PLGA–lecithin–PEG core–shell nanoparticles for controlled drug delivery
Introduction
Over the past two decades, there has been a steady rise in the number of commercially available nanoparticle (NP) therapeutics [1], [2], [3]. Among these products, liposomal drugs [4] and polymer–drug conjugates [5] are two dominant classes, accounting for the majority of clinically approved products. Concurrently, the development of biodegradable polymeric NPs in the ∼100 nm range has become an increasingly exciting field in academic and industrial research [2], [3].
Doxil (liposomal doxorubicin) was the first Food and Drug Administration (FDA)-approved liposomal drug formulation for the treatment of AIDS associated with Kaposi's sarcoma in 1995 [6], [7], [8]. The benefits of liposomal formulations include their ability to encapsulate hydrophilic therapeutic agents at high loading efficiency, shield encapsulated drugs from external conditions, and also be coated with inert and biocompatible polymers such as polyethylene glycol (PEG) for reduced systemic clearance rates and prolonged circulation half-life in vivo [9]. These PEG-end groups may also be functionalized with specific ligands for targeting to specific sites of the cells, tissues and organs of interest [10]. However, liposomes may face some obstacles from their low ability to encapsulate very hydrophobic drugs, burst release of drugs, and having multiple manufacturing steps associated with liposome preparation and purification.
The use of biodegradable polymeric NPs for drug delivery has been gaining momentum and shown significant therapeutic potential [11], [12], [13]. Biodegradable polymers such as poly(d,l-lactic acid), poly(d,l-lactic-co-glycolic acid) and poly(ɛ-caprolactone) and their co-polymers diblocked or multiblocked with PEG have been commonly used to form core–shell structured NPs to encapsulate a variety of therapeutic compounds [14], [15], [16], [17]. These NPs have a number of appealing features: their hydrophobic core is capable of carrying highly insoluble drugs with high loading capacity, while their hydrophilic shell provides steric protection and functional groups for surface modification. Drug release can be manipulated by choosing biodegradable polymers with different surface or bulk erosion rates, and external conditions such as pH and temperature changes may function as a switch to trigger drug release [18]. Worth noting is their ease of manufacturing: amphiphilic co-polymers spontaneously assemble into core–shell structures in aqueous environments [19]. However, to date, polymeric NPs have shown moderate circulation half-lives compared to their liposomal counterparts, despite also being coated with inert and biocompatible polymers such as polyethylene glycol (PEG) [20].
Recently, we developed in our laboratory a protocol for the self-assembly of NPs that combine the properties of liposomes and polymeric NPs [21]. Existing strategies to make such fusion particles (lipopolyplexes) involved multi-step synthesis methods, resulting in inherently inefficient systems which are not easily scalable and show batch-to-batch variation [22], [23]. We engineered a simple, scalable, efficient and more controllable system using a well-defined and predictable formulation strategy. The NPs are formed from three biomaterials: (i) poly(d,l-lactide-co-glycolide) (PLGA) was selected for the hydrophobic core due to its biodegradable nature and ability to encapsulate high amounts of hydrophobic drugs; (ii) soybean phosphatidylcholine (or lecithin) was chosen for a monolayer around the hydrophobic core; and (iii) 1,2-distearoyl-sn-glycero-3-phosphoethanolamine-N-carboxy(polyethylene glycol)2000 (DSPE–PEG–COOH) intersperses in the lecithin monolayer to form a PEG shell which provides electrostatic and steric stabilizations, a longer circulation half-life in vivo as well as functional-end groups for the attachment of targeting ligands such as antibodies, peptides and aptamers.
Herein we report our studies on the systematic preparation and characterization of these PLGA–lecithin–PEG core–shell NPs. We evaluate parameters that affect the core–shell nanostructure to find an optimal formulation, and subsequently characterize the NPs for physical stability, controlled drug release kinetics, post-formulation purification, storage methods and finally material cytotoxicity. We establish the function of the lipid monolayer ring between the PLGA core and the PEG shell in controlled drug release kinetics. This novel NP system may represent a new way of combining existing lipid and polymer classes of materials for controlled drug delivery applications.
Section snippets
Materials
PLGA (poly(d,l-lactide-co-glycolide)) with a 50:50 monomer ratio, ester-terminated, and viscosity of 0.72–0.92 dl/g was purchased from Durect Corporation (Pelham, AL). Soybean lecithin consisting of 90–95% phosphatidylcholine was obtained from MP Biomedicals (Solon, OH), and DSPE–PEG2000–COOH (1,2-distearoyl-sn-glycero-3-phosphoethanolamine-N-carboxy(polyethylene glycol)2000) was obtained from Avanti (Alabaster, AL). Docetaxel (Dtxl) was purchased from Sigma–Aldrich (St. Louis, MO).
Preparation of PLGA–lecithin–PEG NPs
Formulation of PLGA–lecithin–PEG NPs
As shown schematically in Fig. 1A, a modified nanoprecipitation technique was used to prepare the NPs. The aqueous phase consisted of soybean phosphatidylcholine (sPC) and pegylated phospholipids (DSPE–PEG2000–COOH) in the appropriate molar ratio to form the monolayer around the PLGA polymeric core. The lipids were heated at 65 °C before adding the Dtxl/polymer mixture and subsequently vortexed for 3 min. This increase in thermal and mechanical energy allows for the lipids to disperse and
Conclusion
We developed PLGA–lecithin–PEG core–shell NPs which contain a hydrophobic PLGA core, a soybean lecithin monolayer and a hydrophilic PEG shell. The NP formulations were characterized and evaluated for controlled drug release kinetics and physical stability in PBS and plasma. The NPs were well tolerated by human cell line models, HeLa and HepG2. The NPs were prepared by a potentially scalable nanoprecipitation process with predictable and controllable outcomes and thus may be suitable as a
Acknowledgements
We thank the MIT Biopolymers Facility and the MIT-Harvard Nanomedical Consortium Nanomaterial Toxicity Core for their assistance on this project. We also thank Drs. Philip Kantoff and Neil Bander for helpful discussions on this project. This work was supported by National Institutes of Health Grants CA119349 and EB003647 and David Koch – Prostate Cancer Foundation Award in Nanotherapeutics. J.M.C. acknowledges the financial support from the Agency for Science, Technology and Research (A*STAR),
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