On-demand drug delivery from local depots

    On-demand drug delivery from local depots Yevgeny Brudno, David J. Mooney PII: DOI: Reference:

S0168-3659(15)30113-9 doi: 10.1016/j.jconrel.2015.09.011 COREL 7851

To appear in:

Journal of Controlled Release

Received date: Revised date: Accepted date:

15 July 2015 8 September 2015 8 September 2015

Please cite this article as: Yevgeny Brudno, David J. Mooney, On-demand drug delivery from local depots, Journal of Controlled Release (2015), doi: 10.1016/j.jconrel.2015.09.011

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Yevgeny Brudno , David J. Mooney * [1] Wyss Institute For Biologically Inspired Engineering, Harvard University; Boston, MA. 02115 [2] School of Engineering and Applied Sciences, Harvard University; Cambridge, MA. 02138 * E-mail: [email protected]

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Graphical Abstract:

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Abstract Stimuli-responsive polymeric depots capable of on-demand release of therapeutics promise a substantial improvement in the treatment of many local diseases. These systems have the advantage of controlling local dosing so that payload is released at a time and with a dose chosen by a physician or patient, and the dose can be varied as disease progresses or healing occurs. Macroscale drug depot can be induced to release therapeutics through the action of physical stimuli such as ultrasound, electric and magnetic fields and light as well as through the addition of pharmacological stimuli such as nucleic acids and small molecules. In this review, we highlight recent advances in the development of polymeric systems engineered for releasing therapeutic molecules through physical and pharmacological stimulation.

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Keywords drug delivery; depot, polymer, therapeutic, on-demand, ultrasound, magnetic, electric, light, pharmaceutical, nucleic acid, stimulated Introduction Regulated drug presentation made possible through on-demand control over drug release could substantially improve current methods to deliver drugs from local depots. Macroscale drug-delivery devices are gaining attention in medicine due to their ability to exert spatiotemporal control over drug availability locally at a disease site. These drug-releasing depots usually take the form of a polymer or device implanted at a site of need, and release drugs locally. Local, controlled release technology presents a number of advantages over traditional drug administration. The first advantage is finer control over drug concentration at disease sites than is typically possible through oral or intravenous drug dosing. The second advantage is that peripheral, off-target side effects can potentially be avoided because drugs are presented locally. Drug depots can also release drugs for long periods of time (weeks to months), obviating the need for patients to repeatedly dose themselves or come in for a doctor’s visit, thereby helping to improve patient compliance. Examples of currently employed drug delivering depots include drug-eluting stents for the treatment of cardiovascular disease (e.g., the Cypher® stent), Gliadel™, an implantable wafer capable of controlled release of chemotherapeutics, the INFUSE® bone graft, which releases growth factors to stimulate bone growth, and various vaginal and subcutaneous implants for stable, long-term contraception release. The clinical and commercial success of controlled-release devices now opens the opportunity for the next step in controlled drug delivery, the ability for real time control over the timing and dosing of a locally-released therapeutic molecule. The vast majority of drug-releasing depots that are commercially available or currently in development rely on the intrinsic physical and biological environments surrounding the depot to control drug release, but this may not always be appropriate. Most drug-releasing depots rely either on the diffusion of drug out of the depot or the use of local pH, hydrolysis, and enzymatic cleavage to degrade the depot and release the drug. However, the local environment of the disease may dramatically impact these processes, and

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can vary from patient to patient. Further, the release rate in these systems cannot usually be modified after placement. In the treatment of many disease, the required release kinetics and drug dosing may be difficult to predict a priori. In these cases, drug dosing will need to be actively managed. Thus, many diseases could benefit from regulated drug presentation made possible through extrinsic control over drug release. The use of externally applied signals in the form of physical forces or administered signaling molecules could allow for drug presentation at the ideal time and at the ideal dose needed for a specific disease application in a particular patient. For example, on-demand drug delivery systems could allow patients to determine the location, timing and extent of drug release to alleviate pain. In some applications, pulsatile release of therapeutics could result in improved treatment of disease, such as the treatment of diabetes via regulated insulin delivery. Additionally, pulsatile release of therapeutic molecule could more effectively recapitulate innate cyclic biological processes triggered by factors such as growth hormones, sex hormones or their mimics.

