Externally Addressable Smart Drug Delivery Vehicles - ACS Publications

Perspective Cite This: Chem. Mater. XXXX, XXX, XXX−XXX

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Externally Addressable Smart Drug Delivery Vehicles: Current Technologies and Future Directions Somiraa S. Said,†,‡,§ Scott Campbell,†,§ and Todd Hoare*,† †

Department of Chemical Engineering, McMaster University, 1280 Main Street West, Hamilton, Ontario L8S 4L7, Canada Department of Pharmaceutics, Alexandria University, Alexandria 21521, Egypt

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ABSTRACT: Growing clinical evidence around the value of delivering drugs with pulsatile, user-controlled, or patient-specific kinetics has recently resulted in extensive efforts toward the design of new functional materials that can enable on-demand externally triggered drug release using noninvasive or minimally invasive stimuli such as temperature changes, magnetic fields, light, ultrasound, and/or electric fields. While substantial technical improvements have been made toward achieving longer-term and user-defined release kinetics from such materials (based principally on innovation in materials design and the development of various triggering modalities), clinical translation of such technologies remains in its infancy; specifically, multiple residual challenges exist around the chemical complexity, biological responses, and predictability of externally triggered release vehicles when used in vivo. In this perspective, we aim to outline major recent advances in the development of externally triggered drug delivery vehicles with improved performance both in vitro and in vivo and describe ongoing and emerging research challenges that, if overcome, can accelerate the translation of such technologies from the bench to the bedside.

1. INTRODUCTION Since pulsatile release from polymer nanocomposites was first demonstrated in the mid-1980s,1−4 substantial advances in polymer and materials science have enabled the development of new and more effective methods for externally triggering drug release.5 In more recent years, investigations into reversibly triggerable drug delivery systems with the potential to be switched both on and off “on demand” have become an area of particular focus, as outlined in a growing number of review articles6−10 including particularly excellent contributions by Timko et al.,11 Mura et al.,12 and Merino et al.13 Such interest has been motivated by both the need to develop noninvasive or minimally invasive treatments for local diseases (e.g., chronic arthritic pain) as well as a growing appreciation of the clinical relevance of periodic delivery of therapeutics (i.e., chronopharmaceutical drug delivery) for improving treatment outcomes for multiple diseases.14 In this latter case, improved therapeutic responses have been observed with the pulsatile delivery of hormones,15,16 asthma treatments,17 chemotherapeutics,18,19 vaccines,20 anti-inflammatory medications,21 and pain relievers such as nitroglycerin and opiates,22 the latter of which in particular lose efficacy if they are delivered passively over extended periods.11,14,18,23 The potential of externally activated delivery systems to prolong the lifetime of drug delivery vehicles by reducing or eliminating drug release when it is not clinically required, coupled with their potential for facilitating repeated local dosing without the need for repeated injections or other painful administration © XXXX American Chemical Society

methods, offers significant potential for improving patient compliance.11,24 However, despite the obvious therapeutic benefits of externally controllable delivery systems in such applications, with few exceptions clinical treatments are still performed either by directly administering drug at defined times or by using passive drug delivery systems with predetermined release rates independent of changing physiological circumstances or patient needs.25 In the former case, frequent readministration is required, both reducing long-term patient compliance and burdening the time of medical practitioners; in the latter case, once the delivery system is implanted/injected into the body, drug release is completely outside the control of both the practitioner and the patient. This gap between what is known clinically to be a better treatment and the state of currently available clinical therapies presents a tremendous scientific and market opportunity in developing triggerable release systems. However, there remains a significant gap between technologies shown to be effective in the lab or in small animal models and the clinical needs for human therapies in this area. As such, both novel material design approaches as well as improved activation strategies are required to fill this gap and provide the control and biological responses required for human translation. Received: May 7, 2019 Revised: June 21, 2019 Published: June 24, 2019 A

DOI: 10.1021/acs.chemmater.9b01798 Chem. Mater. XXXX, XXX, XXX−XXX

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Chemistry of Materials

Figure 1. Modalities and mechanisms associated with externally triggered drug delivery.

Figure 2. Different mechanisms for enhanced drug release of thermally triggered delivery systems through (a) volume phase transitions for hydrogels, (b) phase transition/melting temperature for lipid-based delivery systems, and (c) glass transition temperature for glassy polymers.

changes in response to that internal stimulus to control drug release (the actuating material). The general design and mechanisms leveraged to generate such materials are illustrated in Figure 1. This combination of signal transduction and the actuation of that signal into a physical change in the material itself differentiates externally triggered drug release vehicles from photodynamic therapy or thermotherapy agents whose primary purpose is only the transduction of the signal into a stimulus (e.g., the generation of radicals or reactive oxygen species,26,27 heat,28,29 or UV irradiation30) that acts directly on a biological target (e.g., local killing of cells by hyperthermia). The choice of the actuator material is thus directly tied to the type of signal that can be generated by the transduction material. 2.1. Thermal Triggering. Thermally triggered drug delivery is the most widely investigated transduction approach for externally activated drug delivery.11,25 Innate local temper-

In this perspective, we seek to outline the state-of-the-art in the field of externally activated drug delivery systems with a particular emphasis on designs that enable better control over the spatiotemporal drug release achieved and/or have demonstrated preclinical evidence of efficacy. Based on those examples, we outline future directions for research on the materials science and biological responses to externally activated drug delivery vehicles that are in our view essential to facilitate the translation of externally activated drug delivery vehicles from the bench to the clinic.

2. CURRENT METHODS FOR EXTERNALLY TRIGGERED DRUG RELEASE In general, externally activated drug delivery vehicles are designed by combining one material that can receive the external signal and convert it into an internal stimulus (the transducing material) and another material that physically B


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A variety of other nanoparticles have also been explored for transducing NIR activation of drug release.44 Carbon-based nanoparticles have attracted particular interest, although the intensity of research around the use of carbon nanotubes for external activation has begun to wane as the health concerns associated with the genotoxicity of such materials have grown.45 Many of the research groups formerly investigating these materials have transitioned to graphene oxide sheets that similarly can generate heat in the presence of NIR irradiation due to their combination of plasmon, fluorescence, and nonlinear emissions,46,47 although the longer term biological safety of graphene derivatives in vivo still requires further investigation. Other metallic-based nanoparticles are increasingly being investigated as thermal transducers for externally mediated release, borrowing from the rapid developments in the photothermal therapy (PTT) literature in developing more efficient and/or more degradable inorganic PTT agents without introducing significant toxicity. For example, Wang et al.48 reported a degradable supramolecular hydrogel nanocomposite made from an α-cyclodextrin (α-CD) and poly(ethylene glycol) (PEG)-modified dendrimer hydrogel encapsulating platinum nanoparticles, a current leading PTT agent.49 Upon NIR irradiation, the platinum nanoparticles induced the photothermally triggered degradation of the hydrogel via dethreading of PEG from the α-CD to induce drug release. Copper sulfide (CuS) nanoparticles similarly show benefits in PTT applications44 and, unlike gold or platinum, can be metabolized in vivo.50 Guo et al. have leveraged these benefits by developing chitosan-coated hollow CuS nanoparticles that can spontaneously self-assemble into immunoadjuvants under NIR irradiation.51 In this approach, the hollow CuS nanoparticles degraded after laser excitation, allowing spontaneous triggered assembly of nanoplexes comprised of chitosan with the cytosine-guanine (CpG) that improved tumor retention of the immunotherapy.51 Other emerging PTT agents such as molybdenum disulfide (MoS2) nanosheets52 and black phosphorus53 may also be relevant to explore for triggered drug release. Organic photodynamic dyes can also be used as transducers for thermally activated drug release. For example, Keurentjes et al.21 reported triggered in vivo release of ibuprofen based on a glass transition switch upon NIR irradiation of poly(D,L-lactic acid) (PLDL) strands dip-coated with quaterrylenebis(dicarboximide) as the organic photodynamic dye. The miscibility and/or degree of immobilization of such organic dyes is often lower than that achieved with inorganic nanoparticles, making this approach less widely pursued at the present time. However, the existing clinical approvals for some of these organic dye agents in PTT applications make them in our view warrant more attention in the design of externally activated delivery vehicles. 2.2. Magnetic Triggering. Magnetic fields represent the next most common transduction method for external activation of drug release vehicles, with both static and alternating magnetic fields (the latter of which ultimately leads to thermal triggering in most cases) having been demonstrated to induce drug release. The vast majority of such devices utilize superparamagnetic iron oxide nanoparticles (SPIONs) as the transducing material, 5−20 nm single crystal nanoparticles demonstrated to have favorable immuno-compatibility, ease of synthesis, and precisely controllable compositions and particle sizes relevant for applications including drug release,54

