WO2008025111A2 - Biodegradable device for intraocular drug delivery - Google Patents

Biodegradable device for intraocular drug delivery Download PDF

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WO2008025111A2
WO2008025111A2 PCT/BR2007/000222 BR2007000222W WO2008025111A2 WO 2008025111 A2 WO2008025111 A2 WO 2008025111A2 BR 2007000222 W BR2007000222 W BR 2007000222W WO 2008025111 A2 WO2008025111 A2 WO 2008025111A2
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pharmaceutical formulation
described
drug
containing corticosteroids
implants
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PCT/BR2007/000222
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French (fr)
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WO2008025111A3 (en
WO2008025111A8 (en
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Armando Da Silva Cunha Junior
Rubens Camargo Siqueira
Silvia Ligorio Fialho
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Universidade Federal De Minas Gerais
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Priority to BRPI0604577 priority patent/BRPI0604577A/en
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Publication of WO2008025111A3 publication Critical patent/WO2008025111A3/en
Publication of WO2008025111A8 publication Critical patent/WO2008025111A8/en

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    • AHUMAN NECESSITIES
    • A61MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61KPREPARATIONS FOR MEDICAL, DENTAL, OR TOILET PURPOSES
    • A61K31/00Medicinal preparations containing organic active ingredients
    • A61K31/56Compounds containing cyclopenta[a]hydrophenanthrene ring systems; Derivatives, e.g. steroids
    • A61K31/57Compounds containing cyclopenta[a]hydrophenanthrene ring systems; Derivatives, e.g. steroids substituted in position 17 beta by a chain of two carbon atoms, e.g. pregnane, progesterone
    • A61K31/573Compounds containing cyclopenta[a]hydrophenanthrene ring systems; Derivatives, e.g. steroids substituted in position 17 beta by a chain of two carbon atoms, e.g. pregnane, progesterone substituted in position 21, e.g. cortisone, dexamethasone, prednisone or aldosterone
    • AHUMAN NECESSITIES
    • A61MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61KPREPARATIONS FOR MEDICAL, DENTAL, OR TOILET PURPOSES
    • A61K9/00Medicinal preparations characterised by special physical form
    • A61K9/0012Galenical forms characterised by the site of application
    • A61K9/0048Eye, e.g. artificial tears
    • A61K9/0051Ocular inserts, ocular implants

Abstract

This invention is related to the production process of a pharmaceutical formulation that can treat retinal degenerations and diseases of the posterior segment of the eye, such as diabetic retinopathy, age-related macular degeneration, retinitis by cytomegalovirus, endophthalmitis and uveitis. Said formulation comprises a monolithic system of the drug composed by a biodegradable matrix of a polymer of the polyester class. The said formulation is applicable for the intraocular route. This biodegradable device can deliver one or more drugs, such as dexamethasone, prednisolone and triamcinolone, nonsteroidal antiinflammatory drugs, antibiotics, immunosuppressive agents and antiproliferative agents.

Description

Title: "BIODEGRADABLE DEVICE FOR INTRAOCULAR DRUG

DELIVERY"

DESCRIPTION

This invention is related to the production of a pharmaceutical formulation that can treat retinal degenerations and diseases of the posterior segment of the eye, such as diabetic retinopathy, age-related macular degeneration, retinitis by cytomegalovirus, endophthalmitis and uveitis. The diseases of the posterior segment of the eye account for most of the irreversible blindness cases in the whole world. Therefore, such a scenario leads to the development of new strategies of treatment for retinal degenerations and other posterior segment diseases of the eye. Successful treatment of such diseases is essentially linked to delivering effective drug doses for the innermost ocular tissues (Geroski D H, Edelhauser H F. Transscleral drug delivery for posterior segment disease. Advanced drug delivery reviews, v. 52, p. 37-48, 2001). The said formulation comprises a monolithic system of the drug composed by a biodegradable matrix of a polymer of the polyester class. The said formulation is applicable for the intraocular route.

This biodegradable device can deliver one or more drugs, such as dexamethasone, prednisolone and triamcinolone, nonsteroidal antiinflammatory drugs, antibiotics, immunosuppressive agents and antiproliferative agents.

The conventional ophthalmic formulations are relatively simple: water-soluble drugs are formulated solutions and water-insoluble drugs are formulated as suspensions and ointments. However, such formulations present the inconveniences of low corneal bioavailability, systemic exposition due to nasolacrimal drainage and reduced efficiency in the posterior segment of the eye (Ding, S Recent developments in ophthalmic drug delivery. Pharmaceutical science and technology today, v.1 , n. 8, p. 328-335, 1998). Despite the easily accessible outer structures of the eye, the ocular physiology and anatomy possess some characteristics, such as relative corneal epithelial impermeability, lacrimal dynamics, nasolacrimal drainage, efficient hematocular barrier, which are protection mechanisms that difficult drug penetration and cause a low ocular bioavailability (Ding, S. Recent developments in ophthalmic drug delivery. Pharmaceutical science and technology today, v.1 , n. 8, p. 328-335, 1998). Therefore, only 5% of the administered dose is absorbed and reaches the intraocular tissues, whereas most of such dose is systemically absorbed (Jarvinen, K; Jarvinen, T; Urtti A Ocular absorption following topical delivery. Advanced drug delivery reviews, v.16, p. 3-19, 1995). However, the action sites of most ophthalmic drugs used are localized in the inner part of the eye, more specifically in the tissues attached to the anterior, posterior and vitreous chambers of the eye (Davies, N M Biopharmaceutical considerations in topical ocular drug delivery. Clinical and experimental pharmacology and physiology, v. 27, p. 558-562, 2000).

Lacrimal drainage is the major responsible for drug loss in the precorneal region, which results in low ocular availability, but it is also one of the most important ways of drug delivery in the systemic circulation (Ding, S Recent developments in ophthalmic drug delivery. Pharmaceutical science and technology today, v.1 , n. 8, p. 328-335, 1998).

The treatment of vitreous diseases and those of retina has been problematic due to the difficult access to these structures (Peyman, G A and Ganiban, G J Delivery systems for intraocular routes. Advanced drug delivery reviews, v.16, p. 107-123, 1995). Although it is the most employed method, the ocular drug administration by means of eyedrops leads to their higher concentration in the anterior tissues (cornea, conjunctiva, sclera, vitreous humor and ciliary body) and minimal effect on the posterior region (lens, vitreous body and retina) and it is usually maintained with frequent applications of the formulation (Jarvinen, K; Jarvinen, T; Urtti A. Ocular absorption following topical delivery. Advanced drug delivery reviews, v.16, p. 3-19, 1995). The direct administration of the drug in the subconjunctival space of the eye is somewhat advantageous in relation to the topical therapy for the treatment of diseases of the anterior segment of the eye, but it does not lead to an effective drug concentration in the posterior segment.

