Rational Design of Cancer Nanomedicine for Simultaneous Stealth

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Rational Design of Cancer Nanomedicine for Simultaneous Stealth Surface and Enhanced Cellular Uptake Qiao Jin, Yongyan Deng, Xiaohui Chen, and Jian Ji ACS Nano, Just Accepted Manuscript • DOI: 10.1021/acsnano.8b07746 • Publication Date (Web): 25 Jan 2019 Downloaded from http://pubs.acs.org on January 25, 2019

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Rational Design of Cancer Nanomedicine for Simultaneous Stealth Surface and Enhanced Cellular Uptake Qiao Jin, Yongyan Deng, Xiaohui Chen, Jian Ji* MOE Key Laboratory of Macromolecule Synthesis and Functionalization of Ministry of Education, Department of Polymer Science and Engineering, Zhejiang University, Hangzhou, 310027, Zhejiang Province, PR China

ABSTRACT

Owing to the complex and still not fully understood physiological environment, the development of traditional nano-sized drug delivery systems is very challenging for precision cancer therapy. It is very difficult to control the in vivo distribution of nanoparticles after intravenous injection. The ideal drug nanocarriers should not only have stealth surface for prolonged circulation time, but also possess enhanced cellular internalization in tumor sites. Unfortunately, the stealth surface and enhanced cellular uptake seem contradictory to each other. How to integrate the two opposite aspects into one system is a very herculean but meaningful task. As an alternative drug delivery strategy, chameleon-like drug delivery systems were developed to achieve long circulation time while maintaining enhanced cancer cell uptake. Such

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drug dnanocarriers can “turn off” their internalization ability during circulation. However, the enhanced cellular uptake can be readily activated upon arriving in tumor tissues. In this way, stealth surface and enhanced uptake are of dialectical unity in drug delivery. In this review, we focused on the surface engineering of drug nanocarriers to obtain simultaneous stealth surface in circulation and enhanced uptake in tumor. The current strategies and ongoing developments, including programmed tumor targeting strategies and some specific zwitterionic surfaces will be discussed in detail. Key words: nanomedicine, nanoparticles, cancer therapy, drug delivery, stealth, stimuliresponsive, enhanced uptake, tumor microenvironment Cancer remains one of the most mortal diseases in the world with more than eight million deaths every year. Chemotherapy, as one of the major categories of medical oncology, is widely adopted for treating various cancers. However, owing to the systemic distribution and lack of tumor targeting ability, the chemotherapeutic agents can also kill healthy cells, leading to severe side effects, including alopecia, myelosuppression, nephrotoxicity, mucositis, and so on. In order to achieve improved therapeutic efficacy as well as minimal adverse effects, Cancer nanomedicine has received much attention in chemotherapy.1-5 Cancer nanomedicine is the application of nanotechnology for cancer treatment, which has received enormous progress in recent years. Taking advantage of the enhanced permeability and retention (EPR) effect,6 nanoparticles are able to accumulate in tumor tissue in a passive mode.7,8 Moreover, Leong et al discovered that nanoparticles can induce micrometer sized gaps between endothelial cells, which is called “nanoparticle induced endothelial leakiness” (NanoEL).9-13 It is very important that NanoEL allows the nanoparticles to control access to the tumor even in the absence of any EPR effect. Compared to conventional chemotherapeutics, cancer nanomedicine shows improved


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drug solubility, prolonged blood circulation, reduced side effects, and so on. In the last decade, huge numbers of drug nanocarriers were developed.14-17 More than 250 nanoparticle-based drug delivery systems are in preclinical and clinical process.18-20 Nanomedicine, which provides the possibility of utilizing nanoparticles to deliver drugs to specific cells has attracted more and more attention.12,21-23 However, how to make full use of “nano” in cancer medicine is not satisfactory and there is still a long way to go. When the nanoparticles are applied for drug delivery, a lot of issues have to be addressed because of the complex and still not fully understood physiological environment in cells, tissue and body. After the nanoparticles are intravenously injected into the body, the drug delivery process can be influenced by a lot of factors. There are a lot of obstacles that should be overcome. The full understanding of nano-bio interactions might be a rationally derived key to achieve the optimal bioavailability of the nanoparticles.24 We should be fully aware of the complexity of the in vivo physiological environment. A myriad of biomolecules such as proteins, lipids, and sugars exist in the biological environments. The interaction between nanoparticles and biomolecules in the biological environments should not be ignored. Proteins may be adsorbed onto the nanoparticles and “protein corona” is formed.25-27 In this regard, although nanoparticles are usually designed with specific surfaces, such surfaces can only exist transiently after exposed to physiological environments due to the formation of protein corona. The protein corona can significantly influence the in vivo fate of nanoparticles. For example, cationic nanoparticles exhibit enhanced cellular uptake due to their affinity to negatively charged cell membrane. Unfortunately, protein corona can be easily formed on cationic nanoparticles and protein corona covered nanoparticles are overall negatively charged, which can hinder the nanoparticle-cell interaction. What’s more, active targeting ligands may also be buried by protein corona. In this situation, the targeting


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ligands will be invalid, leading to the off-target effects.28 The stealth surface engineering is helpful to minimize the formation of protein corona. The remote control of the nanoparticles’ physicochemical properties is especially important to maximize the therapeutic efficacy of drug nanocarriers and achieve the greatest potential especially in disease diagnosis and therapy.29,30 The size, shape, and surface charge of the nanoparticles can greatly influence their in vivo fate by controlling the interactions of nanoparticles with extracellular spaces, cell surface and intracellular components.31,32 For example, the size of the nanoparticles can not only have an important effect on blood circulation and tumor penetration, but also have a profound influence on cellular internalization and metabolism.33,34 The internalization of nanoparticles is both size dependent and shape dependent.35 Smaller nanoparticles can adhere to cells faster and stronger, typically leading to a higher internalization rate. The shape of the nanoparticles is also known to play an important role in cellular uptake. Cylindrical or rod particles can be internalized better than spherical nanoparticles of similar volume.36-38 However, we should keep in mind that the design of nanoparticles for in vivo biomedical applications is not that easy. The delivery of drug nanocarriers is a multistage procedure in vivo. There are many physiological barriers and dilemmas that the drug nanocarriers must overcome for the systemically administered drug delivery systems.39-41 The drugs should not be leaked from the nanocarriers during circulation. However, drugs should be rapidly released from the nanocarriers after internalized by cancer cells. 2) The nanocarriers should have relative large (~ 100 nm) size for long circulation time. However, the nanocarriers also need small size (< 50 nm) for efficient tumor penetration. 3) The nanocarriers should have stealth surface to avoid the clearance by the reticuloendothelial system (RES). However, the nanocarriers should also be efficiently internalized by cancer cells. In this


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review, we will only focus on the up to date approaches to overcome the dilemma of stealth surface in circulation and enhanced uptake in tumor. If the readers are interested in other expects of the dilemmas that the nanocarriers are facing during the delivery procedure in vivo, many excellent and idiographic reviews can be found in literatures.42-53 As discussed above, the nanocarriers should have stealth surface to achieve prolonged blood circulation time, which is the prerequisite of efficient accumulation in tumor tissue by passive targeting or active targeting.54,55 Effective cellular uptake by cancer cells is another key factor to achieve intracellular drug delivery.56-58 Therefore, the surface tailoring of nanocarriers plays a crucial role to improve the therapeutic performance.59 Generally, stealth surface is beneficial for long circulation time and high tumor accumulation.60 However, like a double-edged sword, stealth surface may hinder the uptake of the nanocarriers by cancer cells.61-63 For example, positively charged nanocarriers are believed to be helpful for cellular uptake due to the high affinity to cell membranes. Nevertheless, positively charged nanocarriers may be cleared by RES rapidly, resulting in low accumulation in tumor tissue.64 Similarly, “stealthy” PEGylated drug nanocarriers are well known to achieve long blood circulation time and efficient accumulation at tumor sites by EPR effect. However, PEGylated drug nanocarriers are always suffering from reduced uptake by cancer cells, which is well known as “PEG dilemma”.65,66,67 In order to enhance the uptake of nanocarriers, one of the most effective strategies is modifying the nanocarriers’ surface with targeting moieties.68 Unfortunately, the addition of targeting ligands, e.g. RGD peptide, can greatly influence the circulation time and passive targeting ability.69 Meanwhile, some exposed targeting ligands might be degraded during circulation.70 Moreover, the active targeting ligands may also produce off-target effects since some receptors are expressed on normal cells as well.71 In particular, the utilization of targeting strategies employing


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nonspecific ligands (cell penetrating peptides, RGD, etc.) should be very careful due to the nonspecific interaction with normal cells.72 Accordingly, as a part of “Dr. Ehrlich’s magic bullet” concept, how to achieve long circulation time while maintaining enhanced cancer cell uptake is still a big challenge. Taking advantage of specific tumor microenvironments, such as acidosis, hypoxia, and dysregulated enzymes, endogenous stimuli-sensitive drug nanocarriers were designed to overcome this dilemma for intended optimal cancer therapy.73,74 The specific tumor microenvironments therefore form the basis of smart tumor targeting drug delivery systems.75-77 Various smart drug delivery systems that can switch their surface properties from stealth surface to enhanced cellular uptake in respond to specific tumor microenvironments were fabricated. Many review articles on stimuli-responsive drug delivery systems were reported.78-80 Unfortunately, most of them talked about stimuli-responsive drug release or the combination of stimuli-responsive drug release and stimuli-responsive cellular uptake. It is urgently needed to have a comprehensive description of smart drug delivery systems that can overcome the dilemma of stealth surface in circulation and enhanced uptake in tumor. In this review, we will focus on the current strategies and ongoing developments of the surface engineering of drug nanocarriers to achieve simultaneous stealth surface and enhanced cellular uptake (Figure 1).


