Controlling Drug Release 14

Spatiotemporal control of drug release is another fundamental feature of NPs as a drug carrier. A number of NP designs are predicated on the assumption that a NP

will retain a drug during circulation until it reaches the target cells. Otherwise, the therapeutic effect of NPs may not be different from that of the free drug (Gullotti & Yeo, 2012). However, initial burst release occurs in almost all types of delivery systems, especially in NPs with relatively large surface areas per volume ratios (Bae & Yin, 2008; Dai et al., 2011; Hasan et al., 2007; Yeo & Park, 2004). Therefore, various efforts are made to ensure stable drug retention in NPs in blood. At the same time, stimuli-responsive systems are pursued in parallel to make the carried drugs available at target locations in a timely manner.

Incorporating a diffusion barrier on NP surface. To prevent the initial burst

release and provide sustained drug release, a self-assembled layer of water-insoluble material is added as a diffusion barrier. For example, a lecithin layer was used to coat a PLGA core encapsulating deocetaxel, resulting in a hybrid polymer-lipid NP with an attenuated docetaxel release as compared to bare NPs (Zhang et al., 2008). Similarly, the release of doxorubicin entrapped in PLGA core was suppressed with an external layer of diethylenetriaminepentaacetic acid (DTPA)-gadolinium lipid, which also served as a paramagnetic image contrast agent (Liao et al., 2011).

Crosslinking. Parts of a NP can be crosslinked to improve the stability of self-

assembled NP structures. Polymeric micelles are prone to disassembly and

dissociation upon introduction into the blood stream due to interactions with amphiphilic blood components and dilution below their critical micelle concentrations (Deng et al., 2012). Burt et al studied biodistribution of polymeric micelles based on a block co-polymer of D,L-lactic acid and methoxypolyethylene glycol (PDLLA-MePEG) encapsulating paclitaxel (Burt et al., 1999). They reported that the drug and polymer showed distinct biodistribution profiles, indicating premature drug release and micelle dissociation (Burt et al., 1999). The stability of polymeric micelles can be improved by crosslinking of the core, shell or the interface (Deng et al., 2012). One approach is to crosslink the micelle core made of anionic polymethacrylate via metal cations such as Ca2+ _{(Bronich et al., 2005). In another example, the PLA core of a polymeric micelle was}

stabilized by introducing methacrylol end groups at the terminus of the PLA block, which could be covalently crosslinked after micelle formation (Iijima et al., 1999). These micelles had high stability even in presence of

surfactants (Iijima et al., 1999). Instead of NP core, the interface between hydrophilic shell and hydrophobic core was stabilized via crosslinking mediated by UV irradiation (Yang et al., 2011). For this purpose, photocrosslinkable polymer, poly(acryloyl carbonate), was introduced as the center block of a block co-polymer. Thus formed PEG-poly(acryloyl carbonate)-polycaprolactone (PEG-PAC-PCL) formed more stable micelles in both size and drug retention as compared to those made with non-crosslinkable counterpart (PEG- PCL) (Yang et al., 2011). Due to the improved stability, paclitaxel loaded in the photo- crosslinked PEG-PAC-PCL micelles showed greater in vivo tumor inhibition activity than paclitaxel in PEG-PCL micelles (Yang et al., 2011).

Stimuli-responsive systems. To confine drug release specifically to target

tissues or organelles, stimuli-responsive NPs are developed. NPs are designed to change their physicochemical properties in response to intrinsic conditions of target locations (more details in section 1.4) or external stimuli focused on the targets (R. Cheng et al., 2013). For example, a drug was encapsulated in mesoporous silica NPs, capped with large molecules, such as cadmium sulfide (CdS) nanocrystals (Lai et al., 2003) or PAMAM dendrimers (Gruenhagen et al., 2005) via disulfide bond. These capping materials prevented drug release but were removed in a reductive environment inside the cell. In another example, reduction-sensitive NPs encapsulating doxorubicin were prepared with dextran-lipoic acid derivative crosslinked via disulfide bonds (Y. L. Li et al., 2009). This system showed minimal drug release under extracellular condition and fast release in reductive environments, verified with in vitro release experiments (Y. L. Li et al., 2009). Similarly, cathepsin B, a lysosomal cysteine

proteinase, was used to trigger drug release from NPs in the lysosomes (Meng et al., 2012). Cathepsin B is also overexpressed in some tumor cells, present in the extracellular environment and on the cell surface of tumors (Campo et al., 1994; Sloane, 1990). Therefore, a peptide linker specifically degraded by cathepsin B was incorporated in a PEG NP system via photocrosslinkable diacrylate for tumor selective drug release (Glangchai et al., 2008b). Matrix metalloproteinases (MMPs), another type of enzymes over-expressed in tumor microenvironment, are also used for controlling drug release (Danhier et al., 2010). Doxorubicin was loaded in mesoporous silica NPs, coated with conjugates of PEG diacrylate and MMP-degradable peptides with different

