Theranostic nanosystems for targeted cancer therapy

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Review

Theranostic nanosystems for targeted cancer therapy Homan Kang a,1 , Shuang Hu a,b,1 , Mi Hyeon Cho c , Suk Ho Hong c , Yongdoo Choi c,∗ , Hak Soo Choi a,∗ a b c

Gordon Center for Medical Imaging, Department of Radiology, Massachusetts General Hospital and Harvard Medical School, Boston, MA, 02114, USA Department of Nuclear Medicine, West China School of Medicine, West China Hospital, Sichuan University, Chengdu, Sichuan, 601141, China Biomarker Branch, National Cancer Center, 323 Ilsan-ro, Goyang, Gyeonggi, 10408, South Korea

a r t i c l e

i n f o

Article history: Received 2 July 2018 Received in revised form 26 October 2018 Accepted 2 November 2018 Available online xxx Keywords: Targeted therapy Functional nanoparticle Organ-specific targeting Diagnostic imaging

a b s t r a c t Nanomaterials have revolutionized cancer imaging, diagnosis, and treatment. Multifunctional nanoparticles in particular have been designed for targeted cancer therapy by modulating their physicochemical properties to be delivered to the target and activated by internal and/or external stimuli. This review will focus on the fundamental “chemical” design considerations of stimuli-responsive nanosystems to achieve favorable tumor targeting beyond biological barriers and, furthermore, enhance targeted cancer therapy. In addition, we will summarize innovative smart nanosystems responsive to external stimuli (e.g., light, magnetic field, ultrasound, and electric field) and internal stimuli in the tumor microenvironment (e.g., pH, enzyme, redox potential, and oxidative stress). © 2018 Elsevier Ltd. All rights reserved.

Contents Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 00 Targeting strategy from organ to cell: enhanced targeted therapy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 00 Delivery to the target site: organ-specific targeting . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 00 Delivery to the subcellular and organelle target: Cellular-specific targeting . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 00 Internal stimuli-responsive nanosystems (chemical and biological triggering) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 00 pH gradients . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 00 Enzyme reactions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 00 Redox potential differences . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 00 Reactive oxygen species generation . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 00 ATP-responsive nanosystems . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 00 External stimuli-responsive nanosystems (physical triggering) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 00 Light irradiation . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 00 Photoinduced drug release . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 00 Photothermal therapy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 00 Photodynamic therapy (singlet oxygen generation) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 00 Magnetic fields . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 00 Ultrasounds . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 00 Electric fields . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 00 Ionizing radiation . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 00 Perspectives . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 00 Acknowledgments . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 00 References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 00

∗ Corresponding authors. E-mail addresses: [email protected] (Y. Choi), [email protected] (H.S. Choi). 1 These authors contributed equally to this work. https://doi.org/10.1016/j.nantod.2018.11.001 1748-0132/© 2018 Elsevier Ltd. All rights reserved.

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Introduction Over the past several decades, a variety of nanoparticles (NPs) with different biological functions have been developed for cancer imaging, diagnosis, and treatment by controlling their physicochemical properties including size, shape, composition, and surface functionalization. Among them, NPs that respond to external stimuli (e.g., light irradiation, magnetic field, ultrasound, and electric field) and/or changes in internal biological environment (e.g., pH, enzyme, redox potential difference, and oxidative stress) have received growing attention in the fields of drug delivery, cancer diagnosis, tissue engineering, and other biomedical applications [1–5]. These stimuli-responsive NPs have many advantages as drug carriers over small molecule drugs including greater loading capacity, protectability of the payloads from unwanted degradation, improved drug absorption into the target tissue, multivalent drug payload, and controlled drug release [1,2]. In addition, these NPs as versatile transducers can convert the form of stimuli into another physical quantity [3]. For example, magnetic field or light irradiation can convert such energy source to heat by using appropriately designed metallic or magnetic NPs. However, the lack of understanding in the fate of NPs, which is generally caused by the complexity of biological environments, the diversity of multifunctional NPs, or the lack of research in the field of Materials Science, results in not only less targetability but also aggravates adverse effects on organ, tissue and cellular levels in the body due to their unusual physicochemical properties. Recently, Wilhelm et al. reported that only less than 1% of the administered NPs reach to the site of interest from the survey of literature published during the past decade [4]. Additionally, the in vivo fate (i.e., biodistribution) and/or the clearance pathway of injected NPs have not been intensively investigated [5]. As shown in Fig. 1, therefore, physicochemical properties, such as hydrodynamic diameter (HD), shape, composition, hydrophilicity/lipophilicity, and surface characteristics, are the key design considerations of stimuli-responsive NPs to improve their delivery efficiency and consequences for diagnostic and therapeutic applications. In addition, the nanomaterials for targeted therapy should not only be suitable for the living body but also be capable of interacting with or releasing the therapeutic drugs in response to specific biological stimuli. Researchers can consider making the best selection for each domain such as triggering, imaging, targeting, and therapeutic domains to achieve enhanced targeted therapeutic NPs and their therapeutic efficacy. The details for this modular approach to the design of targeted therapeutic nanosystems are shown in Fig. 1. In this review, we discuss basic design considerations required for the stimuli-responsive NPs to enhance targeted therapy and review the recent state of the art of cancer nanotheranostics.

Targeting strategy from organ to cell: enhanced targeted therapy After systemic administration, the NP is immediately confronted with physiological barriers in the body, such as aqueous blood flow and absorption of cells/proteins and filtrations by the spleen, liver, and kidneys [6]. Consequently, the target efficiency of NPs can often be varied by nonspecific biodistribution and persistent background retention [7,8]. There is a need for NPs that avoid physiological barriers and effectively target only the diseased areas of the body. To achieve favorable targeting beyond biological barriers, a fundamental understanding of the interactions between nanomaterials and proteins/tissues in biological fluids is needed: Once injected, NPs in the blood should first reach the tissue/organ of interest (mode of action) and, sequentially, the drug/action to treat the target tissue must be able to interact with the target cells (mechanism of action)