The systemic route may be used for this purpose, but it leads to a low penetration in the eye due to the existing hematoretinal barrier that difficults the penetration of blood circulation substances to the retina. To reach a drug concentration within the therapeutic range from of this route, high drug concentrations must be administered for a prolonged period, which may cause adverse effects (Peyman, G A and Ganiban, G J Delivery systems for intraocular routes. Advanced drug delivery reviews, v.16, p. 107- 123, 1995).

Intraocular injection is an alternative method for obtaining adequate drug concentration in the vitreous body and retina. Such a route was restricted to endophthalmitis treatment, but it has been used for cases of proliferative vitreoretinopathy, viral retinitis and uveitis (Peyman, G A and Ganiban, G J Delivery systems for intraocular routes. Advanced drug delivery reviews, v.16, p. 107-123, 1995). However, the rapid blood circulation in these areas leads to a reduced half-life of the drugs, which reduces them to subtherapeutic levels. In order to maintain such levels within the therapeutic range, repeated intravitreous injections are then required and that may lead to the patient's discomfort and occasional complications, such as vitreous hemorrhage, infections, cataract, retinal detachment (Kimura, H and Ogura, Y. Biodegradable polymers for ocular drug delivery. Ophthalmologica, v. 215, p. 143-155, 2001 ; 2001 ; Yasukawa, T; Kimura, H; Tabata, Y; Ogura, Y Biodegradable scleral plugs for vitreoretinal drug delivery. Advanced drug delivery reviews, v. 52, p. 25-36, 2001).

Due to the reasons described above, several studies have been carried out viewing to develop systems that are able to maintain drug concentration within the therapeutic range in the posterior segment of the eye for a longer period of time (Colthurst, M J; Williams, R L; Hiscott, P S; Grierson, I. Biomaterials used in the posterior segment of the eye. Biomaterials, v. 21 , p. 649-665, 2000).

Dexamethasone and its derivatives are glucocorticosteroids of prolonged action used in ophthalmology since the 1960's for treating ocular inflammations due to its safeness and anti-inflammatory potency (Leopold I H. Nonsteroidal and steroidal anti-inflammatory agents. In: Sears M L.; Tarkkanen A. Surgical pharmacology of the eye. 1958. New York: Raven Press, p. 83-133; Baeyens V, Kaltsatos V, Boisrame B, Varesio E, Veuthey J L, Fathi M, Balant L P. Gex-Fabrv M. Gurnv R. Optimized release of dexamethasone and gentamicin from a soluble ocular insert for the treatment of external ophthalmic infections. Journal of controlled release, v. 52, p. 215- 220, 1998).

Glucocorticoids are able to greatly reduce inflammatory manifestations (due to their effects on the concentration, distribution and function of peripheral leukocytes) as well as to inhibit the phospholipase A2 activity. They inhibit tissue leukocyte and macrophage functions, by reducing their capacity to respond to antigens and mitogens and limiting the macrophages' capacity to phagocyte and destroy microorganisms as well as to produce interleukin-1 , collagenase, elastase, tumor necrosis factor and plasminogen activator. Besides their immunosuppressive effect, these glucocorticoids have effect on the inflammatory response by reducing the prostaglandin and leukotriene synthesis resulting from the phospholipase A2 activation, increasing the concentration of certain phospholipids that seem to inhibit the prostaglandin and leukotriene synthesis and may reduce the cyclooxygenase expression with the following decrease of enzyme quantity available for prostaglandin formation (Schimmer B P, Parker K L. Hormόnio adrenocorticotrόfico, esterόides adrenocorticals e seus analogos sinteticos, inibidores da sintese e das agόes dos hormόnios adrenocorticorticais. In: Hardman J E, Limbird L E. Goodman & Gilman as bases farmacolόgicas da terapeutica. 1996. Sao Paulo: McGraw Hill. 9. ed., Cap. 59, p. 1082-1102; Goldfien A. Adrenocorticosterόides e antagonistas cortico-supra-renais. In: Katzung B G. Farmacologia basica e clinica. 1998. Rio de Janeiro: Guanabara Koogan. 6. ed., Cap. 38, p. 450-461)

As it happens with other tissues, corticosteroids seem not to cause specific effects on the eye, but they present a wide range of antiinflammatory activity. Positive or negative, the major effects of corticosteroids on ocular tissues include: reduction of cellular immune response, reduction of inflammatory vascular permeability, stabilization of blood-aqueous barrier, inhibition of epithelial proliferation, inhibition of inflammatory corneal neovascularization, reduction of cicatrization time, increasement intraocular pression, and induction of cataract. (Sherif Z, Pleyer U. Corticosteroids in ophthalmology: past-present-future. Ophthalmologica, Basel, v. 216, p. 305-315, 2002).

At present, some patents related to devices for intraocular drug administration can be found.

Patent US 6,001 ,386, entitled Implantable controlled release device to deliver drugs directly to an internal portion of the body, describes a simple and implantable device for controlled drug release with a nucleus containing an effective amount of a low solubility substance and an external polymeric nonbiodegradable membrane permeable to the delivered substance (Ashton P, Pearson P A. Implantable controlled release device to deliver drugs directly to an internal portion of the body. Patent US 6,001 ,386; 1999). This patent also describes a treatment method for mammals by local or systemic physiologic or pharmacological effects caused by the surgically implantation of the device described above.

Patent US 6,331 ,313 entitled, Controlled-release biocompatible ocular drug delivery implant devices and methods, describes biocompatible controlled-release devices that may be implanted in the eye (Wong V G, Hu M W L, Berger D E. Controlled-release biocompatible ocular drug delivery implant devices and methods. Patent US 6,331 ,313; 2001). Such devices have a nucleus composed of a drug and a polymeric membrane impermeable to an external medium and permeable to the drug. In this system, the drug is delivered through the polymeric membrane orifices that cover less than 10% of total surface area of the system. Such devices may be used as vehicle for several drugs with different solubility and molar mass.

Patent US 6,251 ,090, entitled Intravitreal medicine delivery, describes an implant device through which a variety of pharmacologically active substances may be introduced in the vitreous cavity by means of a simple initial surgery for its implantation (Avery R L, Luttrull J K. Intravitreal medicine delivery. Patent US 6,251 ,090; 2001 ). This device and its implantation method reduce surgical incision and prevent eventual and even repeated invasive surgical procedures. Additional drug amounts may be rapidly introduced or medication may be even varied or modified when necessary. Furthermore, this developed device and its implanting method allow for controlling the dose to be released in the vitreous cavity, besides preventing adverse effects on other ocular tissues during drug use or implanting procedure.

