Multi-targeting cancer chemotherapy using temperature-responsive drug carrier systems

Reactive & Functional Polymers 71 (2011) 235–244

Contents lists available at ScienceDirect

Reactive & Functional Polymers journal homepage: www.elsevier.com/locate/react

Multi-targeting cancer chemotherapy using temperature-responsive drug carrier systems Masamichi Nakayama, Teruo Okano ⇑ Institute of Advanced Biomedical Engineering and Science, Tokyo Women’s Medical University, TWIns, Kawada-cho 8-1, Shinjuku-ku, Tokyo 162-8666, Japan

a r t i c l e

i n f o

Article history: Available online 31 August 2010 Keywords: Cancer chemotherapy Temperature-responsive polymer Triggered drug release Liposome Polymeric micelle

a b s t r a c t Recently, a growing number of nano-scale drug carrier systems (e.g., drug–polymer conjugates, liposomes, and polymeric micelles) attract great attention for targeting cancer therapy due to a passively selective accumulation at solid tumor tissues and a subsequent anti-cancer activity. However, for the present drug targeting carrier systems, the target-selective delivery and release of loaded drugs are incapable to control completely. To overcome these current issues, stimuli-responsive drug carriers have been developed as the next-generation drug targeting systems. If drugs can be delivered to target sites via passive targeting of stimuli-responsive carriers and then released from the carriers by external physical signals, the systems are termed ‘‘multi-targeting systems” which are quite attractive for achieving the target site selective pharmaceutical action with reducing adverse effects. As possible external signals, temperature change is one of useful stimuli due to its low invasiveness to living body system and simple site-selective application using medical devices. To install temperature-responsive function to drug carriers, temperature-responsive polymers play significant roles in signal-triggering drug release and carrier-interaction with target cells and tissues. This review introduces several molecular designs for temperature-responsive drug carriers and discusses their potentials as a smart drug targeting system for an effective cancer chemotherapy. Ó 2010 Elsevier Ltd. All rights reserved.

1. Introduction Anti-cancer drug targeting to solid tumors has attracted much attention in efforts to achieve an effective cancer chemotherapy without severe toxic adverse effects. To meet this purpose, several nano-scale drug delivery vehicles, such as polymer–drug conjugates [1,2], liposomes [3,4], and polymeric micelles [5–7], have been extensively studied since the 1990s. Recent developments in both drug carrier design (the optimization of size and surface chemistry) and biocompatible polymer conjugation have led to prolong carrier circulation in the bloodstream with avoiding the body defense system (e.g., reticuloendothelial system, RES) [8] and the renal filtration [9]. Long-term circulating nanoparticles (10–200 nm in diameter) are known to spontaneously accumulate at solid tu-

Abbreviations: RES, reticuloendothelial system; EPR, enhanced permeability and retention; PEG, poly(ethylene glycol); LCST, lower critical solution temperature; IPAAm, N-isopropylacrylamide; DMAAm, N,N-dimethylacrylamide; ELP, elastin-like polypeptide; DOX, doxorubicin; EOEOVE, 2-(2-ethoxy)ethoxyethyl vinyl ether; PBMA, poly(n-butyl methacrylate); PS, polystyrene; PLA, poly(D,L-lactide); PID, poly(N-isopropylacrylamide-co-N,N-dimethylacrylamide); PBzMA, poly(benzyl methacrylate). ⇑ Corresponding author. Tel.: +81 3 5367 9945x6201; fax: +81 3359 6046. E-mail address: [email protected] (T. Okano). 1381-5148/$ - see front matter Ó 2010 Elsevier Ltd. All rights reserved. doi:10.1016/j.reactfunctpolym.2010.08.006

mor sites due to the relatively leaky tumor vasculature and the poor lymphatic drainage of macromolecules (Fig. 1). These phenomena are known as ‘‘enhanced permeation and retention (EPR) effect” [10,11]. This concept of tumor-selective macromolecular accumulation is one of the main backbones for current carrier-based tumor drug delivery. Drug carriers offer advantages including a high drugloading efficiency, reduced drug toxicity, and increased tumorselective drug accumulation. Although several nano-carriers (e.g., poly(ethylene glycol)-grafted liposomal doxorubicin, DoxilÒ) have been evaluated and shown to be partially successful chemotherapeutic drug deliveries in various in vivo models and clinical stages, some challenging issues are still remained to improve the therapeutic effects. Altogether, the present systems relied on the passive drug diffusion of anti-cancer drugs from vehicles [12] and thus are unable to control both target-selective drug delivery and release simultaneously. In other words, incorporated drugs are continually released slowly from carriers at a target site as well as in the bloodstream during drug delivery by a simple diffusion. In addition, the extravasation of drug carriers at tumor sites may varies greatly due to differences in vascular permeability of tumors. To resolve these present problems, the use of stimuli-responsive drug carrier systems with triggering drug release mechanisms at local target sites as a new strategy for site-specific drug delivery

236

M. Nakayama, T. Okano / Reactive & Functional Polymers 71 (2011) 235–244

Fig. 1. Multi-targeting cancer chemotherapeutic system using temperature-responsive drug carriers in conjunction with mild hyperthermia. After the accumulation of temperature-responsive drug carriers at tumor tissues, target-specific pharmaceutical action from the carriers can be performed by the target-selective irradiation of external heating.

and release has been extremely attracted to achieve the higher therapeutic effects of cancer treatments. Stimuli-triggering drug release profiles may give the maximal therapeutic efficacy at target tumor sites, resulting in the locoregional drug accumulation with decreasing unnecessary drug distribution in normal tissues for minimizing adverse effects. Possible stimuli being applied at target sites are target-specific physiological signals (e.g., local pH [13–15] and specific enzyme activity [1,16]) and externally applied physical energies (e.g., heating/cooling [17–20], light [21–23], and ultrasound [24]). Among these stimuli, temperature change is one of valuable stimuli because it is relatively simple, effective, and safe medical applications, such as hyperthermia to damage and kill cancer cells [25]. From this perspective, the present review mainly focuses on the strategy and molecular designing principle of drug carriers with temperature-triggered drug release mechanisms for smart drug targeting applications. The first part of the review describes ‘‘multi-targeting systems” obtained from the fusion between drug carrier technology and external physical signals (a mainly mild heating), and temperature-responsive polymer chemistry allowing the carriers to have thermal responsive function. The other part summarizes and discusses the molecular design of temperatureresponsive drug carriers with respect to triggering release mechanisms and their potentials for cancer therapy with showing typical carrier examples.

