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Файл:Ординатура / Хирургия / Библиотека им академика М.И. Перельмана / Книга_5636_Библиотеки_им_академика_М_И_Перельмана
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9 Advances in Biodegradable Ocular Drug Delivery Systems
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Investigational: GVHD
»3 years Episcleral implant
Duration of
drug release Characteristics Eye diseases
Silicone matrix Cyclosporine
(clinicaltrials.gov
identifier
NCT00102583);
corneal allograft
(0.75-inch) and
15 mg/day
(0.5 in.) versions
20–25 mg/day
day)
(15–25 mg/
trials.gov identifier
NCT00447642)
rejection (clinical-
(phase 3)
Investigational: DME
rod-shaped
(3.5 mm
diameter) intra-
length × 0.37 mm
Nonbiodegradable
Months
18–30
0.2–0.5 mg/
day)
acetonide
(0.59 mg;
Fluocinolone
bioadhesive in
low-dose version)
retinitis pigmentosa
Investigational: ARMD,
polymer
encapsulated
vitreal implant
drug-secreting
Nonbiodegradable,
neurotrophic
factor (up to
15 ng/day)
Ciliary
supported by a
PETP scaffold
cells
half-life
1/2
BioSciences
Lumitect™ NEI and NIH/Lux
Brand name Manufacturer Materials Active agent
Alimera Sciences PVA (with silicone
2009; Kane et al.
2008)
(Kiernan and Mieler
Iluvien™/Medidur™
Neurotech Hollow-fiber membrane
NT-501 (Emerich and
2006)
et al. 2004; Tao et al.
Thanos 2008; Thanos
See individual product labels for complete information
ARMD age-related macular degeneration; CME cystoid macular edema; CMV cytomegalovirus; DME diabetic macular edema; EVA ethylene vinyl acetate;
FDA Food and Drug Administration; GVHD graft-versus-host diseases; NEI National Eye Institute; NIH National Institute of Health; PBMA polybutyl meth-
acrylate; PETP polyethylene terephthalate; PVA polyvinyl alcohol; RVO retinal vein occlusion; t
a
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ophthalmic delivery device. It consists of pilocarpine and alginic acid contained
within a reservoir enclosed by two release-controlling membranes made of EVA
copolymer surrounded by a ring to aid in positioning and placement (Conway 2008;
Ghate and Edelhauser 2006). While this product represented a major innovation in
ophthalmic drug delivery technology, its use was limited by complications associated with device insertion and removal and the later development of more effective
topical antiglaucoma medications for lowering intraocular pressure (Conway 2008;
Kearns and Williams 2009).
9.2.3 Vitrasert
Vitrasert® (Bausch and Lomb, Inc., Rochester, NY, USA) is a nonbiodegradable
polymeric intravitreal ganciclovir implant developed for the treatment of cytomegalovirus retinitis. Ganciclovir is encapsulated within a reservoir, encased by
a PVA/EVA membrane, and diffuses out according to zero-order release kinetics
when fluid enters the device and creates a saturated solution (Yasukawa et al.
2006). The device provides sustained release of ganciclovir for 5–8 months and
achieves higher intraocular drug concentrations as compared with systemic
administration. The ganciclovir implant is ideal for cytomegalovirus retinitis
lesions that pose an immediate risk to vision, and combination treatment with
oral valganciclovir can be used to prevent second-eye involvement (Kedhar and
Jabs 2007).
