Patent Application
20230414496
Known intravitreal implants are generally based on hydrophobic biodegradable polymers, for example lactic acid and glycolic acid-based matrices such as poly-lactic acid (PLA), poly-glycolic acid (PGA), their copolymers and derivatives poly(lactic-co-glycolic) acid (PLGA). The degraded products of these polymers are metabolized to produce carbon dioxide and water. One limitation with the existing hydrophobic polymer matrices (PLA, PGA, and PLGA) is that they do not blend well with hydrophilic drugs, for example methotrexate. Another disadvantage of these hydrophobic matrices is that they degrade very slowly even after the drug has been released, resulting in local toxicity.
The known sustained release intravitreal implants which are also FDA approved include Retiserti™ (Bausch & Lomb) and Ozurdex™ (Allergan). Retisert is a silicone-based disc shaped non-biodegradable implant comprising the cortico steroid fluocinolone acetonide approved to treat uveitis and diabetes macular edema over a period of 30 months. Ozurdex is a pellet shaped PLGA based implant that administers Dexamethasone and is approved to treat uveitis and diabetes macular edema over a period of 6 months. In these exemplary devices, the drug administered is hydrophobic in nature, which binds well with a hydrophobic polymer matrix reservoir made of PLGA or silicones. Since the drug is hydrophobic in nature, it exhibits a sustained release due to an inherent property of limited diffusivity in the vitreous medium of the eye.
The inventors are unaware of any devices similarly effective for sustained release of hydrophilic drugs in the intravitreal domain. Hence, the currently accepted routes of administration for desired hydrophilic agents is generally by intravitreal injection, which does not generally afford an opportunity for sustained-release. Treatments requiring long-term exposure to a therapeutic agent can be highly aversive to a patient.
As such, there remains a need for a sustained release biodegradable implant and methods of using the same that maintains the therapeutic dosage of hydrophilic drugs such as methotrexate, over a prolonged treatment time period, thereby improving the effectiveness and safety of treatment methods of various ocular diseases, including ocular diseases in the vitreoretinal domain such as primary intraocular lymphoma, uveitis, proliferative vitreoretinopathy, age-related macular degeneration.
Accordingly, the present invention provides biodegradable intravitreal implants that provide sustained release of hydrophilic therapeutic agents and methods of making and using the same to treat various ocular disorders. Specific embodiments are directed to sustained release biodegradable PLGA/PLA coated chitosan-methotrexate implants, methods of making and using the same to treat various ocular diseases manifested in the vitreoretinal domain. According to a very specific embodiment, ocular diseases such as primary intraocular lymphoma may be effectively treated.
According to an embodiment, a biodegradable intravitreal implant for sustained release of a hydrophilic therapeutic agent is provided. The implant is comprised of a lyophilized core comprising a porous hydrophilic polymer matrix forming a swellable polymeric core; a hydrophilic therapeutic agent distributed throughout said lyophilized core at a desired concentration, a smooth, non-porous, degradable hydrophobic polymer coating uniformly disposed about the core; and a plurality of nanoparticles encapsulating the therapeutic agent; wherein, said desired concentration of said hydrophilic therapeutic agent is in a range of 10-40% by weight; said nanoparticles are non-metallic nanoparticles; and said biodegradable intravitreal implant is ophthalmically compatible in the eye.
One embodiment is directed to a biodegradable intravitreal implant adapted to provide sustained release of an effective amount of a therapeutic agent to an intraocular region of the eye. The implant is comprised of a swellable polymeric core comprising a hydrophilic therapeutic agent distributed throughout a hydrophilic polymer matrix at a concentration; a biodegradable hydrophobic polymer coating disposed about the surface of the swellable core, the coating being permeable to the therapeutic agent and the coating having a thickness, wherein upon implantation into the eye, the implant is effective to achieve sustained release of the therapeutic agent for a release duration.
Another embodiment is directed toward the preparation of hydrophilic polymeric nanoparticles. The process comprises the steps of making an aqueous solution of hydrophilic nanoparticles; dissolving a pre-defined amount of hydrophilic polymer in a pre-defined amount of distilled water and stirring vigorously until a clear solution is formed; formation of an organic solvent mixture of organic solvents; addition of the hydrophilic polymeric solution to the organic solvent; formation of nanoparticles is achieved through vigorously stirring and addition of the hydrophilic polymeric solution.
Another embodiment is directed toward the preparation of hydrophilic polymeric nanoparticles using ionic crosslinking; the process comprises the steps of making an acidic solution of the hydrophilic polymer; dissolving a pre-defined amount of hydrophilic polymer in an acidic solution; the aqueous solution of a negatively charged anionic molecule is produced; addition of the acidic solution to the aqueous solution through vigorous stirring; cross-linking of the hydrophilic polymer to the anionic molecules occurs to form the nanoparticles.
Another embodiment is directed to a process for making a sustained release biodegradable intravitreal implant. The process comprises the steps of: mixing a hydrophilic therapeutic agent with a hydrophilic polymer matrix; injecting the mixture into medical grade chemically inert flexible tubing; lyophilizing said tubing containing said mixture to obtain hydrophilic agent-hydrophilic polymer fibers; extracting said hydrophilic therapeutic agent-hydrophilic polymer fibers from the tubing; cutting the hydrophilic drug-hydrophilic polymer fibers into a desired implant length to form a swellable polymeric core; dip-coating the core into a hydrophobic coating solution, the hydrophobic coating solution having a concentration; drying the hydrophobic coating to yield a biodegradable sustained release intravitreal implant having a biodegradable hydrophobic polymer coating disposed about a swellable hydrophilic polymeric core, the coating having a thickness and being permeable to the therapeutic agent; injecting the hydrophilic therapeutic agent in a hydrophobic polymer shell; double emulsification or reverse phase evaporation of lipids and/or polymers to form a lipid-based system containing the hydrophilic therapeutic agent in the core; dissolution of lipids in an organic solvent; dissolution of the hydrophilic active ingredient in an ionic solvent; hydrating the lipids in the aqueous media with agitation; evaporating the organic solvent to form lipid-based liposomal formulations; and post-formation processing involving purification.