[9]. This targeting strategy can enhance feasible targetability of NPs at the specific tissue and improve treatment efficacy (Fig. 1). Delivery to the target site: organ-specific targeting Reaching the target site of the NP is the first essential need for the drugs or nanosystems to be functional. If systematically administered NPs cannot reach the target site, they cannot function as designed at the cellular level. Enhanced permeability and retention (EPR) effect is commonly used for delivering the nanosized materials ranging with 10–400 nm in HD and favorable surface properties to accumulate in tumors, owing to leaky vascular structures and poor lymphatic drainage [10,11]. This is firmly associated with intrinsic physicochemical properties of NPs such as HD, charge, and hydrophilicity/lipophilicity as well as the physiological dynamic behavior of the injected NPs [5]. Although the complexity of this factor influences the EPR effect, it is very hard to describe multiple factors in the same basket, so each factor will be discussed one by one. The first and foremost advantage of the NPs in the diagnostic and therapeutic field is the ability to elongate blood circulation, which increases the probability of meeting tumoral tissues. For prolonged blood circulation, the surface charges and hydrophilicity/lipophilicity of NPs are the most important property. NPs normally disperse in aqueous solutions, then meet abundant proteins in the bloodstream including serum opsonin proteins after intravenous administration. NPs that are too charged and/or lipophilic result in undesired aggregations, protein binding, and opsonization, which causes NPs to accumulate in the liver and spleen within several minutes by the immune surveillance. Serum protein binding also results in HD increase of NPs in the blood. After evading the immunological response, the next design consideration of NPs is to have favorable clearance for background reduction [5]. It strongly depends on the pore sizes of fenestrated capillaries or sinusoids in the body blood vascular system. The size to be noticed is 5.5 nm, which is the critical threshold for rapid renal clearance of NPs [12]. Renal clearance, compared with the hepatobiliary clearance route, is preferred for diagnostic and therapeutic NPs because untargeted agents need to be rapidly eliminated from the body with limited cellular internalization/metabolism, thus effectively minimizing their exposure to the immune system [2,5]. Choi et al. reported the tissue- and organ-selective biodistribution and clearance of polyethylene glycol (PEG)-linked NPs [13]. By stabilizing the quantum dot core with various lengths of neutral and hydrophilic PEG ligands (n = 2–22), NPs can not only vary the size ranging from 4.5 to 16 nm in HD, respectively, but also avoid serum protein binding and evade immune surveillance. As a result, enhanced uptake of NPs was obtained at the specific location of the body such as the liver, kidneys, bladder, pancreas, and circulating vasculature. Therefore, after being stabilized in the serum, HD shows a profound effect on the biodistribution and clearance of NPs. This is very important to deliver the NPs to the specific organ in realistic disease models (e.g., orthotopic or genetically engineered tumor model) and, eventually, in the clinic. For example, as shown in Fig. 2, only <5% injection dose (%ID) of NPs is delivered to the brain tissue after administered intravenously, and an even smaller portion (<1%ID) can reach the target tissue, i.e., glioblastoma [14]. Intravenously injected doses, in general, first enter the central compartment including blood, heart, and lungs, then rapidly equilibrate into peripheral compartments such as the liver, kidneys, brain, etc., and finally spread throughout the whole body (slow equilibrating). The first requirement is to avoid unwanted accumulation in the major tissues and organs (heart, lungs, liver, and kidneys) by the biological or immunological barriers (reticuloendothelial system; RES). Then the injected

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Fig. 1. Modular approach to design an enhanced targeted therapeutic nanosystem. A therapeutic nanosystem can be composed of backbone, triggering domain, imaging domain, targeting domain, and therapeutic domain. The final physicochemical properties of a nanosystem are the pivotal modulator for biodistribution and clearance and thus enhanced targeted therapy.

Fig. 2. Schematic illustration of targeting strategy: Organ-specific targeting and cellular-specific targeting. Systemically administered nanosystems should overcome physiological barriers to reach the target organ efficiently (organ-specific targeting) and accumulate in the target site after confronting biological barriers at the cellular/subcellular level (cellular-specific targeting).

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molecules perfuse into the tissues and organs of interest. If the vast majority of injected doses are trapped in the central compartment, the chances are slim to deliver NPs to the site of interest (target). Especially, the brain is protected by a highly selective semipermeable membrane barrier called the blood-brain barrier (BBB), which limits the delivery of multifunctional NPs into the brain [14]. Therefore, as aforementioned in the introduction, most NPs cannot reach the site of interest when they are administrated intravenously and the delivery efficiency of NPs via the EPR effect is <1%ID [4]. In addition, the EPR effect is a crucial concept for solid tumor targeting, but the optimal size or size limitation of NPs is still unclear and controversial [15]. Charges or charge-to-mass ratio is also one of the main considerations for designing multifunctional NPs. For instance, zwitterionic NPs show only minimum adsorption with serum proteins resulting in a rapid systemic circulation and a low nonspecific uptake by the RES (liver, spleen, bone marrow, etc.) and then eliminated efficiently from the body. In contrast, positively charged NPs highly associate with serum proteins resulting in high uptake in the major organs [2,12,16]. As shown in Fig. 1, the physicochemical properties (size, charges, charge-to-mass ratio, and hydrophilicity/lipophilicity) of the final NPs with all 4 functional domains affect the biodistribution and clearance performance for enhanced targeted therapy.

Delivery to the subcellular and organelle target: Cellular-specific targeting After NPs reach the target tissue, the interactions with cells and biomolecules should enhance the retention of NPs on the tumoral membrane or delivery across cell membranes to the intracellular targets [17]. Thus, such active targeting is critical for enhanced targetability where NPs decorated with targeting ligands, such as small molecules, peptides, antibodies, and proteins, specifically bind to cell surface markers present or overexpressed on the extracellular matrix. Active targeting strategies tend to outperform the EPR based passive targeting strategy, affording delivery efficiencies of over 50% increase [4]. However, the physicochemical properties of the targeting ligands conjugated on the surface of NPs can change the overall surface property and HD of the NPs, which need to be carefully selected and the number of ligands should be limited [18]. Another design consideration is the intracellular interactions of NPs with biological barriers at the subcellular level (Fig. 2). This is directly associated with a determination of therapeutic efficacy. Drug carriers are not always helpful for an intracellular drug delivery, especially when the carrier NP is hydrophilic [6]. Therefore, an optimization of the drug release profile is required. For example, a small hydrophobic drug can easily penetrate cell membranes, thus NPs can release drugs at the extracellular space in tumor tissues, and these drugs can subsequently migrate from extracellular space to intracellular space. If the drug has to be transferred to the cell but cannot easily migrate into the cell, the NP can be modified with receptor ligands (for receptor-mediated endocytosis) or cellpenetrating peptide to increase the efficiency of delivery into the cell. Taken together, physicochemical properties (e.g., HD, shape, composition, hydrophilicity/lipophilicity, and surface characteristics) of NP can influence the efficacy of site-specific drug delivery; therefore, these properties should be considered for favorable organ/cellular-specific targeting [5]. In the following section, we describe recent examples of innovative stimuli-responsive nanosystems and their applications to cancer theranostics to achieve favorable tumor targeting and targeted cancer therapy.

