Classification of stimuli-responsive polymers as anticancer drug delivery systems.

http://informahealthcare.com/drd ISSN: 1071-7544 (print), 1521-0464 (electronic) Drug Deliv, Early Online: 1–11 ! 2014 Informa Healthcare USA, Inc. DOI: 10.3109/10717544.2014.887157

CRITICAL REVIEW

Classification of stimuli–responsive polymers as anticancer drug delivery systems Bita Taghizadeh1, Shahrouz Taranejoo2, Seyed Ali Monemian3, Zoha Salehi Moghaddam4, Karim Daliri5, Hossein Derakhshankhah6, and Zaynab Derakhshani7 Drug Delivery Downloaded from informahealthcare.com by Technische Universiteit Eindhoven on 05/25/14 For personal use only.

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Institute of Biochemistry and Biophysics, University of Tehran, Tehran, Iran, 2Chemical Engineering Department, Faculty of Engineering, Monash University, Melbourne, Australia, 3Department of Macromolecular Science and Engineering, Case Western Reserve University, Case Western Reserve Cleveland, OH, USA, 4Department Material Science and Engineering, University of Sheffield, Sheffield, UK, 5Department of Medical Genetics, Shiraz University of Medical Sciences, Shiraz, Iran, 6Department of Pharmaceutical Biomaterials and Nanotechnology, Faculty of Pharmacy, Tehran University of Medical Sciences, Tehran, Iran, and 7Medical Nanotechnology and Tissue Engineering Research Center, Shahid Beheshti University of Medical Sciences, Tehran, Iran Abstract

Keywords

Although several anticancer drugs have been introduced as chemotherapeutic agents, the effective treatment of cancer remains a challenge. Major limitations in the application of anticancer drugs include their nonspecificity, wide biodistribution, short half-life, low concentration in tumor tissue and systemic toxicity. Drug delivery to the tumor site has become feasible in recent years, and recent advances in the development of new drug delivery systems for controlled drug release in tumor tissues with reduced side effects show great promise. In this field, the use of biodegradable polymers as drug carriers has attracted the most attention. However, drug release is still difficult to control even when a polymeric drug carrier is used. The design of pharmaceutical polymers that respond to external stimuli (known as stimuli–responsive polymers) such as temperature, pH, electric or magnetic field, enzymes, ultrasound waves, etc. appears to be a successful approach. In these systems, drug release is triggered by different stimuli. The purpose of this review is to summarize different types of polymeric drug carriers and stimuli, in addition to the combination use of stimuli in order to achieve a better controlled drug release, and it discusses their potential strengths and applications. A survey of the recent literature on various stimuli–responsive drug delivery systems is also provided and perspectives on possible future developments in controlled drug release at tumor site have been discussed.

Anticancer, cancer stimuli–responsive polymers, drug carriers, smart delivery systems, pH-sensitive polymers

Introduction Cancer is a leading cause of death worldwide, and despite the availability of several treatment modalities, cancer therapy remains a great challenge. Chemotherapy is the main treatment modality for cancer patients, and this has led to the introduction of a large number of chemotherapeutic anticancer drugs. However, the major limitation to the clinical application of these drugs is their short half-life and wide biodistribution (Bristow et al., 1978; Praga et al., 1980; Sparano, 1999). The major drawback of conventional drug delivery approaches is their low selectivity; as a result, normal cells are also exposed to the cytotoxic effects of these drugs (Fenton & Longo, 1998). In fact, in most cases, only a small portion of the administered drug reaches the tumor site (Muller & Keck, 2004). This limitation is further compounded by the need to

Address for correspondence: Shahrouz Taranejoo, Chemical Engineering Department, Faculty of Engineering, Monash University, Melbourne, Australia. E-mail: [email protected]

History Received 11 December 2013 Revised 18 January 2014 Accepted 19 January 2014

administer frequent, high doses of these drugs, which adds to their undesired side effects. It is therefore essential to develop new, targeted drug delivery systems to overcome these limitations and to improve the efficacy of cancer treatments. These new systems should deliver the drug to the tumor site and reduce their side effects (Kataoka et al., 2005). Different types of drug delivery vehicles have been developed for different therapeutic applications, including micelles, liposomes, dendrimers, nanoparticles (NPs) and lipoprotein drug carriers, among others. Polymeric carriers are one of the most commonly used approaches (Alamdarnejad et al., 2013; Taranejoo et al. 2014). Soluble polymeric drug carriers can improve drug pharmacokinetics and reduce systemic toxicity by delivering the drug to the tumor tissue, thanks to the higher permeability of tumor vasculature, known as the enhanced permeability and retention (EPR) effect (Matsumura & Maeda, 1986). Tumor tissue exhibits unique characteristics such as EPR, which makes it different from normal tissue. The EPR effect causes tumor tissues to display features such as hypervascularity, larger vascular pore sizes (400–800 nm) compared to normal tissue

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B. Taghizadeh et al.

