Multifaceted Transport Characteristics of Nanomedicine: Needs for

Review pubs.acs.org/molecularpharmaceutics

Multifaceted Transport Characteristics of Nanomedicine: Needs for Characterization in Dynamic Environment Altug Ozcelikkale,† Soham Ghosh,† and Bumsoo Han*,†,‡ †

School of Mechanical Engineering and ‡Weldon School of Biomedical Engineering, Purdue University, West Lafayette, Indiana, United States S Supporting Information *

ABSTRACT: Nanomedicine for cancer, where nanoparticles (NPs) are used to deliver drugs, imaging agents, and heat to tumors, shows great potential of improved therapeutic outcomes. In spite of promising early stage results, its clinical efficacy is still significantly limited due to complex transport barriers in vivo. These transport barriers are associated with tumor microenvironment, which is highly complex and heterogeneous and varies spatiotemporally. Thus, in order to improve the in vivo efficacy of nanomedicine, NPs need to be designed and characterized considering their interaction with these complex transport barriers. In this article, thus, we discuss the multifaceted transport characteristics of NPs and their interaction mechanisms with the tumor microenvironment. We also illustrated that NPs have highly spatiotemporal and multiscale distribution around tumor. This dynamic and complex nature of NP transport needs to be taken into consideration in design strategies and evaluation criteria for successful delivery of cancer nanomedicine. KEYWORDS: nanomedicine, nano particle, cancer therapy, diagnosis, transport characteristics

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INTRODUCTION

small molecular drugs used in chemotherapy (e.g., doxorubicin and paclitaxel),27,28 and a few of these are in clinical use.29,30 Some NPs are used as mediators of localized heating in thermal therapy.31 Upon excitation by external magnetic field,32,33 radiofrequency electric field,20,34,35 or laser irradiation,36−38 NPs can generate heat, and the resulting hyperthermia can kill cancer cells.16,36,38 NPs are also being considered for improving cryosurgery by enhancing the destruction of diseased tissue by freezing.39 In addition to therapeutic tools, NPs are also being developed as diagnostic tools in various imaging modalities. Iron oxide nanocrystals and carbon nanotubes are used to enhance the signal-to-noise ratio in magnetic resonance imaging (MRI).40,41 Gold nanoshells and nanorods, effective in photothermal therapy, are also shown to enhance the signal for optical coherence tomography (OCT).37,42 Quantum dots21,22,43,44 and fluorescent NP probes6 have been commonly used in near-infrared fluorescence imaging. Recently, development of multifunctional NPs is actively pursued by combining therapeutic and diagnostic functions.

Recent advances in nanotechnology provide a wide variety of nanoparticles (NPs). These enable nanomedicine, in which NPs can mediate new and innovative diagnosis and treatment of various diseases.1 Among these, nanomedicine for cancer shows great potential for improved therapeutic outcomes, where NPs are used to deliver drugs, imaging agents, and heat to tumor tissues. The term “nanoparticle” is typically used to refer to particles with sizes between 1 and 100 nm.2 These include organic and inorganic polymers, as well as metals and semiconductors formed into solid particles, porous meshes, hollow shells, or vessels. They can have spherical, hemispherical, disk, rod, tube, or horn shape. Common designs of NPs for cancer nanomedicine are polymeric degradable NPs,3−6 liposomes,7−10 polymeric micelles,4,11,12 carbon nanotubes and nanohorns,13,14 colloidal gold, silver, and iron oxide NPs,15−19 and quantum dots.20−22 Typical size, shape, and applications of those NPs are summarized in Table 1. NPs show great promise as carriers for cancer therapeutic agents. This is because they can offer improved solubility and stability of the drug compounds carried, prolonged circulation in the bloodstream, and controlled release of the drug23 with potential decrease of side effects.3,24 As such, liposomes, polymeric micelle structures, and degradable polymeric nanoparticles3,25,26 have been successfully applied to encapsulate © 2013 American Chemical Society

Special Issue: Emerging Technology in Evaluation of Nanomedicine Received: Revised: Accepted: Published: 2111

October 18, 2012 March 18, 2013 March 21, 2013 March 21, 2013 dx.doi.org/10.1021/mp3005947 | Mol. Pharmaceutics 2013, 10, 2111−2126

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Table 1. Summary of NPs Commonly Used for Cancer Therapy and Diagnosis type of NP

a

size,a D [nm]

available shapes

selected applications

polymer NP

50−300

sphere3,197 rod86,126 disk86

drug delivery3 imaging (fluorescence)197

liposome

100−500

sphere135,193

drug delivery135,193

polymeric micelle

20−200

sphere4,198

drug delivery4,198

carbon NP

d = 0.5−3b l > 10

tube/rod34 sphere/aggregate14

drug delivery13,14 thermal therapy (radiofrequency)20

quantum dot

3−30

sphere20,45,90

drug delivery45 thermal therapy (radiofrequency)20 imaging (fluorescence)20,90

gold NP

5−250

sphere15,57 rod148,163 core−shell36,38

drug delivery15 thermal therapy (photothermal)36,38,148,163 imaging (OCT)148

iron oxide NP

5−250

sphere38

thermal therapy (magnetic)32 imaging (MRI)38,40,90

The values are adapted from Chou et al.199 unless otherwise noted. bd: Cross-section diameter. l: Length.

Figure 1. Overview of NP transport processes to tumor. NPs that are not cleared by the reticuloendothelial system (RES) arrive at the tumor vasculature via blood-borne transport. Then, they interact with the vascular endothelium and selectively extravasate into the tumor interstitium, via transvascular transport. In the subsequent interstitial transport, NPs penetrate to the proximity of cancer cells. During this process, they encounter multiple transport barriers at the interstitium including elevated interstitial fluid pressure, dense extracellular matrix (ECM) structure, and high cell packing density. After the interstitial transport, NPs finally interact with the tumor cells to deliver drugs and imaging agents.

a transport challenge since the delivery and distribution of NPs to and within the tumor are known to be hindered by multiple transport barriers associated with tumor microenvironment.48−50 When NPs are administered to a patient’s bloodstream, they experience complex and multifaceted transport processes as illustrated in Figure 1. First, the NPs will arrive to the tumor vasculature via blood flow driven transport. Then, they interact with the vascular endothelium and selectively extravasate into the tumor interstitium, i.e., transvascular transport. During the subsequent interstitial transport, the NPs encounter multiple transport barriers at the interstitium including elevated interstitial fluid pressure, dense extracellular matrix (ECM) structure, and high cell packing density. After the interstitial transport, the NPs finally

