Recent Advances of Ocular Drug Delivery Systems: Prominence of Ocular Implants for Chronic Eye Diseases

Chronic ocular diseases can seriously impact the eyes and could potentially result in blindness or serious vision loss. According to the most recent data from the WHO, there are more than 2 billion visually impaired people in the world. Therefore, it is pivotal to develop more sophisticated, long-acting drug delivery systems/devices to treat chronic eye conditions. This review covers several drug delivery nanocarriers that can control chronic eye disorders non-invasively. However, most of the developed nanocarriers are still in preclinical or clinical stages. Long-acting drug delivery systems, such as inserts and implants, constitute the majority of the clinically used methods for the treatment of chronic eye diseases due to their steady state release, persistent therapeutic activity, and ability to bypass most ocular barriers. However, implants are considered invasive drug delivery technologies, especially those that are nonbiodegradable. Furthermore, in vitro characterization approaches, although useful, are limited in mimicking or truly representing the in vivo environment. This review focuses on long-acting drug delivery systems (LADDS), particularly implantable drug delivery systems (IDDS), their formulation, methods of characterization, and clinical application for the treatment of eye diseases.

2.1. Dry Eye Diseases Dry eye diseases, which are characterized by symptoms such as ocular surface irritation and vision impairment, are brought about by insufficient tear production or tear hyperosmolarity [ 22 ]. Accordingly, dry eye diseases could be classified into deficient or evaporative diseases [ 23 ]. The tear film’s osmolality rises as a result, and the ocular surface becomes inflamed [ 24 ]. According to estimates, 5–30% of adults over 50 are at risk for developing dry eye diseases [ 25 , 26 , 27 ]. Increasing evidence suggests that ocular inflammation is a major contributor to the pathophysiology of dry eye because it has demonstrated that regardless of the origin of dry eye condition, proinflammatory cytokines and T helper cells are present on the ocular surface [ 28 ]. There are various pathophysiological factors that might trigger dry eye diseases ( Figure 1 ). The major etiological causes are ocular surface injury, meibomian gland dysfunction, and tear film hyperosmolarity and instability [ 29 , 30 ]. Thus, for dry eye to be defined with the greatest degree of precision and to be distinguished from other ocular surface disorders, the etiological factors are essential. In addition, the symptoms dry eye syndrome are associated with malfunction in particular brain regions [ 31 ]. In addition, gut microbiome disturbance or dysbiosis was identified to be associated with the development of dry eye, particularly primary Sjogren’s disease ( Figure 1 ) [ 32 ]. To provide comfort to the ocular surface, tears are replenished with a variety of lubricants. These lubricants, which are called artificial tears, include several polymer solutions, such as hyaluronic acid, carboxymethyl cellulose, polyvinyl alcohol, polyethylene glycol, and polyvinyl pyrrolidone [ 33 ]. Products made of these polymers can be supplemented with additional additives to increase lubrication and prolong their duration in the eye because they do not include any physiologically bioactive molecules such as those found in real tears [ 34 , 35 ]. Diquafosol sodium and other aqueous secretagogues are useful for treating dry eye conditions and promoting mucin and tear production [ 36 ]. Punctal plugs, which are microscopic implants made of silicone or collagen, were initially developed to treat dry eyes by occluding the punctal duct, causing tear fluid to accumulate [ 37 ]. Topical glucocorticoid formulations have gained widespread acceptance as a temporary therapeutic option for dry eye diseases due to their well-recognized anti-inflammatory effects. Topical steroids have been demonstrated to have anti-inflammatory effects on a number of targets associated with the symptoms and signs of dry eye, such as lowering cytokine expression, maintaining the integrity of the corneal epithelium [ 38 , 39 , 40 ], and increasing tear production in animal models [ 41 ]. Topical glucocorticoid drops have been demonstrated to ameliorate symptoms and clinical indicators after a month of usage in various trials, and provided prominent lowering in the level of pro-inflammatory cytokines [ 40 , 41 , 42 , 43 ]. Nonsteroidal anti-inflammatory drugs were also employed topically to treat dry eye syndrome. Topical diclofenac and topical ketorolac have demonstrated enhanced effectiveness against dry eye syndrome [ 44 , 45 ]. Topical cyclosporine A, a typical immunomodulatory, reduces the number of T cells that are activated and the level of inflammatory markers in dry eye syndrome, as well; it controls inflammation, as well as the death of conjunctival epithelial cells [ 46 , 47 , 48 ]. Tacrolimus, which is a 10–100 times more potent immunosuppressant than cyclosporine A, is routinely used to treat dry eye diseases [ 49 ]. The continuous precorneal clearing caused by the dynamic nature of the ocular surface, together with blinking, nasolacrimal discharge and response, and basal tearing, all help to quickly remove foreign particles from the eye. Less than 20% of the applied dosage remains on the ocular surface after a single blink, providing a brief window for drug absorption (5–7 min) [ 50 ]. This is especially true when the quick turnover of tear fluid is taken into consideration. When two or more eye drops are applied at once, there is greater competition for space in the precorneal cavity, which can further reduce precorneal retention time and ocular bioavailability when treating dry eye diseases [ 51 ]. As a result, the development of better drug delivery systems might increase the efficacy of drugs used topically to treat dry eye disorders. Punctal plugs have been proven to increase the action of loaded medications in the treatment of dry eye diseases and to move beyond the ocular barriers [ 52 , 53 , 54 ]. The incorporation of mucin secretagogue rebamipide into nanocarriers significantly increased the activity and penetration into ocular tissues. The optimum use of cyclosporin A to treat dry eye disorders is hampered by its very hydrophobic nature and very

