, the release % of the medium porosity L2 microparticles in the pH 7.4 PBS release medium was evaluated (not shown). For the L2 microparticles, the release % increased from about 8% after 2 hours to about 12% after 4 hours after exposure to the pH 7.4 release media.Example VI - In Vivo Studies of pH Controlled Microparticles Remote Loaded with Setmelanotide

[0146] All animal studies were performed in accordance with the University of Michigan Committee on Use and Care of Animals (IACUC) protocol numbers PR000007890 and PR000009970. All mice (C57BL / 6J- 00664 Jackson Laboratories) for efficacy evaluation were received at 4 weeks of age and immediately placed on a 60% high-fat diet (Research Diets #12047). At 12 weeks of age, mice were single housed to assure accurate measurements of food intake and body weight. Since mice display an acute-social isolation anorexia, the mice were monitored for one week while receiving daily subcutaneous injections of saline for the mice to be acclimated to subcutaneous injection prior to administration of the microspheres. Following acclimation, mice weighing between 26-38 grams were randomly assigned groups. All animals received either a single dose of blank (non-peptide-loaded) PLGA microspheres (n=6), or a single dose of setmelanotide-loaded PLGA microspheres at a dose of 40 mg / kg setmelanotide (n=6). All microspheres were suspended effectively in the aqueous injection vehicle containing 0.5% low viscosity carboxymethyl cellulose (CMC), 0.1% w / v Tween® 80, and 5% D-mannitol, and subcutaneously administered into the back of the mice using a 19-gauge needle at a dose of 40 mg / kg setmelanotide. At predetermined time points (Day 0, 3, 6, 10, 13, 17, 21, 24, 27, 30), the body weight and food intake of every mouse was recorded.

[0147] Mice were procured, fed on a high-fat diet, and pre-conditioned with saline injections as described above with respect to the efficacy study. Setmelanotide-loaded PLGA microspheres (Sample S1) were suspended in the above injection vehicle and subcutaneously administered into the flank of the mice (n=5 or 6) at a dose of 40 mg / kg setmelanotide using a 19-gauge needle. At predetermined time points (Day 1 , 3, 10, 17, 24, 30, 42) blood samples were drawn from the jugular vein or through cardiopuncture into BD Microcontainer® lithium heparin plasma tubes. The blood samples were centrifuged at 13,000 xg for 10 min at 4 °C for plasma separation, which was then stored at -80 °C before extraction and analysis by LC-MS / MS (see below). In addition, a pharmacokinetic evaluation of free setmelanotide was conducted (n=4) to determine useful PK parameters of setmelanotide in the same murine model for deconvolution of the PK profile after microsphere injection. To accomplish this, setmelanotide solution in PBS was administered to the diet induced obese mice without pre-condition by a single intravenous (i.v.) injection at a dose of 15 mg / kg. Plasma levels were collected at 0.25, 0.5, 1, 1.5, 2, 4, 7, 11, 16h and blood sampling was as above.

[0148] Setmelanotide (MW 1117.2 Da) was extracted from plasma after protein precipitation and quantified using liquid chromatography tandem mass spectrometry (LC-MS / MS, Shimadzu 20A HPLC coupled with AB Sciex 5500 Qtrap). Briefly, 30 pL of plasma samples were dispensed into a 96-well plate and then mixed with either 30 pL setmelanotide solution in acetonitrile for calibration standards or 30 pL acetonitrile for sample analysis. All samples were then spiked with 120 pL of cold acetonitrile consisting of internal standard (IS, CTX1228, 100 ng / mL in acetonitrile). The mixture was vortexed for 10 min and then centrifuged at 3,500xg for 10 min. The supernatant was transferred to another injection 96- well plate and injected into LC-MS / MS with an injection volume of 5 pL. Samples were separated using Agilent Poroshell 120 SB-C18 column (4.6 mm x 50 mm, 2.7 pm) through a gradient elution with a 5.5 min run time. A defined peptide compound of similar molecular weight, CTX1228, was used as the internal standard. The concentration-time profile of setmelanotide i.v. injection was fitted by a two-compartment model with the pharmacokinetic parameters estimated in Table 3 for the deconvolution.Table 3. Pharmacokinetic parameters of setmelanotide after intravenous (i.v.) administration in diet-induced obese mice (n=4, mean ± SD).

[0149] The in vivo plasma concentration-time profile of setmelanotide-loaded PLGA microspheres following a single subcutaneous (s.c.) injection in mice at a dose of 40 mg / kg is shown in Figure 13. Pharmacokinetic parameters were estimated by using noncompartmental analysis in the WinNonLin module (Table 4). Overall, the remote-loaded formulation elevated the plasma concentration of setmelanotide for 6 weeks, as predicted by the corresponding in vitro release profiles. Considering the fast elimination rate of setmelanotide in mice (Table 3, Figure 14), the remote-loaded PLGA microsphere showed its ability to deliver setmelanotide for multi-weeks with a Tmaxat day 17 (Table 4, Figure 13). The plasma concentration profile exhibited a delayed biphasic profile with a short initial lag period and a delayed absorption peak.

[0150] The cumulative setmelanotide absorbed in vivo after single subcutaneous injection as a function of time was then determined through deconvolution. The role of deconvolution is to characterize drug input rate. For this purpose, drug disposition needs to be defined by using some unit impulse response, e.g., pharmacokinetic profile of an intravenous injection or an immediate release injection (Tomic, I. et al. European Journal of Pharmaceutics and Biopharmaceutics 125, 21-27 (2018)). In this study, disposition function was estimated after the i.v. administration of setmelanotide solution.Table 4. Pharmacokinetic parameters of setmelanotide-loaded microspheres after subcutaneous injection in mice (mean ± SEM, n = 5 or 6 for each time point).