This type of NP enables monitoring of the progress of drug delivery and the development of adaptive and personalized treatment plans.18,38 For instance, quantum dots were considered for carrying therapeutics and simultaneously monitoring their release in cancer cells.45 Encapsulating NPs with different functions into a single unit also has been proposed to achieve multifunctionality. This approach has been illustrated by a polymeric micelle that encapsulates doxorubicin for cancer therapy and at the same time contains iron oxide particles for MRI and quantum dots for fluorescence imaging.18 Despite very promising initial results, the clinical efficacy of most NP-mediated cancer therapy and diagnostic modalities is still significantly limited.46 One of the main reasons is the poor delivery of NPs to the targeted tumor tissue.47 This is primarily 2112

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interact with the tumor cells to deliver drugs and imaging agents. The excess NPs are thought to be either retained at the tumor or drained to the lymphatics. These processes are highly dynamic and interconnected, and their compounding effects on the in vivo NP transport are not yet fully understood. Limited understanding of how the NP properties affect these transport processes and the lack of proper evaluation methods for the in vivo NP transport characteristics are critical bottlenecks in development of effective nanomedicine. In this article, an overview of recent literature on cancer nanomedicine is provided with focus on the transport characteristics of NPs. This is followed by a discussion of other transport considerations that are relevant to cancer treatment and diagnosis. Those include heat transfer during NP-mediated thermal therapies and light−tissue interactions during in vivo imaging of NPs. Then, current challenges and opportunities for improved nanomedicine are discussed with two case studies on multiscale transport around tumors.

Table 2. Transport Properties of Selected Macromolecules and NPs at Different Stages of in Vivo Transport to Tumor 1. Circulation in Bloodstream

IgGb DOXIL (PEG-liposome) Genexol-PM (PEG-micelle) carbon nanotube

description mouse IgG dextran, 40 kDa dextran, 2 MDa liposome

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TRANSPORT PROCESSES TO TUMOR Blood-Borne NP Transport. After systemic administration (i.e., intravenous injection), NPs circulate in the bloodstream to reach the target tumor. While in circulation, the majority of NPs are eliminated from the bloodstream within minutes or hours after injection.4,51,52 It has been reported that NPs smaller than 5.5 nm are found to pass through glomerular capillary walls at kidneys and undergo rapid renal filtration.53,54 In addition, NPs are coated by serum proteins in the bloodstream. This process, referred to as opsonization, is an adaptation to recognize and eliminate foreign substances from the body. NPs tagged by the serum proteins are removed from the bloodstream by the reticuloendothelial system (RES),1 composed of blood-associated macrophages and other phagocytic cells in liver and spleen. Particles, with 10 to 20 nm hydrodynamic diameter, are uptaken by liver, where they are either internalized and withheld by Kupffer cells or transferred to bile and removed from the body by hepatocytes.54 Larger particles, with sizes up to several micrometers, can be captured by macrophages in spleen.55 There are numerous studies showing a significant portion of injected NP dose to be received by these two organs. In fact, the total amount of NP accumulation at liver and spleen can be more than an order of magnitude greater than what is accumulated in the tumor.9,12,56 From a transport perspective, this is definitely an undesirable outcome since the removal of NPs from the bloodstream precludes any form of delivery to the target tumor. Besides, NPs retained in various organs can induce cytotoxicity and associated side effects. Therefore, a great amount of effort has been devoted to designing NPs with prolonged blood circulation. The plasma half-lifes of some macromolecules and NPs are reported in Table 2. The size of NP can have an effect on its plasma half-life. It was reported that liposomes that are larger than 300 nm or smaller than 70 nm were associated with reduced blood circulation times when compared to intermediate sizes.51 Yet, among gold NPs with diameters between 20 and 100 nm, 60 nm particles had the longest plasma halflife.57 Block-copolymer micelles of 25 nm diameter were also reported to clear from plasma much faster than their 60 nm counterparts.12 While the results of these studies might seem to favor intermediately sized particles, there is little consensus on optimal size for long circulating NPs. In a study that considered polymeric micelles between 30 and 100 nm, the greatest tumor

a

plasma half-life t1/2 [h]

DH [nm]a

description

ref

9.7 80−90

34.7 84.0

200 201

20−50

11.0

201

d = 1, l = 300−1000 2.99−3.52 2. Extravasation at Tumor Vasculature DH [nm]

microvascular permeability Pv [10−9 m/s]

9.88 10 38.4 120 3. Penetration into Tumor

2.82 9.5 1.7 0.155 Interstitium

52

ref 80 82 82 202

description

DH [nm]

effective diffusivity Deff [10−12 m2/s]

ref

IgG dextran, 40 kDa dextran, 2 MDa liposome

9.7 10.0 38.4 152.4

12.7 42.0 1.44 0.297

169 203 169 7

Hydrodynamic diameter. bImmunoglobulin G.

accumulation and anticancer activity was observed for the smallest 30 nm particles.58 Tumor accumulation was also found to be highly sensitive to variations in particle size between 111 and 166 nm.6 While the former accumulated in the tumor, the latter was predominantly located in liver. Taken together, experimental evidence indicates that the plasma half-life of NPs and the associated tumor accumulation amounts indeed depend on the particle size, but identifying the relationships among those three is a part of ongoing research. Some studies in literature suggest that particle shape is also an important factor affecting the plasma half-life. An example is the filament-like polymer-micelles, referred to as filomicelles, that can have 10 times greater circulation time than their spherical counterparts.59 Interactions of NPs with the serum proteins are affected by the NP surface charge as well. It is now well-known that charged particles are readily opsonized by the plasma proteins of opposite charge and, as a result, have a significantly reduced plasma half-life when compared with their neutral counterparts.60−62 Coating of NP surface by poly(ethylene)glycol (PEG) significantly reduces the extent of opsonization and help NPs escape from RES.51,63,64 PEGylation has been shown to improve the plasma half-life of many types of NPs.3,6,51,56,57,61 However, it can have adverse effects on the uptake rate of NPs by cancer cells.17 An alternative approach to provide stealth features to NPs is proposed by mimicking the way red blood cells remain in blood circulation. CD47 is a ligand expressed in red blood cells that interacts with the signal regulatory protein alpha (SIRPα) on macrophages and inhibits phagocytosis. Significant decrease was observed in the phagocytosis of polystyrene microspheres by macrophages in vitro when they were coated with human or mouse CD47. The difference remained in the presence of varying levels of opsonization.65 Whether CD47 can be utilized for cancer nanomedicine remains mostly unknown as relatively few studies have been done in this area. 2113

