Since the concept of protein or drug delivery from polymers was first introduced, research efforts have focused on developing polymer formulations that would be widely applicable for delivery of biologically active agents, such as proteins, peptides, oligonucleotides, DNA, low molecular weight drugs and vaccine antigens. Efforts to this end have intensified recently since hundreds of recombinant proteins and other biotechnological drugs and vaccine antigens are in the pipeline for FDA approval, and the current method of protein delivery generally requires injections on a daily basis. Frequent dosing is clinically undesirable due to patient discomfort, psychological distress, and poor compliance for administering self-injections. To reduce injection frequency, peptide and protein drugs are encapsulated in biodegradable polymers, which are processed into a form that is easily administered through a syringe needle. Current preparations on the market for the delivery of small peptides can reduce the frequency of injections to once every 1-3 months depending on the size and dose of the polymer implant. This incubation time, for which a large globular protein must remain encapsulated in the polymer at physiological temperature, poses significant challenges to retain both the structural integrity and the biological activity of the protein.
Two injectable polymer configurations are currently used to deliver peptides and proteins: spherical particles on the micrometer scale (xcx9c1-100 xcexcm), which are commonly referred to as xe2x80x9cmicrospheresxe2x80x9d, and single cylindrical implants on the millimeter scale (xcx9c0.8-1.5 mm in diameter), which we term xe2x80x9cmillicylindersxe2x80x9d. Both configurations are prepared from the biocompatible copolymer class, poly(lactide-co-glycolide) (PLGA) commonly used in resorbable sutures, and each configuration has distinct advantages and disadvantages.
Once injected into the body, these polymer implants slowly release the biologically active agents, thereby providing desirable levels of the agent over a prolonged period of time. Because of its safety, FDA approval and biodegradability, the poly(lactide-co-glycolides) (PLGAs) are the most common polymer class used for preparing biodegradable delivery systems for biologically active agents. Unfortunately, the microenvironment in PLGA surrounding the encapsulated agent can become highly acidic, causing many of these agents to lose their biological activity. Accordingly, it is desirable to modify the methods that are currently used to prepare polymeric delivery systems which liberate acids during biodegradation, such as PLGA, and to thereby produce a polymeric implant that is capable of releasing the biologically active agent over a prolonged period of time and maintaining the stability of the biologically active agent that is retained in the delivery system during nonenzymatic hydrolysis, hereinafter referred to as xe2x80x9cbiodegradationxe2x80x9d of such a system. Such methods would also be useful for preparing implants that are made from polymers that contain acid that slowly dissolves and lowers the pH of the microenvironment surrounding the encapsulated agent
The present invention provides new methods for reducing or inhibiting the irreversible inactivation of water-soluble biologically active agents in biodegradable polymeric delivery systems which are designed to release such agents over a prolonged period of time, such as PLGA delivery systems. In accordance with the present invention, it has been discovered that, in many instances, the acids that are produced during biodegradation of PLGA can induce an irreversible inactivation or instability of biologically active agents, such as for example proteins, drugs, oligonucleotides and vaccine antigens. It has also been determined that the addition of certain antacids, such as for example MgOH2, to the system will not significantly reduce the acid-induced instability of the biologically active unless the polymer is prepared in a manner which results in the formation of an interconnected network of pores within the polymer. It has also been discovered that the acid-induced instability of biologically active agents encapsulated in PLGA delivery can be inhibited or significantly reduced by preparing PLGA delivery systems whose microclimate, i.e. the pores where the active agent resides, uniformly or homogenously maintain a pH of between 3 and 9, preferably between 4 and 8, more preferably between 5 and 7.5 during biodegradation. Depending on the size of the delivery system, i.e., the weight average particle diameter and the initial bulk permeability of the polymer, this result is achieved by (a) incorporating a water-soluble carrier into the delivery system, (b) incorporating a select basic additive (or antacid) into the delivery system, (c) incorporating both a water soluble carrier and a select basic additive into the delivery system, (d) adding a pore forming molecule for increasing the rate of release of low molecular weight monomers and oligomers into the delivery system, (e) using a PLGA polymer with reduced glycolide content, i.e. PLGA with from 100% to 75% lactide and 0 to 25% glycolide) (f) using a microencapsulation method that yields a more extensive pore-network, e.g. oil-in-oil emulsion-solvent extraction as opposed to water-in-oil-in water-solvent evaporation method, and (g) combinations thereof.