[0151] The deconvolved fraction of setmelanotide absorbed-time profile using the Loo- Riegelman method showed a lag phase of 3 days, followed by continuous release for 27 days, and the release plateau was reached by day 42 (Figure 15). The fraction of setmelanotide absorbed in vivo from PLGA microspheres appeared to be accelerated with a tso (time to reach 50% release) of around 18.7 days post injection and 94.1% absorption after 42 days (both values estimated from sigmoidal nonlinear regression of Figure 15), as compared with the half-life of in vitro cumulative release profile tso 24.1 ± 0.7 days and 95.7 ± 0.8% release after 42 days (both values estimated from sigmoidal nonlinear regression of Figure 12).

[0152] This difference may come from the faster hydrolysis and subsequent erosion of PLGA microspheres in vivo owing to the lower pH condition surrounding the microspheres and biological components in the fluid (e.g., enzymes) at the subcutaneous injection site (Shen, J. et al. J. Controlled Release, 218, 2-12 (2015)). Greater water uptake in vivo may also contribute to the creation of new porous networks and osmotic pressure, increasing aqueous diffusion through pores (Doty, A. C. et al. J. Controlled Release, 256, 19-25 (2017)).

[0153] Despite the difference in release rate and initial burst, the overall slow, continuous setmelanotide in vivo absorption pattern is consistent with the cumulative release of setmelanotide in vitro (time scaled, Figure 15). The linear correlation between in vitro release and in vivo absorption strongly suggests a similar release mechanism both in vivo and in vitro. The remote-loaded formulation (Tmax at day 17) offers distinct advantages in pharmacokinetics relative to current subcutaneous daily injections (median Tmax = 8h in human). The prolonged continuous release demonstrated with the setmelanotide remote loaded pH controlled microparticles effectively mitigated fluctuations in setmelanotide concentration that current pulsed dosing regimens face. Consequently, the demonstrated microparticles maintain setmelanotide concentration consistently within the optimal range over an extended period (42 days).Example VII - In-Vivo Study of Administration of pH Controlled Microparticles Remote Loaded with Setmelanotide

[0154] The optimized remote loaded long-acting setmelanotide PLGA microspheres (Sample S1) was then administered to mice on 60% high-fat-diet (HFD) at a dose of 40mg / kg to study the long-term efficacy. This 1 -month dosing regimen was based on a prior investigation in which the mice were infused with 1200 nmol / kg / day (1.34 mg / kg / day) of setmelanotide using an implanted subcutaneous osmotic pump (Collet, T.-H. et al. Mol. Metab. 6(10) 1321-1329 (2017)). The average release rate of setmelanotide from PLGA microspheres, calculated from Figure 15, is approximately 0.91 mg / kg / day. To assure accurate measurements of food intake and body weight, the mice were single housed one week prior to the microsphere injection, in-tandem with 1-week daily subcutaneous injections of saline for mice to be acclimated to subcutaneous injections. Body weight and food intake were monitored for 30 days. The initial body weights of the vehicle-treated group and the setmelanotide-treated group showed no significant difference (data not shown, unpaired t test, P = 0.2024).

[0155] All drug-free microsphere-treated mice (vehicle-treated controls) gained weight following a high fat diet, while mice treated with setmelanotide-loaded PLGA microspheres exhibited reduced weight relative to the starting baseline until day 21, and reduced weight relative to controls through the end of the study at day 30 (Figure 16). The body weight change profile of the mice receiving the setmelanotide-loaded microspheres was significantly different from that of the negative control group dosed with drug-free microspheres until day 24 (Figure 16, repeated measures two-way ANOVA, P < 0.05 for day 3-24). Figure 17 further indicates that the body weight control of setmelanotide-loaded microspheres appeared to be at least partially mediated by suppressing appetite in comparison to control group food intake, although the effect gradually tailed off after 17 days. Even though the suppression of food intake was temporary, body weight control was persistent over 3 weeks, which is consistent with previous studies in diet induced obese C57BL / 6J mice and rhesus macaques, where transient decreases in food intake were also accompanied with persistent loss of body weight (Kumar, K. G. et al. Peptides 30(10), 1892-1900 (2009); Kievit, P. et al. Diabetes 62(2), 490-497 (2013)). The transient appetite suppression might be related to tachyphylaxis, i.e. , diminishing response to successive exposure of drugs to protect the body, which is frequently noticed in mice in response to continuously administered anorectic drug molecules such as melanocortin peptides (Kumar, K. G. et al. Peptides 30(10), 1892- 1900 (2009); Marsh, D. J. et al. Nat. Genet. 21(1), 119-122 (1999)). Therefore, the suppression of food intake possibly accounted for the initial body weight change but did not elucidate the persistent effect on body weight.

[0156] It is possible that a change in metabolic activity also contributed to long-term body weight control (Kievit, P. et al. Diabetes 62(2), 490-497 (2013)). Interestingly, the difference between average percentage body weight change of the vehicle-treated group and that of setmelanotide-loaded microsphere group (i.e., A control - A Setmelanotide) showed anincreasing trend followed by a gradual decline, with the biggest difference occurring at day 17. The kinetics of this difference correlated well with the bell-shaped logarithm (plasma concentration)-time profile in vivo where Tmax was also located at day 17 (Figure 18, Table 4), indicating that the body weight control effect defending HFD is proportional to the magnitude of drug concentration in the blood.

