Elsevier

Biomaterials

Volume 32, Issue 13, May 2011, Pages 3435-3446
Biomaterials

The effect of surface charge on in vivo biodistribution of PEG-oligocholic acid based micellar nanoparticles

https://doi.org/10.1016/j.biomaterials.2011.01.021Get rights and content

Abstract

To systematically elucidate the effect of surface charge on the cellular uptake and in vivo fate of PEG-oligocholic acid based micellar nanoparticles (NPs), the distal PEG termini of monomeric PEG-oligocholic acid dendrimers (telodendrimers) are each derivatized with different number (n = 0, 1, 3 and 6) of anionic aspartic acids (negative charge) or cationic lysines (positive charge). Under aqueous condition, these telodendrimers self-assemble to form a series of micellar NPs with various surface charges, but with similar particle sizes. NPs with high surface charge, either positive or negative, were taken up more efficiently by RAW 264.7 murine macrophages after opsonization in fresh mouse serum. Mechanistic studies of cellular uptake of NPs indicated that several distinct endocytic pathways (e.g., clathrin-mediated endocytosis, caveolae-mediated endocytosis, and macropinocytosis) were involved in the cellular uptake process. After their cellular uptake, the majority of NPs were found to localize in the lysosome. Positively charged NPs exhibited dose-dependent hemolytic activities and cytotoxicities against RAW 264.7 cells proportional to the positive surface charge densities; whereas negatively charged NPs did not show obvious hemolytic and cytotoxic properties. In vivo biodistribution studies demonstrated that undesirable liver uptake was very high for highly positively or negatively charged NPs, which is likely due to active phagocytosis by macrophages (Kupffer cells) in the liver. In contrast, liver uptake was very low but tumor uptake was very high when the surface charge of NPs was slightly negative. Based on these studies, we can conclude that slightly negative charge may be introduced to the NPs surface to reduce the undesirable clearance by the reticuloendothelial system (RES) such as liver, improve the blood compatibility, thus deliver the anti-cancer drugs more efficiently to the tumor sites.

Introduction

Different types of nanoparticles (NPs), including liposomes, polymeric NPs, micellar NPs, albumin-based particles, inorganic or other solid particles (gold, iron oxide, quantum dots and carbon nanotubes) have been widely used as drug delivery vehicles for the diagnosis and targeted therapy of cancers. Early clinical results suggest that some nanoparticle therapeutics can enhance the therapeutic efficacy of delivered drugs while reducing their side effects, which can be explained by the preferential delivery of loaded drugs to tumor sites via the enhanced permeability and retention (EPR) effect [1]. The physicochemical characteristics of NPs such as composition, particle size, surface charge and surface hydrophobicity may affect their interaction with plasma proteins (opsonins) and blood components (hematocompatibility), uptake and clearance by macrophages, and hence potentially influence their biodistribution and targeted delivery of payload to the intended target sites [2]. The desired particle size of NPs for passive tumor targeting has been reported to be around 10–100 nm [3]. Hydrophilic polymers such as polyethylene glycol (PEG) have been widely used to coat the surface of NPs, in order to minimize the rapid opsonization and subsequent sequestration of NPs by macrophages in the reticuloendothelial system (RES). PEG surface coating can counteract the hydrophobic and electrostatic interactions between NPs and plasma proteins or macrophages, resulting in less RES uptake and prolonged blood circulation time [4], [5], [6]. Surface charge is usually introduced onto certain types of NPs (such as iron oxide and gold) to improve stability and prevent from further aggregation in aqueous solution via the electrostatic repulsion [7], [8]. It has been reported that surface charge is a very important factor to determine the efficiency and mechanism of cellular uptake, and the in vivo fate of NPs [6], [9], [10], [11], [12]. However, the optimum surface charges (e.g. positive, neutral or negative) and charge densities were reported differently for different nanoparticle systems, in order to prolong the blood circulation time, minimize the non-specific clearance of NPs and prevent their loss to undesired locations. For example, Juliano et al. [10] reported that neutral and positively charged liposomes were cleared less rapidly than negatively charged ones, which could be explained by the tendency of negatively charged liposomes to coalesce in the presence of proteins and calcium ion in blood plasma. Conversely, Yamamoto et al. [13] demonstrated that both neutral and negatively charged PEG-PDLLA micelles exhibited no remarkable difference in their blood clearance kinetics; however, negatively charged micelles significantly reduced the non-specific uptake by liver and spleen, compared with neutral micelles, which was attributed to the electrostatic repulsion between negatively charged micelles and cellular surface. The inconsistent results from the above studies may be due to the difference of nanoparticle types, variation in stability of NPs resulted from surface charge, the nature of charged groups, and other confounding factors such as inhomogeneous particle sizes.

