Abstract
A diverse range of drug delivery vehicles have been developed to specifically target chemotherapeutics to solid tumours while avoiding systemic dose-limiting toxicity. Many of these active targeting strategies display limited efficacy because they rely on subtle differences in expression patterns between pathogenic tissue and healthy tissue. In contrast, drug delivery systems that exploit thermoresponsive behaviour allow a clinician to spatially and temporally control the accumulation and/or release of the toxic agents within tumour tissue by simply applying mild hyperthermia (defined as 39–43 °C) to the desired site. Although thermally sensitive materials comprise a significant portion of the literature on novel drug delivery systems, only a few systems have been methodically tuned to respond within this narrowly defined physiological temperature range in an in vivo environment. This review discusses the materials and strategies developed to control the primary tumour through the combined application of hyperthermia and chemotherapy.
Introduction
Chemotherapy has historically been limited by significant off-target toxicity in healthy organs and tissues. Many engineered drug delivery systems seek to minimise these dose-limiting side effects by increasing the fraction of drug delivered to the diseased tissue while reducing harmful interactions with healthy tissue. To this end, effective delivery systems typically display several of the following characteristics: they (1) increase the effective molecular weight and solubility of the drug through physical encapsulation or chemical conjugation, (2) reduce the rate of drug clearance from the plasma, as well as protect the drug from deactivation by plasma proteases and sequestration by the reticuloendothelial system (RES), (3) provide a passive or active mechanism to localise to the tumour site, and (4) release a high concentration of active drug within the tumour or tumour vasculature.
Drug carriers commonly exploit both passive and active targeting strategies to increase the bioavailability of the drug within the tumour. Passive targeting utilises the size of the carrier to passively accumulate in the tumour via the enhanced permeability and retention (EPR) effect [Citation1]. The EPR effect stems from the rapid recruitment of tortuous and hyperpermeable tumour vasculature and the formation of an impaired lymphatic drainage system, which result in the extravasation of sub-100 nm nanoparticles into the tumour and prevent their clearance. Consequently, the treatment efficacy is tied to the carrier size and the length of time the carrier is exposed to the tumour (i.e. the plasma half-life). Active targeting approaches, on the other hand, typically involve decorating the outer surface of the nanoparticle carrier with tumour-specific ligands such as antibodies [Citation2] and aptamers [Citation3] that tightly bind receptors overexpressed by the tumour. While active affinity targeting remains the focus of many delivery strategies, it is limited by the heterogeneity of receptor expression between tumours [Citation4] and even between patients diagnosed with the same cancer [Citation5].
Thermoresponsive drug delivery systems are capable of exploiting an alternative active targeting approach that relies on the localised application of mild hyperthermia to spatially and temporally control the accumulation of a chemotherapeutic agent within a solid tumour. Despite their distinct features, most of these systems are designed to respond to hyperthermia in one of two ways: (1) carriers that physically encapsulate drugs, such as thermally sensitive liposomes [Citation6,Citation7], polymer micelles [Citation8], and hydrogels [Citation9,Citation10] typically release their therapeutic payload upon heating. These carriers are subdivided into fast release vehicles (release in seconds to minutes) that exhibit complete drug release within the first pass through the tumour vasculature, and slow release vehicles (minutes to hours) that exploit the EPR effect to accumulate within the tumour over a period of 24 h and then release the drug within the tumour extravascular space upon heating, and (2) carriers that are physically conjugated to a drug, such as polypeptide–drug conjugates [Citation11] and polypeptide micelles [Citation12,Citation13], undergo a morphological or physicochemical change in response to heat to induce the accumulation of the carrier in the tumour extravascular space. These carriers typically rely on conventional cleavage mechanisms such as hydrolysis or pH to release the drug upon cellular uptake [Citation13,Citation14]. Both targeted accumulation and triggered release are capable of increasing the bioavailability of the chemotherapeutic within the tumour, potentially reducing the tumour burden prior to resection, or even abolishing the tumour entirely. Hyperthermia has also been shown to synergistically boost the effectiveness of chemotherapy by increasing vascular permeability, decreasing drug resistance, and interfering with cell repair mechanisms. There is a significant amount of research being performed in the field of thermally sensitive polymers and delivery systems, which has been reviewed thoroughly [Citation15–18]. This review will focus on the population of carriers that have been tuned to respond to the narrow and physiologically relevant temperature range of 39–43 °C achievable by clinical hyperthermia.
