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. Author manuscript; available in PMC: 2009 Dec 18.
Published in final edited form as: J Control Release. 2008 May 10;132(3):164–170. doi: 10.1016/j.jconrel.2008.05.003

Recent progress in tumor pH targeting nanotechnology

Eun Seong Lee 1,2, You Han Bae 1,*
PMCID: PMC2695946  NIHMSID: NIHMS56768  PMID: 18571265

Abstract

pH-sensitive polymeric micelles and nanogels have recently been developed to target slightly acidic extracellular pH environment of solid tumors. The pH targeting approach is regarded as a more general strategy than conventional specific tumor cell surface targeting approaches, because the acidic tumor microclimate is most common in solid tumors. When nanosystems are combined with triggered release mechanisms by endosomal or lysosomal acidity plus endosomolytic capability, the nanocarriers demonstrated to overcome multidrug resistance of various tumors. This review highlights recent progress of the pH-sensitive nanotechnology developed in Bae research group.

Keywords: Tumor extracellular pH, endosomal pH, pH double targeting, micelle, nanogel, block copolymers, sulfonamide, histidine

1. Introduction

Amongst various cancer therapies, chemotherapy is one of major treatment modalities along with debulking surgery [1-2]. Major challenges in chemotherapy are linked to toxicity on healthy proliferating cells and multidrug resistance (MDR) against anticancer drugs. The life threatening side effects caused by non-specific tissue distribution of the drugs has restricted the systemic high-dose strategy [3, 4]. Cancer cells except those having intrinsic resistance are sensitive to chemotherapy in the beginning but often develop acquired resistance upon repeated chemotherapy cycles [5-7]. The resistance initiated by a cytotoxic agent also extends cross-resistance to a wide range of drugs having different chemical structures and cellular targets [5-7]. Once the resistance appears, systemic high dose therapy becomes ineffective and more toxic and the resistance is further stimulated [5-7].

Tumor targeting approaches have been developed for improved efficacy and reduced toxicity by altering biodistribution of cancer drugs and by using specific cell surface interactions. Solid tumors are often characterized by overexpression of specific antigens or receptors on cell surfaces [8-13]. Antigens and receptors help in transmitting signals from the surrounding environment that are essential for the growth of tumor cells. Targeting antigens or receptors has been extensively investigated as an important delivery mode by using macromolecular or nanosized carriers to tumor cells. Nanodrug carriers attached with surface ligands or antibodies exploit these receptor-mediated uptake pathways that are recognized and internalized by the tumor cells [14-18]. However, these approaches have achieved limited success in clinic, most likely because of significant heterogeneity in both solid tumor cell types and cell surface markers [8-10]. Additionally, both the presence of antigens and the expression of receptors on surface of these tumor cells are transient and dynamic [8-13]. The heterogeneity of cancer cells may explain the reasons for the unexpected results of targeting strategy [19].

The extracellular pH (pHe) of normal tissues and blood pH are kept constant at pH 7.4 and their intracellular pH (pHi) at 7.2. However, in most tumors the pH gradient is reversed (pHi > pHe). Particularly, tumor pHe is lower than normal tissues [20-23]. Although there is a distribution in in vivo, pHe measurements made by using needle type microelectrodes on human patients having various solid tumors (adenocarcinoma, squamous cell carcinoma, soft tissue sarcoma, and malignant melanoma) and in readily accessible areas (limbs, neck, or chest wall), shows the mean pH value to be 7.0 with a range between 5.7−7.8 [22]. This variation is dependent upon tumor histology, tumor volume, and location inside a tumor. Recent measurements of pHe by noninvasive technology such as 19F, 31P, or 1H probes by magnetic resonance spectroscopy in human tumor xenografts and in animals further proved consistently low pHe [23, 24]. Reported pHe data on human and animal solid tumors either by invasive or noninvasive methods showed that more than 80% of all measured values are below pH 7.2 [23, 24]. The primary reason for this imbalance in cancer pH is the high rate of glycolysis in cancer cells, both in aerobic and anaerobic conditions [25-27]. It is also proposed that the acidic milieu benefits the cancer cells by generating an invasive environment that tears down the extracellular matrix and destroys the surrounding normal tissue cells [28].