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Figure 1. Schematic illustration of the design of drug nanocarriers that can achieve simultaneous stealth surface in circulation and enhanced cellular uptake in tumor. As is known, solid tumor tissues often have specific physiological characteristics.81 For example, because of the Warburg effect, the tumor extracellular pH is mildly acidic (~pH 6.56.8).82,83 Matrix metalloproteinases (MMP) are overexpressed in an extracellular tumor microenvironment and play a critical role in tumor progression and metastasis.84 Taking advantage of the tumor microenvironment, a programmed targeting strategy was recently developed by fabrication of drug nanocarriers with stealth surface that can “turn off” the internalization ability during circulation.85,86 After subsequently accumulated in tumor tissue, the dormant targeting ligands can be activated or the surface charge can be changed to positive in response to the tumor microenvironment, which can enable enhanced internalization by cancer cells. Besides, light is another very attractive stimulus for smart surface engineering of nanoparticles due to the excellent spatiotemporal resolution and controllability. Different stimuli-


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sensitive nanocarriers were therefore developed for simultaneous stealth surface in circulation and enhanced uptake in tumor (Table 1). One optional strategy is to shield or cage the targeting ligands during circulation to suppress the nonspecific interactions. Table 1. Design of different stimuli-sensitive nanocarriers for simultaneous stealth surface in circulation and enhanced uptake in tumor. Sensitive linkage/group

Activation approach

Nanocarrier

Example

methacryloyl sulfadimethoxine

pH 6.6

micelle

96

2,3-dimethylmaleamidic amide

pH 6.5-6.8

micelle, polymeric nanoparticle, AuNP, carbon dot

97-99,173-175

2-propionic-3methylmaleic amide

pH 6.5-6.8

micelle

101,102

benzoic imine

pH 6.5-6.8

MSN, micelle

110,112

GPLGIAGQ

MMP

liposome, micelle

118,123

PLGLAG

MMP

polymeric nanoparticle, micelle, dendrimeric nanoparticle, AuNP

119,122,134,140

PLGVR

MMP

CdSe/ZnS quantum dots, MSN

120,139

nitrobenzyl

UV, two-photon NIR

liposome, polymeric nanoparticle, polymerdrug conjugates

128,156,161

coumarin

UV

micelle

158

thioketal

light-triggered ROS

micelle

129

hyaluronic acid

hyaluronidase

nanoparticle

149

pNIPAAm-co-pAAm

NIR-triggered photothermal effect

silica core-gold shell nanoparticles

164


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azide

H2S

micelle

190

histidine

pH 6.8

micelle

208-211

2-(hexamethyleneimino) ethyl methacrylate

pH 6.8

micelle, polymeric nanoparticle

216-218

pHLIPs

pH 6.5

AuNP, mesoporous silica coated gold nanorods, mesoporous organosilica nanoparticles, dendrimerdrug conjugate, liposome

226,228-231

carboxybetaine

pH 6.5

micelle

250,251

alkoxyphenyl acylsulfonamide

pH 6.5

AuNP

252

Stealth corona-detachable drug delivery systems A classic example based on this concept is poly(ethylene glycol) (PEG) detachable drug delivery system since PEG is considered as the golden standard for stealth polymers. As discussed above, PEG can hinder the interaction between drug nanocarriers and cell membrane, which will result in poor cellular internalization and unsatisfactory tumor inhibition.87-91 In order to defuse the conflicts of favorable prolonged circulation time and unexpected poor cellular uptake owing to the “PEG dilemma”, drug nanocarriers with detachable PEG corona have been extensively studied.92-95 As a proof-of-concept research, Bae et al designed cell penetrating TAT conjugated polymeric micelles.96 A pH-sensitive anionic block copolymer poly(methacryloyl sulfadimethoxine)-b-PEG (PSD-b-PEG) was then complexed with cationic TAT micelles via electrostatic interaction to obtain final drug nanocarriers. The complexed drug nanocarriers can be preferentially accumulated in tumor tissues by EPR effect owing to the stealth PEG corona. Afterwards, the PSD block was switched from anionic to neutral upon exposed to extracellular tumor pH (~6.6), which could result in the detachment of PEG corona. As a result, the cell


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penetrating TAT would be exposed. The drug nanocarriers can be transported not only into the cancer cells but also near the nucleus.

Figure 2. (a) Schematic illustration of the preparation of PEG-detachable nanocomplex. (b) Schematic illustration of pH-sensitive charge-reversal and pH-sensitive drug release. (c) Schematic illustration of the dual pH-sensitive nanocomplex for targeted drug delivery, enhanced cellular uptake, and intracellular drug release. Reproduced with permission.100 Copyright 2017 Ivyspring International Publisher 2,3-Dimethylmaleamidic amide (DMMA) linkage, which can be cleaved at pH 6.8, is widely used to construct extracellular tumor pH sensitive drug nanocarriers. DMMA linkage is therefore an ideal candidate to design stealth corona-detachable drug delivery systems. Such drug nanocarriers were developed by the Wang group. In one of their research, they successfully


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fabricated PEG detachable nanoparticles for siRNA delivery.97 The nanoparticles were prepared by introducing DMMA-containing negatively charged copolymer mPEG-b-PAEP-Cya-DMMA to positively charged ssPEI/siRNA complexes via electrostatic interaction. The nanoparticles could be stabilized by stealth PEG corona during blood circulation to reduce nonspecific interactions with serum components. Owing to extracellular tumor pH triggered detachment of PEG corona, the nanoparticles exhibited enhanced cancer cell uptake, which greatly improved the gene silencing efficiency and showed great potential in cancer therapy. Similarly, Wu et al fabricated PEG detachable gene delivery system by conjugation of DMMA functionalized PEG to PEI derivatives/DNA complexes.98 The gene transfection efficacy increased 3 times owing to the enhanced cellular uptake. Similarly, DMMA modified anionic PEG-based block copolymer was coated on the positively charged nanoparticles which were formed by the assembly of βcyclodextrin-PEI600 and miR-34a.99 The obtained ternary nanoparticles with PEG corona exhibited reduced nonspecific adsorption during blood circulation. Due to the charge reversion of PEG-based block copolymer from negative to positive, PEG corona was removed owing to the electrostatic repulsion, which markedly enhanced nanoparticle uptake and promoted intracellular microRNA-34a release, leading to the down-regulation of CD44 expression and inhibition of tumor growth. Dual-pH sensitive PEG-detachable nanocomplex was reported recently for tumor-targeted drug delivery with enhanced anticancer activity (Figure 2).100 Wu et al prepared DOX and TAT conjugated poly(β-L-malic acid) (PMLA), which was assembled with PEI to obtain positively charged nanoparticles (PMLA-PEI-DOX-TAT). The dual pH-sensitive charge-reversal nanocomplex was fabricated by coating of PMLA-PEI-DOX-TAT nanoparticles with DMMA ended comb-like PEG. The as-prepared nanocomplex was covered with PEG corona, which could protect the nanocomplex from the clearance by the RES, leading to


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prolonged circulation time. After arriving at the tumor tissue, PEG corona could be removed upon the hydrolysis of DMMA. In this situation, TAT on the nanoparticles was exposed, thus enhancing the cellular uptake. After cell uptake, DOX could be readily released in the lyso/endosomal compartments of the cancer cells.

Figure 3. PEG detachable polymer-based nanoparticles for siRNA delivery and their change in response to tumor acidity.84 Reproduced with permission. Copyright 2015, American Chemical Society. In addition, since DMMA can only be used to modify polymeric side chains, by replacing the methyl group of DMMA with functional carboxyl group, Wang et al fabricated PEG cleavable siRNA delivery system by introducing 2-propionic-3-methylmaleic anhydride (CDM) in the main chains as a degradable bridged covalent bond (Figure 3).101 The cell-penetrating nanoarginine (R9) was shielded by stealth PEG corona during circulation. After accumulated in tumor tissues by EPR effect, R9 would be exposed for further enhanced cancer cell uptake upon PEG was detached in response to extracellular tumor pH. Thus, the nanoparticles achieved both prolonged circulation time and enhanced cancer cell uptake, which resulted in more effective


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down-regulation of relative mRNA expression. This delivery strategy was extended for effective and safe delivery of chemotherapeutic drugs by the same group.102 They constructed extracellular tumor pH sensitive drug nanocarriers by bridging of U.S. FDA approved PEG and poly(D,L-lactide) (PDLLA) using CDM (PEG-Dlinkm-PDLLA). Better therapeutic efficacy was achieved compared to clinically used PEG-b-PDLLA micelles. The pH-sensitive benzoic imine group, which is formed by the reaction of benzaldehyde and amino groups, is another ideal candidate to fabricate extracellular tumor pH sensitive drug nanocarriers.103-105 Yang et al designed a series of PEG detachable drug nanovarriers via pHsensitive benzoic imine conjugation.106-109 In a typical example, Zhang et al designed tumor acidity-triggered targeted mesoporous silica nanoparticles (MSN) as drug nanocarriers.110 The targeting ligand RGD was anchored on the surface of MSN. Meanwhile, PEG was conjugated to MSN via benzoic imine groups, which was used to shield the RGD peptide to maintain the stealth property during blood circulation. Upon arriving at the tumor tissue, owing to the cleavage of benzoic imine groups, PEG corona can be detached to expose the targeting ligand RGD for targeted cancer cell internalization. Zhang et al also prepared RGD conjugated noncovalently connected micelles.111 PEG was conjugated to the micelles by benzoic-imine bonds. The targeting ligand RGD could be exposed for enhanced cellular uptake with the removal of PEG in acidic tumor microenvironment. Similarly, by conjugation of PEG and lipid with a pH-sensitive benzoic imine bond, Liang et al fabricated detachable PEG-lipid micellar system to defuse the conflicts between excellent stealth surface and poor cancer cell uptake (Figure 4).112 In order to achieve efficient drug release simultaneously, Tan et al constructed multifunctional polyurethane micelles containing GSH-responsive disulfide bonds, targeting ligand folate, and benzoic-imine linked PEG.113 The obtained micelles exhibited excellent


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cytocompatibility, enhanced cellular uptake and efficient intracellular drug release due to the cleavage of pH-sensitive benzoic-imine bonds and GSH-responsive disulfide bonds.