The present invention also relates to PLGA delivery systems prepared by the present method. Such delivery systems have a low porosity (e.g.  less than 50%) and a uniform morphology (e.g. spherical or cylindrical usually with smooth or uniformly rough surfaces, and when particulate, all particles are similar in external and internal appearance under the scanning electron microscope. In addition, the PLGA delivery systems of the present invention have a low initial burst release (e.g.  less than 50% of the drug is released during the 1st hour of biodegradation) Most importantly, during biodegradation, the present PLGA delivery systems maintain a relatively homogenous microclimate pH greater than 3 and less than 9, preferably greater than 4 and less than 8, more preferably greater than 5 and less than 7.5, so that less than 15% of the combined released and residual encapsulated test protein bovine serum albumin forms nonconvalent, water-insoluble aggregates when incubated in a physiological buffer solution for 4 weeks at 37xc2x0 C.
In certain embodiments, the PLGA delivery system comprises bone morphogenetic protein-2, vincristine sulfate, fibroblast growth factor, or tissue plasminogen activator.
The present invention provides methods of preparing PLGA delivery systems which stabilize the soluble biologically active agents that are encapsulated therein. As used herein, the term stabilize refers to an improvement in the stability of the encapsulated agent, which is necessary to approach or achieve a stable state. A stable biologically active agent as used herein refers to a biologically active agent such as a protein, peptide, oligonucleotide, low-molecular weight drug, or vaccine antigen that retains at least 80%, preferably 90%, of its original structure and/or biological activity during its release from the PLGA delivery system. During biodegradation of PLGA delivery systems, soluble agents often undergo acid-induced irreversible instability. Such instability may result from noncovalent aggregation of the agent, peptide-bond hydrolysis, deamidation, isomerization, covalent aggregation, deformylation, depurination, etc. Each of these acid-induced physical or chemical alterations can be monitored using standard techniques known in the art. For example, aggregation can be monitored by loss of solubility, SDS-PAGE, and or size-exclusion chromatography.
The methods of the present invention also provide controlled release PLGA delivery systems. As used herein, controlled release means the release kinetics are engineered into the system such that the agent is released in a manner controlled by the system itself or its surroundings, preferably the system itself. Such controlled release requires that the agent is not all released within a short period of time, e.g., less than one hour, after injection or implantation of the system in a subject. Preferably the agent is released from the implanted system over a prolonged period of time, e.g. 3 days to 1 year. In some cases, the delivery system is designed to release the agent slowly and continuously over this prolonged period of time. In other instances the delivery system is designed to release the agent in multiple phases.
Stabilization of the encapsulated agent is achieved by providing a delivery system whose microclimate, i.e. the pores where the active agent resides, uniformly or homogeneously maintain a pH of greater than 3 and less than 8, preferably greater than 4 and less than 8, more preferably from 5 to 7.5 during biodegradation. To determine if the method has provided a polymeric delivery system whose microclimate homogenously maintains a pH of between 3 and 8, 1% w/w BSA is dispersed in the polymer solution during manufacture by the chosen method and the extent of aggreagation of this protein is assayed after 4 weeks of incubation of the polymeric delivery system in phosphate buffered saline with 0.02% Tween 80 at 37xc2x0 C. If the amount of residual BSA that has formed water insoluble noncovalent aggregates (i.e., soluble in 6 M guanidine hydrochloride or 6 M urea) is less than or equal to 15% of the total BSA in the prepared polymer dosage form, the method has produce a polymeric delivery system whose microclimate homogeneously maintains a pH of between 3 and 8.