Internal stimuli-responsive nanosystems (chemical and biological triggering) Changes in chemical and biological/physiological properties are useful for cancer diagnosis as well as designing stimuli-responsive nanomaterials. For example, tumor microenvironments such as pH gradient, cancer-related enzyme, redox potential difference, and hypoxic conditions can be attractive targets to manipulate for the triggering of the NPs. We discuss several representative examples of internal stimuli-responsive nanosystems and their therapeutic effects (Fig. 3). pH gradients Amongst many internal stimuli used to build environmentresponsive nanosystems, pH gradients have been a popular choice, specifically those that utilize either slightly acidic extracellular pH environment (pH ≈ 6.8) of solid tumor tissue or intracellular pH of cancer cells (i.e., endosomal and lysosomal pH = 5.5–6.0), compared with that in circulating blood (pH ≈ 7.4) [19–21]. In the first case, researchers take advantage of the fact that tumor microenvironment is more acidic than normal tissues, ranging from pH 6.5 to 7.2. This is due to the fact that rapidly growing tumor cells accumulate lactic acid because of their high glucose uptake combined with low rates of oxidative phosphorylation [19]. This phenomenon of high lactate production is also known as the Warburg effect. The reduced blood supply and lymphatic availability of many tumors also add to the acidity of tumor tissues [22]. Therefore, researchers were able to target tumor tissues using pH-responsive systems such that upon contact with low pH environment, the properties of material change and lead to cargo delivery. For example, we reported renal clearable nanocarriers (a.k.a. H-Dots) that deliver anticancer drug imatinib by the host-guest interaction with ␤-cyclodextrins (␤CDs) and releases the imatinib to the gastrointestinal stromal tumor (GIST) in the acidic tumor microenvironment (Fig. 4a) [2]. Since HDots could deliver drugs to the target efficiently with minimum to low background uptake, they successfully used to monitor drug delivery, targetability, pharmacokinetics, and therapeutic efficacy in both GIST-bearing xenograft mice and genetically engineered tumor models. More importantly, imatinib-loaded H-Dots exhibited lower uptake into the immune system, which improved tumor selectivity. Molecules that take on different conformations in different pH environments (due to pH-dependent hydrophobic-to-hydrophilic transitions) have been used for stimuli-responsive therapy in the acidic tumor environment [22]. Park et al. utilized glycol chitosan backbone-conjugated with chlorin e6 (Ce6, photosensitizer) and 3diethylaminopropyl isothiocyanate (pKb ≈ 6.8) for tumor therapy. The chitosan polymer changes its conformation from the coiled structure at a neutral pH to an uncoiled shape at the tumoral site and releases self-quenched Ce6 groups to recover photoactivity [23]. Another usage of pH trigger is to target the low pH of endosomes and lysosomes at the intracellular level. Park et al. synthesized negatively charged NPs composed of hyaluronic acid (HA)-polypyrrole (PPy) and the positively charged anticancer drug doxorubicin (Dox) or photosensitizers via charge-charge interactions [24,25]. When drugs were loaded in HA-PPyNPs, energy transfer from the drugs to PPyNPs was effective and resulted in turn-off of its fluorescence. The charge interaction between HA and Dox was altered in acidic pH after endocytosis to cancer cells, where HA is protonated and its negative charge is diminished. This triggered the release of the drugs from the NPs and spontaneous turn-on of its optical properties [24]. A “gatekeeping” mechanism by a pH-sensitive polymer layer is also used, as is the case in the work with alginate/chitosan multilayer wrapped around iron oxide/gold core and a mesoporous silica layer. Then, algi-

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Fig. 3. Schematic illustration of internal and external stimuli for triggering therapeutic effects of systemically delivered nanosystems.

Fig. 4. Nanosystems triggered by pH gradient and enzyme cleavage for targeted drug delivery. (a) pH-responsive theranostic NPs (H-dots) were composed of renal clearable polymeric backbone, drug delivery motif, and contrast agent. NPs were delivered to the tumor site by avoiding nonspecific uptake and anticancer drugs were released to the tumor cell via a pH gradient. Reproduced with permission from [2]. (b) Enzyme-responsive NPs were designed to be cleaved upon the contact with MMP-2, which releases the photosensitizer from gold nanorod and triggers photoactivity. Reprinted from [30] under (CC BY-NC 4.0).

nate/chitosan layer dissolves in an acidic environment and releases Dox and Ce6; this system was found to be mainly localized in lysosomes and therefore shown to be effective against tumor growth in vivo [26]. Additionally, drug conjugates with pH-sensitive linkers such as hydrazine, acetal, and ortho-ester have been widely used in the field of stimuli-responsive drug delivery [22]. These pH-sensitive NPs undergo faster hydrolytic breakdown at acidic pH relative to neutral physiological pH. Zhao et al. used a polyethylenimine

linked to PEG by a Schiff base, where the PEG group was cleaved off at the acidic pH, leading to the release of drugs (docetaxel and indomethacin) [27]. Enzyme reactions Enzyme-induced cleavages are generally used in bioresponsive systems when the enzymatic activity is associated with a particular tissue or when the enzyme has a higher concentration at a certain

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Fig. 5. Nanosystems triggered by redox and ROS generation for targeted therapy. (a) Chitosan-PEG copolymer coated NPs containing O6 -benzylguanine (BG), and intracellular BG release triggered by glutathione. Reproduced with permission from [36]. (b) H2 O2 -responsive release of photosensitizer and O2 for selective photodynamic cancer therapy. Reprinted with permission from [46].