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This review covers recent developments in methods of controlling the timing and spatial delivery of clinically relevant compounds from depots using externally provided stimuli. The first part of the review will cover therapeutic releases triggered using externally applied physical stimulation, such as magnetic and electric fields, ultrasound and light. The second part of this review covers methods to control drug release through the application of chemical triggers or through the repeated refilling and release of therapeutic molecules from a depot.

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Triggering release with physical stimulation Polymeric systems capable of releasing therapeutic molecules in response to physical stimulation offer a promising avenue for controlling the timing and dose of drugs released locally at a disease site. Some physical stimuli, such as ultrasound and magnetic fields, can penetrate deep into human tissue, allowing for control over drug release even at polymers placed deep within a patient. Additionally, the ability to release therapeutics at a distance may obviate the need for certain invasive procedures. The following sections will review advances using ultrasound, electric and magnetic fields as well as light to control the release of therapeutic molecules from polymeric depots.

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Ultrasound-triggered drug delivery Ultrasound has been traditionally used in medicine as a diagnostic method for observation of internal tissues and therapeutically for its ability to deliver acoustic energy to the body (e.g., by physiotherapists)[1, 2]. Recently, however, ultrasound has been applied to stimulate drug delivery. Ultrasound stimulation with high spatiotemporal resolution can be delivered to tissues of interest deep within a patient. As early as 1989, Langer and coworkers reported that ultrasound accelerates release of small molecules and proteins from polymer matrices [3]. Both biodegradable and non-erodible crosslinked polymer matrices are effective vehicles for on-demand delivery, and the ultrasound causes negligible temperature change and only small changes in convection within the sample, suggesting that enhanced degradation of matrix – rather than increased temperature or enhanced mixing – is mostly responsible for drug release. Ultrasound-mediated degradation permanently damages the delivery matrix in most situations, altering the release rate of therapeutics and potentially accelerating the release achieved with subsequent ultrasound application (Figure 1A). High intensity ultrasound induces polymer strand breaking of both synthetic [3] and biological[4] polymers, and two mechanisms have been suggested to explain this behavior. The first mechanism involves the strong shear forces arising from cavitation collapse, leading to physically-induced rupture of polymer strands. The second proposed mechanism involves the generation of free radicals as a result of local heating, and implicates free-radical-mediated breaking of polymer strands [5] as the major mechanism behind depot

degradation. Whichever mechanism predominates, the resultant lower polymer molecular weight decrease depot cohesion. In an analogy to sonophoresis, in which ultrasound is applied to transiently disrupt skin for transdermal drug delivery[6], recent advances have focused on self-healing polymer matrices, capable of fully reforming gel links after removal of ultrasound stimulation (Figure 1B). One approach to self-healing polymers useful for ultrasound-mediated delivery involves the introduction of a hydrophobic ultrasoundresponsive membrane covering a monolithic gel. The coating is disrupted by ultrasound to release drugs

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and re-forms to return to a low level of baseline release [7]. The coating, composed of methylene carbon chains, was shown to provide a strong barrier to the release of a charged antibiotic as well as active insulin. These kinds of re-sealable coatings may prove effective for on-demand therapy, especially if limitations such as biofouling can be effectively dealt with. Alternatively, injectable monolithic polymer matrices have been developed with encapsulated therapeutics which respond to ultrasound and subsequently self-heal. Epstein-Barash et al. took advantage of hyaluronic acid (HA) gels formed through reversible hydrazone bonds to encapsulate and release dye molecules from liposomal formulations[8]. The injectable gel allows for the incorporation of ultrasound-sensitive microbubbles and shows efficient in vivo delivery over a period of one week post-implantation. The first report of therapeutic use of ultrasound-mediated drug delivery from a depot was reported by Huebsch et al. in which the shear stress upon application of ultrasound disrupts the ionic crosslinks in alginate gels, allowing for enhanced release of the chemotherapeutic mitoxantrone in an in vivo cancer model [9]. The persistence of calcium ions in the gel and surrounding tissues allows the ionic cross-links to reform after cessation of ultrasound. The high-dose “bursts” of drug exposure made possible by ultrasound application enhances the toxicity of mitoxantrone toward triple-negative breast cancer cells in vitro, and xenograft tumors in vivo. Importantly, ultrasound application played a predominant role in drug presentation, as control tumors implanted with hydrogels carrying drugs, but not subjected to ultrasound grew at a significantly faster rate compared to tumors treated with ultrasound-stimulated gels. Self-healing bulk gels were also shown to efficiently release protein therapeutics (SDF-1) and DNA plasmids, without significant degradation of these sensitive biological molecules (Fig 1C). The same alginate bulk gels also demonstrate efficient release of gold nanoparticles in response to ultrasound stimulation[10]. The nanoparticles sizes (30-100nm) significantly exceeded the average gel pore size (~5nm)[11] leading to nearly undetectable levels of nanoparticle release without stimulation. Ultrasound stimulation induced transient gel degradation and release of BMP2 presenting nanoparticles to augment osteogenesis.