Hyperthermia
The local application of mild hyperthermia (temperatures ranging from 39–43 °C) is a powerful adjunctive therapy in the clinic when used in combination with radio- and chemotherapy for the local control of a variety of cancers [Citation19–22]. Heating tumours to supraphysiological temperatures is not only cytotoxic, but it also activates several mechanisms by which tumour cells are sensitised to the primary treatment. For instance, several phase III clinical trials spanning multiple tumour types (cancers of the breast, head and neck, oesophagus, and cervix, as well as melanomas and glioblastomas) indicate that hyperthermia in combination with radiotherapy provides substantial benefits over radiotherapy alone without significantly increasing toxicity or tissue damage to the surrounding region [Citation23]. This combined effect can be attributed to the fact that hyperthermia inhibits DNA repair pathways [Citation24,Citation25], reduces the severity of radio-resistant hypoxic regions throughout the tumour by enhancing tumour oxygenation [Citation26,Citation27], and is particularly cytotoxic to cells in S-phase [Citation28,Citation29], which display the highest resistance to radiation. Similarly, the efficacy of many chemotherapeutics is also improved at elevated temperatures, though the degree of enhancement varies widely between different classes of drugs [Citation30–32].
Polypeptide–drug conjugates as novel heat-triggered drug delivery vehicles
Elastin-like polypeptides (ELPs) are a genetically encoded class of biopolymers that consist of repeats of the pentameric Val-Pro-Gly-Xaa-Gly peptide found in the sequence of mammalian tropoelastin [Citation33,Citation34]. ELPs display a characteristic phase transition similar to the lower critical solution temperature (LCST) observed in some polymer systems in which the ELP is soluble at low temperatures but phase separates into a polymer-rich gel-like phase termed the coacervate above a critical transition temperature (Tt) [Citation35]. The Tt is most commonly tuned by adjusting the composition of the guest residue position (Xaa) [Citation13,Citation36], where hydrophobic amino acids such as valine and leucine reduce the Tt, and hydrophilic amino acids such as glycine or serine increase the Tt [Citation37]. However, several other factors also influence the transition temperature, including the chain length of the polypeptide [Citation36], the solution concentration, the ionic strength and polarity of the solvent [Citation38], and the presence of any fused molecules [Citation13,Citation14], peptides, or proteins [Citation39].
ELPs provide a diverse range of attributes that make them attractive as ‘smart’ drug delivery vehicles. They are intrinsically non-toxic [Citation40] and biodegradable [Citation41,Citation42]. Their recombinant design allows exquisite control over the composition of the polypeptide, the number and distribution of reactive sites along the polypeptide chain, and the biopolymer molecular weight, which is a critical parameter for controlling their architecture, stability, and functional performance [Citation43,Citation44]. Furthermore, unlike synthetic polymers that exhibit a distribution of molecular weights, ELPs – like all recombinant peptide polymers – are monodisperse. Finally, ELPs are produced in high yield in Escherichia coli and are easily purified by exploiting their inverse phase transition behaviour [Citation45].
As elastin-like polypeptides retain their thermal responsiveness following chemical conjugation, ELP–drug conjugates have been rationally designed by optimising their sequence and molecular weight to exhibit a phase transition between 39 °C and 42 °C, a range that is achievable through the application of mild clinical hyperthermia. Triggering the phase transition through the application of mild hyperthermia increased cellular uptake of an ELP conjugated to a fluorescent reporter in three different human carcinoma lines in vitro (ovarian carcinoma SKOV-3, squamous cell carcinoma FaDu, and cervical adenocarcinoma HeLa) compared to a soluble ELP control [Citation46], an effect likely mediated by enhanced interactions between the phase-separated ELPs and the phospholipid membrane of the cells. Meyer et al. observed that the ELP transition could increase total uptake in human ovarian tumours (SKOV-3) implanted within a dorsal skin-fold tumour window chamber model. The window chamber allows real-time fluorescent measurements of the vasculature and extravascular space of an implanted tumour. By heating the window chamber to 42 °C, they observed a 2-fold increase in the intratumoral accumulation of a thermoresponsive ELP conjugate over the course of 1 h when compared to a non-heated control [Citation47].
To further increase the accumulation of ELP–drug conjugates within solid tumours, Dreher et al. proposed a cyclical heating regimen [Citation11]. Applying localised mild hyperthermia to a window chamber tumour model induced the formation of small ELP aggregates that adhered to the vascular walls, increasing in number and size over time (). Upon cessation of hyperthermia, the rapid solubilisation of the ELP aggregates created a steep transvascular gradient that ‘pumped’ the ELP into the tumour across the interstitial fluid pressure gradient. By measuring the fluorescence as a function of time, the authors observed that the thermally sensitive ELP displayed a 2.8-fold concentration increase over a non-heated control and a 1.6-fold increase over a heated control that lacked the phase transition ().
Figure 1. ELP response to clinical hyperthermia visualised with a dorsal skin-fold tumour window chamber. A thermally sensitive (green) and a thermally insensitive (red) ELP in a solid tumour before, during, and following hyperthermia treatment. (A) Prior to heating, the green and red levels were normalised to produce a uniform yellow throughout the vasculature. Between (B) 10 min and (C) 30 min of heating, the thermally sensitive ELPs began to adhere to the vasculature walls, indicated by the green punctate fluorescence. (D) Upon return to normothermia, the aggregates rapidly resolubilised and dissipated, demonstrating the reversibility of the ELP transition. The scale bar represents 100 µm in all images [Citation11].