Drug Deliv, Early Online: 1–11

(5–10 nm), and the absence of a well-defined functional lymphatic drainage system. Together, these characteristics lead to the accumulation of polymer–drug conjugates at the tumor site (Maeda et al., 2000; Maeda et al., 2001; Kopecek et al., 2001; Abulateefeh et al., 2011). Even with polymeric drug carrier systems, it is difficult to control drug release. One solution is to design pharmaceutical polymers that undergo physiochemical changes in response to environmental stimuli such as temperature, pH, electric or magnetic field, enzymes, solvent polarity, etc. (Hoffman, 2004; Duncan, 2006; MacEwan et al., 2010; Aliaghaie et al., 2012). An ideal strategy for targeted drug delivery to the tumor site would be to enclose the drug within a container that remains stable before the drug is released. In addition, the drug should be released at the most appropriate time and dose in a controlled manner that matches the physiological properties of the tumor tissue.

Thermosensitive polymers Thermosensitive polymers are the most widely studied stimuli–responsive polymers, since temperature can easily be used to trigger and control drug release at tumor sites (Gil & Hudson, 2004). The idea of using temperature as a stimulus arose partly from the fact that tumor microenvironments are slightly hyperthermic (i.e. 1–2 C warmer than normal healthy tissue) (Vaupel et al., 1989). Hyperthermia has been used clinically for decades as an adjuvant therapy (combined with radiotherapy or chemotherapy) in the management and treatment of solid tumors (Field & Bleehen, 1979; Law, 1982; Urano et al., 1999; Coffey et al., 2006). Cancer cell death as a result of increased temperature ¨ ber den at the tumor site was first reported in 1866 (U Einfluss, 1866). Hyperthermia preferentially increases the permeability of the tumor tissue vasculature and thus enhances anticancer drug delivery to the tumor site (Issels, 1995; Engin, 1996). These facts suggest that thermal targeting of cancer cells by polymeric drug carriers may have synergistic effects compared to individual treatment modalities. As a drug releasing trigger, heat is applied by the use of temperature-controlled water sacks, radiofrequency oscillators or miniature annular-phased array microwave applicators. Applying hyperthermia using iron oxide nanoparticles (IONPs) via an alternating electromagnetic field has been widely studied. Wang et al. studied a combination of IONPs and Pluronic F-127 (PF 127), a thermosensitive polymer with concentration-dependent gelation properties, for cancer treatment, suggesting that combination treatment would allow short-term, tumorspecific and hyperthermic effects (Kost & Langer, 2001; Wang et al., 2013). A recent work by Meenach et al. investigated the properties of PEG-iron oxide hydrogels for combined hyperthermia and paclitaxel delivery to tumor cell lines in vitro. M059K (glioblastoma), MDA-MB-231 (breast carcinoma) and A549 (lung adenocarcinoma) were exposed to paclitaxel only, hyperthermia only and both paclitaxel and hyperthermia to determine whether a synergistic cytotoxic effect would occur. A cytotoxic synergism was observed for the A549

Figure 1. Schematic illustration of reversible thermal transitions and cellular uptake of micellar-like nanoparticles (Soga et al., 2004).

cells; however, the M059K and MDA MB 231 did not show the same response (Chaterji et al., 2007; Meenach et al., 2013). Pradhan et al. (2010) proved that triggered drug release and hyperthermia induced by magnetic field can be advantageous for thermo-chemotherapy of cancers. Their results demonstrated that folate receptor targeted thermosensitive magnetic doxorubicin-loaded liposome’s (MagFolDox) cytotoxicity toward cancer cells is synergistically increased by magnetic hyperthermia at 42.5 C and 43.5 C (Pradhan et al., 2010). Figure 1 illustrates reversible thermal transitions and cellular uptake of micellar-like NPs. Thermosensitive polymers have a transition temperature at which they become soluble or insoluble. The lower critical solution temperature (LCST) is the temperature at which transition from a more soluble to a less soluble state occurs. Inversely, the temperature in which transition from a less soluble to a more soluble state occurs is called upper critical solution temperature (Kost & Langer, 2001; Abulateefeh et al., 2011). Polymeric drug carriers that undergo LCST phase transition have been studied. They can be injected in the soluble phase and become insoluble and accumulate at the tumor site upon applying local hyperthermia. Tuning the nature and the composition of the carriers can be advantageous for local administration of thermosensitive drug-releasing systems. Thermosensitive delivery systems should have an LCST above body temperature, since hyperthermal treatment in patients with cancer is usually performed at 42 C (Meyer et al., 2001; Chaterji et al., 2007). The materials used in the design of thermosensitive drug delivery systems have to be both safe and sensitive enough to respond to slight temperature changes (Mura et al., 2013). Li et al. synthesized anti-Her2 antibody fab fragment-conjugated immunomicelles (FCIMs) based on