9.2.4 I-vation
I-vation™ (SurModics, Inc., Eden Prairie, MN) is a nonbiodegradable intra-vitreal
implant consisting of a nonferrous metal alloy helix coated with a triamcinolone
acetonide-containing polymer, similar to the design of drug-eluting cardiovascular stents (Kearns and Williams 2009). The device has a sharpened tip, which is
used to make the incision for implantation, and its helical shape maximizes the
surface area for drug coating and enables secure anchoring to the pars plana/sclera
(Conway 2008). The polymers used to manufacture the device are a proprietary
blend of polybutyl methacrylate (PBMA) and PVA, the ratio of which can be customized to vary the drug delivery rate (1–3 mg/day) and corresponding duration of
delivery (6–24 months) (Kiernan and Mieler 2009). I-vation had been under investigation for the treatment of diabetic macular edema, but early trials evaluating the
device reported the incidence of intraocular pressure elevation, conjunctival hemorrhage lenticular opacities, and endophthalmitis, and one phase 2 trial (clinicaltrials.gov study ID NCT00692614) was halted prematurely in 2008 (Kiernan and
Mieler 2009).
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9 Advances in Biodegradable Ocular Drug Delivery Systems
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9.2.5 Iluvien
Iluvien™ (formerly Medidur; Alimera Sciences, Inc., Alpharetta, GA, USA) is a nonbiodegradable intravitreal implant designed to deliver fluocinolone acetonide within a
rod-shaped 3.5 × 0.37-mm reservoir (Kiernan and Mieler 2009). The device is available in two dose formulations: one with a delivery rate of 0.2 mg/day lasting 24–30
months, and the other with a delivery rate of 0.5 mg/day lasting 18–24 months. The
implant’s reservoir is fitted with end caps made from PVA (and a silicone bioadhesive
in the low-dose version) that regulate the rate of drug release and provide nearly zeroorder kinetics, with a slightly higher initial release rate that stabilizes over the long
term (Kane et al. 2008). The implant is inserted using a proprietary 25-gauge injector
system into the inferior vitreous to maximize drug exposure to the retina and minimize exposure to the anterior chamber. It is currently undergoing phase 3 safety and
efficacy trials (Kane et al. 2008); data have not yet been published.
9.2.6 Nonbiodegradable Matrix Implants
9.2.6.1 Lumitect
®
Lumitect® (Lux Biosciences) is an investigational, silicone-matrix episcleral implant
designed for the sustained delivery of cyclosporine for up to 3 years. Two dose formulations have been developed: a 0.75-in. implant that delivers cyclosporine at a
rate of 25 mg/day and a 0.5-in. version with a drug delivery rate of 15 mg/day. The
device is currently undergoing clinical trials for the treatment of ocular graft-versushost disease (clinicaltrials.gov identifier ID NCT00102583) and corneal allograft
rejection (clinicaltrials.gov identifier ID NCT00447642).
9.2.6.2 Punctal Plugs
The Latanoprost Punctal Plug Delivery System (QLT, Inc.) is an experimental sustained-release drug-release implant for the delivery of latanoprost, an analog of
prostaglandin F
that is approved as a topical eye drop (Xalatan®, Merck) to reduce
2
intraocular pressure. The punctal plug is formulated to deliver 44- or 81-mg latanoprost continuously over a 3-month period. Several nonrandomized, open-label,
phase 2 studies evaluating the device in glaucoma and ocular hypertensive patients
have now been completed (clinicaltrials.gov IDs NCT00821002, NCT00845299,
and NCT00820300) but not yet published, and additional phase 2 trials are currently
recruiting subjects (clinicaltrials.gov IDs NCT00967811 and NCT01037036).
Another experimental punctal plug drug system is under clinical development by
Vistakon Pharmaceuticals for the intraocular delivery of bimatoprost, a prostaglandin analog that, like latanoprost, is approved as a topical eye drop for the reduction
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of intraocular pressure. A recently completed randomized, single-blind, paralleldesign, phase 2 study evaluated the efficacy and safety of the bimatoprost punctal
plug in open-angle glaucoma and ocular hypertensive patients (clinicaltrials.gov ID
NCT00824720); the results have not been published.