He et al. systematically studied the effects of particle size and surface charge on cellular uptake and biodistribution of chitosan derivative polymeric NPs [11]. However, the NPs applied in this study had large particle sizes (150–500 nm), which led to significant high liver uptake regardless the surface charges. We have recently developed a novel micellar nanocarrier with desired narrow-dispersed particle sizes of 20–60 nm for effective tumor targeting drug delivery with minimum liver uptake [14], [15], [16]. These NPs are formed by the self-assembly of novel linear-dendritic block copolymers (named as telodendrimer) with engineerable and well-defined structures, comprising polyethylene glycol (PEG) and dendritic cholic acids (CA). PEG5k-CA8 is a representative telodendrimer with optimal properties, where “5 k” represents the molecular weight of PEG (5000 Da) and “8” indicates the number of CA subunits in the telodendrimer. PEG5k-CA8 micelles exhibited high drug loading capacity, outstanding stability, preferential tumor accumulation via EPR effects, and superior anti-tumor effects when loaded with paclitaxel (PTX) in the human ovarian cancer (SKOV-3) xenograft mouse model [14].

To optimize our nanocarriers for efficient in vivo cancer drug delivery, we systematically studied the effects of particle surface charges on their in vitro cellular uptake by macrophages, cytotoxic effects, hemolytic properties and in vivo biodistribution in xenograft models. Different number (n = 0, 1, 3 and 6) of anionic d-aspartic acids (d) or cationic d-lysines (k) were conjugated onto the distal end of PEG chain in PEG5k-CA8 telodendrimer (the micellar subunit) to modulate the surface charge of the micellar NPs. This allowed us to systematically evaluate the effect of surface charge on the cellular uptake and in vivo biodistribution of NPs under the identical conditions, e.g. the same composition and similar particle sizes. The particle sizes and surface charges (zeta potential) of aspartic acids or lysines derivatized NPs were characterized by transmission electron microscopy (TEM) and dynamic light scattering (DLS), respectively. The uptake efficiencies, pathways and intracellular fates of different charged PEG5k-CA8 NPs were examined in RAW 264.7 murine macrophages. The hemolytic properties and in vitro cytotoxicities against RAW 264.7 cells of these nanoparticle preparations were also evaluated. Finally, the in vivo biodistribution and tumor targeting efficiency of different charged PEG5k-CA8 NPs after intravenous administration were investigated in nude mice bearing SKOV-3 human ovarian cancer xenograft via NIRF optical imaging.

Section snippets

Materials

Diamino polyethylene glycol (Boc-NH-PEG-NH2, MW = 5000 Da) was purchased from Rapp Polymere (Tübingen, Germany). Fmoc-d-Asp(Otbu)-OH, Fmoc-d-Lys(Boc)-OH, and Fmoc-Lys(Fmoc)-OH were purchased from Anaspec, Inc. Hydrophobic NIRF dye DiD (1,10-dioctadecyl-3,3,30,30-tetramethylindodicarbocyanine perchlorate, D-307), 4′, 6-diamidino-2-phenylindole (DAPI) and LysoTracker® Green DND-2 were purchased from Invitrogen. Paclitaxel (PTX) was purchased from AK Scientific Inc. (Mountain View, CA). Cholic

Preparation and characterization of different charged PEG5k-CA8 NPs

To systematically modulate the surface charge density of PEG5k-CA8 NPs, different number (n = 0, 1, 3 and 6) of anionic d-aspartic acids (d, negative charge) or cationic d-lysines (k, positive charge) were chemically conjugated to the distal end of PEG strand constituting the shell layer of the micellar NPs, respectively (Fig. 1). After the conjugation of aspartic acids or lysines onto the PEG chain of PEG5k-CA8 telodendrimer, the N-terminal amino group was acetylated. The 1H NMR spectra of

Discussion

For the drug delivery application in cancer therapy, the non-specific uptake of NPs by macrophages in the RES may be minimized through manipulation of particle size and surface charge [19], [22]. For example, we have recently demonstrated that the biodistribution of PEG-oligocholic acid based micellar NPs depends greatly on the particle size [15]. The smaller PTX-loaded NPs (17–60 nm) preferentially accumulated in the tumor site, whereas most of the larger NPs (150 nm) were trapped by the liver

Conclusions

PEG5k-CA8 NPs with high surface charge, either positive or negative, tend to be taken up non-specifically by macrophages in vitro and in vivo, resulting in high uptake of NPs in the liver after systematic administration. In contrast, PEG5k-CA8 NPs with slightly negative charge (one aspartic acid per telodendrimer) demonstrated preferential uptake at the tumor site but not normal organs including liver. We believe that the optimal surface charge of NPs designed for use as drug delivery carrier

Acknowledgments

The authors thank the financial support from NIH/NCI R01CA115483, R01CA140449 and US Department of Defense Breast Cancer Research Program Postdoctoral Training Award (W81XWH-10-1-0817).

References (31)

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1

Kai and Yuanpei contributed equally to this work.

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