![Figure 1. ELP response to clinical hyperthermia visualised with a dorsal skin-fold tumour window chamber. A thermally sensitive (green) and a thermally insensitive (red) ELP in a solid tumour before, during, and following hyperthermia treatment. (A) Prior to heating, the green and red levels were normalised to produce a uniform yellow throughout the vasculature. Between (B) 10 min and (C) 30 min of heating, the thermally sensitive ELPs began to adhere to the vasculature walls, indicated by the green punctate fluorescence. (D) Upon return to normothermia, the aggregates rapidly resolubilised and dissipated, demonstrating the reversibility of the ELP transition. The scale bar represents 100 µm in all images [Citation11].](/cms/asset/60ebc1f5-9274-4033-9bd7-597c5b0702fb/ihyt_a_819999_f0001_b.jpg)
Figure 2. Extravascular accumulation of a thermally sensitive ELP in a heated tumour as a function of time. Data were normalised by the initial vascular intensity and expressed as a percentage of vascular intensity at t = 0 min. The tumour was heated to 41.5 °C for 45 min and then cooled to 37 °C for 15 min for both the thermally sensitive (TS, square) and thermally insensitive (TI, circle) hyperthermia conditions. The tumour was not heated for the 37 °C TS ELP control (diamond). The data are expressed as mean ± SE. *p < 0.05, Fischer’s Protected Least Significant Difference post-hoc test for TS ELP with hyperthermia versus TI ELP with hyperthermia [Citation11].
![Figure 2. Extravascular accumulation of a thermally sensitive ELP in a heated tumour as a function of time. Data were normalised by the initial vascular intensity and expressed as a percentage of vascular intensity at t = 0 min. The tumour was heated to 41.5 °C for 45 min and then cooled to 37 °C for 15 min for both the thermally sensitive (TS, square) and thermally insensitive (TI, circle) hyperthermia conditions. The tumour was not heated for the 37 °C TS ELP control (diamond). The data are expressed as mean ± SE. *p < 0.05, Fischer’s Protected Least Significant Difference post-hoc test for TS ELP with hyperthermia versus TI ELP with hyperthermia [Citation11].](/cms/asset/b1d4bbd7-e6cd-4135-824d-1532208db26a/ihyt_a_819999_f0002_b.jpg)
This technique was also used in vivo to induce tumour regression in a murine E0771 breast cancer model (). Thermally sensitive ELP–doxorubicin (DOX) conjugates were fused to the terminal cell-penetrating peptide (CPP) SynB1 to enhance cellular uptake in both tumour and vasculature cells. The DOX moiety was attached through a pH-sensitive hydrazone linker that rapidly cleaves in the low pH environment of the late endosome [Citation48,Citation49]. The authors observed that thermally cycling (20 min heat; 10 min cool; 4 cycles) the tumour following each of the four SynB1–ELP–DOX administrations (at 2-day intervals) resulted in a statistically significant 4-fold reduction in tumour volume by day 14 over normothermia treatment [Citation50].
Figure 3. Tumour regression in E0771 breast cancer model. E0771 cells were injected subcutaneously into the mammary fat pad of C57BL/6 mice. The mice were treated on day 0 (tumour volume = 150 mm3) with saline, free DOX, or SynB1–ELP–DOX with and without heat. Arrows indicate treatment on days 0, 2, 4, and 6. **A statistical difference was observed between SynB1–ELP–DOX hyperthermia group and saline (p < 0.0001), SynB1–ELP–DOX without hyperthermia (p < 0.0001), and free DOX without hyperthermia (p < 0.001) between day 7 and day 14 [Citation50].
![Figure 3. Tumour regression in E0771 breast cancer model. E0771 cells were injected subcutaneously into the mammary fat pad of C57BL/6 mice. The mice were treated on day 0 (tumour volume = 150 mm3) with saline, free DOX, or SynB1–ELP–DOX with and without heat. Arrows indicate treatment on days 0, 2, 4, and 6. **A statistical difference was observed between SynB1–ELP–DOX hyperthermia group and saline (p < 0.0001), SynB1–ELP–DOX without hyperthermia (p < 0.0001), and free DOX without hyperthermia (p < 0.001) between day 7 and day 14 [Citation50].](/cms/asset/bcc3cb96-cbc6-4fc6-9b15-4ded2c2a6e89/ihyt_a_819999_f0003_b.jpg)
While the delivery strategies described above rely solely on the application of mild hyperthermia to the tumour site, the diversity in structure and composition available with polypeptide carriers enables elastin-like polypeptide conjugates to utilise a vast array of ‘smart’ triggers in conjunction with heat, including pH-sensitivity [Citation51], ligand binding [Citation52], metal ion concentration [Citation53], and light [Citation54].