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microvascular wall and the drug molecule or NP under consideration. Early studies for molecular drugs have reported that the MVP of tumor vasculature does not depend on the particle size as long as the particle is much smaller than the pore cutoff diameter of the vascular wall. Pore cutoff diameter is reported to be between 300 and 700 nm, while in rare occasions it can be up to 1.2 to 2 μm.80,81 Therefore, P values for small macromolecules such as BSA and IgG do not differ significantly from each other.80 However, as the molecular weight of the agent is further increased, P decreases by about 2 orders of magnitude.82 This is supported by earlier findings that the dominant mode of molecular transport across the microvascular wall is diffusion, which is a molecular weightdependent process, rather than convection.83 In contrast to molecular agents, direct measurement of P for NPs is not common in literature, possibly because the extravasation of NPs into tumor interstitium can be highly heterogeneous, making it difficult to determine a representative MVP value.9,82 Many studies in this area are, therefore, based on the measurement of NP accumulation at the tumor.6,12 In that case, the effects of the NP size on the MVP cannot be independently determined, since tumor accumulation depends on many other factors including plasma half-life. In a recent study, these two effects were separated by a linear regression analysis and, for a given plasma half-life, the tumor accumulation of gold NPs between 20 and 100 nm was found to be inversely related to the particle size.57 This finding implies that the MVP for NPs decreases with increasing particle size possibly in a similar trend to large macromolecules.82 Particle surface charge and functionalization also have significant effects on NP transvascular transport. Positively charged macromolecules, such as cationized BSA and IgG, have been shown to extravasate approximately two times faster than their negatively charged counterparts.8 While it is not yet clear whether this mechanism holds for NPs, cationic liposomes are known to preferentially bind to tumor vascular endothelial layer,10,84,85 resulting in improved extravasation. Furthermore, the charge of liposomes can be adjusted to optimize the accumulation of NPs at the tumor site.9 Shape is another factor that affects the interactions between the NP and vessel wall. In vitro experiments show that spherical asymmetry can promote the deposition of particles on vessel walls and may provide a mechanism for enhancing tumor accumulation of those particles. Disk shaped particles can have increased adhesion when compared to rods and spheres since they have larger surface area for adhesion and experience less drag from shear flow.86,87 The EPR effect has provided some means of passive targeting of tumor and has been used as a major design strategy for cancer nanomedicine. The improvement of drug accumulation at the tumor by NPs has been reported, but there is still room to improve as only about 1 to 10% of the administered drugs are delivered to the intended site.20,23,64 In addition to the EPR effect, other aspects of the tumor microenvironment should be considered. One approach is the use of active targeting in which the NP surface is functionalized by ligands such as antibodies and growth factors that can bind to specific receptors overexpressed by cancer cells or tumor-associated endothelial and stroma cells. Improved accumulation of NPs by active targeting has been reported in various tumor xenograft studies.4,78,88−90 The functionalization of block-copolymer micelles by endothelial growth factor could improve the tumor accumulation of the NPs.12 In addition, it was reported

New studies on protein−NP interactions draw attention to the distinction between synthetic and biological identity of NP.60,66 Synthetic identity here describes the “factory settings” of the NP, e.g., the size, shape, surface properties, and functionalization before introduction into the in vivo system. When a NP in the bloodstream attains a corona of native proteins by opsonization, it develops a biological identity that is different from its synthetic counterpart. Biological identity governs NP’s interaction with the cell and tissue components, and can also be effective on its transport in various ways. Apart from the known role of opsonization during NP clearance by RES, protein corona can also change the effective size and surface charge of the NP. The thickness of the protein corona has been reported to vary between 20 and 35 nm for metal and polymer NPs of 30 to 200 nm sizes.60 This means there is a significant increase in the NP diameter when exposed to in vivo environment. Incubation of NPs with serum results in a zeta potential of −10 to −20 mV irrespective of initial NP surface chemistry and size.67 This is thought to be due to the negative charge of most serum proteins in physiological pH.60 Transvascular Transport. NP transport across tumor vasculature is determined by the unique features of tumor vasculature that is composed of a highly disorganized network of blood vessels with functional defects.68,69 An incomplete formation of the endothelial cell layer and underlying basement membrane on the tumor vascular walls leads to intercellular gaps, called fenestrations. These defects are heterogeneously distributed across tumor vasculature.70 The presence of fenestrations in tumor vessels implies that tumor vasculature is leaky compared to that of normal tissues. Thus, the tumor vasculature is hyperpermeable to physiological fluids, macromolecules, and possibly NPs.71,72 The main reason behind the formation of this leaky, dysfunctional vasculature is thought to be the abnormally high levels of vascular endothelial growth factors (VEGFs) secreted by cancer cells.70,73 While lymphatic vessels at the periphery of the tumor are functional and can facilitate drainage of interstitial fluids,74 the function of lymphatic vessels at the tumor interior is impaired. The lack of functional lymphatics results in poor drainage of fluid from the tumor interstitium.75 The combination of leaky vasculature and poor lymphatic drainage leads to the well-known enhanced permeability and retention (EPR) effect. The EPR effect has been used to explain the preferential accumulation and prolonged presence of anticancer drugs in the tumor when compared to the rest of the body.76,77 Now, this concept is being extended to the design of various types of nanomedicine. The effectiveness of tumor targeting by the EPR effect, sometimes referred to as “passive targeting”, is significantly affected by the microvascular permeability (MVP) of tumor vasculature.78 MVP relates the rate of extravasation of particles to the concentration difference between the vascular and interstitial spaces. Transvascular transport in tumors is often mathematically described by the Kedem−Katchalsky formulation, a phenomenological law originally developed for diffusion of species across a semipermeable membrane79 in a simple form, Js = P(Cv − C i)

(1)

Here, P is the MVP, Js is the transvascular flux per unit area, and C is the concentration, with subscripts v and i standing for vascular and interstitial spaces respectively. Values of MVP for some macromolecules and NPs are reported in Table 2. MVP is a characteristic of both the tumor 2114