[0157] Overall, setmelanotide-loaded microspheres are capable of treating obese mice as evidenced by i) a sustained control of absolute body weight over 21 days (Figure 21); ii) a reduction in food intake over 17 days when comparing to controls (Figure 17); and iii) an improvement of body weight control over vehicle-treated mice for the entire 30 days (Figure 18 which offsets for the natural weight gain due to HFD over time).Example VIII - Preparation & Remote Loading of Nanoparticles

[0158] PLGA nanoparticles were prepared by a single water-oil (W / O) emulsion method. Uncapped PLGA (75 / 25) at 3.5% w / v was dissolved in dichloromethane and 5% w / v PVA was added prior to single emulsification using a homogenizer. The resulting emulsion was then transferred into a 0.5% PVA bath and mixed for 3 h to allow for solvent evaporation. After mixing, nanoparticles were centrifuged for collection and washed 3 times with water before freeze drying for at least 24 h.

[0159] The resulting blank PLGA nanoparticles were then incubated with an aqueous therapeutic agent loading solution at 37 °C for 24 h to form nanoparticles loaded with therapeutic agent. These loaded nanoparticles were centrifuged for collection, washed 3 times with water, and freeze dried. The loading percentage and encapsulation efficiency for the loaded nanoparticles was determined by a multiple extraction protocol and UPLC. At predetermined time points, samples were centrifuged at 8000 rpm for 5 min and 0.4 mL was collected for analysis by UPLC. The loading, encapsulation efficiency, and initial pH for each sample were measured and are presented in Table 5. The nanoparticles samples were loaded with MOG 38-50, except for RL3 which was loaded with NRPA7.Table 5. Remote loaded nanoparticle loading and encapsulation efficiencies for net positively charged peptides, MOG 38-50 and NRPA7.Example IX - Study of pH Controlled Remote Loading of Microparticles with Setmelanotide under Varying Buffer Concentration & In Vitro Study of Initial Burst Release Profile

[0160] Pre-formed blank PLGA microparticles were prepared using PLGA with different weight average molecular weights and different porosities. Specifically, three remote loaded PLGA microsphere formulations (A, B, and C) were prepared from PLGA (50 / 50, 7-17 kDa or 5-10 kDa) and 200 pL or 300 pL, aliquots of a trehalose porosigen solution (500 mg / mL). Similar to the results shown in Example II, the use of a higher porosigen concentration during microparticle formation results in higher porosity in the formed microparticles. The PLGA and amount of trehalose added are presented in Table 6.Table 6. Microparticle formulations with PLGA molecular weight used and amount of trehalose added.

[0161] Microparticles (180 mg / mL) of each formulation were incubated in a solution of setmelanotide (20 mg / mL) for 24 h at 37 °C in different concentrations of HEPES buffer solution. Specifically, the concentration of HEPES buffer was varied to 0.05 M, 0.1 M, 0.2 M, and 0.3 M. Prior to remote loading, the setmelanotide loading solution had a pH of 7.4. The drug loading and encapsulation efficiency percentages for formulations A, B, and C were observed and are shown in Figure 19.

[0162] The pH values of the initial loading solution and final supernatant were measured using a Fisher Scientific Accument® AE150 benchtop pH meter. The pH drop was calculated by subtracting the pH of the final supernatant from the initial pH of the loading solution. The pH drop and initial burst release for each sample are presented in Table 7 and shown in Figure 20.Table 7. Remote loaded microparticle loading solution pH drop values and initial burst release for the net positively charged peptide, setmelanotide.ainitial burst release describes the amount of drug released from t = 0 days to t = 7 days;bas measured per 10 mg of microspheres.

[0163] For each formulation, the pH drop of the loading solution decreased as the concentration of HEPES buffer increased. For example, each formulation in Table 7 exhibited a pH drop of greater than 2 pH units when the concentration of HEPES buffer in the loading solution was 0.05 M. As the concentration of HEPES buffer in the loading solution was increased to 0.3 M, each formulation in Table 7 exhibited a relatively lesser pH drop of 1.2 pH units or less. This is consistent with the understanding that as the buffer concentration decreases, the capacity of the buffer to resist pH changes is diminished. As a result, the loading solution with a lower concentration of buffer, such as 0.05 M HEPES, is less resistant to pH changes than the loading solution with a higher concentration of buffer.

[0164] Further, formulation C exhibited a relatively larger pH drop than formulations A and B across all four concentrations of HEPES buffer studied. Additionally, formulations A and B exhibited similar drops in pH values after loading across all four concentrations of HEPES buffer studied, despite the different porosities for the microparticles of formulations A and B. All three formulations were observed to have increased drug loading and encapsulation efficiencies as the concentration of the HEPES buffer increased (Figure 19). In particular, formulation C was shown to approach 100% encapsulation efficiency (the particles were substantially saturated with setmelanotide), while formulations A and B were shown to approach 75% encapsulation efficiency, as the concentration of the HEPES buffer increased from 0.05 M to 0.3 M.

[0165] Without being bound by theory, the drop in pH is attributed, in part, to uncapped, and acidic, polymer chain ends associating with the peptide through ion-pairing. As the encapsulation efficiency increases, the ion-pairing between the polymer chain ends and peptides also increases, which increases the amount of acid generated in the release media. Moreover, at a substantially constant particle mass, the number of polymer chains in each particle is expected to increase as the molecular weight of the polymer decreases. As the number of polymer chains in each particle increases, the number of polymer chain ends in each particle will also increase. Further, since the pH drop after loading is attributed to uncapped, and acidic, polymer chain ends associating with the peptide through ion-pairing, it follows that as the number of polymer chain ends increases, the pH drop after loading will also increase. Therefore, the increased pH drop observed for formulation C as compared to formulations A and B is due to the near 100% encapsulation efficiency shown for formulationC and may further be the result of a larger number of acidic polymer chain ends in the lower molecular weight PLGA.