Liposomes
Since their debut in 1965 [Citation55], liposomes have comprised a substantial fraction of known drug delivery vehicles and currently represent one of the most extensively studied systems in the literature. Liposomes consist of a stable phospholipid bilayer encompassing an aqueous core. The aqueous compartment is capable of encapsulating water-soluble therapeutics whereas the phospholipid membrane can solubilise lipophilic molecules. Incorporation of small molecule therapeutics into liposomes improves drug performance by extending the plasma half-life [Citation56], increasing the solubility of the drug, enhancing accumulation in solid tumours via the EPR effect [Citation57,Citation58], and shielding the drug from deactivating plasma proteases [Citation59]. These benefits are also reflected in the clinic where there are currently multiple systems in clinical trials and on the market, including the chemotherapeutics Doxil and Caelyx (pegylated liposomal doxorubicin) and DaunoXome (liposomal daunarubicin) [Citation60]. Despite the significant increase in tumour accumulation, these formulations continue to display poor drug bioavailability because they suffer from slow and incomplete drug release. Furthermore, the extended plasma half-life of the drug combined with non-specific targeting can result in a number of systemic side effects such as hand and foot syndrome [Citation61]. Using clinical hyperthermia to trigger thermoresponsive liposomal drug release has emerged as a viable strategy to address these issues and allows the rapid release of the entire drug payload within the tumour vasculature. As the field of thermally sensitive liposomes is well developed, we refer the reader to several very thorough reviews on this subject [Citation62,Citation63].
Micelles
Diblock polymers, exhibiting a large solubility difference between the two blocks, form core shell micelles above their characteristic critical micelle concentration (CMC) to reduce the unfavourable interactions between the hydrophobic block and the aqueous environment. The spontaneous self-assembly of micelles creates a hydrophobic, water-excluding core stabilised by a soluble corona. First popularised by Kataoka and coworkers in 1990 [Citation64], drug-encapsulating micelles remain a common therapeutic delivery vehicle because they exhibit many desirable characteristics for treating solid tumours. The micellar core provides a high loading capacity for small and lipophilic molecules (i.e., most of the chemotherapeutic drugs on the market), while the hydrophilic corona –when comprised of specific polymers – can extend the plasma half-life of the vehicle by evading uptake by the RES [Citation65]. The sub-100 nm size and narrow size distribution is ideal for accumulating within the tumour interstitium by the EPR effect and avoiding renal filtration. Actively targeted micelles – whether through the addition of targeting ligands or the localised application of hyperthermia or magnetic fields – represent a new class of drug delivery vehicles that can increase the therapeutic payload reaching the tumour while decreasing non-specific uptake in healthy tissues. To this end, micelles have been modified through the incorporation of thermally sensitive polymers, resulting in two basic classes of thermosensitive micelles that can be actively targeted to a solid tumour: micelles with thermoresponsive coronas and micelles with thermoresponsive cores.
Micelles with thermosensitive coronas
The most commonly used polymer in the design of thermosensitive coronas is poly(N-isopropylacrylamide) (pNIPAAm). This synthetic polymer displays a distinct lower critical solution temperature (LCST) at 32 °C; pNIPAAm is a soluble polymer below this temperature but undergoes a coil-to-globule phase transition above 32 °C that makes it insoluble. The LCST phase transition is thermally reversible. The LCST is frequently tuned by copolymerising pNIPAAm with a hydrophilic polymer to increase the LCST or a hydrophobic polymer to decrease the LCST [Citation66]. Despite the large compositional diversity of thermally sensitive micelles in the literature, there are only a few examples that are responsive within the physiological range of 39–43 °C () [Citation8,Citation67–73].
Table I. Thermosensitive therapeutic micelles that respond to clinical hyperthermia.
In general, these micelles encapsulate their payload at physiological temperatures (37 °C) and accumulate within the tumour via the EPR effect (). The subsequent application of hyperthermia to the tumour tissue induces two effects that enhance tumour drug uptake. The first is that the thermoresponsive corona (formed by copolymerising pNIPAAm with acrylamide or dimethylacrylamide) collapses into a hydrophobic aggregate, which destabilises the micelle structure and drastically enhances the rate of drug release within the tumour interstitium (). illustrates that three different p(NIPAAm-co-DMAAm)-PLGA micelle formulations release DOX slowly and incompletely at 37 °C, a temperature below the transition temperature, but rapidly release DOX at and above the transition temperature of 39.5 °C [Citation71]. The release of free drug from the micelles allows the therapeutic drug to penetrate deeply into the tumour due to its lower molecular weight and higher diffusivity. The second effect is that the collapsed hydrophobic corona enhances intracellular uptake by interacting with the cellular membrane. illustrates this point by suggesting that intracellular uptake of p(NIPAAm-co-DMAAm)-PDLLA micelles is negligible unless the temperature is above the 39 °C transition temperature [Citation73].