2. Multi-targeting cancer chemotherapy and hyperthermia Since the 1990s, tumor-selective carrier-based drug targeting has attracted great attention for cancer chemotherapy. Presentlyapplied drug carrier systems are mainly attributed to a passive targeting methodology. This methodology improves a target-selective drug delivery through reduced non-specific interactions with nontarget tissues and organs by optimizing carrier design, such as size (appropriate size range of 10–200 nm for long-term circulation) [26] and surface chemistry (e.g., surface potential and hydrophobicity) [27,28]. Historically, two great developments in drug targeting contributed to the successful tumor-selective drug delivery. Since the 1970s, the conjugation of low toxic and hydrophilic polymer, ‘‘poly(ethylene glycol) (PEG)”, to proteins (e.g., L-asparaginase and interferon-alpha) [29] and drug carriers (e.g., PEGgrafted liposomal doxorubicin, DoxilÒ) [4] have been extensively developed for improving their pharmacokinetics. Grafting of PEG

(PEGylation) can provide several significant pharmacological advantages including an improved carrier solubility, reduced immunogenicity and antigenicity, and enhanced protection from proteolytic degradation. Consequently, PEGylation allows drug carriers to prolong their circulatory time with avoiding the RES’s uptake and the renal excretion (i.e., ‘‘stealth” function). On the other hand, in 1986, H. Maeda et al. first presented ‘‘enhanced permeability and retention (EPR) effect”, which is a novel concept for macromolecule-based tumor targeting via tumor-specific macroscopic properties: (1) the hyperpermeability of tumor vasculature by various vascular permeability factors and (2) the poor lymphatic drainage of macromolecules from tumor tissues [10,11]. Thus, macromolecular matters are significantly localized at tumor interstitium tissues for a prolonged time. Various EPR-mediated passive targeting approaches using macromolecules (e.g., polymer–drug conjugates [1,2]) and nano-particulates (e.g. liposomes [3,4] and polymeric micelles [5,6]) to solid tumors have been successfully achieved. Clinical trials for tumor-selective and efficient drug delivery of anti-cancer drugs have been already performed, and some of them have been now approved for clinical use [30]. Drug targeting includes two components: (1) a carrier-based site-specific drug delivery and (2) a target-selective drug release from carriers. However, most current carrier systems are unable to control target-selective drug delivery and release. For example, drugs should be unreleased from carriers during delivery, while drugs should be released at their target sites effectively. To resolve these contradictory issues, stimuli-responsive carrier systems based on signaled release at local target sites are quite attractive way [31– 33]. A drug carrier technology combining two or more targeting methodologies to improve selectivity is defined as ‘‘multi-targeting systems”. To develop a multi-targeting system giving target-selective delivery (spatial control) and release (temporal control), stimuli-responsive carrier systems have been extensively studied for drug targeting on demand in response to internal and/or external stimuli. Specific chemical and/or physical stimuli at target sites include internal signals (e.g., local pH [13–15] and specific enzymatic activity [1,16]) or external applied signals (e.g., temperature variation [17,19,20], light [21–23], and ultrasound [24]). Among these possible signals, mild heating is one of the most preferable stimuli because of its easy-to-use approach, a low invasiveness to normal tissues, and a widely used technique in medical practice. A potential cancer therapy, ‘‘mild hyperthermia”, involves heating tumor tissues above body temperature (up to 41–43 °C) to damage cancer cells and/or for enhancing the therapeutic outcome

M. Nakayama, T. Okano / Reactive & Functional Polymers 71 (2011) 235–244

of combine cancer chemotherapy [25,34]. Mild hyperthermia is known to promote blood flow and vascular permeability at tumor tissues compared to normal tissues [25]. In addition, hyperthermia affects the biological functions of cancer cells (e.g., inhibited DNA synthesis and repair, disrupted microtubules, and altered receptor expression) without any biological influence on normal cells [34]. Current hyperthermia technology realizes region-controlled heating through focusing ultrasound irradiation or radiofrequency electrode implantation at solid tumor sites. The fusion between temperature-responsive drug carrier technology and locally controlled mild hyperthermia at target tumor tissues can be a potential tumor treatment (Fig. 1). The local heating of target tumors increases tumor blood flow and vascular permeability causing enhanced carrier accumulation and thermally induced pharmacological action for significant therapeutic effects. 3. Temperature-responsive polymers For constructing smart materials showing a thermo-responsive function, temperature-responsive polymers, which change their conformation and physical properties in response to temperature change, have been extensively utilized in biomaterials and drug delivery fields [31]. Some polymers undergo coil-to-globule transitions in water above a specific temperature, called ‘‘lower critical solution temperature (LCST)”. In other words, the polymer can change self-conformation from a hydrophilic state (random coil) to a hydrophobic state (globule and aggregation) across its LCST [35]. A series of poly(N-substituted acrylamide) derivatives are known to demonstrate thermal phase transitions at various temperatures, dependent on their individual chemical structures [36]. Among the LCST-type polymers, polymers exhibiting their LCST around typical human body temperature have been used in various smart thermoresponsive applications. Especially, poly(N-isopropylacrylamide) (PIPAAm) which exhibits its LCST at 32 °C in water [37,38]. Therefore, it has been nowadays extensively investigated for novel applications such as drug delivery systems [39,40], bioseparation [41,42], and tissue engineering [43]. The important and valuable property of temperature-responsive polymers allows us to design and tailor LCST through their hydrophilic/hydrophobic comonomer composition [44,45], end-functionalization [46,47], stereochemistry [48]. Using these temperature-responsive polymers with controlled LCST around body temperature, they can be applied for specific biomedical applications. 4. Temperature-responsive polypeptide–drug conjugates Elastin-like polypeptides (ELPs) as artificial temperatureresponsive biopolymers have been developed by Urry [49]. A series of Val–Pro–Gly–X–Gly (VPGXG) (where X, as a guest residue, is any amino acid except for Pro) pentapeptide repeated-polymers exhibit their individual LCSTs, which are controllable by substituting the guest residues (X) in the repeat sequence to other amino acids with different hydrophobicity [20,49]. Thus, the LCST of ELP can be adjusted to 41 °C from 27 °C (where is the transition temperature of natural elastin sequence, X: Val) through the substitution with more hydrophilic units (X: Ala or Gly) [20]. The great advantages of ELPs include uniform molecular weights and the location-controlled introduction of reactive amino acid residues (e.g., lysine and cysteine) due to recombinant synthesis in Escherichia coli by the overexpression of artificial genes. By using ELPs showing a LCST of 41 °C, Chilkoti and Raucher have investigated temperature-triggered drug targeting using macromolecular carriers to solid tumors in conjunction with mild hyperthermia [20,50–52]. ELP in combination with mild hyperthermia (up to 42 °C) showed a 2-fold increase in the carrier accu-