9.3 Medical Applications for Biodegradable Polymers
Biodegradable polymers have been intensively investigated as surgical biomaterials and in the formulation of drug delivery devices. These materials have included
both synthetic polymers, such as polyester derivatives, and natural polymers (biopolymers), such as bovine serum albumin, human serum albumin, collagen, gelatin, chitosan, and hemoglobin (Jain 2000; Gorle and Gattani 2009). Synthetic
biodegradable polymers have been successfully used as biomaterials for several
decades, being employed initially to produce absorbable sutures, fixation devices
for orthopedic surgery (e.g., bone pins and screws), and stents (see Sect. 9.4).
Biodegradable polymers can be utilized in a variety of ways to increase the residence time of drugs, slow drug clearance, and enhance drug absorption in the eye.
Polymer-based biodegradable ocular drug delivery systems investigated to date
include implantable sustained-release drug pellets and inserts, viscosity enhancers, mucoadhesive agents, drug-releasing contact lenses, and injectable formulations such as hydrogels and liposomes as well as microemulsions, microsuspensions,
microspheres, microcapsules, and their nanoscale counterparts (see Conway
Ghate and Edelhauser 2006; Kearns and Williams 2009; Wadhwa et al. 2009;
Gaudana et al. 2009 for review).
Biodegradable polymer-based drug delivery systems show considerable promise for the treatment of ocular diseases and offer a potential solution to many of the
limitations of conventional (i.e., systemic, oral, and topical) methods for the administration of ophthalmic drugs (Table
retinal diseases (Yasukawa et al. 2006; Ghate and Edelhauser 2008; Wadhwa et al.
2009; Gaudana et al. 2009). Biodegradable polymer-based drug delivery can be
used to achieve prolonged therapeutic drug concentrations in ocular target tissues,
such as the retina, that are not readily accessible by conventional means, and with
drugs that may be poorly absorbed by other routes of administration. Drugs formulated with these polymers can be released in a controlled manner, in which the drug
concentration in the target tissue is maintained within the therapeutic range (Park
et al. 2005), thereby avoiding both insufficient efficacy and high peak drug concentrations associated with pulsed dosing. The tissue specificity of biodegradable drug
delivery systems can potentially limit side effects that may otherwise occur with
systemic drug exposure. The polymers used in biodegradable drug delivery systems
are biocompatible, undergoing biodegradation to nontoxic metabolites or polymers
that solubilize in vivo and can be eliminated safely by endogenous metabolic
pathways in the human body without eliciting permanent, chronic foreign-body
9.3), particularly for treating sight-threatening
2008;
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9 Advances in Biodegradable Ocular Drug Delivery Systems
Characteristic Implications
Target specificity Local, sustained administration can achieve therapeutic
concentrations with drugs that otherwise may be poorly
absorbed and in tissues that are difficult to target by
conventional routes of administration (e.g., systemically/
topically). Minimizes drug exposure and side effects in
nontarget tissues
Mechanical strength Devices can maintain structural integrity under conditions of
mechanical stress
Biodegradability Eliminates the need for removal of inserted materials and risks
associated with such procedures
Biocompatibility Absence of foreign-body reactions
Drug compatibility Nonreactive with drug; stability of drug not affected
Controlled polymer degradation
and drug release rates
Allows for a range of delivery durations from weeks to years.
Long-term drug delivery systems eliminate the need for
frequent retreatment of chronic diseases
Consistent drug release Drug concentrations remain within the therapeutic range for
the desired time without significant variation (e.g.,
minimal burst effect)
Noninvasive procedures Placement causes minimal injury
Patient compliance One-time procedure; self-administration not required
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Table 9.3 Ideal characteristics of polymeric ocular drug delivery systems
reactions (Chu 2008). They also provide sufficient mechanical strength to withstand physical stress in vivo, which enables structural integrity to be maintained
for a sufficient duration during therapy. The release rates of drugs from biodegradable systems can be manipulated by choosing polymers with the desired biodegradation kinetics, physicochemical properties, and thermodynamic characteristics, as
well as by varying the shape of the delivery system (Park et al. 2005). Biodegradable
delivery systems with a long duration of action eliminate the need for frequent
retreatment, as occurs with intravitreally injected drugs; unlike nonbiodegradable
systems, they do not require removal once the drug supply is exhausted, thereby
eliminating the risks associated with such procedures (Conway 2008; Kimura and
Ogura 2001). Lastly, such devices remedy the issue of poor compliance with medication treatment regimens, which is problematic with orally and topically administered ocular medications.