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thermosensitive poly(N-isopropylacrylamide-co-N,N-dimethy lacrylamide)118 (PID118) copolymers, and used this system to deliver doxorubicin (Dox). Their approach doubled the Dox loading capacity and enhanced cytotoxicity due to the combined effects of increased temperature and the fab antibody moiety attached to the surface of the micelles. Furthermore, serum stability and intratumoral accumulation were significantly improved (Li et al., 2012). Rejinold et al. (2011) designed a highly stable curcuminloaded thermosensitive NP drug delivery system based on chitosan-g-poly(N-vinylcaprolactam) (TRC-Nps), which showed preferential cytotoxicity toward cancerous cells and seemed to be more beneficial when combined with hyperthermia. Gong et al. (2009a) introduced a thermosensitive polyethylene glycol-poly("-caprolactone)-polyethylene glycol (PEG-PCL-PEG, PECE) hydrogel, which remains in the soluble phase at low temperatures (10 C) but forms a gel at body temperature upon injection. Thermosensitive PECE hydrogel has been used for the controlled delivery of anticancer drugs such as bFGF, honokiol and 5-Fu (Gong et al., 2009b, 2010; Fang et al., 2009; Wang et al., 2010). Recently, the histone deacetylase agent suberoylanilide hydroxamic acid and cisplatin were delivered to an oral squamous cell carcinoma tumor (in a model system) with PECE hydrogels; the results showed synergistic cytotoxic effects that were evident for more than 14 days after injection, and suggested that this novel chemotherapy protocol may have promise in cancer treatment (Zheng et al., 2012). Thermosensitive polyethylene glycol-b-poly[N-(2-hydroxypropyl) methacrylamide-lactate] (mPEG-b-p(HPMAmLacn)) diblock copolymers have also been extensively studied as a delivery system for different anticancer drugs including PTX and the photosensitizer Si(sol)2Pc. This thermosensitive copolymer consists of a hydrophilic (PEG) and a thermosensitive (p(HPMAm-Lacn)) block which self-assembles into micellar structures in aqueous solution when heated above its critical micelle temperature (Soga et al., 2004, 2005; Rijcken et al., 2005, 2007; Van Nostrum et al., 2006; Talelli et al., 2010a).

pH-sensitive polymers The lower extracellular pH of the tumor microenvironment distinguishes it from normal tissue, and the difference is attributed to the higher rate of aerobic and anaerobic glycolysis in cancer cells. This gradient has been used to design pH-responsive tumor-targeted drug release systems based on pH-sensitive polymers (Vaupel et al., 1989; Stubbs et al., 2000; Lee et al., 2008; Fleige et al., 2012). The pH values measured in most solid tumors range from 5.7 to 7.8, with a mean of 7.0 (Ojugo et al., 1999). The essential part of a pH-sensitive drug delivery system is a pH-triggering group (an ionizable weak acidic or basic moiety) that is attached to the backbone. pH-sensitive polymers are polyelectrolytes whose coiled chains expand upon ionization, in response to changes in environmental pH, and their hydrodynamic volume increases dramatically due to electrostatic repulsion from the charges generated (anions or cations) along the backbone (Figure 2)

Classification of stimuli–responsive polymers

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(Qiu & Park, 2001; Majedi et al., 2013). Two main strategies exist in the design of pH-sensitive drug delivery systems: the use of polymeric systems with ionizable groups attached to the backbone that their conformational or solubility properties changes in response to changes in pH values and the use of polymers bearing acid-sensitive bonds whose cleavage triggers the release of drug molecules attached to the polymer backbone, modification of the polymer charge or the exposure of targeting ligands (n ¼ 4 independent experiments) (Majedi et al., 2013). When pH-sensitive polymers are cross-linked to form hydrogels, their behavior is influenced by the ionizable groups, polymer composition, hydrophobicity of the polymer backbone and cross-linking density. Preferred pH-responsive systems for drug release at tumor site must give a sharp response to the pH change in the tumor extracellular microenvironment. Another pH-responsive strategy utilizes the buffering capacity of the cationic polymer polyethylenimine (PEI). PEI’s buffering capacity is a result of its different types of amine groups with different pKa values, which can become protonated at different levels at the given pH (Zhu & Mahato, 2010). Following endocytosis, the protonation of amine groups increases the influx of protons and chloride ions as well as water into the endosome resulting in swelling and rupture of the endosome which results in the release of endosomal contents (Boussif et al., 1995, Sonawane et al., 2003). PEI’s buffering capacity as well as the positive charge has made it a universal vector for both in vitro and in vivo delivery of nucleic acids (Boussif et al., 1995; Urban-Klein et al., 2005). The pH-responsive swelling has been observed in various natural polymers as well, including albumin and gelatin. These linear polymers could form helices that function as cross-links that hold the amorphous regions together. These natural proteins have a minimal surface charge at their isoelectric point (pI) and show extensive swelling at a pH away from their pI as a result of the generation of a high surface net charge and subsequent increase in the electrostatic repulsive forces (Welz & Ofner, 1992; Park et al., 1998). The significantly lower pH in subcellular compartments (including the endosomes) can also be used as a route for delivering anticancer drugs via their pH-initiated release from endocytosed drug carriers (Kopansky et al., 2011). The pH decreases from 7.4 to about 5 or 6 in endosomes and about 4 or 5 in lysosomes; in these environments, an acid-sensitive linker could provide a tool for intracellular drug release. Acid-responsive polymers have provided enhanced endosomal delivery of drugs. Two types of acid-sensitive linkers are commonly used for this purpose: hydrazone and cis-aconityl linkers. Both are relatively stable at physiological pH and can release the bound drugs under low pH conditions. Some examples of polymer–drug conjugates containing hydrazone linkages are HPMA-Dox, PEG-Dox (Lu et al., 2011; Dong et al., 2013) PEG-epirubicin (Takahashi et al., 2013) and PEG-PTXL (Alani et al., 2010). Zhu et al. recently reported their in vivo study of acidsensitive (hydrazone linker) and acid-insensitive PEG-based lysosomal delivery of a prodrug of the nucleoside analogue gemcitabine (GemC18). They showed that GemC18 in the