ature increases with the potential to self-trigger drug delivery can be in some cases observed in vivo as a consequence of a local disease state (e.g., cancerous tissues with heightened metabolic activity31 or infection sites32), offering potential for targeting therapeutics to such local sites using only the actuator component of the externally triggered drug delivery system. However, there is frequent variation between patients as to the magnitude of the temperature differences that can be achieved,33 making external triggering of temperature changes typically preferred for ensuring local on-demand delivery. Regardless of the method of heating, the actuator for thermally triggered delivery vehicles is typically a biomaterial that can retain its payload at physiological body temperature (∼37 °C) but rapidly deliver the drug upon localized heating (∼40−42 °C), as illustrated in Figure 2.12 Among the most widely used actuator materials are those that can undergo some kind of phase transition between normal physiological temperature and a physiologically safe and accessible higher temperature, including polymers that exhibit a lower critical solution temperature (LCST) such as poly(N-isopropylacrylamide) (PNIPAM)34 or poly(oligoethylene glycol methacrylate) (POEGMA),35 hydrogels comprised of the same materials that exhibit a volume phase transition temperature (VPTT),36 block copolymer micelles that conformationally reorient or invert upon heating,37 or liposomes that undergo a phase transition and thus associated conformational variations in the lipid bilayer upon heating.38 This latter category of thermally activated liposomes has in particular attracted interest given that the choice of phospholipid(s) used to assemble the liposome can result in the creation of either leaky membrane structures38 or effective solubilization of the entire nanostructure39 depending on the type of release kinetics desired. As an alternative to phase transition-based approaches, glass transitions can also be used to mediate release. For example, Viger et al. developed electrosprayed drug-loaded poly(lactide-co-glycolide) (PLGA) particles containing confined water pockets that can be rapidly heated upon exposure to 980 nm near-infrared (NIR) light,40 triggering the glass transition and thus promoting drug release (in this case achieving an impressive 25:1 on:off release ratio).40 Indeed, in our view, the glass transition represents a relatively underexplored mechanism for mediating externally triggered release that offers the potential for substantially lower off state release since diffusion of drug through a glassy polymer is significantly slower than that through a collapsed hydrogel that still has significant porosity. The transducer component of thermally triggered delivery vehicles can be any material that can generate heat in situ in response to an external signal, the most popular of which are NIR-responsive nanomaterials. Anisotropic gold nanoparticles are most commonly used given the generally recognized bioinert properties of gold nanoparticles in the body coupled with the capacity of anisotropic gold nanoparticles to generate heat in the presence of NIR based on surface plasmon resonance (SPR) effects. Many examples of the use of gold nanorods or nanoshells for triggering remote-control drug delivery have been reported, as outlined in recent reviews by Timko et al. and Tong and Kohane.11,41 Gold nanorods and gold nanostars are particularly relevant since they can facilitate photothermal heating in the 800−900 nm range by adjusting the aspect ratio of the nanorod42 or the length of the arms in the nanostars,43 enabling deeper tissue activation than with other geometries that undergo SPR at lower wavelengths. C


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Figure 3. Examples of high-frequency AMF-induced magnetic drug release. (a) Nanocomposite ethyl cellulose membranes containing thermoresponsive nanogels and SPIONs demonstrate on−off drug release upon the application of an oscillating magnetic field. (b) Injectable carboxymethyl cellulose/dextran hydrazone cross-linked hydrogels embedded with SPIONs and thermosensitive microgels facilitate pulsatile drug release upon AMF application. Modified with permission from refs 72 and 76. Copyright 2011 and 2015 American Chemical Society.

magnetic resonance imaging (MRI),55 cell tracking,56 and hyperthermia.57 Of particular importance, SPIONs (unlike most nanoparticles used for transduction of external signals) can be degraded into nontoxic iron species in vivo.58,59 Furthermore, based on their superparamagnetic properties, SPIONs can be used to both activate drug release (static or alternating fields) and direct the delivery vehicles to a desired site of action (static fields), offering an attractive combination of both on-demand drug delivery as well as site-specific drug delivery. Such combined targeting has been demonstrated with a variety of magnetic nanoparticles,60 nanocarriers,61 liposomes,62 and nanocapsules.63 The MRI activity of SPIONs further allows for facile fabrication of “theranostic” vehicles that can both deliver a drug and assess the efficacy of that delivered drug in treating the targeted disease.64,65 Drug release can be magnetically triggered from delivery systems containing SPIONs using either static or alternating magnetic fields. The application of a static magnetic field can be used to mechanically deform or move all or parts of a composite delivery system to induce triggered release. Most commonly, magnetically induced compression of a soft SPION-containing elastomer or hydrogel has been used to effectively squeeze the carrier to convectively enhance drug release.19,66,67 For example, triggered drug release has been demonstrated following the exposure of a ferrogel composed of SPIONs, Pluronic-F127, and a hydrophobic drug (indomethacin) to a static magnetic field;68 even larger release was achievable using macroporous ferrogels that can be further compressed in response to the same magnetic field strength, enabling on-demand release of various biological agents such as enzymes, chemokines, and even cells.66 Second, alternating magnetic fields (AMFs) can drive the release of drugs from SPION−polymer composites via one of two mechanisms. The application of a low frequency (Hz) alternating magnetic field (AMF) can locally shear the material adjacent to the SPION, either dynamically or irreversibly creating damage in the carrier that leads to enhanced drug

release. As particularly relevant examples, Cezar et al. developed biphasic macroporous alginate/SPION ferrogels that could greatly deform upon exposure to a weak AMF (a 1 Hz field for 2 min) to enable effective on-demand release of both drugs (mitoxantrone) and cells from the materials,69 Luo et al. reported triggered doxycycline delivery from layer-bylayer assembled magnetic microcapsules using low frequency AMF (50 Hz for 30 min) to dynamically alter microcapsule permeability,70 and Nappini et al. demonstrated that lowfrequency (0.2−5.2 kHz) AMF could alter the permeability of the lipid bilayer of magnetoliposomes loaded with cobalt ferrite magnetic nanoparticles to enhance long-term drug release.71 Alternately, the application of a high frequency (tens to hundreds of MHz) AMF results in the local generation of heat via hysteresis losses and/or Néel relaxation that, when coupled to thermoresponsive materials described in the previous section72−74 or temperature-dependent biological polymers such as gelatin that can change conformation upon heating,75 can induce drug release via thermal triggering. The SPION transducer and thermoresponsive polymer actuator can be incorporated into multiple types of delivery vehicles. Membrane-based devices have been among the most popular, in which the collapse of the thermoresponsive polymer or micro/nanogel reversibly creates free volume inside a porous continuous phase to induce release. Among the first examples enabling such reversible activation was an ethyl cellulose-based membrane into which both SPIONs and thermoresponsive microgels were embedded; upon the application of the highfrequency AMF, the heat generated by the SPIONs is transferred to the microgels to drive a volume phase transition, opening the pores to induce up to a 20:1 on−off gradient in drug release from the drug reservoir.76 Our group has since demonstrated that using an injectable hydrogel as the continuous phase in place of ethyl cellulose enables the injection of the vehicle without the need for surgical implantation, albeit with a somewhat reduced on:off gradient D


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Figure 4. (a) Schematic of the composition of SPION-loaded microgel particles and their thermosensitive sol−gel transition; (b) Heating of SPION microgels for multiple magnetic hyperthermia therapy (MHT, heating to 39−45 °C) and magnetic thermal ablation (MTA, heating >45 °C) cycles using an AMF. Modified with permission from ref 84. Copyright 2016 Elsevier.