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A wide range of polymer matrices and hydrogels have been used with ultrasound mediated drug delivery, including poly-lactide-go-glycolides, alginates, hyaluronic acid, and methacrylates. These and other polymers differ in their susceptibility to physical and chemical degradation in response to ultrasound as well as in their ability to “heal” from an ultra-sound induced cavitation. Additional work remains to be done to identify physical and chemical properties of materials that make them well suited to ultrasoundmediated drug release. Despite the promise of ultrasound-mediated drug delivery, some challenges remain in this approach to triggered drug delivery. Efficient polymer disruption is still dependent on internal cavitation from dissolved gases within the matrix or the incorporation of long-lasting microbubbles, at least in certain situations[1214]. Kost et al. reported that polymer matrices within rigorously degassed buffers have significantly lower

Figure 1: Ultrasound-mediated therapeutic delivery. A: Ultrasound can mediate the fracture of polymeric depots, releasing entrapped cargo. B: Self-healing gels can release cargo in response to ultrasound and will re-gel, trapping unreleased drugs. This healing allows for repeat stimulation without permanent depot breakdown.

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ultrasound-induced drug release[3], highlighting the importance of small amounts of dissolved gases for on-demand drug release. Epstein-Barash et al. purposefully incorporated microbubbles into hydrogels, but this encapsulation translates to enhancement in drug delivery only in the first three days after gel implantation. These findings raise important questions about the need to reproducibly and stably incorporate dissolved gasses or stabilized microbubbles into clinically-relevant systems; the development of long-lived microbubbles is likely to be a challenge to efficient ultrasound-mediated delivery [3, 12, 14].

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Magnetic-stimulated Drug Delivery Application of magnetic fields may also enable one to trigger drug release from a depot, and is under active investigation. One approach involves the synthesis of nanocarrier heterostructures with drugcontaining reservoirs and magnetic-stimulated particles. Exposure to high-frequency magnetic fields either introduces grain boundaries or dissolves whole sections, releasing bursts of encapsulated compounds. Several excellent reviews on magnetically actuated drug delivery have recently been published. These reviews cover the use of magnetic nanoparticles[15] and magnetic liposomes[16] in therapeutic applications. This review will therefore focus on the incorporation of magnetically responsive elements into macroscopic drug depots for on-demand therapeutic release.

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Ferrogels made from magnetic particles embedded in polymer gels have attracted attention due to the promise of combining a relatively safe physical stimuli with magnetic particles generally considered biocompatible [17-24]; these approaches have progressed to clinical trials with certain systems [25, 26]. Magnetic oxide nanoparticles present favorable biocompatibility, and have been widely applied in applications such as magnetic resonance imaging (MRI)[27-31], cell tracking in vivo[32-35], and hyperthermia treatments [36-38]. Iron oxide nanoparticles are readily synthesizable and have predictable physical properties [39, 40]. Ferrite and maghemite forms of iron oxide have gained favor in the field due to their ready availability and precise control over composition and size. Finally, iron oxide nanoparticles can be chemically functionalized with biological ligands or coated with polymers.

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A wide range of nanoparticle sizes can be used for these systems, with sizes ranging from 5nm to 500nm, with the most appropriate size dependent on the mechanism of drug release. When Liu et al. used magnetic fields to align particles and prevent drug diffusion, 500nm particles led to the greatest level of drug release [21]. However, the vast majority of studies limit particle sizes to a 520nm range. In an application where magnetic-moment flipping is important (i.e. use of high frequency magnetic fields), a size limit of <30nm is generally acknowledged to achieve superparamagnetism [41]. Moreover, the use of particle sizes of less that 20nm reduces eddy currents and limits the heating of the particles [41].

