A review of stimuli-responsive nanocarriers for drug and gene delivery

Abstract

Nanotechnology has shown tremendous promise in target-specific delivery of drugs and genes in the body. Although passive and active targeted-drug delivery has addressed a number of important issues, additional properties that can be included in nanocarrier systems to enhance the bioavailability of drugs at the disease site, and especially upon cellular internalization, are very important. A nanocarrier system incorporated with stimuli-responsive property (e.g., pH, temperature, or redox potential), for instance, would be amenable to address some of the systemic and intracellular delivery barriers. In this review, we discuss the role of stimuli-responsive nanocarrier systems for drug and gene delivery. The advancement in material science has led to design of a variety of materials, which are used for development of nanocarrier systems that can respond to biological stimuli. Temperature, pH, and hypoxia are examples of “triggers” at the diseased site that could be exploited with stimuli-responsive nanocarriers. With greater understanding of the difference between normal and pathological tissues and cells and parallel developments in material design, there is a highly promising role of stimuli-responsive nanocarriers for drug and gene delivery in the future.

Introduction

With parallel recent breakthroughs in molecular understanding of diseases and controlled manipulations of material at the nanometric length scale, nanotechnology offers tremendous promise in disease prevention, diagnosis, and therapy [1]. Among the various approaches for exploiting developments in nanotechnology for biomedical applications, nanoparticulate carriers offer some unique advantages as delivery, sensing and image enhancement agents [2]. Many bioactives used for pharmacotherapy, while have a beneficial action, can also exhibit side-effects that may limit their clinical application. There has long been the desire to achieve selective delivery of bioactives to target areas in the body in order to maximize therapeutic potential and minimize side-effects. For example, cytotoxic compounds used in cancer therapy can kill target cells, but also normal cells in the body resulting in undesired side-effects. For achieving better therapeutic application, nanocarriers are considered for target-specific delivery of drugs and gene to various sites in the body in order to improve the therapeutic efficacy, while minimizing undesirable side-effects. Improvements in target-to-non-target concentration ratios, increased drug residence at the target site, and improved cellular uptake and intracellular stability are some of the major reasons for greater emphasis on the use of nanoparticulate delivery systems. With nucleic acid-based therapeutic modalities, there is substantial need for the therapeutic molecules to be delivered to desired sub-cellular compartments in an efficient and reproducible manner [2].

Nanoparticulate carriers can be made from a variety of organic and inorganic materials including non-degradable and biodegradable polymers, lipids (liposomes, nanoemulsions, and solid-lipid nanoparticles) self-assembling amphiphilic molecules, dendrimers, metal, and inorganic semiconductor nanocrystals (quantum dots) [1], [3]. The selection of material for development of nanoparticulate carriers is mainly dictated by the desired diagnostic or therapeutic goal, type of payload, material safety profile, and the route of administration. Preponderance of literature on nanocarrier systems is based on the use of polymeric, lipid, self-assembling, and a variety of inorganic nanoparticulate carriers [1], [3], [4].

The use of stimuli-responsive nanocarriers offers an interesting opportunity for drug and gene delivery where the delivery system becomes an active participant, rather than passive vehicle, in the optimization of therapy. Several families of molecular assemblies are employed as stimuli-responsive nanocarriers for either passive or active targeting. Liposomes, polymeric nanoparticles, block copolymer micelles and dendrimers are colloidal molecular assemblies (Fig. 1). The composition of each class of these molecular assemblies can be manipulated to obtain nanocarrier of desired stimuli-responsive property. The benefit of stimuli-responsive nanocarriers is especially important when the stimuli are unique to disease pathology, allowing the nanocarrier to respond specifically to the pathological “triggers”. Select examples of biological stimuli that can be exploited for targeted-drug and gene delivery include pH, temperature, and redox microenvironment [2], [5], [6], [7], [8]. The extracellular and intracellular pH profile of biological system is greatly affected by diseases. For instance, in solid tumors, the extracellular pH tends to be significantly more acidic (∼ 6.5) than the pH of the blood (7.4) at 37 °C [9]. In addition, the pH values of endosomal and lysosomal vesicles inside the cells are also significantly lower that the cytosolic pH. By selecting the right material composition, it is possible to engineer nanocarriers that can exploit these pH differences and allow for delivery of the encapsulated payload to specifically occur in select extracellular or intracellular sites. Temperature is another variable that can be exploited in specifically releasing the nanocarrier-delivered drugs or genes to a select target site [10]. For instance, using temperature-sensitive nanocarriers one could envision a delivery system that will only release the payload at temperatures above 37 °C. Such a system would keep the toxic drug encapsulated in the systemic circulation or upon contact with non-targeted tissue. However, on application of hyperthermic stimuli to the disease area, the drug would be readily available in a localized region [10]. Lastly, intracellular glutathione (GSH) levels in tumor cells are 100–1000 fold higher than the extracellular levels [11]. This concentration gradient can be exploited using disulfide cross-linked nanocarriers that will release the payload inside the cell. Such a system is especially relevant in delivery of nucleic acid-based therapies, such as plasmid DNA, small interference RNA, or oligonucleotide, since these molecules have to reach intracellular targets in a stable form for efficient therapeutic effect.

