11.2.2  Liposomes and Lipid Nanoparticles

(Walker et al. 1997). The content within the spherical compartment of the liposome is largely dictated by the contents of the medium in which the liposomes are formed. The liposomal structure can be small, large or giant, and can be unilamellar (comprised of one lipid bilayer) or multilamellar (comprised of layers of lipid bilayers) depending on the method of preparation. Acronyms such as SUV (small unilamellar vesicles), LUV (large unilamellar vesicles), and MLV (multilamellar vesicles) are used to describe the nature of the vesicle. Liposomes can be synthesized using a variety of methods. Typically, the lipids obtained from a commercial source are dried under N2 followed by rehydration in a buffer (Wu et al. 2008). The buffer can consist of the drug desired to be encapsulated in the liposome or it can consist of other small molecule agents. Rehydration is performed under heating and vigorous vortexing in a buffer over approximately 4–16 h to allow for the lipids to enter solution. The temperature is set above the phase transition temperature of the lipid from the solid crystalline phase to the fluid crystalline phase. The size of the lipid can be greatly reduced by sonication after rehydration or by extrusion through a porous membrane with the pore diameter close to the desired liposome diameter. Thus, liposomal preparation can be a simple process whereby liposomes are ready within a day assuming the complexity of the liposome is minimal. For liposomes with targeting ligands or multiple compartments, synthesis will likely be much more complex and time-consuming.

Evading the reticuloendothelial system (RES) is a major hurdle in the development of nano-drug delivery systems. RES is responsible for the detection and elimination of foreign objects including nano-drug delivery systems by the phagocytic cells, macrophages, and primary monocytes. Several investigators have focused on determining the optimal particle size and composition of delivery devices that bypass the RES. Litzinger et al. in 1994 demonstrated that PEG-coated liposomes with a diameter of 150–200 nm have optimal blood retention levels, whereas liposomes within the range of 200–300 nm experienced highest uptake by the RES and were found in the spleen (Litzinger et al. 1994).

STEALTH liposomes, such as liposomes coated with PEG that are capable of evading RES at least in part, are FDA approved and have been implicated to improve target tissue availability of cytotoxic drugs (e.g., doxorubicin, Doxil®), while minimizing their nontarget tissue delivery. The PEG moiety provides for increased retention in the circulation as well as tissue compartments by preventing premature lipid degradation by lipsases (e.g., phospholipase) and providing a steric barrier that prevents liposomes from interacting with opsonins and macrophages that are involved in liposomal clearance (Papahadjopoulos et al. 1991). However, the precise mechanism by which PEG increases circulation time and tissue drug retention is highly debated among scientists in this field. Interestingly, PEG-coated 100 nm liposomes (DoxilTM) can sustain the release of doxorubicin across sclera when compared to the plain drug (Kim et al. 2009a).

Several investigators have realized the potential for liposomes as nonviral vectors. As early as 1998, it was discovered that positively charged lipids are capable of binding to deoxyribonucleic acid (DNA) and releasing them into the cell for gene expression (Koltover et al. 1998). It was later discovered that PEG-coated

lipid–DNA complexes had extended circulation times and tumor-selective gene expression due to increased blood circulation time to reach the tumor site (Ambegia et al. 2005). Most recently, several of the primary investigators in this field including Drs. Ian MacLachlan and Pieter Cullis have developed a rational design approach to develop lipids for siRNA delivery (Semple et al. 2010). The use of liposomes for gene delivery and transfection is a prominent field and although there have been several breakthroughs, the field is still in its infancy and little is known regarding the mechanism of transfection and cellular uptake. The potential for liposomes to be marketed is largely hindered by this “black box” in which controlling the retention and cell uptake properties are poorly understood.