Figure 4. Schematic illustration of benzoic imine-based PEG detachable drug nanocarriers. Synthesis and self-assembly of pH-sensitive DSPE-Hy-PEG copolymer; Schematic illustration of tumor acidity-sensitive removal of PEG for enhanced cellular uptake.112 Reproduced with permission. Copyright 2016, The Royal Society of Chemistry. In addition to tumor acidity, up-regulated enzymes are also widely used for the design of PEG detachable drug nanocarriers.114-116 In particular, MMP is overexpressed in many different cancers and has attracted much attention for programmed cancer targeting. MMP triggered detachment of stealth PEG was first developed by the Torchilin group.117 In their research, PEG was conjugated to a TAT functionalized liposomal nanocarrier by MMP-2 substrate peptide


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GPLGIAGQ. The cationic TAT can hence be shielded by PEG corona. After pretreated with MMP-2, cell-penetrating peptide TAT would be exposed, which resulted in 2-fold increase of cellular uptake. Using the same concept, they explored in vivo performance of MMP2-sensitive PEG-drug conjugates.118 MMP-2 cleavable peptide GPLGIAGQ was used as a linker to synthesize PEG-paclitaxel (PTX) conjugate (PEG2000-peptide-PTX). The prodrug nanocarries were

prepared

by

co-assembly

of

PEG2000-peptide-PTX

and

TATp-PEG1000-

phosphoethanolamine (PE). As expected, MMP-2 triggered cleavage of long chain PEG facilitated TAT-mediated cancer cell uptake. Compared with non-sensitive controls, the MMP-2 sensitive drug nanocarriers achieved much better anticancer efficacy, which might be ascribed to the collaborative functions of stealth PEG corona and MMP-2 triggered TAT targeting. Gu et al constructed tumor-specific multiple stimuli-activated dendrimeric nanoassemblies with a metabolic blockade to overcome chemotherapy resistance.119 The stealthy dendritic PEGylated corona could be removed under the stimulation of MMP-2 in tumor microenvironment, which was very important for deep tumor penetration, enhanced cellular uptake, cytoplasmic redoxsensitive disintegration, and subsequent lysosome acid-triggered nucleus delivery of anticancer drugs. Very recently, a dual-enzyme-sensitive gemcitabine (GEM) delivery system was developed for programmed pancreatic cancer therapy.120 The dual-enzyme-sensitive GEM nanocarriers were prepared by conjugation of MMP-9 detachable PEG, cathepsin B cleavable GEM, and CycloRGD to CdSe/ZnS quantum dots. The GEM nanocarriers with PEG outer shell could avoid nonspecific interactions and exhibit prolonged blood circulation time. After accumulation in tumor tissue by EPR effect, the PEG outer shell can be removed by overexpressed MMP-9 in tumor tissue and RGD would be exposed, which was capable of facilitating cellular internalization. Once internalized into the pancreatic cancer cells, the up-


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regulated lysosomal cathepsin B could further promote the release of GEM. By employing cascade enzymatic reactions, the GEM nanocarriers could achieve prolonged circulation time while maintaining enhanced cellular internalization and effective drug release. Meanwhile, compared to free GEM, the deamination of GEM into inactive 2’,2’-difluorodeoxyuridine (dFdU) can be greatly suppressed, while the concentration of the activated form of GEM (gemcitabine triphosphate, dFdCTP) was significantly increased in tumor tissue, thus exhibiting superior tumor inhibition activity with minimal side effects. Using the same concept, dual-enzyme sensitive MSN was developed for tumor targeting and controlled drug release using MMP-2 and hyaluronidase.121 As a programmed targeting strategy, MMP-2 triggered detachment of PEG was also utilized for gene delivery122 and gene/drug co-delivery.123 A very interesting example was reported by Li et al.124 They designed tumor microenvironment-adaptive nanocarriers loaded with PTX and Twist-targeting siRNA. The nanocarriers consisted of a pH-responsive core, a cationic shell, and a MMP-cleavable PEG corona. PEG could be cleaved by MMP in the tumor tissue, which endowed the nanocarriers with smaller size and higher positive charge, leading to enhanced cellular uptake and intra-tumor accumulation of both PTX and siRNA. Moreover, endo/lysosomal pH-triggered drug release was achieved owing to the pH-sensitive core. As a result, the MMP and pH dual-sensitive nanocarriers could significantly inhibit tumor growth and pulmonary metastasis. The PEG-detachable drug nanocarriers could also be fabricated by iminobiotin-neutravidin non-covalent interaction. The iminobiotin- neutravidin interaction is a pH-dependent bond which is stable in physiological environment and can be easily decomposed in acidic environment.125 Hammond et al prepared layer-by-layer nanoparticles with pH-sheddable PEG corona for targeted cancer therapy.126 The poly-L-lysine (PLL) coated nanoparticles were modified with


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iminobiotin, followed by the linker neutravidin. Biotin end-functionalized PEG was then attached to the nanoparticles. Upon the nanoparticles were accumulated in tumor tissue, the iminobiotin-neutravidin bonds between PEG and PLL were cleaved, resulting in the exposure of cationic PLL. The exposed PLL layer could improve cellular uptake of drug nanocarriers. Unmasking of PEG corona can also be achieved by light-responsive drug nanocarriers. Löwik et al developed constrained and UV-activatable TAT for intracellular delivery of liposomes.127 They incorporated UV-activatable TAT into a loop anchored to PEGylated liposomes by two alkyl-chains of which one contained a light-cleavable nitrobenzyl group. The PEG corona could successfully shield TAT with stealth behavior. However, after the liposomes were arrived at the targeted site, nitrobenzyl linkage could be cleaved after UV irradiation, followed by loopopening and exposure of TAT. Therefore, simultaneous stealth behavior and TAT-mediated cellular uptake was achieved by UV activation. Using the same idea, an ABA-typed polymer, octadecyl-polyethylene glycol (biotin)-(o-nitrobenzyl)-octadecyl ester (CPB-p-C) with an onitrobenzyl group inserted between PEG and octadecyl ester was synthesized.128 The targeted biotin moieties could be exposed after light cleavage of nitrobenzyl linkages due to the stretching process of PEG segments, leading to receptor mediated targeted delivery. In another fantastic example, Wang et al prepared photoinduced PEG detachable nanocarriers for on-demand drug delivery (Figure 5).129 The reactive oxygen species (ROS)-sensitive thioketal (TK)-linked block copolymers of PEG and polylactide (PLA) were synthesized and then self-assembled into micelles. The photosensitizer Chlorin e6 (Ce6) and paclitaxel (PTX) were co-encapsulated into the micelles. The outer PEG shell rendered the nanoparticles long circulation time. After 660 nm laser irradiation, the generated ROS could cleave the TK linkages and PEG corona would be removed, leading to light-triggered enhanced cellular uptake and improved antitumor efficacy.


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Figure 5. Schematic illustration of the ROS-sensitive thioketal (TK) linked copolymer nanocarriers for on-demand drug delivery. After 660 nm laser irradiation, the generated ROS could cleave the TK linkages and PEG corona would be removed for enhanced cellular uptake.129 Copyright 2018, Elsevier B.V. Besides PEG, polyanionic peptide provides an alternative to fabricate stealth drug nanocarriers.130 However, it encounters the same predicament as PEG owing to its poor cellular uptake. The envelope-type nanocarriers were applied to balance this contradiction. In an pioneering research, Tsien et al designed MMP-2 activatable cell penetrating peptides (CPPs).131 The cationic CPPs were shielded by polyanionic peptides via MMP cleavable linker PLGLAG. On this occasion, the cell penetrating function of CPPs was inhibited by the outer polyanionic peptides. However, if polyanionic peptides were detached by overexpressed MMP in tumor sites, the cell penetrating function of CPPs would be restored. Surprisingly, more than 10-fold cellular uptake was observed upon linker cleavage. Therefore, stealth polyanionic peptides and targeted CPPs were perfectly integrated together to achieve stealth surface as well as enhanced cellular uptake. Such MMP activatable CPPs were followed and well developed by many researchers.132-


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134

The nanocarriers were used for targeted delivery of doxorubicin (DOX),135 paclitaxel

(PTX),136 DNA,137 docetaxel (DTX),138 and so on. The targeting ligand RGD can be activated by the detachment of polyanionic peptides as well. Zhang et al designed multifunctional envelopetype MSN for tumor-triggered targeted drug delivery to cancer cells (Figure 6). In this research, DOX was encapsulated in the mesoporous silica core. β-Cyclodextrin (β-CD) was anchored on the surface of MSNs as a gatekeeper via disulfide bond. RGD was then introduced via host-guest interaction. Furthermore, RGD was buried by polyanionic peptides via MMP cleavable peptide PLGVR.139 The polyanionic peptides endowed MSN with stealth ability and prevented them from uptaken by normal cells. However, the polyanionic peptide layer could be detached by MMP-2 triggered cleavage of PLGVR peptide, leading to the exposure of targeting ligand RGD and subsequently enhanced cancer cell uptake. After cellular uptake, DOX could be released since the β-CD gatekeeper could be removed in response to high concentration of glutathione. Such envelope-type MSN achieved enhanced cell growth inhibition efficacy in cancer cells. Very recently, MMP-2 triggered programmed tumor targeting was applied in photodynamic therapy.140 In this design, the nanocarriers were prepared by the co-modification of gold nanoparticles (AuNPs) with hydrazone-linked 5-aminolevulinic acid (ALA) and MMP-2activatable

CPPs.

CPP

was

shielded

by

zwitterionic

stealth

peptide

EKEKEKEKEKEKEKEKEKEK (EK10) via MMP-2 substrate peptide PLGLAG. The functional shift of stealth ability of EK10 to cell penetrating ability of R8 was achieved upon exposed to tumor-microenvironment-overexpressed MMP-2. After cellular uptake, ALA were converted to photosensitizer protoporphyrin IX (PpIX), exhibiting superior photodynamic therapeutic efficacy. Near infrared (NIR) activatable CPP was another good choice to avoid the nonspecific cellular uptake.141 Polyanionic peptide was conjugated to CPP with a photocleavable


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4,5-dimethoxy-2-nitrobenzyl group. After systemic administration, the nanocarriers showed excellent stealth ability since the cell penetration ability of CPPs was shielded by outer polyanionic peptide. However, when the tumor tissue was irradiated with NIR laser, the cell penetration ability of the nanocarriers could be effectively activated upon the outer polyanionic peptide was removed because of the light cleavage of nitrobenzyl linkages.

Figure 6. Design of multifunctional envelope-type MSN for enzyme-triggered detachment of polyanionic peptides for targeting drug delivery.139 Reproduced with permission. Copyright 2013, American Chemical Society. Another activatable protecting strategy is the use of biodegradable coating which can be specifically degraded in tumor microenvironment. For example, gelatin coated mesoporous silica nanocarriers were designed for programmed MMP-2 triggered tumor targeting.142 Hyaluronic acid (HA) is another ideal candidate since the concentration of hyaluronidase (HAase) in tumor extracellular matrix is 20-1000 times higher than that in normal tissues.143,144 Moreover, the


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excellent properties of non-toxicity, non-immunogenicity, and active targeting to CD44 receptor increase the potential application of HA in drug delivery.145-147 A typical example is the nanocarrier with a liposomal core and a cross-linked HA shell for sequential and site-specific drug delivery.148 Tumor necrosis factor-related apoptosis inducing ligand (TRAIL) and doxorubicin (Dox) were load into CPPs modified liposomes. Negatively charged HA modified with polymerizable acrylate groups was then attached to the liposomal surface by electrostatic interaction. The nanocarriers were cross-linked after UV irradiation to improve the stability. After intravenous administration, the nanocarriers exhibited a considerable accumulation in tumor tissues due to the passive and active targeting. The HA shell could undergo enzymatic degradation by overexpressed HAase and CPPs would be exposed, leading to enhanced cancer cell uptake and strong tumor inhibition. Zhang et al designed a more sophisticated ZnO based nanococktail with programmed targeting ability for self-synergistic cancer therapy.149 The positively charged nanoparticles (ZnO-DOX/R8) were protected by stealthy HA shell. HA was purposed to shield the positive charge to prolong circulation time. After accumulated in tumor tissue, the degradation of protective HA shell occurred, which could lead to the exposure of CPP R8 to facilitate cellular uptake. Subsequently, accompanied with endo/lysosomal pH triggered degradation of ZnO, reactive oxygen species (ROS) were produced in situ. ROS could not only lead to serious photodynamic cytotoxicity, but also result in ROS-triggered on-demand drug release, thus achieving enhanced anticancer efficacy synergistically. In order to decrease the affinity of targeting ligands for biological molecules during circulation, light-activatable caged ligands were introduced to drug delivery systems.150-152 Light-responsive drug delivery system is a very interesting research topic due to the excellent spatiotemporal resolution and controllability, which could be easily controlled by changing light intensity,