Uveitis is the inflammation of the uveal tract. The uveal tract, which is the middle part of the eye, is located between the retina on the inside and the sclera, conjunctiva, and anterior chamber on the outside and consists of the ciliary body, the choroid, and the iris [ 79 ]. Uveitis is considered the fourth most common reason for acquired blindness, especially for chronic uveitis, and is characterized by a high rate of related complications [ 80 , 81 , 82 , 83 ]. Uveitis is subcategorized according to the inflamed anatomical section into either the anterior, intermediate, or posterior, where the inflammation and accompanied leucocytes are present in the iris, vitreous humor, or choroid, respectively [ 79 ]. The concurrent presence of anterior, intermediate, and posterior uveitis is called panuveitis. Anterior uveitis, which is far more common than intermediate, posterior, or panuveitis, accounts for around 85% of all incidences of uveitis [ 84 ]. Corticosteroids, such as fluocinolone acetonide, difluprednate, fluormetholone, and triamcinolone acetonide, as well as immunomodulatory medications, including rapamycin, infliximab, and methotrexate, may be applied topically to treat anterior uveitis. The inadequate bioavailability and the ocular tissues’ barrier properties prevent the transfer of administered medications to deeper ocular tissues, which may lead to the failure of the uveitis treatment. Consequently, the use of an efficient drug delivery system could enhance the bioavailability and improve the activity of ocularly applied corticosteroids and immunomodulatory medications [ 85 , 86 ]. As seen in Table 1 , the use of different drug delivery systems led to the formation of more efficient treatment choices.

Cytomegalovirus (CMV) retinitis is still the most common ocular-invading virus in patients with acquired immunodeficiency syndrome (AIDS) [ 96 , 97 ]. Patients continue to be at risk for developing CMV retinitis predominantly as a result of either a delayed diagnosis of HIV infection or as a result of noncompliance, intolerance, or resistance to antiretroviral therapy [ 98 ]. Even though the prevalence of CMV retinitis has significantly decreased due to development of more effective treatments, CMV retinitis is still a major contributor to vision loss in AIDS patients managed with antiretroviral drugs [ 99 ]. Therefore, understanding the incidence rate and risk factors associated with the development of CMV retinitis is essential for both patients and medical professionals. CMV retinitis could be controlled via intravitreal injection of antiviral drugs, such as ganciclovir, foscarnet, and cidofovir [ 100 , 101 ]. These drugs were produced in noninvasive sustained release nanocarrier formulations employing a range of drug delivery systems ( Table 1 ).

