Extracellularly activatable nanocarriers for drug delivery to tumors

Abstract

Introduction

Nanoparticles for drug delivery to tumors need to satisfy two seemingly conflicting requirements: they should maintain physical and chemical stability during circulation and be able to interact with target cells and release drug at desired locations with no substantial delay. Unique microenvironment of tumors and externally-applied stimuli provide a useful means to maintain a balance between the two requirements.

Areas covered

We discuss nanoparticulate drug carriers that maintain stable structures in normal conditions but respond to stimuli for spatiotemporal control of drug delivery. We first define the desired effects of extracellular activation of nanoparticles and frequently used stimuli and review examples of extracellularly activated nanoparticles.

Expert opinion

Several challenges remain in developing extracellularly activatable nanoparticles. First, some of the stimuli-responsive NPs undergo incremental changes in response to stimuli, losing circulation stability. Second, the applicability of stimuli in clinical settings is limited due to the occasional occurrence of the activating conditions in normal tissues. Third, the construction of stimuli-responsive nanoparticles involves increasing complexity in nanoparticle structure and production methods. Future efforts are needed to identify new targeting conditions and increase the contrast between activated and non-activated NPs, while keeping the production methods simple and scalable.

1. Introduction

Nanoparticulate drug carriers can offer several features useful for the delivery of chemotherapeutic drugs. For example, nanoparticles (NPs) made of amphiphilic polymers can be used to solubilize hydrophobic drugs in aqueous media. NPs that securely encapsulate a drug can protect it from hydrolytic or enzymatic degradation and the loss of biological activity. NPs with an optimal size can modify tissue distribution of a drug and reduce systemic toxicity [1, 2]. NPs decorated with specific ligands can facilitate cellular uptake of a drug and help bypass drug efflux pumps [3, 4]. In order to realize these potentials it is critical that a NP remain stable during circulation, without interacting with healthy cells, releasing drug, or entering off-target organs and tissues. On the other hand, once a NP manages to arrive at intended targets, the NP should release the encapsulated drug in the vicinity of tumor cells or enters the cells to unload it inside. A balance between the circulation stability and the reactivity in tumors is, therefore, one of the most important properties of an ideal NP. For this purpose, many carriers are developed with materials that form stable, long-circulating NPs and maintain the chemical structures in normal physiological conditions but change their properties by chemical or mechanical stimuli to make the encapsulated drug available to target cells. Various types of stimuli-responsive NPs are extensively reviewed elsewhere [5-10]. In this review, we focus on nanocarriers that are activated in the extracellular matrix (ECM) of tumors to bring about drug release, cellular uptake, or intratumoral transport (Figure 1). We briefly define the desired effects of extracellular NP activation and frequently used stimuli. We then review examples of extracellularly activated NPs or NP-related systems, ending with a discussion about remaining challenges.

Figure 1.

Schematic diagram of NPs activated in the tumoral extracellular matrix in response to internal or external stimuli. Specifically, the diagram illustrates that circulating NPs extravasate at tumors, undergo structural changes to release drug, interact with cells, and/or change the particle size, according to various stimuli, such as light, ultrasound, magnetic field, temperature, hypoxia, low pH, reductive potential, or increased enzyme levels.

2. Desired effects of extracellular activation of NPs

Systemic NP delivery to tumors is a three-step process: blood-borne delivery, extravasation, and then passage through ECM to tumor cells [11]. Extravasation of NPs in tumors occur relatively selectively, due to the difference between normal tissues and tumors in vascular permeability, a common feature shared by many solid tumors – part of the so-called enhanced permeability and retention effect (EPR effect) [12]. The contribution of the EPR effect to the NP delivery to tumors has been well documented in preclinical animal studies, especially in xenograft models [13, 14]. However, the universal utility of the EPR effect in human patients has recently been questioned, due to significant heterogeneity within and between tumor types [15, 16] and the lack of clinical evidence supporting the benefits [13, 17]. Nevertheless, the EPR effect is arguably the most dominant mechanism by which NPs access solid tumors [17]. Once the NPs arrive at tumors, they are expected to distribute evenly in the tumor mass and release the payload either inside or outside of the tumor cells. Extracellular activation of NPs is widely explored to facilitate the post-extravasation events.

2.1 Tumor-specific drug release

Once NPs are introduced to circulation, it typically takes 1-2 days for the NPs to achieve maximum tumor accumulation via the leaky vasculature [18]; therefore, it is critical to keep drug release from NPs in blood to the minimum during this period. On the other hand, excessive attenuation of drug release also negatively influences the therapeutic effect. An ideal NP should keep the drug inside in a normal physiological condition but have a built-in mechanism to trigger drug release in a timely manner at intended targets, such as tumor ECM and/or intracellular organelles.

2.2 Cellular uptake

NPs may release a drug outside the cells or enter the cells and unload the drug at desired intracellular locations [19]. Either scenario will work if the drug can freely traverse the cell membrane. However, there are situations that NPs need to be internalized by the cells before they release the payload: when the drug is unable to cross the cell membrane efficiently (e.g., nucleic acids, peptides, or proteins) or when the drug is readily removed from the cells due to drug efflux pumps in the cell membrane [20, 21]. In these cases, it is often advantageous to encapsulate drug in NPs, as they can help bypass the cellular barriers [22]. To facilitate the cellular uptake of NPs, their surfaces are decorated with cell-interactive ligands, such as small molecules, peptides, antibodies, or nucleic acids, which allow them to enter cells via specialized endocytosis pathways. On the other hand, the ligand-modified NPs face a greater risk of removal by the mononuclear phagocyte system [23, 24]. Therefore, NPs are designed to circulate as ‘stealth’ NPs (surface-protected with hydrophilic polymers to prevent opsonization) but expose the cell-interactive ligands or charges in response to the applied stimuli after they arrive at tumors [25].