According to another embodiment, a method of treating an ocular condition of an eye of a patient is provided. The method comprises placing a sustained release biodegradable intravitreal implant into an intraocular region, the implant comprising a swellable polymeric core of hydrophilic therapeutic agent distributed throughout a hydrophilic polymeric matrix in a concentration, said core coated with a hydrophobic polymer permeable to the therapeutic agent, said coating having a thickness, wherein the therapeutic agent is delivered to the intravitreal region through a combination of diffusion through the permeable membrane, swelling of the core, and degradation of the coating, for a release duration effective to treat the ocular condition.
Particular details of various embodiments of the invention are set forth to illustrate certain aspects and not to limit the scope of the invention. It will be apparent to one of ordinary skill in the art that modifications and variations are possible without departing from the scope of the embodiments defined in the appended claims. More specifically, although some aspects of embodiments of the present invention may be identified herein as preferred or particularly advantageous, it is contemplated that the embodiments of the present invention are not necessarily limited to these preferred aspects.
Unless defined otherwise, all technical and scientific terms used herein have the same meaning as commonly understood by one of ordinary skill in the art to which the presently-disclosed subject matter belongs.
In certain embodiments, a biodegradable intravitreal implant adapted to provide sustained release of an effective amount of a therapeutic agent to an intraocular region of the eye is provided. “Intravitreal implant” refers to a device or element that is sized, structured, or otherwise configured to be placed in an eye and that can release a therapeutic agent over a sustained period of time, including days, weeks, and even months. Intravitreal implants can be placed in an eye without disrupting vision of the eye, and intravitreal implants are generally biocompatible with physiological conditions of the eye and do not cause adverse side effects.
The disclosed implants are comprised of a swellable polymeric core comprising a hydrophilic therapeutic agent distributed throughout a hydrophilic polymer matrix at a concentration. In some embodiments, the swellable polymeric core comprises hydrophilic therapeutic agent-hydrophilic polymer fibers. The hydrophilic therapeutic agent may be homogenously distributed throughout the core of the implant. As used herein, a “hydrophilic therapeutic agent” refers to a portion of the intravitreal implant comprising one or more hydrophilic substances used to treat a medical condition of the eye. The hydrophilic therapeutic agent may be any hydrophilic pharmacologically active agent, either alone or in combination, for which sustained and controlled release is desirable and may be employed. The hydrophilic therapeutic agents are typically ophthalmically compatible, and are provided in a form that does not cause adverse reactions when the implant is placed into the eye. In some embodiments, the hydrophilic therapeutic agent is selected from the group comprising of methotrexate, carboplatin, cisplatin, cladribine, cyclophosphamide, cytarabine, doxorubicin, floxuridine, fluorouracil, gemcitabine hydrochloride, hydroxyurea, ifosfamide, mechlorethamine hydrochloride, mitomycin, topotecan, and hydrophilic proteins such as aflibercept, bevacizumab, ranibizumab, combinations thereof. In certain embodiments, the hydrophilic therapeutic agent is methotrexate. In some embodiments, the swellable polymeric core comprises 10%, 25%, or 40% by weight hydrophilic agent.
The rate of release and the release duration of the hydrophilic therapeutic agent can be controlled by the loading concentration of the hydrophilic therapeutic agent, the weight and size of the hydrophilic therapeutic agent, and the solubility of the hydrophilic therapeutic agent.
The term “hydrophilic polymer matrix” refers to a hydrophilic polymer or polymers which degrade in vivo, and wherein the erosion of the hydrophilic polymer or polymers over time occurs concurrent with the subsequent release of the hydrophilic therapeutic agent. The term includes hydrophilic polymers, which act to release the hydrophilic therapeutic agent through polymer swelling. A hydrophilic polymer matrix may be a homopolymer, copolymer, or a polymer comprising more than two different polymeric units. In some embodiments, the hydrophilic polymer matrix is selected from the group comprising chitosan, hydroxyethylcellulose, hydroxypropylmethylcellulose, and hydroxypropylcellulose, and mixtures thereof. In certain embodiments, the hydrophilic polymer matrix comprises chitosan.
The rate of release and release duration of the hydrophilic therapeutic agent will be controlled in part by the rate of transport through the hydrophilic polymeric matrix of the implant, and thus will be affected by the rate of swelling of different hydrophilic polymers and combinations thereof upon water absorption so as to make the hydrophilic polymer matrix more permeable to the hydrophilic therapeutic agent. Thus, the rate of release and the release duration of the hydrophilic therapeutic agent from the hydrophilic polymer matrix can be controlled by the use of different hydrophilic polymers and combinations thereof. The selection of a particular hydrophilic polymer matrix composition will vary depending on the desired release kinetics of the hydrophilic therapeutic agent and compatibility with the therapeutic agent, as well as the nature of the disease being treated, the implantation site, and the like.
A biodegradable hydrophobic polymer coating is disposed about the surface of the swellable core, with the coating having a thickness and being permeable to the therapeutic agent. As used herein, a “hydrophobic polymer coating” refers to a hydrophobic polymer or polymers which degrade in vivo and refers to a portion of the intravitreal implant that is effective to provide a sustained release of the hydrophilic therapeutic agents of the implant. The erosion of the hydrophobic polymer or polymers over time occurs concurrent with the subsequent release of the hydrophilic therapeutic agent. Besides imparting hydrophobicity to the surface of the implant, the hydrophobic polymer coating prevents the entry of water into the hydrophilic polymer matrix, thereby reducing the rate of swelling of the hydrophilic polymer matrix and subsequent hydrophilic therapeutic agent release. A hydrophobic polymer coating may be a coating covering a core region of the implant that comprises a hydrophilic therapeutic agent distributed throughout a hydrophilic polymer matrix. A hydrophobic polymer coating may be a homopolymer, copolymer, or a polymer comprising more than two different polymeric units. A hydrophobic polymer coating may be polylactic acid, poly(lactic-co-glycolic) acid, polyanhydride, polycaprolactone or polyorthoesters, and mixtures thereof.
The rate of release and release duration of the hydrophilic therapeutic agent can be effected by the degradation and erosion rate of the hydrophobic polymer coating. Thus, the rate of release and release duration of the hydrophilic therapeutic agent can be controlled by the use of different hydrophobic polymers and mixtures thereof. Additionally, the thickness of the hydrophobic polymer coating can be used to control the rate of release and release duration of the hydrophilic therapeutic agent, and in some embodiments the release duration is inversely proportional to the hydrophobic polymer coating thickness. The thickness of the hydrophobic polymer coating can be controlled by several factors, including the molecular weight of the coating polymer or polymers, crystallinity of the coating polymer or polymers and the concentration of the coating solution used to make the hydrophobic coating. Thus, the selection of a particular hydrophobic polymer coating composition will vary depending on the desired release kinetics of the hydrophilic therapeutic agent and compatibility with the therapeutic agent, as well as the nature of the disease being treated, the implantation site, and the like.