Diabetic retinopathy, which is a microvasculature diabetes-related problem, continues to be a predominant cause of vision loss and preventable blindness in individuals aged 20 to 74, especially in middle- and high-income nations [ 120 ]. Globally, an estimated 415 million people had diabetes in 2015, and by 2040, that figure is projected to increase to 642 million [ 121 ]. The number of those who have diabetic retinopathy and visual impairment is growing globally due to the rising incidence of diabetes and the increasing number of diabetics living longer [ 122 ]. The primary pathophysiology of diabetic retinopathy is a combination of changes brought on by hyperglycemia that leads to neovascularization ( Figure 4 ). The neovascularization is caused by increased retinal vascular permeability, increased thickness of the retinal capillary basement membrane, inadequate blood supply to the tissues, and the release of numerous vasoactive molecules. The neovasculature is usually fragile, unstable, and leaky, causing retinal detachment and vitreous bleeding. Diabetic retinopathy combined with neovasculature is usually referred to as proliferative diabetic retinopathy, which can ultimately cause vision loss. Contrarily, the subtype of diabetic retinopathy known as non-proliferative diabetic retinopathy lacks neovascularization in the early stages. The development of microaneurysms and the minor dilatation of retinal blood vessels, which are recognized as early clinical indications of diabetic retinopathy, are common features of non-proliferative diabetic retinopathy [ 123 , 124 ]. The most common cause of visual loss in those with diabetic retinopathy is diabetic macular edema. In diabetic macular edema, the macula swells or thickens as a result of fluid building up sub- and intra-retinally and inside the macula as a result of the collapse of the blood–retinal barrier [ 125 ]. Antiangiogenics, steroids, anti-inflammatories, and antioxidant medications are the most popular treatments for diabetic retinopathy. However, most of these medicaments have poor ocular penetration and require implantation via surgery. Therefore, the development of more advanced drug delivery technologies may improve the potency of currently prescribed drugs, as shown in Table 1 . pharmaceutics-15-01746-t001_Table 1 Table 1 Chronic eye conditions, available therapies, and drug delivery systems and their merits. Disease Treatment Drug Delivery System Platform Advantages of Delivery Systems In Vivo Refs. Dry eye syndrome Tear substitutes Hypromellose Solution [ 126 ] Methylcellulose and derivatives Solution [ 127 ] hyaluronic acid Solution [ 128 ] Aqueous secretagogues Diquafosol sodium Solution [ 129 ] Punctal plugs Collagen and atelocollagen In situ hydrogel Prolonged activity [ 49 , 50 , 130 ] methacrylate-modified silk fibroin In situ hydrogel Prolonged activity [ 54 ] Mucin secretagogues Rebamipide Nanoparticles Sustained release [ 131 ] Liposomes Improved activity [ 132 ] Micelles Improved penetration [ 133 ] Anti-inflammatory and immunomodulatory drugs Cyclosporine Micelles Improved activity [ 134 ] Self-nanoemulsifying Improved efficacy [ 135 ] Liposomes Improved activity [ 136 ] Nanoparticles Improved activity [ 137 ] Nano-emulsion Improved penetration [ 138 ] Solid lipid nanoparticles Controlled release [ 139 ] In situ hydrogel Improved activity [ 134 ] Epigallocatechin gallate Nanoparticles Extended activity [ 140 ] In situ gels Enhanced efficacy [ 141 ] Lactoferrin Nanoparticles Enhanced efficacy [ 142 ] Nanocapsules Controlled release [ 143 ] Liposomes Reduced irritation [ 144 ] Nanostructured lipid carriers Controlled release [ 145 ] Vitamin A Liposomes Improved activity [ 146 ] Tacrolimus Nanoparticles Improved penetration [ 147 ] Progylcosomes Improved activity [ 148 ] Microcrystals Improved efficacy [ 149 ] Liposomes Improved retention time [ 150 ] Micelles Prolonged activity [ 151 ] Nanocapsules Improved activity [ 152 ] Corticosteroids Dexamethasone Dendrimer Improved activity [ 153 ] Nano-wafer Improved activity [ 154 ] Nanostructured lipid carriers Improved activity [ 155 ] Nanoparticles Improved penetration [ 156 ] Micelles Release modulation [ 157 ] Nanosuspension Prolonged activity [ 158 ] Nano emulsion Improved activity [ 159 ] Nanosponges Improved permeability [ 160 ] Fluorometholone Nanoparticles Improved activity [ 161 ] Triamcinolone acetonide Micelles Release modulation [ 60 ] Nanoparticles Improved activity [ 162 ] Hydrocortisone Nanosuspension Prolonged activity [ 158 ] Micelles Improved targeting [ 163 ] Nanoparticles Improved penetration [ 163 ] Nanosuspension Prolonged activity [ 158 ] Prednisolone Nanoparticles Prolonged activity [ 164 ] Nano capsules Reduced toxicity [ 165 ] Lotep rednol etabonate Nanoparticles Improved penetration [ 166 ] Non-steroidal anti-inflammatory drugs Diclofenac sodium Nanoparticles Improved bioavailability [ 167 ] Nanosuspension Prolonged activity [ 168 ] Pranoprofen Nanosuspension I