One method for preparing a delivery system which stabilizes the agent encapsulated therein during biodegradation comprises adding a poorly soluble, mildy strong basic additive to a solution comprising the biologically active agent and the polymer. Except for CaOH2, the basic additive has a solubility and basicity comparable to the solubility and basicity of the compounds shown in Table I below.
Suitable basic additives are magnesium carbonate, magnesium hydroxide, magnesium oxide, magnesium trisilicate, zinc carbonate, zinc hydroxide, zinc phosphate, aluminum hydroxide, basic aluminum carbonate, dihyroxyaluminum sodium carbonate, dihydroxyaluminum aminoacetate, ammonium phosphate, calcium phosphate, calcium hydroxide, magaldrate. Preferably, the polymer comprises from 50% to 100% lactide or lactic acid, which may be a D isomer, L-isomer, or a D-,L-racemic mixture, and from 50% to 0% of a glycolide or glycolic acid. The polymer has an inherent viscosity of from 0.1 to 2.0 dl/g.
The polymer solution comprises from 0.1 to 20% of the biologically active agent or a composition comprising the biologically active agent and a carrier. In those instances where the amount of biologically active agent incorporated into the polymer solution is sufficient to promote formation of an interconnected network of pores, addition of carrier to the polymer solution is optional. In those cases where the amount of bioligically active agent incorporated into the polymer solution is low (e.g., due to cost, toxicity, etc.), it is preferred that a carrier be added. Examples of suitable carriers are albumin, gum arabic, gelatin, dextran, a water soluble amino acid, a monosaccharide, a disaccharide, and combinations thereof.
The polymer solution comprises from 0.5 to 20% of the basic additive. In those cases where the amount of basic additive dispersed in the solution is low, i.e. from 0.5% to 3% w/w, it is preferred that the porosity of the polymeric delivery system be increased. Methods for increasing the porosity include adding a pore-forming agent to the polymer solution, increasing the amount of biologically active agent or the composition comprising the biologically active agent and carrier to a value of 5 to 20% (w/w), or using a low concentration of polymer, e.g. 40-300 mg/ml of polymer in the organic solvent. In those cases where the polymer concentration is high, e.g. 1200 mg/ml or the inherent viscosity is high, it is preferred that the polymer solution comprise from 3 to 20% by weight of the basic additive.
Another method of preparing biodegradable polmeric delivery systems for stabilizing the biologically active agents encapsulated therein involves blending a pore-forming agent with a polymer which comprises from 50% to 100% lactide or lactic acid and from 50% to 0% glycolide or glycolic acid. Examples of suitable pore-forming agents are polyethylene glycol (PEG) and water soluble poloxamers. Preferably, the pore-forming agent has a molecular weight of from 500 to 30,000, more preferably from 4000 to 10,000.
The methods of the present invention are suitable for preparing large delivery systems having a weight average diameter of 5 to 500 mm, intermediate-sized delivery systems having a weight average diameter of 100 to 5000 xcexcm, and small delivery systems having a weight average diameter of from 10 nm to 100 xcexcm. The delivery systems of the present invention encompass spheres, including microspheres and nanospheres, cylinders, including millicylinders, and particles.
When aqueous soluble compounds are encapsulated in PLGA delivery systems, they are typically distributed throughout the polymer. However, for many processes that are used to prepare PLGA delivery systems, there is a large difference in content of the encapsulated compound at the surface of the polymer relative to the bulk. This phenomenon, the presence of acidic impurities in the polymer, and erosion events (e.g., water uptake, acid-catalyzed polyester hydrolysis, sequestration of low-molecular-weight acids, polymer permeability changes, pH-gradients, polymer glass transition changes, etc.) often result in a lowering of microclimate pH in PLGAs.