237

mulation at the heated tissues of human tumors implanted in nude mice, compared to that of normothermic tissues [50]. In addition, the 2–3-fold greater internalization of ELPs inside various human cancer cells have been also performed by combination with mild hyperthermia [51]. ELPs occurred the thermal phase transition above its LCST and formed hydrophobic polymer aggregates in diameters of 0.1–1 lm. Thus, ELP aggregates hydrophobically interacted with cell membranes, and subsequent intracellular uptake was promoted, especially approximately 0.1 lm particles were significantly internalized inside the cells [51]. However, the magnitude of thermally induced cellular uptake depends on cancer cell types. Therefore, cell-penetrating Tat peptide was terminally introduced to ELP for improving the extent of cellular uptake and achieved the 20-fold enhancement of reducing cancer cell viability in conjunction with hyperthermia [52]. As the molecular design of macromolecular drug carriers, ELPs conjugated with anti-cancer drug, doxorubicin (DOX), were investigated (Fig. 2). DOX were often conjugated to ELP using a low pH-cleavable chemical linker (e.g., hydrazone bond) [20] or enzymatic cleavable tetrapeptide, Gly–Phe–Leu–Gly (GPLG) [52]. In both case, after the cellular uptake of ELP-drug conjugates, drugs can be accelerated to release from the ELP backbones through the cleavage of environment-sensitive linkers in lysosomal/endosomal environment, such acidic pH (<6.5) or lysosomal enzyme degradation. Based on these results, ELP-drug conjugates will be a promising candidate for smart tumor targeting combined with hyperthermia treatment.

5. Temperature-responsive liposomes Since first discovery in the 1960s, liposomes as phospholipidbased nanovesicles have been studied extensively for drug delivery vehicles, especially for tumor-selective drug targeting, and some of them have been already in clinical uses [4]. Recently, functional liposomal carriers with drug release triggered by physical and/or chemical stimuli have been developed as an attractive strategy for smart drug targeting [53,54]. In the late 1970s, Yatvin and Weinstein have pioneered temperature-responsive liposomal drugs using a phospholipid mixture of dipalmitoylphosphatidylcholine/distearoylphosphatidylcholine (the ratio: 3/1), which shows a gel-to-liquid crystalline phase transition at a temperatures of around 41 °C for combining with mild hyperthermia [55,56]. These liposomes showed an accelerated release of liposome-entrapped water-soluble drugs, caused by the permeability enhancement of liposomal membranes through the lipid phase transition (Fig. 3A) [57,58]. Temperature-responsive liposomal drugs have been currently developed by providing long-circulating properties via the PEGylation technology [59], and by incorporating additional lipids (e.g., lysolipid [17]) that improve the thermal induced membrane permeability. More recently, lysolipid formulated liposomes have been further evolved as temperature-responsive liposomal doxorubicin (ThermoDoxÒ) by a biopharmaceutical company (Celsion), and the global clinical trials are performed for primary liver cancer (phase III) and recurrent chest wall breast cancer (phase I). In the present medical protocol of ThermoDoxÒ for liver cancers, radiofrequency ablation (RFA) as a heating treatment is carried out after the accumulation of liposomal drugs at their target tumor tissues. High intensity focused ultrasound (HIFU) is also investigated as other candidate for hyperthermic treatments. Another approach for liposome with thermally triggering drug release function involves temperature-responsive polymers grafted liposomal carriers. Various LCST-type polymers have been employed for fixation onto liposome membranes for achieving temperature-controlled structural change and a subsequent drug release from the liposomal carriers [54,60–62]. At a temperature

238

M. Nakayama, T. Okano / Reactive & Functional Polymers 71 (2011) 235–244

Fig. 2. Doxorubicin-conjugated elastin-like polypeptides (ELP) for a thermal drug targeting; (A) ELP conjugated with doxorubicin via an acid-sensitive hydrazone linker, (B) Tat-ELP conjugated with doxorubicin via a lysosomal enzyme-cleavable (GFLG) linker.

Fig. 3. Schematic representation of drug release mechanisms from temperature-responsive liposomes; (A) drug release based on thermally enhanced membrane permeability changes via lipid phase transition, (B) drug release based on a membrane destabilization via phase transition of temperature-responsive polymers grafted to liposome.

below the LCST, temperature-responsive polymers grafted on the membranes show a hydrophilic property with extending polymer conformation, while the polymers undergo their thermal phase transition to hydrophobic and conformational change to an aggregated state above the LCST. This mechanical distortion derived from polymer phase transition strongly destabilizes liposomal membranes and induces accelerated the release rates of drugs entrapped inside liposomes (Fig. 3B). As possible temperatureresponsive polymers, PIPAAm derivatives possessing hydrophobic moieties act as polymer–anchors are most frequently used, since Ringsdorf and co-workers first reported [60]. Chemical structures and characteristic of temperature-responsive polymers significantly affect on the thermo-responsive function of grafted liposomes. For example, the anchor positions in polymer chain are one of considerable factors for constructing liposomes with a higher thermal response. Liposomes grafted with

polymers possessing terminal anchors showed a sharper release profile of entrapped drugs in a narrow temperature range than those grafted via the anchors located at polymer chains randomly [62]. This possible reason is that polymer chain fixed at terminal positions bring about their thermal coil-to-globule transitions effectively due to a high conformational flexibility, compared to that fixed at the arbitrary and plural points of polymer chain. As another candidate for temperature-responsive polymers, poly [2-(2-ethoxy)ethoxyethylvinylether(EOEOVE)] , which exhibits its LCST around 40 °C, was investigated to give a sharp thermal response and drastic membrane destabilization in a narrow temperature range [63]. Recently, Kono et al. reported a promising tumor-specific chemotherapy using DOX-loaded liposomes grafted with poly(EOEOVE)-b-poly(octadecyl vinyl ether) (the latter block act as hydrophobic anchors to liposomal membranes) [64]. Although these liposomes were stable and entrapped DOX at