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wavelength, or irradiation area.153-155 o-Nitrobenzyl and coumarin are the most widely used photo-caged groups. As a proof-of-concept, Kohane et al prepared light-targeted nanoparticles by caging of targeted YIGSR peptide with a nitrobenzyl group.156 After irradiation with UV light (365 nm), the caged nitrobenzyl group was released and the targeting ligand YIGSR became active, leading to light-targeted enhanced cellular uptake. Some similar designs were well developed by caging of nonspecific CPPs and RGD with light-cleavable nitrobenzyl and coumarin groups.157-169 However, UV light is not an ideal candidate for biomedical applications due to its very limited tissue penetration and light toxicity. NIR laser is much more promising due to its minimal tissue absorption as well as much deeper tissue penetration. Mei et al prepared a series of two-photon-sensitive CPPs modified nanoparticles for NIR triggered targeted drug delivery.160-162 Up-conversion nanoparticles (UCNP) provide another important research direction to fabricate NIR-responsive drug delivery systems. In order to improve the targeting selectivity of folic acid (FA), Yeh et al designed UCNP-based drug delivery systems with NIR photocontrolled targeting, bioimaging, and chemotherapy (Figure 7).163 DOX and nitrobenzyl caged FA were conjugated to UCNPs. After irradiation with 980 nm NIR laser, the emitted UV light from UCNPs could cleave photolabile nitrobenzyl groups, resulting in the recovery of the targeting ability of FA. Therefore, selective targeting was achieved with reduced adverse effects in cancer chemotherapy.


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Figure 7. Schematic illustration of photo-caged UCNPs. The cargo molecules could be removed upon NIR laser irradiation and subsequent targeting of cancer cells was achieved.163 Reproduced with permission. Copyright 2013, American Chemical Society. In order to design stealth corona-detachable drug delivery systems, photo-activatable shielded ligands were developed based on the photothermal effect. In a very interesting work, Kohane et al conjugated targeting ligand YIGSR and thermo-responsive poly(N-isopropylacrylamine-coacrylamide) (pNIPAAm-co-pAAm) to the silica core-gold shell nanoparticles.164 The targeting ligand YIGSR immobilized on gold nanoshells was buried by thermo-responsive pNIPAAm-copAAm to avoid the off-target effects during circulation. Upon NIR irradiation, the nanoparticles would convert NIR light into heat, leading to the collapsion of pNIPAAm-co-pAAm and subsequent exposure of YIGSR. Light-targeted cellular uptake would then be achieved. Importantly, this approach showed great potential to obtain spatiotemporal control of cellular uptake, resulting in enhanced therapeutic efficacy and minimal side effects. The localized photothermal heating was also used to realize light-activable cancer-targeting of aptamer modified nanoparticles.165 In this research, aptamers

that

were

caged

by

their


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complimentary sequences were conjugated to gold nanorods (GNRs). After NIR laser irradiation, photothermal heating would induce the dehybridization of the double-strand DNA and uncage the aptamer structures. The aptamer conjugated GNRs could then recognize specific types of cancer cells after dehybridization, which was very promising for highly specific and selective cancer therapy. Charge-reversal drug delivery systems The surface charge of drug nanocarriers can greatly influence their in vivo fate.166 Generally, stealth negatively charged nanoparticles can resist protein adsorption and exhibit prolonged blood circulation time, thereby facilitating tumor accumulation by the EPR effect. However, because of the weak interaction between negatively charged nanoparticles and anionic cell membranes, the internalization of negatively charged nanoparticles is a big challenge.167 In contrast, positively charged nanoparticles show enhanced cell uptake owing to the strong affinity to anionic cell membranes.168,169 However, positively charged nanoparticles always suffer from high cytotoxicity, low plasma stability, and fast clearance by RES, which greatly limit their biomedical applications.170 Considering the prolonged circulation time of negatively charged surface and enhanced cell uptake of positively charged surface, it will be particularly attractive if the nanoparticles have stealth negatively charged surface during circulation. After accumulated in tumor tissue by EPR effect, the nanoparticles can reverse their surface charge from anion to cation in response to tumor microenvironment to achieve both prolonged circulation time and enhanced cell uptake. The as called “charge-reversal drug delivery systems” were therefore developed.167,171,172 This concept was firstly used in programmed tumor targeting by Wang’s group.173 They prepared amino functionalized poly(2-aminoethyl methacrylate) (PAMA) nanogels. The amino groups were then reacted with DMMA to generate carboxyl groups. The


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obtained amide bond is stable in physiological environment (pH 7.4), but can be cleaved in tumor extracellular environment (pH 6.8) to recover the positive amino groups. The DMMA functionalized nanogels with stealth negatively charged surface showed low protein adsorption and poor cell uptake at pH 7.4. However, enhanced cell uptake was observed when incubated at pH 6.8, which could be ascribed to the charge reversal ability. Furthermore, in vivo experiments also proved that such charge-reversal nanogels could be effectively uptaken by cancer cells using non-sensitive nanogel as a control. This charge-reversal concept was also applied for enhanced drug delivery to cancer stem cells. Wang et al designed dual pH-sensitive polymer-DOX conjugates for efficient drug delivery (Figure 8).174 They firstly synthesized a block copolymer of PEG and amino functionalized polyphosphoester. DOX was then conjugated to the polymer by an acid-labile hydrazone bond, which can be cleaved at endo/lysosomal pH (pH 5.0). The unreacted amino groups were further reacted with DMMA, which resulted in charge-reversal ability in acidic tumor extracellular environment (pH 6.8). After incubation of the nanoparticles at pH 6.8, the zeta potential of the nanoparticles changed from negative to positive within 10 min. Compared to the non-sensitive controls, DMMA modified nanocarriers can be easily uptaken by cancer stem cells at pH 6.8 as confirmed by confocal laser scanning microscopy (CLSM) and fluorescent activated cell sorting (FACS). After cell uptake, the covalently conjugated DOX could be released with the cleavage of hydrazone bond at endo/lysosomal pH and then localized in the nuclei, which exhibited enhanced cytotoxicity in drug resistant cancer stem cells. DMMA was also used to modify carbon dots.175 In this research, a tumor microenvironment-sensitive drug nanocarrier based on cisplatin(IV) prodrug-loaded chargereversal carbon dots (CDs-Pt(IV)@PEG-(PAH/DMMA)) was developed for imaging-guided drug delivery. The positively charged carbon dots were encapsulated into a nanocarrier by an


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anionic polymer with dimethylmaleic acid (PEG-(PAH/DMMA)). After the nanocarriers were accumulated in tumor tissue, anionic PEG-(PAH/DMMA) could undergo charge reversion to cationic PEG-PAH in acidic tumor microenvironment (pH 6.8), leading to the release of positively charged CDs-Pt(IV) due to the strong electrostatic repulsion. Moreover, positively charged CDs-Pt(IV) displays enhanced cancer cell internalization and effective activation of cisplatin(IV) prodrug in the reductive cytosol.

Figure 8. (a) Chemical structure of the dual pH-responsive polymer DOX Conjugate (PPC-HydDOX-DA) and schematic illustration of its pH triggered cellular internalization and intracellular drug release. (b) Zeta potential change of PPC-Hyd-DOX-DA NPs after incubation at pH 7.4 or 6.8 for different time periods. (c) Time-dependent cumulative release of DOX from PPC-HydDOX-DA NPs at different pH values.174 Reproduced with permission. Copyright 2011, American Chemical Society.


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As discussed above, the DMMA-based charge-reversal drug nanocarrier is considered as a powerful strategy to realize programmed tumor targeting. Many interesting trials can be found in the literatures.176-182 Some impressive examples are DMMA modified CPPs reported by Shen et al and Zhang et al.183-185 In their design, the amino groups in CPPs were reacted with DMMA to achieve deactivation of a CPP in blood circulation and subsequent activation in tumor tissue for enhanced cancer cell uptake. Taking advantage of the adjustable physicochemical properties of the nanoparticles by different stimuli, Wang et al explored the tumor acidity and near-infrared (NIR) light activated transformable nanoparticle

DATAT-NP

IR&DOX

using a tumor acidity-

activatable DMMA modified TAT.186 The physicochemical properties of

DATAT-NP

IR&DOX

can

be well controlled in a stepwise fashion using tumor acidity and NIR light, leading to adjustable nano-bio interactions in vivo. At first,

DATAT-NP

IR&DOX

was in “stealth” state during blood

circulation owing to the masking of TAT by DMMA. Once accumulated in the tumor tissues, DATAT-NP

IR&DOX

could be transformed into positively charged “recognize” state by the removal

of DMMA, which resulted in enhanced cellular internalization. Furthermore,

DATAT-NP

IR&DOX

could be transformed into “attack” state after internalization under NIR irradiation, achieving efficient DOX release. Recently, the concept of DMMA-based charge-reversal was also introduced to design pHresponsive light-up nanoparticle bioprobes with aggregation-induced emission (AIE) characteristics.187 The DMMA conjugated AIE-based nanoparticles were negatively charged and showed very weak fluorescence in physiological environment. However, the surface charge of the nanoparticle bioprobes would be switched to positive after DMMA was removed in extracellular tumor pH, leading to enhanced cellular uptake, significant fluorescence light up, and selective suppression of cancer cells.