Amphiphilic polymers can self-assemble into different structures, known as micelles ( Figure 5 ). The formation of micelles within the nanometer range can efficiently improve the aqueous stability and enhance cell permeability. Prior research has demonstrated that the use of a nanomicelle formulation increased medication absorption in the eye [ 320 , 321 ]. Nanomicelle formulations are primarily used to improve the solubility of medications with low solubility and subsequently improve their bioavailability. Hesperidin, Sirolimus, Voriconazole, and Sunitinib are only a few of the medications that were made into polymeric nanomicelles with better solubility and therapeutic efficacy ( Table 1 ).

Lipids have been used to ameliorate the limited water solubility of several lipophilic drugs and adapt them as a drug delivery system [ 329 ]. Müller and Lucks initially developed solid lipid nanoparticles (SLNs, Figure 5 ) in 1996, which received the attention of scientists as a popular, stable, safe, and effective nanoscale drug delivery device. A surfactant layer that surrounds a solid lipid core in SLNs stabilizes and holds the medication [ 330 ]. Drug molecules can be found mostly in the center of particles or molecularly scattered throughout the matrix, depending on the drug solubility and the drug/lipid ratio [ 331 ]. SLNs are considered an efficient system intended for ocular drug delivery. SLNs can improve corneal drug absorptivity, enhance ocular tissue penetration and bioavailability, prolong residence time, and provide extended drug release properties [ 332 ]. SLNs were efficiently used to improve the delivery of bimatoprost [ 185 ], ofloxacin [ 333 ], and dorzolamide [ 334 ], as shown in Table 1 .

Dendrimers ( Figure 5 ) are globular, negatively, positively, or neutrally branch-like nanostructured polymers. They derive their net charge from the functional groups, which are located at the ends of their branches [ 339 ]. These molecules consist of a fundamental unit called the “core”, which comprises the major component, and side chain units called “dendrons” [ 340 ]. Drugs may be conjugated to the ligands on the dendrimer surface or may be retained in the dendrimer core. Dendrimer manufacturing, generation, surface characteristics, and conjugation technique all have an impact on the drug-loading and drug-release kinetics of dendrimers [ 341 ]. As a result of their ability to selectively target inflammatory cells while causing no harm to healthy tissue, dendrimers have proven to be a viable drug delivery vehicle for the treatment of inflammatory eye conditions. The capacity to lower medication toxicity off-target is the key advantage of dendrimers’ targeting abilities [ 153 ]. Utilizing dendrimers effectively can increase the therapeutic effectiveness of various active pharmaceutical compounds ( Table 1 ), including pilocarpine [ 241 ], tropicamide [ 241 ], dexamethasone [ 342 ], brimonidine, and timolol [ 343 ].

Cubosomes ( Figure 5 ) are made up of two inner aqueous pathways that are separated into two arched interpenetrating lipid bilayers, which are structured in three dimensions resembling honeycombs [ 351 ]. These pathways can be occupied by a variety of bioactive molecules, including natural bioactives, chemical pharmaceuticals, peptides, polypeptides, and proteins [ 352 ]. Cubosomes are thought to be promising delivery systems because of their special characteristics, including thermodynamic stability, bioadhesion, the capacity to encapsulate different types of drugs, and their potential to control drug release [ 353 ]. Active medicines and macromolecules can successfully be applied topically to the posterior portion of the eye using cubosomes ( Table 1 ). These drugs include beclomethasone [ 352 ], flurbiprofen [ 354 ], timolol [ 199 ], and brimonidine [ 355 ].

An emulsion ( Figure 5 ) is a uniform dispersion system that is formed upon mixing two or more immiscible liquids under certain circumstances [ 365 ]. Lipid-based emulsions have become a potential vehicle for ocular medication administration. The emulsions enhance ocular delivery using one of two major strategies, either by improving ocular permeability or by lengthening the period the formulation is retained on the ocular surface [ 366 ]. Both hydrophilic and lipophilic drug types may be loaded into emulsions [ 367 , 368 ]. Emulsions have been successfully used to create more effective formulations for several medications used intraocularly that have increased absorption and therapeutic effectiveness. These drugs include cyclosporine A [ 369 ], coumarin-6 [ 370 ], azithromycin, and disulfiram [ 371 ].