site of tissue. Enzymes are also beneficial because enzymes themselves are a useful target for diagnostics, and are very efficient in catalyzing a reaction [28]. Tumor cells typically overexpress several enzymes, such as matrix metalloproteinases (MMPs), cathepsin B, and hyaluronidase (HAdase). These enzymes are used as targets [29]. Jang and Choi conjugated photosensitizers to gold nanorods via a protease-cleavable peptide linker [30]. The photoactivity of NPs was dramatically suppressed because its excitation energy was transferred to gold nanorods in its proximity. Once the linker was cleaved by MMP-2, photosensitizer was released from the surface of gold nanorods and their fluorescence and phototoxicity were regained thereafter (Fig. 4b). The efficacy of this probe was demonstrated on MMP-2 positive and negative cancer cell lines. In the following paper, Choi et al. also designed enzyme activatable graphene oxide (GO)- photosensitizer complexes, where Ce6-conjugated HA was physically adsorbed onto GO surface as an energy quencher [31]. This complex becomes highly fluorescent and phototoxic inside cancer cells, and the decomposition of HA by HAdase helps the release of HA-Ce6 fractions from the GO surface. In another example, cathepsin B-cleavable peptide linker was used to build a diblock copolymeric micelle [32]. Bcl-2 homology domain peptide containing middle block was cleaved by cathepsin B and released therapeutic drugs to the target cells, which induced target-specific apoptosis.

and non-phototoxic even upon light irradiation. GO-Ce6 exhibited fluorescence recovery and singlet oxygen generation only inside cancer cells, where glutathione initiates to activate the photosensitizer. This could be suited for selective fluorescence imaging and photodynamic therapy. Another example is a 7.5 nm iron oxidebased NP conjugated with tumor targeting peptide chlorotoxin through a redox-responsive disulfide linker (Fig. 5a) [36]. The surface of this NP was coated with the chitosan-PEG copolymer and functionalized with O6 -benzylguanine (BG). Under the reductive tumor intracellular condition, the NPs released therapeutic BG via disulfide cleavage, which blocked the DNA repair proteins and potentiated the cytotoxicity of temozolomide both in vitro and in vivo. This strategy could improve the efficacy of temozolomide chemotherapy for glioblastoma multiforme by overcoming the resistance derived from upregulation of DNA repair proteins. Separately, a redox-responsive nanoassembly of amphiphilic polyprodrug was prepared using 10-hydroxycamtothecin as a model drug [37]. The disulfide bonds of the self-assembled NPs were incorporated into each repeating unit of amphiphilic polyprodrug, which were attacked by intratumoral glutathione and the polyprodrug segment underwent a redox-responsive elimination reaction, resulting in a chain breakage patterned release of the intact drugs. Reactive oxygen species generation

Redox potential differences Redox potential change is one of the major signatures occurring at the tumor microenvironment due to the hypoxia and overproduction of reductive biomolecules, including reductase and glutathione [33–38]. To take advantage of this signature, reductively cleavable bonds such as disulfide and diselenide bonds have been developed to attain redox sensitivity [34,39]. The intratumoral region possesses high glutathione concentrations (2–10 mM) compared with the extracellular microenvironment (2–10 ␮M) and disease-free tissue, which have been used as an internal trigger to achieve redox-responsive drug delivery [35–37, 40]. NPs containing therapeutic molecules could be designed either through direct conjugation [36,40] or embedding [37] via selectively responsive chemical bonds or physical intermolecular interactions. Choi et al. designed nanosized graphene oxide (GO) conjugates with photosensitizer through a redox-responsive cleavable disulfide linker [40]. GO-Ce6 conjugate was non-fluorescent

The enrichment of reactive oxygen species (ROS) is another characteristic found in the tumor microenvironments [41]. The aberrant metabolism of rapidly proliferating cancer cells leads to significantly elevated cellular concentration of ROS such as hydrogen peroxide (H2 O2 ) and free radicals. The increased levels of ROS could be considered as not only biomarkers for diagnosis and prognosis of diseases, including cancer and inflammation but also triggers for the stimuli-responsive delivery [42]. Various ROSresponsive intelligent NPs have been developed in recent years for ROS-triggered drug release and imaging [43]. Most of them are based on either ROS-induced solubility switch [44] or ROS-induced degradation [45]. For example, Jeong et al. developed a chondroitin sulfate (CS)-anthocyanin (ATC)-based nanocomplex as an efficient ROS scavenging material and loaded Dox in the nanocomplex [44]. When Dox-loaded CS-ATC nanocomplex reacted with H2 O2 , the ATC was decomposed into two hydrophilic products such as 2,4,6-trihydroxybenzaldehyde and protocatechuic acid. It

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Fig. 6. Nanosystems triggered by ATP for controlled drug delivery. Based on the conformational change from the duplex to the tertiary aptamer structure, the intercalated anticancer drug (Dox) of this nanosystem could be selectively released inside the ATP-rich intracellular environment. Reproduced with permission from [50].

induced Dox release from the nanocomplex in the ROS-enriched tumor region. In another example, polymeric NPs of Ce6-HA conjugate showed peroxynitrite-mediated degradation of the polymer backbone, the release of the photosensitizers from the NPs, and subsequent turn-on in its fluorescence and phototoxicity, thereby enabling ROS-mediated selective fluorescence imaging and photodynamic therapy of activated macrophage cells [45]. To take advantage of O2 -generation from the reaction between catalase and H2 O2 , Chen et al. developed H2 O2 -activatable and O2 evolving NPs for selective and enhanced photodynamic therapy of hypoxic tumors [46]. Biodegradable polymeric NPs were loaded with methylene blue (MB) and catalase in the core, and black hole quencher-3 (BHQ-3) in the polymer shell, and functionalized with a tumor-targeting ligand c(RGDfK). Once intracellular H2 O2 penetrated the core of the NP, it was catalyzed by catalase to generate O2 gas, causing the rupture of polymer shell and subsequent release of MB. Then, the released MB molecules became highly phototoxic at the tumor environment. This NP platform can achieve H2 O2 responsive phototoxicity as well as increase O2 levels in hypoxic tumor region for selective and enhanced photodynamic cancer therapy (Fig. 5b). Recently, manganese dioxide (MnO2 ) NPs have also been used for H2 O2 -responsive theranostic materials [47,48]. MnO2 NPs have the ability to react with H2 O2 under acidic conditions and simultaneously produce Mn2+ ions and O2 , which greatly enhance T1 - and T2 -MRI performances for selective tumor imaging. Furthermore, O2 generation from MnO2 NPs enhanced radiotherapy by modulating the tumor hypoxic status. ATP-responsive nanosystems Recently, adenosine-5 -triphosphate (ATP) has been suggested as a new trigger for cancer-specific controlled drug delivery [49–53]. The concentration of extracellular ATP is considerably lower (<0.4 mM) compared to intracellular ATP (1–10 mM) and the ATP level in cancer cells is significantly higher than that in normal cells due to the increased glycolysis in cancer cells [49,54,55]. These