In one specific embodiment, a hydrophobic PLA coating is 100 μm thick. Additionally, different hydrophobic polymers can be selected for appropriate hydrophobic surface properties, time dependent degradation properties (biodegradation) and biocompatibility. In some embodiments the hydrophobic polymer coating is selected form the group comprising polylactic acid, poly(lactic-co-glycolic) acid, polyanhydride, polycaprolactone, and polyorthoesters. In other embodiments, the hydrophobic polymer coating comprises polylactic acid. In some embodiments, the hydrophobic polymer is poly(lactic-co-glycolic) acid comprising of polylactic acid and poly(glycolic) acid (PGA) in a desired ratio of PLA:PGA is 50:50, 65:35 or 75:25.
In some embodiments, a biodegradable intravitreal implant for sustained release of a hydrophilic therapeutic agent is provided. The implant is comprised of a lyophilized core comprising a porous hydrophilic polymer matrix forming a swellable polymeric core; a hydrophilic therapeutic agent distributed throughout said lyophilized core at a desired concentration, a smooth, non-porous, degradable hydrophobic polymer coating uniformly disposed about the core; and a plurality of nanoparticles encapsulating said therapeutic agent; wherein, said desired concentration of said hydrophilic therapeutic agent is in a range of 10-40% by weight; said nanoparticles are non-metallic nanoparticles; and said biodegradable intravitreal implant is ophthalmically compatible in the eye.
In a specific embodiment, the non-metallic nanoparticles is selected from silica nanoparticles, graphene, nanodiamonds, fullerene, carbon nanotube, quantum dots, colloidal apatite nanoparticles or hydroxyapatite particles. The nanoparticles are in the shape of a cylinder or a sphere. The therapeutic agent is encapsulated in the liposomal non-metallic nanoparticles wherein the amount of nanoparticles is equivalent to the amount of chitosan in the implant. Said liposomal nanoparticles are embedded in the cylindrical implant that further aids in sustained release of therapeutic agent through permeation process.
In a specific embodiment, the nanoparticles are gold nanoparticles capable of activation through enhanced thermal therapy including activation through ultrasound, laser activation or activation by light.
In a specific embodiment, the process of preparation of the implant involves dissolution of lipids in an organic solvents that includes ethanol, methanol, transcutol, iso-propanol, chloroform and the dissolution of hydrophilic active ingredient takes place in an ionic solvent that includes tetraethyl ammonium, tetra butyl ammonium, 1-butyl-2,3-dimethylimidazolium (BMMIM or DBMIM), 1-dodecyl-3-methyl-docecyl (MIM).
Upon implantation into the eye, the implant is effective to achieve sustained release of the therapeutic agent for a release duration. As mentioned previously, the rate of release and the release duration of the therapeutic hydrophilic agent are controlled by a variety of factors, including but not limited to, the loading concentration of the hydrophilic therapeutic agent, the size of the hydrophilic therapeutic agent, the solubility of the hydrophilic therapeutic agent, the use of different hydrophilic polymers and combinations thereof, the rate of diffusion of the hydrophilic therapeutic agent through the hydrophilic polymers, the rate of swelling of the hydrophilic polymers, the degradation and erosion rate of the hydrophobic polymer coating, the thickness of the hydrophobic coating, and the size and shape of the implant. In some embodiments, the release duration is inversely proportional to the hydrophobic polymer coating thickness. In certain embodiments, the release duration is about one month, while in other embodiments the release duration is about 8-10 weeks. In certain specific embodiments, the rate of release of the hydrophilic therapeutic agent methotrexate is 0.2-2.0 μg/day.
The therapeutic agent release rate data of certain specific embodiments of a PLA coated chitosan-methotrexate implants of the present invention were fitted to pharmacokinetic models to interpret the therapeutic agent diffusion kinetics. Therapeutic agent release data of all methotrexate loadings (10%, 25%, and 40% by weight of the swellable polymeric core) of the coated implants were fitted to zero order equation, first order equation, Higuchi model and Korsmeyer-Peppas model in order to analyze the mechanism of drug release and diffusion kinetics. The fitting of each model is evaluated based on correlation coefficient (R2) values. The R2 values of each model fitting are reported in Table 1.
The Korsmeyer-Peppas model provides an insight into the type of drug release mechanism taking place from swellable polymeric devices. The ‘n’ of the Korsmeyer Peppas model is estimated from the linear regression fit of the logarithmic release rate data. n>1 suggests super case II transport relaxational release and also indicates zero order kinetics. The generic equation for the Korsmeyer Peppas model is as follows:
F=(Mt/M0)=Kkptn (1)
where M0 is the initial amount of drug, Mt is the amount of drug released in time t, F is the fraction of drug released at time t, Kkp is the Korsmeyer Peppas release constant and n is estimated from linear regression of log F versus log t; n suggests the type of diffusion. Consistent R2 values ˜0.99 and ‘n’ values ˜1.2 were obtained by fitting the first 60% of drug release rate data to the Korsmeyer Peppas model (
The zero order release equation represents a process when the release rate of the drug is independent of the concentration of the drug in the system and the generic equation for the zero order equation is as follows:
M
t
=M
0
+K
0
t (2)
The first order release equation represents a system where the release rate of the drug is dependent on the concentration of the drug in the system and the generic equation for the first order equation is as follows:
log Mt=log M0+K1(t/2.303) (3)
The Higuchi release equation predicts that the drug release is caused primarily by diffusion mechanism and the generic equation for the Higuchi model is as follows:
M
t
=K
H
t
1/2 (4)
Therefore, it can be concluded that the drug release mechanism primarily follows i) Korsmeyer Peppas model, and zero order model for the first ˜8 days where the initial burst takes place and 60% of the drug is released due to swelling of the polymer matrix; and ii) first order and Higuchi model from the 10th day till the end of drug release signifying the drug release mechanism being concentration dependent and is primarily caused by diffusion mechanism, as shown in
In some embodiments of the presently-disclosed subject matter, a process for making a sustained release biodegradable intravitreal implant is provided. In certain embodiments the process comprises mixing a hydrophilic therapeutic agent with a hydrophilic polymer matrix and injecting the mixture into medical grade chemically inert flexible tubing. The tubing containing said mixture is lyophilized to obtain hydrophilic agent-hydrophilic polymer fibers, and the hydrophilic therapeutic agent-hydrophilic polymer fibers are extracted from the tubing. The hydrophilic drug-hydrophilic polymer fibers are then cut into a desired implant length to form a swellable polymeric core. The core is then dip-coated into a hydrophobic coating solution having a certain concentration. The coated core is then dried to yield a biodegradable sustained release intravitreal implant having a degradable hydrophobic polymer coating disposed about a swellable polymeric core, the coating having a thickness and being permeable to the therapeutic agent.