Patent US 6,719,750, entitled Devices for intraocular drug delivery, describes devices for the delivery of therapeutic agents in hardly accessible regions of the body, such as the posterior chamber of the eye and the internal region of the ear (Varner S E, DeJuan E, Shelley T, Barnes A C, Humayun M. Devices for intraocular drug delivery. Patent US 6,719,750; 2004). This invention's devices are less invasive and may be recharged and easily fixed on the treatment site. Such systems are able to release one or more substances in a prolonged period of time.

Patent US 6,726,918, entitled Methods for treating inflammation- mediated conditions of the eye, describes treatment methods for ocular inflammation-mediated conditions by means of a biodegradable device implant placed in the vitreous body, comprised of an antiinflammatory steroid and a biodegradable polymer (Wong V G, Hu M W L. Methods for treating inflammation-mediated conditions of the eye. Patent US 6,726,918; 2004). This implant can release a substance in the vitreous body in a sufficient amount to reach a drug concentration equivalent to at least 0.05 mcg/ml of dexamethasone within 48 hours and can maintain a drug concentration equivalent to at least 0.03 mcg/ml of dexamethasone for about three weeks.

These devices are made of different polymers, that can be biodegradable and non-biodegradable. They may also be of two kinds: matrix (or monolithic) and reservoir (Dash A K, Cudworth Il G C. Therapeutic applications of implantable drug delivery systems. Journal of pharmacological and toxicological methods, v. 40, p. 1 -12, 1998; Kimura H, Ogura Y. Biodegradable polymers for ocular drug delivery. Ophthalmologica, v. 215, p. 143-155, 2001). In the matrix system, the substance is dispersed in the polymeric matrix. In the case of biodegradable system, drug release is achieved by diffusion through the matrix pores, by polymer degradation or by a combination of the two mecanisms. When non-biodegradable polymers are used, drug is released by the matrix through a slow diffusion process. In the reservoir system, made of biodegradable or nonbiodegradable polymers, the substance is found in a central cavity enveloped in a polymeric membrane, which controls the drug release rate. (Dash A K, Cudworth Il G C. Therapeutic applications of implantable drug delivery systems. Journal of pharmacological and toxicological methods, v. 40, p. 1 -12, 1998; Kimura H, Ogura Y. Biodegradable polymers for ocular drug delivery. Ophthalmologica, v. 215, p. 143-155, 2001). As an implant-shaped drug delivery system must be biocompatible with the organism, its components must be noncarcinogenic, hypoallergenic, mechanically stable and non-inductive of inflammation-mediated response on the application site. Additionally, chemical and physical characteristics should not be modified by the local tissue (Athanasiou K A, Niederauer G G, Agrawal C M. Sterilization, toxicity, biocompatibility and clinical applications of polylactic acid/polyglycolic acid copolymers. Biomaterials, v. 17, p. 93-102, 1996).

Implants made of polymeric systems may be applied in different ocular regions. From the most superficial to the deepest one, the regions of the eye can be ordered as follows: the subconjunctival region, subtenonian region, the sclera and the interior of the ocular bulb (anterior chamber and vitreous body). Generally, the deepest the ocular region, the more delicate the procedure and the less effective the drug concentration in the vitreous body and retina (Kimura H, Ogura Y. Biodegradable polymers for ocular drug delivery. Ophthalmologica, v. 215, p. 143-155, 2001 ; Fialho S L, Rego M G B, Cardillo J A, Siqueira R C, Jorge R, Silva-Cunha A. lmplantes biodegradaveis destinados a administraςao intra-ocular. Arquivos brasileiros de oftalmologia, v. 66, p.891 -896, 2003). Low-release nonbiodegradable implants have been approved for use in the United States of America (USA): Ocusert ® (Alza, USA), a conjunctival device delivering pylocarpin; Vitrasert ® (Bausch & Lomb, USA), an intravitreous implant containing gancyclovir that has been used in patients having acquired immunodefficiency syndrome for treating retinitis caused by cytomegalovirus and; Retisert ® (Bausch & Lomb, USA), an intravitreous implant containing fluocinolone indicated for treating chronicle non-infectious uveitis (U.S. Food and Drug Administration. In: FDA news. Product approvals. Available in: < http://www.fda.gov/opacom/7approvl.html>. Accessed on: Aug. 8, 2005). The problem arising from the use of such implants is that they must be removed after the drug has been competely delivered as the polymers are not biodegradable or, in case they are not removed, they should remain on the site until they are taken out by means of surgical procedures.

Natural and synthetic biodegradable polymers have been observed as components of drug delivery systems, but only some of them have shown to be really biocompatible. Natural polymers, made of bovine and human albumin, collagen and gelatin show some constraints as they present questionable purity and remarkable antigenic activity in some cases. The synthetic ones, however, such as polyamids, polyaminoacids, polyalkylcyanoacrylates, polyesters, polyorthoesters, polyuretans and polyakrilamids have held growing interest as drug delivery systems. Polyesters are now the most used biodegradable polymers, such as the poly(ε-caprolactone), poly(D,L-lactic) (PLA) and the copolymers derived from lactic and glycolic acids (PLGA), the latter two being widely used (Jain R, Shah N H, Malick A W, Rhodes C T. Controlled drug delivery by biodegradable poly(ester) devices: different preparative approaches. Drug development and industrial pharmacy, v. 24, p. 703-727, 1998).

So far the state-of-the-art does not clearly describes a biodegradable system of an intraocular implant allowing a prolonged drug delivery conveyance, specifically of a dexamethasone salt. An alternative for such a formulation consists in the use of implants in the form of matrix or monolithic systems. This invention consists in the use of a matrix system compatible with an intraocular drug delivery for humans and animals that may delivery different drugs, specially dexamethasone acetate. The polymer poly D,L-lactic-co-glycolic is used as the polymeric matrix for the system composition.

The polymers of the polyester class, such as those derived from lactic and glycolic acids, have been studied since the decades of 1960 and 1970 for the production of suture threads. Results have proved that they show good mechanical properties, low allergenic capacity, low toxicity, excellent biocompatibility and a predictable kinetics of biodegradation and have called the attention of several researchers for their possible applications in pharmaceutical technology. The use of such polymers was approved by the Food and Drug Administration (FDA) in drug delivery systems and several studies show their low toxicity (Jain R, Shah N H, Malick A W, Rhodes C T. Controlled drug delivery by biodegradable poly (ester) devices: different preparative approaches. Drug development and industrial pharmacy, v. 24, p. 703-727, 1998). Polymers and copolymers derived from lactic and glycolic acids are synthetized by a condensing reaction through an opening of cyclic dimers (lactic acid and/or glycolic acid). Polymerization generally occurs within a period of two to six hours under a temperature of about 175° C, by using a catalyzer (Jain R, Shah N H, Malick A W, Rhodes C T. Controlled drug delivery by biodegradable poly (ester) devices: different preparative approaches. Drug development and industrial pharmacy, v. 24, p. 703-727, 1998).