Figure 4. Diblock micelles with thermosensitive coronas. (A) At temperatures below the transition temperature of the thermosensitive polymer (orange; pNIPAAm), the hydrophilic thermosensitive corona and the hydrophobic core (blue; poly(ε-caprolactone)) spontaneously assemble into micelles that can encapsulate drugs. (B) At temperatures above the LCST of pNIPAAm, the corona collapses into a hydrophobic coacervate, thereby destabilising the structure and releasing the drugs.
Figure 5. (A) Release profiles of DOX from DOX-loaded p(NIPAAm-co-DMAAm)-PLGA micelles at 37 °C (open markers; below the LCST) and 39.5 °C (closed markers; above the LCST) as a function of time [71]. (B) Intracellular uptake of p(NIPAAm-co-DMAAm)-PDLLA micelles labelled with a fluorophore as the temperature is cycled between 37 ° and 42 °C. Data represents mean ± SD. Reprinted with permission from J. Akimoto, M. Nakayama, K. Sakai, T. Okano, Mol Pharmaceut 2010;7:926–935 [Citation73]. Copyright 2010 American Chemical Society.
![Figure 5. (A) Release profiles of DOX from DOX-loaded p(NIPAAm-co-DMAAm)-PLGA micelles at 37 °C (open markers; below the LCST) and 39.5 °C (closed markers; above the LCST) as a function of time [71]. (B) Intracellular uptake of p(NIPAAm-co-DMAAm)-PDLLA micelles labelled with a fluorophore as the temperature is cycled between 37 ° and 42 °C. Data represents mean ± SD. Reprinted with permission from J. Akimoto, M. Nakayama, K. Sakai, T. Okano, Mol Pharmaceut 2010;7:926–935 [Citation73]. Copyright 2010 American Chemical Society.](/cms/asset/1be010c8-9c37-46f9-a323-0d5dd806f333/ihyt_a_819999_f0005_b.jpg)
Polypeptide micelles with thermally sensitive coronas have also been designed using chimeric polypeptides. CPs consist of two primary components: (1) a thermally responsive elastin-like polypeptide, and (2) a cysteine-rich peptide fused to the C-terminus of the ELP domain to which small molecule therapeutics can be chemically conjugated. When the conjugated molecules are sufficiently hydrophobic, the CP spontaneously self-assembles into star-like micelles that retain the thermal responsiveness of the parent ELP in that they undergo a phase transition from soluble nanoparticles to a hydrophobic coacervate above their transition temperature. The thermal transition characteristics of these CP micelles deviate from those of the elastin-like polypeptide drug conjugates in two important ways. First, the Tt of micelles formed by conjugation of small hydrophobic molecules to a CP is independent of the structure of the conjugated molecule. This allows a large range of therapeutic molecules to be conjugated to a CP carrier without regard to how the molecule will influence the transition temperature. In contrast, ELP small molecule conjugates that do not form micelles are sensitive to even small changes in the structure of conjugated moieties and can display unpredictable fluctuations in their transition temperature [Citation14]. Second, the transition temperature of CP small molecule conjugate micelles is nearly independent of the CP concentration above their CMC (). In stark contrast, the Tt of unconjugated CPs displays a logarithmic dependence as a function of CP concentration (). The dependence of the Tt on concentration has an important consequence for thermally targeted drug delivery because in vivo, concentration and time are tightly coupled. Upon in vivo injection, ELPs and unconjugated CPs are, like all soluble polymers, rapidly diluted in circulation due to tissue uptake and renal clearance. Because their Tt is an inverse function of their plasma concentration, the time window over which they remain within the range of concentration that corresponds to a Tt that is between body temperature and 42 °C is rather short, so that hyperthermia is only effective for a short duration of time. In contrast, the self-assembly of CPs into micelles dramatically flattens the Tt versus concentration profile so that even upon significant dilution in vivo after administration, the micelles will undergo their phase transition from micelles to aggregates over a much wider concentration range and hence temporal widow [Citation13]. This larger temporal window over which CP–drug micelles are responsive to mild hyperthermia makes it possible to design CP–drug micelles that can respond to mild hyperthermia in a clinically relevant setting.
Figure 6. Thermal properties of micelles of chimeric polypeptide-doxorubicin (CP-DOX) conjugates. (A) Relationship between the inverse transition temperature of a CP-DOX micelle (red squares) and an unconjugated CP of the same composition (blue circles). The inset displays a magnified view of the low concentration regime for each construct. (B) The CP-DOX micelle (red circles) transition temperature shows minimal dependence on the concentration, and thus remains within the hyperthermia window (solid black lines) over a wide concentration regime, whereas the ELP monomer used in previous hyperthermia studies (blue triangles) shows a strong concentration dependence [Citation13].