In some embodiments of a process for making a sustained release biodegradable intravitreal implant, the hydrophobic coating solution comprises a polymer selected from the group consisting of polylactic acid, poly(lactic-co-glycolic) acid, polyanhydride, polycaprolactone, and polyorthoester. In certain embodiments, the hydrophobic coating solution concentration is proportional to the thickness of the hydrophobic polymer coating. In other embodiments, the hydrophobic coating solution concentration is 40 mg/ml.
In some embodiments of a process for making a sustained release biodegradable intravitreal implant, the hydrophilic polymer matrix is selected from the group comprising chitosan, hydroxyethylcellulose, hydroxypropylmethylcellulose, and hydroxypropylcellulose. In certain embodiments, the hydrophilic therapeutic agent is selected from the group comprising methotrexate, carboplatin, cisplatin, cladribine, cyclophosphamide, cytarabine, doxorubicin, floxuridine, fluorouracil, gemcitabine hydrochloride, hydroxyurea, ifosfamide, mechlorethamine hydrochloride, mitomycin, topotecan, and hydrophilic proteins such as aflibercept, bevacizumab, ranibizumab, combinations thereof. In other embodiments, the swellable polymeric core comprises 10%, 25%, or 40% by weight hydrophilic therapeutic agent.
For certain specific embodiments of polylactic acid (PLA) coated chitosan-methotrexate implants of the present invention, the PLA coating is about 100 μm thick and the length and diameter of the PLA coated implant are 4.2±0.03 mm and 0.9±0.04 mm, respectively.
In another embodiment of the presently-disclosed subject matter, a method of treating an ocular condition of an eye of a patient is provided. The term “treating” or “treat” as used herein, refers to the level or amount of agent required to treat an ocular condition, or reduce or prevent ocular injury or damage without causing significant adverse side effects to the eye or region of the eye. As used herein, an “ocular condition” is a disease or ailment which affects or involves the eye or one or more regions of the eye. In some embodiments, the ocular condition is selected from the group consisting of intraocular lymphoma, primary central nervous system lymphoma, primary vitreo-retinal lymphoma, proliferative vitreo-retinopathy, uveitis, and retinal detachment, while in certain embodiments the ocular condition is intraocular lymphoma.
In some embodiments of a method of treating an ocular condition of an eye of a patient, a sustained release biodegradable intravitreal implant is placed into an intraocular region of the patient. The implant comprises a swellable polymeric core of hydrophilic therapeutic agent distributed throughout a hydrophilic polymeric matrix in a concentration. In some embodiments, the swellable polymeric core comprises hydrophilic therapeutic agent-hydrophilic polymer fibers. The core is coated with a hydrophobic polymer permeable to the therapeutic agent, with the coating having a thickness. The therapeutic agent is delivered to the intraocular region through a combination of, but not limited to, diffusion through the permeable hydrophobic polymer coating, swelling of the core, and degradation of the hydrophobic polymer coating, for a release duration effective to treat the ocular condition.
In certain embodiments of a method of treating an ocular condition of an eye of a patient, the swellable polymeric core comprises hydrophilic therapeutic agent-hydrophilic polymer fibers. In other embodiments the hydrophobic polymer coating is selected form the group comprising polylactic acid, poly(lactic-co-glycolic) acid, polyanhydride, polycaprolactone, and polyorthoesters. In other embodiments, the hydrophobic polymer coating comprises polylactic acid. In some embodiments, the hydrophilic polymer matrix is selected from the group comprising chitosan, hydroxyethylcellulose, hydroxypropylmethylcellulose, and hydroxypropylcellulose, and mixtures thereof. In certain embodiments, the hydrophilic polymer matrix comprises chitosan.
In additional embodiments of a method of treating an ocular condition of an eye of a patient, the hydrophilic therapeutic agent is selected from the group comprising methotrexate, carboplatin, cisplatin, cladribine, cyclophosphamide, cytarabine, doxorubicin, floxuridine, fluorouracil, gemcitabine hydrochloride, hydroxyurea, ifosfamide, mechlorethamine hydrochloride, mitomycin, topotecan, and hydrophilic proteins such as aflibercept, bevacizumab, ranibizumab, combinations thereof. In certain embodiments, the therapeutic agent is methotrexate. In other embodiments, the swellable polymeric core comprises 10%, 25%, or 40% by weight hydrophilic therapeutic agent.
In some embodiments of a method of treating an ocular condition of an eye of a patient, the release duration is inversely proportional to the hydrophobic polymer coating thickness. In certain embodiments, the release duration is about one month, while in other embodiments the release duration is about 8-10 weeks.
The following examples are given by way of illustration and are in no way intended to limit the scope of the claims of the present invention.
This example illustrates particular embodiments of the process for making sustained release biodegradable intravitreal implants of the present disclosure.
Methotrexate (MP Biomedical) is mixed with low molecular weight chitosan (M.W 50,000-190,000 and DA %>75%) (Sigma Aldrich) in dilute HCl to make different mixtures of 10%, 25% and 40% w/w drug loadings. These mixtures are then injected into Tygon® tubing ( 1/16 in I.D). The tubes containing the mixture are lyophilized at a temperature below −40° C. and pressure below 1200 mTorr for 2 hours (Millrock BT48A, Millrock Technology) to obtain chitosan-methotrexate fibers. The chitosan-methotrexate fibers extracted from the Tygon® tubing are cut into desired implant lengths using a surgical knife under an optical microscope to ensure accurate dimensions of the implant.