The polymeric biodegration occurs by erosion through the cleavage of polymeric chain by hydrolysis, which releases lactic and glycolic acids. As they are natural metabolites in the organism, such acids are eliminated by the Krebs cycle in the form of carbonic gas and water (Athanasiou K A, Niederauer G G, Agrawal C M. Sterilization, toxicity, biocompatibility and clinical applications of polylactic acid/polyglycolic acid copolymers. Biomaterials, v. 17, p. 93-102, 1996). The enzymes role in biodegradation of PLGA and PLA is not clearly defined yet, although data described in the literature state that this process does not imply any enzymatic activity, being hydrolysis the only mechanism involved (Jain R A. The manufacturing techniques of various drug loaded biodegradable poly (lactide-co-glycolide) (PLGA) devices. Biomaterials, v. 21 , p. 2475-2490, 2000).

The presence of methyl group (CH3) in a polymer derived from lactic acid gives it a higher hydrophobic^ as compared with a polymer derived from glycolic acid (PGA). Therefore, as it is quite sensitive to hydrolysis, PGA is not adequate for being used in drug delivery systems. As to PLGA, the higher the proportion used of lactic acid, the higher the hydrophobicity of the copolymer as it absorbs less water, and hence a lower degradation rate. Furthermore, the molar mass and crystallinity degree may affect mechanic properties, hydrolysis capacity and degradation rate of such polymers (Lewis D H. Controlled release of bioactive agents from lactide/glycolide polymers. In: Chasin M, Langer R. Biodegradable polymers as drug delivery systems. 1990. New York: Marcel Dekker, p. 01 -41 ; Blanco- Prieto M J, Fattal E, Puisieux F, Couvreur P. The multiple emulsion as a common step for the design of polymeric microparticles. In: Grossiord J L, Seiller M. Multiple emulsions: structure, properties and applications. 1998. Paris: Editions de Sante, p. 397-435).

The glass transition temperature (T9) in the different PLA and PLGA can be found above the physiologic temperature (37°C) and, under such condition, they are shown in a crystalline form. Thus, their chains are presented as relatively rigid structure, with a significant mechanic property and allowing their formulation as drug delivery systems. This feature is also a determining factor of polymeric degradation rate as it is related to the crystallinity degree and organization of polymeric chains (Jain R, Shah N H, Malick A W, Rhodes C T. Controlled drug delivery by biodegradable poly (ester) devices: different preparative approaches. Drug development and industrial pharmacy, v. 24, p. 703-727, 1998).

The described polymers may be used in the preparation of implants that are usually presented in the form of sticks, disks or membranes. The methods for obtaining such systems include: molding, extrusion and film preparation. For implant molding, the mixture of powders (polymer and drug) is placed in an implant-shaped mold, and heating and pressure may be used while preparing it. In implant extrusion, the equipment continuously pushes the powder mixture, and it passes through high temperature and pressure regions where it is melted and compacted then taking its implant definitive shape. Preparing films may be achieved through melting and pressing the powder mixture or by adding a solution. The solution adding method is more widely used and, in this method, components are dissolved by using an appropriate solvent, which produces a solution that is then launched onto a smooth and nonadherent surface. The solvent evaporates and the film is formed and then removed from the said surface (Kimura H, Ogura Y. Biodegradable polymers for ocular drug delivery. Ophthalmologica, v. 215, p. 143-155, 2001).

This invention is now further described by the following examples: Example 1 - Preparation of the biodegradable implants containing dexamethasone acetate

Two different methods, compression and hot molding, were used for the development of the biodegradable implants. PLGA was used as the polymeric matrices and dexamethasone acetate as the drug. All systems presented 27.7% w/w of the drug, equivalent to 25% w/w of dexamethasone, and 72.3 % w/w of polymer.

Firstly, the polymer and dexamethasone acetate were dissolved in a mixture of acetonitrile and distilled water (2:1). The resultant solution was filtered through a 0.22mcm sterile filter and then.lyophilized (E-C MODULYO, E-C Apparatus Inc., USA) to obtain a homogeneous cake. This cake was used for the development of the implants by compression and hot molding techniques.

For the compression technique, the implants were prepared using a Carver hydraulic press at a pressure of one metric ton, using a stainless steel system, specially developed for this purpose, composed of a set of 1 mm diameter cylindric punches (unpublished work). For the hot molding technique, the homogeneous cake previously prepared was molded into rods using a Teflon® sheet heated on a hot plate at a temperature of 100 to 120°C. To obtain a maximum of uniformity, the polymeric systems were weighed and measured after preparation.

Immediately after being developed, the implants composed of PLGA, prepared from compression and hot molding techniques were compared. Macroscopically, all of them were smooth and similar in appearance, presenting somewhat brittle characteristics. The implants presented as a monolithic device, where the drug was dispersed within the polymeric matrices. The mean weight of implants was 4.5 ± 0.5 mg, the mean diameter was 1.0 ± 0.1 mm and the mean length was 4.0 ± 0.1 mm (n = 8). The small batch-to-batch coefficient of variation (2.79%) was indicative of the reproducibility of the techniques. The same amount of powders was used for the preparation of the implants by compression and hot molding, in order to be possible the comparison of the two methods. Example 2: Differential scanning calorimetrv (DSC) analysis

The crystallinity of the powders and the implants was evaluated by DSC. The thermograms of PLGA powders as received, lyophilized and mixed with 27.7% w/w of dexamethasone acetate were recorded. The thermograms of both kinds of final implants were also analyzed. Differential scanning calorimetry (DSC) (TA Instruments, model

2910 Modulated DSC, USA) technique was used, initially, to find information of the powders about residual solvent after the lyophilization process, drug and polymer stability in the temperature of 100 to 120°C used in the preparation of the implants by the hot molding technique, and the possibility of interactions between the drug and the employed polymer.

For the powders analysis, 3-5 mg of the samples were sealed in aluminium pans. Calibration of the system was performed using indium standard The following procedure was used for all samples, under nitrogem atmosphere: 1 ) heating from 30 to 100°C at the rate of 10°C/min

2) isotherm at 1000C for 5 minutes

3) cooling from 100 to 2O0C 4) heating from 20 to 2500C at the rate of 10°C/min Afterwards, the final implants were also evaluated by DSC. For this analysis, the developed systems were sealed in aluminium pans and then heated from 20 to 250°C, under a nitrogen atmosphere at the rate of 10°C/min.