sensitivities (Singh et al., 2011). Drug release and cytotoxicity of doxorubicin-loaded NPs were proportional to degradability of the peptide linker, and the MMP-sensitive NPs resulted in greater cell death in MMP-overexpressing tumors as compared to MMP- insensitive PEG-coated NPs (Singh et al., 2011). Acidic pH of endo/lysosomes is frequently used to achieve intracellular drug release. Mesoporous silica NPs were capped with calcium phosphate, soluble in pH 4-5 but insoluble in pH 7.4. To limit the calcium phosphate deposition on the surface, the silica NPs were conjugated with urease, which hydrolyzed urea and created a high pH zone around the NPs forming a layer of calcium phosphate (Rim et al., 2011). The coated NPs significantly attenuated drug release at pH 7.4 but released drug intracellularly, indicated by sustained nuclear localization of doxorubicin (Rim et al., 2011).

Multiple-stimuli responsive systems. To increase the selective reactivity of the

NPs, two or more mechanisms are utilized either simultaneously or sequentially. NPs are engineered to respond to dual stimuli, such as pH and temperature, pH and reductive potential, or temperature and enzyme (R. Cheng et al., 2013; Dai et al., 2011;

Sankaranarayanan et al., 2010). For example, polymeric NPs with a dual pH-sensitivity were formulated using a random co-polymer that degraded via both bulk dissolution and surface degradation at weakly acidic pH (Sankaranarayanan et al., 2010). The advantage of this system is the good stability at physiological pH and quick onset of degradation, which are often conflicting with each other (Sankaranarayanan et al., 2010). In another study, polymeric micelles dual-responsive to acidic endosomal pH and intracellular reductive potential were prepared using a tri-block copolymer made of a pH-sensitive hydrophobic block, disulfide-crosslinkable middle block, and PEG (Dai et al., 2011). This micelle system showed minimal drug release at pH 7.4 and increasing drug release in response to dithiothreitol (10 mM) and/or pH 5 (Dai et al., 2011). Taking this approach a step further, triple-stimuli responsive micellar system was developed using a block copolymer with a pH-sensitive hydrophobic block and a temperature-sensitive hydrophilic block connected v ia a reduction-sensitive disulfide linker (Klaikherd et al., 2009). The block-copolymer lost amphiphilic properties in response to temperature increase, acidic pH, or high reductive potential, allowing for tunable control of drug release with single stimulus or simultaneous

multiple stimuli. Notably, individual stimulus caused slow or incomplete dye release, but combined stimuli resulted in a significantly faster and greater drug release (Klaikherd et al., 2009).

Covalent conjugation of a drug to carrier. For stable drug encapsulation, a drug

may be conjugated to a NP matrix via a linker, which may be hydrolyzed or cleaved in a stimuli-responsive manner. pH-activatable NPs was made with paclitaxel conjugated to PEG-poly(acrylic acid) via an acid-labile acetal linker (Gu et al., 2013). The NPs showed >80% drug release in pH 5 in 2 days, while they released only 29% of the total drug at pH 7.4 (Gu et al., 2013). In another study, drug-polylactide conjugates were synthesized using a drug (paclitaxel, docetaxel, and camptothecin) as an initiator of polymerization (Tong & Cheng, 2008). The conjugates formed NPs with high drug contents (5-36 wt%), which showed minimal initial burst release followed by gradual drug release over a week in vitro (Tong & Cheng, 2008). Similarly, docetaxel was conjugated to PEGylated

carboxymethylcellulaose (CMC) to assemble into 120 nm NPs releasing the drug in a controlled release manner (Ernsting et al., 2011). Notably, the CMC-based NPs had an anti-stromal effect, increasing tumoral perfusion and lowering the IFP, with a greater anti- metastatic effect than Abraxane (Murakami et al., 2013). In another study, camptothecin was conjugated to a β-sheet-forming peptide to make drug amphiphiles, which assembled into supramolecular structures (nanotubes) with definite structure (Cheetham et al., 2013). Serving as a part of the carrier building block, the drug molecules constituted up to 38% of the nanotubes (Cheetham et al., 2013). An important consideration in design of drug- polymer conjugates is that cleavage of the conjugate should restore a pharmacologically active drug (Stella & Nti-Addae, 2007). Although it is an efficient way of controlling drug release, worth mentioning is that drug-polymer conjugate is considered a new chemical entity, which needs a new FDA approval for clinical use (Kim et al., 2009).