Figure 2: Magnetic-stimulated drug release. Macroscale polymers can be embedded with iron nanoparticles to make the magneto-responsive. A: In the presence of a high-frequency magnetic fields, the polymeric system heats up, leading to gel melting and drug release. B: A constant magnetic field gradient can induce polymer deformation with expulsion of water and drug molecules with each application.

One mechanism for magnetic field stimulated drug release involves high-frequency magnetic pulses that are converted to heat when applied to nanoparticles, releasing drugs through local de-gelling (Figure 2A). This de-gelling can be mediated through dissociation of bound nucleic acid strands[42] or through melting of temperature-

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responsive polymers[43], which undergo a phase or volume transition at a critical temperature. Polymers displaying lower critical solution temperatures (LCST) slightly above that of body temperature are of particular interest for this mode of magnetically-actuated drug delivery due to the promise of coupling drug delivery to hyperthermic cell ablation or to the induction of a heat-shock response. Increases in temperature lead to matrix collapse and a fast release of trapped molecules. Poly(N-isopropylacrylamide) (PNIPAM) composites can be tuned to display an LCST in the 30-40 degree range. However, the collapse of PNIPAM hydrogels under sustained magnetic fields need not always release incorporated drug. In some cases, hydrogel collapse has been shown to suppress drug release due to reduced diffusion through the gel[44]. Pulsed AC magnetic fields can also effect burst release of drugs from PNIPAM ferrogels through heat-induced melting of PNIPAM disk-shaped ferrogels. Temperature-sensitive biological polymers, such as gelatin, can be harnessed in the same way, in which high frequency magnetic field-mediated heating causes local unwinding of the gelatin triple-helix, leading to on-demand release of vitamin B12[24]. Similarly, magnetic nanoparticles embedded within PNIPAM microbeads allow magnetic fields to collapse the microbead shell and trigger release of compounds[45]. Finally, local magnetic-field-mediated heating can increase hydrogel degradation rates, accelerating, rather than actuating drug release[46]. These systems all can potentially enable post-implantation tuning of drugrelease kinetics on a patient-by-patient basis.

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On-demand drug delivery from a ferrogel depot through direct actuation is an alternative mechanism to provide on-demand release, without depending on temperature changes in the system (Figure 2B). This approach was first reported through the incorporation of a millimeter-scale magnetic bar in a polyvinyl alcohol gel. Baseline release rates of small molecule and peptide drugs increased by five- to ten-fold under magnetic activation in vitro, and short term control of glucose levels through insulin release was demonstrated in vivo[18, 20, 47]. A number of subsequent reports have extended the technology to the release of model compounds and therapeutic agents in injectable ferrogels[48] as well as the extension of ferrogels to incorporate biological polymers[49]. One limitation to the use of ferrogels is limited gel deformation upon application of the magnetic field. Since pore sizes in most ferrogels, as in many hydrogels, are in the nanometer range, the gels resist compression and consequently the transport of large molecules and cells through the gels is very slow. In order to create scaffolds sensitive to magnetic fields, capable of on-demand drug and cell delivery, Zhao et al. fabricated ferrogels with threedimensionally connected macropores. The use of macroporous, rather than nanoporous, ferrogels results in large gel deformations upon stimulation with a magnetic field, leading to more exaggerated and rapid carrier collapse. The resultant convection accelerates release of encapsulated therapeutics, including chemotherapeutic drugs, plasmid DNA, proteins and stem cells[50-52]. Significant progress has been made to advance magnetic field-triggered release towards therapeutic use, though challenges remain. The availability of clinically-compatible magnetic nanoparticles is an advantage to this approach, but the biocompatibility of polymer matrices with magnetic actuation still needs further study. The use of magnetic stimulation for heat induction and delivery of a therapeutic compound, rather than the more typical use of magnetic field to induce cell ablation, limits the dosing to relatively small temperature changes (3-5 degrees Celsius) and thus potentially limits the dosing achievable. Additionally, the heat-inductive drug delivery systems must be stable to natural temperature changes in patients, such as those resulting from fevers. Additionally, both heat-induction and physical compression of polymeric systems induces local physical stresses, and devices will need to be designed that do not suffer from significant degradation over numerous induction rounds as that could unpredictably change device functionality. Electric-stimulated drug delivery Similarly to magnetic field-responsive drug depots, electro-responsive drug depots have been used for delivery of therapeutic molecules from gels matrices. Drug delivery from wirelessly-controlled, multireservoir microchip systems is a promising area and is in clinical trials; this work is reviewed elsewhere [53, 54].