M. Nakayama, T. Okano / Reactive & Functional Polymers 71 (2011) 235–244

239

physiological temperatures, upon heating above 40 °C across the LCST, encapsulated DOX release was enhanced through the membrane destabilization caused by polymer phase transition. Furthermore, independently grafted PEG chains to the poly(EOEOVE)liposomes provide a long-circulating property and pharmacokinetics similar to that of PEG-grafted liposomes [65]. The combination therapy with poly(EOEOVE)-grafted liposomal drugs and tumorselective mild hyperthermia with a radiofrequency oscillator (45 °C for 10 min) to tumor-bearing mice showed a significantly suppressive effect on tumor growth [65]. These results indicate that temperature-responsive liposomal drugs in conjunction with tumor-selective heating have a great potential to improve the therapeutic effects of current liposomal-based cancer chemotherapy. 6. Temperature-responsive polymeric micelles In the past two decades since Kataoka and co-workers first proposed an anti-cancer drug delivery using PEG-block copolymer micelles, many types of micellar vehicles have been continuously investigated for improving in vivo pharmaceutical activity to solid tumors. Polymeric micelle carrier system is a promising candidate as tumor-selective drug delivery tools due to (1) nano-scale sizes (typically 10–100 nm) for providing a long-term circulation in the bloodstream, (2) a high loading capacity of anti-cancer drugs inside the inner cores, and (3) densely packed hydrophilic polymer structures as outer coronas, which reduce possible interactions with serum components and escape body’s defense systems such as RES uptake [5,6,66]. Based on these characters, passive targeting of polymeric micelles can be achieved through the EPR effect of solid tumors. Already, some clinical trials for the micellar vehicles containing various anti-cancer drugs (e.g., doxorubicin, paclitaxel, and cisplatin) have been performed at both domestic and foreign sites [67,68]. Temperature-triggering function to polymeric micelle carriers can be added by applying temperature-responsive polymers as one part of block copolymers. Thus, temperature-responsive polymeric micelles are classified into temperature-responsive ‘‘cores” and ‘‘coronas” systems. In either way, temperatureresponsive polymeric micelles demonstrate thermally modulated structural and characteristic changes significantly for possible temperature-triggering applications. 6.1. Polymeric micelles with temperature-responsive ‘‘cores” Block copolymers comprising hydrophilic segments and LCSTbased polymer segments (e.g., PEG-b-PIPAAm [18]) form polymeric micelles possessing temperature-responsive inner ‘‘cores” as shown in Fig. 4A. Below the LCST, micelle-forming block copolymers are isolated as dispersed polymer chains due to the watersoluble property of both blocks. However, upon heating above the LCST, the temperature-responsive blocks show their phase transitions and become hydrophobic to associate with each polymer chain, resulting in a subsequent formation of core-corona micellar structures. A possible advantage of polymeric micelles having temperature-responsive cores is a rapid drug release with applied stimuli. This is due to the collapse of hydrophobic inner core as a drug reservoir by reducing environmental temperature below the LCST, and this event induces a quick drug release without a diffusion process of drugs through the core-forming polymer matrix. Another possible advantage is that drugs can be incorporated just by heating aqueous polymer solution containing drug molecules above the LCST. Therefore, no organic solvent is required for incorporating drugs inside the cores, unlike conventional drug loading method into polymeric micelles [69]. On the contrary, the disadvantages involve an issue, ‘‘how to decrease temperature in living system” for releasing entrapped drugs. In medical applica-

Fig. 4. Thermally induced drug release systems from temperature-responsive polymeric micelles; polymeric micelles with temperature-responsive inner core (A) and outer corona (B).

tion, temperature decrease treatment, ‘‘hypothermia” can be applied only to body surfaces or limited regions around surface such as the skin, blood vessels, and tissue near body surface. In addition, the selection and molecular design of core-forming polymer blocks is limited. There is a restriction on the strength of possible interactions for drug incorporation. The nature of these interactions is of paramount importance, since it may determine the drug-loading efficiency, drug release profile, and thermal response of polymeric micelles. To overcome these limitations, Hennink’s group presented one promising strategy using hydrolytically sensitive micelles comprising PEG-b-poly[IPAAm-co-N-(2-hydroxypropyl) methacrylamide(lactate)n] [70,71]. Hydrolytically sensitive oligo lactide (OLA) side chains act as both lower LCST shifters and drug loading enhancers of core-forming blocks. After hydrolysis of OLA side chain, the resultant monomer units showed more hydrophilic than IPAAm units, resulting in the elevated LCST of core-forming polymer segments to 45 °C. Consequently, the block polymers become watersoluble at normal body temperature [71], resulting in micelle dissociation. This strategy may improve temperature-responsive core systems for in vivo applications. 6.2. Polymeric micelles with temperature-responsive ‘‘coronas” Alternatively, polymeric micelles with temperature-responsive ‘‘coronas” can be fabricated using block copolymers composed of thermoresponsive blocks (mainly PIPAAm derivatives) and hydrophobic blocks as depicted in Fig. 4B. Thermal transition temperature can be modulated to desirable temperatures (for example, 40 °C for combination with mild hyperthermia) by introducing hydrophilic comonomers such as N,N-dimethylacrylamide (DMAAm) [44,72]. Below the LCST, corona-forming temperature-responsive polymer chains stabilize a clearly-separated core-corona micelle structure due to their highly hydrated and hydrophilic features. However, upon raising temperatures above the LCST, the outer coronas switch to be hydrophobic, resulting in the collapse of micellar coronas and the aggregation of micelles. The significant micellar property changes may be utilized as a trigger for releasing entrapped drugs. A possible advantage of these systems as smart drug delivery vehicles is a target-specific pharmaceutical action