properties make ATP as a potential internal stimulus for selective cancer imaging and therapy. For example, Mo et al. developed Doxloaded polymeric nanocarriers functionalized with ATP-binding aptamers, in which the Dox intercalated with the aptamers could be selectively released from the NPs upon a conformational change from the duplex to the tertiary aptamer structure inside an ATPrich intracellular environment of cancer cells (Fig. 6) [50]. This ATP-responsive nanocarrier showed a 3.6-fold increase in the cytotoxicity and significantly enhanced tumor growth inhibition compared to the non-ATP-responsive nanocarriers. Using the ATPbinding aptamer moiety, Shen et al. reported an ATP-activatable nanophotosensitizer (Apt-HyNP/BHQ2 ) for selective fluorescence imaging and photodynamic cancer therapy [51]. Since the surfaces of the photosensitizer-loaded hybrid quantum dot NPs are conjugated with the BHQ2 -labeled ATP-binding aptamers, fluorescence emission and singlet oxygen generation from the NPs are efficiently inhibited by the BHQ quenchers on the surface. Fluorescence and phototoxicity of the nanophotosensitizers could be selectively turned on when the BHQs were dissociated from the aptamer inside the target cancer cells. While a few examples have been reported so far, various designs of ATP-responsive nanosystems are expected to be developed in the future. Compared with the other types of internal stimuli-responsive systems (e.g., redox-responsive disulfide bonds and pH-cleavable linkers), ATP-responsive nanosystems currently use relatively expensive and large aptamers (∼30 bases) as a basic unit to obtain the ATP-responsive property. Therefore, to achieve success from the benchtop to the clinic, we would need to overcome this main challenge of developing a simpler design that is more cost-efficient for the scale-up chemistry of ATP-responsive nanosystems. External stimuli-responsive nanosystems (physical triggering) Since there are too many biological/physiological factors that influence the therapeutic efficacy of multifunctional NPs in the

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Fig. 7. Nanosystems triggered by light irradiation for targeted drug delivery. (a) Light is absorbed by the gold nanocage and converted to heat, which opens pores and releases payload drugs by folding thermosensitive polymer chains on the surface of nanocage. Reproduced with permission from [60]. (b) NIR light triggers drug release from gold NPs. After cancer-specific targeting, NIR light generates heat on the drug-loaded Au/Ag hollow nanoshells, followed by releasing anticancer drugs into the cell that induces effective cell death. Reproduced with permission from [61].

body including tumor heterogeneity, researchers have been trying to use external stimuli to make assured changes to enhance drug delivery and therapeutic efficacy instead of using microscopic changes for targeting. Various external stimuli can be utilized such as light irradiation, magnetic field alteration, ultrasound sonography, and electric field treatment as illustrated in Fig. 3. This section describes some representative examples of external stimuli-responsive nanosystems and their therapeutic effects in the living organism. Light irradiation Light is a part of electromagnetic radiation, which ranges from ultraviolet to infrared spectrum (wavelength from 10 nm to 1 mm). The light in this range does not transmit or ionize, unlike gamma rays and x-rays. The intensity of a laser light for therapy is much lower than that of a surgical laser; therefore, there is no hot feeling or damage during the procedure. In the near-infrared (NIR) window region (650-1,700 nm), the light has lower photon scattering, light absorption, and autofluorescence than those for ultraviolet and visible light in endogenous tissue [56]. Therefore, fluorescence light in the NIR window is suitable for in vivo applications. Light-sensitive NPs combined with NIR fluorescence are multifunctional due to their unique optical properties and various surface functionalities. In this section, we discuss light irradiation-triggered heat generation, drug release, and ROS generation for tumor diagnosis and therapy. Photoinduced drug release Tumoral temperature is considered to be 1–2 ◦ C higher than that of normal tissues due to the increased metabolic rate [57], which has been used for developing thermosensitive drug delivery systems using such endogenous temperature variations [58]. However, it seems that utilizing tumoral temperature is not a practical strategy because not only the target temperature is too close to 37 ◦ C, but there is also a wide variation of temperature values in different parts of the human body as well as at different periods of the

daily cycle [57]. Therefore, inducing a large temperature gradient between tumor and normal tissues by utilizing an external energy source is more strategic for tumor-specific drug delivery and treatment than employing the endogenous tumoral temperature, while avoiding unwanted drug release in normal tissues. Strategies for effective drug delivery have been intensively investigated by using nanocarriers, such as hollow nanoshells, metal nanocages, and mesoporous silica. Among them, plasmonic noble metal NPs have been widely used as light-sensitive NPs for drug delivery and photothermally induced drug release. The major advantage of NPs that differentiates from a bulk material is a high surface-to-volume ratio, resulting in a further increase of drug payload inside the NP. In addition, appropriately designed carriers can increase drug solubility, stability, and apparent molecular weight to safely prolong the presence of a drug in the circulation [6]. For the light-stimulated drug release, metal NPs are often incorporated with other functional materials such as mesoporous silicas or thermosensitive polymers. For instance, Yang et al. demonstrated NIR light-sensitive gold NPs incorporated within a mesoporous silica framework [59]. The surface of the mesoporous silica-coated gold NPs was functionalized with aptamer DNA. Upon irradiation of NIR light, the photothermal effect of the gold NPs led to increasing the local temperature, resulting in the dehybrization of the linkage DNA duplex that anchored the G-quadruplex DNA cap the pores of the mesoporous silica, allowing the release of the drug inside the silica pores. In another example, photosensitive gold nanocage decorated with thermal sensitive polymer was utilized (Fig. 7a) [60]. Gold nanocages have large hollow interiors and porous walls for drug payloads and release. The pre-loaded drug can be released by photothermal induced polymer shrinkage using a NIR laser. Separately, the combination of photothermal therapy and heatinduced drug release systems can produce excellent synergies for cancer treatment [61,62]. Noh et al. used Au/Ag nanoshells and employed EGFR antibody and Dox on the nanoshells for lung cancer targeting and treatment, respectively [61]. The hollow interiors of the multifunctional nanosystem were made by galvanic replacement reaction, and heat-induced photothermal drug delivery was