Figure 5. Schematic of the core−shell implant design (left) and the on-demand release under periodic 1 h AMF exposures (right) from a cylindrical polymer composite loaded with 5 wt % ibuprofen and 50% SPIONs. Modified with permission from ref 73. Copyright 2012 Royal Society of Chemistry.

undergo a sol−gel transition upon injection into the body to create a bulk gel.84 The resulting highly uniform SPION content achieved via the concurrent collapse of the PEG chains and the formation of physical bridges between the L-isoleucine ethyl ester groups allows for highly consistent remotely actuated hyperthermia within the now-macroscale device (Figure 4), although the use of this strategy for regulating drug release has not yet been demonstrated.84 Indeed, there remain few reported studies in which all the components of an injectable delivery vehicle activatable by magnetic fields are fully degradable in vivo, an issue that needs further attention for translating such vehicles into the clinic. Developing effective strategies for magnetic activation upon application by more widely available lower power magnetic fields with frequencies on the few kHz range, rather than the high power and hundreds of kHz/MHz frequency magnetic fields now typically used, is also important for improving the potential translation of such devices into clinics or, in the future, into patients’ homes. One recently reported approach of interest to address this challenge is that of Yassine et al.,85 who incorporated anisotropic iron nanowires (FeNWs) with dimensions of 500 nm length × 45 nm diameter into PNIPAM microgels. AMF-driven cumulative drug release of approximately 70% can be achieved following repeated pulses using an AMF frequency of 20 kHz and 5-fold lower power than required for conventional AMF-driven activation, exploiting the use of frictional heat generation rather than hysteresis losses/Néel relaxation to enable activation. Alternate heating mechanisms can also be employed to expand the scope of materials useful for AMF-induced release beyond SPIONs, whose potential for oxidation and degradation in vivo can limit their utility as components of long-term drug release systems.86 For example, gold nanoparticle films

(4:1) and duration of action achievable (∼1 week) given the enhanced porosity of the continuous phase (Figure 3).72 Static magnetic fields can also be used in conjunction with membrane-based devices to facilitate magnetically triggered release. For example, static magnetic fields have been used to induce the movement of magnetic nanoparticles up or down to either block or unblock the pores of a polycarbonate membrane to enable on−off switching of drug release.77 Similar gating-based release has also been demonstrated on injectable nanoscale delivery vehicles based on mesoporous silica nanoparticles, using thermally driven disassembly of inclusion complexes for single-pulse delivery (e.g., uncapping of cucurbit[6]uril from a silica molecular thread),78 thermocleavable cross-linked polymers (e.g., azo-containing poly(ethylene glycol)),79 or thermoresponsive polymer grafting to the pores (e.g., linear PNIPAM switches).80 However, challenges persist with such vehicles in terms of designing efficient reversibility coupled with sufficient drug dosing for practical use as well as achieving clear in vivo proof-of-concept. In particular, long-term reproducibility of such devices is typically low given the potential for membrane fouling over time. Injectability can also be achieved by reducing the size of the SPION−thermoresponsive polymer composites to the microscale or nanoscale,81−83 although practical use of such vehicles would require that (1) the micro/nanoparticles are designed to degrade over time and (2) the particle design can effectively prolong drug release, given that release of small molecules from micro/nanogel particles is typically fast even in the absence of activation. In this context, one of the more promising materials reported recently is Zhang and Song’s SPION-loaded degradable poly(organophosphazene) microgels coated with PEG and L-isoleucine ethyl ester that can spontaneously E


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Chemistry of Materials embedded in PNIPAM hydrogels can generate heat under similar AMF conditions used for SPION heating via the induction of eddy currents within the gold film, a result of the collective conductivity of the supra-structure of spherical gold nanoparticles formed.87 In this context, the improved longterm stability of gold in vivo should improve long-term triggering while the AMF responsiveness of the gold film overcomes the penetration depth issues associated with NIR activation of gold.87,88 Miniaturization of such films to generate similar effects in micro/nanoscale delivery vehicles would further enhance the potential of this technique. Similar to NIR activation, not only a thermal conformational phase transition but also the melting point (Tm) or the glass transition temperature (Tg) of polymers can be utilized to remotely activate drug release. As an interesting example of the former, Müller et al. recently reported biobased hydrophobic dextran ester derivatives with physiologically relevant melting points that could induce well-defined drug release in response to an external AMF.89 As an example of the latter, Rovers et al. reported an implantable poly(methyl methacrylate) p(MMA)co-SPION core/poly(butyl methacrylate(BMA)-co-MMA) shell cylindrical implant in which the shell was tuned to have a Tg of ∼52 °C, enabling significantly enhanced drug release upon AMF exposure and the resulting local heating of the shell above its Tg (Figure 5).73 The proven record of tissue compatibility of the pMMA/pBMA polymers90 that can achieve such physiologically relevant Tg values makes the latter approach of particular interest in terms of achieving externally triggered on-demand release, especially for highly potent drugs for which sufficiently large local doses can be delivered through the rubbery phase of the polymer over short AMF exposure times. However, the nonbiodegradable nature of polymethacrylate copolymers necessitates the surgical removal of the implant after the delivery of the drug payload. While this may not be problematic for long-term implants, creating biodegradable analogues of these materials would be clinically beneficial for shorter or moderate-term applications. Increasing interest is also being directed to tethering drugs to carriers using thermo-reversible bonds that can only be cleaved upon AMF heating, providing an alternate solution to the leakage issues that can limit the functional lifetimes of reservoir-based delivery vehicles. Such conjugates have been achieved using Diels−Alder chemistry (reversible via the retroDiels−Alder pathway upon heating)91 or azo chemistry (reversible via RNNR′ bond cleavage of the aliphatic azo compounds with the release of N2 upon heating).92 A recent example of this strategy is the work of Fuller et al., who fabricated nanoparticles consisting of SPIONs and a poly(hexamethyacrylate-co-maleimide methacrylate) surface stabilized with a poly(ethylene glycol)-block-poly(lactic acid) block copolymer.61 Conjugation of a furan-containing drug to the maleimide functional side groups via a Diels−Alder reaction is reversible upon AMF application to release the drug ondemand, with the presence of the maleimide functional groups over the full length of the polymer enabling significantly enhanced total drug loading relative to chain end-only functionalized materials. The value of such an approach in a practical application would depend on whether the likely improved on-state triggering and longer drug release lifetimes of the device would supersede the synthetic and regulatory complications introduced by the need to conjugate a drug to the vehicle. Analogous approaches in which the gel itself can degrade upon AMF heating have also been explored,93