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The main mechanism by which drugs are released in these systems involves the forced convection of drug out of the gel as it de-swells in response to imposed electric fields (Figure 3A). Polyelectrolyte hydrogels under an electric field generally undergo de-swelling as water is synersed from the gel[55]. Common medical equipment allow for precise control over the magnitude of electric field pulses, duration of pulse and interval between pulses. Additionally, a large body of literature exists cataloging safe levels of electrical field strengths, and commercially approved iontophoresis devices make electrical fields accessible in the clinic. Much of the field has focused on electro-responsive hydrogels because they resemble biological tissue and can encompass large amounts of water or biological fluids, allowing them to swell significantly. The large water content of hydrogels (up to 99%) allows for efficient transport of electrical ions in response to an electric field. Electro-responsive gels typically contain repeat monomer units that are ionized at neutral pH, and thus have a high number of a single charge spread out throughout the polymer backbone. Examples of poly-anionic polymers used include hyaluronic acid, hydrolyzed polyacrylamide, alginate, chondroitin sulphate and agarose. In comparison, electricallyresponsive gels made from poly-cationic or neutral polymers are relatively rare[56].

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The amount of drug released in response to application of an electric field tends to correlate closely to the loss of gel mass[57-60]. Responsive polyelectrolyte gels synthesized from hyaluronic acid or chondroitin are particularly common for the release of small molecule drugs and biological molecules [61-63], and allow a pulsatile “on” and “off” profile for drug release. The size, hydrophilicity and charge of released drug, as well as the drug’s affinity for the encompassing hydrogel play an important role in determining release profiles; one must also take into account the possibility of backflow of drug into the gel upon cessation of electric field and re-swelling[57, 59, 64]. In some systems, application of the electric field induced de-swelling only at the surface of the hydrogel, leading to a suppression, rather than induction, of drug release[65, 66]. For example, hydrocortisone release from calcium-crosslinked alginate hydrogels could can be both enhanced and diminished, depending on the system design. The rate of de-swelling, and thus drug release, is often limited because the shrinkage of the gel is diffusion controlled, meaning that equilibrium is reached slowly as fluid diffuses out of the gel. In most cases, the gels used in these systems are nanoporous and well solvated, creating stiff gels that resist compression.

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In theory, gel de-swelling is reversed once the electric field is turned off. In practice, a number of variables may impact the efficient reversibility of gel de-swelling. First, the viscoelasticity of the hydrogel will control the rate of reswelling, and this could be slow enough to appear non-reversible. Additionally and as discussed below, pH and chemical changes induced by the electric field and the presence of electrodes can lead to irreversible gel degradation[67].

Figure 3: Electric field-stimulated drug delivery. A: Polyelectrolyte gels de-swell towards oppositely charged nodes in the presence of electric fields. During de-swelling, water and therapeutic compounds are expelled. The de-swelling is a passive, diffusioncontrolled process. B: pH and chemical changes accompanying an electric filed can degrade or de-gel the depot, leading to release of therapeutic compounds.

In some cases, rather than relying on deswelling, the electric field can induce gel erosion (Figure 3B). One example of gel erosion-controlled release utilizes a drug depot composed of two polymers that interact through hydrogen bonds or ionic crosslinking to form a gel. This complex formation is pH dependent, and upon a pH change caused by electric stimulation, the gel converts from a solid state to solution,

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resulting in the release of drug through the disintegration of the polymer complex [68, 69]. Disintegration of the gel occurs through surface erosion, and gel degradation can be designed to follow zero-order kinetics. Electric stimulation can also break weak electrostatic or reducible bonds at the surface of a device, allowing for release of small molecules and protein therapeutics[70, 71].