![Figure 6. Thermal properties of micelles of chimeric polypeptide-doxorubicin (CP-DOX) conjugates. (A) Relationship between the inverse transition temperature of a CP-DOX micelle (red squares) and an unconjugated CP of the same composition (blue circles). The inset displays a magnified view of the low concentration regime for each construct. (B) The CP-DOX micelle (red circles) transition temperature shows minimal dependence on the concentration, and thus remains within the hyperthermia window (solid black lines) over a wide concentration regime, whereas the ELP monomer used in previous hyperthermia studies (blue triangles) shows a strong concentration dependence [Citation13].](/cms/asset/c021f405-1487-4b91-be3d-73c78568906b/ihyt_a_819999_f0006_b.jpg)
These two characteristics provide the foundation for the development of a thermally sensitive CP-DOX micelle formulation that transitions between 38–42 °C in a physiological milieu over a concentration range spanning two orders of magnitude (). Although the ability to thermally target this system has not yet been assessed in vivo, it represents a robust strategy to impart a precise thermal trigger to a large variety of hydrophobic therapeutics.
Micelles with thermally sensitive cores
Thermally sensitive micelles can also be designed with the thermoresponsive polymer in the core. In contrast to micelles that disassemble and release their payload in response to hyperthermia, micelles that contain thermoresponsive cores spontaneously assemble when heated. This triggered assembly is uniquely suited to allow the locoregional modulation of ligand density in response to heat [Citation12,Citation74,Citation75]. In a recent example of the utility of this approach, we have synthesised diblock ELPs that self-assemble into monodisperse micelles in response to a thermal trigger, wherein self-assembly in the narrow temperature range between 37 °C (normal body temperature) and 42 °C (highest temperature approved for mild clinical hyperthermia) results in the presentation of a functional cell-penetrating peptide (CPP) motif on a ELP nanoparticle [Citation12]. In this system, nanoparticles were assembled by fusing two ELPs: one with a hydrophilic composition and one with a hydrophobic composition, resulting in a diblock ELP (ELPBC) that exhibited two independent LCST transitions (Tt1 = 39 °C and Tt2 = 56 °C). At systemic physiological temperatures (37 °C) the ELPBC was in a unimer state – soluble polymer chains – that display five arginine (Arg) residues on one end of the polymer chain (). We selected five Arg residues because previous studies have shown that there is minimum threshold of six consecutive Arg residues required to create a functional CPP [Citation76], so we hypothesised that at 37 °C where the ELPBC is in a unimer state, cell uptake should be low. When heated to the range of clinical hyperthermia (42 °C) the hydrophobic ELP (Tt1 = 39 °C) collapsed into a hydrophobic core, whereas the hydrophilic ELP formed a soluble corona decorated with a high local density of Arg (). This high density presentation of arginine in the heated Arg5–ELPBC resulted in an 8-fold increase in cellular uptake in vitro compared to the 37 °C control (). Uptake was negligible for ELPBCs without fused CPPs () and CPP–ELPs that did not self-assemble (). The presentation of a functional CPP only at 42 °C, but not at 37 °C opens the way to convert CPPs which are powerful yet promiscuous agents to promote the uptake of drugs into cells, into an exquisitely targeted system for cancer drug delivery via the application of focused external hyperthermia to solid tumours. While it remains to be seen whether this strong enhancement of cellular uptake in vitro translates in vivo, it represents a potentially powerful strategy to overcome the problem of non-specific targeting of CPPs and promote the tissue-specific delivery of therapeutics.
Figure 7. Diblock micelles with thermosensitive cores. (A) A diblock ELP construct that consists of a conjugated drug (red circle), a hydrophobic block with a low Tt (blue), a hydrophilic block with a high Tt (orange), and a dysfunctional cell-penetrating peptide (CPP; green triangle). At physiological temperatures, the diblock is a soluble chain and the low arginine (Arg) density on the terminus of the ELPBC does not promote cellular uptake. (B) At ∼40 °C, desolvation of the hydrophobic block results in the assembly of spherical micelles that display a high density of Arg residues, thus creating a functional CPP motif on the surface of the nanoparticle. This results in the enhanced uptake of the nanoparticles by cells heated to the clinically relevant mild hyperthermia temperature of 42 °C.
Figure 8. Modulating cellular uptake of thermosensitive ELPBC micelles. At 37 °C, (A) Arg5-ELPBC, (B) ELPBC, and (C) Arg5-ELP exist as unimers and do not exhibit significant cellular uptake. (D) At 42 °C, the temperature-triggered high-density presentation of Arg5 on the micellar surface results in enhanced cellular uptake (green). (E) The ELPBC micelle control lacking the Arg5 and (F) the Arg5-ELP unimer controls do not show enhanced uptake at 42 °C. Green, ELP; red, cellular membrane; blue, cell nuclei; and scale bars represent 25 µm [Citation12].