although the potential for rigorous on:off switching is lower than with direct drug conjugation given the “off”-state leakage associated with hydrogel matrices. 2.3. NIR or UV/Visible Light Triggering. While light in the UV (10−400 nm) and visible (390−700 nm) regions cannot practically be used for deep-tissue triggering since it is strongly absorbed by skin and other tissues and can damage tissues at much lower powers than can NIR, near-surface tissues such as the skin, the ear, or the back of the eye and/or delivery strategies such as transdermal delivery remain potential candidates for UV/visible light-triggering systems given that penetration depth is not an issue in such cases. Expanding the available range of frequencies for light irradiation expands the potential mechanisms available to enact on-demand release, including introducing various photolabile bond-cleavage or bond-isomerization chemistries such as coumarins, spiropyrans, azobenzenes, and o-nitro benzyl derivatives that can actuate in the UV or visible range but are typically not operable in the NIR range or via magnetic/heat activation.94 Such photolabile chemistries can be applied to change the degree of cross-linking and/or ionization locally within a gel to enhance diffusion-based release11,95,96 and/or release bound drug via photocleavage of the linker between the drug and the vehicle.97,98 However, challenges persist with both tissue compatibility as well as the relatively low degree of bond cleavage often achieved at biologically acceptable powers/wavelengths for many typical photosensitive agents.99 In this context, the use of the hydrophobic−hydrophilic transition that accompanies the trans−cis photoisomerization of azobenzenes,100−102 in particular to deassemble amphiphilic block copolymers on demand, is particularly attractive, although azobenzenes are not without their own in vivo cytotoxicity concerns. The use of NIR light (650−900 nm) has more advantages for phototherapy and externally triggered drug delivery over the use of UV/visible light due to deeper tissue penetration with higher biosafety,103 although the penetration distances even for NIR light are still relatively low (∼2 mm through skin104) relative to the needs for deep tissue therapies. NIR light-triggered drug delivery vehicles comprise three main types: (1) NIR-photothermal converting nanoparticles such as precious metal-based nanoparticles with NIR plasmon resonance,105 carbon-based materials,106 and semiconductor nanocrystals with NIR absorption,107 as discussed in the previous section; (2) upconversion nanoparticles that convert NIR light to UV in vivo to allow deeper tissue penetration but maintain the potential for UV-mediated processes, as triggered via triplet−triplet annihilation-based upconversion108 or lanthanide-doped upconversion nanoparticles;109 and (3) optical strategies to enable deeper light penetration, such as two-photon absorption110 and second harmonic generation.111 Recently, the second NIR window (NIR-II, 1000−1700 nm) has attracted particular attention due to deeper tissue penetration compared to the first NIR window (NIR-I, 650− 900 nm),112 with bioimaging in the NIR-II demonstrated to result in better spatial resolution and feature contrast at deeper tissue depth due to reduced autofluorescence and photon scattering;113 as a result, this band is also of key emerging research interest for triggered-drug delivery and associated in vivo monitoring of long-term toxicity and pharmacokinetics of the NIR-responsive drug delivery vehicles.114,115 The recent development of self-immolative polymers stimulated by light activation represents an alternate and F


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HIFU that resulted in higher drug accumulation and better antitumor efficacy in vivo compared to nonirradiated controls.127 Cavitation can be achieved with lower intensity ultrasound frequencies (in the kHz range) and is already used clinically in the context of microbubble-based therapies in which generally perfluorocarbon (PFC)-based nanoemulsions are converted into microbubbles under the action of therapeutic ultrasound to enable triggered cellular uptake and/or release of the entrapped drugs at the site of interest.128 Localized cavitation of the surrounding fluid can also be used to drive triggered release from other types of carriers via the generation of highly localized shear that mechanically disrupts the delivery vehicle. For example, Zhang et al. reported the triggered release of Nile red dye from poly(lactic acid)/ poly(ethylene glycol) block copolymer micelles using HIFU to mechanically disrupt the hydrophobic interactions holding the particle together.129 However, aside from the use of ultrasound as the heating source for the thermally activated delivery vehicles previously described, most reported ultrasound-based delivery vehicles are functionally irreversible, making ultrasound in most cases useful only for activating release rather than enabling the reversible on−off control facilitated by other triggering techniques. Some recent work has attempted to address this challenge of reversibility using novel materials design. In particular, selfhealing polymers offer appealing properties in the context of ultrasound-mediated delivery given that the mechanical disruption enacted by ultrasound can be healed in the absence of the ultrasound trigger to recover an “off” state release. Ratner’s group pioneered this idea by demonstrating switchable on−off release from a 2-hydroxyethyl methacrylate-based hydrogel coated with C12 n-alkyl methylene chains whose selfassembly in water can be transiently disrupted by ultrasound.130 More recently, Huebsch et al. showed that the transient disruption of ionic calcium−alginate cross-links under ultrasound allowed for enhanced release of mitoxantrone in an in vivo cancer model131 while Sun et al. disclosed the use of double network hydrogels in which a photopolymerized ultrasound-stable network maintains structural rigidity (and thus a degree of reversibility) while a phenylboronic acid/ polyphenol secondary dynamic network provides ultrasoundmediated release.132 Given recent advances in developing selfhealing polymers via various mechanisms (many of which are compatible with in vivo use), this approach appears to be one worthy of additional investigation, particularly if less permeable/soluble polymers that could further reduce the “off” state release are used. Alternatively, ultrasound-mediated gating approaches also offer the potential to expand the scope of release kinetics achievable upon ultrasound activation. For example, Birajdar et al. reported effective nanoparticle release from nanoparticleimpregnated core−shell nanofibers by exploiting the different mechanical indices of the embedded nanoparticles and the surrounding fiber to allow the nanoparticles to “pop out” (akin to corks on a wine bottle) without disrupting the fiber structure upon the application of ultrasound.133 Furthermore, the “uncorking” as a result of ultrasound activation creates templated pores in the core−shell fiber that were shown to subsequently mediate zero-order drug release instead of the typical burst release observed using ultrasound-based triggering (Figure 6).133 Miniaturization of this vehicle offers another appealing potential design strategy for achieving nonburst release kinetics using ultrasound as a trigger.

appealing strategy to allow for shorter or less intense irradiation to actuate a large change in drug diffusion. For example, Fan et al. reported on-demand drug delivery based on UV light-cleavage of the linker end-caps of self-immolative poly(ethyl glyoxylate)−poly(ethylene oxide) nanoparticles.116 Similar self-immolation has been demonstrated using NIR light as the degradation trigger,117 expanding the use of this approach for deeper tissue penetration. The work of Gillies’ group in developing polyglyoxylate-based self-immolative polymers is particularly relevant from a translational perspective since, unlike most reported self-immolative polymers that release potentially toxic small molecule byproducts, the byproducts of polyglyoxylate self-immolation are ethanol and a naturally occurring metabolic intermediate,118 enabling safer use of such a triggering strategy in vivo. Photocaging strategies in which a photoactive functional group is used to dynamically mediate the activity or diffusibility of a drug also represent an interesting approach in which a relatively small amount of photoactivation can induce relevantly large biological effects. In particular, the use of photocleavable groups like o-nitro benzyls119 or photodimerizing−cleaving groups like thymine120 have demonstrated efficacy for opening and closing pores on mesoporous silica nanoparticles, enabling reversible gating of larger doses of drug release from the mesopores. Photolabile caging groups have also been widely exploited to turn the biological activity of biomacromolecules on and off on-demand; as one example, lipid vesicles filled with the cellular machinery responsible for transcription and translation (including amino acids, ribosomes, and DNA) caged with a photolabile 1-(4,5-dimethoxy2-nitrophenyl) diazoethane protecting group have been shown to enable spatial and temporal control over protein synthesis both in vitro and in vivo using metered doses of millisecondlong UV irradiation.121 The practical use of such systems in the clinic, however, remains challenging given the complexity of the preparation conditions used to prepare photocaged therapeutics. Light has also been used in nontraditional contexts to enable controlled release. For example, SPION-loaded PNIPAM hydrogel beads intended to work within transdermal patches have been successfully heated using visible light absorbed by the embedded SPIONs, enabling triggered dexamethasone release at a rate controlled by the intensity of the light applied.122 In principle, this approach enables the use of magnetic composites that could still be directed to a target site by a static magnetic field but remove the need for an often expensive AMF source for triggering the device once at the target site. Ultimately, however, this and other broader uses of light-based triggering requires the development of strategies to safely enhance the penetration of such light sources into tissues. 2.4. Ultrasound Triggering. Ultrasound triggering is characterized by relatively high spatiotemporal resolution and, unlike light-based sources, highly effective penetration into soft tissues of interest, making it a recent target for triggered release applications.123−125 Ultrasonic waves can remotely trigger drug release through either thermal or mechanical effects generated by cavitation phenomena or acoustic convection forces.125,126 Thermal heating is typically accomplished using high-intensity focused ultrasound (HIFU) and is often used as an alternative activation method for thermally triggered release vehicles (Section 2.1). For example, Ranjan et al. reported triggered release of doxorubicin from thermosensitive liposomes using G