There are a variety of mechanisms associated with MDR cells that need to be circumvented for a successful tumor treatment [29-41]. At unicellular level, ATP dependent drug-efflux pumps of P-glycoprotein (Pgp), multidrug resistance protein (MRP), lung resistance protein (LRP), antiapototic (or survival) bcl-2 gene, and altered expression of Topoisomerase II interfere with a sufficient intracellular drug dose and decrease the effectiveness of drug in killing tumor cell [29-36]. In clinical setting, additional tumor microenvironmental factors such as epidermal growth factor, fibroblast growth factor, insulin-like growth factor, and extracellular matrix components are strongly associated with survival mechanisms of cancer cells under cytotoxic drug treatment [37-40].

To tackle the multifaceted MDR mechanisms, it is hypothesized that a focal high dose strategy works at cellular level rather than at systemic level. In addition, local high dose may overwhelm most resistant mechanisms, which might have their intrinsic limitation in defense capability because even extremely resistant experimental MDR cells are killed at high drug concentrations [42-46]. Intracellular organelles in parental drug-sensitive cells are characterized to have somewhat acidic, diffuse pH profiles inside cells [47, 48]. MDR cancer cells develop more acidic organelles (recycling endosome and lysosome) than those in sensitive cells, which are more acidic than cytosolic pH and nucleoplasmic pH [49, 50]. This results in acid-induced sequestration of anticancer drugs. Acidic organelles in MDR cells contribute to developing resistance to chemotherapeutic drugs [51]. Since most anticancer drugs are in an ionizable form, the pH of extracellular matrix and intracellular compartments are critical factors in determining drug partitioning and distribution [52]. The low pH in tumor extracellular space or in various subcellular organelles is a significant signal for targeting.

This review highlights recent progress of the pH-sensitive tumor-targeting nano-carriers developed in Bae research group. For a more general literature overview, interested readers are referred to a most recent review by Oh et al. [53].

2. Tumor extracellular pH (pHe) Targeting

The approaches to target various solid tumors by pHe include micelle systems with a triggered drug release mechanism, and exposing nonspecific cationic TAT (HIV transactivator of transcription) peptide by a shielding/deshielding mechanism or by a pop-up mechanism. These systems have utilized the pH-sensitivity of poly(L-histidine) or polysulfonamide. The imidazole ring of a polyHis (pKb ∼ 7.0; polyHis is the most effective pH-buffering agent in a physiological system) has lone pairs of electrons on the unsaturated nitrogen that endow pH-dependent amphoteric properties [43-46]. Particular polysulfonamides (pKa ∼ 6.8) are negatively charged at blood pH (i.e., pH 7.4) and can be neutralized at acidic pH (e.g., tumor pHe) [41]. In addition, these polymers demonstrated a strong endosomolytic property by proton sponge effect and/or interactions with the anionic phospholipids of endosome [41, 43-46]. These properties of polyHis or polysulfonamide have contributed to the development of smart nanosystems designed for targeting tumor pH.

2.1. Triggered tumor pHe drug release

It has been proposed that pH-induced anticancer drug release from pH-sensitive liposomes, which are stable at neutral pH but leaky under mild acidic condition (pH 4.5−6.0), could be a new mode for tumor treatment. However, due to the lack of response to tumor acidity (pH 6.5−7.2), these carriers are not optimal for pHe targeting [54-58]. Recently, the smart polymeric micelles, which have the capability of responding to tumor pHe, have been designed [59-64]. These polymeric micelles are physically destabilized and thus accelerate the anticancer drug release at tumor pHe.