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Taking advantage of strong proton-buffering capacity of poly(ethylenimine) (PEI), Shuai et al developed a class of charge-reversal drug delivery system.188 They synthesized triblock copolymer PEG-PAsp(MEA)-PEI, which can form nanoparticles with a pH-buffering core of PEI complexed with siRNA. By control the N/P ratio during preparation of the nanoparticles, the surface charge of the nanoparticles could be adjusted from negative to positive. Interestingly, owing to the reversible protonation of PEI, when the N/P ratio was 3, the nanoparticles were negatively charged at pH 7.4 but positively charged at pH 6.8, achieving tumor acidity-triggered charge reversal. In this situation, compared to the positively charged controls, the charge-reversal nanoparticles exhibited significantly longer circulation time and much better accumulation in tumor site via EPR effect. Meanwhile, the gene silencing efficacy of the charge-reversal nanoparticles was pH dependent. When incubated at pH 6.8, more effective gene silencing was observed because of the enhanced cell uptake after charge reversal at pH 6.8. Owing to pHinduced surface charge reversal from “negative” to “positive”, highly effective siRNA delivery leading to targeted gene silencing was achieved in animal studies, revealing great potential as an ideal nucleic acid delivery vector. Besides acidic tumor microenvironment triggered charge reversion for enhanced cellular uptake, hydrogen sulfide (H2S) triggered charge-reversal micelles were fabricated for cancertargeted drug delivery and imaging.189 The H2S sensitive micelles are very advantageous since cancer cells can produce large amounts of H2S.190 In this study, the authors prepared a series of N-(2-hydroxyethyl)-4-azide-1,8-naphthalimide ended amphiphilic diblock copolymer poly(2hydroxyethyl methacrylate)-block-poly(methyl methacrylate) (N3-Nap-PHEMA-b-PMMA-N3) micelles. Because azido groups can be selectively reduced by H2S, the surface charge of N3-NapPHEMA-b-PMMA-N3 micelles reversed from negative to positive after monitoring H2S in tumor


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tissue. Owing to the H2S triggered charge reversal, the internalization of DOX-loaded N3-ended micelles was much improved by electrostatic attraction mediated targeting. Therefore, N3-ended micelles can be served as a promising platform for tumor diagnosis and therapy. After accumulated in tumor tissue, it is extremely important that the nanocarriers are able to penetrate deep into the tumor to reach tumor cells. The penetration ability of the nanocarriers plays an important role in improving the therapeutic efficacy.191 It is reported that nanoparticles with large size favor blood circulation and tumor accumulation. However, the nanoparticles must be small to obtain excellent tumor-penetration ability.192,193 The surface charge of the nanocarriers could also influence the tumor penetration. Positively charged nanocarriers exhibited better tumor penetration ability, while neutral or negatively charged nanocarriers favor blood circulation.194,195 In order to overcome the above multiple physiological barriers imposes conflicting requirements for size and charge, the nanoparticles which can realize size reduction and charge reversal in response to the tumor microenvironment were developed for simultaneous improved tumor penetration and enhanced cellular uptake.196,167 Small sized poly(amidoamine) (PAMAM) dendrimers were frequently used to fabricate such size and charge adjustable nanocarriers.198 In a typical example, Wang and Nie et al synthesized pH sensitive PCL-CDMPAMAM/Pt by covalent conjugation of poly(ε-caprolactone) (PCL) and platinum (Pt) prodrug linked PAMAM with CDM linkage (Figure 9).199 The tumor acidity-sensitive clustered nanoparticles (iCluster/Pt) were elaborately prepared by co-assembly of PCL-CDM-PAMAM/Pt, PCL, and PEG-b-PCL. Upon incubated at pH 6.8, Pt linked PAMAM could be removed from iCluster/Pt due to the break of CDM linkage. In this situation, iCluster/Pt with the diameter of 104.1 nm were changed to very small Pt conjugated cationic PAMAM dendrimers. The small Pt linked PAMAM dendrimers could penetrate deep in tumor tissue. The Pt content in tumor tissue


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was much higher than pH-stable clustered nanoparticle. At the same time, the neutral PEG surface was changed to positively charged amino groups of PAMAM, leading to enhanced cellular uptake. After internalized by cancer cells, Pt could be rapidly released from PAMAM in reductive environment. It is very interesting that iCluster/Pt nanoparticles showed prolonged blood circulation, effective tumor penetration, enhanced cellular uptake, and rapid drug release, resulting in superior therapeutic efficacy. In another fantastic study, tumor acidity sensitive micelleplexes were prepared by self-assembly of DMMA modified poly(ethylene glycol)-blockpoly(2-amino ethyl aspartamide)-block-poly(ε-caprolactone) with Pt conjugated cationic PAMAM.200 After the cleavage of DMMA linkage at pH 6.8, Pt conjugated cationic PAMAM could be released from micelleplexes due to the electrostatic repulsion. Therefore, simultaneous size reduction and charge reversal was achieved under the stimulation of tumor acidity, leading to improved tumor penetration and enhanced cellular uptake. Similarly, multistage drug nanocarriers based on simultaneous size reduction and charge reversal were designed for programmed nuclear targeting.201 The nuclear-homing cell-penetrating peptide R8 conjugated DOX was grafted to N -(2-hydroxypropyl) methacrylamide (HPMA) copolymer by a hydrazone linkage, and then self-assembled with anionic DMMA grafted HPMA copolymer. The negatively charged nanocarriers with the diameter of 55 nm were obtained. After accumulated in tumor tissue, charge reversal of DMMA grafted HPMA copolymer occurred due to the cleavage of DMMA linkage, leading to the disassembly of the nanocarriers. DOX grafted HPMA copolymer with reduced size as well as positive charge could then penetrate deep to tumor tissue and then uptaken by cancer cells. After cellular uptake, R8 conjugated DOX was removed from HPMA copolymer for nuclear targeting due to the instability of hydrazone bond at endo/lysosomal pH. Very recently, Chen et al designed a shell-stacked nanoparticle (SNP) with DMMA modified


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pH-sensitive thick shell and reduction-sensitive cross-linked core.202 After accumulated in acidic tumor tissue, SNP could undergo surface charge reversal from -7.4 to 8.2 mV and size reduction from about 145 to 40 nm. By this design, SNP showed enhanced tumor penetration and could be uptaken by cancer cells in deep tumor tissue. Enzyme-responsive nanocarriers with simultaneous size reduction and charge reversal were reported by Gao et al.203 The MMP-2 responsive nanoparticles were prepared by conjugation of DOX loaded PAMAM and hyaluronic acid (HA) with PLGLAG peptide. The size could be decreased from ~200 nm to ~10 nm in the presence of MMP-2, accompanied with charge reversal. The enzyme-responsive size shrinkage and charge reversal could achieve faster nanoparticle penetration and cellular uptake. The intelligent design of such tumor microenvironment sensitive clustered nanoparticles is a very important advance in drug delivery area since they can achieve simultaneous improved tumor penetration and enhanced cellular uptake.

Figure 9. (A) Chemical structure of PCL-CDM-PAMAM/Pt. (B) Self-assembly of iCluster/Pt nanoparticles and their structural change in response to tumor acidity and intracellular reductive environment. (C) The hydrodynamic diameter of iCluster/Pt nanoparticles determined by DLS.


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(D) TEM images of iCluster/Pt nanoparticles incubated at pH 6.8 for 0, 4, and 24 h. (E) Quantitative determination of Pt content in tumor tissue cells by ICP-MS.199 Reproduced with permission. Copyright 2016, National Academy of Sciences. Tumor acidity-protonated drug delivery systems As described previously, neutral or negative nanoparticles have stealth surface and exhibit prolonged blood circulation time. Positively charged nanoparticles exhibit enhanced cellular uptake. Therefore, some protonated groups which can generate cations in response to tumor extracellular pH are introduced to fabricate drug delivery systems with stealth surface and enhanced cellular uptake.204,205 The imidazole group with pKa of approximately 7 is a very attractive candidate because of its reversible protonation ability. One typical example is the amino acid histidine (H), which offers the buffering effect upon protonation of the imidazole groups. As discussed above, the application of CPPs is greatly hampered by extensive penetration during circulation without proper selectivity. One way to address this dilemma is to develop pH-induced CPPs. Such pH-induced CPPs have weak cell penetrating ability during circulation and show strong cell penetrating ability after activation in the tumor microenvironment. Histidine-containing peptides have been well studied to fabricate pH-induced CPPs. For example, RRRRRRHHHH (R6H4) offers pH-dependent penetrating potential.206,207 In this design, arginine (R) can provide R6H4 with strong cell penetrating ability whereas protonatable histidine (H) endows R6H4 with pH-responsive cellular uptake. The balance between R and H in R6H4 can therefore be used as an important factor in programmed tumor targeting for improved cancer therapy. As an example, 5-aminolevulinic acid (ALA) pseudopolyrotaxane prodrug micelles with dual pH-responsiveness were constructed by hostguest interaction.208 The photosensitizer precursor ALA was conjugated to α-cyclodextrin (α-CD)


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through an acid-labile hydrazone bond. The pH-induced CPPs R6H4 was conjugated to PEG to form PEG-R6H4. The pseudopolyrotaxane prodrug micelles can be obtained owing to the inclusion complexation between α-CD and PEG. The cell-penetrating ability of the pseudopolyrotaxane prodrug micelles was deactivated during circulation to avoid nonspecific uptaken by normal cells. After arriving at tumor sites, the cell-penetrating ability can be activated owing to the protonation of histidine in acidic tumor extracellular environment. Therefore, the ALA prodrug micelles exhibited enhanced uptaken by cancer cells because of the tumor acidityresponsive R6H4. After internalization, ALA could be released by the cleavage of the hydrazone bond at endo-/lysosomal pH and converted to PpIX for targeted photodynamic therapy (PDT). Histidine-grafted chitosan was also prepared to fabricate stepwise pH-/reduction-responsive nanocarriers.209 The nanocarriers were negatively charged at physiological pH, which can reverse to positive at extracellular tumor microenvironment. Meanwhile, rapid DOX release was observed when incubation at endolysosome pH or high-concentration GSH. The prepared nanocarriers could achieve excellent stability during circulation, effective accumulation in tumor tissue, enhanced cellular uptake, and efficient intracellular DOX release based on the stepwise pH-/reduction responsiveness. Another typical example of this concept is the poly-L-histidine (polyHis)-based nanoparticle developed by Bae et al.210 PolyHis is composed of pH-sensitive imidazole groups and shows reversible protonation and deprotonation with a pKa of ~7.0. They prepared DOX loaded polyHis-b-PEG micelles. Because of partial protonation of polyHis block at pH 6.8, in vitro DOX uptake by A2780 cells at pH 6.8 was 5 times more than that of pH 7.4 for initial 20 min. More interestingly, a super pH-sensitive core-shell mixed micellar system was prepared by coassembly

of

polyHis-b-PEG

and

PLLA-b-PEG-b-polyHis-biotin.211

In

physiological


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environment (pH 7.4), polyHis is hydrophobic and in deprotonated state. The micellar core was composed of polyHis block and PLLA block. The targeting group biotin was hence hidden in the micellar core. Since the nonspecific targeting ligand biotin was not exposed, the nonspecific interaction with normal cells can be inhibited during circulation owing to the stealth PEGylated surface. In the acidic tumor microenvironment, polyHis block became hydrophilic and would be migrated from inner core to outer shell because of the protonation of histidine groups. Most biotin ligands would be exposed on the micellar surface, facilitating the uptaken of micelles by cancer cells. Furthermore, after internalized into endosome by endocytosis, the micelle showed pH-dependent dissociation, causing accelerated release of DOX in endosomal pH. At the same time, the dissociated micellar components could disrupt endosomal membrane owing to the proton sponge effect. DOX would be more avidly located in the nucleus, resulting in significant cytotoxicity. Using the same idea, TAT could also be hidden in the micellar core during circulation while exposed on the micellar surface after arriving in tumor sites upon acid-triggered protonation of polyHis.212 The polyHis-based tumor acidity-protonated drug nanocarriers could also be one-pot synthesized by simply mixing metal ions and organic ligands together with polyHis-b-PEG, in which the imidazole groups on the pHis chain could strongly bind with metal ions.213 Dual-pH sensitive enhanced cellular uptake as well as smart drug release was achieved in this research. Poly(2-(hexamethyleneimino) ethyl methacrylate) (PC7A) is an ultra pH sensitive polymer with a transition pH of about 6.9.214,215 In a typical example, Chen et al designed a hierarchical tumor microenvironment-responsive nanomedicine (HRNM) for programmed drug delivery (Figure 10).216 The HRNMs were prepared by the self-assembly of RGD conjugated block copolymer,

poly(2-(hexamethyleneimino)ethyl

methacrylate)-poly(oligo-(ethylene

glycol)