2.3 Extracellular particle transport

NPs arriving at tumors are expected to penetrate into the interior of the tumor mass and completely kill the tumor cells. In reality, NP distribution is limited to the periphery of the tumor mass close to the vasculature [26, 27], while central regions of the tumor remain unaffected [28, 29] and become a potential source for tumor relapse or metastasis. Difficulties in NP penetration into tumors stem from at least 2 abnormal features: increased stiffness of tumor ECM [30] and relatively high interstitial fluid pressure (IFP) [31-37]. Approaches to overcome these challenges involve pre- or co-treatment of tumors with enzymes to degrade the ECM [29, 38-41], priming tumors with an apoptotic-inducer [42-45], or employing external stimuli to increase the mobility of NPs in tumors [46] or to disrupt the ECM [47-52]. In recent efforts, various stimuli are used to reduce the particle size, thereby enhancing intratumoral NP distribution.

3. Stimuli

3.1 Internal stimuli

Tumor cells initiate several changes in the stroma to support their growth and progression, creating unique microenvironment distinguished from normal tissues, such as hypoxia, acidity, and overexpression of proteolytic enzymes [53, 54]. Such differences have widely been used to induce tumor-specific activation of NP drug carriers.

3.1.1 Oxygen level

Hypoxia, inadequate oxygen supplies to the interior of tumors, results from fast unorganized expansion of tumors and inadequate vascularization [54-56]. More than half of locally advanced solid tumors have regions of hypoxia, heterogeneously distributed throughout the tumor mass [54]. Hypoxia leads to several changes in cell metabolism and gene regulation, responsible for increasing resistance to chemo- or radiation therapy [57]. Tumor hypoxia induces upregulation of signaling pathways involved in survival of hypoxic cells, such as hypoxia-inducible factors (HIFs), unfolded protein response (UPR), and mammalian target of Rapamycin (mTOR) [57]. While these changes are exploited as direct targets for cancer therapy, tumor hypoxia also takes part in chemical changes serving as molecular cues to activate nanocarriers, such as acidic pH and reductive environment.

3.1.2 pH

Mildly acidic pH of the tumor microenvironment is one of the most widely used features for the extracellular activation of nanocarriers [6]. The reported range of tumor extracellular pH varies with studies: Some report a median value of 7.0 [58], 6.8-7.2 [56] or ~7.03 [59], as compared to 7.4-7.5 in normal tissues. A study on 67 tumor samples from 58 patients revealed that tumor extracellular pH ranged from 5.66 to 7.78 with an average of 7.01 [60]. The acidity of a solid tumor is attributable to metabolic abnormalities in tumor cells, including the high rate of aerobic and anaerobic glycolysis, which leads to accumulation of lactic acid [61, 62], and the increased proton-pump activities in the plasma membrane, which promote the secretion of acidic metabolites to extracellular milieu [61]. Moreover, the acidified tissues do not readily return to neutral pH due to the reduced blood flow in tumors [6]. In designing stimuli-sensitive drug carriers, the acidic pH is used to change the ionization status of the carrier molecules [63-66] or induce cleavage of acid-labile linkers [6, 67]. A challenge in using acidic pH of tumors is the small difference from the normal pH of 7.4, which require high sensitivity of the carrier molecules to pH change [6].

3.1.3 Reductive potential

Difference in reductive potential is typically used for intracellular drug delivery [10]. The inside of the cells has glutathione (GSH) in millimolar range, kept reduced by NADPH and glutathione reductase [68], whereas extracellular GSH concentration is around 10 μM [69, 70]. Such a difference in reductive potential across the cell membrane is useful for intracellular activation of drug carriers, where the carriers with labile linkers like disulfide [71] or dithiobenzyl carbamate [72] are reduced in the cells to release drug and/or undergo matrix degradation [71, 73]. Normal extracellular matrix maintains a relatively more oxidized state than intracellular environment as a function of redox-modulating proteins, extracellular thiol/disulfide couples, and reactive oxygen/nitrogen species that travel across cell membranes [74, 75]. This balance is perturbed in some tumors, resulting in elevation of the extraceullar reductive potential. For example, an aggressive prostate cancer cell line (WPE1-NB26) showed twice as high extracellular GSH/GSSG ratio as that of nonmalignant prostate epithelial cells [74, 75]. In addition, in vivo electron paramagnetic resonance spectroscopy revealed that the GSH level in radiation-induced fibrosarcoma tumors was four times higher than that in normal muscles [76]. In these tumors, the reductive potential may be used for extracellular activation of drug carriers as well.

3.1.4 Enzyme level

Enzymes overexpressed in tumors, such as matrix metalloproteinases (MMPs), constitute another class of stimuli used for extracellular activation of carriers. The expression of these enzymes, which is tightly regulated in normal tissues, is upregulated in invasive tumors due to the increased need for ECM degradation [77, 78]. Cathepsin B, a lysosomal cysteine proteinase, is also overexpressed in the ECM and cell surface of some tumors [79-81].

3.2 External stimuli

While the internal stimuli are very useful for inducing disease-specific activation of drug carriers, not all the diseases have specific internal molecular triggers. In this case, non-invasive external stimuli such as light, ultrasound, magnetic field, and temperature may be employed to attain spatiotemporal control of drug delivery. Advantages of this approach include a high level of control over the duration and extent of stimuli and the possibility of combining multiple stimuli to increase the sensitivity of the system.

3.2.1 Light

Ultraviolet (UV), visible, and near infrared (NIR) lights are widely used as an external stimuli to trigger drug release and structural changes of nanocarriers. Light-stimulated systems are of particular interest because of the non-invasiveness and the ease of controlling the intensity and duration [82]. UV and short visible (<410 nm) lights are used as an energy source to destabilize caged (deactivated) compounds, but their utility is limited to thin objects such as skin surface or external layers of organs due to the short penetration depth [83]. Longer visible and NIR (>650 nm) lights, which can reach deeper tissues on the orders of hundreds of micrometers to centimeters, have thus gained increasing interest for in vivo applications [83-86]. On the other hand, lights with longer wavelengths cannot afford sufficient energy to initiate cleavage of chemical linkers directly. Therefore, long visible or NIR lights are combined with compounds such as gold NPs that absorb the lights and generate heat, which then trigger structural changes in drug carriers [83, 84, 87]. Alternatively, long visible or NIR lights are used with an agent that generate reactive oxygen species (ROS) upon radiation (photosensitizer), which help enhance drug release and/or intracellular trafficking of a drug [83].

