ReviewNanomedicines for advanced cancer treatments: Transitioning towards responsive systems
Introduction
The development of new medicinal compounds to face current healthcare challenges, especially in oncology, is commonly hindered by their poor solubility in water and non-specific cytotoxicity resulting in adverse side effects (Crawford et al., 2004, Gale, 1985, Gharib and Burnett, 2002, Bhattacharyya et al., 2015). Chemotherapeutics impair cell mitosis and thereby target fast dividing cells like cancer cells (Hanahan and Weinberg, 2000). A high dose of these drugs needs to be administered to achieve therapeutic levels, leading to various adverse effects because also other fast dividing, healthy cells are affected. Examples of adverse effects of chemotherapeutic drugs are neuropathy, nausea, general discomfort, myelosuppression (suppression of the bone marrow's production of blood cells and platelets), alopecia (hair loss), nephrotoxicity and cardiotoxicity (Crawford et al., 2004, Gale, 1985, Gharib and Burnett, 2002). Importantly, these adverse events often limit the dose and duration of the administered drugs. Another major problem is that cancer cells can become resistant towards chemotherapeutic drugs (Baguley, 2010, Szakács et al., 2006). Formulating cytostatic drugs in liposomal carriers or polymeric micelles, as well as in other nanoparticulate drug delivery systems, has demonstrated successful results in terms of therapeutic efficacy and tolerability, giving rise to several commercial products (Wicki et al., 2015a, Pérez-Herrero and Fernández-Medarde, 2015).
In general, drug delivery systems modify the pharmacokinetics of the loaded drug enabling a more favorable distribution (Nunez-Lozano et al., 2015). Given that drug carriers have sizes in the submicrometer range and they are surface modified by hydrophilic polymers, such as polyethylene glycol (PEG), they can avoid renal clearance and rapid uptake by cells of the reticuloendothelial system (RES) (Davis et al., 2008), therefore maximizing blood circulation time. Consequently, passive accumulation mechanisms can lead to high drug concentration in certain tissues. An example of this effect, attributed to certain tumor regions, (Kobayashi et al., 2014) inflammation (Yuan et al., 2012) and sites of bacterial infection (Maeda, 2013), is the enhanced permeation and retention (EPR) (Maeda and Matsumura, 1989a, Maeda et al., 1992, Matsumura and Maeda, 1986a, Seymour, 1992), by which different intravenously injected drug carriers can extravasate and later accumulate in the tissue, which is characterized by a leaky vasculature and poor lymphatic drainage (Maeda et al., 2000a, Maeda and Matsumura, 1989b, Matsumura and Maeda, 1986b). As a result, drug delivery systems enhance the therapeutic potential of the drug while reducing systemic side effects.
Despite the fact that PEGylated drug nanocarriers have a higher probability of ending up in tumors due to their extended time in the systemic circulation, there are a number of eventualities that may reduce the chances for an enhanced drug efficacy. For instance, in the case of polymeric micelles, instability might cause disassembly and thus hinder passive accumulation by the EPR effect (Chen et al., 2008, Savic et al., 2006). In fact, drug delivery to the tumor is rarely reported to be higher than 10% of the injected dose (Park, 2010, Bae and Park, 2011a, Ferris et al., 2016).
The conjugation of targeting moieties to the surface of drug nanocarriers has been suggested as a means to improve disposition in tumor and therapeutic efficacy. The overexpression of some receptors on the outer membrane of tumor cells facilitates recognition of drug delivery systems functionalized with biomolecules such as folic acid (FA), Arg-Gly-Asp (RGD) peptide or some other specific peptides or antibodies (Ruoslahti et al., 2010, Sudimack and Lee, 2000). This strategy has been broadly validated in in vitro experiments, and has also shown a significantly higher uptake of drug delivery systems in cancer cells within tumors in vivo (Kirpotin et al., 2006). However, active drug targeting generally speaking does not improve overall tumor accumulation of systemically administered nanomedicines in vivo (Kunjachan et al., 2014a), and with the exception of antibody-drug conjugates no drug delivery system employing this approach has yet been approved for clinical use (van der Meel et al., 2013). Further, it has been reported that nanoparticles or macromolecules with sizes of a few tens of nanometers could circulate in blood for long periods (half-life up to several hours), but they hardly penetrate into the tumor interstitium due to a high cell density and interstitial fluid pressure (Popovic et al., 2016, Dreher et al., 2006). For this reason, active targeting might not lead to improvements in the outcome over passively targeted nanomedicines. Active targeting however is often required for therapeutics, such as nucleic acid based drugs, which have their action intracellularly and as such do not spontaneously pass cellular membranes. The targeting ligand facilitates internalization of the drug-loaded nanoparticles leading to intracellular action of the drug after being released from its carrier system. It is worth mentioning that some stimuli-sensitive nanosystems have been designed to enhance tumor penetration after micelle disruption (Callahan et al., 2012) or cleavage of amide bonds (Li et al., 2016a) at the low pH characteristic of tumor sites.
Another challenge in local drug delivery is associated with the fact that the retention of the drug within the carrier should be strong enough to avoid premature leakage, but on the other hand the drug should be released from the carrier once the nanomedicine has reached its target site. Even in those cases in which a drug-loaded nanocarrier is able to reach the tumor region, an insufficient release of the cargo may hinder therapeutic efficacy (O’brien et al., 2004, Lammers et al., 2012). Overall, given the various constraints affecting drug accumulation in tumors, it is not surprising that most currently approved drugs do not substantially increase clinical performance after formulating them in nanocarriers (Venditto and Szoka, 2016).