240

M. Nakayama, T. Okano / Reactive & Functional Polymers 71 (2011) 235–244

by combination with target-selective heating, such as local hyperthermia and focused ultrasound irradiation. As other advantages, various hydrophobic polymer candidates can be chosen for coreforming segments, compared with temperature-responsive core systems described in the previous Section 6.1. Our group first reported functional polymeric micelles with temperature-triggered drug release mechanisms using block copolymers comprising PIPAAm and various hydrophobic polymers such as poly(n-butylmethacrylate) (PBMA) and polystyrene (PS) [73,74]. These works presented that the physico-chemical properties (e.g., hydrophobicity and flexibility/rigidity) of coreforming hydrophobic polymers is quite significant factors for achieving temperature-induced release modulation of incorporated drugs. For example, when PIPAAm-b-PS was used for micelle formation, the release of entrapped anti-cancer drugs, DOX, was suppressed at temperatures both above and below the PIPAAm’s LCST of 32 °C (Fig. 5) [74]. By contrast, the PIPAAm-b-PBMA mi-

Fig. 5. (A) Entrapped drug release modulations from temperature-responsive polymeric micelles loaded with doxorubicin (DOX) in response to temperature cycles between 4 and 40 °C.

celles showed a distinctive drug release profile. At a temperature below the LCST, the release of encapsuled DOX was at the minimum value, less than 10%. However, increasing temperature above the LCST initiated to accelerate DOX-release rate [74]. In addition, the temperature cycles induced on–off drug release modulation from PIPAAm-b-PBMA micelles between 4 and 40 °C (Fig. 5) [74]. The difference in thermal response of cores is probably related to the physico-chemical properties of respective core-forming polymers. In fact, the glassy transition temperature (Tg) of PBMA (Tg: ca. 20 °C < LCST) is lower than that of PS (Tg: ca. 60 °C > LCST). Therefore, the thermal triggering drug release may be caused by the structural destabilization of rubber-like PBMA core, which is induced by the aggregation of dehydrated outer coronas. Namely, the aggregation of covalently-connected temperature-responsive polymers probably gives a mechanical distortion to the inner cores. Consequently, the distortion allows water molecules to penetrate into the cores, resulting in producing channels for drug release. Temperature-responsive micelles with biodegradable cores comprising poly(D,L-lactide) (PLA), poly(e-caprolactone) (PCL), and their copolymers have been further developed [72,75]. The biodegradable system may allow the polymers to be rapidly excreted by the kidney after the degradation of hydrophobic blocks (if the diameters of non-degradable segments are below a critical value for the renal excretion of approximately 3 nm [9]). In aqueous environment, P(IPAAm-co-DMAAm)-b-PLA block copolymers incorporated DOX and formed polymeric micelles with 69 nm in an averaged diameter. In addition, approximately 6-fold enhancement of entrapped drug release have been achieved by heating from normal body temperature (37 °C) to 42.5 °C across the micellar LCST of 40 °C [72]. Moreover, DOX-loaded polymeric micelles combined with mild hyperthermia demonstrated a significant cytotoxicity against cultured endothelial cells [72]. Stimuli-triggered active interactions of polymeric micelles with target cells and tissues are quite valuable for the effective intracellular delivery of hydrophobic drugs or bioactive molecules. Recently, our groups reported a temperature-triggering intracellular uptake system with PIPAAm-based polymeric micelles (Fig. 6) [76,77]. At below the LCST of 40 °C, densely packed and hydrated thermoresponsive corona-forming polymers reduced possible interactions of the hydrophobic cores with cell surfaces. Upon raising temperatures above the LCST (e.g., 42 °C), the corona-forming polymer chains collapsed via the dehydration of IPAAm units. Be-

Fig. 6. Confocal images of P(IPAAm-co-DMAAm)-b-PLA micelles (LCST: 39.5 °C) localized within cultured cells after incubation for 9 h (a) below the LCST (37 °C) and (b) above the LCST (42 °C). The nuclei and cytoplasm were stained with Hoechst 33258 (blue) and Cell Tracker Red (red), respectively. Green fluorescence was derived from fluorescently labeled micelles. Scale bars: 50 lm. Reproduced with permission from [77]. Copyright 2009 American Chemical Society.

M. Nakayama, T. Okano / Reactive & Functional Polymers 71 (2011) 235–244

cause of these conformational changes, the hydrophobic interactions between polymeric micelles and cell membranes might increase significantly. Consequently, the adhesion of polymeric micelles to cell surfaces might be promoted, followed by an enhancement of intracellular uptake. As a result, approximately 16-fold greater internalization has been achieved by mild heating above the LCST, compared with that below the LCST (Fig. 7). By contrast, an extremely low internalization of PEG-corona micelles was observed, regardless of temperature changes. Based on these results, temperature-responsive ‘‘corona” micelles may be a promising intracellular drug delivery tool combining with mild thermal stimuli. 6.3. Surface functionalization of thermoresponsive polymeric micelles Nowadays, a recently developed living radical polymerization provides various polymer architectures with controlled molecular weights, and the method also allows polymer termini to be converted with various functional groups easily [78,79]. Outermost

Fig. 7. Time-dependent cellular uptake of fluorescently labeled P(IPAAm-coDMAAm)-b-PLA micelles (LCST: 39.5 °C) at temperatures below (37 °C, the open circle) and above (42 °C, the closed circle). Reproduced with permission from [77]. Copyright 2009 American Chemical Society.