3.2.2 Ultrasound

Ultrasound refers to acoustic sound with high frequencies (>20 kHz) above those of audible sound, which penetrate deeper into inner organs than light [5, 88]. Ultrasound generates various effects (heating, acoustic cavitation, and acoustic radiation forces) useful for diagnosis and physical therapy of diseases [89]. These effects are lately used as external stimuli for controlling drug delivery [90]. For example, acoustic cavitation, growth and collapse of microbubbles in blood, induces reversible changes in nanocarriers and trigger drug release [91, 92]. Ultrasound applied at high amplitudes also produces mechanical actions called radiation forces, increasing extravasation and interstitial transport of drug and the carriers [89].

3.2.3 Magnetic field

External magnetic field is used in combination with magnetically-responsive carriers as a way of positioning them in specific organs or tissues and triggering drug release [93, 94]. A magnetic carrier should have superparamagnetism, the ability to strongly magnetize (align all magnetic moments of atoms parallel to the direction of a magnetic field) when exposed to a magnetic field and show no residual magnetization (remanence) upon the removal of the magnetic field [93, 94]. Superparamagentic iron oxide nanoparticles (SPIONs) based on γ-Fe₂O₃, α-Fe₂O₃, and Fe₃O₄ are most commonly used as a magnetic carrier [87]. When used for external control of drug release, SPIONs are incorporated in polymer matrices or liposomes that can be deformed by heating or the movement of the magnetized particles [87].

3.2.4 Temperature

Most of the above-mentioned external stimuli generate mild heat, which provides a useful control over drug release when combined with thermo-sensitive materials [95]. For example, high intensity focused ultrasound (HIFU) can produce local heating by focusing multiple ultrasound waves to deposit a high acoustic intensity in the focal volume [96]. Ideal temperature range for hyperthermia-triggered drug delivery is 41-42°C; above this range, vascular coagulation and tissue damage may occur [95, 97]. For this reason, most thermo-sensitive liposomes are made of dipalmitoylphosphatidylcholine (DPPC), which undergoes phase transition at 41.5 °C [95, 98-101]. For polymeric NP systems, poly(N-isopropylacrylamide) (PNIPAM), which changes hydrophilicity according to the temperature, is typically used as a thermosensitive component. The transition temperature can be controlled by the polymer concentration [102], molecular weight of the polymer [103], and type and content of additional blocks [104, 105]. While thermal stimulus is mainly used to trigger drug release by causing structural changes of the carriers, it also contributes to drug delivery by increasing vascular permeability [97, 106-108] and/or decreasing high interstitial tumor pressure [109], thereby enhancing extravasation and intratumoral transport of NPs.

4. Extracellularly-activatable nanocarriers

Extracellular activation has been employed in various types of NP systems including inorganic or polymeric NPs, liposomes, and dendrimers. Such systems respond via drug release and cellular interactions through different mechanisms. The following section introduces recent examples of extracellularly-activated NPs, classified by the consequences of the stimuli-triggered activations. While drug carriers are the main focus of this review, imaging agents are also mentioned when relevant in principle. The readers are also advised to note that in many cases the NP activation was demonstrated in relatively extreme in vitro conditions (e.g., pH 5.5) than those faced in vivo to represent clear contrasts between the activated and non-activated status.

4.1 Control over drug release

4.1.1 Internal stimuli

pH

For pH-triggered drug release, acid-labile linkers such as hydrazone, acetal, or ester bonds are frequently used, although the triggering pHs for these linkers are somewhat low for extracellular drug release. Polyhistidine (pHis) is another chemical moiety widely used for pH-sensitive drug carriers. The pH sensitivity of pHis comes from the imidazole group, which protonates in acidic pH with a pKa value of ~6 [110]. Polymeric micelles prepared with a block-copolymer of pHis and PEG showed higher drug release at acidic pH as pHis block turned hydrophilic with protonation [110, 111]. More recently, a pHis-based AB₂-miktoarm polymer (mPEG-b-pHis₂) was designed to form polymersomes, thin-walled polymer vesicles similar to liposomes [112]. Below pH 7.4, the polymersomes underwent conformation changes to cylindrical micelles, spherical micelles, and finally to unimers, showing increasing drug release [112].

Alternatively, chitosan and its derivatives are used as a component of NPs for pH-triggered drug release. Magnetic nanocrystals and DOX were encapsulated in micelles made of amphiphilic chitosan derivative, N-naphthyl-O-dimethymaleoyl chitosan (N-nap-O-MalCS) with an average size of 158.8 nm at pH 7.4. Exposure to acidic medium induced hydrolysis of maleoyl group, which caused the loss of amphiphilicity and destabilization of micelle structure, as evidenced by a significant increase of particle size at pH 6.5 [113]. The pH-induced change caused an abrupt DOX release (90% release in 24 hours) at pH 5.5 as compared to 20% at pH 7.4 [113]. In another study, chitosan NPs were used as a pH-sensitive carrier of methotrexate (MTX). MTX-loaded chitosan NPs (MTX-CS-NPs) were prepared by ionic gelation of chitosan via tripolyphosphate and an anionic surfactant (77KL), which has a membrane-lytic activity [114]. MTX release from the NPs increased with pH decrease, due to the protonation of 77KL and MTX, leading to decreased electrostatic interactions with chitosan and destabilization of the NPs. Consequently, MTX-CS-NPs showed enhanced cytotoxic effect on MCF-7 cells at pH 6.6 as compared to at pH 7.4, while free MTX did not show such a pH sensitivity [114].