A trace amount of organic solvent remaining in the sample shifts Tg to lower temperatures (Hatakeyama T, Quinn F X. Applications of thermal analysis. In: Hatakeyama T, Quinn F X. Thermal analysis - fundamentals and applications to polymer science. 1994. London: John Willey & Sons Ltd, p. 65-105). DSC results of the mixtures of polymer and drug showed no alteration in Tg, suggesting the absence of residual solvent in the samples after the lyophilization process.

When PLGA was mixed with the drug, a decrease in the melting temperature of dexamethasone acetate was observed, but only a very small relative crystallinity reduction occurred. As it was not found any significant Tg alteration of the polymer in this mixture, we can suppose that there is no evidence of any important physico-chemical interaction among the drug and this polymer in the mixture.

Finally, the analysis made on the final implants showed that the fabrication processes (particularly the hot molding technique) did not induce negative effect on the formulation, considering the stability of the drug and of the polymer. Tg of the polymer, melting temperature of dexamethasone and the enthalpy of fusion of the drug showed no significant alteration after the preparation processes. Example 3: Evaluation of the surface morphology of the implants prepared according to Example 1

The surface morphology of compressed and molded implants was analyzed by scanning electron microscopy (SEM) using a JSM-35C scanning microscope (Jeol, Japan) operating at 15kV. Immediately after being prepared, the implants were randomly selected and mounted on aluminium stubs horizontally, using double-sided adhesive tape.

Prior to microscopical examination the samples were sputter- coated with a gold layer under argon atmosphere for 1 minute (accessory DSV 203 of the equipment BAF 300, Balzers). The implants surfaces were viewed under 300 X and 1000 X magnifications and the images were transferred to the computer by means of a Digital Image Transference Interface (DITI). The photomicrographs were adjusted using the softwares Adobe Photoshop 6.0 and Adobe Illustrator 9.01 (Adobe Systems Incorporated, 2000, USA).

The scanning electron photomicrographs showed that differences were observed in the surface morphology of implants prepared with both methods employed. The surface of the devices prepared by compression was extremely irregular with too many pores and channels. The systems prepared by the hot molding technique had a smoother surface, with little evidence of cracks and pores. The hot molded implant surface was more homogeneous in appearance than the surface of the ones prepared by compression.

Example 4: In vitro degradation of the implants developed in Example 1

The in vitro degradation of both types of implants was monitored measuring the percent of mass loss. Pre-weighed PLGA implants prepared by compression and hot molding were placed in individual jars and immersed in 200 ml of 0.1 M phosphate-buffered solution (PBS), pH 7.4. This study was carried out in a water bath BD R02020 (Lauda, Germany) at constant temperature of 37°C. At predetermined intervals, the medium containing the incubated implants prepared with PLGA by the two techniques (n = 4) were centrifuged and the liquid was then removed. The remaining sample was then dried for 48 hours in a vacuum desiccator at ambient room temperature. After that, the final dry weight was recorded and the percentage mass loss was calculated as follows: % mass loss = (initial weight - final dry weight) / initial weight. The obtained profiles are typical of these polymers in that there was one onset time before which no mass loss occurred, followed by a rapid mass loss phase, which was described by pseudo-first-order kinetics. Statis- tical analysis using the unpaired Mest by the software GraphPad Prism®, version 3.00 (GraphPad Software Incorporated, 1994-1999) showed that in vitro degradation of the implants prepared by different techniques were not significantly different for P<0.05 (P = 0.1811 for PLGA implants and P = 0.3125 for PLA implants).

The onset time observed was of 4 weeks for PLGA implants prepared by compression, compared to 6 weeks for PLGA implants prepared by hot molding. The lag time observed of 2 weeks for PLGA implants prepared by compression, compared to 4 weeks for PLGA implants prepared by hot molding prior to mass loss decay is proportional to the water influx into hydrophobic PLGA and PLA matrices. As the compressed systems presented a more porous polymeric matrix, the water influx was facilitated and so, contributed to its faster degradation when compared to the molded ones. Since PLGA and PLA undergo bulk erosion, both implants prepared with these polymers presented a lag time phase independent of the manufacturing technique employed for their development. The initial lag phase, typical for the degradation of polyesters implants, has often been explained with the slow penetration of water into hydrophobic matrices. The second phase can be attributed to the hydrolysis of the polymer chain, which occurs by random scission (Witt C, Mader K, Kissel T. The degradation, swelling and erosion properties of biodegradable implants prepared by extrusion or compression molding of poly(lactide-co-glycolide) and ABA triblock copolymers. Biomaterials, v. 21 , p. 931 -938, 2000).

Example 5: in vitro release study of the implants developed as described in Examplei .

The in vitro release study was performed in 200 ml of 0.1 M phosphate-buffered solution (PBS), pH 7.4, under sink conditions. This study was carried out in a water bath BD R02020 (Lauda, Germany) at constant temperature of 37°C, coupled with a stir plate. Four each of pre-weighed PLGA 50:50 and PLA implants prepared by compression and four each of pre-weighed PLGA 50:50 and PLA implants prepared by hot molding were added to individual small leaky polypropylene vials in order to allow the water entrance and to not cause the implants movement through the medium. The vials were then added to individual sealed amber jars. This procedure was realized in order to better approximate to in vivo conditions, where the implants do not move within the ocular medium. At set time intervals, 2 ml of the incubation medium was sampled and 2.0 ml_ of fresh medium was immediately added to each sample jar. The amount of the drug released into the medium was measured by high- performance liquid chromatography (HPLC) using the method described in Example 6. The in vitro release profiles obtained from the developed implants showed that dexamethasone acetate was released slowly. In the beginning, probably, the drug deposited on the surface and in the water channels in the matrix was released. Next, a slow drug release stage, attributed to diffusion through the initial pores already present in the matrices and the new channels formed during the polymer degradation process, was observed.

After 4 weeks, PLGA prepared devices presented changes in the implants initial structure and reduction in strength. The systems developed with this polymer showed one release burst probably due to the drug diffusion through the increased number of pores and channels formed in the matrix during its degradation, which allows faster diffusion of the drug to the incubation medium. The comparison between the implants prepared by different techniques showed that the compressed systems promoted one faster release of dexamethasone acetate than the molded ones. It can be attributed to the higher water uptake in the compressed devices, due to its greater number of pores and channels that seems to influence the degradation rate of the devices.

Statistical analysis using the unpaired f-test by the software GraphPad Prism®, version 3.00 (GraphPad Software Incorporated, 1994- 1999) showed that in vitro release of DA from the implants prepared with the same polymer and by different techniques were significantly different for P<0.05 (P = 0.0437 for PLGA implants and P = 0.072 for PLA implants). The compressed and molded systems showed a maximum percent release of 93 % and 76 % in 25 weeks, respectively.