241

surface functionalized polymeric micelles can be formed through the molecular assemblies of end-functionalized amphiphilic block copolymers. Functionalization of micellar surface gives additional unique functions to polymeric micelles. For instance, the conjugation of target-specific bioactive molecules (e.g., galactose [80] and cyclic RGDS peptide [81]) to micellar surfaces increases the active interactions of polymeric micelles with target cells and tissues (Fig. 8). Our group has demonstrated the quite unique thermal phase transitions of PIPAAm-based corona micelles with outermost surface functional groups. Hydrophobicity and polarity of surface functionality significantly affected the thermoresponsive behavior of polymeric micelles with constant particle sizes and critical micelle concentration [82]. Poly(N-isopropylacrylamide-co-N,N-dimethylacrylamide)-b-poly(benzyl methacrylate) (PID-b-PBzMA) with various PID chain lengths by reversible addition–fragmentation chain transfer radical (RAFT) polymerization [83]. RAFT polymerization has a great advantage that functional groups can be introduced by reacting selected coupling agents with exposed polymer terminal thiol groups via the aminolysis of polymer growth-terminal [84]. Only hydrophobic phenyl (Phe)-surface PID-b-PBzMA micelles exhibited significant LCST shifts to lower temperatures than those of hydrophilic hydroxyl (OH)-surface micelles, although both have the same chemical composition of diblock copolymers except for end functionality (Fig. 9) [82]. Furthermore, lower LCST shifts were promoted with reducing the molecular weights of temperature-responsive PID chains. Our group have previously reported that hydrophobic semi-telechelic PIPAAm exhibited a different thermoresponsive behavior, dependent on polymer molecular weights. The freely-mobile hydrophobic end-groups promote the dehydration of proximal IPAAm units and further disrupt the polymer hydration, resulting in the lower LCST shifts below pure PIPAAm’s LCST [46]. Similarly, for the Phe-surface micelles, the lower LCST shifts were attributed to the influence of surface hydrophobic groups, which located at the micellar surface (temperature-responsive polymer termini). However, the large magnitude of LCST shifts were observed for the micelle system, compared to linear polymers. This reason is ascribable to the architecture of polymeric micelles, which possess densely packed polymer brush corona structures. Close-packed temperature-responsive polymers facilitate a cluster effect for concentrated hydrophobic groups on the micelle surfaces. Accumulation of hydrophobic groups surrounding polymeric micelles increased the overall hydrophobicity of micelle systems. As a result, the total enhancement of hydrophobic effect induced a drastic micellar LCST shift to lower temperatures, rather than that of dispersed linear polymers in water [82].

Fig. 8. Schematic representation and functions of outermost surface-functionalized thermoresponsive polymeric micelles formed with end-functional thermoresponsive block copolymers.

242

M. Nakayama, T. Okano / Reactive & Functional Polymers 71 (2011) 235–244

Fig. 10. LCST profiles for polymeric micelles comprising the blends of block copolymers with phenyl (Phe) and hydroxyl (OH) termini at various ratios in wt.%; the closed circle, Phe/OH: 100/0; the open circle: 90/10; the closed diamond: 75/25; the open diamond: 50/50; the closed triangle: 25/75; the open triangle: 0/100. LCST profiles were determined by transmittance changes at 600 nm in phosphate buffer saline (pH 7.4) at a heating rate of 0.1 °C/min and the polymer concentration: 10 mg/ml. Reproduced with permission from [85]. Copyright 2008 American Chemical Society.

tions cooperatively (Fig. 11). Furthermore, a series of micelles with various Phe/OH end-group ratios showed individual phase transitions, dependent on their blended compositions (Fig. 10) [85]. In future work, this unique transition behavior of surface-functionalized thermoresponsive coronas might be potentially promising for the creation of novel smart materials including targetable drug carriers and/or nano-scale sensors, which respond to multiple stimuli such as surface chemistry modulations with applied signals such as pH, light, and biomolecular interactions. 7. Conclusions

Fig. 9. LCST profiles for the dispersions of micelles having (A) PID hydroxyl terminal groups and (B) PID phenyl terminal groups on the PID-b-PBzMA micelle surfaces. Number-averaged molecular weight (Mn) of PBzMA is 2700. Micelles comprising diblock copolymers with various Mn of thermoresponsive PID segments (the closed circle: 6500, the open circle: 8500, the closed diamond: 14,000, and the open diamond: 36,000). LCST profiles were determined by transmittance changes at 600 nm in phosphate buffer saline (pH 7.4) at a heating rate of 0.1 °C/min and the polymer concentration: 10 mg/ml. Reproduced with permission from [82]. Copyright 2005 American Chemical Society.

More interestingly, block copolymer-mixed micelles comprising phenyl- and hydroxyl-block copolymers (Phe/OH-surface PID-bPBzMA micelles) showed only one sharp transition between the individual LCSTs of both homogeneous surface micelles (Fig. 10) [85]. These unique phenomena are probably due to that respective end-functional temperature-responsive polymer chains in one close-packed environment are influenced by surrounding polymer chains and/or end-functional groups, resulting in thermal transi-

Substantial progresses in nano-carrier technology, cancer biology, and polymer chemistry enable a valuable carrier-based tumor-specific drug targeting. However, current vehicles still remain some issues to be improved for effective site-specific pharmaceutical activities. As leading cancer treatments of the 21st century, external energy-triggered release systems from stimuliresponsive nanovehicles are quite prospective because of a higher efficacy in suppressing cancer and reduced drug accumulation in normal organs. Especially, a possible strategy using temperatureresponsive carriers in conjunction with local mild hyperthermia is greatly promising, because tumor tissues can be attacked by both a tumor-specific pharmaceutical activity and a thermo-therapeutic effect. Current and future developments in this area will contribute to open up a new frontier in the next-generation cancer chemotherapy. Acknowledgments The authors are grateful to Dr. N. Ueno (Tokyo Women’s Medical University) for his valuable comments and manuscript editing. This work was partially supported by NEDO Special Courses for

M. Nakayama, T. Okano / Reactive & Functional Polymers 71 (2011) 235–244

243

Fig. 11. Schematic illustration of thermal phase transition of homogeneous and heterogeneous end-functional thermoresponsive polymers under close-packed conditions produced in the micellar corona. Block copolymer-mixed micelles comprising phenyl- and hydroxyl-block copolymers (Phe/OH-surface micelle) showed only one sharp transition between the individual LCSTs of both homogeneous surface micelles due to the influence of surrounding polymer chains and/or end-functional groups.

Development of Innovative Drug Delivery Systems from New Energy and Industrial Technology Development Organization (NEDO), Japan, and Formation of Innovation Center for Fusion of Advanced Technologies in the Special Coordination Funds for Promoting Science and Technology from the Ministry of Education, Culture, Sports, Science and Technology (MEXT), Japan.