[0166] The initial burst release profiles of the remote loaded formulations A, B, and C were also characterized in PBS release medium after 7 days to evaluate the effect of buffer concentration used in the loading solution on the initial burst release profile and are shown in Figure 20. However, the release data for formulation C with a buffer concentration of 0.3 M was omitted as the loading % observed was substantially similar to the loading % observed for formulation C with a buffer concentration of 0.2 M, which indicated that encapsulation efficiency of the particles was near 100% and believed to be at the upper limit.

[0167] As shown in Figure 20, the cumulative amount of setmelanotide released after day 7 decreased as the pH drop increased, across all formulations. For example, formulations that were loaded in the 0.3 M HEPES buffer solution exhibited pH drops of approximately 1.0 and cumulatively released approximately 150 pg of setmelanotide into the PBS release medium, while formulations that that were loaded in the 0.05 M HEPES buffer solution exhibited pH drops of approximately 2.3 and released approximately 50 pg setmelanotide into the PBS release medium.

[0168] Formulations A and B were shown to release a lower cumulative amount of setmelanotide after day 7 as compared with formulation C. Since formulations A and B included different amounts of porosigen, yet displayed similar initial burst release, the data shown in Figure 20 and Table 1 indicated that the presence of differing levels of pores in the PLGA matrix had a negligible impact on the initial burst release profile in vitro. Rather, the presence of differing concentrations of buffer in the loading solution was shown to have an impact on the initial burst release profile in vitro.

[0169] Therefore, the loading and encapsulation efficiency %’s and initial burst release studies using different concentrations of buffer corroborate our understanding that increased porosity is found in microspheres in the first state and that the porosity can be modulated downwardly by changing the pH around the microspheres in a way that can affect loading as well as release.Example X - Preparation of Microparticles with Varying PLGA Molecular Weight & Remote Loading with Setmelanotide

[0170] Pre-formed blank PLGA microparticles were prepared with varying PLGA polymers to evaluate the effect of adjusting the weight average molecular weight and ratio of lactic to glycolic acid (L:G ratio) in the PLGA matrix on remote loading of setmelanotide. Specifically, five remote loaded PLGA microsphere formulations (F4, F15, F26, F27, and F17) were prepared from PLGA (50:50, Evonik Resomer® RG502H, 7-17 kDa), PLGA (50:50, CDbioparticles CDP081 , 5-10 kDa), PLGA (60:40, CD bioparticles CDP068, 5-10 kDa), PLGA (70:30, CD bioparticles CDP258, 5-10 kDa), and PLGA (75:25, CD bioparticles CDP073 , 5- 10 kDa), respectively. Microparticles (180 mg / mL) of each formulation were incubated in a solution of setmelanotide (20 mg / mL) in 0.2 M HEPES buffer at pH 7.4 for 24 h at 37 °C.

[0171] The pH values of the initial loading solution and final supernatant were measured using a Fisher Scientific Accument® AE150 benchtop pH meter. The pH drop was calculated by subtracting the pH of the final supernatant from the initial pH of the loading solution. The pH drop and setmelanotide loading percentage for each formulation are presented in Table 8 and shown in Figure 21.Table 8. Remote loaded microparticle loading solution pH drop values and drug loading for the net positively charged peptide, setmelanotide.N = 1 for formulations F4 and F17.

[0172] After remote loading, the loading solutions from formulations F4, F15, F26, and F27 exhibited pH drops of 1.0 or greater, while formulation F17 had a pH drop of 0.32. However, formulations F15, F26, and F27 were found to be loaded with 9.88 %, 9.19 %, and 9.88 % setmelanotide respectively, while formulations F17 and F4 were found to be loaded with 7.36 % and 6.60 % setmelanotide peptide, respectively. Thus, while formulation F4 was observed to have a relatively higher pH drop as compared to the other formulations tested, formulation F4 was also shown to have the lowest setmelanotide loading percentage of any formulation tested. Further, formulations F15 and F27 were shown to approach 100% encapsulation efficiency (Figure 21). Since the formulations including 5-10 kDa molecular weight PLGA (F15, F26, F27, F17) had higher setmelanotide loading percentages than the formulation with 7-17 kDa PLGA (F4), the data shown in Table 8 and Figure 21 suggest that the molecular weight of the PLGA included in the formulation has a significant impact on remote loading of setmelanotide.

[0173] The pH drop after remote loading was evaluated by comparing the results observed for formulations F4 and F15. As shown in T able 8, F15 was shown to have a pH drop of 1.68 and a setmelanotide loading of 9.88 ± 0.05%, while formulation F4 was shown to have a pH drop of 1.53 and a setmelanotide loading of 6.60%. These findings indicated that formulation F4 exhibited a lower setmelanotide loading than formulation F15 as well as a smaller pH drop after remote loading.

[0174] The L:G ratio of the PLGA matrix was evaluated by comparing the results observed for formulations F15, F26, F27, and F17. As shown in Table 8, the L:G ratio of the PLGA matrix had a significant impact on pH drop and a negligible impact on peptide loading. For example, formulations F15, F26, F27, and F17 show that as the L:G ratio increases from 50:50 to 75:25, that the pH drop of the loading solution decreases from 1.68 to 0.32. Thus, as the relative content of lactic acid in the PLGA matrix decreases, the pH drop observed in the corresponding PLGA particles is expected to increase in a given incubation media, i.e., the pH drop is expected to increase as the L:G ratio decreases.

[0175] Meanwhile, setmelanotide loading percentages do not show a clear correlation with the L:G ratio, where formulation F17, which has the highest L:G ratio of 75:25, was shown to have a lower setmelanotide loading percentage than formulations F15, F26, and F27, which had lower L:G ratios and higher setmelanotide loading percentages. Therefore, the L:G ratio does not have a significant impact on peptide loading.