Example 6: High performance liquid chromatography method for dexamethasone acetate determination The determination of dexametasone acetate was realized measured by high-performance liquid chromatography (HPLC) using the method described in the United States Pharmacopoeia 24 (The United States Pharmacopoeia 24 ed. - NF 19. Rockville: United States Pharmacopoeia Convention Inc., 2000. CD-ROM - Insight Publishing Productivity) by a Waters apparatus equipped with a 717plus autosampler model (Waters, USA). A pump (model 515, Waters, USA) was used at a constant flow rate of 1.2 ml_/min. A C-18 reversed-phase column (3.9 mm x 150 mm) filled with octadecyl silane chemically bonded to porous silica (5 μm, Nova-Pak, Waters, USA) was used. The mobile phase was a mixture of acetonitrile and ultrafiltrated water (45:55). An ultraviolet detector (model 2487, Waters, USA) was used at a wavelength of 254 nm. Method validation showed that media compounds and the polymer did not interfere with dexamethasone acetate retention time, eliminating the risks of overestimation. Example 7: In vivo release study The implants used in this example were prepared using the hot molding method, in accordance with the technique described in Example 1.

Sixty male New Zealand white rabbits, weighing approximately 2.0 to 2.5 kg, were studied. Throughout the observation period, the animals were maintained in the animal facility of the Faculty of Veterinary Medicine, University of Rio Preto, Sao Jose do Rio Preto, Brazil. They were kept in individual cages in a quiet and climatically controlled environment (250C average temperature, air conditioning and exhaustion) and luminosity varying in accordance with the solar light. The animals had free access to standard rabbit chow and water The animals were divided into two groups. In group 1 , devices containing dexamethasone acetate were implanted surgically into the vitreous of the right eye of 38 rabbits. In group 2, control rabbits (22 animals) re- ceived the intravitreous implant without the drug also in the right eye. The experiments were carried out in accordance with the guidelines set forth by the Association for Research in Vision and Ophthalmology (ARVO) for the use of animals in ophthalmic and vision research. The study was approved by the Institutional Animal Care and Use Committee of the School of Medicine of Ribeirao Preto (University of Sao Paulo, Sao Paulo, Brazil) and by the Ethics Committee in Animal Experimentation of the Federal University of Mi- nas Gerais (BeIo Horizonte, Brazil).

Surgical procedures were necessary for implanting the developed devices. The rabbits were anesthetized with an intramuscular injection of ketamine hydrochloride (30 mg/kg, Ketamin® 50 mg/ml, Cristalia, Brazil) and xylazine hydrochloride (4.0 mg/kg, Coopazine® 2.0 g/100ml, Schering-Plough Coopers, Brazil) with additional doses of ketamine chloride when necessary. The ocular surface was then anesthetized by topical instilla- tion of 0.4% oxybuprocaine hydrochloride (Oxinest; Latinofarma, Sao Paulo, Brazil). A 5-mm peritomy was made in the superotemporal quadrant of the right eye and a 2-mm sclerotomy was created 2 to 3 mm from the limbus. The implant was then inserted into the vitreous cavity through the sclerotomy and positioned without suture. The sclerotomy wound and the peritomy were closed with 7-0 Vicryl sutures (Johnson & Johnson, Sao Jose dos Campos, Brazil).

Four animals (group 1) and two animals (group 2) per week were euthanized with a dose of 100 mg/kg intravenous sodium thiopental (Pento- thai sodium; Abbott Laboratories, Abbott Park, IL, USA) for up to 8 weeks after implantation, and their right eyes were immediately enucleated. The vitreous of the 32 animals of group 1 and of 16 animals of group 2 were completely removed and immediately stored at -80°C until the analysis of dexamethasone acetate concentrations. The implants were also retrieved from the enucleated eyes for the determination of the amount of dexamethasone remaining and for the biodegradation study.

The amount of dexamethasone acetate released into the vitre- ous was measured by a competitive enzyme linked immunosorbent assay (ELISA) as described in Example 8.

The results of the in vivo drug release study were consistent with the three-phase drug delivery profile: firstly, a small peak was observed; a second phase followed, caused by drug diffusion before the start of polymeric erosion and; a final peak came about, caused by disintegration of the polymeric matrix. The dexamethasone acetate levels started to reduce from the seventh week. Animals of group 2, whose eyes received implants with no drug, the dexamethasone acetate was not detected in the vitreous. The in vitro release profile obtained from such implants, previously described in Example 5, was similar to that found in the in vivo study. Initially, a slow drug release stage was observed, attributed to diffusion through the initial pores already present in the matrices and the new channels formed during the polymer degradation process. A release burst noted after 4 weeks both in vitro and in vivo was probably due to the drug diffusion through the increased number of pores and channels. Thereafter, the implant released the drug in vivo faster than it did in vitro, a fact that can be attributed to the environment surrounding the implant in the vitreous, which is not the same as the in vitro environment. Drug movement in the vitreous body and the elimination profile of the drug in the rabbit eye contribute to the faster drug release. The nature of the vitreous and surrounding tissue barriers creates concentration gradients within the vitreous that must be accounted for when developing ophthalmic drug therapy (Friedrich S, Saville B, Cheng Y L. Drug distribution in the vitreous humor of the human eye: the effects of aphakia and changes in retinal permeability and vitreous diffusivity. Journal of Ocular Pharmacology and Therapeutics, v. 13, p. 445-459, 1997; Tojo K, Isowaki A. Pharmacokinetic model for in vivo/in vitro correlation of intravitreal drug delivery. Advanced drug delivery review, v. 52, p. 17-24, 2001).

Effective dexamethasone concentrations for suppressing various inflammatory processes range from 150 to 4,000 ng/ml (Culpepper J A, Lee F. Regulation of IL 3 expression by glucocorticoids in cloned murine T lymphocytes. Journal of Immunology, v. 135, p. 3191 -3197, 1985; Lewis G D, Campbell W B, Johnson A R. Inhibition of prostaglandin synthesis by glucocorticoids in human endothelial cells. Endocrinology, v. 119, p. 62-69, 1986; Grabstein K, Dower S, Gills S, Urdal D, Larsen, A. Expression of interleukin 2, interferon^, and the IL 2 receptor by human peripheral blood lymphocytes. Journal of immunology, v. 136, p. 4503-4508, 1986; Knudsen P J, Dinarello C A, Strom T B. Glucocorticoids inhibit transcriptional and post- transcriptional expression of interleukin 1 in U937 cells. Journal of imunology, v. 139, p. 4129-4134, 1987; Lee S W, Tsou A P, Chan H, Thomas J, Petri K, Eugui E M, Allison A C. Glucocorticoids selectively inhibit the transcription of the interleukin 1 β gene and decrease the stability of interleukin 1 β mRNA. Proceedings of the National Academy of Sciences of the United States of America, v. 85, p. 1204-1208, 1988). Equivalent concentrations were achieved by our implant during the 8-week period of the study.