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monomethyl ether methacrylate)-poly[reduction-responsive camptothecin] (PC7A-POEGPssCPT). The shielding of RGD by POEG during circulation enabled HRNMs with long circulation time. After accumulated in tumor tissue, PC7A was protonated and converted from hydrophobic to hydrophilic by the acidic tumor microenvironment, leading to the exposure of RGD and enhanced tumor retention and cancer cell internalization. Moreover, intracellular GSH would trigger CPT release for cancer chemotherapy. Li et al reported a tumor acidity-triggered ligand-presenting (ATLP) nanoparticle for cancer therapy.217 PC7A-based acid-sensitive block copolymer was used as a sheddable matrix and iRGD-modified polymeric prodrug of doxorubicin (iPDOX) was used as an amphiphilic core. The ATLP nanoparticles could be specifically accumulated in the tumor tissue. Acid-triggered dissociation of the polymer matrix was observed due to the protonation of hydrophobic PC7A. iRGD ligand was then exposed, which was helpful for improved tumor penetration and enhanced cellular uptake. PC7A-based acid-sensitive block copolymer was also used to protect small interfering RNA (siRNA) for multistaged delivery.218 The charge-mediated complexes of siRNA and tumor cell-targeting- and penetrating-peptide- amphiphile (TCPA) was encapsulated by PC7A-based block copolymer. Because of the protonation of PC7A in acidic tumor microenvironment, siRNA-TCPA complexes could be released for specific targeting of tumor cells and cytosolic transport. Much more efficient gene silencing was observed due to the multistaged delivery strategy. Poly(βamino ester) (PAE)

is another smart polymer which can be protonated in tumor

microenvironment.219-221 Shi et al prepared mixed-shell micelles by co-assembly of PEG-b-PCL, PAE-b-PCL, and RGD-PAE-b-PCL.222 In physiological environment (pH 7.4), PAE was hydrophobic and RGD was hidden by PEG, which endowed the micelles with stealth property and 66% longer blood circulation half-life. Upon accumulated in tumor site, PAE was in a


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protonated hydrophilic state and the PAE chains were stretched, leading to the exposure of RGD for enhanced cancer cell internalization.

Figure 10. Schematic illustration of hierarchical tumor microenvironment-responsive nanocarriers for programmed drug delivery.216 Reproduced with permission. Copyright 2018, Wiley-VCH. PC7A was also adopted to prepare clustered nanoparticles which can change their size and surface charge in acid environment. In a pioneered research, Wang et al grafted PEG-b-PC7A to Pt loaded PAMAM.215 Because of the hydrophobic interaction of PC7A block, pH-sensitive clustered nanoparticles were obtained. However, the disintegration of the clustered nanoparticles into small PAMAM nanoparticles was observed at tumor acidic pH due to the protonation of PC7A. As expected, the very small Pt loaded PAMAM nanoparticles showed excellent penetration ability in poorly permeably pancreatic tumor models. Meanwhile, positively charged Pt loaded PAMAM nanoparticles could be easily uptaken by cancer cells. Furthermore, BLZ-945,


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a small molecule inhibitor of tumor-associated macrophages, could be physically encapsulated into

PEG-b-PC7A

conjugated

Pt

loaded

PAMAM

nanoparticles

for

cancer

chemoimmunotherapy.223 In this situation, BLZ-945 could be released after the disintegration of the clustered nanoparticles. Therefore, the strategy of differential delivery with simultaneous change of size and charge can achieve prolonged circulation, improved tumor penetration, enhanced cellular uptake, and site-specific differential drug release, representing a strategy to optimize the therapeutic efficacy. Protonation can not only change the surface charge, but also change the hydrophobicity, which may improve the cell penetrating ability. Reshetnyak et al reported pH (Low) Insertion Peptides (pHLIPs) with pH-dependent transmembrane activity.224 At neutral pH, pHLIP is water-soluble in an equilibrium state. However, the protonation of negatively charged residues (Asp or Glu) of pHLIP occurs at low pH, which enables pHLIP hydrophobic, increasing the affinity of the peptide for the lipid bilayer and triggering peptide folding and subsequent membrane insertion. Therefore, pHLIP can be used as a powerful ligand for targeting acidic diseased tissue and therefore enriches tumor-targeting delivery strategies.225-227 Engelman et al innovatively prepared pHLIP conjugated AuNPs, which could target various imaging agents to acidic tumors.228 The pHLIP technology could substantially enhance the accumulation and retention of AuNPs in tumors, which is extremely important for disease diagnosis and therapy. pHLIP was also used to fabricate acidic pH-targeted theranostic nanoplatform for the treatment of orthotopic pancreatic tumors.229 pHLIP was conjugated to chitosan-capped mesoporous silica coated gold nanorods and gemcitabine was then encapsulated into mesoporous silica. Thanks to the acidic pH triggered targeting specificity of pHLIP, the nanoparticles could be efficiently uptaken by cancer cells in tumor microenvironment, leading to significantly greater cytotoxicity. More importantly,


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pHLIP-conjugated nanoparticles were preferentially accumulated in the tumor tissue with more than 20 times stronger MOST signal than that of untargeted nanoparticles. Therefore, such pHsensitive targeting of pHLIP is particularly promising to fabricate smart nanoplatforms for simultaneous in vivo tumor imaging and drug delivery. For example, a pHLIP-dendrimer-drug conjugate system was recently reported for pH-triggered direct cytosolic delivery.230 Interestingly, pHLIP conjugated DOX could be inserted into membrane bilayers in a pHdependent manner, achieving low pH enhanced cellular uptake. Similarly, pHLIP conjugated mesoporous organosilica nanoparticles was synthesized by Lu et al for tumor acidic microenvironment targeted drug delivery.231 The engineered pHLIP modified nanoparticles exhibited higher cellular uptake at pH 6.5 than PEG decorated nanoparticles. Moreover, DOX which was loaded in the nanoparticles could be efficiently released in high GSH and low pH environment. The in vivo experiments showed that the pHLIP conjugated nanoparticles were more effective to be accumulated in the orthotopic breast cancer via targeting to acidic tumor microenvironment. Taking advantage of the pHLIP technology, pHLIP conjugated gold nanostars (GNS-pHLIP) were reported with pH-dependent transmembrane activity.232 Compared to PEG conjugated gold nanostars, GNS-pHLIP showed more efficient tumor accumulation and cellular uptake owing to the tumor acidity-triggered transmembrane activity. Moreover, the pHresponsive tumor targeting of GNS-pHLIP could significantly improve the photothermal therapy (PTT) efficacy with low side effects to normal tissues. In order to address the major challenges in photodynamic therapy, pHLIP was recently adopted to fabricate pH-driven photosensitizer nanocarriers. Li et al successfully absorbed pHLIP and photosensitizer Ce6 onto the surface of hollow gold nanospheres (HAuNS), which could be used for pH-driven and photo-responsive stepwise antitumor treatment.233 After arriving in tumor tissue, pHLIP could spontaneously form


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helical structures and swiftly translocate the nanocarriers into the cell because of the low pH of tumor microenvironment. Under laser irradiation, Ce6 was unloaded and dequenched from HAuNS owing to the photothermal effect of HAuNS, leading to enhanced photodynamic ablation of cancer cells. Zhang et al designed a self-transformable pH-driven membrane anchoring photosensitizer (pHMAPS) by direct conjugation of PpIX to the end of pHLIP for membrane localized PDT (Figure 11).234 In acidic tumor microenvironment, pHMAPS was able to form α-helix structure, which enabled successful insertion of pHMAPS into membrane lipid bilayer, leading to tumor-specific accumulation and in situ PDT on tumor cell membrane to maximize the therapeutic potency.

Figure11. Schematic illustration of the self-transformable pH-driven membrane-anchoring photosensitizer for pH-sensitive tumor-targeting delivery to achieve effective membrane localized PDT in fighting against cancer.234 Reproduced with permission. Copyright 2017, Wiley-VCH.


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Specific zwitterionic drug delivery systems As an alternative to PEGylation, Zwitterion is recently well established for their construction of stealth nanoparticles for drug delivery.235-238 Zwitterions exhibit better hydration ability than PEG, which can bind water molecules via electrostatic interactions, more strongly than those relying on hydrogen bonding. Zwitterions are well known for their outstanding non-fouling ability, long circulation time, low immunogenicity, and high biocompatibility. Very surprisingly, some zwitterions exhibit special cell uptake behavior because of their biomimetic molecular structures. Zwitterion might be one of the most simple and effective ways to realize both nonspecific resistance during circulation and enhanced cancer cell uptake in tumor sites. Phosphorylcholine (PC) is a biomimetic zwitterionic polar group of phospholipids on cell membranes. PC-based small molecular ligands and polymers were well developed for the fabrication of stealth inorganic nanoparticles, micelles, and nanogels.239-243 Ji et al prepared PC modified gold nanorods (GNRs) using 11-mercaptoundecylphosphorylcholine (HS-PC).244 The PC modified GNRs were very stable in PBS solution, blood plasma and cell culture medium. They showed strong resistance of protein adsorption owing to the stealth surface. Moreover, the cellular uptake difference of PC modified GNRs by normal cells and cancer cells was investigated using PEG modified GNRs as controls. Interestingly, the ICP-MS results reflected that PC modified GNRs can be internalized by cancer cells much more than that of PEG modified GNRs. Meanwhile, the amount of PC modified GNRs uptaken by cancer cells was 4 times more than that of normal cells. The cell uptake difference of PC modified GNRs between cancer cells and normal cells was observed in different kinds of cancer cell lines, suggesting PC as a universal targeting ligand of cancer cells. The mechanism of PC-based cancer cell targeting is still unknown. It might because most cancer cells present much higher level of choline


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phospholipid metabolism than normal cells.245 The elevation of phosphorylcholine provides a molecular target for different cancers. Similarly, Ishihara et al found amphiphilic phospholipid polymers were capable to provide energy-independent rapid cancer cell uptake via direct penetration across HepG2 cell membrane.246 Furthermore, Ji et al prepared PC conjugated DOX prodrug by conjugating DOX to 11-mercaptoundecyl phosphorylcholine via an acid-labile hydrazone linker.247 The PC prodrug micelles had exact and very high drug loading content (56.2%). They can strongly minimize nonspecific phagocytosis by macrophages owing to the stealth PC shell. Meanwhile, they exhibited better ability to be internalized by cancer cells than that of the PEG prodrug micelles, which indicated that the PC shell facilitated cancer cell internalization. The in vivo studies demonstrated that PC prodrug micelles showed significantly slower blood clearance and better tumor accumulation than PEG prodrug micelles. Very recently, a pH-sensitive poly(choline phosphate) prodrug was reported for rapid cellular internalization.248 Poly(2-(methacryloyloxy)ethyl choline phosphate)-b-poly(2-methoxy-2-oxoethyl methacrylatehydrazide-doxorubicin) (PCP-DOX) was synthesized with choline phosphate (CP) groups and DOX conjugation. Excellent non-fouling ability, rapid cellular uptake and efficient drug release could be easily realized by such simple molecular design. Because of the strong interaction between CP groups and cell membranes, PCP-DOX can be easily and rapidly internalized by various cancer cells.