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damage.138 Furthermore, both nonspecific fouling and bodyinduced (instead of controlled) redox chemistry can interfere with the in vivo performance of most conductive polymers, potentially leading to inflammation/fibrosis in the former case and loss of activity (particularly over the longer term) in the latter case. Electric field-induced deswelling of hydrogels to drive forced convection of a drug out of a hydrogel matrix offers another strategy for electrically triggered release that, based on the protein-repellent properties of most hydrogels, can address at least the fouling issues associated with other electroactive release vehicles. Polyelectrolyte hydrogels under an electric field generally undergo electro-responsive deswelling,139 with the amount of drug released typically proportional to the degree of deswelling observed.140 In particular, electroresponsive gels synthesized from hyaluronic acid or chondroitin sulfate have been shown to facilitate pulsatile release of small molecule drugs, proteins, and peptides, albeit with substantial maintained “off” state release due to the inherent porosity of the swollen gel.141 Doping of conductive nanoparticles into the hydrogel can further improve the speed and/or the magnitude of the electroactive responses, although potentially also introducing stability or fouling challenges into the material design.142 Electrically activated microchips represent another interesting strategy for triggerable pulsatile release of tunable drug combinations with high spatial and temporal precision and/or integrating a degree of automation that is not easily achievable using other pulsatile delivery strategies.138 The most advanced of such devices are silicon chips that contain multiple drug reservoirs capped with polymeric or gold seals; upon application of a localized current addressed to a particular reservoir, ablation of the gold film or electrically stimulated degradation of a thin polymer film opens the seals on individually addressable reservoirs to release a drug on demand.143−145 Other types of multireservoir approaches have also been reported, including using reduced graphene oxide (rGO)−poly(vinyl alcohol) (PVA) hydrogels assembled into a chip-like device that released enhanced quantities of lidocaine upon cyclic exposure to an external electrical field.146 All of these approaches, however, require surgical implantation to enable function, limiting their use to applications in which

Figure 6. Ultrasound-triggered drug release from core−shell nanofibers by sonication-induced uncorking of silica nanoparticles: (a) attachment of nanoparticles onto the core−shell nanofibers; (b) embedding the nanoparticles by solvent−vapor annealing; (c) detaching the nanoparticles by ultrasonication to trigger drug release from the core. Adapted with permission from ref 133. Copyright 2016 Elsevier.

2.5. Electric Field Triggering. External weak electric fields can be used to achieve on-demand drug release via a range of different actuation mechanisms. Most commonly, an applied electric field is used to perform reversible redox chemistry that can transiently alter the chemistry of a delivery vehicle to promote drug release. For example, a range of different delivery vehicles based on the conductive polymer polypyrrole show electric field-induced changes in net charge due to redox chemistry, enabling on-demand and reversible changes in the release of encapsulated drugs (particularly in cases in which the drugs are themselves charged).134 Similarly, reversible electric field-driven scission of supramolecular vesicles formed via host−guest complexation between the β-cyclodextrin and ferrocene end groups has been demonstrated to control drug release by regulating the assembly/disassembly of the vesicle according to the applied voltage,135 while an oxidizing voltage has been demonstrated to stimulate a reversible structural reconformation of aqueous vesicles composed of redoxresponsive self-assembled amphiphilic rod−coil tetraaniline and poly(ethylene glycol) into smaller puck-like micelles.136 Redox-stimulated polymer degradation also offers on-demand release potential similar to that of self-immolative polymers,137,138 although in this case providing only burst release control rather than on/off control. However, all these approaches are practically restricted in a clinical setting due to the limited tissue penetration of electric fields (in most cases requiring the use of implantable electronics) and the need to limit the applied voltage to avoid undesired local cell/tissue

Figure 7. (a) Schematic representation of the developed multifunctional doxorubicin-loaded microbubbles conjugated with SPIONs. (b) Deposition of the externally triggered multiresponsive system into the brain tissue was induced by focused ultrasound, and targeting was enhanced using an external magnet. DSPC: 1,2-distearoyl-sn-glycero-3-phospho-choline, DSPG: 1,2-dioctadecanoyl-sn-glycero-3-phospho-(1′-rac-glycerol), DSPE-PEG2000: 1,2-distearoyl-sn-glycero-3-phosphoethanolamine-N-[methoxy(poly(ethylene glycol))-2000]. Adapted with permission from ref 148. Copyright 2013 Elsevier. H


Breast cancer

ThermoDox with hyperthermia treatment

Osteoporosis

Breast cancer

Heat-activated target therapy (radiotherapy + hyperthermia + ThermoDox)

Electrical-triggering of teriparatide (FORSTEO) microchip-based implant. The implant integrated two drugcontaining microchip assemblies on the surface of an ablatable titanium/platinum housing that contained control and communication electronics

Liver tumor

ThermoDox activated by mild hyperthermia using focused ultrasound (FUS)

Phase I completed

Hepatocellular carcinoma

ThermoDox plus standardized RFA

Phase III completed Phase I completed Phase II suspended Phase I terminated

Painful bone metastases; Breast carcinoma; Non-small and small cell lung cancer; Adenocarcinoma Colon cancer liver metastasis

Hepatocellular carcinoma; Liver neoplasms


ThermoDox plus HIFU

Phase II withdrawn

conditions

Hepatocellular carcinoma

Breast cancer

Phase II terminated


Phase I completed

Phase 2 study of ThermoDox as adjuvant therapy with thermal ablation (RFA) in treatment of metastatic colorectal cancer (ABLATE) Phase 3 study of ThermoDox with radiofrequency ablation (RFA) in treatment of hepatocellular carcinoma (HCC) Targeted chemotherapy using focused ultrasound for liver tumors (TARDOX) Heat-activated target therapy of local-regional relapse in breast cancer patients (EURO-DIGNITY) Temperature-sensitive liposomal doxorubicin and hyperthermia in treating women with locally recurrent breast cancer The pharmacokinetics of human parathyroid hormone fragment (1−34) delivery with MicroCHIPS’ implantable reservoir array device

ThermoDox plus standardized RFA using standardized treatment dwell time for solitary HCC lesions

Phase III completed

treatment/intervention ThermoDox (thermally sensitive liposomal doxorubicin) in combination with microwave hyperthermia

status Phase I/II completed

study title

Phase 1/2 study of ThermoDox with approved hyperthermia in treatment of breast cancer recurrence at the chest wall (DIGNITY) Study of ThermoDox with standardized radiofrequency ablation (RFA) for treatment of hepatocellular carcinoma (HCC) (OPTIMA) A study of ThermoDox in combination with radiofrequency ablation (RFA) in primary and metastatic tumors of the liver MRI guided high intensity focused ultrasound (HIFU) and ThermoDox for palliation of painful bone metastases