Enzyme level

When the enzyme level in ECM is used to trigger drug release, enzyme-cleavable peptide substrates are used as a structural component of a nanocarrier [67, 115-117]. MMP-sensitive liposomes were developed using a lipopeptide with a cleavage site for MMP-9 [115]. The lipopeptide was mixed in the lipid bilayer of a liposome generating a triple helical structure, which was destroyed by MMP-9 and caused the release of liposomal contents [115]. A model compound carboxyfluorescein loaded in the liposomes was released according to the concentration of MMP-9 in release medium (40% in 200 nM MMP-9 and 100% in 2.3 μM MMP-9) [115]. Peptide cleavage was MMP-9-specific, with none of MMP-7, MMP-10, or trypsin able to trigger liposomal destabilization and payload release [115].

Reductive potential

Most nanocarriers using the redox potential difference as a trigger of drug release contain disulfide bond, which is cleaved in a relatively reductive environment. While the increased reductive potential is a potentially useful cue for extracellular activation of NPs, most examples in the literature have been evaluated in the context of intracellular drug delivery. Collagen, a natural component of ECM, was conjugated on the MSN surface via disulfide bond to serve as a capping material [118]. The collagen-capped MSNs released only 7% of the encapsulated fluorescein isothiocyanate (FITC) in PBS (pH 7.4) in 3 hours but released additional 80% of FITC after the addition of dithiothreitol (DTT) to 30 mM [118]. It is worthwhile to note that, despite the widespread use as a synthetic substitute for GSH, DTT may not result in similar drug release patterns as GSH, depending on the nature of drug-carrier interactions. Nanohydrogels composed of poly(methacrylic acid) (PMAA) crosslinked via N,N-bis(acryloyl)cystamine were prepared as a redox-sensitive carrier of DOX [119]. With the addition of 10 mM DTT to the medium, DOX release increased from 15% to 27% after 24 hours, whereas with 10 mM GSH drug release dramatically increased it to 80% [119]. This difference is explained by the electrostatic interaction between the hydrogel and DOX, which was effectively displaced by partially protonated GSH but not by DTT.

Hypoxia

2-nitroimidazoles, which undergo selective bioreduction in hypoxic conditions, have been used for the synthesis of hypoxia-activated prodrugs for cancer therapy [57, 120]. Using the same principle, a 2-nitroimidazole was conjugated to a hydrophilic carboxymethyl dextran backbone to produce a hypoxia-sensitive polymeric system [121]. This polymer conjugate formed a NP with a hydrophobic 2-nitroimidazole core, loaded with a base form of DOX. Under hypoxic conditions, 2-nitroimidazole underwent a series of bioreductions to form more hydrophilic 2-aminoimidazole, destabilizing the NP structure and releasing the loaded DOX [121]. In vitro release studies were performed in hypoxic (degassed PBS containing NADPH as electron donor) versus normoxic conditions (PBS containing NADPH with no degassing); after 12 hours, complete DOX release was observed in hypoxic medium, but 49% in normoxic medium. The DOX loaded NPs showed greater inhibition of SCC7 cancer xenograft growth than an equivalent dose of free DOX [121].

4.1.2 External stimuli

Temperature

Thermosensitive liposomes (TSLs) are frequently used for heat-induced drug release control. Several TSLs have been reported using DPPC as a main component. For example, TSLs prepared with DPPC, HSPC (hydrogenated soy sn-glycero-3-phosphocholine), cholesterol, and DSPE-PEG (1,2-distearoyl-sn-glycero-3-phosphoethanolamine-N-PEG 2000) released drug upon heating at 42-45°C for 30 min [122]. Various lipid compositions have been explored to control the thermosensitivity of TSLs. TSLs prepared with DPPC, MSPC (1-stearoyl-2-hydroxy-sn-glycero-3-phosphatidylcholine), DSPE-PEG released drug completely at >40°C in 10 min [123]. TSLs more quickly responding to temperature were prepared by replacing MSPC and DPSE-PEG with a non-ionic surfactant (Brij78) [123]. TSLs responding to a relatively low temperature (39-40 °C) for a short period of time (low TSL) were developed using a mixture of DPPC, MPPC (1-palmitoyl-2-hydroxy-snglycero-3-phosphocholine), and DSPE-PEG [122] and used with HIFU-induced hyperthermia [124]. The Low TSLs showed minimal drug release in 2 min at 37°C but 50% drug release at 42°C during the same period. The Low TSLs exposed to pulsed HIFU (1300 W/cm², 0.1 sec on and 0.9 sec off for 120 times) showed significantly faster and greater drug delivery to tumors as compared to non-TSLs [124]. The growth of tumors in animals treated with repeated IV injections of Low TSLs and HIFU was significantly delayed as compared to no HIFU group [124].

Ultrasound

Acoustic cavitation is used to trigger drug release from nanocarriers. In a recent example, ultrasound was applied to localize drug release from liposomes on tumors [125]. Liposomes were prepared with DSPE, cholesterol, 1,2-distearoyl-sn-glycero-3-phosphocholine (DSPC), and DSPE-PEG, including luciferin as a model active ingredient. The liposomes circulated with no significant leakage of luciferin. When ultrasound was applied to tumors in the presence of SonoVue®, phospholipid-stabilized microbubbles to enhance cavitation effects, the liposomes were destabilized and released luciferin, producing 16 times higher luminescence signal in tumors than non-stimulated ones [125]. An important feature of this approach is that with the aid of SonoVue the ultrasound activation could be carried out under conditions used in diagnostic ultrasound scanners. This is a significant improvement over previous approaches as it alleviates the need to use a high intensity ultrasound wave that may generate damages to non-target tissues [125].

Light

To use NIR as an external stimulus to cause drug release, a reservoir-type drug carrier was developed using a nanocomposite ethylcellulose membrane, containing gold nanoshells and PNIPAM-based thermosensitive nanogels, as a drug diffusion barrier [126]. When triggered by a NIR laser (808 nm), gold nanoshells generated heat beyond a critical temperature of nanogels, causing them to shrink and leave pores in the membrane, through which the encapsulated drug could be released [126]. This system allowed for NIR-triggered release of aspart, a fast-acting insulin analog, from the subcutaneously implanted device, effectively reducing blood glucose over 14 days. On the other hand, a critical challenge is to increase the response temperature substantially higher than 37°C and prevent accidental drug release due to fever or hot weather but to keep it low enough to avoid tissue damages (<43°C [127]).