Nanocapsules ( Figure 5 ) are a subtype of nanoparticles that are comparable to vesicular systems, in which a medicine is contained in a hollow vessel with an inner liquid core encircled by a polymeric coating [ 378 ]. Nanocapsules are well-known to be retained in the cornea for a prolonged time and to enhance penetration throughout the deep ocular tissues [ 152 ]. Thus, the development of topically applied drug-loaded nanocapsules could reduce uncomfortable intravitreal injections and systemic delivery, which have serious side effects [ 152 ]. The therapeutic action of several medications was effectively potentiated via formulation in the form of nanocapsules ( Table 1 ). These drugs include bevacizumab [ 379 ], prednisolone [ 165 ], tacrolimus [ 152 ], brinzolamide [ 227 ], and cyclosporine [ 380 ].

4.1. Solid Devices Solid ocular devices are applied to the eye in a solid form and include inserts, implants, contact lenses, and films. Ocular inserts are objects that could be loaded with therapeutic drugs and inserted into the conjunctival sac for extending the duration of medicine delivery. Based on their physicochemical characteristics, the inserts are divided into three categories: bioerodible, soluble, and insoluble [ 385 ]. Soluble and erodible devices gradually dissolve while dispensing the medication and require no need for removal, while insoluble inserts can typically distribute medications at a controlled, predetermined rate through reservoir and matrix systems, but they must be removed from the eye [ 386 ]. The system prolongs drug activity, increases drug residency, improves bioavailability, and prevents crest and trough release profiles to subvert the negative effects that go along with those features [ 385 ]. Bimatoprost [ 387 ], acyclovir [ 388 ], triamcinolone acetonide [ 389 ], voriconazole [ 390 ], ketorolac [ 391 ], azithromycin [ 392 ], and dorzolamide [ 223 ] are just a few of the medications that have been delivered non-invasively to the eye using ocular inserts. A list of the commercially available long-acting drug delivery systems is shown in Table 2 . Ocular implants are solid devices that are used as medication delivery systems to slowly release molecules from polymeric matrices that are either biodegradable or not over the course of months to years. Contrary to non-biodegradable implants, which must be surgically removed after treatment, biodegradable solid implants are made utilizing biodegradable polymers, including polycaprolactones, polyglycolic acid, polylactic acid, and polylactic-co-glycolic acid, and polyanhydrides. However, these implants can have unpredictable drug release characteristics [ 431 ]. These implants can be placed at different sites in the eye, including the cameral, vitreal, scleral, episcleral, and subconjunctival areas. Implants have a number of benefits over more conventional means of administering medication to the eye, including bypassing the blood–ocular barrier and delivering a defined drug amount directly to the target site for a long period. Therefore, the danger of infection or retinal detachment may be reduced with the use of implants placed intravitreally, which may also localize therapy to the vitreous with minimal exposure to the systemic circulation [ 412 ]. In addition, implants minimize the need for repeated treatments by continuously supplying medication over a long period and are consequently suitable as a treatment of long-term eye disorders. Durasert TM is a solid polymer implant technology in which small drug molecules can be released for up to three years. Three FDA approved implants using this technology including Iluvien ® , Retisert ® , and Vitrasert ® [ 432 ]. Drug-eluting contact lenses are solid dosage forms that have high potential to produce prolonged drug residence and close drug contact with the cornea, resulting in a major enhancement of drug bioavailability [ 433 ]. Consequently, drug-laden contact lenses provide several advantages, such as lowering the overall amount of medication required, decreasing dosing frequency, and diminishing the quantity of medication lost through systemic absorption [ 434 ]. Molecular imprinting, supercritical fluids, ion ligation, and colloidal polymeric nanoparticles are a few techniques that have been designed to payload pharmaceuticals into contact lenses [ 435 ]. Many medications have been placed into contact lenses in an effort to increase their pharmacological activity and move past the eye’s barriers, which hinder drug delivery, especially for the posterior chamber of the eye. These medications include dexamethasone [ 436 ], phomopsidione [ 437 ], latanoprost [ 438 ], a timolol–bimatoprost combination [ 433 ], and flurbiprofen [ 439 ]. Before some of these technologies can be used in clinical settings and made commercially available, a number of problems need to be resolved, including protein adhesion, diversity in the swelling ability, changes in water content, opacity, surface integrity, strength properties, ion and oxygen permeation, and drug leakage during manufacturing and storage. Ocular films are solid sterile dosage forms that are applied topically on the eye sac to improve ocular bioavailability and remove barriers to ocular drug delivery [ 440 ]. The use of ocular films improves therapeutic efficacy, reduces systemic adverse effects, and minimizes dose frequency. In order to maximize the therapeutic response and patient compliance, ocular films could present intriguing prospects as a vehicle for the administration of therapies. They could thus replace the conventional dosage forms. However, the designation of efficient films for the ocular delivery of therapeutic medications depends on a thorough understanding of the medication, the restrictions of drug permeation to ocular