[0176] Without being bound by theory, the relative amounts of polymer chain ends vs. bulk polymer are also expected to affect the PLGA matrix. At a substantially constant particle mass, the number of polymer chains in each particle is expected to increase as the molecular weight of the polymer decreases. As the number of polymer chains in each particle increases, the number of polymer chain ends in each particle will also increase. Further, since the pH drop after loading is attributed to uncapped, and acidic, polymer chain ends associating with the peptide through ion-pairing, it follows that as the number of polymer chain ends increases, the pH drop after loading will also increase. Therefore, the increased pH drop observed for formulation F15 as compared to formulation F4 may be attributable to a relatively larger number of acidic polymer chain ends in the lower molecular weight PLGA matrix of formulation F15.Example XI - In Vitro Study of pH Controlled Microparticles with Varying PLGA Molecular Weight & Remote Loaded with Setmelanotide

[0177] The five formulations from Example X were tested in vitro to characterize the longterm release profile in PBS release medium for 8 weeks. As shown in Figure 22, formulations F15, F26, F27, and F4 showed triphasic release behavior with a small initial burst release, followed by a lag phase of around 1 week during which substantially little peptide is released. Following this lag phase, setmelanotide release was continuous for 1-2 weeks, and was complete by day 28 for formulations F15, F26, and F4, and day 49 for formulation F27. Formulation F17 also showed small initial burst release but did not complete setmelanotide release after 56 days. Accordingly, formulation F27 was shown to have high setmelanotide loading, near 100% encapsulation efficiency, (99% observed), loading of 9.9%, and sustained release over 5 weeks with low initial burst. This wassurprising, as it is well known in the art that relatively lower weight average molecular weight PLGA (less than 10 kDa) is expected to have a much faster release, particularly in the first week.

[0178] Initial burst release is affected by the dissolution of immediately available peptides near the surface, peptides in the existing pores as well as polymer pore healing. Without being bound by theory, the low initial burst release (less than 5% in day 1 for formulations F15, F26, F27, and F4 ) of setmelanotide observed may be due to an ion-pairing association of peptide with two distinct polymer chains within the swollen polymer phase (corresponding to the first state of the particles). Therefore, the release profile supports the understanding that setmelanotide was successfully absorbed into the polymer phase via remote loading (which enables sustained release for long-term administration of the therapeutic agent), rather than merely adsorbed onto functional groups at the surface of the microsphere or dwelling in a porous channel thereof.

[0179] After remote loading peptide, some passive polymer healing, i.e. , which was visibly confirmed by rounding and / or disappearance of tiny holes, was observed via SEM and shown in Figures 23(a-c) for formulations F17, F27, and F26, respectively. The particle size and size distribution had negligible change before and after the incubation of blank microspheres in peptide loading solution. Further, formulation F17, which was shown to have higher initial release (> 5%) than formulations F27 and F26, showed less polymer healing after remote loading, than formulations F27 and F26. Therefore, the SEM images and initial release study indicate that upon exposure to a solution at the second pH, the surface area of the particle decreases as the pores and channels of the particle become less interconnected, thereby trapping the peptide in the “closed” pores and channels of the particle in the second state through polymer self-healing. Further, these data indicate that when the particles in the second state are exposed to a solution at the first pH, the particles transition from the second state to the first state, promoting the formation of interconnected pores and thereby promoting the release of the loaded peptide.Example XII - In Vitro Study of pH Controlled Microparticles Remote Loaded with Leuprolide & Suspended in pH Buffered Hydrogels

[0180] The medium porosity formulation (L2) from Example V was tested in vitro to characterize the burst release in PBS release medium over 2 h by adjusting the procedure described in Example IV. The remote loaded microparticles were suspended in hydrogels buffered to different pH values, prior to exposure to the PBS release medium. Immediately following suspension in the pH buffered hydrogel, the PBS release media (12 mL, pH 7.4) was gently layered on top of the hydrogel. PLGA microparticles without gel were used as control (PBST pH 7.4).

[0181] Hydrogels were prepared from mixtures of water with a poloxamer where x is 101 and y is 56 and having the following structure:and was commercially available from Sigma Aldrich (Pluronic® F-127).

[0182] The amount of pH buffered hydrogel provided was tested at 15% and 18%. At each of these concentrations, the pH values of the pH buffered hydrogels on initial release were varied at pH 7.4, 6.5, and 5.5 using microparticles L2 and the release was measured after 2 h and 4 h and are reported in Table 9. The cumulative release of L2 in the 15% hydrogel formulation at each pH after 2 h and 4h are presented in Figures 24 and 25, respectively. The cumulative release of L2 in the 18% hydrogel at each pH after 2 h and 4h are presented in Figures 26 and 27, respectively. For all formulations suspended in hydrogel, the initial release was less than 6% across all pH values tested after 2 h, and less than 12% across all pH values tested after 4 h, while the formulation with only PBST buffer at pH 7.4 exhibited a burst release of about 11% after 2 h and about 12% after 4 h.Table 9. Remote loaded microparticle initial release % and for the net positively charged peptide, leuprolide, after 2 and 4 hours in release media.

[0183] For the 15% hydrogel formulation, the burst release % was about 5% when the pH of the hydrogel formulation was 5.5, while the burst release % was 4% or less when the pH of the hydrogel formulation was raised to 6.8 or more, after 2 h in the PBS release medium (Figure 24). After 4 h in the PBS release medium, the burst release % was about 9% when the pH of the 15% hydrogel formulation was 5.5, while the burst release % increased to 10% or more when the pH of the 15% hydrogel formulation was raised to 6.5 or more (Figure 25).