Controlled-release systems for proteins and peptides using poly(lactide-co-glycolide) (PLGA) have been studied for more than one decade. Although this type of biodegradable polymer has been successful in delivery of small peptides such as LHRH analogues, the delivery of large globular proteins in PLGA has been limited because of the irreversible inactivation of these therapeutic agents prior to their release in vivo. Previous work from our group has shown that encapsulated bovine serum albumin (BSA) in PLGA systems forms insoluble non-covalent aggregates and is hydrolyzed after incubation in a physiological buffer at 37xc2x0 C. for 28 days. The acidic pH and intermediate water content existing in the polymer were implicated as two major factors causing instability of the encapsulated protein, and the BSA was stabilized by co-encapsulating poorly water-soluble basic inorganic salts such as Mg(OH)2 The incorporation of the basic additive in the formulation was also successful in stabilizing therapeutic proteins such as recombinant human basic fibroblast growth factor and bone morphogenetic protein-2.
In this study, to further characterize the stabilization mechanism by co-encapsulation of Mg(OH)2, the effect of basic additive type and content on protein stability and release kinetics in PLGA delivery devices was studied. Since acid-induced inactivation pathways (e.g., at pH less than 3) are common for most proteins, BSA was selected as a model protein. BSA undergoes unfolding from its F to E form at pH 2.7, and forms non-covalent aggregates in PLGA presumably due to this unfolding. The influence of Mg(OH)2 on the delivery system such as pH change in the release medium, polymer degradation and water uptake kinetics was also examined. In addition, the basicity of the salt as well as the loading of base and protein were examined for their effects on BSA aggregation.
Our results confirm that below a critical loading of either basic salt or protein, both acidic and neutral pH regions in the polymer are present. Successful neutralization by the salt requires selection of the appropriate base as well as the appropriate combination of base and protein loading, which allows the base to diffuse to all the protein-containing pores and neutralize all the acidic regions in the polymer.
Materials and Methods
Chemicals
Poly(DL-lactide-co-glycolide) 50/50 with inherent viscosity of 0.23, 0.41, and 0.63 dl/g in hexafluoroisopropanol were purchased from Birmingham Polymers, Inc. (Birmingham, Ala.). Bovine serum albumin (A-3059, Lot 32H0463) was purchased from Sigma Chemical Co. (St. Louis, Mo.). Poly(vinyl alcohol) (80% hydrolyzed with Mw range of 8,000-9,000), Mg(OH)2, Ca(OH)2, and Ca3(PO4)2 were obtained from Aldrich Chemical Co. (Milwaukee, Wis.). ZnCO3 was from ICN Biopharmaceuticals Inc. (Aurora, Ohio). All these salts were fine powders ( less than 5 xcexcm) and were used as received.
Preparation of PLGA Cylindrical Implants
A solvent extrusion method similar to that used previously by our group for intraocular implants was used to prepare the PLGA cylinders with a diameter on the millimeter scale, which we term millicylinders. Briefly, a uniform suspension of sieved protein powder ( less than 90 xcexcm) with or without basic salt in 50% (w/w) acetone-PLGA 50/50 solution was loaded in a syringe and extruded into a silicone tubing (I.D. 0.8 mm) at about 0.1 ml/min. The solvent extruded suspension was dried at room temperature for 24 h and then dried in a vacuum oven at 45xc2x0 C. for another 24 h before testing. The protein loading was calculated as the percentage of amount of BSA versus the total weight of mixture (i.e., protein, polymer, and salt).
Evaluation of BSA Release From PLGA Implants
Release of protein was carried out in PBST (which consists of PBS (7.74 mM Na2HPO4, 2.26 mM NaH2O4, 137 mM NaCl, and 3 mM KCl, pH 7.4), and 0.02% w/v Tween(copyright) 80) at 37xc2x0 C. under perfect sink conditions. Millicylinders (10xc3x970.8 mm, 5-10 mg or microspheres (about 20 mg) were placed in I ml of the release medium and the medium was replaced at each time point. The protein content was determined by using Coornassie plus protein assay reagent, which is also compatible with denaturing agents (e.g., 6 M urea) and reducing agents (e.g., 10 mM DTT).