![Figure 8. Modulating cellular uptake of thermosensitive ELPBC micelles. At 37 °C, (A) Arg5-ELPBC, (B) ELPBC, and (C) Arg5-ELP exist as unimers and do not exhibit significant cellular uptake. (D) At 42 °C, the temperature-triggered high-density presentation of Arg5 on the micellar surface results in enhanced cellular uptake (green). (E) The ELPBC micelle control lacking the Arg5 and (F) the Arg5-ELP unimer controls do not show enhanced uptake at 42 °C. Green, ELP; red, cellular membrane; blue, cell nuclei; and scale bars represent 25 µm [Citation12].](/cms/asset/1ba220e7-c866-4ac3-ac2f-8d9896f96acd/ihyt_a_819999_f0008_b.jpg)
Branched structures
Dendrimers are a specialised class of synthetic polymers that consist of a spherical topology with a well-defined tree-like architecture. Polymer branches grow outward from a central core in a stepwise fashion such that each new branch represents a new generation. Thus, dendrimers are commonly described according to the number of generations, or branches, that they contain, which also correspond to their macromolecular size. In contrast to polymeric micelles, dendrimers are single molecules. Their interior contains numerous nanoscopic cavities that are capable of hosting guest molecules [Citation77]. The specific environment of these cavities can be tuned to favour certain molecules by modifying the terminal end groups of the central polymer [Citation78]. Dendrimers also offer a small size distribution relative to multimolecular micelles, can be easily functionalised, and are stable at very high dilutions. As they do not display a critical aggregation concentration (CAC), they do not suffer from structure destabilisation and premature drug release upon systemic administration as do multimolecular micelles.
Stratified dendrimers can be generated by growing different polymer layers off the terminal end groups. Thermally sensitive dendrimers have been created in this manner by decorating the outer layer with pNIPAAm. Increasing the temperature above the LCST of the dendrimer has been shown to increase the release rate of hydrophobic molecules encapsulated in the core [Citation79,Citation80]. These dendrimers theoretically operate in a manner similar to thermosensitive micelles; the dendrimers accumulate in a solid tumour via the EPR effect and are then triggered to release the encapsulated drug by hyperthermia. One example of this mechanism is the hyperbranched block copolymer of H40-poly(ε-caprolactone) and poly(N-isopropylacrylamide-co-acrylamide) that encapsulates paclitaxel [Citation81]. Upon the administration of hyperthermia, the release rate of paclitaxel doubled and the hydrophobicity of the outer layer increased cellular uptake of the polymer.
Surprisingly, it is also possible to design dendrimers that release their payload in response to hyperthermia without triggering an LCST phase change. Chandra et al. demonstrate that an oligo(ethylene glycol)-grafted amidoamine can encapsulate doxorubicin through a series of hydrogen bond interactions [Citation82]. While these bonds result in a stable encapsulation at 37 °C with only a basal 10% release over 24 h, they rapidly break and release doxorubicin when heated to 43 °C leading to 90% release in 2–3 h in vitro. With the continued development of stimuli-responsive branched structures, this thermally triggered release mechanism will undoubtedly become more prominent.
Hydrogels
Polymers can be cross-linked into three-dimensional hydrogel networks that can absorb large quantities of water, ranging from ∼10% to thousands of times their own weight. Hydrogels have become a popular material for drug delivery for several reasons: (1) the cross-links between polymers vastly increase the physiological stability of the gel, (2) hydrogels can be loaded with aqueous drugs or vehicles that are released in a sustained manner via diffusion or matrix degradation, (3) their environmental sensitivity can be precisely controlled through the choice of polymer, and (4) many of the physical properties such as elasticity, degradation rate, swelling ratio, and hydrophilicity can be tailored by altering polymer parameters such as block length, polymer, cross-linker ratio, etc.
Recently, nanohydrogels – hydrogels that range in size from tens to hundreds of nanometers – have begun to receive attention in the field of drug delivery. Similar to many polymer micelles, they are made through common self-assembly techniques such as solvent emulsion, diffusion, and precipitation [Citation83]. Nanohydrogels retain the material benefits of a hydrogel and gain access to the broad range of applications available to nanoparticles; their small size promotes cellular uptake and tumour accumulation while avoiding renal filtration, and the outer shell can be decorated with ligands to target specific tissues and stealth polymers to extend the plasma half-life by evading the RES.
Thermally sensitive nanohydrogels are formed by cross-linking polymers that display LCST behaviour, most commonly p(NIPAAm) copolymers. Nanohydrogels that have been tuned to respond to mild hyperthermia maintain a swelled state below the characteristic transition temperature (37 °C) but experience a volume phase transition when heated above physiological temperatures (42 °C). The hydrophobic collapse of the polymer chains results in a rapid shrinking of the nanoparticle, aggregation, and deposition within the heated tissue. The reduced volume may result in the expulsion of the entrapped water and/or molecules within the plasma during the transit through the tumour and following nanoparticle deposition within the heated tissue.