DL-PLA (M.W 150,000) (Lactel Biodegradable Polymers) is mixed in Dicholoromethane (Fisher Sci.) to synthesize a 40 mg/ml coating solution. The chitosan-methotrexate implants are then dip coated in the PLA coating solution for a hydrophobic surface coating. The dip coating protocol is carried out on both longitudinal directions of the implant to ensure uniform coating on the surface and on two ends of the implant. Each implant is dipped in the PLA solution for 5 sec and dried at room temperature for 2 min. This process is carried out 3 times in each direction, longitudinally. Subsequently, the implants are dried overnight at room temperature in dark conditions. After initial drying, the implants are vacuum dried overnight at 45° C. to evaporate the dichloromethane from the implant.
This example illustrates the appearance, dimensions, microstructure morphology, and hydrophobicity of the PLA coating of certain embodiments of the sustained release biodegradable intravitreal implants of the present disclosure. Optical microscopy and SEM techniques were utilized to assess the implant's material properties, including appearance, dimensions and microstructure morphology. Hydrophobicity of the PLA coating is evaluated using Time of Flight-Secondary Ion Mass Spectroscopy (ToF-SIMS) and Differential Scanning Calorimetry (DSC) studies.
Optical Microscopy (Keyence Digital Microscope, VHX-600) is used to assess the implant's dimensions and appearance. Scanning Electron Microscopy (SEM) (FEI XL 30-FEG, FEI) is used to assess the microstructure and morphology using an accelerating voltage of 15 KV. The implant samples are sputter coated prior to the SEM analysis in Argon plasma using an Au—Pd target for 1 min to cause them to be conductive.
A summary of the implant dimensions is provided in Table 2. For implant samples (n=9; 3 samples and 3 readings per sample), the dimensions of the uncoated type and the PLA coated type are measured using an optical microscope. The length and cross-sectional diameter of the uncoated implant are 4±0.04 mm and 0.7±0.03 mm, respectively. The length and cross-sectional diameter of the PLA coated implant are 4.2±0.03 mm and 0.9±0.04 mm, respectively.
The optical microscopy images of surfaces of the PLA coated and the uncoated implants are shown in
SEM images showing the longitudinal view of the surface of the uncoated and PLA coated implants are shown in
ToF-SIMS is used to assess the hydrophobic modification of the implant's surface. ToF-SIMS is performed using a ToF-SIMS IV instrument (IONTOF Inc.). Secondary ions are produced from a Ga+ primary ion source at 15 KV accelerating voltage and 1.5 pA current raster over a 200 μm by 200 μm area of the sample. The secondary ions produced are analyzed in high-current bunched mode with analyzer energy of 2 KV. The ion peaks are assigned using SurfaceLab 6 software (IONTOF Inc.). DSC is used to measure thermal properties of the implants at physiological temperature ˜38° C. DSC is performed at the heating rate of 10° C./min. (DSC6200, Seiko Instruments Inc.).
ToF-SIMS spectra of PLA (MW 150,000), PLA coated 40% chitosan-methotrexate implant surface and uncoated 40% chitosan-methotrexate implant surface are reported in
The spectrum of the uncoated implants does not show the same characteristic peaks (56 [C3H4O+], 71 [C3H3O2+], 73 [C3H5O2+], 127 [C6H7O3+], 128 [C6H8O3+], 129 [C6H9O3+], 143 [C6H7O4+] and 145 [C6H9O4+]) as that of pure PLA mass fragments and PLA coated implant. However, in the spectrum of uncoated implants, there is a match with the spectra of pure PLA mass fragments and PLA coated implant at mass fragment 43 [C2H3O+], but with a much higher relative intensity than the spectra of the pure PLA mass fragments and PLA coated implant. The higher relative intensity from the uncoated implants is probably due to the mass fragment 43 [C2H3O+] being generated from the chitosan and methotrexate present on the surface of the uncoated implants. Therefore, the spectra of
If the coating polymer PLA undergoes glass transition in the physiological conditions, then the PLA coating would soften, affecting the structural properties of the implant, thus leading to faster drug release. A DSC plot of one of the PLA coated implants is shown in
This example illustrates the rate of release and the release duration of the hydrophilic therapeutic agent from particular embodiments of the biodegradable intravitreal implants of the present disclosure.
The implants are kept in vials containing 5 ml of phosphate buffered saline (PBS; pH 7.4). Each implant weighs ˜1 mg. The implants with 40% w/w methotrexate contain ˜400 μg of methotrexate, the implants with 25% w/w methotrexate contain ˜250 μg of methotrexate, and the implants with 10% w/w methotrexate contain ˜100 μg of methotrexate. The vials are slowly stirred in a water bath maintained at 38° C. 1 ml of release media sample (PBS) containing methotrexate is taken out at pre-determined time intervals. 1 ml of fresh PBS is added to maintain sink conditions. The concentration of methotrexate present in 1 ml of release media is assayed using a UV-Visible Spectrophotometer (Cary 50-Bio UV-Vis Spectrophotometer, Varian) at the characteristic methotrexate wavelengths (258,302 and 372 nm) (Perron, M. J., and Page, M., 1994, “Measurement of the Enzymatic Specificity of Carboxypeptidase—A by Capillary Zone Electrophoresis,” J. Chromatogr. A., 662(2), pp. 383-388). The calibration of methotrexate absorbance in the UV-Visible Spectrophotometer is done using methotrexate standard concentrations in PBS. A calibration curve is derived from the absorbance readings obtained from the methotrexate standards and the molar absorbtivity of methotrexate is determined.
Release rate profiles of methotrexate from the uncoated implants are shown in
The mean release rate of the uncoated chitosan-methotrexate implants is 88.9±4.8 ag/day, 188.0±7.9 μg/day and 372.6±7.5 μg/day for the 10%, 25% and 40% w/w drug loadings respectively as mentioned in Table 4. The total release duration is defined as the duration from the start of drug release till the time it remains in the therapeutic window. The total release duration for 10%, 25% and 40% w/w chitosan-methotrexate implants is 19, 29, and 32 hours respectively. The 10% w/w, 25% w/w and the 40% w/w implants remain in the therapeutic window between 12th to 19th hour, 22nd to 29th hour and 25th to 32nd hour respectively as shown in
The mean release rate of the PLA coated chitosan-methotrexate implants is 1.8±0.4 ag/day, 3.2±0.1 μg/day and 6.6±0.3 μg/day for the 10%, 25% and 40% w/w drug loadings respectively as mentioned in Table 5. The total release duration for 10%, 25% and 40% w/w PLA coated chitosan-methotrexate implants are 58, 74 and 66 days, respectively.