Biodegradable devices are most often constructed from synthetic aliphatic polyesters of the poly-a-hydroxy acid family, which include poly(glycolic acid) (PGA),
poly(lactic acid) (PLA), and the PGA/PLA copolymer poly(lactic-co-glycolic acid)
(PLGA) (Fig. 9.1). Other biodegradable polymers used as biomaterials include
poly(e-caprolactone) (PCL), poly(glycolide-co-lactide-co-caprolactone) (PGLC),
poly(ortho esters) (POE), polyanhydrides (PAH), polymethylidene malonate
(PMM), polypropylene fumarate (PPF), and poly(N-vinyl pyrrolidone) (PVP).
These polymers have been explored for ocular drug delivery systems, but the technologies are still in the early stages of development and none have been marketed
commercially (see Sect. 9.6).
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Fig. 9.1 Chemical structures
of biodegradable polymers
commonly used in
biomedical applications
9.3.1 Polylactic Acid, Polyglycolic Acid,
and Polylactic-Co-Glycolic Acid
The aliphatic poly-a-hydroxy acids – PLA, PGA, and their copolymer PLGA – are
the most widely studied of the synthetic biodegradable polymers. This family of
polymers shows suitable biocompatibility and biodegradability and is versatile for a
range of biomedical applications. Polymers from this family have been approved by
the US FDA for drug delivery use (Jain et al. 1998).
The discovery and synthesis of lactide- and glycolide polymers was first reported
several decades ago, and during the late 1960s and early 1970s their application as
suture biomaterials was first described (Jain 2000). The polymers showed several
useful characteristics such as good mechanical properties, low immunogenicity
and toxicity, excellent biocompatibility, and predictable biodegradation kinetics
(Jain 2000). The widespread use of lactide/glycolide polymers as suture materials
led to interest in their use for other biomedical applications – initially as orthopedic
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surgical biomaterials (e.g., bone screws and pins) and later for drug delivery
systems (e.g., implants, hydrogels, micro/nanoparticles).
PGA is a semicrystalline polymer with a high melting temperature and low solubility (Bowland et al. 2008). PGA is synthesized using toxic solvents (hexafluoroisopropanol or hexafluoroacetone sesquinhydrate) which can limit its potential as a
biomaterial for drug delivery, since drugs could react with residual solvent, or trace
amounts of solvent could be incorporated into the end product (Bowland et al. 2008).
PLA is a hydrophobic synthetic polyester, which, unlike PGA, can exist as an optically active stereoregular polymer (l-PLA aka PLLA) and an optically inactive racemic
polymer (d,l-PLA aka PDLLA) (Jain et
al. 1998; Chu 2008). PLLA is semicrystalline because of the high regularity of its polymer chain and has an extremely slow
biodegradation rate in vivo, while the amorphous nature of PDLLA results from
irregularities in its polymer chain structure. Therefore, PLLA has higher mechanical
strength than noncrystalline PDLLA and is used mainly for surgical fixation devices
such as pins and screws, while PDLLA is preferable over PLLA for drug delivery
because PDLLA enables more homogeneous dispersion of the drug in the polymer
matrix (Jain et al. 1998). In comparison with PGA, PLA is more hydrophobic and
degrades more slowly due to the presence of methyl side groups.