In recent years, experts have predicted that 3D printing will revolutionize the pharmaceutical industry since it can generate specific doses of individualized medications with novel designs [ 456 ], drug mixtures [ 457 ], and targeted drug release properties [ 458 ]. Additionally, 3D printing could be employed for the development of highly accurate, individualized medical instruments [ 459 ]. Over the past 10 years, 3D printing has been heavily utilized in the fields of contact lens manufacturing, drug delivery to ocular tissues, implants, ocular research, and diagnostic models production [ 460 ]. Ocular prostheses, which aid ophthalmic patients in restoring the symmetry of their face, were successfully developed throughout 3D bioprinting technology with minimal cytotoxic effects. The 3D-printed prosthesis showed no negative effects on the conjunctival sac or membrane and provided the best resemblance to the look of a human eye, including iris color, sclera, and vascular structures [ 461 ]. Ocuserts made by 3D printing were also used to modulate the pharmacokinetics of ganciclovir-loaded glycerosomes, resulting in prolonged release, enhanced tissue penetration, and therapeutic potential [ 272 ]. Three-dimensional (3D) printing was additionally incorporated into the development of prosthetic corneal structures in an effort to bypass religious restrictions and drug histories. Artificial corneas created by 3D printing technology were proven a reliable, quick, convenient, and useful choice [ 462 , 463 ]. Gelatin, collagen, polyvinyl alcohol, and sodium alginate are the primary materials used to create the 3D-engineered corneas because they are biodegradable, translucent, permeable to oxygen and nutrients, able to endure shear stress, and sufficiently robust mechanically [ 464 ]. The development of artificial retinas is essential for the design of more efficient systems for drug delivery, research into disease causes, and the development of cutting-edge therapeutic choices. Artificial retinas with the best cytocompatibility were created via 3D bioprinting, simulating the natural structure of the human retina [ 465 ]. Moreover, human retinal progenitor cells were effectively maintained in vitro by the use of 3D-printed polymeric scaffolds. The subretinal implantation of the cell-free scaffolds into retinitis pigmentosa porcine models did not result in inflammation, infections, or cytotoxicity, supporting the possibility that they may be used in preclinical studies [ 466 ]. Dexamethasone-loaded punctal plugs created by 3D printing demonstrated sustained drug release for 1 to 3 weeks, depending on the particular polymer or blends of polymers chosen [ 37 ]. Ocular 3D-printed patches have successfully been designed to hold various pharmacological active constituents, and they may be adjusted to release varying amounts based on the patient’s demands [ 467 ]. Timolol maleate-loaded 3D-printed contact lenses were successfully utilized to treat glaucoma in patients who did not take their prescribed glaucoma medications [ 468 ]. The lenses had a smooth surface with high printing quality and released timolol maleate steadily over a period of three days [ 468 ]. Three-dimensional micro-stereolithography has been enrolled in the production of therapeutic devices for controlling intraocular pressure and, consequently, glaucoma. It combines the advantages of both digital light processing and stereolithography technologies. Over the past ten years, minimally invasive glaucoma devices have been designed to boost aqueous humor discharge in an effort to control glaucoma [ 469 ]. With the use of 3D printing techniques, a complex surgical device can be produced with significant flexibility while maintaining functionality [ 470 ].