For chronic posterior segment disorders, mainly uveitis, non- biodegradable implants presented a prolonged release of steroids over several years (Jaffe G J, Ben-Nun J, Guo H, Dunn J P, Ashton P Fluocinolone acetonide sustained drug delivery device to treat severe uveitis. Ophthalmology, v. 107, p. 2024-2033, 2000.). However, as a biodegradable system, our system does not have to be removed after complete drug relea- se as it occurs with non-biodegradable systems.

Example 8: The ELISA method for dexamethasone acetate determination

The amount of dexamethasone acetate released into the vitreous from the biodegradable implants developed was measured by a competi- tive enzyme linked immunosorbent assay (ELISA) for corticosteroid using a commercial kit specific for dexamethasone determination (Corticosteroid EIi- sa Kit, DM 2156, Randox Laboratories Ltd., London, UK). A photometer (MuI- tiskanMS; Titertek.Huntsville, AL1USA) was used to read the plates. The vitreous retrieved weekly from the rabbits of group Il was also analyzed and served as control.

The samples were thawed out at ambient temperature and submitted to analysis without pretreatment according to the kit manufacturer's instructions. The vitreous samples were used in duplicate after homogeniza- tion and diluted to fit the calibration curve. The amount of dexamethasone acetate was expressed as dexamethasone acetate equivalent concentration (ng/ml of vitreous). The limit of detection of the Corticosteroid Elisa Kit for dexamethasone acetate was 0.25 ng/ml in the tested media.

Data obtained for the calibration curves (r2 = 0.98, range = 0.20 to 4.98 ng/ml) were analyzed by linear regression (GraphPad Prism 3.00;

GraphPad Software Incorporated, San Diego, CA, USA). The dilution factor was multiplied in all samples to obtain the real dexamethasone acetate con- centrations.

Example 9: Percentage of dexamethasone acetate remaining in the devices implanted according to Example 7

The amount of dexamethasone acetate remaining in the implants was obtained by estimating the percentage of the drug versus the initial content in the implant.

For this procedure, three of the retrieved implants per week were gently washed with distilled water and then dissolved in a fixed volume of acetonitrile. The amount of dexamethasone acetate was measured by high- performance liquid chromatography using the method described in Example 6

The percentage of dexamethasone acetate still present in the implants by the sixth week was approximately 40%, which means that the devices released in vivo approximately 60% of the drug within six weeks. After 6 weeks, it was not possible to obtain the remaining drug because the implant was so degraded that it could not be removed from the enucleated eyes.

Example 10: In vivo biodegradation study of the devices implanted according to Example 7

Morphological changes of the surface of the dexamethasone a- cetate-loaded implants retrieved from the rabbits' vitreous were analyzed by scanning electron microscopy (SEM) using a Zeiss DSM 950 microscope (Carl Zeiss NTS GmbH, Oberkochen, Germany) operating at 15 kV. Retrieved implants from each week were selected at random. Before visualization, the implants were gently washed with distilled water, blotted with wipes to dry off excess water, and then dried for 72 hr in a vacuum desiccator at room temperature. After drying, they were mounted on aluminum stubs. Prior to microscopic examination, the samples were sputter- coated with a gold layer under an argon atmosphere for 1 min (accessory DSV 203 of the equipment BASF 300; Balzers, Inc., Elgin, IL.USA). The implant surfaces were viewed at 20χ to 1000χ magnification, and the images were transferred to the computer by means of a digital image transfer interfa- ce (DITI). The photomicrographs were adjusted using the software Adobe Photoshop 6.0 and Adobe Illustrator 9.01 (Adobe Systems Inc., San Jose, CA, USA). Implants not placed within the eye were also analyzed for comparison using the same protocol as described above.

The scanning electron photomicrographs showed typical chan- ges in the surface and shape of the developed devices during biodegradation in the eye. After 5 weeks, it was not possible to study the implant surface because the implants were so weak due to the degradation process that they could not resist the electron beam used in the procedure.

The surface morphology of polymeric systems plays an impor- tant role in degradation and drug delivery (Dash A K, Cudworth Il G C. Therapeutic applications of implantable drug delivery systems. Journal of pharmacological and toxicological methods, v. 40, p. 1 -12, 1998). The pores and channels in the matrices allow drug diffusion possibly not dependent u- pon polymer degradation. PLGA matrices degrade by bulk hydrolysis of ester bonds and break down to their constituent monomers, lactic and glycolic a- cids (PARK T G. Degradation of poly-lactic-co-glycolic acid microspheres: effect of copolymer composition. Biomaterials, v. 16, p. 1123-1130, 1995). Some studies have revealed that the degradation of these polymers proceeds faster in the center of the device than on the surface (Kunou N, Ogura Y, Hashizoe M, Honda Y, Hyon S H, lkada Y. Controlled intraocular delivery of ganciclovir with use of biodegradable scleral implant in rabbits. Journal of controlled release, v. 37, p. 143-150, 1995; Kunou N, Ogura Y, Honda Y, Hyon S H, lkada Y. Biodegradable scleral implant for controlled intraocular delivery of betamethasone phosphate. Journal of biomedical materials research, v. 51 , p. 635-641 , 2000). Thus, water channels are formed during the degradation process, connecting the surface to the inner part of the im- plant and allowing drug diffusion throughout the water channels of the polymer matrix (Kunou N, Ogura Y, Yasukawa T, Kimura H, Miyamoto H, Honda Y, lkada Y. Long-term sustained release of ganciclovir from biodegradable scleral implant for the treatment of cytomegalovirus retinitis. Journal of Controlled Release, v. 68, p. 263-271 , 2000). The pores and channels in the matrices may promote an increased water uptake by the implants, which may consequently accelerate the degradation process. The surface of the dexamethasone acetate-loaded implants was initially smooth, with no evidence of pores or channels. The pores started to appear 1 week after implantation and were increased throughout the study. The observed pores can be attributed to voids left behind by the release of the drug or to the absorption of water.

Example 11 : Toxicity studies of the implants developed by hot molding according to Example 1

Firstly, a clinical examination of the rabbits' eyes was performed. Six animals from each group were examined weekly by two masked observers.

Clinical evaluation included ocular inspection and binocular indirect ophthalmoscopy preoperatively and weekly after surgery until week 8. The following signs were evaluated: conjunctival hyperemia and discharge, corneal clarity, hypopyon, cataract, vitreous opacity, and retinal detachment. The intraocular pressure (lOP) of both eyes was measured at baseline and at weeks 1 , 4, and 8 after surgery. Before IOP measurements, 0.5% proxyme- tacaine hydrochloride solution was instilled topically.