Liposomal delivery devices for ocular drug delivery have been under investigation at least since 1981 and various routes of administrations have been used in the comparison of liposomal formulations vs. traditional therapeutic formulations (Meisner and Mezei 1995). Smolin et al. in 1981 found that topical administration of idoxuridine in the liposomal formulation was much more effective than in its marketed formulation in the treatment of acute and chronic herpetic keratitis in rabbit (Smolin et al. 1981). Peyman et al. further validated this study in 1986 and confirmed that liposomal formulations of idoxuridine were able to penetrate the cornea much better than the marketed formulation (Dharma et al. 1986). The administration of liposomal formulations by subconjunctival (SC) injection in rabbit was also evaluated by other investigators in 1991 (Hirnle et al. 1991). Liposomal formulations made with phosphatidylcholine (PC) cholesterol (Chol) dicethyl phosphate (DP) lipids encapsulating carboxyfluorescein (CF) were injected subconjunctivally and 30 min after injection, ~87.4, 5.8 and 2.3% of CF was distributed to the sclera, retina and choroid tissues totaling in 95.5% of total CF distribution in the intraocular tissues of the posterior segment. The following data describe the concentrations of CF in each tissue after administration of the liposomal formulation. After 1 h, the amount of CF at the site of injection dropped from ~600 to 275 mg/g. In the sclera, the amount of CF at 1 h was ~40 mg/g and at 6 h reduced to ~15 mg/g and in the retina-choroid after 1 h ~8 mg/g of CF was present and after 6 h, ~6 mg/g was present. However, CF levels as high as 5 mg/g persisted in the retina-choroid for at least 7 days. At 1 h only ~7.5 mg/g of CF was found in the cornea and decreased to ~2 mg/g at 6 h. In the iris-ciliary body, at 1 h ~8 mg/g of CF was present and decreased to ~1.5 mg/g at 6 h. Less than ~2 mg/g of CF was in the vitreous and less than ~2.5 mg/g of CF was in the anterior chamber over the entire time period 0–6 h. Concentrations of CF after administration of aqueous CF (nonliposomal) were much less in all ocular tissues. After subconjunctival injection of CF, the concentration dropped to ~5 mg/g after 1 h; at 1 h, the concentration in the sclera, retinachoroid, cornea, iris-ciliary body, anterior chamber and vitreous was ~8, 2, 0.2, 2, 0.2, and 0.2 mg/g, respectively. The liposomal formulation of CF was found to have higher drug retention in all ocular tissues compared to aqueous CF, but the concentrations within the posterior segment of the eye was highest.

Liposomal formulations were shown to be much more effective than unencapsulated drugs as early as 1986 and continue to show superiority over small molecule ocular therapeutics. Topical administration of positively charged and neutral

acetazolamide MLVs consisting of phosphatidylcholine:cholesterol:stearylamine in a molar ratio of 7:4:1 were much more effective in lowering intraocular pressure (IOP) than negatively charged MLVs of the same composition and aqueous acetazolamide (Hathout et al. 2007). In 2009, another study investigated the potential use of a chitosan-coated liposomes encapsulated with diclofenac sodium for sustaining drug release (Li et al. 2009). The authors used 0.25 and 0.5 mol% 540 kDa chitosan and the diameter was 82.4 ± 2.2 and 84.0 ± 4.7 nm, respectively. Both liposomal formulations released only 50% of the drug over 24 h, whereas the nonchitosancoated diclofenac liposomes and aqueous diclofenac released 60% and 90%, respectively, over 24 h. The concentrations of diclofenac in precorneal regions after 360 min for chitosan-coated diclofenac liposomes (both 0.25 and 0.5 mol% chitosan), noncoated diclofenac liposomes, and aqueous diclofenac dropped from 16 to 6 mg/ml, 15 to 2 mg/ml, and 13 to 1.5 mg/ml, respectively. The cumulative penetration for the 0.25 mol% chitosan-coated diclofenac liposomes, 0.5 mol% chitosancoated diclofenac liposomes, noncoated diclofenac liposomes, and aqueous diclofenac was 35, 32, 22 and 25 mg/cm2, respectively. From these studies, the importance of liposomal surface charge and composition in determining tissue penetration, drug retention, and drug concentration in ocular tissues is apparent.

Similar to polymeric systems, laser irradiation-triggered release is also feasible with liposomal systems. Zasadzinski et al. have showed that liposomal release can be controlled by triggering the release by activating gold nanoparticles encapsulated within the liposome (Wu et al. 2008). Further, by modifying lipids with targeting elements, targeted liposome delivery systems can be prepared. Thus, liposomes offer a commercially viable approach for encapsulation and sustained drug delivery.