Table 1. Externally Addressable Drug Delivery Systems in the Pipeline for Regulatory Approval

Chemistry of Materials Perspective

I


Perspective


enhance drug diffusion by up to 25-fold) well ahead of other systems. At least two Phase III studies on such materials have already been completed, with many others underway using various heating methods and combinations of different therapy types (e.g., chemotherapy combined with radiation). In addition, a Phase I trial has been successfully completed on the MicroCHIPs technology for treating osteoporosis, a promising development toward the commercialization of that technology. However, relative to the broad scope of materials and targeting strategies discussed in the previous section, clinical translation and commercialization of externally addressable drug delivery systems has been slow. Of particular note, the chemical simplicity of the ThermoDox formulation relative to most of the more experimental systems previously described stands out, making the manufacturing, reproducibility, quality control, and ultimate regulatory approval of that type of material substantially simpler than for most other systems reported in the literature.

the accurate temporal control of combinations of drugs was crucial to the success of a given therapy. Electric field-based release vehicles may also offer unique potential for integrating electrochemical sensors together with a controlled release vehicle to facilitate closed-loop control over drug release without any need for external monitoring or intervention. As an example, a gold electrode coating consisting of an insulin-impregnated rGO thin film loaded with Ni(OH)2 was reported to enable both glucose sensing and electrically triggered release of insulin upon the application of a potential to the interface.147 While such devices have obvious attraction in the context of self-metering therapeutics, fouling (and its well-recognized effects on limiting signal:noise ratio in biosensors) and device lifetime (in terms of both biomolecule stability as well as the total drug dose that can be delivered) are both significant challenges to clinical translation. 2.6. Multiple Triggering. Responsiveness to more than one stimulus can enable synergistic control over drug delivery kinetics while also providing potential multiuse of the device for both therapy and imaging. As a representative example of materials enabling such multistimulus control, Fan et al. reported multifunctional doxorubicin-loaded microbubbles conjugated with SPIONs that can open the blood−brain barrier while delivering the drug upon focused ultrasound exposure, act as a dual MRI and ultrasound contrast agent, and facilitate additional spatial magnetic drug targeting (Figure 7).148 Many such examples in the literature consist of multiple material components that can be individually addressed by various activation stimuli. For example, Xie et al. reported triple stimuli responsive self-assembled micelles consisting of a thermosensitive shell of tetraethylene glycolyl poly(trimethylene carbonate), a light sensitive core of poly(2 nitrobenzyl methacrylate), and a side chain linked by redoxresponsive disulfide bonds.149 Such approaches also enable simultaneous delivery of different types of drugs in a manner that is individually addressable. For example, Cao et al. reported the use of nanogels composed of hydrophilic pH- and thermoresponsive poly(2-(dimethylamino)ethyl methacrylate) and hydrophobic photocleavable o-nitrobenzyl linkages that enabled co-delivery of hydrophobic cargos associated with the lipophilic nanogel core (via temperature, pH, and UV light activation) and hydrophilic cargos linked to the nanogel via disulfide linkages (via redox activation).150 Despite the attractive versatility of these multistimuli responsive systems, their inherent complexity in terms of their synthetic preparation, release mechanism, and control over release profiles may make them challenging for practical use (see Section 4.3 for further comments on these vehicles).

4. CURRENT CHALLENGES AND FUTURE PERSPECTIVES Based on the limited current translation activity in this area, it is clear that there are several key challenges associated with externally triggered drug delivery systems that need to be addressed to facilitate their translation from bench to bedside. The following section highlights what we perceive to be the five key challenges to overcome, coupled with some perspectives on emerging strategies to address such challenges. 4.1. Penetration. Effective penetration of the various activation signals into various sites of targeted action in the body remains a central challenge for many activation strategies. Materials design can help to address some of these issues; for example, up-conversion nanoparticles have enabled UVtriggered technologies to function much deeper into tissues than would otherwise be possible by enabling their activation with near-infrared light.151,152 However, this and other lightbased triggering techniques still offer only limited penetration potential (at best on the order of several millimeters)153 given the combination of high tissue attenuation and the potential for skin burns at higher incident powers.154,155 In addition, ultrasound-triggered techniques remain largely limited to soft tissues given the high attenuation of ultrasound signals in bone.156 However, technologies developed primarily for diagnostic imaging show promise for in part overcoming these challenges. Emerging developments from the field of photodynamic therapy in terms of improving light penetration into tissues using fiber optic technology,157 pulsed or two photon lasers that can achieve higher powers at a particular focal point,158 or nanoscintillators that can locally convert more penetrative X-ray irradiation into visible light159 all offer promise, although the surgical demands of the first strategy, the technology accessibility challenges of the second strategy, and the potential materials toxicity associated with the third strategy are all potential barriers to clinical translation. In the same way, recently developed techniques for enabling ultrasound imaging through bone may offer potential to improve the penetrability associated with ultrasound-mediated delivery vehicles.160−162 Integrating these and other recent innovations in the medical imaging field to improve the penetration depth and/or reproducibility of externally targeted drug delivery vehicles is a research area of substantial promise that has to-date been only peripherally explored. Alternately, magnetic field-based activation (which does not suffer from the

3. CLINICAL TRIALS OF EXTERNALLY ADDRESSABLE DRUG DELIVERY SYSTEMS While many externally activated systems have proven to be effective in animal models (including many of those already discussed in this perspective), relatively few have proceeded to the stage of clinical testing. Table 1 lists the externally addressable drug delivery systems that are currently in clinical trials. Thermoresponsive liposomes are the leading formulations in a translation context, with the ThermoDox system consisting of a lipid mixture of dipalmitoylphosphatidylcholine, myristoylstearoylphosphatidylcholine, and 1,2-distearoyl-snglycero-3-phosphoethanolamine-N-[amino(polyethylene glycol)-2000] that enables a structure change upon heating to 40−45 °C to create channels in the liposome bilayer (and thus J


Perspective


ing the speed of the response of the delivery vehicle to an external signal (since drug released from a triggered nanoparticle vehicle also has to diffuse through the entrapping material to be released72,169) and/or suppressing the degree of drug leakage at the untriggered “off” state72 can be addressed. Developing new ways to form injectable hydrogel scaffolds that are either much more permeable (if the entrapped particle itself regulates release and the purpose is immobilization) or much less permeable (if the porosity of the injectable phase is itself intended to mediate release) than existing scaffolds could address these challenges. In addition, recent work in the cell encapsulation field in which specific microparticle sizes (>1 mm diameter) and interfacial chemistries have been identified to significantly mitigate fouling and thus inflammation170 may help to inform the design of new mitigation strategies for the fibrotic responses to microparticle-based devices. 4.3. Control over Activation Signal in Vivo. Controlling and metering the intensity of the external signal required to safely deliver a desired drug dose in vivo remains a key challenge. For example, while the maximum temperature achieved in thermally activated drug delivery systems can be estimated in vitro using phantom studies at different incident laser or magnetic field powers, the differences in the local environment once the device is injected (e.g., thermal conductivity, fluid perfusion, etc.) can make such predictions inaccurate in the clinical case.21 Overshooting the target temperature can lead to substantial local tissue damage, while undershooting the target temperature would lead to substantially lower or perhaps no significant drug release achieved. While some studies have sought to couple magnetic activation with self-learning controllers171,172 and/or complementary imaging approaches that can either dynamically meter or report in real time the local temperature at the target site,173,174 such approaches require extensive instrumentation availability in addition to introducing complications in the clinical approval process. Alternate approaches to tag the loaded drug with radiolabel(s) or fluorophore(s) to enable real-time quantification of released drug can be effective but also require access to additional (and co-located) advanced imaging tools as well as customized therapeutics that pose both logistical and regulatory challenges.175 Differences in local attenuation due to the natural heterogeneity of tissues can similarly make accurate prediction of the strength of the transducer signal challenging for nonthermal stimuli. Recent work has suggested promising avenues for enabling improved control over the activation via advanced materials design. In the context of magnetically driven transitions, the use of magnetic nanoparticles with a Curie temperature (i.e., ferromagnetic to paramagnetic transition) just above the targeted activation temperature can inherently self-limit local heating. However, at the present time such self-regulation is only practically achievable using magnetic nanoparticle compositions that are less inherently tissue compatible or degradable compared to conventional SPIONs.176 Alternatively, coupling magnetically driven release vehicles with phase change materials that have a melting point corresponding to the target temperature of the stimulus offers an attractive strategy for limiting the maximum temperature to which the body is exposed. For example, Wang et al. demonstrated that Fe3O4@SiO2 particles entrapped in a polyurethane/paraffin phase change material could accurately maintain a target temperature in the safe but effective range for thermotherapy (41−47 °C) for several hours,177 an approach that could be