Another possible strategy is physical targeting of drugs and genes by external stimuli (magnetic field, ultrasound, light and heat) [12], [13], [14], [15], [16]. An interesting example is targeted delivery of iron oxide nanoparticles using magnetic field. Upon the administration, the drug immobilized magnetite carrier can accumulate at targeted site under the direction of external magnetic field [14]. During the last decade, ultrasound has attracted growing attention in the targeted-drug delivery. Ultrasound has been used to achieve the targeted delivery to the tumor by local sonication after the injection of micellar encapsulated drugs [15], [16]. In addition to tumor uptake, this technique also allows the uniform distribution of micelles and drug throughout the tumor tissue [17]. Light-responsive nanocarriers have also gained recent attention. Designing of light sensitive polymeric systems that undergo reverse micellization/disruption under the action of light is an attractive idea that would allow external control of drug release [18]. Discussion of extensive current literature on the external stimuli is beyond the scope of the present review. Several examples illustrating approaches to designing external stimuli are available for further reading from the references [12], [13], [14], [15], [16].

For systemic therapy, passive and active targeting strategies are utilized. Passive targeting relies on the properties of the delivery system and the disease pathology in order to preferentially accumulate the drug at the site of interest and avoid non-specific distribution. For instance, poly(ethylene glycol) (PEG)- or poly(ethylene oxide) (PEO)-modified nanocarrier systems can preferentially accumulate in the vicinity of the tumor mass upon intravenous administration based on the hyper-permeability of the newly-formed blood vessels by a process known as enhanced permeability and retention (EPR) effect, schematically illustrated in Fig. 2. Maeda and colleagues [19], [20] first described the EPR effect in murine solid tumor models and this phenomenon has been confirmed by others. When polymer–drug conjugates are administered, 10–100 fold higher concentrations can be achieved in the tumor due to EPR effect as compared to administration of free drug [21]. The EPR effect has been also present in other diseases such as chronic inflammation and infection. Thus, the application of nanocarriers is expected to have therapeutic benefits for treating these diseases as well [22]. The tendency of nanocarriers to localize in the reticuloendothelial system also presents an opportunity for passive targeting of bioactives to the macrophages present in the liver and spleen. For example therapies can be used to treat intracellular infections such as candidiasis, leishmaniasis and listeria; where macrophages are directly involved in the disease process [23]. Other approaches for passive targeting involve the use of specific stimuli-sensitive delivery system that can release the encapsulated payload only when such a stimuli is present. For instance, the pH around tumor and other hypoxic disease tissues in the body tends to be more acidic (i.e., ∼ 5.5 to 6.5) relative to physiological pH (i.e., 7.4). Using pH-sensitive poly(beta-amino ester) (PbAE) nanoparticles, we have found significant enhancement in drug delivery and accumulation in the tumor mass as compared to drug administration in PCL nanoparticles, a non-pH-sensitive polymer, and in aqueous solution [5], [6], [7]. Further approaches for passive targeting involve size of the nanocarriers and surface charge modulation. Nanoparticles of < 200 nm in diameter and those with positive surface charge are known to preferentially accumulate and reside in the tumor mass for longer duration than either neutral or negatively charged nanoparticles [21]. Recently, we have examined the role of combination paclitaxel and the apoptotic second messenger, C₆-ceramide, when administered concurrently in PEO-modified PCL nanoparticles to overcome multidrug resistance in cancer [21], [24].

Furthermore, we have shown that passively-targeted delivery of type B gelatin-based nanoparticles has been very effective in systemic gene delivery to solid tumor [25]. Type B gelatin (pI ∼ 4.5) in nanoparticulate formulations is able to physically encapsulate plasmid DNA at neutral pH [26]. The physically encapsulated DNA in PEG-modified gelatin nanoparticles was found to be more effective in vitro and in vivo in transfection of reporter plasmid DNA expressing green fluorescent protein and beta-galactosidase [27], [28]. Upon systemic administration in C57/BL6J mice bearing Lewis lung carcinoma, the PEG-modified gelatin nanoparticle afforded long-lasting transfection (up to 96 h) upon intravenous and intratumoral delivery. Recently, we have shown that PEG-modified thiolated gelatin nanoparticles could also encapsulate DNA and transfect tumor cells in response to higher intracellular glutathione levels [8], [29]. When PEG-modified thiolated gelatin nanoparticles encapsulated with plasmid DNA encoding for soluble vascular endothelial growth factor receptor 1 (sVEGFR-1 or sFlt-1), highly efficient transgene expression was observed in human breast cancer cells and in vivo in an orthotopic tumor model. In addition, the expressed sFlt-1 was very effective in suppressing tumor growth and angiogenesis [29].