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and consumption of the drug by the cells, which further limits the penetration of the particles into the tumor.112 Quantification of Deff for NPs is rare in literature, yet NPs also are reported to have difficulties in penetrating into tumor interstitium.9,12,57 Intratumoral distribution of NPs is heterogeneous and positively correlates with tumor vasculature.4,9 This indicates poor distribution of NPs in the interstitial space. For PEGylated gold NPs whose final diameters are within the range of 20−100 nm, the greatest amount of penetration was observed for the smallest particles.57 Likewise, between blockcopolymer micelles with 25 and 60 nm sizes, 25 nm NP was found to diffuse farther away from the capillary than the 60 nm one.12 When comparing 30 and 70 nm polymeric micelles, no significant difference was found for murine adenocarcinoma,58 which is well vascularized with hyperpermeable vessels. However, 30 nm NP was observed to penetrate more than 70 nm in the case of pancreatic tumor, which has a vasculature with reduced microvascular permeability and a thick interstitial environment.58 These results suggest that the effect of NP size on interstitial transport depends on the tumor microenvironment.58 Surface charge-wise, neutral and negatively charged liposomes penetrate into tumor interstitium more effectively than their positively charged counterparts.9,10,62,85,113,114 To improve the interstitial penetration of nanomedicine, several approaches have been proposed. One of the recent approaches that has shown some clinical success115 is the normalization of the abnormal tumor vasculature.116 This approach aims to change the structure, organization, and transport properties of the tumor vasculature to those of the normal tissues using antiangiogenic factors which bind to VEGF receptors of tumor-associated endothelial cells.117 In particular, the decrease in the microvascular hydraulic conductivity has the net effect of lowering the IFP. This in return causes the development of a net transvascular pressure gradient that improves the extravasation of therapeutics.101 Furthermore, depending on the normalization time window, IFP gradients can develop within the tumor (rather than only the periphery) and can lead to enhanced particle penetration into tumor by convection.117 However, a recent study shows that the improvement of tumor penetration by this approach might be limited to small NPs only (e.g., 12 nm), as the blood vessels, having reduced pore size, can prevent the extravasation of larger NPs (125 nm in the study). Other remedies to improve the penetration of NPs into tumor mainly focus on mitigating the interstitial diffusion barriers caused by the dense tumor interstitium. A technique that is coming forward is tumor priming, in which the cancer cell density is decreased by apoptosis inducing drugs prior to administration of NPs.118,119 This is illustrated by a study where systemic administration of paclitaxel increased the interstitial space and blood perfusion into the tumor xenograft and the following delivery of 100 and 200 nm latex NPs (but not 500 nm) and 85 nm liposomes was improved.120 Yet another effort to improve interstitial transport is the degradation of ECM components by digestive enzymes.106−108,121,122 The use of collagenase to reduce the collagen matrix in subcutaneous tumors has been shown to increase the effective diffusivity of 500 kDa macromolecules up to 10 times.122 Cellular Uptake. The cellular uptake of NPs is an essential step especially in therapeutic applications since most anticancer drugs need to reach a specific intracellular target, including mitochondria, cytosol, and nucleus, in order to elicit a cytotoxic effect. While free drugs can be transported into the cell by

that iRGD, which is one of the tumor-homing CendR (C-EndRule) peptides,89 could be a new targeting ligand. iRGD targets the tumor-associated endothelial cells and also results in an increase of vascular and interstitial permeabilities within the tumor. iRGD has been shown to significantly improve the penetration of both conjugated and coadministered drugs into the tumor.88,89 However, there are also reports of targeting modalities that did not work as expected, and the use of some tumor-specific markers resulted in even lower tumor accumulation than untargeted controls.91 One of the reasons for the poor performance is suggested to be the binding-site barrier.92 This denotes the lack of mobility and tumor penetration of antibody conjugated NPs since they directly bind to their intended target upon extravasation. The unexpected outcomes can also be considered in the context of “targeting dilemma”93 stating that selection between passive and active targeting for NP design requires a compromise since the functionalization of NPs with targeting ligands usually conflicts with PEGylation necessary for prolonged circulation. Detailed review of current status in active targeting modalities is provided elsewhere.94,95 Interstitial Transport. Tumor tissue architecture is marked by a tightly packed compartment of cancer and stromal cells and an interstitium that features relatively stiff ECM with high collagen and proteoglycan content.50 Furthermore, due to leaky vasculature and poor lymphatic drainage, tumors are associated with elevated levels of interstitial fluid pressure (IFP). IFP in many types of tumors falls into the range of 20 to 50 mmHg,96,97 but much higher values, 75−130 mmHg, are reported in pancreatic cancer.98 These are dramatically higher than IFP in normal tissues, e.g., 0 to 0.5 mmHg in skin and breast.96 IFP is relatively uniform within the tumor. Therefore, the main mode of interstitial transport is thought to be diffusion, rather than convection. However, at the tumor periphery, IFP rapidly decreases to the normal tissue level and steep pressure gradient presents. This pressure gradient induces a net outward flow, by which particles accumulated near the periphery are carried away from the tumor.99−101 Tumor IFP is found to be comparable to the microvascular pressure,102 so that the transvascular pressure gradient is small. Thus, the extravasation of NPs is also expected to occur mainly through diffusion.83,103 At the interior of tumor, interstitial transport is modeled reasonably well by diffusion processes as follows:104,105 ∂C = Deff ∇2 C ∂t

(2)

where the right-hand side represents the time rate of change of interstitial concentration, C, and the left-hand side is the spatial variation of C. The proportionality constant, Deff, is the effective diffusivity. Deff is a measure of mobility of particles, and small values of Deff indicate high hindrance of diffusion in interstitial space. Deff values for several macromolecules and NPs are reported in Table 2. The dense tumor ECM is the leading barrier for the penetration of particles into tumor interstitium.106−108 The effective diffusivity of a particle in tumor can be more than 2 orders of magnitude lower than its counterpart in solution.7,106,109,110 The interstitial concentrations of even small particles such as free drugs decrease exponentially away from the blood vessels and might not reach the hypoxic regions occupied by cells resistant to drugs.111 The concentration gradient that drives the diffusion is also reduced by the binding 2115