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made after effective targeting and triggering by NIR light irradiation (Fig. 7b). Another example is activatable fluorescence imaging and chemo/photothermal dual therapy of Dox-loaded PPy theranostic NPs [62]. This nanosystem was used for treating triple-negative breast cancer, where Dox is the only treatment option since there are no available targets for treatment. Despite reasonable response rates to the initial Dox therapy, long-term treatment with Dox resulted in the development of a drug-resistant cellular phenotype in patients and poor clinical outcome. Park et al. showed that Dox could be released from the NPs repeatedly every time when NIR light was applied. The combination of stimulated drug release and hyperthermia upon NIR light irradiation greatly improved therapeutic efficacy whereas Dox-resistant triple-negative breast cancer does not respond to free Dox even at high concentrations [62].

Photothermal therapy Hyperthermia (a.k.a. thermotherapy) is a well-established cancer treatment modality in clinical practice along with surgery, chemotherapy, radiotherapy and biological therapy [63]. Radiofrequency, ultrasound, and microwave can be used as heat-inducing stimuli for relatively large sized tumor masses. However, the main obstacles of these stimuli are 1) poor tumor specificity to energy absorption resulting in unwanted damage to the surrounding normal tissues and 2) some parts of the tumors may be intact even after hyperthermia due to insufficient heat induction in the site, making it difficult to achieve homogenous heat distribution in tumors. Recently photothermal therapy using plasmonic metal NPs have been widely explored because the region of heat generation could be confined to tumor tissues with a high spatial resolution [64–68]. Plasmonic metal NPs have strong electric fields at the surface, so the absorption and scattering of electromagnetic radiation have the opportunity to considerably improve the specificity of cancer ablation by acting as antennas for accepting external light [64,65,69]. Many researchers have been pioneering on the development of the plasmonic nanocrystals, such as nanoshells, nanocages, and nanorods for photothermal therapy [66,67]. For example, anti-epidermal growth factor receptor (antiEGFR) monoclonal antibodies were conjugated to gold NPs for cancer cell diagnosis and selective photothermal therapy [64]. The anti-EGFR antibody-conjugated nanorods bind specifically to the surface of the malignant type cells overexpressed EGFR on the cytoplasmic membrane. Using darkfield microscopy, the malignant cells were clearly visualized and diagnosed from the nonmalignant cells. Upon the exposure of laser at 800 nm, malignant cells were photothermally destroyed [64]. In addition, Bhatia et al. reported PEG-protected gold NPs that exhibit superior spectral bandwidth, photothermal heat generation, and extended blood circulation [65,68]. A single intravenous injection of PEGylated gold NPs enabled the destruction of all irradiated human xenograft tumors in mice. The NPs were also labeled with Raman reporters, providing an integrated platform for multiplexed Raman detection and remote-controlled photothermal heating. For efficient photothermal therapy, it is prerequisite to obtain a uniform distribution of high temperature at a low concentration of gold NPs (i.e., ≤ 0.5 nM) [70]. This could result in effective and homogeneous temperature for hyperthermia treatment of tumors that could be achieved with a low concentration of plasmonic nanomaterials. Both operating temperature and maintenance time of the temperature in tumors must be considered to achieve effective thermal ablation of cancer cells. That is, the instantaneous induction of protein coagulation and subsequent cell death could occur when the temperature is higher than 60 ◦ C [71]. However, it takes 60 min to induce irreversible cellular damage at 46 ◦ C [72].

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Photodynamic therapy (singlet oxygen generation) Photodynamic therapy involves an administration of tumor localizing photosensitizers, which transfer the energy acquired by absorbing photons to the surrounding oxygen molecules [73]. Similar to the photo-induced drug release, researchers deliver hydrophobic photosensitizers using NPs and activate them by external light irradiation. For example, photosensitizer-loaded plasmonic gold NPs assembly was reported for effective cancer imaging and treatment [74]. Ce6 photosensitizers were encapsulated into the assembly of gold NPs and NIR laser irradiation released the encapsulated molecules into the tissue because of its strong absorbance in the NIR range. In addition, the assembly of Ce6-loaded gold NPs has multifunctionality, including trimodal NIR fluorescence/thermal/photoacoustic imaging and synergistic effect on photothermal/photodynamic therapy [74]. Likewise, hyperthermia increases tumor temperature and can therefore directly kill cancer cells. It may also sensitize the cells to other treatment modalities such as chemotherapy and photodynamic therapy. Tian et al. developed a photosensitizer molecule (Ce6) loaded on PEGfunctionalized GO complex [75]. The Ce6-GO complex can generate cytotoxic singlet oxygen in addition to heat. This local heating promoted the delivery of Ce6 molecules to the target cancer cells under the light excitation and improved photodynamic cancer cell killing efficacy. Magnetic fields Magnetic fields have excellent tissue penetration (15 cm at 400 kHz) [76] and are widely used for medical imaging in the clinic. Magnetic NPs are a major class of multifunctional nanoscale materials with the potential to be applied in hyperthermia treatment, drug and gene delivery, and imaging of cancer cells [77,78]. These NPs can integrate diagnostics and therapy into one system to achieve a theranostic approach. Since a magnetic core with appropriate coating shells is basic components for magnetic NPs, thermal and/or mechanical effects can be generated on the magnetic NPs depending on the strength and status of magnetic fields [79]. In addition, therapeutic drugs are often loaded in the shell and targeting moieties and/or optical dyes can be added on the surface as well. These NPs will spin into a large overall magnetic moment in one direction, responding to an external magnetic field [80]. When alternating magnetic fields are applied, a magnetization reversal process occurs in the magnetic particles resulting in heat is released [81]. The thermal effects increase with the frequency and amplitude of the fields [82]. The heat then acts as the source for hyperthermia therapy on one side and/or realize thermal-sensitive drug release on the other [63,83,84]. For example, as shown in Fig. 8a, exchange-coupled magnetic NPs induce an excellent antitumor therapeutic effect on a tumor-bearing mouse under an external magnetic field. Fig. 8b represents that tumors on the mice treated with magnetic NPs were eliminated over time, while the control tumors (untreated, hyperthermia only with Feridex, and Dox only) increased during the same period of treatment time [85]. Efficient treatment of breast cancer xenografts could be strongly enhanced by multi-functionalized iron NPs (MF66+N6L + Dox) and magnetic hyperthermia [86]. In addition, magnetic iron oxide NPs conjugated with a peptide that contains the amino-terminal fragment urokinase plasminogen activator (uPA) can target tumors overexpressing uPA receptors, which enabled enhanced tumoral imaging and drug delivery in a cancer animal model [87]. Thermotherapy under the magnetic field also showed promising results in some clinical studies. Patients with recurrent glioblastoma received thermotherapy using magnetic NPs in conjunction with a reduced radiation had longer overall survival compared to conventional therapies [88]. In addition, the control of cell death signaling can be made by a magnetic switch. As shown in Fig. 8c, when a magnetic field was applied to the