Magnetic field

The same principle was applied using an oscillating magnetic field as a release trigger [128, 129]. For this purpose, the diffusion barrier was prepared with paramagnetic magnetite NPs as the triggering entity, which generated heat (+2 °C) by an external oscillating magnetic field and induced nanogel shrinkage and drug diffusion from the device [128]. Depending on the phase transition temperature and loading density of nanogels as well as the membrane thickness, this device showed 6-15 times increase in drug release rate by heating and maintained the zero-order release kinetics through the duration of on-state [129].

4.1.3 Multiple stimuli

Several stimuli may be used in combination to facilitate the formation of NPs, increase the flexibility in release control, or increase the selective reactivity of a system to stimuli [5, 130-132]. For example, pH- and temperature-sensitive micelle system was produced using a block copolymer consisting of poly(NIPAM-co-acrylic acid) and polycaprolactone [133]. The hydrophilic block of this polymer, poly(NIPAM-co-acrylic acid), imparted the sensitivity to temperature and pH, based on the phase transition of PNIPAM in increasing temperature and protonation of acrylic acid in acidic pH [133]. Paclitaxel (PTX) encapsulated in this micelle system was released most quickly when both conditions were met [133].

A mesoporous silica material capped with boronic acid-modified gold NPs is another pH-/temperature-sensitive system [134]. A saccharide derivative was anchored on the external surface of mesoporous silica-based material (MCM-41), where multiple alcohol groups of the saccharide derivative had reversible interactions with boronic acid of gold NPs by forming boronate esters [134]. The boroester bonds were hydrolyzed at pH 3 or cleaved thermally by plasmonic heat emitted from NIR-irradiated gold NPs, removing the capping gold NPs and releasing molecules entrapped in the pores of the silica device [134]. While this approach opens up a new possibility of designing stimuli-induced release systems, the present format has a limited utility for drug delivery purposes, due to the low trigger pH (pH 3) and the heat generated by the long wavelength NIR laser itself.

A polythioether ketal-based NP system was produced to activate drug release in response to acidic pH and ROS, which changed hydrophilicity and degradation rate, respectively (Figure 2) [135]. Upon the exposure to ROS, thioether groups in the polymer backbone underwent oxidation to sulfone, turning the polymer from hydrophobic to hydrophilic. On the other hand, ketal groups in the backbone underwent rapid acid-catalyzed hydrolysis in mildly acidic condition [135]. When both conditions were provided, the polythioether ketal-based NPs showed complete degradation and drug release in 24 hours [135]. In contrast, the NPs in neutral pH with no oxidative reagents showed minimal release of model drugs over the same period of time [135].

Figure 2.

(a) Degradation mechanism of polythioether ketal. Hydrogen peroxide and acidic pH stimulate the degradation of the polymeric nanoparticles in tandem. (b) Schematic diagram of polythioether ketal-based NP system which releases drug in response to acidic pH and ROS. Reprinted with permission from [135]. Copyright (2011) American Chemical Society.

An important advantage of multiple-stimuli-sensitive system is the potential to precisely regulate the drug release according to the combination of stimuli. To enhance the precision of control, a polymeric micelle system responding to triple stimuli (temperature, pH, and reductive potential) was developed using a block-copolymer comprising of an acid-sensitive hydrophobic core (poly(hydroxyethyl methacrylate)), temperature-sensitive hydrophilic shell (PNIPAM), and redox-sensitive interface (disulfide linker) [136]. The decrease of pH converted acid-sensitive hydrophobic block to hydrophilic one, temperature increase made PNIPAM hydrophobic, and a reducing environment induced cleavage of block-copolymer to individual homopolymers, all contributing to disassembly of the micelle system. Notably, individual stimulus caused slow or incomplete release of the encapsulated dye, but combined stimuli led to a significantly faster and greater release [136].

4.2 Control over cellular uptake

Promoting selective interaction of nanocarrier systems with tumor cells is achieved by removing protective surface layers from the carriers or transforming the surface properties in a tumor-specific manner. Most nanocarriers are protected by non-ionic hydrophilic polymers to avoid non-specific interactions with immune cells and normal tissues during circulation. Removal or transformation of such a protective surface results in exposure of cationic charges or cell-interactive ligands, thereby allowing for electrostatic or ligand-mediated interactions with the cell membrane. This effect can be achieved by employing a stimulus-sensitive linker as a component of the carrier polymer or using a stimulus-sensitive polymer for a protective surface.

4.2.1 Internal stimuli

pH

To increase cellular uptake at acidic tumoral pH, a polymeric micelle system was developed using a blend of polyhistidine (polyHis)-based amiphiphilic polymers [137]. The micelle was made of a blend of polyHis_5kD-b-PEG and PLA-b-PEG-b-polyHis_2kD-TAT, where polyHis_5kD and PLA blocks from each polymer formed a hydrophobic core and PEG formed a shell. At pH 7.4, PEG from the latter polymer formed a loop as polyHis_2kD block remained unionized and associated with the hydrophobic core, keeping the cell-interactive TAT away from the surface. Below pH 7.2, polyHis_2kD started to ionize, exposing the TAT on the surface to facilitate cellular uptake of the polymeric micelles [137]. Due to the enhanced cellular uptake, the micelles carrying DOX showed a greater cytotoxic effect on drug-resistant NCI/ADR-RES cells upon the acid-triggered activation [137]. A consistent result was observed in a mouse model of a drug-resistant ovarian cancer xenograft [137].