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Challenges to the clinical application of electric fields for drug delivery remain significant. Slow gel deswelling kinetics make practical application of electro responsive depots difficult, although this could be overcome by decreasing gel size or reducing gel stiffness through the use of macroporous hydrogels[72]. Beyond this concern, the Figure 4 Light-induced drug delivery. A: When stimulated with light, lightresponses of polymer absorbing molecules can heat up polymer depots, melting the polymer or matrices to electric fields are increasing diffusion and releasing therapeutics. B: Polymers linked through dependent on the precise photo-cleavable linkers can be degraded through light-mediated cleavage. C: Alternatively, photo-cleavable linkers can link therapeutic molecules to polymer spatial arrangement of the gels, the electrodes and the so that photo-irradiation releases molecules upon application of light. field gradient. Further, electro-responsiveness has been shown be affected by many factors, including charge density, polymer hydrophilicity, composition of aqueous medium, concentrations of electrolytes and the presence of ionizable molecules. How all of these properties interact with the in vivo environment and how electroresponsiveness, and thereby drug delivery, would change as a consequence of biofouling, immune response and local pH changes is a subject of on-going research. Some early studies in vivo demonstrated the ability of electric fields to release labeled glucose[73] and insulin[74], and provide proof of concept validation for use of electric stimuli for on-demand drug delivery from drug depots. However, outside of multi-reservoir microchip sensors, very few studies have reported the use of electrically responsive depots in animal models. Electrically-induced changes in local pH at the electrodes due to electrolysis of water[75] are an additional source of concern for translating this technology, as some therapeutics will be incompatible with sharp pH changes, and this could also induce the degradation of the polymer[76, 77]. Additional research and, crucially, significant in vivo testing is needed to fully explore the potential of electro-responsive drug depots as viable on-demand drug delivery systems. Light-triggered drug delivery The use of light to trigger drug release is an especially attractive possibility as it can potentially be remotely applied with high spatial and temporal precision. This stimuli is tunable across a wide range of parameters, including wavelength, intensity, and duration, which may be adjusted to afford desired drug delivery profiles. Although radiation below 650 nm cannot penetrate deeper than about one centimeter into tissue due to high scattering and endogenous absorption, NIR light (650–900 nm) can penetrate up to one decameter into tissue and causes minimal tissue damage at the site of application. Excellent

ACCEPTED MANUSCRIPT reviews covering light-triggered drug release from nanoparticles have been published [78, 79] and this review will instead focus on light-triggered drug delivery from macroscopic depots.

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Multiple chemical transformations can be triggered with light to manipulate drug depots in situ (Figure 4). Polymers cross-linked with photo-labile groups can be adopted for on-demand drug release, as UV irradiation, for example, can release covalently bound drugs from polymeric materials created with photocleavable cores or surface functionalized with photoactivatable compounds[80, 81](Figure 4A). Selfimmolative chain fragmentation of dedrimers initialized by a single photo-cleavage event can release multiple drug molecules on-demand[80]. The strategy of photo-caging therapeutic compounds such as small molecule drugs[82-84] and nucleic acids[85-88] is a widely applied strategy for inducing on-demand release. Photo-caging involves the inactivation of a molecule through covalent conjugation of a key functional group to a photo-cleavable moiety. In cases where a photo-cleavable linker anchors the caged therapeutic, the caged molecule remains inert until light-induced photo-cleavage liberates and releases the therapeutic (Figure 4C). Reversible isomerization of chemical double bonds upon irradiation with near-UV and visible light can also be used to disrupt macroscale gel structure, leading to release of drugs. Such photo-isomerization reactions usually involve azobenzenes. Molecules such as azodiaryls transition from the trans to the cis confirmation with UV light, and isomerize to the trans through exposure to longer wavelength blue light. Liposomes and lipid membranes incorporating azobenzene groups can be induced to release drugs by taking on a more permeable conformation. Illumination by visible light isomerizes the bonds back to trans, closing the pores and inhibiting drug release[89]. Photo-isomerization of diaryl-aza compounds at the end of a pore can control the porosity of biological and synthetic channel structures and this has been applied to release chemotherapeutic drugs from depot pores into living cells. In addition to gel degradation and chemical isomerization, light irradiation can induce phase transitions in gels made of natural and synthetic polymers, accompanied by reversible volume changes[90]. These volume changes are induced through osmotic changes of liberated ions created during illuminations [90].