PLGA, a copolymer of PLA and PGA, is the most widely utilized biodegradable
polymer for drug delivery. PLGA is synthesized by a random ring-opening copolymerization of the cyclic dimers of glycolic acid and lactic acid, whereby successive monomeric units of PGA or PLA are linked together by ester linkages. One
of the primary advantages of PLGA over other biodegradable synthetic polymers
is that the ratio of PLA to PGA used for the polymerization can be adjusted to
alter the biodegradation rate of the product (Fig. 9.2). The rate of drug release from
PLGA-based drug delivery implants depends on several factors, including the total
surface area of the device, the percentage of loaded drug, the water solubility of
the drug, and the speed of polymer degradation (Shive and Anderson 1997). The
three main factors that determine the degradation rate of PLGA copolymers are the
lactide:glycolide ratio (Fig.
9.2), the lactide stereoisomeric composition (i.e., the
amount of l-lactic acid vs. d,l-lactic acid), and the molecular weight (Chu 2008;
Avgoustakis 2008).
The lactide:glycolide ratio and stereoisomeric composition are most important for
PLGA degradation as they determine polymer hydrophilicity and crystallinity. These
properties can be manipulated during manufacturing to produce PLGA copolymers
with degradation times ranging from weeks to years (Avgoustakis 2008). Lactic acid
is more hydrophobic than glycolic acid; therefore, lactide-rich PLGA copolymers are
less hydrophilic, absorb less water, and subsequently degrade more slowly (Jain 2000).
The most widely used PLGA composition of 50:50 has the fastest biodegradation rate
(50–60 days) of the d,l-lactide/glycolide polymers and degrades faster than either
PLA or PGA (Yasukawa et al. 2006). PLGA copolymers with lactide:glycolide ratios
of 65:35, 75:25, and 80:20 have progressively longer in vivo lifetimes (Mundargi et al.
2008). Lactide-rich PLGA copolymers (up to 70%) are typically amorphous and
degrade more rapidly. As the molecular weight of the polymer decreases, the degradation rate increases due to the higher content of carboxylic groups at the end of the
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Fig. 9.2 Relationship between lactide/glycolide content and degradation half-life of PLGA. Graph
illustrates how PLA and PGA polymer ratios can be adjusted to result in a specific biodegradation
half-life and drug release from the implant
polymer chain, which accelerate hydrolysis (Park et al. 2005). PLGAs prepared from
PLLA and PGA are crystalline copolymers, and those from PDLLA and PGA are
amorphous (Jain et al. 1998). PLGA is degraded in vivo, through hydrolysis and enzy-
matic action, to lactic acid and glycolic acid which are ultimately converted via the
citric acid cycle to water and carbon dioxide (Fig. 9.3).
The versatile properties of PLGA have led to its use in the production of a variety
of biomedical devices. PLGA can be used to construct biomaterials of various types
and shapes such as resorbable suture materials; rods, screws, plates, and pins for
orthopedic surgery; vascular grafts and stents; and surgical meshes and scaffolding
for tissue regeneration (see Sect.
9.3). PLGA has also been widely utilized in the
development of various types of drug-release systems (e.g., drug implants and
micro/nano-particulates) for ocular diseases of the posterior and anterior segments.
PLA, PGA, and PLGA are cleaved predominantly by nonenzymatic hydrolysis
of their ester linkages throughout the matrix in the presence of water in the surrounding tissues. This process, referred to as bulk erosion (Fig. 9.4a–f), is distinguished from surface erosion (Fig. 9.4g–l) of the drug/polymer matrix surface
(Yasukawa et al. 2006) occurring with polymers such as PAH and POE. Drug release
from PLA- and PLGA-based matrix drug delivery systems generally follows pseudo
first-order or square-root kinetics, and the release rate is influenced by many factors
including polymer type, drug load, implant morphology, and porosity. In general,
drug release from PLGA-based implants, which is depicted in Figs. 9.4a–f and 9.5,
occurs in three phases:
1. Burst release: Drug release from the implant surface occurs, creating a short
period of high drug release.