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transported to lymph nodes. Considering the fact that lymph nodes are typical starting points of tumor spread and metastasis, understanding of the lymphatic transport is highly desired.136 Typically the lymph nodes have been targeted via local injection route, but targeted drug delivery to the lymph nodes is still under development. Although it is thought that the lymphatic transport has a different allowable size window of NPs, compared to the transvascular transport, the characteristics of the lymphatic transport are not well studied. Heat Transfer around NPs. In addition to delivering anticancer drugs, NPs can also be used to locally heat the tumor while sparing surrounding healthy tissues. As the use of NPmediated hyperthermia emerges as a new cancer therapy option, it becomes increasingly important to consider not only the transport of NPs to tumor site but also intratumoral transport of heat. This heat transfer problem is concerned with the spatial and temporal variation of temperature, and requires the analysis of heat transfer at multiple length scales, from tissue level (macroscale) to NP level (nanoscale). Tissue level heat transfer has been modeled by the Pennes bioheat equation.137 Once the temperature field within the tumor is known, it can be related to cancer cell survival based on hyperthermic injury mechanisms.138,139 Nanoscale energy conservation, on the other hand, is an active area of research that is primarily concerned with the energy conversion mechanisms of NPs and their thermal interactions with the local environment.140 The bioheat equation represents the conservation of energy at tissue level as a partial differential equation in the following form:

diffusion or endocytosis, for NPs, which are relatively larger, the rate of transmembrane transport is primarily determined by the endocytosis.15 Once inside the cell, the therapeutic agent needs to be released from the NP and escape the endosome before it merges with a lysosome for digestion of its contents.23 Studies with gold NPs between 10 and 100 nm consistently report that optimal particle size for cellular uptake by receptor mediated endocytosis is around 50 nm.15,17,123 However, future research is necessary as those in vitro studies were conducted on immortalized cell lines and the optimal particle size for cellular uptake will likely depend on the cell type and culture conditions.124 At this size range, spherical particles perform better than rods. However, for particles larger than 100 nm, rods show better uptake than spheres.125,126 In fact, even the orientation of the particle at the time of attachment to the cell surface can determine the cellular uptake outcome. Whether macrophages phagocytize rod NPs is thought to be determined at least partly by the local geometry of the particle that is felt by the cell.127 Once internalized, localization of NPs in cell might also depend on their size and shape. When endothelial cells were exposed to spherical and rod microparticles with size range between 1 and 8.75 μm in pairs, the regions closest to the nucleus were occupied by, in order of preference, large spherical, small spherical, large rod, and small rod particles.128 Whether such patterns of segregation exist for smaller NPs still needs to be investigated. As for the surface properties, cationic particles are generally associated with higher uptake rates,61,129 and there is evidence that PEGylation reduces cellular uptake of gold nanoparticles17 and liposomes.61 One of the main barriers against the intracellular accumulation of free chemotherapy drugs is multidrug resistance (MDR), a characteristic developed by some cancer cell lines that greatly diminishes the success of chemotherapy.130,131 MDR is primarily mediated by a family of membrane transport proteins, ABC (ATP binding cassette) proteins,132 that pump the internalized chemotherapy drugs out of the cell. Thus, the relative rates of cellular uptake of NP carrier and MDR mediated efflux determine the intracellular concentration of therapeutic agents. It has been extensively studied to modulate these membrane transport mechanisms. Some research has focused on the alteration of membrane biophysical properties, including membrane fluidity and permeability to enhance the cellular uptake of drugs.133,134 Another approach has been the coadministration of the anticancer drug with MDR suppressants (or chemosensitizers) such as verapamil or tariquidar to inhibit the MDR processes.25,26 NPs are considered to be suitable platforms for this purpose since they can be used to carry and release both agents simultaneously and can achieve high levels of colocalization inside the cell. Improved tumor-specific targeting of NPs is also a desirable feature since ABC transporter proteins have vital functions in normal physiology of various tissues as well (e.g., blood−brain barrier, kidneys, and intestines) and the suppression of the drug efflux mechanisms in normal cells can lead to loss of that functionality.131 Recently, various NP-based multidrug delivery systems have been proposed to overcome MDR and have shown promising results.5,25,26,135 However, their selective delivery to the target tissue still remains to be a primary challenge.

ρc

∂T = ∇·(keff ∇T ) + ρb cbω(Tb − T ) + qṁ‴ + qṡ ‴ ∂t

(3)

where T is the temperature across the tissue, ρ and c are the density and specific heat representative of the tissue respectively, and keff is the effective thermal conductivity that mainly accounts for the heat diffusion. The second term on the right-hand side represents the effect of heat supplied by blood perfusion to the tissue and is commonly modeled by the approach of Pennes.141 Here, ω is the blood perfusion rate per unit volume of tissue, ρb and cb are the density and specific heat of blood, and Tb is the temperature of the blood in the artery entering the tissue. q̇m ‴ is the metabolic heat generation rate. Finally, q̇‴ s represents heat generated by the specific mode of thermal therapy and depends on the nanoscale energy conversion mechanism and intratumoral distribution of NPs. Numerous applications of bioheat equation to model NPmediated hyperthermia treatment are available elsewhere.142−144 Energy generation mechanisms of NPs depend on the nature of external excitation. In photothermal conversion, heat is generated by resonant oscillation of electrons in NP when excited by an electromagnetic irradiation of specific wavelength, referred to as local surface plasmon resonance.145 In the case of magnetic heating of superparamagnetic NPs, the main mechanism of heat generation is the relaxation losses that occur as a particle’s internal magnetic field is aligned with the external alternating magnetic field or the viscous dissipation that occurs as the NP rotates with respect to the surrounding fluid under the action of the external magnetic field.146 The origin of radiofrequency heating is the eddy currents that are induced in conductive NPs by the external electric field.147 Regardless of the mechanism of heating, energy conversion is

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TRANSPORT PROCESSES AFTER THE DELIVERY After delivery to the tumor, excess NPs and the interstitial fluid are drained to the lymphatic vessels, and are eventually 2116

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tissue components, NPs have also been used to improve the outcomes of cryosurgery, a treatment modality based on destruction of cancerous tissue by freezing. In NP-mediated cryosurgery, NPs can improve the therapeutic efficiency by preconditioning tumor cells to undergo necrosis at elevated freezing temperatures152,153 or enhancing the thermal conductivity of the tissue.154,155 We refer the interested reader to a review on this matter.39 Light Transport around Photonic NPs. If photonic NPs are delivered to target tissue, light from/to these NPs interacts with the surrounding tissues. The interaction of light with tissue is generally described by the radiative transfer equation (RTE)156 as given below. The equation describes spatiotemporal variation of light intensity across the tissue. μ + μs s ⃗ · ∇I = − (μa + μs )I + a ∫4π p(s ⃗ , s ⃗′) I( r ⃗ , s ⃗′, t ) dΩ′ 4π

almost instantaneous, e.g., occurs within on the order of femtoto picoseconds.148 Some part of the generated heat is dissipated to the surroundings, mostly by diffusion, while the rest accumulates in the NP, increasing its temperature. The time scale of the heat diffusion in the vicinity of NP is represented by the relaxation time,149 τrel, which is given by τrel =

d2 27α

(4)