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Systems for light-induced drug delivery have focused on UV or blue light mainly because these higherenergy wavelengths promote the degradative and isomerization chemistry that translates to bulk changes in the properties of a drug carrier. As UV and blue light do not penetrate deep into the body, this has created a challenge to translate these approaches to disease therapy. The application of visible-lightresponsive hydrogels to date have focused on the incorporation of chromophores that convert light into elevated local temperatures, inducing gel melting or swelling and drug release[91]. Recent advances in harnessing far-red light have focused on photo-oxidative cleavage of small molecules[92], two photon absorbance[93] or plasmonic nanocavities to upconvert NIR light into usable blue light[94]. To date, the field has been hampered by a lack of an efficient photo-caging system capable of release triggered by near-IR light. Chemically-triggered drug delivery In contrast to the use of external physical cues, pharmacological stimuli created by chemical substances are governed by the pharmacokinetics and biodistribution of these molecules. The ability of chemical substances to diffuse throughout the body created an orthogonal stimulatory profile and potentially overcomes some limitations of therapeutic release stimulated by physical stimuli. The advantages of small molecules to stimulate on-demands release are several-fold. First, externally applied forces (magnetic fields, light) may have difficulty in accessing deep tissues, whereas chemical substances are able to diffuse to any site accessible to the blood. Second, induction by physical cues often requires sophisticated equipment to localize the signal which may not be widely accessible, requiring specialized patient visits and making it potentially difficult for patients to self-administer drugs in their home. Pharmacologically triggered drug release could allow for control over drug presentation, with the caveat that any such system is still subject to the limitations of systemically applied molecules, including their pharmacokinetics, metabolism, biodistribution and clearance. Nucleic acid stimulated release One common method for proof of concept pharmacological-induced drug delivery has been the use of nucleic acids as the stimuli. The use of DNA nanotechnology to stimulate drug delivery is promising due to the predictable nature of nucleic acid binding kinetics and the specificity of recognition between interacting partners. Additionally, nucleic acids can fold into secondary structures, aptamers, which

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specifically recognize and bind small molecule drugs. Nucleic acids have been used to disrupt drug depots in order to cause them to release their contents. Examples include inducing oligonucleotidecrosslinked hydrogels to swell or shrink through administration of complementary oligonucleotides[95, 96](Figure 5A). Ma and co-workers published other methods for the selective release of drugs in response to the application of oligonucleotides[97]. These methods mainly take advantage of nucleic acid recognition to bring a conjugated (and inactive) drug in close proximity to a catalyst capable of releasing the drug. Batting et al [98] took the concept one step further through the use of aptamers that selectively bind biotherapeutic growth factors. The growth factors are released through the addition of complementary nucleic acids that disrupt aptamer structure (Figure 5B). The use of a nucleic acid trigger to induce spatial deformation and allow drug release has been additionally used to open pores in mesoporous silica particles [99]. Several researchers have taken advantage of the significant structural changes upon aptamer-substrate binding to create releasable systems based on this approach. The ATP-aptamer has proven particularly fruitful [100-103] due to the significant steric changes the aptamer undergoes upon binding its target.

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Although modifying nucleic acids to improve stability can often disrupt aptamer recognition or fine-tuned DNA binding kinetics, serumstabilized nucleic acids can be advantageous for specific in vivo recognition. Consequently, Figure 5: Nucleic acid-mediated drug delivery. A: Nucleic acid-bound drug-carrying nucleic acid hydrogels can be disrupted through strand displacement by competing nucleic mimics such as morpholinos acids, releasing encapsulated therapeutic cargo. B: aptamer-therapeutic and peptide nucleic acids have molecule complexes can be disrupted by complementary nucleic acids, been used to recognize cancer releasing bound therapeutic cargo. C: Nucleic acid-mediated refilling of cells “pre-targeted” with polymeric depot. An implanted depot captures nanoparticle refills from the antibodies carrying circulation, allowing temporal control over therapeutic presentation. complementary nucleic acids[104]. In a variation on the theme of pre-targeting a site for nucleic acid recognition, Brudno et al. reported the development of refillable drug delivery devices that take advantage of nucleic acid binding to temporally control drug presentation (Figure 5C)[105]. In this system, a tumorally embedded nucleic acid-functionalized hydrogel locally releases chemotherapeutic drug. Once the hydrogel releases the drug, the gel serves as a homing beacon for drug refills administered systemically. The device can be refilled multiple times and this approach has shown efficacy over nanoparticle delivery systems in a mouse tumor model of triple-negative breast cancer. Many significant challenges remain to using oligonucleotide structure to provide dynamic drug delivery. First and foremost, many of the systems use nucleic acids selected in their native DNA or RNA state. These nucleic acids suffer from rapid in vivo degradation and may serve as potent immunostimulators. It bears notice that many of the anti-cancer systems use the intercalation of anthracyclin drugs into CpG islands[106, 107], the very sequences most likely to stimulate the immune system[108]. Moreover, nucleic acid formulations may be incompatible with chemotherapeutic drugs specifically designed to interact