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Figure 2. Controlled release of PTX from microfl uidically synthesized nanoparticles (FL ¼ 0.031) after media pH change, at 37 C (mean ± SD, n ¼ 4 independent experiments).

acid-sensitive micelles was more toxic toward cancer cells compared to acid-insensitive micelles, as a result of the acidsensitive drug release in lysosomes (Zhu et al., 2012a). pH-sensitive N-(2-hydroxypropyl) methacrylamide-Dox (HPMA-Dox) conjugates bearing an acid-responsive hydrazone linker in their structure have also been widely studied as anticancer drug delivery systems. They have led to a significant increase in therapeutic efficacy in different in vitro and in vivo cancer models. The hydrazone linker is

stable at pH 7.4 and becomes rapidly cleaved under low pH conditions (which occur in endosomes and tumor tissue). Through the hydrazone linker, the drug is released in the acidic tumor microenvironment or in the acidic organelles after cellular uptake by endocytosis (Etrych et al., 2001; Vetvicka et al., 2009; Sirova et al., 2010; Talelli et al., 2010b). The half-life of the cis-aconityl linker is about 3 h at pH 4.0. HPMA-Dox (Ulbrich et al., 2003), polyamidoamine-Dox (Lavignac et al., 2009) and polyamidoamine (PAMAM)-Dox

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(Zhu et al., 2010) are examples of polymers that use this linker in the design of anticancer drug delivery systems. In a recent study, chitosan (CS) and stearic acid copolymers conjugated with PEG via an acid-sensitive cis-aconityl linkage to create acid-sensitive PEG-CS micelles loaded with Dox were compared to acid-insensitive micelles which conjugated PEG directly to CS (PCS). The PEG-CS micelles had a greater Dox loading capacity, less cytotoxicity toward normal cells, enhanced cellular uptake and better accumulation in tumor tissue compared to acid-insensitive PCS micelles (Hu et al., 2012).


Enzyme-sensitive polymers Most invasive diseases including cancer are characterized by the upregulation and overexpression of different enzymes that are either secreted or membrane bound, including proteases, matrix metalloproteinases and cathepsins (Moin et al., 1998; Park et al., 1999; Utani et al., 2003; Deryugina & Quigley, 2006). The use of enzymes as biological stimuli to trigger selective drug release offers new possibilities in the design of drug delivery systems for cancer therapy. Most of enzymeresponsive drug delivery systems utilize the presence of enzymes in the extracellular environment. A peptide linker mimicking the enzyme substrate situated between the drug and the polymeric carrier can provide a useful drug release mechanism. The concept of utilizing enzymes that are consistently overexpressed in cancer tissues to activate prodrugs has been exploited since the early 1970s. Meanwhile, more recent work with prodrugs activated by tumor-secreted enzymes or transported to and released in tumor cells and then activated by cellular enzymes has been promising. For instance, colonic drug delivery is potentially useful to treat colon cancer by exploiting the colon microflora (Yang et al., 2002). Microbial populations in the colon may express reductive or hydrolytic enzymes and are able to degrade various types of polysaccharides including pectin, CS, cyclodextrin and dextrin. Because the microbial enzyme dextranase can degrade the polysaccharide dextran, it has been exploited in formulations of drug delivery systems for the colon (Hovgaard & Brøndsted, 1995; Jain et al., 2007). Two commonly used approaches for the enzyme-specific release of conjugated drugs at the tumor site are antibodydirected enzyme prodrug therapy and polymer-directed enzyme prodrug therapy (Baghshawe & Begent, 1996, Grifith & Hansen, 2002). The latter is a two-step approach: injection of the polymer–drug conjugate is followed by the administration of a polymer–enzyme conjugate that accumulates in tumor tissue because of the EPR effect. The linker in the polymer–drug conjugate is designed to be cleaved by the foreign enzyme. A limitation of this approach is that the introduction of a foreign enzyme will induce immunogenicity (Duncan et al., 2001; Duncan & Satchi-Fainaro, 2002). Many anticancer drugs have been used in the design of enzyme-sensitive polymeric delivery vehicles, including cathepsin B-sensitive PEG-PTX conjugates (Liang et al., 2012) and dextran–methotrexate conjugates sensitive to matrix metalloproteinase (Chau et al., 2006). Matrix metalloproteinases (MMPs) are among the important enzymes that have been considered as biomarkers for diagnostics and