The ability to trigger drug release on demand might therefore offer a renovated hope on the use of nanocarriers. A promising alternative to achieve site- and time-controlled release of therapeutics resides in the development of stimuli-responsive drug delivery systems. Ideally, such systems would entrap the drug in such a way that no premature release occurs at undesirable places in the body. Subsequently, drug release would be triggered upon stimulation by external means (heat, ultrasound, light, magnetic field, etc.) or physiological cues (pH, redox potential, metabolite or enzyme concentration, etc.) at the site of action. ThermoDox®, a heat-activated liposomal formulation of an approved and frequently used drug for the treatment of a wide range of cancers, doxorubicin (DOX), exemplifies this concept (May and Li, 2013, Needham et al., 2013). The release of DOX from this formulation occurs upon heating to temperatures above the transition temperature of the used phospholipids (ca. 41 °C), at which the permeability of the liposome membrane is enhanced and the drug is released from the aqueous core into the outer medium. ThermoDox® together with radiofrequency ablation recently completed a phase III clinical trial in hepatocellular carcinoma patients (Needham, 2013).
Several types of stimuli-responsive materials have been proposed in the literature. In the field of polymeric drug delivery systems, sensitive hydrogels have been widely investigated to provide a triggered release of therapeutic agents. The mechanism of action of these systems is based on abrupt changes in volume in response to temperature, pH, electric field or protein concentration (Yang et al., 2015a, Griset et al., 2009, Shirakura et al., 2014). Inorganic nanocarriers have also been developed for this purpose, providing a stable structure to host the drug and the necessary anchoring groups to which a variety of responsive building blocks can be coupled (Vallet-Regi et al., 2011). Among the most intensively reported, silica-based materials, in particular mesoporous silica nanocarriers, have been chemically modified and combined with different inorganic nanoparticles (such as quantum dots and magnetic nuclei) or macromolecules to allow the construction of stimuli-responsive systems. (Li et al., 2014a, Han et al., 2015, Díez et al., 2014, Knezevic et al., 2013, Martin-Saavedra et al., 2010, Vallet-Regi and Ruiz-Hernandez, 2011)
In this article, we summarize recent developments on nanocarriers for anti-cancer drug delivery, with a particular focus on stimuli-responsive systems. Different material configurations have been attempted to enable these carriers to release the loaded drugs due to internal cues intracellularly or in the tumor microenvironment as well as external application of physical stimuli. We will discuss the use of biomaterials as carriers for drugs, the different options for triggered release and the interactions of these drug delivery systems with the biological milieu.
Section snippets
Liposomes
Nanosized drug delivery systems like liposomes have been developed for the encapsulation of chemotherapeutic drugs to improve their therapeutic efficacy and to reduce adverse events. Liposomes are unilamellar or multilamellar vesicles, which were first described in the 1960s by Bangham (Fig. 1A) (Bangham et al., 1965, Bangham and Horne, 1964). Liposomes consist mainly of phospholipids, which spontaneously form a lipid bilayer surrounding an aqueous core when dispersed in water via non covalent
Polymeric micelles
Amphiphilic block copolymers with e.g. an A-B or A-B-A architecture spontaneously form micelles with size ranging from 5 to 100 nm when these polymers are dispersed in an aqueous medium. The outer shell is formed by the hydrophilic block of the polymer while the inner core is formed by the hydrophobic part of the block copolymer. The inner core can be used to solubilize hydrophobic drugs (Kowalczuk et al., 2014) and because of this and their small size, polymeric micelles are very attractive
Mesoporous silica nanoparticles
Silica-based materials have been proposed for a variety of biomedical applications (Vallet-Regi and Ruiz-Hernandez, 2011, Alcaide et al., 2012, Kinnari et al., 2009, Serrano et al., 2008). Mesoporous silica can be synthesized as ordered pore arrays with hexagonal or cubic symmetries. SBA15- or MCM41-type structures were discovered in the early nineties of last century following a template-assisted process with the use of surfactants as structure-directing agents (Kresge et al., 1992,
Ultrasound
Ultrasound is an imaging modality which is cost effective, patient convenient and is used in the clinic for real time imaging. Furthermore, ultrasound can be used to purposefully destabilize drug delivery systems. Due to these advantageous properties and broad applicability it is possible to use ultrasound as a trigger for drug release and enhance its delivery to tumors.
Delivery of chemotherapeutic agents and therapeutic genes to the central nervous system through the blood brain barrier (BBB)
Image guided chemotherapy
Image guided drug delivery has gained substantial interest in recent years and significant progress has been made in this field. Preclinical, image guided drug delivery is used to monitor the pharmacokinetics, distribution, target site accumulation and treatment efficacy of nanoparticles. Furthermore, imaging can be used to visualize and quantify drug release mainly for triggered nanosystems. In a clinical setting, imaging can be used for preselection of patients who will likely benefit from a
Perspectives
Cancer chemotherapy is currently limited by a high systemic toxicity that occurs as a result of high systemic dosing and the failure of non-targeted drugs to reach therapeutic doses at the target site either due to barriers around the tumor (aberrant vasculature), increased efflux of molecules from cancer cells (P-gp based efflux pumps) and the development of resistance to chemotherapeutic agents through co-activation of alternative signalling pathways (Lissa Nurrul Abdullah, 2013).
Acknowledgements
We thank the following for funding this work: RCSI Seed funding Scheme, EU FP7 AMCARE, H2020 FP7 DRIVE. This publication has emanated from research supported in part by a research grant from Science Foundation Ireland (SFI) under Grant Number SFI/12/RC/2278.
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