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irreversibly with DNA, such as cisplatin and possibly doxorubicin[109]. The use of modified nucleic acids, such as 2’-fluro and 2’-methoxy, as well as more intense modifications such as morpholinos and peptide nucleic acids (PNAs) may overcome these problems, but with few exceptions[110, 111] aptamer selection and dynamic DNA nanotechnology has been limited to unmodified nucleic acids. Likely due to these issues, oligonucleotide-mediated drug releasing technologies have only rarely been tested in in vivo disease settings.

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Small molecule triggered drug release Rather than use nucleic acids, researchers can use small molecules to effect on-demand drug delivery from drug depots. Webber and colleagues have reported a number of promising hydrogel formulations capable of in situ decomposition in response to the presence of orally bioavailable pharmacological substances (Figure 6A, B). This system utilizes a polymer-bound bacterial gyrase subunit B (GyrB) gelled through the addition of the dimeric antibiotic coumermycin, resulting in hydrogelation. Addition of the monomeric, clinically-approved antibiotic novobiocin destabilizes the gel, leading to release of gelencapsulated drugs such as VEGF[112] or a vaccine formulation[113]. Similarly, FK506 will dissociate hydrogels formed through the interaction of homodimeric FK-binding protein 12 domains to release VEGF[114]. Tetracyclin can disrupt the interaction between TetR, a bacterial transcription factor and the tetO DNA motif to release interleukin 4[115].

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Instead of actuating the decomposition of a gel, small molecule reactivity can directly lead to the release of drugs from pro-drug formulations through clever use of copper-free click chemistry (Figure 6C). Copper-free bioorthogonal chemistry[116], whether through Staudinger ligation, azide cyclooctyne chemistry or the newer tetrazine – alkene reaction, have been used in vivo to target small molecules to specific cells[117], promote antibody “pretargeting”[116, 118, 119] and target drugs and small molecules to disease sites[120, 121]. Through taking advantage of the conversion of azides to amine during the biocompatible Staudinger reaction, van Brakel et al. report the liberation of doxorubicin upon addition of triphenylphosphine[122]. The slow kinetics of Figure 6: Small molecule-mediated drug delivery. Polymeric matrices linked through Staudinger ligation and protein-protein or protein-small molecule bindings can be disrupted through incomplete application of small molecule to disrupt protein duplexes formed around a small molecule dimer (A) or a specific disruptor of protein-protein interactions (B). Polymer biocompatibility of triphenylphosphine likely disruption leads to therapeutic release of entrapped therapeutic molecules. C: limit the utility of this Bioorthogonal, “click,” chemistry can be used to target therapeutic molecules to system in vivo. polymeric depot and release them on site. However, a more recent report from the same

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The use of pharmacological substances, whether nucleic acids or small molecules to specifically control the release of therapeutic agents is still very much in its infancy. Many challenges remain for the technology, including the need to demonstrate for many technologies that they can obtain favorable in vivo reaction kinetics, PK/PD profiles, and exhibit appropriate biocompatibility to make them strong candidates for therapy in disease models.

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Conclusion There has been much exciting progress in the use of physical and pharmacological stimuli to effect ondemand drug release from macroscale depots. In order to be truly translatable to a clinical setting the depot must likely confirm to several rules. First, the depot should be non-toxic and a non-irritant when implanted in vivo, and should elicit a minimal physiological response except for those explicitly designed into its function. Second, the depot should have a biodegradation profile tuned to the specific application to stably persist in the implanted location for as long as necessary, but slowly degrade into safe subunits when no longer needed. Finally, the carrier should be easy to administer and ideally would be readily refillable for repeat on-demand administration. The field is steadily progressing to reach these requirements, and it is an exciting time for on-demand drug delivery research.

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Acknowledgments This work was supported by the Wyss Institute for Biologically Inspired Engineering. Y.B. gratefully acknowledge funding support from the Wyss Technology Development Fellowship.

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