[0184] For the 18% hydrogel formulation, the initial release % was less than 1% when the pH of the hydrogel formulation was 6.5 or below, while the initial release % was about 1% when the pH of the hydrogel formulation was raised to 7.4, after 2 h in the PBS release medium (Figure 26). After 4 h in the PBS release medium, the initial release % was less than 0.3% when the pH of the 18% hydrogel formulation was 5.5, while the initial release % was less than 1.5% when the pH of the 18% hydrogel formulation was raised to 6.5 or more (Figure 27).

[0185] Additionally, Figure 28 shows that formulations including leuprolide were prepared from hydrogels without microparticles released significantly more leuprolide after 2 h than formulations including microparticles. In particular, the hydrogel formulations without microparticles were shown to have an initial release of about 45% at pH 7.4, which increased as the pH of the hydrogel decreased to about 70% at pH 5.5 (Figure 28). After 4 h, all formulations without microparticles were observed to have full release of leuprolide. Therefore, it is believed that at pH 5.5, the hydrogel is more permeable than at pH 7.4. Accordingly, the increased release % observed for the 15% hydrogel formulation is believed to be caused by the increased permeability of the hydrogel at pH 5.5 as compared to the 15% hydrogel formulation at pH 6.5 or 7.5.

[0186] These experiments demonstrate that suspending the microparticles in a hydrogel formulation before exposing to the release medium was shown to have an impact on the initial release profile in vitro. The 18% hydrogel formulation exhibited lower release as compared to the 15% hydrogel formulation after 2 h and 4 h, at each pH value tested. Buffering the hydrogel gel to a specific pH was also shown to impact release behavior. The hydrogel formulations buffered to higher pH values approaching 6.8 and greater consistently exhibited larger release % between 2 h and 4 h as compared to the formulations with hydrogel gel buffered to lower pH values.

[0187] Further, these experiments demonstrate that the release profile of the microparticles can be affected by several factors including the concentration of the buffer in the loading solution and the pH of the release media through influencing the pH of the loading solution after remote loading and during release. Since the particles adopt the first and second states upon exposure to the first and second pH values, respectively, control of the pH in the incubation solution after remote loading dictates whether the microparticles adopt the second state after remote loading.

[0188] Regarding buffer concentration in the loading solution, the drop in pH of the loading solution after remote loading was shown to decrease and the initial release % of the microparticles was shown to generally increase as the concentration of the buffer increased.

[0189] In particular, the formulations prepared in loading solutions (which can become solutions at a second pH as described herein) which did not have pH values below 6.5 after remote loading demonstrated significant initial release, while the formulations prepared in loading solutions which did have pH values below 6.5 after remote loading demonstrated relatively low initial release. Further, the formulations prepared in loading solutions which had drops in pH of roughly 1.0 or more demonstrated relatively little initial release (< 5%) compared to formulations prepared in loading solutions observed to have changes in pH of less than 1.0, which generally demonstrated more significant initial release. These trends corroborate our understanding that the particles adopt the first state when the particles are exposed to the first pH, adopt the second state when the particles are exposed to the second pH, and a sufficient change in pH is required for the particles to effectively transition between the first state and the second state and provide the advantageous combination of both sustained release and low initial burst.

[0190] Regarding the pH of the release media, the release % of the microparticles was shown to generally increase as the pH of the release media increased to pH 6.8 or greater, which corroborates the understanding that the particles adopt the first state when exposed to a pH of 6.8 or greater. These trends also corroborate our understanding that when the particles in the second state are exposed to a solution at the first pH, the particles at some point transition from the second state to the first state which promotes release of the therapeutic agent but only after a lag time during which the transition occurs, thereby still avoiding an initial burst release. Accordingly, the concentration of the buffer in the loading solution and the pH of a second incubation solution (e.g., a release media), can be adjusted to control the remote loaded microparticle formulations that can provide sustained release of net positively charged therapeutic agents with low initial release.

Claims

What is claimed is:

1. A method of preparing a particle for sustained release of a therapeutic agent, comprising: providing one or more biodegradable particles comprising a polymer matrix, wherein the polymer matrix comprises an uncapped polymer comprising PLGA or PLA; incubating the one or more biodegradable particles with a therapeutic agent in a loading solution at a first pH between about 6.8 and about 8.0; and incubating the one or more biodegradable particles in a solution at a second pH between about 4.5 and about 6.5; wherein: the uncapped polymer comprises free carboxyl groups such that the biodegradable particle includes free carboxyl groups; the particle has an average particle diameter in the range of about 10 nm to about 100 pm; and the therapeutic agent comprises a protein or peptide with a net positive charge in the loading solution at the first pH.

2. The method of claim 1 , further comprising preparing the solution at the second pH by adding an acid to the loading solution.

3. The method of claim 2, wherein the acid comprises phosphoric acid, trifluoroacetic acid, or acetic acid.

4. The method of claim 1 , further comprising: collecting and drying the one or more biodegradable particles, after incubating the one or more biodegradable particles at the first pH; and providing the one or more biodegradable particles in the solution at the second pH, thereby incubating the one or more particles at the second pH.

5. The method of any one of claims 1 to 4, wherein the therapeutic agent comprises a protein or peptide with a net positive charge greater than about +1 , for example, a protein or peptide with a net positive charge greater than about +1.5 or a protein or peptide with a net positive charge greater than about +2, in the loading solution at the first pH.

6. The method of any one of claims 1 to 5, wherein the biodegradable particle has a therapeutic agent content in the range of about 2% weight percent (wt.%) to about 20% wt.%, based on the entire weight of the biodegradable particle, after incubating in the solution at the second pH.

7. The method of any one of claims 1 to 6, wherein the encapsulation efficiency of the therapeutic agent is in the range of about 40% to about 100%, after incubating in the solution at the second pH.