Many chronic ocular illnesses necessitate the use of implanted drug-delivery systems or devices (IDDS) for management or therapy. IDDS are made to be implanted in order to regulate the drug efflux and, as a result, lengthen the time that the disease condition is under control. IDDS have significant benefits over conventional systemic administration. Higher medicament concentrations in the intended locations can be achieved via site-specific implantation, which can avoid oral absorption and distribution phases [ 477 ]. Additionally, IDDS increases patient compliance, minimizes parenteral treatment pain, and sustains the drug concentration in the therapeutic window by a continuous controlled release of the loaded medication [ 478 ]. As a result, IDDS were successfully used in the production of a number of authorized marketed medications to control a variety of chronic diseases, including eye chronic disorders, which include glaucoma, uveitis, endophthalmitis, dry eye diseases, AMD, and diabetic retinopathy. These products include Ozurdex ® (Allergan Co., Ltd.), Retisert ® (Bausch&Lomb), Vitrasert ® (Bausch&Lomb), and I-vation ® (Surmodics Inc.). The technologies or techniques used to generate IDDS and characterize these products are discussed in the following sections. 5.1. Polymers Used to Formulate IDDS The choice of polymer is essential for adjusting the release profile of IDDS. Polymers used for intraocular IDDS might be biodegradable or nonbiodegradable. In the next section, we will discuss the polymers often used to formulate ocular IDDS. 5.1.1. Nonbiodegradable Polymers The virtue of nonbiodegradable polymers, which are used to formulate nonbiodegradable IDDS, is that they may achieve very long-term release and have high biocompatibility [ 479 ]. On the other hand, the matrix polymer needs to be surgically removed after drug exhaustion. These polymers include EVA, polyimide, polyethylene terephthalate, and silicones. Several intraocular IDDS are commercially available, including Retisert ® , Vitrasert ® , Iluvien ® , and Renexus ® .

5.2.1. Solvent Casting For the production of polymeric inserts and implants, solvent casting is an efficient and scalable technique. Various experimental conditions, such as heating and lyophilization, were used to produce and cast polymeric solutions containing drug(s) and plasticizer(s). The type of drug loaded, as well as its thermal stability, play a major role in the choice of the condition. In an effort to improve the stability and therapeutic activity and prolong residency, this approach was used to formulate inserts or implants for several therapeutic medications that had received approval for ocular use. These drugs include dexamethasone [ 480 ], acetazolamide [ 234 ], bimatoprost [ 387 ], etoposide [ 481 ], and dorzolamide [ 482 ]. Table 3 displays a list of the FDA-approved polymers or copolymers used in ocular preparations. Figure 7 a displays a schematic illustration showing the solvent casting process.

Electrospun inserts and implants are automatically generated utilizing a system that includes a syringe pump, collector electrode, and high voltage generator ( Figure 7 c). Several factors may have an impact on the manufactured inserts or implants, including the polymeric solution pump rate, the distance between the syringe tip and collection electrode, and the applied volage [ 485 ]. Electrospinning became popular due to its benefits, including simple control of the shape, diameter, surface properties, and porosity, and the simplicity of achieving nanosized inserts/implants [ 486 ]. Moreover, electrospinning enables the administration of many medicaments at once.