For the 10% coated chitosan-methotrexate implant, there is an initial burst release on the 4th day (
For the 25% coated chitosan-methotrexate implant, an initial burst release is seen on the 3rd day (
In the case of 40% coated chitosan-methotrexate implant, a significant initial burst release is noticed on the 3rd day (
Thus, the data demonstrates that uncoated chitosan-methotrexate implants are able to administer the drug for approximately 1 day. This rapid release of methotrexate is expected because of the similar hydrophilic nature of both chitosan and methotrexate. However, the presently disclosed data demonstrates that a PLA coating imparts hydrophobicity to the surface of the chitosan-methotrexate implant, and that the PLA coated chitosan-methotrexate implants are able to administer the therapeutic release rate of 0.2-2.0 μg/day of methotrexate for more than 50 days.
The PLA coating plays an important role in sustained release administration of methotrexate and also influences the initial burst release or the peak release rate of methotrexate. Besides imparting hydrophobicity to the surface of the implant, the PLA coating prevents the entry of PBS into the chitosan matrix, thereby reducing the rate of swelling of the chitosan matrix and subsequent methotrexate release. The presently disclosed data further demonstrates that the sustained release of methotrexate from the PLA coated implants can also be attributed to the degradation rate of PLA coating. Thus, the presently disclosed data demonstrates that sustained release biodegradable intravitreal implants that consist of a degradable hydrophobic polymer coating disposed about a swellable polymeric core comprising a hydrophilic therapeutic agent distributed throughout a hydrophilic polymer matrix at a concentration, can be used as an alternative to intravitreal injections for sustained release of the therapeutic agent and potentially better tolerance and improved efficacy in treating ocular diseases, including ocular diseases in the vitreo retinal domain, using minimally invasive surgical methods.
This example illustrates particular embodiments of the process for making sustained-release biodegradable intravitreal implants of the present disclosure.
PLGA of different PLA/PGA ratios and different molecular weight of PLAs were used to coat the Chitosan-Methotrexate implant. The different PLGA used for hydrophobic coating were PLGA 5050 (PLA/PGA ratio=50:50); PLGA 6535 (PLA/PGA ratio=65:35); PLGA 7525 (PLA/PGA ratio=75:25). The different molecular weight of PLA used for coating were 100 kDa (PLA 100) and 200 kDa (PLA 200). The implants were fabricated based on the method described in the earlier study. Briefly, CS-MTX fibers were obtained by freeze-drying the mixture of CS and MTX in 0.1 N HCl in Tygon® tubing (Saint-Gobain, Malvern, PA, USA) of a 1/16-inch internal diameter. Subsequently, these fibers were cut to desired lengths under a microscope, and then a 200 μm coating of PLGA/PLA was applied on the surface using dip-coating techniques (see
The microstructure and the morphology of the micro-implant were evaluated using optical microscopy (Keyence Digital Microscope, VHX-600, Osaka, Japan) and scanning electron microscopy (SEM) (FEI XL 30-FEG, FEI, Hillsboro, OR, USA). The final dimensions of the PLGA-coated CS-MTX micro-implant containing 40% w/w MTX were ˜4.3 mm in length and 1.2 mm in diameter. The weight of the micro-implant containing 40% w/w MTX was 1 mg.
This example illustrates particular embodiments of the statistical analysis of the ERG data.
The ERG study was conducted using a portable ERG machine (HMsERG system, Ocuscience LLC, Henderson, Nevada). The ERG measurement on each rabbit for both prior to surgery (PS) and prior to euthanasia (PE) were conducted during the daytime, between 6:00 a.m. and 2:00 p.m. The same systemic anesthesia (a mixture of xylazine hydrochloride (5 mg/kg) and ketamine hydrochloride (45 mg/kg) by intramuscular injection followed by Isoflurane 1-2.5%) was used at each time point for ERG recording. Following anesthetizing the animals, electrodes between the rabbit and the HMsERG system were connected. Standard procedure reported within Surgery sub-section was followed. Animals were occasionally kept on a Bair Hugger blanket to maintain their body temperature at 37° C. A droplet of 2.5% hypromellose ophthalmic demulcent solution (Goniovisc™, Hub Pharma., Rancho Cucamonga, CA) was applied on the concave side of the ERG-Jet contact lens electrode (Fabrinal SA, La Chaux-de-Fonds, Switzerland), and then the contact lens electrode was placed on the cornea. A stainless-steel needle electrode was inserted subcutaneously at the base of the ear, which was the reference electrode and another similar needle electrode was inserted subcutaneously on top of the forehead (midline), which served as the ground electrode. Animals were kept in a completely dark room for 30 min without any light to ensure dark adaptation of the eyes before the data was acquired for scotopic protocol. On a similar note, the eyes were exposed to 10 min of light adaptation at 30,000 mcd s/m2 by the HMsERG unit for the light adaptation of the photopic protocol
The amplitude and the implicit time of the A-wave and the B-wave were recorded for each intensity, each day, each protocol for a) prior to surgery (PS) and b) prior to euthanasia (PE) conditions. As an indicator of the retinal functional integrity, the ratio of the B-wave amplitude to the A-wave amplitude, known as the B/A ratio (Damico et al., 2012; De Paiva et al., 2019; Lam, 2005), was also recorded for each intensity for both the PS and PE conditions.