References [1] D. Putnum, J. Kopecek, Adv. Polym. Sci. 122 (1995) 55–123. [2] R.B. Greenwald, T.H. Choe, J. Mcguire, C.D. Conovor, Adv. Drug Deliv. Rev. 55 (2003) 217–250. [3] O. Ishida, K. Maruyama, K. Sasaki, M. Iwatsuru, Int. J. Pharm. 190 (1999) 49–56. [4] A. Coukell, C. Spencer, Drugs 53 (1997) 520–538. [5] K. Kataoka, G.S. Kwon, M. Yokoyama, T. Okano, Y. Sakurai, J. Control. Release 24 (1993) 119–132. [6] G.S. Kwon, K. Kataoka, Adv. Drug Deliv. Rev. 16 (1995) 295–309. [7] A.V. Kabanov, V.Y. Alakhov, Crit. Rev. Therm. Drug. Carrier Syst. 19 (2002) 1– 72. [8] R. Gref, Y. Minamitake, M.T. Peracchia, V. Trubetskoy, V. Torchilin, R. Langer, Science 263 (1994) 1600–1603. [9] L.W. Seymour, R. Duncan, J. Strohalm, J. Kopecek, J. Biomed. Mater. Res. 21 (1987) 1341–1358. [10] Y. Matsumura, H. Maeda, Cancer Res. 46 (1986) 6387–6392. [11] H. Maeda, L.W. Seymour, Y. Miyamoto, Bioconjugate Chem. 3 (1992) 351–362. [12] H. Maeda, J. Wu, T. Sawa, Y. Matsumura, K. Hori, J. Control. Release 65 (2000) 271–284. [13] K. Na, E.S. Lee, Y.H. Bae, J. Control. Release 87 (2003) 3–13. [14] E.S. Lee, K. Na, Y.H. Bae, Nano Lett. 5 (2005) 325–329. [15] Y. Bae, N. Nishiyama, S. Fukushima, H. Koyama, M. Yasuhiro, K. Kataoka, Bioconjugate Chem. 16 (2005) 122–130. [16] R. Duncan, S. Dimitrijevic, E.G. Evagoroa, S.T.P. Pharm. Sci. 6 (1996) 237–263. [17] D. Needham, G. Anyarambhatla, G. Kong, M.W. Dewhirst, Cancer Res. 60 (2000) 1197–1201. [18] M.D.C. Topp, P.J. Dijkstra, H. Tapsma, J. Feijen, Macromolecules 30 (1997) 359– 373. [19] M. Nakayama, T. Okano, J. Drug Deliv. Sci. Technol. 16 (2006) 35–44. [20] A. Chilkoti, M.R. Dreher, D.E. Meyer, D. Raucher, Adv. Drug. Deliv. Rev. 54 (2002) 623–630. [21] C.M. Peterson, J.M. Lu, Y. Sun, C.A. Peterson, J.-G. Shiah, R.C. Straight, J. Kopecek, Cancer Res. 56 (1996) 3980–3985. [22] J. Jinqiang, X. Tong, Y. Zhao, J. Am. Chem. Soc. 127 (2005) 8290–8291. [23] N. Nishiyama, Y. Nakagishi, Y. Morimoto, P.-S. Lai, K. Miyazaki, K. Urano, S. Horie, M. Kumagai, S. Fukushima, Y. Cheng, W.-D. Jang, M. Kikuchi, K. Kataoka, J. Control. Release 133 (2009) 245–251. [24] Z.-G. Gao, H.D. Fain, N. Rapoport, J. Control. Release 102 (2005) 203–222. [25] A.M. Ponce, Z. Vujaskovic, F. Yuan, D. Needham, M.W. Dewhirst, Int. J. Hyperthermia 22 (2006) 205–213. [26] D.C. Litzinger, A.M. Buiting, N.C. Rooijen, L. Huang, Biochim. Biophys. Acta 1190 (1994) 99–107. [27] Y. Takakura, M. Hashida, Pharm. Res. 13 (1996) 820–831. [28] L. Illum, S.S. Davis, R.H. Müller, E. Mak, P. West, Life Sci. 40 (1987) 367–374.