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Figure 12. Schematic illustration of a pH-dependent zwitterionic-to-cationic charge conversion system.252 Reproduced with permission. Copyright 2015, Wiley-VCH. Carboxybetaine (CB) is another interesting zwitterionic molecule containing a positive quaternary amine group and a negative carboxylate group. CB modified nanoparticles own stealth surface and cannot induce antibody production. They exhibited much longer circulation time than PEG modified nanoparticles.249 Because of the protonation/deprotonation of carboxylic groups, CB is pH-sensitive. Ji et al prepared polyion complex micelles with CB shell for efficient protein delivery.250 The micelles had stealth surface in physiological pH (7.4). However, if the solution pH decreased to tumor extracellular pH (6.5), the carboxylate groups would be partial protonated. The zeta potential of CB modified micelles would be more positive at pH 6.5, which would result in enhanced cellular uptake. The flow cytometry results confirmed that the internalization of CB modified micelles at pH 6.5 was enhanced compared with that at pH 7.5, which might be attributed to the increased positive charge of the micellar surface. In another research, Chen et al designed a polymeric prodrug micelle system with a stealth poly(carboxybetaine) (pCB) shell.251 They found this drug delivery system had an instant and sensitive affinity switch from strong resistance in physiological pH to a high affinity to tumor cell membranes in the slightly acidic tumor extracellular pH by the protonation/deprotonation of CB groups in the outmost layer. Very recently, Rotello et al developed a pH-dependent zwitterionic-to-cationic charge conversion system based on the alkoxyphenyl acylsulfonamide groups (Figure 12).252 Acylsulfonamide-functionalized zwitterionic gold nanoparticles (AuNPs) were prepared which showed low cellular uptake at pH 7.4 owing to the stealth zwitterionic surface. When the acylsulfonamide-functionalized AuNPs were incubated at tumor extracellular pH (pH 6.5), the zeta potential increased from -2 mV to 17 mV, indicating a sharp transition


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from neutral to cationic. As a result, concomitant enhancement of cellular uptake and cytotoxicity on cancer cells were achieved at tumor extracellular pH. Very recently, such chargeswitchable zwitterionic nanoparticles with pH sensitive cellular uptake were successfully used for biorthogonal imaging of biofilm-associated infections.253

Figure 13. Schematic illustration of the targeting of acidic tumor microenvironment by pHresponsive mixed-charge zwitterionic AuNPs.255 Reproduced with permission. Copyright 2013, American Chemical Society. Although zwitterion has received considerable attention for stealth surface tailoring, the synthesis of zwitterionic molecules is always not as easy as expected. Mixed-charge surface engineering provides an easy and robust way to construct zwitterionic nanomedicine platforms by simply mixing cationic ligands and anionic ligands together. Especially, Mixed-charge ligands are very advantageous to fabricate stealth nanoparticles with switchable surface charge for adjustable cellular uptake.254 As an example, pH-responsive mixed-charge AuNPs were prepared by surface modification with strong electrolytic (10-mercaptodecyl)-trimethylammonium bromide and weak electrolytic 11-mercaptoundecanoic acid (Figure 13).255 By


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tailoring surface ratio of carboxylic groups to quaternary ammonium groups on AuNPs, AuNPs could respond to different pH ranges. The pH-sensitivity of the mixed-charge AuNPs were proved to be fast, reversible and ultrasensitive within very narrow pH range by protonation/deprotonation of carboxylic groups. When the feed ratio of carboxylic ligands and to quaternary ammonium ligands was 1:1, the AuNPs showed good stability, non-fouling ability, and prolonged circulation time during blood circulation. After arriving in tumor sites, the carboxylic groups would be partial protonated responding to the acidic tumor extracellular pH (pH 6.5), resulting in the aggregation and change of surface charge of AuNPs. Therefore, enhanced cell uptake in tumor sites was achieved. Moreover, the aggregation of mixed-charge AuNPs would lead to prolonged retention time in tumor sites due to the large size of aggregates and enhanced cellular uptake. By utilizing the reversible protonation of carboxylic groups, Lu et al prepared PEGylated mixed-charge gold nanostars using long-chain amine/carboxyl-terminated PEG.256 By optimizing the composition of amine/carboxyl-terminated PEG, the gold nanostars with zwitterionic stealth surface showed low cell affinity during circulation (pH 7.4). After arriving in tumor sites, the gold nanostars exhibited high cell affinity due to the break of surface charge balance by protonation of carboxylic groups. Conclusions and future perspectives The past decade has witnessed extensive exploration of cancer nanomedicine accompanied by the rapid development of nanotechnology, materials science, biology, pharmaceutics, and medicine. Thousands of publications suggested that nano-sized drug delivery systems are very effective in cancer treatment. Drug nanocarriers are expected to improve the biodistribution and pharmacokinetics of small drug molecules. Therefore, a large dose of drug can be delivered to tumor tissue for improved therapeutic index and reduced systemic toxicity. Unfortunately, only


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very few nanomedicines have been successfully applied in clinic due to the unsatisfactory clinical outcomes.257,258 We should be fully aware of the complexity and challenges of the drug delivery process. Moreover, the complexity of tumor tissues is far beyond our imagination. A full understanding of nano-bio interactions is necessary for the discovery of safer and more efficient nanomedicines. How can we make full use of “NANO” to solve the clinical problems in cancer treatment? There is still a long way to go to make the dream a reality. This review aims to present a survey on the design of drug nanocarriers for simultaneous stealth surface in circulation and enhanced uptake in tumor. Nanomedicine has been considered as a dimension to address the limitations of conventional chemotherapy.

There are many

obstacles that the nanocarriers should overcome to achieve improved therapeutic index. It is the premise that nano-sized drug delivery systems should have stealth surface to prolong the circulation time and accumulate in tumor sites. At the same time, efficient uptake by cancer cells is another prerequisite to improve the therapeutic efficacy. During the past decade, great efforts have been made to overcome the dilemma of stealth surface and enhanced cellular uptake. To realize effective accumulation in tumor sites and efficient uptake by cancer cells, we believe that two directions are particularly promising in the near future. Firstly, there is a great need to develop tumor microenvironment ultrasensitive drug nanocarriers which can be activated for programmed tumor targeting within a very narrow range of acidity or enzymatic activity. The differences in the physiological and biological characteristics between the tumor and normal tissues are being increasingly investigated as the basis for the design of tumor microenvironment sensitive drug nanocarriers. Although tumor microenvironment has been studied for many years, how to design nanocarriers to make full use of tumor microenvironment is still in its infancy. Because of the individual differences of different people, the tumor microenvironment might


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also be different. There are many challenges that we have to face. To this end, stealth coronadetachable drug delivery systems, charge-reversal drug delivery systems, and tumor acidityprotonated drug delivery systems were developed with activatable tumor targeting ability. However, owing to the complex structure design of tumor microenvironment sensitive drug nanocarriers, the production of these nanocarriers on an industrial scale is very difficult. In addition, the response time of the drug nanocarriers to tumor microenvironment is another important factor that should be considered. In this respect, compared to the cleavage of covalent bonds, protonation/deprotonation might be more advantageous because of the superfast response time. Secondly, because of the biomimetic zwitterionic structures, zwitterionic drug delivery systems were recently developed to obtain stealth drug nanocarriers with enhanced cancer cell uptake. It might be the easiest way to resolve this real conundrum owing to the simple molecular design strategy. In this respect, zwitterionic drug nanocarriers might be more practical for their application from animal to human models. On the other hand, endogenous signals including pH or enzymatic gradient are almost impossible to control. They might also be changed in different individuals. The design of stealth drug nanocarriers with enhanced cellular uptake is very challenging to achieve ideal therapeutic efficacy. As discussed above, drug delivery is a multistage procedure. Although the main objective of this review is to present the up to date approaches to overcome the dilemma of stealth surface in circulation and enhanced uptake in tumor, there are many other factors that can influence the drug delivery procedure. For example, protein corona is a very important issue that we should particularly pay attention to. Because of the existence of protein corona, the nanoparticles may not be pristine nanoparticles any more upon systematically administrated. The surface engineering is therefore extremely important to minimize the formation of protein corona. The


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stability of the nanocarriers is another critical issue that should not be ignored. Otherwise, drugs will be leaked from the nanocarriers during circulation and the nanocarriers will lose their functions. If very high concentration of nanocarriers is accumulated in tumor tissue, is it good enough for high efficacy? We should also consider the penetration ability of the drug nanocarriers. Last but not the least, drugs must be released from the nanocarriers after arriving in tumor tissue, which will also greatly influence the therapeutic efficacy in cancer treatment. All in all, it is still very far to resolve this matter once and for all. The future is bright; the road is tortuous.