For each rabbit, at every time point and for each of the intensities, the relative B/A ratio was computed for both protocols. The relative B/A ratio, (B/A)rel, is defined as:
(B/A)rel=(B/A ratio prior to euthanasia)/(B/A ratio prior to surgery) (5)
The effect of the intensities, the days (observation time-points), the protocols on the (B/A)rel were studied using a mixed-effects model. This model used 288 data points (For day 1, 3 and 7, i.e., 3 days×3 rabbits per day×2 eyes per rabbit×2 protocols per eye×3 intensities per protocol=108 data points; and for day 14, 28 and 56, i.e., 3 days×5 rabbits per day×2 eyes per rabbit×2 protocols per eye×3 intensities per protocol=180 data points). Subsequently, the effect of the protocols, the intensities, the days along with their interaction on the (B/A)rel was examined using a 3-way analysis of variance (ANOVA) model, which is defined as:
(B/A)rel)ijkl=μ+Ii+Dj+Pk+(ID)ij+(IP)ik+(DP)jk+εijkl (6)
where i=3,000, 10,000, and 25,000 mcd s/m2; j=days 1, 3, 7, 14, 28 and 56; k=scotopic and photopic protocols; l=1-5 rabbits (replication number—1, 2, 3 for day 1, 3 and 7; replication number—1, 2, 3, 4, 5 for day 14, 28 and 56); μ=overall effect; Ii=effect of ith level of intensity; Dj=effect of the jth level of day; Pk=effect of the kth level of protocol; (ID)ij=interaction between the ith level of intensity and jth level of days; (IP)ik=interaction between the ith level of intensity and kth level of protocol; (DP)jk=interaction between the jth level of days and kth level of protocol; and eijkl=random error ˜N(0, s). The responses to the intensities>1000 mcd s/m2 were not recorded for both protocols as the light stimulus due to low intensities did not yield significant measurable A-wave responses. The p-value>0.05 was adopted as indicator of the statistically insignificant effect of the intensities, days, protocols, and their interaction on the (B/A)rel. Multiple comparisons of means on protocols, days and intensities were obtained using Tukey contrasts.
Oscillatory potentials (OPs) were recorded by using a band-pass filter between 34 and 300 Hz, like our prior study (Manna et al., 2016a). For each protocol, the OPs were retrieved for intensities 3,000, 10,000, and 25,000 mcd s/m2 on day 1, day 3 and day 56 for eye receiving the MTX micro-implant and the placebo micro-implant. As per the ISCEV (International Society for Clinical Electrophysiology of Vision), OP amplitude is considered the difference between the positive peak following the negative peak, and OP implicit time is the time where the OP amplitude peak is observed. The amplitude and the implicit time of the first 5 OPs after stimulation were recorded for each rabbit.
Consequently, for each time point (days: 1 and 3), n=15 (5 OPs×3 rabbits) for each intensity for each eye at each protocol; and for day 56, n=25 (5 OPs×5 rabbits) for each intensity for each eye at each protocol. For each protocol, each time point, each intensity and in each eye, a 2-tailed Student's t-test was conducted to compare the mean OP amplitude and OP implicit time between PS and PE conditions, where a p>0.05 for mean comparisons of OP amplitude and OP implicit time between PS and PE conditions is considered to be statistically insignificant.
Based on the statistical analysis on the (B/A)rel, it was demonstrated that the PLGA-coated CS-MTX micro-implant maintained the integrity of retinal functional in presence of micro-implants over the entire duration of the experiment.
This example illustrates particular embodiments of material characterization of the micro-implant.
The molecular weight of the lipophilic polymers used for coating.
The molecular weight of the different polymers used for lipophilic surface modification, as obtained from the GPC analysis, is presented in Table 6.
The number averaged molecular weight (Mn) and the weight averaged molecular weight (Mw) of all the polymers are in direct proportion of the reported inherent viscosity of the polymers, as provided by the manufacturer. Furthermore, the polydispersity index (PDI) provides a measure of the distribution of the molecular weight for each polymer used for lipophilic coating. The PDI of all the polymers, except that of PLGA 5050, is in the range of ˜1.5-1.6.
The high PDI of PLGA 5050 (2.78) is indicative of a wide range of molecular weight distribution in the polymer, which is expected to influence the degradation of the polymer coating in the simulated vitreous conditions (Proikakis et al., 2006). DL-PLA with an inherent viscosity of 0.67 dL/g and 1.16 dL/g will be referred to as PLA-100 and PLA-250, respectively, hereafter.
FTIR is used to evaluate the bonding between the CS-MTX matrix and the PLGA/PLA coatings. The characteristic IR bands for CS (
The characteristic IR bands for MTX (
The mean swelling profile of the PLGA/PLA-coated micro-implants is presented in
The PLGA 5050-coated micro-implant and the PLGA 6535-coated micro-implant disintegrate after the 66th and 106th day, respectively. There is no disintegration observed in micro-implants coated with PLGA 7525, PLA 100, and PLA 250. The high PDI of PLGA 5050 (2.78), which is indicative of a wide range of molecular weight distribution in the polymer, is also expected to influence the degradation of the polymer coating in simulated vitreous conditions (Proikakis et al., 2006).
This indicates there is no chemical bonding or complex formations between CS and MTX. The characteristic IR spectra for all the combinations of PLGA and PLA show the IR bands around 2996 cm-1 and 2943 cm-1 (symmetrical and asymmetrical stretchings of alkyl groups, respectively), 1747 cm-1 (carbonyl C═O stretching vibrations) and 1180 cm-1 (C═O stretching vibrations). The IR spectra obtained for PLGA and PLA polymers are consistent with the observations of Marques et al. (Marques et al., 2013). In the IR spectra of the PLGA/PLA-coated micro-implant, the IR bands observed are around 2996 cm-1 and 2943 cm-1 (symmetrical and asymmetrical stretchings of alkyl groups, respectively), 1747 cm-1 (carbonyl C═O stretching vibrations) and 1180 cm-1 (C═O stretching vibrations), which represent the lipophilic PLGA/PLA coating of the CS-MTX micro-implant. Furthermore, in PLGA/PLA-coated micro-implant, an IR band around 1602 cm-1 (amine N—H bending vibrations) is also observed, which represents the CS-MTX matrix of the micro-implant.
Therefore, in the IR spectra of the PLGA/PLA-coated micro-implant, the characteristic bands of both the coating polymers and the uncoated CS-MTX micro-implant remain unchanged, which indicates the lipophilic coating of PLGA/PLA does not have any chemical bonding with the CS-MTX matrix
Both PLA 100 and PLA 250-coated micro-implant shows significantly reduced (˜1.9 times) swelling compared to that of PLGA-coated micro-implants, which can be attributed to the pure PLA polymer coating. Lastly, the peak swelling of PLA 250 (2.1 times) is lower compared to that of PLA 100 (2.2 times), which could have been caused by a higher molecular weight of the PLA 250.
This example illustrates particular embodiments of the Pharmacokinetics Study.