Alternative lipid-based nanosystems include solid lipid nanoparticles and nanostructured lipid carriers (Pardeike et al. 2009). In this case, similar to PLGAand PLA-based nanoparticles, lipids are formulated into solid nanoparticles or nanoparticles with defined nanostructure, without the core (liquid) and coat (lipid) structure of liposomes. Similar to polymeric systems as well as liposomes, solid lipid nanoparticles and nanostructured lipid carriers can also be surface modified to improve circulation time or to enhance cell uptake.

11.2.3  Micelles

Similar to liposomes, micelles are formed by self-assembly of amphiphilic molecules; however, micelles exhibit a slightly different higher-ordered structure (Israelachvili et al. 1975). Micelles may assemble in three different distinct structures (Trivedi and Kompella 2010): (1) standard micelles, where the hydrophilic portion is oriented toward the solvent to form the shell of the micelle and the hydrophobic portion is oriented away from the solvent to form the core of the micelle,

(2) inverse micelles, where the hydrophobic parts form the shell of the micelle and the hydrophilic parts form the core of the micelle, and (3) unimolecular micelles,

where one molecule with hydrophilic and hydrophobic blocks forms the entire micelle. On exposure to aqueous solvent, the hydrophobic parts are clustered in the center of the structure to form the core of the micelle and the hydrophilic parts are primarily on the outside of the structure to form the shell of the micelle. Depending on the solvent and type of amphiphilic molecules used, different structures will predominate. For example, in hydrophobic solvents, the inverted micelle will predominate so that the hydrophobic interactions between the hydrophobic parts of the micelle and the solvent are maximized while the hydrophilic portions of the molecule are oriented away from the solvent.

The propensity of an amphiphilic molecule to form a micelle will largely depend on the characteristics of the molecule (Lukyanov and Torchilin 2004). Polyethyleneglycol (PEG) is a polymer commonly used in the synthesis of lipid drug carriers due to its ability to reduce particle aggregation and prolong systemic exposure time. PEG between the molecular weights of 750 and 5,000 when ligated to phosphatidylethanolamine (PE) lipids can spontaneously form micelles at a concentration above the critical micellar concentration (CMC). For example, PEG(750)- distearoyl-phosphatidyl-ethanolamine (DSPE) has a CMC of 0.1 mM and forms 7–15-nm-sized micelles. These micelles assemble spontaneously when PEG–DSPE dry lipid is hydrated for several hours (Lukyanov and Torchilin 2004). Methods such as sonication or extrusion can be performed to further control the size of the micelles.

Several authors (Kataoka et al. 2000; Liu et al. 2000; Yoo and Park 2001) have demonstrated the ability of different micelle structures to sustain the release of drug molecules over a few days. A prodrug of paclitaxel loaded in PEG-b-poly (e-caprolactone) micelles (PEG-b-PCL) of 27–44 nm diameter resulted in ~40% cumulative release of the prodrug in 14 days (Forrest et al. 2008). These paclitaxel prodrug-loaded micelles were also shown to have a slightly higher anticancer activity than free paclitaxel as well as prodrug cremophor formulation in breast cancer cell lines. In serum, the total concentration of paclitaxel was reduced threefold in ~25 h for the cremophor formulation of paclitaxel prodrug and in ~50 h for the PEG-b-PCL micellar formulation of paclitaxel prodrug. These paclitaxel prodrugloaded PEG-b-PCL micelles were fabricated by dissolving PEG-b-PCL and paclitaxel with a minimal volume of acetone and adding it dropwise with the help of a syringe pump to vigorously stirred distilled water. The organic solvent was removed by stirring under air purge and the micelles were purified by extruding through a membrane filter with a 0.22-mm pore size.

Indomethacin-loaded micelles using methoxy-PEG/e-caprolactone (e-CL) block copolymers (MePEG/e-CL) have been shown to be superior over plain drug in sustaining drug release (Kim et al. 1998). The indomethacin-loaded MePEG(5000)/e-CL micelles were synthesized by dissolving the MePEG/e-CL block copolymer in an organic solvent and then indomethacin was added to the mixture and stirred at room temperature. The micelles were dialyzed and then centrifuged to remove the unencapsulated drug. The size of the resulting micelles ranged from 54 to 180 nm and the size depended on the molecular weight of the copolymer and the solvent in which it was dissolved to form micelles. As the molecular weight of the copolymer