Author information Corresponding Authors *E-mail: [email protected] (J. Ji). Acknowledgements Financial supports from Zhejiang Science and Technology Project (2016C04002) and the National Natural Science Foundation of China (21774110, 21574114), are gratefully acknowledged. VOCABULARY Nanomedicine, the medical applications of nanoscience and nanotechnology to achieve improved patient outcomes


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Drug delivery, refers to the approaches, formulations, and technologies for transporting drugs to the area where they are needed in the body. The drugs can therefore achieve desired therapeutic efficacy. Stealth, the naonparticles cannot be recognized by reticulo-endothelial system during circulation and have long blood circulation time. PEGylation, the nanoparticles are covered by PEG corona, which can endow the nanoparticles with stealth properties to improve the drug delivery efficiency. Tumor microenvironment, is the cellular environment in which the tumor exists, including surrounding blood vessels, Warburg effect induced acidic pH, hypoxia, specific up-regulated enzymes, and so on. References 1. Shi, J.; Kantoff, P. W.; Wooster, R.; Farokhzad, O. C. Cancer Nanomedicine: Progress, Challenges and Opportunities. Nat. Rev. Cancer 2017, 17, 20-37. 2. Hartshorn, C. M.; Bradbury, M. S.; Lanza, G. M.; Nel, A. E.; Rao, J.; Wang, A. Z.; Wiesner, U. 2. B.; Yang, L.; Grodzinski, P. Nanotechnology Strategies to Advance Outcomes in Clinical Cancer Care, ACS Nano 2018, 12, 24-43. 3. Peer, D.; Karp, J. M.; Hong, S.; Farokhzad, O. C.; Margalit, R.; Langer, R. Nanocarriers as an Emerging Platform for Cancer Therapy. Nat. Nanotechnol. 2007, 2, 751-760. 4. Sun, Q.; Zhou, Z.; Qiu, N.; Shen, Y. Rational Design of Cancer Nanomedicine: Nanoproperty Integration and Synchronization. Adv. Mater. 2017, 29, 1606628. 5. Wicki, A.; Witzigmann, D.; Balasubramanian, V.; Huwyler, J. Nanomedicine in Cancer Therapy: Challenges, Opportunities, and Clinical Applications. J. Control. Release 2015, 200, 138-157. 6. Matsumura, Y.; Maeda, H. A New Concept for Macromolecular Therapeutics in Cancer Chemotherapy: Mechanism of Tumoritropic Accumulation of Proteins and the Antitumor Agent Smancs, Cancer Res. 1986, 46, 6387-6392. 7. Maeda, H. Toward a Full Understanding of the EPR Effect in Primary and Metastatic Tumors as Well as Issues Related to Its Heterogeneity. Adv. Drug Deliv. Rev. 2015, 91, 3-6.


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8. Prabhakar, U.; Maeda, H.; Jain, R. K.; Sevick-Muraca, E. M.; Zamboni, W.; Farokhzad, O. C.; Barry, S. T.; Gabizon, A.; Grodzinski, P.; Blakey, D. C. Challenges and Key Considerations of the Enhanced Permeability and Retention Effect for Nanomedicine Drug Delivery in Oncology. Cancer Res. 2013, 73, 2412-2417. 9. Setyawati, M. I.; Tay, C. Y.; Bay, B. H.; Leong, D. T. Gold Nanoparticles Induced Endothelial Leakiness Depends on Particle Size and Endothelial Cell Origin. ACS Nano 2017, 11, 5020-5030. 10. Tay, C. Y.; Setyawati, M. I.; Leong, D. T. Nanoparticle Density: A Critical Biophysical Regulator of Endothelial Permeability. ACS Nano 2017, 11, 2764-2772. 11. Setyawati, M. I.; Mochalin, V. N.; Leong, D. T. Tuning Endothelial Permeability with Functionalized Nanodiamonds. ACS Nano 2016, 10, 1170-1181. 12. Wang, J.; Zhang, L.; Peng, F.; Shi, X.; Leong, D. T. Targeting Endothelial Cell Junctions with Negatively Charged Gold Nanoparticles. Chem. Mater. 2018, 30, 3759-3767. 13. Setyawati, M. I.; Tay, C. Y.; Chia, S. L.; Goh, S. L.; Fang, W.; Neo, M. J.; Chong, H. C.; Tan, S. M.; Loo, S. C. J.; Ng, K. W.; Xie, J. P.; Ong, C. N.; Tan, N. S.; Leong, D. T. Titanium Dioxide Nanomaterials Cause Endothelial Cell Leakiness by Disrupting the Homophilic Interaction of VE-cadherin. Nat. Commun. 2013, 4, 1673-1684. 14. Trantidou, T.; Friddin, M.; Elani, Y.; Brooks, N. J.; Law, R. V.; Seddon, J. M.; Ces, O. Engineering Compartmentalized Biomimetic Micro- and Nanocontainers. ACS Nano 2017, 11, 6549-6565. 15. Chen, Y.; Wu, Y.; Sun, B.; Liu, S.; Liu, H. Two-Dimensional Nanomaterials for Cancer Nanotheranostics. Small 2017, 13, 1603446. 16. Grimaldi, N.; Andrade, F.; Segovia, N.; Ferrer-Tasies, L.; Sala, S.; Veciana, J.; Ventosa, N. LipidBased Nanovesicles for Nanomedicine. Chem. Soc. Rev. 2016, 45, 6520-6545. 17. Li, Z. H.; Wang, H. B.; Chen, Y. J.; Wang, Y.; Li, H.; Han, H. J.; Chen, T. T.; Jin, Q.; Ji, J. pH- and NIR Light-Responsive Polymeric Prodrug Micelles for Hyperthermia-Assisted Site-Specific Chemotherapy to Reverse Drug Resistance in Cancer Treatment. Small 2016, 12, 2731-2740. 18. Cabral, H.; Kataoka, K. Progress of Drug-Loaded Polymeric Micelles into Clinical Studies. J. Control. Release 2014, 190, 465-476. 19. Satalkar, P.; Elger, B. S.; Hunziker, P.; Shaw, D. Challenges of Clinical Translation in Nanomedicine: A Qualitative Study. Nanomedicine 2016, 12, 893-900. 20. Min, Y.; Caster, J. M.; Eblan, M. J.; Wang, A. Z. Clinical Translation of Nanomedicine. Chem. Rev. 2015, 115, 11147-11190. 21. Rodríguez-Nogales, C.; González-Fernández, Y.; Aldaz, A.; Couvreur, P.; Blanco-Prieto, M. J. Nanomedicines for Pediatric Cancers. ACS Nano 2018, 12, 7482-7496. 22. Deng, Z. J.; Morton, S. W.; Ben-Akiva, E.; Dreaden, E. C.; Shopsowitz, K. E.; Hammond, P. T. Layer-by-Layer Nanoparticles for Systemic Codelivery of an Anticancer Drug and siRNA for Potential Triple-Negative Breast Cancer Treatment. ACS Nano 2013, 11, 9571-9584. 23. Deng, Y.; Jia, F.; Chen, S.; Shen, Z.; Jin, Q.; Fu, G.; Ji, J. Nitric Oxide as an All-Rounder for Enhanced Photodynamic Therapy: Hypoxia Relief, Glutathione Depletion and Reactive Nitrogen Species Generation. Biomaterials 2018, 187, 55-65.


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24. Björnmalm, M.; Thurecht, K. J.; Michael, M.; Scott, A. M.; Caruso, F. Bridging Bio–Nano Science and Cancer Nanomedicine. ACS Nano 2017, 11, 9594-9613. 25. Ke, P. C.; Lin, S.; Parak, W. J.; Davis, T. P.; Caruso, F. A Decade of the Protein Corona. ACS Nano 2017, 11, 11773-11776. 26. Setyawati, M. I.; Tay, C. Y.; Docter, D.; Stauber, R. H.; Leong, D. T. Understanding and Exploiting Nanoparticles' Intimacy with the Blood Vessel and Blood. Chem. Soc. Rev. 2015, 44, 8174-8199. 27. Docter, D.; Westmeier, D.; Markiewicz, M.; Stolte, S.; Knauer, S. K.; Stauber, R. H. The Nanoparticle Biomolecule Corona: Lessons Learned - Challenge Accepted? Chem. Soc. Rev. 2015, 44, 6094-6121. 28. Salvati, A.; Pitek, A. S.; Monopoli, M. P.; Prapainop, K.; Bombelli, F. B.; Hristov, D. R.; Kelly, P. M.; Åberg, C.; Mahon, E.; Dawson, K. A. Transferrin-Functionalized Nanoparticles Lose Their Targeting Capabilities When A Biomolecule Corona Adsorbs on the Surface. Nat. Nanotechnol. 2013, 8, 137-143. 29. Lane, L. A.; Qian, X.; Smith, A. M.; Nie, S. Physical Chemistry of Nanomedicine: Understanding the Complex Behaviors of Nanoparticles in Vivo. Annu. Rev. Phys. Chem. 2015, 66, 521-547. 30. Mitragotri, S.; Lahann, J. Physical Approaches to Biomaterial Design. Nat. Mater. 2009, 8, 15-23. 31. Tay, C. Y.; Setyawati, M. I.; Xie, J.; Parak, W. J.; Leong, D. T. Back to Basics: Exploiting the Innate Physico-chemical Characteristics of Nanomaterials for Biomedical Applications. Adv. Funct. Mater. 2014, 24, 5936-5955. 32. Wang, J.; Mao, W.; Lock, L. L.; Tang, J.; Sui, M.; Sun, W.; Cui, H.; Xu, D.; Shen, Y. The Role of Micelle Size in Tumor Accumulation, Penetration, and Treatment. ACS Nano 2015, 9, 7195-7206. 33. Jiang, W.; Kim, B. Y. S.; Rutka, J. T.; Chan, W. C. W. Nanoparticle-Mediated Cellular Response is Size-Dependent. Nat. Nanotechnol. 2008, 3, 145-150. 34. Chithrani, B. D.; Ghazani, A. A.; Chan, W. C. W. Determining the Size and Shape Dependence of Gold Nanoparticle Uptake into Mammalian Cells. Nano Lett. 2006, 6, 662-668. 35. Jiang, W.; Kim, B. Y. S.; Rutka, J. T.; Chan, W. C. W. Nanoparticle-Mediated Cellular Response is Size-Dependent. Nature Nanotechnol. 2008, 3, 145-150. 36. Lamprecht, A.; Schäfer, U.; Lehr, C.-M. Size-Dependent Bioadhesion of Micro- and Nanoparticulate Carriers to the Inflamed Colonic Mucosa. Pharm. Res. 2001, 18, 788-793. 37. Barua, S.; Yoo, J.-W.; Kolhar, P.; Wakankar, A.; Gokarn, Y. R.; Mitragotri, S. Particle Shape Enhances Specificity of Antibody-Displaying Nanoparticles. Proc. Natl. Acad. Sci. U. S. A. 2013, 110, 3270. 38. He, Y.; Park, K., Effects of the Microparticle Shape on Cellular Uptake. Mol. Pharmaceutics 2016, 13, 2164-2171. 39. Sun, Q.; Zhou, Z.; Qiu, N.; Shen, Y. Rational Design of Cancer Nanomedicine: Nanoproperty Integration and Synchronization. Adv. Mater. 2017, 29, 1606628. 40. Yang, R.; Wei, T.; Goldberg, H.; Wang, W.; Cullion, K.; Kohane, D. S. Getting Drugs Across Biological Barriers. Adv. Mater. 2017, 29, 1606596.


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Rational Design of Cancer Nanomedicine for Simultaneous Stealth

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