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prognostics due to the changes in their composition and expression levels in different types of cancers. MMPs have also provided opportunities for enzyme-mediated drug delivery to tumor site (Meers, 2001; Mansour et al., 2003; Gialeli et al., 2011; Zhu et al., 2012b). In a study by Terada et al. (2006), an MMP2-cleavable lipopolymer was incorporated into targeted liposomes for the delivery of N4-octadecyl-1-b-D-arabinofuranosylcytosine to tumor cells. A dextran–peptide–methotrexate conjugate was also designed to achieve tumor-targeted delivery of the anticancer drug. The peptide linker was optimized to allow drug release in the presence of MMP-2 and -9. A significant enhancement of tumor reduction was observed in MMP expressing tumor models compared to no tumor growth inhibition by the conjugate in non-MMP overexpressing tumor models (Chau et al., 2006). Recently, a novel smart and multifunctional liposomal drug delivery system of polyethylene glycol (PEG)-lipid conjugates was developed based on MMP-2 overexpression in tumor cells. The functionalized liposome was further modified with the tumor cell-specific antinucleosome monoclonal antibody (mAb 2C5) to enhance targeted delivery, while an MMP 2-sensitive bond was placed between PEG and lipid and a cell-penetrating peptide (TATp) was also incorporated to the liposome to enhance the intracellular delivery of the system. Their study showed that such a design can enhance the targetability and internalization of nanocarriers in cancer cells (Zhu et al., 2012b). Some researchers have developed enzyme-sensitive polymeric micelles for the codelivery of PTX and DDP. This multidrug vehicle is based on PEG, glutamic acid and phenylalanine (PEG–PGlu–PPhe) triblock copolymers, and is degraded in the presence of cathepsin B. An attractive characteristic of this multidrug polymer-based micellar delivery system is that it offers the benefits of synergistic cytotoxicity and superior antitumor activity compared to individual drug-loaded micelles or free drugs and has thus shown promise for combinatorial drug delivery to cancer cells (Desale et al., 2013). Enzyme-sensitive delivery systems have also been designed for the site-specific delivery of photosensitizer prodrugs to tumor tissue (Gabriel et al., 2011). Recently, the photosensitizer agent pheophorbide was attached to a polyL -lysine backbone via urokinase-like plasminogen activator (uPA) peptide substrates. The protease uPA is overexpressed in prostate cancer. Polymeric photosensitizer prodrugs (PPPs) designed by Zuluaga et al. are nontoxic in their native state but become toxic after activation by uPA. The selective accumulation and activation of PPPs was observed in prostate tumor xenografts; this finding shows that a more selective treatment approach may be possible for localized prostate cancer (Zuluaga et al., 2012). In an interesting study in 2013, Zhu et al. introduced a unique tumor-targeted micellar enzyme-sensitive drugdelivery system using a nanocarrier composed of an MMP 2-sensitive self-assembly PEG 2000-paclitaxel conjugate, transactivating transcriptional activator peptidePEG1000-phosphoethanolamine (PE) (a cell-penetrating enhancer), and PEG1000-PE (a nanocarrier building block) with potential for effective intracellular delivery of drug into

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cancer cells, superior cytotoxicity, tumor targeting and antitumor efficacy in vitro and in vivo compared to its nonsensitive counterpart, free paclitaxel and conventional micelles (Zhu et al., 2013). Despite the large number of representative examples of enzyme-dependent drug delivery systems, work is still needed to determine the exact levels of the target enzyme at tumor site in order to fine-tune the drug release systems, control cell uptake and to make sure that drug release at tumor site is correlated to target enzyme activity.