same attenuation issues) remains a particularly attractive research area in this context. Remote electrical stimulation may also be assisted by recent advances in soft and flexible electronics that both minimize the substantial inflammation and fouling responses commonly observed with traditional conductive materials in vivo163 as well as offer improved integration with native tissues. However, improved strategies must still be developed for remotely applying the local voltages required to drive most electrically induced release strategies to enable electrical stimulation in a fully or even minimally noninvasive manner. Ultimately, magnetic fields offer the best combination in our view of being highly penetrative, safe, and free of the need for implantable devices, although improving the oxidative stability of magnetic nanoparticles in the body is important for fabricating longer-term triggered delivery vehicles. 4.2. Delivery to and Retention at the Active Site. Even if an actuating signal can reach the target site, the externally activatable delivery vehicle must also be able to reach and be retained at the disease site for the therapy to be effective. Challenges exist on this point at all size scales of materials. While nanoparticle delivery vehicles can be injected directly at the target site (or, if targeted to bind specifically to an appropriate diseased cell or tissue, circulate and concentrate at the target site), subsequent retention at the target site often proves problematic even in cases in which effective biological targeting is achievable.164,165 Given the inherent site specificity achievable with many of the available transducing strategies, movement of the nanoparticle vehicles away from the target site is not necessarily problematic as long as the drug either does not release at all or releases very slowly at nonfocused sites. However, if long-term therapy is the goal, the reduction of the total drug dose at the target site over time coupled with the often decreasing release efficacy often observed upon repeated activation may prove to be problematic.61,72,73 Alternately, bulk scale implants or membrane-based devices typically cannot be injected, requiring surgical implantation and thus significantly reducing the practical clinical translatability of the device. Integrating drug reservoirs that can be externally refilled as needed is a potentially useful strategy to extend the functional lifetime of an externally triggerable device and thus makes surgical implantation a more clinically realistic option, although currently available percutaneously refillable reservoir devices come with their own challenges.166,167 Microparticle-based vehicles offer a desirable combination of immobilization at the target site and injectability, but they are often observed to induce significant inflammatory responses (e.g., frustrated phagocytosis for larger microparticles) and ultimately fibrotic responses that can significantly alter long-term device performance.168 Several approaches may be envisioned for addressing these challenges. Among them, in situ gelling formulations in which nanoparticles or soluble polymers can be injected as low viscosity solutions/suspensions but rapidly gel or solidify in the body following injection can address the injectability challenges of bulk devices, the immobilization challenges of nanoparticle-based devices, and (if the injectable phase is an antifouling hydrogel) the inflammation challenges of microparticle-based devices. In particular, use of the growing library of in situ-gelling hydrogels to replace the “hard” noninjectable membrane or solid scaffolds widely used to entrap nano/ microscale externally triggered delivery vehicles offers significant promise, provided that complications around regulatK


Perspective


known about the long-term fate or degradation pathways of materials such as upconversion nanoparticles or emerging magnetic nanomaterials that typically contain one or more rare earth elements that may be of concern from both an acute toxicity standpoint as well as a chronic standpoint given the uncertain elimination pathways of such materials;185−187 even if the toxicity proves to be acceptable, the potential impact of biological processes (e.g., oxidation, enzymatic degradation, etc.) on the long-term in vivo performance of externally activated release vehicles that often contain metallic or redoxactive organic components as activation agents remains largely an open question. Thorough standardized toxicity studies are essential to inform best practices in nanomaterial design to ensure that externally activated delivery vehicles containing such nanoparticles are safe and clinically practical to use over the long-term. Improvements in instrumentation to facilitate effective triggered release using more portable, less expensive, and ideally user-controllable devices would also substantially improve the potential impact of externally triggered therapeutics to patients. At the present time, most of the approaches described in this perspective require specialized medical instrumentation to administer the drug pulse, restricting patient access to such therapies while placing excessive strain on existing clinical imaging facilities that are already heavily used. Long-term, the coupling of lower-cost activation devices suitable for at-home use (either fully patientdirected or physician-directed via either preprogramming or telemetry linkages) with externally triggered drug delivery vehicles that could be activated using the lower powers achievable with such lower-cost devices offers tremendous potential to replace regular oral pills with more targeted and thus more effective therapeutics. 4.5. Translational Models and Safety. The wellestablished poor correlation between in vitro models and in vivo animal studies as well as the biological incompatibilities between the animal models used in preclinical studies and humans have led to thousands of drug delivery systems failing at various stages of the translational pathway. The inverse is likely true too, as systems that may have been effective in humans may have been phased out earlier on when they exhibited insufficient efficacies in vitro or in vivo.188 This problem is in our view likely to be even more acute for externally activated delivery systems given that variations in signal penetration depths (UV, NIR), heat transfer due to variations in blood supply or simply animal size (magnets, NIR), or other physiological differences offer an additional translational barrier beyond the typical material and pharmacokinetics considerations around the translation of controlled release vehicles. Understanding the interactions between the native biology and the on-demand drug release technology, including the influence of disease pathophysiology on delivery system accumulation, distribution, retention, and efficacy as well as the biopharmaceutical correlations between the delivery system properties and the in vivo performance in animals versus humans, is crucial for successful clinical translation.189 In particular, to mitigate the challenges with interspecies translation as well as the heterogeneity of many diseases (in particular cancer) in humans,190 (1) different animal models should be simultaneously investigated to attempt to mimic disease complexity in the human population,191 (2) the drug dose and administration route should be assessed based on in