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Fig. 8. Nanosystems triggered by magnetic field changes for apoptosis. (a) Schematic of magnetic field induction to the tumor-bearing mouse injected with magnetic NPs. (b) Nude mice xenografted with cancer cells before treatment (top row) and 18 days post-treatment (bottom row) with hyperthermia using magnetic NPs, Feridex, and doxorubicin, respectively. Reproduced with permission from [85]. (c) Focused magnetic field triggers magnetic NPs bound to the death receptor 4 and turns on apoptosis signaling in a zebrafish model. Magnetic field switching results in apoptosis and morphological changes of cells. Reproduced with permission from [89].

zinc-doped iron oxide magnetic NPs with a targeting antibody for death receptor 4, apoptosis signaling was successfully turned on in cancer cells and animals [89]. This magnetic switch was operable precisely at the micrometer scale. Very recently, Kandasamy et al. successfully developed hydrophilic superparamagnetic iron oxide NPs [90]. These NPs were effectively internalized by the HepG2 liver cancer cells during cytotoxicity and in vitro magnetic fluid hyperthermia experiments. This is a promising heat-inducing agent for hyperthermia operable under a magnetic field.

Ultrasounds Ultrasound is routinely used for diagnostic imaging in the clinic and is also actively investigated as an external stimulus for drug delivery and therapeutic applications [91–95]. Tissue penetration depth can be adjustable by tuning frequency, cycles, and duration of exposure. Ultrasound allows for spatiotemporal control at the desired site with millimeter precision, thus preventing side effects in healthy tissues [91–93]. As an ultrasound wave propagates through tissue in the body, several physical effects such as simple pressure variation, acoustic fluid streaming, cavitation, and local hyperthermia occur, which can be used as a source for ultrasoundmediated drug release (Fig. 9) [94]. Cavitation, hyperthermia, and acoustic fluid streaming are three major mechanisms by which ultrasound realizes drug release from nanocarriers [95]. Cavitation involves either rapid growth and collapse of bubbles (inertial cavitation) or sustained oscillatory motion of bubbles (stable cavitation). Both types of cavitation induce strong physical, chemical, and biological effects in tissues and it is a phenomenon involving the nucleation, growth, and oscillation of gaseous cavities [96]. NPs displacements and fluid currents, namely “acoustic streaming”

caused by the force of the ultrasound field, may assist in the mixing and delivery of drugs [97]. The multiplicity of its interactions with cells and tissues is the key to the versatility of ultrasound in medical therapies. Ultrasound can render cell membranes more permeable to facilitate the delivery of drugs across some barriers like the skin and BBB, and the most common ultrasound-sensitive nanocarriers are liposomes and lipid nanobubbles encapsulating air or perfluorinated hydrocarbon [98]. Compared with other approaches, ultrasound is especially beneficial for the gene or protein delivery, attributed to its ability of pore formation in the cell membrane, increasing the permeability of the cell membrane and allowing direct access to the cytosol bypass the degradative endocytotic pathway [99].

Electric fields Electric field has been successfully utilized to trigger the release of molecules via conducting polymeric materials or implantable electronic delivery devices [100–103]. Novel electro-responsive delivery systems with the higher responsiveness and drug loading are achieved by implementing concepts of nano-engineering into the structure [104]. Redox reaction or ionization of the host or guest molecule can be induced under the certain magnitude of voltage or current. The gain or loss of electrons can change the hostguest interaction system, thus leading to bond cleavage or particle deformation thus drug release [105–107]. Polyelectrolytes containing large numbers of ionizable groups confer responsiveness to an electrical stimulus through deformation of the polymers [108,109]. For instance, Ying et al. developed an electro-responsive hydrogel NP and polyelectrolyte poly(sodium 4-vinylbenzene sulfonate) is used as the responsive material to facilitate the release of anti-

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Fig. 9. Nanosystems trigged by ultrasound for targeted drug delivery. (a) Liposomes composed of phospholipid bilayer membrane, aqueous core, and anticancer drugs. Hydrophobic drugs can be loaded on the lipophilic bilayer and released by ultrasound (b–d). (b) Ultrasound-induced cavitation generates violent collapse of gas bodies near the liposome. Microbubble nucleation and growth may also occur within the lipid bilayer because lower energies are required to disrupt intermolecular van der Waals forces between lipid chains. (c) Focused ultrasound increases the local fluid temperature to promote drug release from temperature-sensitive liposomes. (d) Acoustic radiation-force streaming increases the frequency of particle collisions and thus facilitates material transfer. Reproduced from [94].