8. The method of any one of claims 1 to 7, wherein the biodegradable particle has an initial burst release of the therapeutic agent of about 5% or less, after 24 hours in a phosphate-buffered saline solution as determined by the amount of therapeutic agent released from t = 0 hours to t = 24 hours in a phosphate-buffered saline buffer at a pH of about 7.4.

9. The method of claim 1 , wherein the one or more biodegradable particles are prepared by a process comprising: dissolving the uncapped polymer in an organic solvent thereby forming a solution; adding polyvinyl alcohol to the solution, thereby forming the polymer matrix; and removing the organic solvent, thereby forming the plurality of biodegradable particles.

10. A formulation for sustained release of biodegradable particles containing a therapeutic agent, comprising: a plurality of biodegradable particles comprising a polymer matrix, wherein the polymer matrix comprises an uncapped polymer comprising PLGA or PLA; a therapeutic agent absorbed and encapsulated by the polymer matrix; and a solution with a pH of about 4.5 to about 6.5; wherein: the biodegradable particle has an average particle diameter in the range of about 10 nm to about 100 pm; the therapeutic agent comprises a protein or peptide with a net positive charge in the solution.

11. The formulation of claim 10, wherein the therapeutic agent comprises a protein or peptide with a net positive charge greater than about +1 , for example, a protein or peptide with a net positive charge greater than about +1 .5 or a net positive charge greater than about +2, in the solution.

12. The formulation of claims 10 or 11 , wherein the biodegradable particle has a therapeutic agent content in the range of about 2% weight percent (wt.%) to about 20% wt.%, about 4 wt.% to about 18 wt.%, about 6 wt.% to about 16 wt.%, about 8 wt.% to about14 wt.%, or about 10 wt.% to about 12 wt.%, based on the entire weight of the biodegradable particle.

13. The formulation of any one of claims 10 to 12, wherein the biodegradable particle has an initial burst release of the therapeutic agent of about 5% or less, after 24 hours in the solution as determined by the amount of therapeutic agent released from t = 0 hours to t = 24 hours in a phosphate-buffered saline buffer at a pH of about 7.4.

14. The formulation of any one of claims 10 to 13, wherein the biodegradable particle has a particle diameter in the range of about 10 pm to about 100 pm, about 15 pm to about 90 pm, or about 20 pm to about 80 pm.

15. The formulation of claim 14, wherein the biodegradable particle is a microsphere.

16. The formulation of claim 14 or 15, wherein the therapeutic agent is setmelanotide or leuprolide.

17. The formulation of any one of claims 10 to 13, wherein the biodegradable particle has a particle diameter in the range of about 30 nm to about 10 pm, about 100 nm to 1 pm, and / or about 500 nm to about 900 nm.

18. The formulation of claim 17, wherein the biodegradable particle is a nanosphere.

19. The formulation of claim 17 or 18, wherein the therapeutic agent is a MOG peptide or NRPA7.

20. A method of preparing a biodegradable particle for sustained release of a therapeutic agent, comprising: providing one or more biodegradable particles comprising a porous polymer matrix, wherein the polymer matrix has a first state with relatively more interconnected pores and a second state with relatively less interconnected pores; incubating the one or more biodegradable particles with a therapeutic agent in a loading solution at a first pH between about 6.8 and about 8.0, thereby encapsulating the therapeutic agent in the first state of the porous polymer matrix; and incubating the one or more biodegradable particles in a solution at a second pH between about 4.5 and about 6.5, thereby transitioning the polymer matrix from the first state to the second state; wherein: the polymer matrix comprises an uncapped polymer comprising PLGA or PLA and including free carboxyl groups such that the biodegradable particle includes free carboxyl groups;the therapeutic agent comprises a protein or peptide with a net positive charge in the loading solution at the first pH; and the biodegradable particles have an initial burst release of the therapeutic agent of about 5% or less after 24 hours when the biodegradable particles are in the second state as determined by the amount of therapeutic agent released from t = 0 hours to t = 24 hours in a phosphate-buffered saline buffer at a pH of about 7.4.

21. The method of claim 20, wherein the therapeutic agent comprises a protein or peptide with a net positive charge greater than about +1 in the loading solution at the first pH.

22. The method of claim 20 or 21 , wherein the therapeutic agent comprises a protein or peptide with a net positive charge greater than about +1.5, or a net positive charge greater than about +2, in the loading solution at the first pH.

23. The method of any one of claims 20-22, wherein the one or more biodegradable particles are prepared by a process comprising: dissolving the uncapped polymer and a porosigen in an organic solvent thereby forming a solution; adding polyvinyl alcohol to the solution, thereby forming the porous polymer matrix; and removing the organic solvent, thereby forming the one or more biodegradable particles.

24. The method of any one of claims 20 to 23, wherein the initial burst release of the therapeutic agent is in the range of about 0% to about 5%, about 1% to about 5%, about 1% to about 4%, about 1 % to about 3%, and / or about 1% to about 2%, after 24 hours when the particle is in the second state.

25. A biodegradable particle prepared according to the process of any one of claims 20 to 24, wherein the particle has an average particle diameter in the range of about 10 nm to about 10 pm, about 10 nm to about 1 pm, about 100 nm to about 10 pm, about 100 nm to about 1 pm, about 10 nm to about 100 nm, and / or about 100 nm to about 10 pm.

26. The biodegradable particle of claim 25, wherein the initial burst release of the therapeutic agent is in the range of about 0% to about 5%, about 1% to about 5%, about 1% to about 4%, about 1 % to about 3%, and / or about 1% to about 2%, after 24 hours when the particle is in the second state as determined by the amount of therapeutic agent released from t = 0 hours to t = 24 hours in a phosphate-buffered saline buffer at a pH of about 7.4.