Two different p(NIPAAm) copolymers have been used in the synthesis of thermally sensitive nanohydrogels. Zhang et al. showed that p(NIPAAm-co-AAm) nanohydrogel particles with a 50 nm diameter and a volume phase transition between 37 and 43 °C were able to deliver a near-infrared fluorophore (NIRF) to a heated tumour [Citation84]. No tumour accumulation was observed for the dye by itself (with hyperthermia) or for the nanohydrogels without hyperthermia. The same group later showed that this methodology could also be used to deliver the chemotherapeutics docetaxel and 5-fluorouracil [Citation10]. In this study, nanohydrogels loaded with docetaxel displayed significant tumour inhibition against the S180 murine sarcoma line (78%) when coupled with mild hyperthermia, compared to nanohydrogels without hyperthermia (49%) and free docetaxel (40%). Peng et al. synthesised p(NIPAAm)-co-poly(2-(dimethylamino)ethyl methacrylate) (PDMAEMA) hydrogel nanoparticles, which experienced a volume phase transition that caused the 140 nm diameter hydrogel particles to shrink to 100 nm at 41 °C in vivo [Citation9]. Loaded with the highly active SN-38 metabolite of the chemotherapeutic, irinotecan, these nanoparticles were used to treat the C26 murine colon carcinoma by spacing five treatments (20 mg/kg) over 15 days. In combination with hyperthermia, these nanoparticles induced a significant decrease in the tumour burden after 30 days post-treatment compared to the nanoparticles without hyperthermia (2.5-fold decrease) and free irinotecan (2.5-fold decrease). The extensive cross-linking found throughout these hydrogel systems provides a high degree of stability against the large mechanical and biological stresses of circulation that simply cannot be matched by classical polymer micelles. This unique characteristic of nanoscale hydrogels, which significantly reduces premature and off-target drug release, will likely become a commonly adopted motif in thermally targeted systems.
Conclusions and future perspectives
Many actively targeted strategies rely on the unique characteristics of tumours, such as drug release in response to tumour pH or ligands that bind up-regulated receptors or enzymes. While these strategies can be effective, these features dramatically vary between tumours and even between patients diagnosed with the same cancer. In contrast, thermoresponsive drug delivery systems in combination with focused mild hyperthermia have the potential to circumvent the limitations of other active targeting approaches. Hyperthermia mediated targeting can remotely trigger the rapid release of therapeutics, induce drug accumulation, and/or enhance cellular uptake in a site-specific manner.
The thermally sensitive delivery systems that have been optimised to exploit this strategy typically display two common traits: (1) at the physiological temperature of 37 °C, the carrier protects the therapeutic from clearance, inactivation, and harmful interactions with healthy tissue, and (2) at temperatures above 39 °C but below 43 °C, the carrier undergoes a significant morphological change that results in rapid drug release or accumulation at the target site. Both of these effects result in increased drug bioavailability at the site of the tumour, thereby addressing the principle limitation of many chemotherapeutic regimens. Although the current state of the art exclusively utilises polymers displaying LCST-type behaviour to release drugs in response to hyperthermia, we anticipate that the use of UCST (upper critical solution temperature) polymers that are insoluble at low temperatures and soluble at high temperatures will soon be integrated into these smart systems. This will open up new applications such as micelles that simply dissociate at high temperatures to release a payload and micelles that invert (the hydrophobic block becomes hydrophilic and the hydrophilic block becomes hydrophobic) past a specific transition temperature to outwardly present a ligand or CPP that was previously protected in the core.
One of the challenges in the field at the synthesis level is the fact that most thermally sensitive systems are not designed to display a morphological transition in the clinically relevant temperature range of 39–43 °C in blood where the many cosolutes present can influence their phase transition behaviour in unanticipated ways. Hence, while many such systems have been reported, very few have been shown to explicitly exhibit their thermal phase transition in plasma or blood. Ideally, an engineered carrier that responds to mild hyperthermia would experience a highly cooperative phase transition (1–2 °C range) in blood and hence have the ability to fully release its payload during its transit through the tumour vasculature. The cooperativity of this transition can be enhanced by minimising the polydispersity of the polymer either by recombinant synthesis of peptide polymers that are completely monodisperse or alternatively using precision polymerisation techniques for the chemical synthesis of polymers with low polydispersity [Citation85–88]. The continued development of thermosensitive polymers and approaches to thermally trigger the targeted in vivo assembly and disassembly of drug delivery systems promises to provide exciting new tools for cancer chemotherapy.
Declaration of interest
The authors report no conflicts of interest. The authors alone are responsible for the content and writing of the paper.
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