In vitro testing and characterization of drug delivery systems or devices is a key element in pharmaceutical development’s quality control process for evaluating and determining the best formulation(s). One of the most crucial characterization criteria, in vitro testing for dissolution, is utilized to develop in vitro–in vivo relationships, which aid formulation marketing and reduce medication costs. However, the idea of developing in vitro–in vivo correlations becomes more difficult for IDDS, owing to the sophistication of ocular physiological conditions, ocular tissue barriers, and the insertion site of the implant. Therefore, the design of practical in vitro testing for drug release and dissolution from IDDS remains challenging. Several in vitro simulation experiments have been developed to mimic the in vivo insertion of IDDS in the eye tissue. The static diffusion system ( Figure 8 a) was developed to investigate the in vitro efflux rate of drugs formulated as IDDS. In this system, the appropriate release media is chosen and directly incubated with the IDDS and kept at a standard temperature with or without mechanical agitation [ 498 ]. The amount of medication released is then measured at predetermined intervals of time. Despite the static diffusion method’s widespread use, it was restricted in its capacity to investigate IDDS because it lacked ocular flow modelling and the capacity to control diffusion layers [ 499 ]. The agar diffusion system was also adopted to assess the drug release from IDDS under very viscous conditions [ 500 ]. The procedures involved inserting IDDS into agar gel and, using the proper analytical technique, the gel was evaluated for the amount of drug dispersed at predetermined time intervals [ 500 ]. Figure 8 b displays a schematic illustration of the procedures involved. However, the implants are made to be inserted in certain environments; this technique does not accurately reflect such environments. Additionally, this method only relies on a diffusion mechanism to control the drug outflow from the implant, avoiding any potential impact from the actual vitreous environment [ 499 ]. The dialysis bags system is the most straightforward method for predicting the in vitro dissolution and release of therapeutic drugs from IDDS. This technique employed a dialysis bag, which was closed on both sides once the implant was inserted ( Figure 8 c). Drug molecules should be able to pass through the specified molecular weight cutoff for the dialysis bag. After that, the dialysis bag is placed in the release medium solution, which is constantly agitated at standard temperature. At regular intervals, samples from the release media were obtained, analyzed, and quantified [ 501 ]. The Franz diffusion cells or modified Franz diffusion cells with a modified curved donor compartment to accommodate the curvature of the excised corneal tissues operates with the same principle of the dialysis bag, but with a more consistent and reproducible surface area for drug diffusion; they have the capacity to hold ocular tissues. The pharmacokinetic eye model is a more complex system that simulates drug clearance via the anterior chamber, including intraocular aqueous outflow. The apparatus has two compartments that are partitioned by a dialysis membrane that simulates the posterior and anterior ocular chambers. It is hypothesized that this model may also be used to determine how much of the drug would be released from IDDS that are placed in the cavity of the vitreous [ 502 ]. The device was designed to mimic the actual insertion operation of IDDS into the eye. Both the injection and the aqueous inflow ports were positioned inside the replicating vitreous cavity, while a single output port was positioned in the simulating anterior chamber. The device proved highly effective in evaluating the in vitro release studies of various commercially available medicines, including Kenalog ® , Avastin ® , and Lucentis ® . The eye movement system model was created as an in vitro simulation system to imitate the vitreous body, as well as environmental stimuli that move the eyes, such as head movement [ 503 ]. IDDS were inserted inside the chamber, which imitates eye and head movements. The release medium is refreshed every 24 h, and the drug concentration is assessed using the proper analytical technique. The method may demonstrate how the vitreous body’s gelled region, together with conscious motions of the head and eyes, influence the release of produced IDDS. Despite the wide advancements of in vitro testing, there is still no in vitro experimental design that accurately mimics the factors that determine release in an in vivo setting. This happens as a result of the drug’s distribution and penetration process in the eye being more difficult to simulate than with other routes. Furthermore, it would be unethical to repeatedly monitor drug levels in a living eye in order to demonstrate an in vitro–in vivo association relat

The FDA classifies IDDS as either class II or class III medical devices, which, respectively, denote intermediate and notably higher risk levels, because of their direct and persistent contact with the living tissues [ 507 ]. In order for IDDS to be approved by the FDA and marketed in the USA, it needs to obtain the pre-market notification 510(K). A 510(K) is a premarket application submitted to the FDA to prove that the product being marketed is essentially identical to, or equally safe and effective to, a product that has previously gained FDA approval [ 508 ]. If there is no comparable product on the market, the innovative device must receive pre-market approval with sufficient reliable scientific data that must prove that it is effective and safe for the intended usage(s) [ 509 ]. For medical device manufacturers to follow guidelines while designing, producing, packing, and distributing their products, the FDA introduced “Design Control Guidance for Medical Device Manufacturers”. The FDA periodically inspects manufacturers to ensure that they adhere to the required good manufacturing practice requirements [ 510 ]. For IDDS approval, further laboratory tests, including those for sterility, biocompatibility, and material characterization, are required. Sterilization assures patient safety during implantation procedures via the lack of live microorganisms on the device. The FDA recommends terminal sterilization using either ethylene oxide or gamma radiation. A vital component of good terminal sterilization is the packaging mechanism, which must permit gas penetration and radiation to reach the biomaterial. The FDA’s primary criteria are equipment validation, microbiological testing, and sterilization testing [ 511 ]. The chosen material must also be biocompatible and should not result in any undesirable unfavorable biological reactions when in contact with the human body. The material’s biocompatibility must be verified with tests for cytotoxicity, hemocompatibility, pyrogenicity, sensitization, genotoxicity, and carcinogenicity [ 511 ]. The physical, chemical, and mechanical characteristics should be determined for biomaterials allowed to generate IDDS. The pore size, pore size distribution, structure, and connectivity are examples of physical characteristics. The potential for toxicity, carcinogenicity, and immunogenicity are all factors of chemical characterization, and compressibility and mechanical strength are examples of mechanical properties [ 512 ].