In another study, a layer-by-layer (LbL) approach was used to make multilayered ~80 nm NPs with a fluorescent core and PEG-coated surface [138]. The core particle was a carboxyl-functionalized quantum dot (QD). The negatively charged core particle was first coated with poly-L-lysine (PLL)-iminobiotin conjugate and neutravidin, which was then coated with mPEG-biotin conjugate. Under acidic pH, the iminobiotin-neutravidin linker was decomposed due to the reduced affinity of the protonated iminobiotin for neutravidin. The decomposition of iminobiotin-neutravidin bonds were demonstrated over a range of pH 4-7.4, where the decomposition rate increased linearly with the decrease of pH. This caused the external PEG layer to be shed and a cationic PLL layer exposed, thereby facilitating cellular uptake of QD cores in acidic environment. After incubation in a pH 5.5 medium, NPs showed significantly higher cellular uptake in five different cancer cell lines as compared to those in pH 7.4 [138]. As a consequence of pH-sensitive tumoral uptake, the NPs modified with iminobiotin showed longer retention in tumors than control NPs with biotin [138].

pH-sensitive polymers have been used to form a surface layer that becomes more cell-interactive in tumoral pH. For example, low molecular weight chitosan (LMWC) was used as a pH-sensitive surface coating. Chitosan is a polysaccharide with primary amines, which impart a unique pKa of 5.5-6.5 [139], matching the weakly acidic pH of tumor tissues [63]. Due to the reduced molecular weight (<6500 Da), LMWC remains neutral yet hydrophilic at pH 7.4, thus qualifying for a stealth polymer. LMWC was conjugated to poly(lactic-co-glycolic acid) (PLGA), yielding a polymeric NP with a PLGA core and a LMWC surface [63]. The PLGA-LMWC NPs showed a slightly negative charge at pH 7.4 but acquired a positive charge in acidic pH. Consequently, PLGA-LMWC NPs showed greater interactions with SKOV-3 cells at pH 6.2 than at pH 7.4, whereas the unmodified PLGA NPs showed limited cellular uptake irrespective of the pH [63].

Another example involves a cationic polyamidoamine (PAMAM) dendrimer coated with a zwitterionic chitosan derivative (ZWC) [66]. Amine-terminated PAMAM dendrimer is an attractive carrier of drug and gene therapeutics due to the well-defined structure and functionalization potential; however, the utility is limited because of undesirable cytotoxic effects [140]. Created by partial amidation of chitosan, ZWC showed a negative charge in a relatively basic condition and positive charge in an acidic condition, where the transition pH is readily tunable according to the extent of amidation [141]. The PAMAM dendrimer was electrostatically coated with ZWC, which was anionic in neutral pH, reducing toxicity associated with the cationic charge of the dendrimer and preventing cellular uptake of PAMAM dendrimers (Figure 3). On the other hand, in mildly acidic pH such as pH 6.5 or lower, where ZWC acquired more positive charges, the PAMAM dendrimer was no longer protected and allowed to interact with the cell membrane and enter the cells (Figure 3) [66].

Figure 3.

Cellular uptake of PAMAM or ZWC(PAMAM) at pH 7.4 (left) and pH 6.4 (right). Green: PAMAM dendrimers; blue: nuclei. Adapted with permission from [66]. Fluorescently labeled PAMAM dendrimers (green) appeared in or on the cells irrespective of the pH. In contrast, PAMAM dendrimers coated with ZWC showed minimal cellular interactions at pH 7.4, whereas strong green signals were observed in or on the cells at pH 6.4, where the dendrimers were no longer protected by ZWC. Copyright (2013) American Chemical Society.

Peptides are another class of pH-sensitive materials that can be used to promote pH-sensitive cell interactions [142]. A pH low insertion peptide (pHLIP) is a pH-sensitive peptide made of 38 amino acids with moderate water solubility [142]. As pH drops from 7.4 to 6.5, it becomes more hydrophobic with protonation of Asp and Glu residues and inserts its tail into the cell membrane lipid bilayer, helping NPs modified with the peptide to enter cells [142, 143]. pHLIP was conjugated with MSN, an inorganic drug carrier, via a disulfide bond [143]. When placed in mildly acidic pH, the pHLIP-conjugated MSNs were readily taken up by cells, in which the disulfide linker was reduced and the loaded DOX released. Due to the pH-induced cellular uptake, this system showed greater cytotoxic effects in both drug sensitive (MCF-7) and resistant (MCF-7/ADR) cell lines at pH 6.5 relative to those at pH 7.4 [143]. A similar approach was used to enhance tumor uptake of gold NPs [142].

Enzyme level

Removal of a protective layer can be induced by enzymes abundant in tumor ECM. PLGA NPs were dual-coated with a cell penetrating peptide (TAT peptide) and PEG for promoting cellular uptake and preventing non-specific exposure of the TAT peptide, respectively [67]. The surface PEG was conjugated to the NP surface via a MMP-2 cleavable peptide linker, so that it could be removed in a MMP-2 rich environment such as tumoral ECM, allowing the TAT peptide to promote cellular uptake of NPs. The dual-modified PLGA NPs showed minimal uptake by SKOV-3 cells in MMP-2 free medium but significantly enhanced cellular uptake after treatment with MMP-2. In contrast, NPs with non-cleavable PEG showed minimal cellular uptake, irrespective of the presence of MMP-2. Consistent with the MMP-2-dependent cellular uptake, the dual modified NPs loaded with PTX resulted in a greater cytotoxic effect after MMP-2 treatment than non-treated ones, although the difference was not as clear as the cellular uptake, due to the high initial burst release of the drug [67].

4.2.2 Multiple stimuli

A dual pH-sensitive drug conjugate was developed to produce polymeric NPs that respond to tumor ECM pH (6.5) to facilitate cellular uptake and then to lower lysosomal pH (5-5.5) to enhance intracellular drug release [144]. A DOX-polymer conjugate (PPC-Hyd-DOX-DA) was synthesized by conjugating 3-dimethylmaleic anhydride (DA) and DOX to a block-copolymer of mPEG and cysteamine-modified poly(allyl ethylene phosphate) (mPEG-b-PAEP-Cya, PPC), via amide and hydrazone bond, respectively. Here, the amide bond between with β-carboxylic acid of DA and amino group of cysteamine is cleavable at slightly acidic conditions like pH 6.8 [145], whereas hydrazone bond is cleaved at lower pHs. The polymer conjugate formed self-assembled NPs with an average diameter of ~27 nm that underwent two levels of changes according to the pH [144]. First, the β-carboxylic amide bond between DA and polymer cleaved at pH 6.8, increasing the cationic charge density and, thus, enhancing cellular uptake of NPs. After the uptake and endosomal localization, the hydrazone link between DOX and polymer was cleaved off and the drug was released [144]. Due to the pH-sensitive enhancement of cellular uptake and intracellular drug release, the NPs achieved greater cytotoxicity in drug-resistant cancer stem cells than free DOX at pH 6.8 [144].