27. The biodegradable particle of claim 25 or 26, wherein the uncapped polymer comprises PLGA.

28. The biodegradable particle of any one of claims 25 to 27, wherein the uncapped polymer has a lactic acid content in the range of 25% to 100% and a weight average molecular weight in the range of about 2 kDa to about 50 kDa, about 3 kDa to about 45 kDa, about 5 kDa to about 40 kDa, about 7.5 kDa to about 35 kDa, and / or about 10 kDa to 20 kDa.

29. The biodegradable particle of any one of claims 25 to 28, wherein the therapeutic agent has a weight average molecular weight in the range of about 500 Da to about 5 kDa, about 1 kDa to about 4.5 kDa, about 1.5 kDa to about 4 kDa, and / or about 2 kDa to about 3.5 kDa.

30. The biodegradable particle of any one of claims 25 to 29, wherein the therapeutic agent comprises at least two positively charged amino acid residues at the first pH.

31. A formulation for sustained release of biodegradable particles containing a therapeutic agent, comprising: a plurality of particles comprising a porous polymer matrix, wherein the porous polymer matrix comprises an uncapped polymer comprising PLGA and PLA; a therapeutic agent absorbed and encapsulated by the polymer matrix; a media capable of providing a pH of about 4.5 to about 6.5 within an unstirred boundary layer of the particles upon hydration; wherein: the porous polymer matrix has a first state with more interconnected pores and a second state with less interconnected pores; the biodegradable particles transition from the first state to the second state when exposed to a pH between about 4.5 to about 6.5; the biodegradable particles transition from the second state to the first state when exposed to a pH between about 6.8 to about 8.0; the particles are dispersed in the media; and, the therapeutic agent comprises a protein or peptide with a net positive charge at the pH between about 4.5 to about 6.5.

32. The formulation of claim 31 , wherein the therapeutic agent comprises a protein or peptide with a net positive charge greater than about +1 at the pH between about 4.5 to about 6.5.

33. The formulation of claim 31 , wherein therapeutic agent comprises a protein or peptide with a net positive charge greater than about +1.5, or a net positive charge greater than about +2, at the pH between about 4.5 to about 6.5.

34. The formulation of any one of claims 31 to 33, wherein the biodegradable particle has an initial burst release of the therapeutic agent of about 5% or less after 24 hours in a phosphate-buffered saline solution when the particle is in the second state as determined by the amount of therapeutic agent released from t = 0 hours to t = 24 hours in a phosphate- buffered saline buffer at a pH of about 7.4.

35. The formulation of any one of claims 31 to 34, wherein the media comprises a thermoreversible gel.

36. The formulation of claim 35, wherein the thermoreversible gel comprises PEG- chitosan, chitosan-polyethylenimine, chitosan-arginine or glycol-chitosan-spermine.

37. The formulation of any one of claims 31 to 34, wherein the media comprises an acid source coated on the particle.

38. The formulation of claim 38, wherein the acid source has low solubility in water, for example, the acid source has a solubility in water of less than 1 mg / mL or a solubility in the range of about 1 mg / mL to about 10 mg / mL.

39. The formulation of claim 39, wherein the acid source has a solubility in water of less than 1 mg / mL.

40. The formulation of any one of claims 31 to 34, wherein the media comprises a granule or particle capable of releasing an acid source.

41. The method of any one of claims 1 to 9, or the method of any one of claims 20 to 24, wherein: the difference between the first pH and the second pH is about 1 .0 to about 3.0 pH units, about 1 .2 to about 2.8 pH units, or about 1.5 to about 2.5 pH units.

42. The method of any one of claims 1 to 9, or the method of any one of claims 20 to 24, wherein: the solution at the second pH is formed in situ; and the loading solution further comprises a buffer.

43. The method of claim 42, wherein the buffer has a concentration of about 0.01 M to about 0.3 M in the loading solution.

44. The method of any one of claims 1 to 9, or the formulation of any one of claims 10 to 19, or the method of any one of claims 20 to 24, or the biodegradable particle of any one of claims 25 to 29, or the formulation of any one of claims 31 to 40, wherein: the uncapped polymer has a lactic acid content in the range of 50% to 70%, relative to the total amount of lactic acid and glycolic acid in the uncapped polymer and a weight average molecular weight in the range of about 2 kDa to about 20 kDa, about 5 kDa to about 17 kDa, about 5 kDa to about 15 kDa, about 5 kDa to about 10 kDa, about 7 kDa to about 17 kDa, and / or about 8 kDa to about 12 kDa.

45. The formulation of claim 35, wherein the thermoreversible gel comprises a poloxamer having a structure of Formula (I):wherein: x can range from 2 to 130; and y can range from 15 to 67.

46. The formulation of claim 45, wherein the poloxamer is present in in the range of 15% to 20% (m / v).

47. The formulation of claims 45 or 46, wherein x is 70 to 130 and y is 50 to 60.

48. The formulation of claim 47, wherein x is 101 and y is 56.

49. The method of any one of claims 1 to 9, or the method of any one of claims 20 to 24, or the method of any one of claims 41 to 44, wherein: at least one of the loading solution and the solution at the second pH is at a temperature ranging from 30 °C to 55 °C, preferably in the range of 35 °C to 40 °C.

50. The method of any one of claims 1 to 9, or the formulation of any one of claims 10 to 19, or the method of any one of claims 20 to 24, or the biodegradable particle of any one of claims 25 to 29, or the formulation of any one of claims 31 to 40, or the method of any one of claims 41 to 44, or 49 orthe formulation of any one of claims 45 to 48, wherein: the uncapped polymer has a weight average molecular weight in the range of about 5 kDa to about 10 kDa.

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