Active targeting to the disease site relies, in addition to PEG modification of nanocarriers to enhance circulation time and achieve passive targeting, coupling of a specific ligand on the surface that will be recognized by the cells present at the disease site [30]. Using solid tumor as an example again, there are several strategies that can be adopted for surface modification of nanocarrier systems for effective targeted delivery to the tumor cells or to endothelial cells of the tumor blood vessels. Since tumor cells are rapidly proliferating, they over-express certain receptors for enhanced uptake of nutrients, including folic acid, vitamins, and sugars. When the surface of nanocarriers is modified with folic acid, they can be targeted to the tumor cells that over-express folate receptors (Fig. 3). In addition, Fig. 3 also illustrates the intracellular delivery of folate anchored nanocarrier through endocytosis process and releasing its contents in response to internal stimuli. Tumor and capillary endothelial cells also express specific integrin receptors, such as α_vβ₅ or α_vβ₃ that can bind to arginine–glycine–aspartic acid (RGD) tripeptide sequence. RGD-modification, therefore, has been used to direct nanocarriers to tumor cells and capillary endothelial cells of the angiogenic blood vessels. The phage display method has been used to identify specific peptide sequences that can be used for targeting tumors and other disease areas in the body [31]. For example, Schluesener et al. [32] used in vivo phase display of recombinant M13 phages as a tool to select peptides targeting pathological endothelium of experimental rat brain tumors. One of the FDA approved targeted therapeutics is Adalimumab® antibody; a human anti-TNF IgG1 used against rheumatoid arthritis is generated by phage display technique [33]. Recently, Farokhzad et al. [34] have elegantly described the use of aptamers, nucleic acid constructs that specifically recognize prostate membrane antigen on prostate cancer cells. The aptamer technology provides an additional strategy for active targeting of tumor cells in the body. Development of monoclonal antibodies against specific epitopes present only on tumor cells allows for other targeting strategies. For instance, HER2 specific antibody (Trastuzumab® or Herceptin®) modified nanoparticles were able to localize and deliver the therapeutic payload specifically in HER2 expressing tumor cells [35]. Using a monoclonal antibody 2C5 that specifically recognizes anti-nuclear histones, Torchilin's group has developed various strategies for active targeted delivery of drugs to the tumor mass using liposomes and micellar delivery systems [36], [37]. Other groups have used transferrin, an iron-binding protein, for surface modification of nanocarriers for delivery to tumors [38]. Epidermal growth factor receptors are over-expressed in breast or prostate cancers, making it a good candidate for targeting gene-delivery complexes [39]. Apart from tumor, epithelial surfaces of the lungs and gastrointestinal tract, endothelial cells lining the blood vessels, muscle myoblasts and skin fibroblasts are also potential targets for gene delivery [40]. Along with targeting ability, the nanocarriers also can be tailored of stimuli-responsive property to enhance the transfection ability of the carriers. For example, Oishi et al. [41] showed that pH-responsive and PEGylated nanogels bearing a lactose group at the PEG end display endosomolytic abilities and achieving the enhanced transfection efficiency.

Once the nanocarriers are delivered to the specific diseased organ or tissue, they may need to enter the cells of interest and ferry the payload to sub-cellular organelles (Fig. 4). In this case, non-specific or specific cell penetrating strategies need to be adopted [42]. Non-specific cell uptake of nanocarriers occurs by endocytotic process, where the membrane envelops the nanocarriers to form a vesicle in the cell called an endosome. The endosome then shuttles the content in the cell can fuse with lyososomes, which are highly acidic organelles rich in degrading enzymes. Endocytosed nanocarriers usually travel in a specific direction and converge at the nuclear membrane. Endosomal acidic condition is deterrent to therapeutic molecules present in the nanocarrier. This bottleneck in gene delivery can be responsible for the degradation of > 99% of the internalized DNA. Efficient gene delivery is achieved by buffering the endosomes for safe release of its contents. For example, the buffering capacity of the polycationic carriers can hamper the acidification of the endosomes, causing it to swell and burst, as a consequence safe release of trapped contents [43], [44]. Specific cellular uptake can occur through receptor-mediated endocytosis, where upon binding of the ligand-modified nanocarrier with the cell-surface receptor leads to internalization of the entire nanocarrier–receptor complex and vesicular transport through the endosomes. Following dissociation of the nanocarrier–receptor complex, the receptor can be re-cycled back to the cell membrane. Recently, in order to enhance cellular uptake, a surge of research effort has been directed towards development of arginine-rich cell penetrating peptides (CPP's) [45]. Based on the pioneering work of Dowdy's group HIV-1 Tat peptide was identified to promote non-specific intracellular localization of various molecules upon systemic delivery [46]. This observation has been supported by other groups and a number of cationic peptides have been identified, including penetratin, to enhance intracellular delivery. The exact mechanism of how Tat and other CPP's enhance cell permeation is still a subject of controversy, but recent data show that it may be through endocytosis as well [47]. Following cellular internalization, stability of the payload in the cytosol and uptake by specific organelle, such as the nucleus, is also essential for nucleic acid-based therapeutics. Weissig's group has attempted to direct various nano-sized delivery systems to mitochondria using delocalized cationic amphiphiles and other mitochondriotropic vector systems [48]. For efficient systemic gene therapy using non-viral vectors, nuclear import of plasmid DNA in non-dividing cells is considered to be the major limiting factor.