similarly applied for externally activated drug delivery. Such an approach is also useful for limiting the total time over which the patient needs to be exposed to the external activating stimulus to achieve a targeted drug release rate, as the phase change material can store input energy during activation and then slowly release it to maintain a high local temperature for a designed period of time. Given the relatively low drug doses delivered by most on-demand drug release systems (and the corresponding long treatment times that would be required to achieve clinically relevant dosing), such prolonged dosing by energy storage is in our view a particularly attractive solution. However, the formulation of biologically acceptable phase change materials into relevant nanoscale or injectable delivery vehicles remains an outstanding challenge. 4.4. Materials and Instrumentation Complexity. As outlined earlier in this perspective, externally activated drug delivery vehicles typically require the incorporation of multiple chemical components: at minimum, a transducing component to receive the external signal and an actuating component that converts that signal into a physically significant change in the material that regulates drug release. However, from a regulatory perspective, the need for multiple materials (coupled with one or more therapeutic(s)) represents a significant clinical translation barrier. Indeed, while the current trend in the literature appears to be to fabricate more and more complex nanocomposite vehicles that can meter release via multiple internal or external signals,148,178,179 it is hard to envision how such complex materials will ultimately see clinical use, particularly given the inherent challenges around correlating performance in vitro and in vivo as outlined in Section 4.2. Interactions between the various stimuli-control units may also produce negative interaction effects in addition to the positive synergistic effects typically reported in the literature that may be hard to predict outside of the complex physiological environment. As such, if clinical use is indeed the goal, there is a clear need to simplify the design and fabrication of externally activated devices. The use of more reservoir-based delivery strategies in which potential material−drug interactions are minimized and the delivery vehicle is clearly a device rather than a new therapeutic entity as well as materials that are already approved or are on the pathway to approval for other in vivo uses may be positive steps toward such simplifications. Aside from the inherent regulatory challenges with the complexity of most delivery vehicles, there is also a clear deficit of knowledge around the long-term fate of many of the typical inorganic nanoparticles used for external signal actuation. This is true even for routinely used activating agents such as SPIONs, which are to-date likely the most studied externally activatable nanoparticle from a clearance/degradation standpoint. While degradation is achievable for the majority of SPIONs via intracellular pathways (and, importantly, tunable in terms of rate according to the morphology and geometry of the nanoparticle used),180,181 at least a fraction of intact nanoparticles are observed in vivo one year after administration.182 As such, small changes in particle structure and/or surface chemistry may significantly affect the long-term fate and thus tissue responses associated with such materials in a way that needs to be further understood. Similarly, while gold nanoparticles of various sizes and shapes have been widely applied in a variety of externally activated delivery approaches,183 there is still controversy as to their ultimate fate and long-term toxicity in vivo.184 Even less is currently L


Perspective

Chemistry of Materials vitro experiments and precisely extrapolated for animal testing and clinical trials, ideally through predictive modeling frameworks of potential cytotoxicity,192 and (3) effective methods should be developed for in vivo monitoring of the long-term toxicity and pharmacokinetics of externally triggered drug delivery systems.112 While externally activated delivery vehicles pose additional challenges with respect to safety, the inherent capacity of many triggered delivery vehicles to be imaged also offers unique opportunities for better monitoring the fate of drugs and the drug carriers in vivo. For instance, Dai and co-workers studied the in vivo pharmacokinetics of Ag2S and PbS@CdS quantum dots using fluorescence in the NIR-II region and determined biliary clearance as their main clearance pathway,193,194 while Wang et al. combined NIR imaging probes with mesoporous silica by fabricating the SiO2@NaGdF4@mSiO2 nanoparticles to achieve real time tracking of the nanocarrier fate and drug release in vivo.114 More effectively leveraging these in vivo tracking technologies offers significant potential to address the real concerns over safety and efficacy in externally activated drug delivery systems.

Todd Hoare: 0000-0002-5698-8463 Author Contributions §

(S.S.S. and S.C.) Both authors contributed equally to the manuscript Notes

The authors declare no competing financial interest. Biographies Somiraa S. Said is an Assistant Professor in the Department of Pharmaceutics, Alexandria University, Egypt. Dr. Said was awarded her Ph.D. in 2016 from Western University, based on her work in studying the controlled delivery of growth factors from electrospun fibers for therapeutic angiogenesis. Somiraa joined the Hoare Lab as a Natural Sciences and Engineering Research Council postdoctoral fellow in 2017, developing magnetically actuated hydrogel nanocomposites and nanofibers for drug delivery and tissue engineering applications. Scott Campbell is a Senior Scientist at Avro Life Science, focussing on the development of novel transdermal systems for delivering a range of different drugs from antihistamines to antipsychotics. He completed his B.A.Sc. at Queen’s University in 2010 and his Ph.D. at McMaster University in 2016, where he worked on externally addressable smart drug delivery nanocomposite systems in Todd Hoare’s laboratory as a Natural Sciences and Engineering Research Council Vanier Scholar.

5. CONCLUSIONS In summary, the potential of externally triggered drug delivery vehicles to achieve on-demand, repeated, and reproducible dosing at a targeted site offers tremendous promise to reduce side effects, increase efficacy, and allow for precise tuning of drug therapies according to patient needs. In this context, externally triggered drug delivery strategies are an essential component toward enabling an era of truly “personalized” medicine in which localized drug doses can be dynamically designed according to the changing needs of each patient. However, in our view, there is still (at least in the broad sense) a minimal focus on clinical translatability in this field, particularly with regards to designing delivery vehicles that could practically pass a regulatory approval process, understanding the long-term biological fate of the components of ondemand delivery vehicles, and ensuring that the instrumentation needs around the activation of any externally triggered therapeutic match the practical clinical realities around the accessibility of such instruments. In this sense, while recent advances in materials chemistry have enabled the development of a rich variety of remote-controlled therapeutic strategies, this field would now particularly benefit from more direct involvement by clinicians to clearly link the new synthetic and physical chemistry innovations with the needs and limitations of clinical practice. Indeed, the well-documented challenges encountered with translating even simple nanomedicines to the clinic195 are likely to be even more challenging to overcome with more complex and multicomponent externally triggered delivery vehicles. Collaboration with both preclinical and clinical researchers with experience in overcoming such challenges is the best way forward to enable externally activated drug delivery systems make the leap from bench to bedside and deliver on their promise to improve patient comfort and outcomes.

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Todd Hoare is a Professor and Canada Research Chair in Engineered Smart Materials in the Department of Chemical Engineering at McMaster University. He completed his B.A.Sc. at Queen’s University in 2001, his Ph.D. at McMaster University in 2006, and a postdoctoral fellowship in Robert Langer’s lab at the Massachusetts Institute of Technology in 2008. His lab develops hydrogels on multiple length scales for applications in medicine, agriculture, and personal care products: http://hoarelab.mcmaster.ca.

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ACKNOWLEDGMENTS The authors thank the J.P. Bickell Foundation (Medical Research Grant Program), the Vanier Scholarship program (S.C.), and the Natural Sciences and Engineering Research Council of Canada (PDF fellowship to S.S.S. and Discovery Grant to T.H.) for funding.

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REFERENCES

(1) Edelman, E. R.; Kost, J.; Bobeck, H.; Langer, R. Regulation of drug release from polymer matrices by oscillating magnetic fields. J. Biomed. Mater. Res. 1985, 19, 67−83. (2) Bae, Y. H.; Okano, T.; Hsu, R.; Kim, S. W. Thermo-sensitive polymers as on-off switches for drug release. Makromol. Chem., Rapid Commun. 1987, 8, 481−485. (3) Miyazaki, S.; Yokouchi, C.; Takada, M. External control of drug release: controlled release of insulin from a hydrophilic polymer implant by ultrasound irradiation in diabetic rats. J. Pharm. Pharmacol. 1988, 40, 716−717. (4) Kost, J.; Leong, K.; Langer, R. Ultrasound-enhanced polymer degradation and release of incorporated substances. Proc. Natl. Acad. Sci. U. S. A. 1989, 86, 7663. (5) Hauser, A. K.; Wydra, R. J.; Stocke, N. A.; Anderson, K. W.; Hilt, J. Z. Magnetic nanoparticles and nanocomposites for remote controlled therapies. J. Controlled Release 2015, 219, 76−94. (6) Campbell, S. B.; Hoare, T. Externally addressable hydrogel nanocomposites for biomedical applications. Curr. Opin. Chem. Eng. 2014, 4, 1−10. (7) Brudno, Y.; Mooney, D. J. On-demand drug delivery from local depots. J. Controlled Release 2015, 219, 8−17.

AUTHOR INFORMATION

Corresponding Author

*(T.H.) E-mail: [email protected]. ORCID

Somiraa S. Said: 0000-0002-7347-1319 M


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