The concentration of MTX in the vitreous samples obtained from the eyes receiving the MTX micro-implant and the placebo micro-implant for each time-point was analyzed using high-performance liquid chromatography (H PLC). The HPLC method was carried out as described in the United States Pharmacopeia assay for Methotrexate (MTX). This method has been previously described in our previous publication on in vivo study of a similar PLA-coated CS-MTX micro-implant. Briefly, the Agilent® 1100 HPLC system (Agilent Technologies, Santa Clara, C, USAA) with a diode array detector was used for the HPLC analysis. A C-18 column measuring 150 mm×4.6 mm with a pore size of 80 Å was used. The column temperature was set at 23° C. Acetonitrile and phosphate/citrate buffer (pH 6.0) mixed in the ratio of 10:90 was used as the mobile phase. A flow rate of 1 mL/min of the mobile phase was used with an injection volume of 10 μL. The characteristic UV wavelength of 302 nm was used for the detection of MTX. For each time-point, (a) n=3 vitreous samples obtained from the eye receiving the MTX micro-implant and (b) n=3 vitreous samples obtained from the eye receiving the placebo micro-implant were analyzed.
The globes were grossed and sectioned to display the micro-implant and surgical wound in pupil-optic nerve (P-ON) sections. The globes were then processed and stained as previously described. The stained histopathology slides were then evaluated to evaluate any potential toxicity or complications.
The concentration profile of MTX in the vitreous of the eye post-implantation on Days: 1, 3, 7, 14 and 56 is shown in
The details of the histopathological findings have been reported in the non-invasive study of toxicity and performance of the same micro-implant on the same rabbits, as used in this study. Briefly, the issues identified in the histopathology analysis were associated with surgical procedures. Otherwise, the micro-implant showed no signs of toxicity and appeared to be safe for application in the VR domain. Histopathology findings include focal vitreoretinal traction without any apparent predominance between MTX and placebo micro-implants (
Release rate profiles of MTX from the PLGA/PLA-coated micro-implants containing 40% w/w of MTX, are shown in
The cumulative release profiles of MTX from the coated micro-implants are shown in
The mean release rate (Mean±SD) of the PLGA-coated micro-implants is 5.4±0.1 μg/day (PLGA 5050), 5.7±0.5 μg/day (PLGA) 6535) and 3.4±0.6 μg/day (PLGA 7525), as reported in Table 2. Furthermore, the peak release rate of MTX observed from the PLGA-coated micro-implants is 48.6±20.1 μg/day (PLGA 5050, 4th day of release), 31.4±3.5 μg/day (PLGA 6535, 8th day of release) and 15.5±12 μg/day (PLGA 7525, 14th day of release). The total release duration is defined as the duration from the start of drug release till the time it remains in the therapeutic window. The total release duration of MTX from the PLGA-coated micro-implants is 82 days (PLGA 5050), 82 days (PLGA 6535), and 138 days (PLGA 7525). MTX release from PLGA-coated micro-implants remains in the therapeutic window from the 22nd day to the 82nd day (PLGA 5050); from the 26th day onward up to the 82nd day (PLGA 6535); and from the 42nd day up to the 138th day (PLGA 7525).
The mean release rate of the PLA-coated micro-implants is 3.3±0.3 μg/day (PLA 100) and 1.8±0.1 μg/day (PLA 250). Furthermore, the peak release rate of MTX observed from the PLA-coated micro-implants is 15.5±5 μg/day (PLA 100, 8th day of release) and 3.4±0.5 μg/day (PLA 250, 86th day of release). The total release duration of MTX from the PLA 100-coated micro-implant is 138 days. The total release duration of MTX from the PLA 250-coated micro-implant could not be obtained as the study was truncated after 5 months. MTX release from PLA 100-coated micro-implants remains in the therapeutic window from the 42nd day to the 138th day. Furthermore, PLA 250-coated micro-implant exhibits therapeutic release of MTX for the entire duration of the study. It is observed, that with an increase in the ratio of PLA content in PLGA and molecular weight of PLA: a) the mean release rate and the peak release rate of MTX reduce, and b) the total duration of MTX release along with the duration of therapeutic release of MTX increase. Thus, the therapeutic MTX release from all the PLGA/PLA-coated micro-implants is exhibited for an extended period of ˜3-5 months, as compared to 58-74 days in the prior study.
The mechanism of MTX release and diffusion kinetics from all the coated micro-implants is determined from fitting the MTX release data to the release kinetics models. The fitting of each model is evaluated based on correlation coefficient (R2) values (Table 9).
On fitting the first 60% of MTX release rate data from all the PLGA/PLA-coated micro-implants to the Korsmeyer Peppas model (
When the MTX release data is fitted to the zero order equation, the R2 values obtained for the entire duration of drug release (
The R2 values are ˜0.9 when the whole range of MTX release data from the coated micro-implants is fit to the first-order equation (
When the drug release data for the whole range of data is fitted to the Higuchi model, the R2 values obtained are 0.79 (PLGA 5050), 0.88 (PLGA 6535), 0.98 (PLGA 7525), 0.89 (PLA 100), and 0.93 (PLA 250) (
During this phase, MTX is released due to swelling of the polymer matrix and diffusion of loosely bound MTX particles on the surface causing an initial burst release. Overall, the mechanism of the drug release appears to be governed by a combination of: a) diffusion process in the initial phase and b) hydrolysis of the coating polymers in the latter phase. A biphasic release system is observed in the release profiles of micro-implants coated with PLGA 5050, PLGA 6535 and PLA 100. In comparison with the release profiles of micro-implants coated with PLGA 7525 and PLA 250, the rate of diffusion of MTX can be reduced by increasing the a) the PLA content in PLGA and b) molecular weight of PLA, as observed from the Zero order, First order and Higuchi model fits.
All documents cited are incorporated herein by reference; the citation of any document is not to be construed as an admission that it is prior art with respect to the present invention.
Having described embodiments of the present invention in detail, and by reference to specific embodiments thereof, it will be apparent that modifications and variations are possible without departing from the scope of the embodiments defined in the appended claims. More specifically, although some aspects of embodiments of the present invention are identified herein as preferred or particularly advantageous, it is contemplated that the embodiments of the present invention are not necessarily limited to these preferred aspects.
This application is a continuation in part application of U.S. patent application Ser. No. 16/882,850, filed on May 26, 2020 that claims benefit to U.S. Provisional Application Ser. No. 61/712,337, filed on Oct. 11, 2012, which is incorporated by reference herein in its entirety.
| Number | Date | Country | |
|---|---|---|---|
| Parent | 16882850 | May 2020 | US |
| Child | 18172360 | US |