[29] F.M. Veronese, J.M. Harris, Adv. Drug Deliv. Rev. 54 (2002) 453–456. [30] R. Duncan, Nature Rev. Drug Discov. 2 (2003) 347–360. [31] T. Okano, N. Yui, M. Yokoyama, R. Yoshida, Advances in Polymeric Systems for Drug Delivery, Gordon and Breach Science Publishes, Switzerland, 1994. [32] S. Cammas, K. Suzuki, Y. Sone, Y. Sakurai, K. Kataoka, T. Okano, J. Control. Release 48 (1997) 157–164. [33] K.T. Oh, H. Yin, E.S. Lee, Y.H. Bae, J. Mater. Chem. 17 (2007) 3987–4001. [34] R.K. Jain, Cancer Metastasis Rev. 6 (1987) 559–593. [35] S. Fujishige, K. Kubota, I. Ando, J. Phys. Chem. 93 (1989) 3311–3313. [36] Y.H. Bae, T. Okano, S.W. Kim, J. Polym. Sci. Part B: Polym. Phys. 28 (1990) 923– 936. [37] M. Heskins, J.E. Guillet, J. Macromol. Sci. Chem. A2 (1968) 1441–1455. [38] H. Schild, Prog. Polym. Sci. 17 (1992) 163–249. [39] A.S. Hoffman, J. Control. Release 6 (1987) 297–305. [40] A. Kikuchi, T. Okano, Adv. Drug Deliv. Rev. 54 (2002) 53–77. [41] A. Kikuchi, T. Okano, Prog. Polym. Sci. 27 (2002) 1165–1193. [42] D. Umeno, T. Mori, M. Maeda, Chem. Commun. (1998) 1433–1434. [43] N. Matsuda, T. Shimizu, M. Yamato, T. Okano, Adv. Mater. 19 (2007) 3089– 3099. [44] Y.G. Takei, T. Aoki, K. Sanui, N. Ogata, T. Okano, Y. Sakurai, Bioconjugate Chem. 4 (1993) 341–346. [45] H. Feil, H.Y.H. Bae, J. Feijen, S.W. Kim, Macromolecules 26 (1993) 2496– 2500. [46] J.E. Chung, M. Yokoyama, T. Aoyagi, Y. Sakurai, T. Okano, J. Control. Release 53 (1998) 119–130. [47] F. Kohori, K. Sakai, T. Aoyagi, M. Yokoyama, Y. Sakurai, T. Okano, J. Control. Release 55 (1998) 87–98. [48] B. Ray, Y. Okamoto, M. Kamigaito, M. Sawamoto, K. Seno, S. Kanaoka, S. Aoshima, Polym. J. 37 (2005) 234–237. [49] D.W. Urry, Prog. Biophys. Mol. Biol. 57 (1992) 23–57. [50] D.E. Meyer, A.G.A. Kong, M.W. Dewhirst, M.R. Zalutsky, A. Chilkoti, Cancer Res. 61 (2001) 1548–1554. [51] D. Raucher, A. Chilkoti, Cancer Res. 61 (2001) 7163–7170. [52] G.L. Bidwell, I. Fokt, W. Priebe, D. Raucher, Biochem. Pharmacol. 73 (2007) 620–631. [53] M.B. Yatvin, W. Kreutz, B.A. Horwitz, M. Shinitzky, Science 210 (1980) 1253– 1255. [54] K. Kono, Adv. Drug Deliv. Rev. 53 (2001) 307–319. [55] M.B. Yatvin, J.N. Weinstein, W.H. Dennis, R. Blumenthal, Science 202 (1978) 1290–1293. [56] J.N. Weinstein, R.L. Magin, M.B. Yatvin, D.S. Zaharko, Science 204 (1979) 188– 191. [57] D. Papahadjopoulos, K. Jacobson, S. Nir, T. Isac, Biochim. Biophys. Acta 311 (1973) 330–348. [58] M.C. Blok, E.C.M. van der Neut-Kok, L.L.M. van Deenen, J. de Gier, Biochim. Biophys. Acta 406 (1975) 187–196. [59] S. Unezaki, K. Maruyama, N. Takahashi, M. Koyama, T. Yuda, A. Suginaka, M. Iwatsuru, Pharm. Res. 11 (1994) 1180–1185. [60] H. Ringsdorf, J. Venzmer, F.M. Winnik, Angew. Chem. Int. Ed. Engl. 30 (1991) 315–318. [61] X.S. Wu, A.S. Hoffman, P. Yager, Polymer 33 (1992) 4659–4662. [62] A. Polozova, F.M. Winnik, Biochim. Biophys. Acta 1326 (1997) 213–224. [63] S. Aoshima, S. Sugihara, J. Polym. Sci. Part A: Polym. 38 (2000) 3962–3965. [64] K. Kono, T. Murakami, T. Yoshida, Y. Haba, S. Kanaoka, T. Takagishi, S. Aoshima, Bioconjugate Chem. 16 (2005) 1367–1374.

244

M. Nakayama, T. Okano / Reactive & Functional Polymers 71 (2011) 235–244

[65] K. Kono, T. Ozawa, T. Yoshida, F. Ozaki, Y. Ishizaka, K. Maruyama, C. Kojima, A. Harada, S. Aoshima, Biomaterials 31 (2010) 7096–7105. [66] M. Yokoyama, T. Okano, Y. Sakurai, H. Ekimoto, C. Shibazaki, K. Kataoka, Cancer Res. 51 (1991) 3229–3236. [67] M. Matsumura, T. Hamaguchi, T. Ura, K. Muro, Y. Yamada, Y. Shimada, K. Shirao, T. Okusaka, H. Ueno, M. Ikeda, N. Watanabe, Brit. J. Cancer 91 (2004) 1775–1781. [68] T. Hamaguchi, Y. Matsumura, M. Suzuki, K. Shimizu, R. Goda, I. Nakamura, I. Nakatomi, M. Yokoyama, K. Kataoka, T. Kakizoe, Brit. J. Cancer 92 (2005) 1240– 1246. [69] G. Gaucher, M.-H. Dufresne, V.P. Sant, N. Kang, D. Maysinger, J.-C. Leroux, J. Control. Release 19 (2005) 169–188. [70] D. Neradovic, C.F. Van Nostrum, W.E. Hennink, Macromolecules 34 (2001) 7589–7591. [71] D. Neradovic, O. Soga, C.F. Van Nostrum, W.E. Hennink, Biomaterials 25 (2004) 2409–2418. [72] F. Kohori, K. Sakai, T. Aoyagi, M. Yokoyama, M. Yamato, Y. Sakurai, T. Okano, Colloids Surf. B: Biointerfaces 16 (1999) 195–205. [73] J.E. Chung, M. Yokayama, M. Yamato, T. Aoyagi, Y. Sakurai, T. Okano, J. Control. Release 62 (1999) 115–127.

[74] J.E. Chung, M. Yokayama, T. Okano, J. Control. Release 65 (2000) 93–103. [75] M. Nakayama, T. Okano, T. Miyazaki, F. Kohori, K. Sakai, M. Yokoyama, J. Control. Release 115 (2006) 46–56. [76] M. Nakayama, J.E. Chung, T. Miyazaki, M. Yokoyama, K. Sakai, T. Okano, React. Funct. Polym. 67 (2007) 1398–1407. [77] J. Akimoto, M. Nakayama, K. Sakai, T. Okano, Biomacromolecules 10 (2009) 1331–1336. [78] K. Matyjaszewski, J. Xia, Chem. Rev. 101 (2001) 2921–2990. [79] M.H. Stenzel, Chem. Commun. (2008) 3486–3503. [80] Y. Nagasaki, K. Yasugi, Y. Yamamoto, A. Harada, K. Kataoka, Biomacromolecules 2 (2001) 1067–1070. [81] N. Nasongkla, X. Shuai, H. Ai, B.D. Weinberg, J. Pink, D.A. Boothman, J. Gao, Angew. Chem. Int. Ed. 43 (2004) 6323–6327. [82] M. Nakayama, T. Okano, Biomacromolecules 6 (2005) 2320–2327. [83] C.L. McCormick, A.B. Lowe, Accounts. Chem. Res. 37 (2004) 312–325. [84] C.W. Scales, A.J. Convertine, C.L. McCormick, Biomacromolecules 7 (2006) 1389–1392. [85] M. Nakayama, T. Okano, Macromolecules 41 (2008) 504–507.