where d is the characteristic length of the system considered which, in nanoscale, is the size of the NP and, in macroscale, the size of the tissue of interest. α is the thermal diffusivity of the system. τrel represents the time required to dissipate a considerable amount of heat accumulated in the particle by diffusion. If the external excitation of the NP lasts for a duration, t, that is shorter than the relaxation time (t < τrel) then all the heat will accumulate in the particle, i.e., without any appreciable diffusion to surroundings. Such a state is marked by the condition of “thermal confinement”149 and stands for the case of intense temperature increase in the vicinity of NP only. Thermal confinement effect can be useful in nanoscale targeted photothermal therapy where small macromolecules or intracellular organelles are selectively exposed to NP-mediated heating by use of a short-pulse laser irradiation. Recently, intense localized heating of gold NPs in endosome has been used to rupture colocalized doxorubicin loaded liposomes and facilitate endosomal escape of the drug.135 If the excitation duration is longer than the thermal relaxation time, bulk heating becomes the main mode of heat delivery. In that case, while the temperature difference between the NP and its surroundings remains small, the temperature of the whole tissue can be elevated significantly.150 For a case of photothermal therapy,36 the temperature difference between the NP and its surroundings was estimated to be only 0.003 °C, which is hardly enough to induce any sort of thermal damage in the NP’s local environment. The whole tumor tissue, on the other hand, could be heated up to a temperature 31 °C greater than its surroundings.150 Altogether, depending on the relative magnitudes of the heat diffusion and accumulation, the NP heating effect can be either bulk, i.e., sensed throughout the tissue containing heated NPs, or localized, i.e., limited to the vicinity of the NP only. The outcome depends on not only the thermal properties of the NP and the tissue but also the type, duration, and intensity of the external excitation.140,146 Hyperthermic injury of the tissue has been characterized by the thermal damage function, Ω,151 which is related to the duration and temperature of the exposure by an Arrhenius relationship,138,139 dΩ = A exp( −Ea /RT ) dt

(7)

In the RTE, I is directional spectral intensity, s ⃗ is the unit direction vector, μa and μs are absorption and scattering coefficients respectively, and p(s,⃗ s′⃗ ) is the scattering phase function, which takes care of multiple scattering. Other probability based methods (Monte Carlo and molecular dynamics simulations) are also used for predicting the photons’ interaction with tissue. Since human tissue is a turbid media (i.e., scattering dominant), the RTE can be reduced to the diffusion approximation as given in eq 8,157 −∇·(D∇Φ) + μa Φ = 0

(8)

where Ω is photon density (W/mm2), g is the anisotropy parameter, and D = 1/[3[μa + (1 − g)μs]]. The optical parameters, μa, μs, and g, are very much dependent on tissue type and frequency of light used. Values for those parameters in different tissues are reported in the literature.157,158 Figure 2 shows an example of light−tissue interaction during quantum dot mediated fluorescence thermometry.159−161 This thermom-

(5)

where A is the frequency (amplitude) factor and Ea is the activation energy for thermal damage inducing processes. The fraction of undamaged tissue is

Fs = exp( −Ω)

(6)

which can be used to quantify the extent of thermal damage. Thermal damage parameters, A and Ea, for different types of tissues are available in literature.139 While the focus here has been on approaches that utilize elevated temperature to cause injury to tumor cells and other

Figure 2. An example of quantum dot mediated fluorescence thermometry. The quantum dot (QD) fluorescence intensity decreases with increasing temperature and gets diffuse with increasing tissue thickness. Adapted from Ghosh et al.159 2117

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Figure 3. (A) Illustration of the problem setup in case study (1). (B) Spatiotemporal GNS concentration across tumor and normal tissues. (C) Variation of mean concentration of GNS accumulated in tumor and normal tissues.

factors that promote the removal of NPs after delivery, can cause the NP accumulation in tumor to be highly spatiotemporal, i.e., NP amounts that are greatly changing by time and location in the tumor. In order to better illustrate this point, we present below two simple case studies based on transport modeling. The first case study is for NP transport on the whole-tumor level, while the second case study considers the same problem at a smaller length scale, at the proximity of a tumor capillary. Spatiotemporal Accumulation NPs in Tumor. The work presented here is based on a previous experimental work that involved the delivery of gold nanoshell (GNS) particles (110 nm diameter) to mammary carcinoma in mouse and application of laser-induced photothermal therapy.36 The problem setup and major aspects of the transport model are illustrated in Figure 3A. The study employed an embedded tumor model where the tumor was surrounded by the normal tissue. GNS transport to the tumor was then modeled by considering the species conservation law:105

etry technique has been developed to intraoperatively visualize the local temperature of a thermal lesion using temperaturedependent fluorescence intensity of quantum dots. One of the critical bottlenecks to extend this technique for in vivo imaging is the development of methods that enable proper analysis of the fluorescence signal based on light−tissue interaction. To get a predictive nature of therapy and diagnostics it is important to understand the optical parameters of the nanoparticle embedded biological system. The other inherent complications of an in vivo system, e.g., anisotropy of tissues,162 blood flow, and multiple compositional parts of a given tissue, should also be accounted.

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CHALLENGES AND OPPORTUNITIES Recent development of nanotechnology and understanding of physiological barriers of NP transport has made a wide variety of NPs available. These NPs enable numerous innovative NPmediated therapies, and their promising outcomes have been reported. However, their success in the laboratory setting is not well translated into the clinical setting. This significantly impaired outcome in vivo is primarily due to the limited understanding and lack of predictive measure of their in vivo behavior. Complex and Multiscale Transport Barriers. The transport processes taking place at different stages of NP delivery are highly complex and interconnected. The interactions of transport barriers, like plasma clearance, resistances against transvascular and interstitial transport, and

2 ∂Ci 1 ∂(r vr, ifi Ci) + 2 ∂r ∂t r ⎛ C ⎞ 1 ∂ ⎛ ∂C ⎞ = Deff, i 2 ⎜r 2 i ⎟ + Pi(S /V )i ⎜Cv − i ⎟ KAV ⎠ r ∂r ⎝ ∂r ⎠ ⎝

(9)

The first three terms in eq 9 represent the time rate of change of concentration and convective and diffusive interstitial 2118

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Figure 4. (A) Illustration of the problem setup in case study (2). (B) Spatiotemporal nanoparticle concentration near tumor capillary vessel for three different particle sizes.