Ultrasound-sensitive polymers Although ultrasound is considered a therapeutic tool to induce hyperthermia for tumors that are not located near critical healthy tissues (Kennedy et al., 2003), drug-delivery approaches involving ultrasound energy also benefit from the nonthermal effects of ultrasonic waves (Saad & Hahn, 1989). Sonication is considered a noninvasive technique that can penetrate deeply into the body tissues. The main thermal and physical actions of ultrasound are a local increase in temperature, perturbation of the cell membrane, and increased permeability of the blood vasculature (Gao et al., 2005). Ultrasound-dependent drug release can be achieved through the thermal and/or mechanical effects generated by cavitation phenomena or radiation forces. It has been shown that physical forces associated with cavitation can cause nanocarrier destabilization and subsequent drug release (Schroeder et al., 2009) and also that a transient increase in vessels permeability by the application of therapeutic ultrasound lead to twofold increase in tumor accumulation of anticancer drug molecules (Kheirolomoom et al., 2010). Ultrasound triggered cavitation can be easily achieved in low frequencies. However, enhancement of vessel permeability by ultrasound can result in possible problems in the case of metastatic dissemination (Garcıá-Romań & Zentella-Dehesa, 2013). Combination of thermo-responsive carriers and highintensity focused ultrasound (HIFU) enables triggered drug release with an increase in temperature. For instance, Ranjan et al. showed that by using the magnetic resonance (MR)guided HIFU, mild hyperthermia (40–41 C) was applied to the tumors and thus triggered the release of doxorubicin from doxorubicin encapsulated low temperature sensitive liposomes (LTSLs); which resulted in higher drug accumulation in an experimental tumor model compared with nonirradiated controls (Ranjan et al., 2012). Although drug delivery triggered by nonthermal ultrasound has been studied since the 1950s (Kassan et al., 1996), ultrasound-triggered drug release in tumors was investigated by Rapoport et al. in more recent years. Their drug delivery technique was based on the release of the encapsulated drug from polymeric micelles following the application of ultrasound focused on the tumor (Rapoport et al., 2003). The activation of photosensitive chemicals by ultrasound is termed sonodynamic therapy. Certain light-sensitive, nontoxic drugs can create reactive oxygen species and kill cancer cells following exposure to ultrasonic waves (Yumita et al., 1994; Tachibana et al., 1997; Tachibana et al., 1999). Numerous carriers for ultrasonically activated drug delivery have been


explored, including microtubules, micelles and nanosized particles (Fong et al., 2008; Villa et al., 2013; Yin et al., 2013). Microtubules (used as an ultrasound contrast agent) are spherical tubes ranging in diameter from 1 to 10 mm that are filled with gas and coated with polymers or phospholipids. After exposure to ultrasound, the microtubules break apart and can release the drugs enclosed within them or attached to their surface. Site-specific actions of microtubules can also be achieved by attaching specific ligands (antibodies or peptides) to their surface (Lanza et al., 1996; Dayton et al., 2004). When the ultrasound wave interacts with an attached microtubule, an extremely high-velocity jet develops (Bourne & Field, 1999), which travels to the microtubule end (toward the attached cell). The impacting shockwave creates intense stress, which in turn causes the cell membrane to deform or even rupture, affecting membrane integrity and permeability. It has been shown that ultrasound application facilitates the uptake of drug-loaded polymeric micelles by cancer cells, allowing greater retention (Wan et al., 2012). Gao et al. demonstrated that ultrasonic irradiation of the tumor triggered the release of an encapsulated drug from polymeric micelles and altered cell membrane permeability, resulting in greater drug uptake by cancer cells. Furthermore, ultrasound application enhanced drug diffusion through the tumor tissue and reduced drug concentration gradients (Gao et al., 2004, 2005). Some kinds of ultrasound responsive polymers used for cancer therapy are multisensitive polymeric drug or gene carriers. For example, Chen et al. reported a novel polymer vesicle sensitive to both physical (ultrasound) and chemical (pH) stimuli based for controlled drug release based PEO43b-P(DEA33-stat-TMA47) block copolymer vesicles (Figure 3) where PEO is short for poly(ethylene oxide), DEA for 2-(diethylamino)ethyl methacrylate and TMA for (2-tetrahydrofuranyloxy)ethyl methacrylate (Chen & Du, 2013). All energy-based tumor treatments, including ultrasoundmediated drug delivery, require primary tumor imaging. Combining techniques for imaging and treatment is always attractive. A number of combined imaging–therapeutic approaches have been introduced and shown to be both time and cost efficient (Tachibana et al., 2004; Rapoport et al., 2007; Stride & Coussios, 2010).

Magnetic-responsive polymers Early studies of controlled drug release from polymer matrices by using magnetic force were reported by Edelman et al. (Edelman et al., 1987); however, their approach was not applicable in practice. More recently, advances in nanotechnology have led to application of magnetic NPs as drug delivery vehicles. Main magnetic responses that might occur following the application of magnetic fields can be a magnetic guidance under a permanent magnetic field, a temperature increase when an alternating magnetic field is applied or both when alternately used. Furthermore, it is possible to perform MR imaging and associate diagnostics and therapy within a single system (Yang et al., 2011). NPs used in drug delivery systems must present three basic characteristics: their surface must be biocompatible and


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Figure 3. Formation of ultrasound and pH dually responsive PEO43-b-P(DEA33-stat-TMA47) vesicle and controlled drug release triggered by ultrasound radiation or decreasing pH value (Chen & Du, 2013).