Similarly, NPs responding to multiple levels of pH were developed to induce NP-cell interaction and drug release according to the environment. Dual-pH-sensitive polymeric micelles were prepared using two block copolymers, poly(L-histidine)-b-short branched polyethyleneimine (PHis-b-sbPEI) and mPEG-b-polysulfadimethoxine (mPEG-b-PSDM) [64]. PHis-b-sbPEI self-assembled to yield core polymeric micelles, where PHis formed a pH-sensitive hydrophobic core and sbPEI formed a cationic shell. The core micelles were coated with mPEG-b-PSDM via pH-sensitive electrostatic interactions between sbPEI and PSDM. The pH-sensitivity of the shielded micelles came from PSDM, negatively charged at neutral pH but uncharged in slightly acidic conditions. At mildly acidic pH, the micelles lost a mPEG-b-PSDM layer due to the weakening electrostatic interactions with the PEI shell, and the exposed cationic surface then interacted with tumor cells surface. Once taken up by cells, the core PHis-b-sbPEI micelles were destabilized and released the loaded PTX in the low endo/lysosomal pH as the PHis block started to protonate. The shielded micelles were stable in serum due to the protective effect of PEG and showed minimal cellular uptake at pH 7.4 but significantly enhanced uptake at <pH 6.6. When injected IV in MCF-7 tumor bearing mice, the shielded micelles showed superior tumor growth inhibition as compared to free PTX or the unshielded core micelles [64].

In a recent example, pH and ultrasound were used in combination for the delivery of oncolytic adenovirus to tumors [146]. Adenovirus was coated with a new stealth polymer, N-(2-hydroxypropryl) methacrylamide copolymer containing pH-sensitive hydrazone bonds, which degraded at acidic intratumoral pH to restore the ability of virus to bind and infect specific cancer cells. Focused ultrasound was used along with SonoVue to produce an intense cavitation effect, which enhanced penetration of viral particles into the acidic interior of tumors. The combination of ultrasound-mediated tumor penetration and pH-sensitive deshielding of virus led to significant improvement in tumor growth inhibition and survival of tumor-bearing mice as compared to non-ultrasound stimulated ones [146].

4.3 Control over extracellular particle transport

While the approved NP products have particle sizes ranging from 100 nm or higher (e.g., Doxil), recent animal studies find that sub-100 nm sizes are required for effective tumor penetration. Kataoka et al report that only 30 nm polymeric micelles can penetrate poorly permeable pancreatic tumors to achieve an anti-tumor effect [147]. Similarly, Allen et al find that 25 nm, but not 60 nm, polymeric micelles penetrate into breast tumor xenografts [2]. Chan et al also showed using PEGylated gold NPs that smaller particles (20 nm) penetrated better into tumor matrix than larger NPs (40-100 nm) [148]. On the other hand, such a small NP is not necessarily beneficial for the accumulation and retention in tumors [148] and the drug loading capacity [149]. To address these conflicting needs, NPs are engineered to circulate as relatively large particles and be reduced to smaller NPs by internal and external stimuli. Alternatively, stimuli are also used to increase the particle size after arrival at tumors [150]. In this case, NPs are delivered as relatively small particles to take advantage of the EPR effect but swell in the tumoral ECM so that their retention in the intended locations may be improved.

4.3.1 Internal stimuli

Enzyme level

Wong et al. reported a NP system changing the size in tumor ECM according to high MMP levels [11]. The NPs consist of crosslinked gelatin core with an average diameter of 100 nm, on which 10 nm quantum dot (QD) NPs were covalently bound on the surface. The 100 nm size allowed the NPs to circulate and reach tumor via the EPR effect; once in the tumor with an elevated level of MMP-2, the gelatin core degraded to release 10 nm QD NPs that could better penetrate into tumors. When injected intratumorally in mice bearing HT-1080 tumors with high MMP-2 activity, the gelatin-QD NPs showed greater penetration into tumor tissues than silica-QD NPs that did not change the size [11].

4.3.2 External stimuli

Light

To improve tissue penetration and drug release in target tissues, Tong et al developed a lipid-based NP system, composed of DSPE-PEG, lecithin, and a spiropyran-alkyl conjugate (SP-C9), which reduced the average size from 150 to 40 nm upon the exposure to UV light (365 nm) (Figure 4) [151]. The photo-triggered shrinkage resulted from the isomerization of hydrophobic SP to zwitterionic merocyanine (MC) and subsequent movement of MC to hydrophilic PEG layer, which let alkyl chains of DSPE and lecithin form tighter assembly in the hydrophobic core [151]. The shrunken NPs swelled back with the removal of UV as MC reverted to SP and translocated to the core. After photo-triggering, NPs encapsulating fluorescent dyes showed greater penetration into a dense collagen gel and the cornea than free dyes and NPs with no UV trigger [151]. While the use of UV light may limit the utility of this system in drug delivery to tumors, it is conceivable to introduce similar features in other types of NPs to enhance drug delivery into the tumor interior, which can be hard to reach due to the dense interstitial matrix.

Figure 4.

(a) Structure and photoisomerization reaction between spiropyran (SP) and merocyanine (MC). (b) Abbreviations for SP and MC derivatives. (c) Scheme of photoswitching SP NP composed of SP-C9 and DSPE-PEG. Yellow oval, SP molecule: blue line, alkyl chain (R) in SP; red, lipid part; green line, PEG. SP NP are converted to MC NP (purple oval, MC molecule) by UV light irradiation; the reversible photoisomerization from MC NP to SP NP happens in dark but is accelerated by visible light (500-600 nm). (d) Dynamic light scattering measurement of size changes of SP NP composed of SP-C9, DSPE-PEG, and lecithin. Inset: the solution of NPs before and after UV irradiation. Reprinted with permission from [151]. Copyright (2012) American Chemical Society.

