The efficacies of conventional drugs can usually be increased through the use of drug delivery systems (DDS). DDS intended for systemic administration include micro- and nanoparticles (capsules, shells, spheres, drug-polymer conjugates), liposomes, micelles and dendrimers. Such carrier systems are advantageous as they modify the biodistribution and pharmacokinetics of their cargo by granting protection from degradation, prolonging circulation time, sustaining drug effects in target tissue, enhancing solubility and reducing harmful side effects [85-89].
The small sizes of these DDS along with the increased circulation time allow passive targeting occur which help these DDS to reach their target sites via the enhanced permeability retention (EPR) effect [90-92]. The EPR effect occurs because of leaky vasculature present at tumour sites or inflamed tissues which leads to the gradual accumulation at these locations. Such drug delivery systems may also employ active targeting whereby a chemical moiety, which specifically interacts with target cells, is conjugated to the carrier. All these lead to greater drug efficacy at the desired sites of action, which negates the need for high doses of free-drug and thus reducing the consequential harmful side effects.
Nanoparticles have great potential for in vivo therapeutic applications (e.g. drug delivery) as well as diagnostic applications (e.g. cell tracking) [93]. The rates of drug release from their carrier enhance the effectiveness of the drug delivery system. Drugs within carriers are inactive and need to be released to have any therapeutic
effect. If the drug is released too quickly in the body, the pharmacokinetics and distribution of the DDS might be similar to administration of the drug in its ‘free’ form. And if the drug is released too slowly, the delivery system may have lower therapeutic effects than unencapsulated drug [94,95]. For active or passive targetted DDS, it is important that the particles have long enough half-lives in vivo to accumulate at their target tissue. Therapeutic levels of free drug must also be maintained and sustained for a period of time once at the target site. Different drugs have different mechanisms of action and will require tailored release profiles. DDS need to be initially assessed in vitro to determine their release rates, half-lives, how cells interact with them and their toxicity before they are deemed safe for in
vivo studies.
Some drugs need to be delivered into the cell rather than released from their carrier extracellularly [96]. This would require passage through biological barriers such as epithelia and mucosa, to reach target tissues. The encapsulated drug would then have to enter the cell through the plasma membrane and to the target organelle. Cells take up nanoparticles via several different processes. Most cell types can take up nanoparticles through endocytosis; clathrin-mediated, caveolae- mediated, macropinocytosis and others (clathrin-caveolae-independent endocytosis).
Clathrin-mediated endocytosis (CME) may be receptor-dependent or receptor- independent. Receptor-dependent CME provides a way of targetted entry into cells through the attachment of specific ligands to the surfaces of nanoparticles such as
low density lipoprotein (LDL), epidermal growth factor (EGF) or folic acid (FA) [97]. Folic acid binds with a low affinity to the folate receptor present in most cells, but with a high affinity to the glycosylphosphatidylinositol-linked folate receptor often overexpressed in cancer cells [98,99]. This makes FA a popular ligand for surface modification of anticancer nanoparticles. CME intake of material end up in lysosomes containing acid hydrolases and low pH. This can be utilised as a release mechanism through carrier biodegradation within lysosomes, though the drug molecule must be tailored to resist the harsh lysosomal environment. Caveolae- mediated endocytosis (CvME) leads to the formation of cytosolic caveolar vesicles, which provides a ‘safe’ route into the cell, avoiding the lysosomes. Macropinocytosis is a non-selective, clathrin-independent endocytic mechanism which involves micrometer-sized endocytic vesicles. These large vesicles acidify and may fuse with lysosomes. Phagocytosis is another mechanism of uptake but is limited to certain cell types (macrophages, monocytes, dendritic cells, neutrophils, fibroblasts, epithelial and endothelial cells).
Nanoparticle size can affect entry into cells. Polystyrene nanoparticles within the size range of 20 – 1000 nm were not endocytosed by human umbilical vein endothelial cells, whereas 20 – 100 nm particles were endocytosed by HepG2 hepatocytes [100]. Caco-2 cells take up 100 nm PLGA particles more readily than 500 – 10000 nm PLGA particles [101]. Drug-carrying nanoparticles can be made more effective by surface charge manipulation. Due to the negative charge on cell membranes, positively-charged carriers show greater cell interactions and uptake. PLA-PEG nanoparticles with positively-charge stearylamine were taken up by HeLa
cells at a faster rate and to a greater degree than the uncoated, negatively-charge PLA-PEG nanoparticles [102]. Uptake by phagocytic cells can be altered by adjusting the hydrophilicity/hydrophobicity of the nanoparticle surface to influence interactions with opsonins, and in turn, phagocytosis.
Thus, nanoparticles can be modified in many ways to increase the efficiency of drug delivery, but this has to be balanced with any potential toxic effects. Testing nanoparticles in models of biological barriers and tissues would allow the assessment of naoparticles in an in vivo-like environment for the development of nanoparticles with predictable and reproducible behaviour and low toxicity.