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9 Advances in Biodegradable Ocular Drug Delivery Systems
Polylactic-co-glycolic acid
Polylactic acid
Polyglycolic acid
Glycolic acid
Lactic Acid
Acetyl-CoA
Citric acid cycle
H
2
OCO
2
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Fig. 9.3 Biodegradation of
poly(lactic-co-glycolic acid)
(PLGA) to glycolic acid and
lactic acid monomers, which
then enter the citric acid
cycle, and produce water and
carbon dioxide as byproducts
2. Diffusion and chain scission: Diffusional drug release, which is governed by the
3. Biodegradation and mass loss: Drug release is associated with biodegradation of
of continuous lower dose release makes such delivery systems ideally suited for
acute-onset diseases that require an initial loading dose of drug followed by tapering
doses over several months (Lee et al. 2008). More recent advancements in PLGAbased drug delivery systems have allowed for biphasic release characteristics with
an initial high (burst) rate of drug release followed by sustained zero-order kinetics
release; that is, the drug release rate from matrix is steady and independent of the
drug concentration in the surrounding milieu over longer periods (Kiernan and
Mieler 2009).
inherent solubility of the drug in the surrounding media, occurs. Random chain
scission of polymers occurs by hydrolytic cleavage, which increases the porosity
and surface area for drug diffusion.
the polymer matrix, mass loss initially occurring in the central core of the implant,
and a final burst in some delivery systems (Conway 2008; Kearns and Williams
2009; Yasukawa et al. 2006; Gaudana et al. 2009).
The rapid achievement of high drug concentrations followed by a longer period
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S.S. Lee et al.
Drug
SURFACE EROSIONBULK EROSION
a
b
c
d
e
f
g
h
i
j
k
l
BURST
BURST
DIFFUSION AND CHAIN SCISSION
BIODEGRADATION AND MASS LOSS
SURFACE EROSION AND MASS LOSS
SURFACE EROSION
H
2
O
Drug
H
2
O
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Fig. 9.4 Drug-release mechanisms and biodegradation of matrix implants. (a–f) Illustrate the bulk
erosion process: (a) dry implant prior to implantation shows porous structure of PLGA. Magnified
view (right) shows drug molecules (red spheres) interspersed in the pores and skeleton of the PLGA
polymer. (b) Burst phase of drug release is the short phase occurring immediately after implantation. Magnified view (right) shows water penetrating the pores of PLGA (black curved arrows and
squares) on the surface and drug molecules diffusing out of the implant on the surface (red arrows
and red spheres). (c) Diffusion and random chain scission phase of drug release. Implant swells
slightly as water molecules penetrate deeper into the core of the implant. The polymer is undergoing
random chain scission, where the long PLGA chains are cleaved at random locations. Magnified
view (right) shows water molecules (black curved arrows and squares) and drug molecules (red
arrows and spheres) entering and exiting the implant from the core, respectively. (d) Biodegradation
and mass loss phase is when the polymer begins to structurally break down from internal cavitation.
Magnified view (right) shows that much of the drug molecules have diffused out from the cavity.
(e) Continued biodegradation causes structural changes that alter the shape of the implant. Magnified
view (right) shows that water is still passing through the polymer and less drug is available for
release. (f) Implant fragments towards the end of biodegradation. Magnified view (right) shows that
water is still passing through the smaller polymer skeleton and even less drug is available for release.
(g–l) Illustrate the surface erosion process: (g) dry implant prior to implantation. Magnified view
(right) shows drug molecules (violet spheres) interspersed in the pores and skeleton of the polymer.
(h) Burst phase of drug release is the short phase occurring immediately after implantation.
Magnified view (right) shows water penetrating the pores of the polymer (green curved arrows and
squares) on the surface and drug molecules diffusing out of the implant on the surface (violet arrows
and violet spheres). (i) Surface erosion begins shortly after the burst phase. Drug and polymer are
solubilized only on the surface of the implant. (j–l) Continued surface erosion results in mass loss
from the surface of the implant. Drug and polymer are released and solubilized from the surface of
the implant, and implant volume and surface are gradually reduced over time
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