Temperature

Thermosensitive composite microparticles were developed to control the surface adhesion of circulating particles by thermal stimulus [152]. The microparticle consisted of a PNIPAM core particle containing SPIONs and a surface stabilized with SiO₂ particles. Heating SPIONs with the magnetic field resulted in a reduction of the particle size, which in turn reduced the drag force on the particles in the flow and helped them remain better adherent to a surface [152]. This principle may be useful for controlling distribution and retention of the circulating nanocarriers.

4.3.3 Multiple stimuli

A dual temperature- and pH-responsive polymeric micelle system based on mPEG-b-P(HPMA-Lac-co-His) was developed to facilitate drug encapsulation and micelle retention in tumors [150]. The hydrophobic segment of this polymer consisted of poly(N-(2-hydroxypropyl) methacrylate dilactate (HPMA-Lac) and pHis, which provided thermosensitivity and pH-sensitivity, respectively [150]. P(HPMA-Lac-co-His) had a low cloud temperature (<10 °C), allowing for efficient micelle formation and drug encapsulation by quick heating [150]. At physiological temperature, particle size gradually increased and drug release occurred due to the accelerated hydrolysis of lactic acid side chains, and further accelerated at relatively acidic pH of tumor ECM [150]. The authors propose that the micelles initially maintain a desirable size (<60 nm) for circulation and tumor accumulation but are better retained in tumors due to the increasing size (>80 nm) in a condition best met by tumor ECM (physiological temperature and acidic pH).

5. Conclusion

Success of NP-based drug delivery depends on the circulation stability of NPs and their ability to deliver drug at the right location and time. Unique microenvironment of tumors and externally-applied physical stimuli provide a useful means to maintain a fine balance between the two properties. Recent studies show examples of NP systems that respond to a wide variety of internal and external cues in the tumor ECM. The majority of these systems are studied at preclinical levels, and a lot remains to be done to translate the technical potential to clinical benefits. Future efforts are required to increase the choices of stimuli-sensitive biomaterials. Equally important is the advancement in imaging technology to locate tumor lesions and technologies to apply external stimuli in a focused and non-invasive manner.

6. Expert opinion

Extracellularly activatable nanocarriers have shown a great potential to achieve drug delivery to tumors in a target-specific manner, but several challenges remain.

First, some of the stimuli-responsive NPs undergo incremental changes in response to stimuli, which often translates to low circulation stability. For example, DOX conjugated to MSNs via a pH-sensitive hydrazone linker showed a linear increase in drug release with the decrease of the solution pH [153], and MMP-sensitive liposomes showed gradual increase in the release of a model compound with the increase of the MMP-9 concentration [115]. When the responses are linearly proportional to the intensity of stimuli, there is a good chance that the NPs undergo inadvertent changes upon small fluctuations in environmental conditions. If the changes are irreversible, their circulation stability will be significantly compromised. The instability issues tend to be underestimated during development because of the overly simplified test conditions (e.g., buffered saline in lieu of blood) or the lack of analytical tools to predict the stability of NPs in complex fluids. Many NP systems with promising in vitro effects fail at later stages of development due to the insufficient improvement from free drugs, resulting from their instability during circulation [154]. On the other hand, an effort to improve the NP stability can lead to poor responsiveness to the stimuli. In this regard, it is worthwhile to note recent approaches that combine multiple stimuli to increase the contrast between responses to normal and tumoral conditions [135, 136].

Second, while many internal stimuli inherent to tumors have been identified in the literature, their applicability in clinical settings is often questioned as such conditions occasionally occur in normal tissues. For example, acidic environment can develop in conditions like ischemia or inflammation [155, 156], and small fluctuations in body temperature can occur due to fever or hot weather. It is also suggested that MMP-activation occurs in circulation rather than in tumor tissues [157]. Moreover, conditions to activate NPs may not be readily met if they are located far away (e.g., acidic and hypoxic regions) from the perivascular regions where NPs typically accumulate [158], or parts of the NP construct interfere with the access of the activating conditions (e.g., interference of PEG with an enzyme access to the cleavage site). In order to further advance the field of stimuli-responsive nanocarriers, it is necessary to make interdisciplinary efforts to identify new targeting stimuli and understand their physiological backgrounds.

Third, the need to address the intricate nature of tumor biology often leads to increasing complexity in NP structure and production methods. Technical complexity is seldom considered an issue in academia but rather encouraged for the advancement of materials science. However, if the complexity is not justified by the substantial improvement in therapeutic benefits, it is difficult to gain significant attention from the consumers. In particular, the growing complexity leads to increasing difficulties in quality control of the production, toxicological studies of the product, and regulatory approval processes, which pose significant obstacles in the late stage of new product development. While multiple chemical functionalities may be an integral part of a versatile carrier, a conscious effort toward a simple and scalable method should be made in tandem in the early phase of development.

Article highlights.

• For successful drug delivery to tumors with nanoparticulate carriers, they should remain stable during circulation, without interacting with healthy cells, releasing drug, or entering off-target organs and tissues. Once nanoparticles arrive at intended targets, they should release the encapsulated drug in the vicinity of tumor cells or enters the cells to unload it inside.

• Features of tumoral microenvironment and the externally-applied physical stimuli provide a useful means to maintain the balance between circulation stability and reactivity in tumors.

• Unique features shared by many solid tumors, such as hypoxia, acidity, and overly-expressed enzymes, are used as internal stimuli for the activation of nanocarriers.

• Non-invasive external stimuli such as light, ultrasound, magnetic field, and temperature are also used for spatiotemporal control of drug delivery.

• A single stimulus or combinations of multiple stimuli are employed for the extracellular activation of various types of nanoparticle systems such as inorganic or polymeric nanoparticles, liposomes, and dendrimers.