Visual inspection and indirect ophthalmoscopy showed no evi- dence of drug toxicity or media opacity in the animals from either group during the 8-week period of the study. Some surgical complications were found in the experimental eyes of two animals in group 1 : one of them had severe inflammation 4 days after implantation, suggesting a diagnosis of endophthalmitis, and the other had retinal detachment after 3 weeks, which was considered to be caused by a retinal tear at the site of sclerotomy. These two animals were excluded from the study. The implant remained at the scleroctomy site throughout the period of the study, and showed signs of degradation from the fifth week.

In the intraocular pressure evaluation, one animal in group 1 had an IOP increase of 10 mmHg in the right eye after 8 weeks (25.33 mmHg) compared with baseline. In the remaining rabbits of thetwo groups in which IOP was measured, there was no increase in pressure, which was lower than 20 mmHg by week 8. There was no statistically significant difference in IOP when baseline values were compared with those obtained at each time point (1 , 4, and 8 weeks) in group 1 (p = 0.0544) or group 2 (p = 0.2500). The statistical analysis of IOP between group 1 and group 2 also showed no signifi- cance preoperatively (p = 0.4705) or after 1 week (p = 0.8852), 4 weeks (p = 0.3123), or 8 weeks of implantation (p = 0.5637).

Retinal function was evaluated by electroretinography (ERG) in both eyes of the same 6 animals of groups 1 and 2 as used in the clinical examination using an LKC model EPIC 2000 unit with a Ganzfeld flash unit (VPA-10; Caldwell Laboratories, Inc., Kennewick, WA, USA). Electroretino- grams were recorded at baseline and 8 weeks after placement of the implants. Scotopic ERGs were performed after at least 30 min of dark adaptation at a frequency of 0.34 Hz, and photopic ERGs were performed at a frequency of 2.8 Hz after at least 10 min of light adaptation. By ERG, no signs of retinal toxicity were observed in the experimental eyes of any of the six animals of group 1 that could be caused by the presence of drug or the polymeric system in the vitreous. At 8 weeks, the A-wave amplitude (p = 1.000) and the B-wave amplitude (p = 0.5637) were not significantly different from baseline. Preoperative and postoperative median values were 83.75 and 76.65 μV for A-wave amplitude and 154.50 and 158.00 μV for B-wave amplitude, respectively. The empirical confidence interval (Cl) for A-wave was 64.00-91.80 before implantation and 74.30- 96.70 by week 8; for B-wave, Cl was 107.00-208.00 preoperative^ and 155.00-220.00 by week 8.

Finally, a histopathoiogical study was carried out. Six rabbits each from groups 1 and 2 used for ERG analysis were sacrificed at 8 weeks for histopathologic analysis. The eyes were immediately enucleated and prepared for light and electron microscopy. The posterior segments of the eyes were immersed for 5 hours in 4% formaldehyde (freshly prepared from paraformaldehyde) in 0.1 M Sorensen's phosphate buffer, pH 7.2, at 40C. Next, small pieces were cut from all regions, re-fixed in 2.5% glutaraldehyde in the same buffer for 3 hours at 40C, and postfixed in 1 % osmium tetroxide for 2 hr at 4°C. They were then dehydrated in graded ethanol, cleared in propylene oxide, and embedded in LX 112 resin (Ladd Research Ind., Burlington, VT, USA).

Semithin sections (0.5 μm) were stained with toluidine blue for examination by light microscopy (Carl Zeiss), whereas ultrathin sections were contrasted with uranyl acetate and lead citrate and examined with a transmission electron microscope.

Histopathologic examination of the experimental eyes showed no signs of retinal toxicity or inflammatory cell infiltration. No structural abnorma- lities were noted at week 8 by light (paraffin and resin sections) or transmission electron microscopy. In short, the morphological features of the retina in the experimental and control eyes did not differ from those of any normal retina processed under the same technical conditions.

In addition, our implant proved to be biocompatible because no substantial toxic reactions were observed by ERG or histopathology.

Example 12: "Mini-device" developed bv the hot molding technique according to Example 1

Biodegradable devices containing dexamethasone acetate were developed by the hot molding technique as described in Example 1. These systems presented an average weight of 1.5 ± 0.2 mg and

8.0 ± 0.3mm of length and 0.40 ± 0.03 mm of diameter and are referred to in this report as "mini-devices". They released 86% of the drug in vitro in 42 days.

Next, the mini-devices were implanted in the rabbits' eyes using a 25-gauge transcleral cannula, with no need of surgical procedure. Dexamethasone acetate was released within the therapeutic range for a period of 21 days. It was not observed retinal histological changes and/or increased intraocular pressure.

Claims

1. A production process of a pharmaceutical formulation containing corticosteroids in solid devices.
2. A pharmaceutical formulation as described in claim 1 containing corticosteroids, in which the corticosteroid is the dexamethasone acetate.
3. A pharmaceutical formulation as described in claim 1 containing corticosteroids, that can be administered by the intraocular route.
4. A pharmaceutical formulation as described in claim 1 containing corticosteroids, in which such formulation is composed of a biodegradable polymer of the polyester class, such as the copolymer poly- lactide-co-glicolide.
5. A pharmaceutical formulation as described in claim 4 containing corticosteroids with a diameter between 0,12 and 1 ,2 mm and length between 1 ,0 de 4,0 mm.
6. A pharmaceutical formulation as described in claim 4 containing corticosteroids that does not present signals of physico-chemical interactions between the polymer and the drug.
7. A pharmaceutical formulation as described in claim 4 containing corticosteroids that presents an in vitro release profile of the drug of, at least, 25 weeks.
8. A pharmaceutical formulation as described in claim 4 containing corticosteroids that presents an in vivo release profile of the drug within the therapeutic range of, at least, 8 weeks.
9. A pharmaceutical formulation as described in claim 4 containing corticosteroids that does not present signals of toxicity or opacity of the ocular media during the 8-week period of the study.
10. A pharmaceutical formulation as described in claim 4 containing corticosteroids that does not present intraocular pressure alterations in rabbits' eyes
11. A pharmaceutical formulation as described in claim 4 containing corticosteroids that does not present signals of retinal toxicity, evaluated by electroretinography.
12. A pharmaceutical formulation as described in claim 4 containing corticosteroids that does not present structural abnormalities or inflammatory cell infiltration, evaluated by histological studies.
13. A pharmaceutical formulation similar to that described in claim 4, containing corticosteroids, presenting a diameter between 0,2 and 0,5mm and length between 5,0 and 9,0mm that can be administered by the intraocular route through a 25-gauge transcleral cannula.
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