epilepsy drug under an electric field [110]. Both in vivo and in vitro results of this hydrogel showed the potential application in epilepsy treatment. In the field of voltage-responsive polymer systems, ␤-CD, calixarene, and cyclophane can be used host molecules and ferrocene (Fc), cobaltocenium and viologen can be used common guest molecules [111]. Association and dissociation effect, inducing a desirable assembly and disassembly of vesicles would happen during the voltage change as shown in Fig. 10. Poly(styrene) with ␤-CD and PEG with Fc can self-assemble into deblock copolymer and form supramolecular vesicles in the aqueous solution. This supramolecular vesicular system showed voltage-responsive reversible assembly and disassembly by the external electric stimuli [102]. Rhodamine B was used as a model drug in this system to test the release efficacy in vitro where the release speed could be

controlled by the applied voltage strengths (32–450 min). In the following study, the same group successfully demonstrated a release of paclitaxel under a +1.0 V stimulus and no release without stimulus with using PEG-modified with ␤-CD and the poly(L-lactide) with Fc [111]. However, intrinsic properties of electric fields such as cell damage and poor tissue penetration limit the application in deep tissues. Ionizing radiation Ionizing radiation (gamma ray and X-ray) as an external trigger has benefits such as high tissue penetration and clinical utility [112]. Metallic NPs, such as gold NPs, have been reported to enhance the radiation therapy on cancer cells (radiosensitization) and decrease side effects on surrounding normal tissue. For exam-

Fig. 10. Nanosystems trigged by voltage changes to control assembly and disassembly of PS-␤-CD/PEO-Fc supramolecular vesicles. Reprinted with permission from [102].

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ple, functional gold NPs can enhance both apoptosis and necrosis significantly in cancer cells when treated with ionizing radiation compared with cells treated only with ionizing radiation [113]. Separately, Heinfeld et al. evaluated the synergistic effect of gold NPs and X-rays in the survival rate of tumor-bearing animals compared with the control group that received irradiation only [114]. The combination of gold NPs and X-rays has also been used in the controlled release of chemotherapeutics. Dox was conjugated to the complex of DNA strands and gold NPs, where X-rays irradiation induced the release of Dox by breaking DNA strands [115]. These gold NPs could be used for CT imaging in addition to providing a theranostic platform in cancer management. Ionizing radiation can also induce reactive oxygen species (singlet oxygens) for cancer therapy similar to photodynamic therapy. For example, CeF3 NPs and SiC/SiOx nanowires were used to enhance radiation therapy in cancer treatments by generating singlet oxygens upon ionizing radiation [116,117]. Additionally, X-ray irradiation can induce hypericin (photosensitizer) in the lanthanide (Gd3+ ) micelles, generating singlet oxygen species [118]. Gd3+ can be also utilized for MRI imaging, thus, lanthanide micelles are used an efficient theranostic nanosystem.

Perspectives Because nanomaterials have been revolutionary in cancer imaging, diagnosis, and treatment over the past decades, stimuliresponsive nanosystems will continuously contribute to the field of drug delivery and targeted cancer therapy. For enhanced targeted cancer therapy, however, a nanomaterial administered into the bloodstream should travel to the target site efficiently without getting trapped in the physiological barriers. This is the first and foremost requirement for site-specific drug delivery and targeted cancer therapy. To achieve desirable theranostic targeting beyond biological barriers, therefore, a fundamental understanding in the intrinsic physicochemical properties such as HD, composition, charge, charge-to-mass ratio, and hydrophilicity/lipophilicity of NPs as well as their physiological dynamic behavior is essential. Based on this prerequisite knowledge, multifunctional nanosystems can be designed to reach specific target organs and be sequentially activated by internal or external stimuli. Although there are many reports that tumor-specific delivery and targeting can be achieved by using small microenvironmental changes such as pH gradient, cancer-related enzyme, redox potential difference, and ROS, it seems like such internal stimuli are not sufficient enough to activate the nanosystem to be functional. Rather than using small changes occurring at the tumor microenvironment, nanosystems triggered by large environmental changes such as from physiological pH (7.4) to the lysosomal pH (≤ 5.0) would be more effective for tumor targeting. Additionally, triggering by external energies including light irradiation, magnetic fields, ultrasounds, electric fields, and ionizing radiation can directly activate the nanosystem at the specific tumor site. Although elevated temperatures using high frequency or ultrasonic waves have the advantage of inducing hyperthermia over a larger area, it is challenging to completely avoid normal tissue damage owing to the difficulty in separation of tumors from normal tissue. It is possible to precisely control the region where temperature increases using plasmon nanomaterials for photothermal therapy. However, the tissue penetration depth of the therapeutic light is limited for photoinduced nanosystems. Therefore, new nanosystems or technologies that can induce hyperthermia over an extended body surface and are capable of precise temperature control should be developed in the future. Although a large number of stimuli-responsive nanosystems using different mechanisms and structures have been reported,

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Shuang Hu, M.D. received her medical degree in Radiology and Nuclear Medicine from West China School of Medicine, West China Hospital, Sichuan University in China. She just finished her 2-year research study at the MGH and HMS through the Global Joint Program of Sichuan University. Her research focuses on the advanced clinical imaging such as the application of PET/MRI in radiotherapy and clinical translation of novel targeted molecular imaging for cancer theranostics.

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Mi Hyeon Cho, Ph.D. received her B.S. degree in chemistry from Sookmyung Women’s University in 2007. She received her M.S. and Ph.D. degrees in chemistry from Yonsei University in 2011 and 2017, respectively. She is currently a postdoctoral research fellow at the National Cancer Center (NCC), Republic of Korea. Her current research interest is on the design of nanomaterials for the control of cellular activities.

Yongdoo Choi, Ph.D. is a Principal Research Scientist and chief of Nanochemistry Laboratory at NCC. His laboratory focuses on the development of novel molecular imaging agents, activatable photodynamic therapy agents, and theranostic nanomedicines for enhanced photothermal and radiotherapy.

Suk Ho Hong, M.S. is a former researcher at NCC. His main research topics at NCC included designing targeted drug delivery systems and theranostic probes for imaging and multimodal therapy. He is currently pursuing his Ph.D. in Chemistry at Columbia University in New York, NY.

Hak Soo Choi, Ph.D. is an Associate Professor of Radiology at HMS, Associate Chemist at MGH, and faculty of Dana Farber/Harvard Cancer Center. Dr. Choi is currently Director of Bioengineering & Nanomedicine Program of MGH. His laboratory focuses on the development of novel targeted contrast agents for image-guided surgical interventions to solve important problems in oncology and biomedical research, with an emphasis on molecular imaging and tissue-specific targeting.

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