nontoxic, and the particles must be well-dispersed without forming aggregates (Liz-Marzań et al., 1999). The size, charge and surface chemistry of NPs strongly affect their biodistribution. Magnetic NPs must be magnetized either by incorporation of ferrimagnetic (or magnetite) or by direct modification of the magnetized metal particles with biocompatible polymers. Ferrimagnetic particles have been widely studied in a wide range of biomedical applications including drug delivery and magnetic hyperthermia (Torchilin et al., 2009). Lower magnetic field intensity of magnetite compared to ferromagnetic materials makes it more suitable for clinical applications due to reduced potential health risks caused by magnetic fields and a lower chance of agglomeration due to magnetostatic interactions during preparation and storage of magnetic NPs. The ability of magnetic NPs to generate heat and serve as a transducer upon applying magnetic field, has made them been extensively used for the selective hyperthermia of tumors. Today, the main objective is to optimize magnetic particles to provide increased magnetic NP concentrations, reduce early clearance from the body, minimize nonspecific cell interactions and increase their uptake by target cells (DouziechEyrolles et al., 2007; Mahmoudi et al., 2011). In this connection, the combination of magnetic NPs with polymers has attracted much interest. The polymer shell can interact with the environment and act as a compatibilizer, while the magnetic properties of the NPs allow magnetically triggered drug release (Schmidt, 2007; Liu et al., 2008). IONPs encapsulated in a polymeric shell have been extensively studied as magnetic-responsive drug delivery systems for cancer therapy (Akbarzadeh et al., 2012; Glover et al., 2013; Filippousi et al., 2013; Oh et al., 2013; Qu et al., 2013). Superparamagnetic IONPs (SPIONs) coated with polymers have also yielded powerful vehicles for targeted delivery (Wahajuddin & Arora, 2012). SPIONs are small

synthetic g-Fe2O3 (maghemite), Fe3O4 (magnetite) or a-Fe2O3 (hermatite) particles with a core size ranging from 10 to 100 nm. Mixed oxides of iron with transition metal ions such as copper, cobalt, nickel and manganese also exhibit superparamagnetic properties and fall into the category of SPIONs. Drug-loaded SPION–polymer conjugates can be guided by an external magnet to the target tissue, where the SPION part of the system exhibits the phenomenon of superparamagnetism, i.e. when an external magnetic field is applied, they become magnetized up to their saturation magnetization, and once the magnetic field is removed, they no longer exhibit any residual magnetic interaction and are dispersed easily. This superparamagnetism characteristic is unique to NPs (Schmidt, 2007). SPIONs have been used for both therapeutic and imaging purposes in cancer therapy (Sengupta & Sasisekharan, 2007; Ruoslahti et al., 2010). One of the main limitations of using stimuli–responsive gels is the fact that structural changes in their shape and volume (swelling or deswelling) in response to external stimuli occur relatively slowly. In order to accelerate the response to stimuli, magnetic field-sensitive gels (ferrogels) have been developed (Szabo´ et al., 2000). In recent work, Chiang et al. introduced a hybrid nanogel prepared by co-assembly of citric acid-coated SPIONs with a graft copolymer comprising acrylic acid and 2-methacryloylethyl acrylate units as the backbone, and PEG and pNIPAm as the grafts in the aqueous phase (pH 3.0) in a hybrid vesicle structure. The resulting hollow nanogel was shown to be pH-responsive and retained its structure at pH 7.4, thus making MR imaging pH-tunable. Doxorubicin-loaded nanogels showed increased drug release rates in response to both pH reduction and temperature increase. By magnetic transport guidance toward the target and subsequent exposure to an alternating magnetic field, their Dox-loaded nanogel system demonstrated capabilities of hyperthermia and


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magnetic-triggered drug release in a cancer cell model in vitro (Chiang et al., 2013). In the case of unremovable tumors by surgical operation (because they are too hemorrhagic or localized in tissues with high risk of healthy-tissue injury), magnetic-responsive systems represent a promising therapeutic option. However, major limitation in magnetic guidance by NPs is the complexity in the setup of external magnetic fields, which needs adequate focusing and deep penetration into the tissues to reach the desired area with sufficient strength, and work is still needed to identify better magnetic technologies (Mura et al., 2013). Sensitivity of drug-delivery systems to more than one stimulus can even further improve targeted drug delivery to tumors. Enhancement in triggered drug-release due to dual responsiveness to pH and temperature has been shown for ionically self-assembled NPs and liposomes (Ta et al., 2010; Cui et al., 2012). Another system has shown responsiveness to both ultrasound and enzymes to enhance drug release from bubble liposomes (Nahire et al., 2012). Nano-carriers responding to dual-stimuli have also been studied including pH and temperature responsive P(NIPAAm-co-DMAAmco-UA) and P(NIPAAm-co-DMAAm-co-UA)-g-cholesterol NPs loaded with DOX (Soppimath et al., 2005, 2007), paclitoxel-loaded P(NIPAAm-co-AA)-b-PCL NPs (Zhang et al., 2007), mPEG-b-P(HPMA-Lac-co-His), mPEG-b-PLA and cy5.5-PEG-PLA mixed micelles carrying DOX (Chen et al., 2012), PLA-g-P(NIPAAm-co-MAA) and P(NIPAAmco-DMAAm)-b-PCL/PLA micelles carrying adriamycin (Li et al., 2011). A number of examples of pH and magnetic sensitive drug releasing systems include Fe3O4 nanocarriers coated with peptide mimic polymers loaded with DOXHCI (Barick et al., 2012) DOX-tethered Fe3O4 conjugates (Zhao et al., 2012) and mPEG-b-PMAA-b-PGMA-Fe3O4 NPs conjugated with adriamycin (Guo et al., 2008). Examples of studied temperature and magnetic-responsive systems include Pluronic-Fe3O4 NPs loaded with doxorubicin (Guo et al., 2008). Several multiresponsive carriers have also been developed recently. Thayumanavan and co-workers designed a triplestimuli sensitive NPs based on an acid-sensitive tetrahydropyran-protected 2-hydroxyethyl methacrylate (THP-protected HEMA) hydrophobic block, thermo-sensitive PNIPAAm hydrophilic block and an intervening redox-sensitive disulfide linker (Klaikherd et al., 2009).