pressure in the tumor induce a net interstitial flow from the center to the periphery of the tumor.99 As a result, the GNS starts to accumulate in the peripheral regions. Due to the interplay of plasma clearance, interstitial diffusion, and convection, the mean concentration of GNS in both tumor and normal tissues first rapidly increases, then decreases. Besides the size window by the EPR effect, this simple model illustrates that significant changes can occur in the accumulation amount over short periods of time, leaving only a narrow optimal time window for NP-mediated therapy and imaging applications. Similar time-dependent trends have been observed for model polymer drugs,163 targeted and untargeted liposomes,78 gold NPs,57,164 and GNS under similar settings.165 In the last case, the peak GNS accumulation occurred at 24 h postinjection (among measurements at 1, 4, 24, and 48 h). The difference between the experiments and the simulations might be due to the lack of proper knowledge of transport properties of GNS, i.e., microvascular permeability and effective diffusivity for the GNS in tumor tissue as well as other properties associated with tumor microenvironment like hydraulic conductivity, and vascularization level. Some other limitations of the model are listed in the following section. Size Dependent Dynamics of NP Transport near Tumor Microvasculature. In this study, the extravasation and interstitial transport of NPs with three different diameters (d = 2, 38, and 100 nm) are modeled in the proximity of a tumor capillary vessel with radius Rc = 5 μm. The selected sizes are based on the hydrodynamic diameters of nanomedicine commonly studied. In particular, 2 nm is for free drugs,110 and 38 nm is for large macromolecules,7 which can also

transport of GNS respectively. The last term on the right-hand side is a source term that accounts for the gradual accumulation of GNS in tumor tissue by transvascular transport. Here, r and t are the spatial and temporal coordinates, Ci is the GNS concentration in tissue domain i (either tumor (T) or normal (N) tissue), Cv is the concentration of GNS in the bloodstream, vr is the interstitial fluid velocity, f is the retardation coefficient, Deff is the effective diffusivity of GNS in the interstitium, P is the microvascular permeability to GNS, S/V is the surface-tovolume ratio for transvascular transport, and KAV is the available volume fraction of the tissue. In developing eq 9, the tumor was assumed to be a sphere with radius RT and the GNS concentration was assumed to vary in the radial direction only. The binding and cellular uptake of GNS were neglected for simplicity. Furthermore, both tumor and normal tissue vasculatures were assumed to be uniformly distributed. The problem was solved numerically using finite difference method. Further details of the model are provided in the Supporting Information. The spatial and temporal variations of GNS concentration are shown in Figure 3B and Figure 3C. It is seen that the GNS concentration varies significantly over time in both tissues. During the first 4 h of injection, the GNS particles gradually accumulate in the tumor. The amount of accumulation is greater than for normal tissue due to higher transvascular permeability of leaky tumor vasculature. This is an example of passive targeting achieved through the EPR effect.77 However, this accumulation is followed by a decay of bloodstream concentration and rapid wash out of GNS from the tumor tissue. Furthermore, the elevated levels of interstitial fluid 2119

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represent small NPs such as functionalized quantum dots.166 The size of 100 nm is well within the range for gold NPs167 and liposomes61,62 and can be considered as a nominal size for large NPs. An illustration of the problem is shown in Figure 4A. A Krogh cylinder approach168,169 was adopted for this study by segmenting the tumor microvasculature into long cylindrical domains each having a tumor micro capillary vessel at its center. The governing equation for the problem is a simplified form of species transport equation based on the assumptions stated in the previous section. As a result, the interstitial transport is governed by the diffusion equation: ∂C 1 ∂ ⎛⎜ ∂C ⎞⎟ = Deff r ∂t r ∂r ⎝ ∂r ⎠

for

Rc < r < Ro

tumor interstitium throughout the 30 min duration of experiments. At the end, larger particles were found to have penetrated more into the tumor interstitium.82 Reversed diffusion into the bloodstream can be one of the ways NPs are removed from tumor after delivery. At locations near the periphery of the tumor, NPs may also be removed from tumor by IFP driven outflow and/or drainage to functional lymphatics. It is important to note that the simple models presented here by no means fully address the complex NP transport problem. Most notably, being one-dimensional, they do not capture the actual tumor geometry and spatial heterogeneity like the uneven distribution of vasculature and ECM components. These models do not take into account the variation in the vessel pore size and its effect on transvascular permeability. The depletion of mobile NPs in interstitium by cellular uptake or binding site barrier are not considered either. For more advanced modeling of transport in tumors, the reader is referred to other works on this topic.114,169−174 Need for New Evaluation Strategies. As described so far, the interactions of different transport barriers result in dynamic accumulation of NPs in both space and time. Unless these dynamics are properly taken into account, they can confound the evaluation of NP performance. The main platforms used for preclinical testing of cancer nanomedicine primarily are in vitro systems that involve static culture of cancer cells and tissues, and animal xenograft models realized by the implantation and/ or growth of human tumors in immunodeficient mouse.175 In vitro systems include cell monolayers on substrate, threedimensional spheroids, and engineered tissue scaffolds.176−178 While they are convenient and can enable high-throughput screening, they are also dramatically limited in capturing the NP in vivo behavior since they fail to mimic the complex tumor microenvironment. These systems lack the barriers that would normally be encountered by NP in vivo such as plasma clearance and interstitial fluid pressure and flow. Many of the key features of the tumor microenvironment missing in the in vitro systems are present in animal models, making them a valuable tool for characterizing NP in vivo behavior. Yet, animal models also do not offer an easy way to control tumor microenvironmental parameters. Therefore, the information for tumor accumulation of NPs is currently available only at a caseby-case basis, and shows significant variation across different research groups and experimental settings. This variation makes it difficult to translate the outcomes of studies with these systems into clinical use with human patients. In addition, despite the recent advances in in vivo imaging,179,180 many features of tumor environment remain inaccessible during the experiments. Therefore, the characterization in these systems are commonly done at a whole organ basis, and at select time points.4,52,167,181 As a result, animal models are limited to provide a proper characterization of spatiotemporal aspects of NP transport. There appears to be a great need for flexible, accessible, and high-throughput techniques for the evaluation of cancer nanomedicine. The recent emergence of in vitro tumor models based on microfluidics145,182−186 can help to close the gap between the current static culture systems and animal models by enabling systematic and independent control of tumor microenvironmental parameters while maintaining a realistic representation of tumor environment such as the vascular and interstitial compartments and the presence of interstitial fluid pressure driven flow. These models offer a promising avenue of research

(10)

And the transvascular transport is included in the model as a boundary condition at r = Rc, ⎛ ∂C C ⎞ = P ⎜C v − ⎟ at ∂r KAV ⎠ ⎝ ∂C − Deff = 0 at r = R o ∂r

−Deff

r = Rc ,

and

(11)