Evaluation of BSA Stability Within PLGA Implants
Protein stability was assessed by the percentage of water insoluble non-covalent BSA aggregates generated within the implants versus the initial encapsulated protein. Protein stability within PLGA implants was analyzed as follows: First, millicylinders with a length of 1 cm were incubated under 80% and 96% relative humidity (RH) at 37xc2x0 C. for 21 days. Then, the polymer was dissolved in acetone and centrifuged to spin down the protein. The remaining protein pellet was washed three times with acetone and then air-dried. The final protein pellet was analyzed as in Analysis of the Protein Extracted from PLGA Implants. The protein remaining in PLGA implants after release in PBST at 37xc2x0 C. for 28 days was also extracted similarly and analyzed as above.
Analysis of the Protein Extracted From PLGA Implants
The BSA pellet extracted from PLGA implants was first reconstituted in PBST and incubated at 37xc2x0 C. overnight to determine the soluble protein fraction remaining in the polymer. Any remaining aggregates were collected by centrifugation again, and brought up in the denaturing solvent (PBST/6 M urea/1 mM EDTA) and incubated at 37xc2x0 C. for 30 mm to dissolve non-covalent bonded BSA aggregates. Then, any final undissolved BSA aggregates were collected again and dissolved in the reducing solvent (the denaturing solvent plus 10 mM DTT) to dissolve any disulfide-bonded aggregates.
Protein Assay
For quantitation of soluble BSA, a modified Bradford assay was used as follows: 10 xcexcl of standard or sample in PBST was added to 250 xcexcl of Coomassle reagent/well on a 96-well plate and then the plate was read at 595 nm using a Dynex MRX microplate reader (Dynex Technology, Inc., Chantilly, Va.). The concentration range of the standard curve was 50 to 1000 xcexcg/ml. For quantitation of non-covalent and covalent BSA aggregates, the solvents used for preparation of standards and samples were 6 M urea and 6 M urea/10 mM DTT, respectively.
Measurement of Water Uptake in PLGA Millicylinders
After incubation either in PBST or under relative humidity at 37xc2x0 C., the millicylinders were blotted with tissue paper and weighed immediately. They were then freeze-dried. The water uptake of millicylinders was calculated by:
Water uptake (%)=(W1xe2x88x92W2)/W2xc3x97100% 
Where W1 and W2 are the weights of the fully hydrated millicylinders and the dried millicylinders, respectively.
Measurement of Molecular Weight of PLGA
Weight-averaged molecular weight (Mw) of the degraded polymers was measured by gel permeation chromatography (GPQ on a Styragel(trademark) HR 5E column (7.8xc3x97300 mm, Waters, Milford, Mass.), which was performed on a HPLC system (Waters, Milford, Mass.) equipped with a refractive index detector (Hewlett Packard). The mobile phase was tetrahydrofuran with a flow rate of 1 ml/min. Mw was calculated based on polystyrene standards (Polysciences Inc., PA) using Millenium Software Version 2.10.
SEM Image Analysis of PLGA Implants
Images of PLGA millicylinders were obtained by using a Philips XL30 field emission gun scanning electron microscope (SEM). Samples were coated with conductive-old palladium prior to the analysis.
pH Measurement of Saturated Basic Salts in Water
Basic salts (i.e., Mg(OH)2, Ca(OH)2, ZnCO3 and Ca3(PO4)2) in excess of their solubility were added to 5 ml of distilled water. The suspension was then incubated at 37xc2x0 C. for 7 days. The pH of the supernatant was determined with a Corning 430 pH meter (Corning Inc., NY).