Remained challenges Despite the large number of studied polymers for this purpose, their clinical application is limited due to a number of factors including potential toxicity, their accumulation in the body and biodegradability. For instance, acrylamide or acrylic acid-derived polymers are not hydrolytically degradable and more research on their possible adverse effects, due to neurotoxicity of AAm monomer (LoPachin, 2005), seems mandatory. Furthermore, many of the polymeric carriers most effectively reach their targets when higher molecular weights are used, thus they are not readily excreted via the kidneys after delivering the drug and accumulate in the body (Hoffman, 2013).


Stimuli–responsive drug carriers show improved control of drug release and internalization in tumor cells and their microenvironment. However, work is still needed on the sensitivity of these drug carriers to the applied stimuli. Some major drawbacks in the field may include the off-tumortargeting caused by non-specific stimuli such as overexpression of target enzymes in non-cancer cells (Spallarossa et al., 2006; Cheng et al., 2009) inaccurate utilization of stimuli such as uncontrolled depth of magnetic field and heating (Sawant et al., 2006; Zhu et al., 2009; Deng et al., 2011) and slow response to applied stimulus (Zhu et al., 2008; Wong et al., 2011).

Conclusion For tumors, it is important for smart delivery systems to respond to the specific changes in cancer cells rather than normal cells. Specific enzymatic reactions, metabolic changes and changes in pH magnitude and gradient are particular features of cancer cells and tumors that can be exploited when designing cancer-specific smart drug or gene delivery systems. Stimuli-sensitive amphiphilic block copolymers are of particular interest since they can self-assemble into various supramolecular structures, and thus provide an interior space to encapsulate the drug molecule. It is expected that the next generation of stimuli-sensitive drug delivery systems for cancer therapy will use a combination of two or more stimuli in a single system, and thus provide an opportunity to fine-tune drug release at a specific site. However, despite the fact that these systems can be appealing for more precise and controlled drug release, they also happen to be too complicated and possible advantages of many of these systems still remain to be proofed.

Acknowledgements The authors thank K. Shashok (Author AID in the Eastern Mediterranean) for his help to revise the manuscript.

Declaration of interest The authors report no conflicts of interest. The authors alone are responsible for the content and writing of this article.

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Molecular imprinted polymers as drug delivery vehicles.

Supramolecular hydrogels as drug delivery systems.

Albumin nanostructures as advanced drug delivery systems.

Particles from preformed polymers as carriers for drug delivery.

Apomaghemite as a doxorubicin carrier for anticancer drug delivery.

Drug delivery systems 5A. Oral drug delivery.

Drug delivery systems. 6. Transdermal drug delivery.

Bioadhesive polymers in peptide drug delivery.

Implantable drug-delivery systems.

Mucoadhesive polymers for buccal drug delivery.

Mathematical modeling analysis of intratumoral disposition of anticancer agents and drug delivery systems.

Self-nanoemulsifying drug delivery systems as novel approach for pDNA drug delivery.

Stimuli sensitive polymers and self regulated drug delivery systems: a very partial review.

Fluorescence optical imaging in anticancer drug delivery.

Nano-niosomes as nanoscale drug delivery systems: an illustrated review.

Cubic and hexagonal liquid crystals as drug delivery systems.

Nanoscaled Zinc Pyrazolate Metal-Organic Frameworks as Drug-Delivery Systems.

Advances in polyelectrolyte multilayer nanofilms as tunable drug delivery systems.

Sequential polydepsipeptides as biodegradable carriers for drug delivery systems.

Interpenetrating polymer networks as innovative drug delivery systems.

Polybutylcyanoacrylate nanocarriers as promising targeted drug delivery systems.

Heparin-based nanocapsules as potential drug delivery systems.

Nanovesicular carriers as alternative drug delivery systems: ethosomes in focus.

Lipid-Based Drug Delivery Systems.