The first condition stands for the transvascular flux driven by the concentration gradient across the vessel wall. The second condition represents the lack of mass flux due to symmetry across cylindrical compartments. Further details of model parameters are provided in the Supporting Information. The spatial and temporal evolution of nanoparticle concentration for different particle sizes is shown in Figure 4B. It is immediately noticed that as the particle size increases, both the accumulation and penetration into the tissue decreases. This is because diffusion both across vascular wall and into the interstitium slows down with increasing particle size. However, as NPs are cleared out of the bloodstream, the concentration gradient driving the diffusion into the interstitium is reversed and particles move back into the bloodstream. Smaller particles are removed from tumor interstitium faster than the larger particles as they diffuse through interstitium, and move between vascular and interstitial compartments more easily. After 24 h all cases show similar concentration profiles, with smallest particle concentration being somewhat higher than the others. This example illustrates that NPs have the tendency to accumulate near tumor vasculature rather than getting uniformly distributed in the interstitium. The extent of tumor penetration can be affected by NP characteristics, like size; however, it is also subject to change by time as the NPs are cleared from the bloodstream. Therefore, the size effect might be significant at a certain time after injection and might not be at a later time point. Similar size dependent dynamics have been observed previously in experiments with dextran as a model macromolecular drug.82 Comparisons were done on different molecular weights ranging from 3.3 kDa to 2 MDa, that corresponded to hydrodynamic diameters of 1 to 40 nm. The smallest 3.3 kDa particles accumulated in tumor interstitium rapidly, and achieved the highest amounts of accumulation. Nevertheless, after 5 to 10 min, a majority of particles got cleared from the bloodstream and the concentration at tumor interstitium also decreased rapidly. This was in contrary to larger 40 and 70 kDa particles that remained in the bloodstream longer and continued to accumulate in the 2120

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CONCLUDING REMARKS The interactions of transport barriers often result in highly dynamic accumulation and distribution of NPs in tumor. The dependence of these dynamics on NP transport characteristics like size, shape, and surface properties have been investigated extensively, but it remains as a challenging question to answer whether these characteristics can be tuned so that NP-mediated strategies overcome the transport barriers effectively. This challenge cannot be overcome by just synthesis of new types of NPs. Since NPs are exposed to a complex and dynamic in vivo environment where each stage of transport requires different, sometimes contradicting, properties of NPs, changes are required in design strategies. New approaches should consider the multifaceted interactions of NPs with cells, ECM, and fluid. Novel NP design modalities that are being developed to fulfill this need include active targeting and environment/stimuliresponsiveness. These approaches have brought new interactions of NPs with the biological system that need to be quantitatively understood in terms of transport characteristics reviewed here. Changes are also needed in evaluation criteria for successful delivery. The new evaluation techniques need to take into account the dynamic environment of tumor, yet allow for rapid characterization and tunable tumor environmental parameters for screening. Moreover, a knowledge-base on how to interpret the results of in vitro assays to predict NP in vivo behaviors in animals, not to mention in humans, is necessary.

for understanding the dynamics of NP transport and evaluating the efficiency of newly developed NP formulations. At the end, a proper understanding of the dynamic nature of NP transport can be one of the ways to improve the evaluation of NP systems currently used in cancer diagnosis and treatment. Environment and Stimuli Responsiveness. Multiple properties of NPs affect different stages of NP delivery in different ways. An NP with certain attributes, while being transported efficiently at one stage of delivery, can show significantly poor performance in another. While vascular targeting and the rate of cellular uptake favor positive surface charge,10,61,82,84,85 neutral and sometimes negatively charged particles are likely to penetrate into tumor interstitium better.10,85,113 Similarly, PEGylation, while greatly improving the plasma half-life, also results in a reduced cellular uptake.17,61 Large NPs (∼100 nm) have rather prolonged blood circulation, and they successfully accumulate at tumor vasculature through the EPR effect;124 however, their large size is also the main reason for their poor diffusion into the interstitium.57,187 Apparently, in NP design, one size does not fit to overcome all transport barriers, and the design of effective NP-mediated cancer biotechnologies requires integration of those multiple design requirements together. This significantly complicates the problem of optimizing NP designs for delivery. In order to address this problem, a new generation of NPs is being proposed, whose configuration changes in response to tumor environmental cues and/or external stimuli.188−190 One of the recent works in this line of thought proposes a multistage delivery system composed of 100 nm degradable gelatin NPs embedded with 10 nm quantum dots. The agent, thanks to its large size, remains in the bloodstream for extended periods of time, enabling preferential accumulation of the agent at tumor site through the EPR effect. Enhanced interstitial transport is achieved by disassociation of the gelatin structure upon encounter with matrix metalloproteinases to release quantum dots into the tumor interstitium.187 Another example of this design approach is polystyrene NPs conjugated with collagenase on their surface. This formulation achieved enhanced penetration into multicellular spheroids in vitro when compared to unconjugated controls.108 Other applications include the delivery of therapeutic payload by NPs that are sensitive to tumor-specific microenvironment such as low pH.5 External stimuli like light,192 heat,163,193 and pressure waves generated by ultrasound194 have also been studied to trigger drug delivery or mediate targeting for cancer. Guidance of magnetic nanoparticles to tumor site under the action of a localized external magnetic field have also shown promising results.195 An example of environment responsiveness is the shape-shifting polymeric NPs that can change between spherical and needle-like (elongated) configurations in response to elevated heat or decreased pH. Those particles can be effective in accumulating in tumor cells by initially having the elongated configuration to escape the clearance by RES and then change to the spherical configuration for faster cellular uptake.196 While these examples provide demonstrations of promising ideas, the design of environment and stimuli responsive NPs that function in vivo will also need to be based on the knowledge of spatiotemporal dynamics of tumor microenvironment. A recent detailed review of environmentresponsive NPs is provided elsewhere.190

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ASSOCIATED CONTENT

S Supporting Information *

Additional description of transport models used in case studies. This material is available free of charge via the Internet at http://pubs.acs.org.

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AUTHOR INFORMATION

Corresponding Author

*585 Purdue Mall, West Lafayette, IN 47906, United States. Phone: +1-765-494-5626. Fax: +1-765-496-7535. E-mail: [email protected]. Notes

The authors declare no competing financial interest.

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ACKNOWLEDGMENTS This work is partially supported by grants from the National Science Foundation (CBET-1009465) and Purdue Research Foundation.

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REFERENCES

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Review

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dx.doi.org/10.1021/mp3005947 | Mol. Pharmaceutics 2013, 10, 2111−2126