A mixed pH-sensitive micelle (PHSM) system with folate (PHSM/f), created from poly(L-histidine) (polyHis) (Mw 5,000)-b-poly(ethylene glycol) (PEG) (Mw 2,000) and poly(L-lactic acid) (PLLA) (Mw 3,000)-b-PEG (Mw 2,000)-folate (0−25 wt%), showed a gradual destabilization below pH 7.0 due to the ionization of the polyHis block in the micelle core [60]. In the drug release studies using PHSM/f containing 25 wt % PLLA-b-PEG showed a favorable pH-dependency; such that within 24 hours 32 wt% of doxorubicin (DOX) was released at pH 7.0, 70 wt% of DOX at pH 6.8, and 82 wt% at pH 5.0. Consequently this enhanced the killing effect on sensitive cancer cells below pH 7.0. Furthermore, the DOX loaded PHSM/f (equivalent DOX=10 mg/kg) exhibited significant inhibition (P<0.05 compared with free DOX or saline solution) on the growth of s.c. MCF-7 xenografts [44]. The tumor volume of mice treated with the PHSM/f (P<0.05 compared with free DOX) was approximately 4.5 to 3.6 times smaller than those treated with saline solution or free DOX after 6 weeks. When tested with DOX loaded polyHis(Mw 5,000)-co-PEG (Mw 3000) micelle, without folate and mixing with PLLA-b-PEG, in MDA 231 MD breast tumor-bearing mice model, the time-dependent DOX accumulation was visualized using skinfold window chamber model (Figure 1). The intensity of DOX fluorescence carried by PHSM was significantly more intense and spread in the tumor site than that carried by a control pH-insensitive (PLLA-b-PEG) micelle [65]. This is because once the micelles are exposed to pHe, the micelles dissociated and released the payload. The micelle dissociation may also help the extravasation of next arriving micelles by providing space. Therefore, the pH induced micelle destabilization and triggered release of DOX by tumor pHe, after the accumulation of the micelles in the tumor sites via enhanced permeability [66], presented a more effective modality of chemotherapy for sensitive tumors by providing higher local concentrations of the drug at tumor sites and minimal release of the drug from micelles during blood circulation (pH 7.4).

Figure 1.

Figure 1

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Mouse dorsal skin fold window chamber made of two symmetrical titanium frames. Tumor piece was inoculated into nu/nu mouse window chamber (A). After implanting MDA 213 breast cancer tumor piece, tumor blood vessels were growing in the window chamber Day 1 (B) and Day 15 (C). Normal blood vessel images after IV injection of DOX loaded pH-sensitive micelles (polyHis (Mw 5,000)-b-PEG (Mw 3,000)) 5 min and 60 min were shown in (D) and (E), respectively. The tumor blood vessels after i.v. injection of DOX loaded pH-insensitive and pH-sensitive micelles at the time course for 60 min were present in row (F) and row (G). The bright color is from DOX fluorescence. Reproduced with permission from reference [64].

The detailed physicochemical characteristics of a mixed PHSM, including the size, pH-dependent size change-dissociation kinetics, stability, the compatibility of core forming polymer blocks are reported in reference [67] and PK data using polyHis-b-PEG micelles were described in reference [63]. In addition, the brief toxicity results of the micelles and MTD were reported in reference [68].

2.2. TAT exposure by shield/deshielding mechanism

The shield/deshielding mechanism by positive charges of TAT, a cell penetrating peptide, on micelle surfaces controlled by the pH difference between 7.4 and pHe was designed [69] (Figure 2). Poly(methacryloyl sulfadimethoxine)-b-PEG is negatively charged and interacts electrostatically with TAT molecules (shielding) at pH 7.4. However, charge density on this polymer decreases by decreasing pH. Below pH 6.8, due to destabilized electrostatic interactions, the TAT will be deshielded. The zeta potential measurements on micelle comprising of PLLA-b-PEG-TAT demonstrated the shield/deshielding process. It was shown that the zeta potential is close to zero between pH 8.0 to 6.8, which indicates complete shielding of TAT, and from pH 6.6 to 6.0 it increased to 6.0 mV, which is close to the measured zeta potential for TAT decorated micelles without masking. When the shielded and unshielded TAT-micelles were tested for tumor cell internalization at pHs 7.4 and 6.6 by incubating for an hour, unshielded micelles were internalized into both the cells and its nucleus at pH 7.4 and 6.6. However, the micelle shielded with poly(methacryloyl sulfadimethoxine)-b-PEG was not internalized at 7.4 indicating TAT was masked even as it internalized into cells and the nucleus at pH 6.6. This shield/deshielding mechanism suggests that an optimized pHe targeting system with an appropriate sulfonamide polymer is feasible. The optimized system can be further tested in vivo.

Figure 2.

Figure 2

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Schematic concept for a proposed drug delivery system: the carrier system consists of two components, poly(L-lactic acid)-b-PEG-TAT micelles and pH-sensitive poly(methacryloyl sulfadimethoxine-b-PEG. a) At normal blood pH, polysulfonamide is negatively charged, and when mixed with TAT, polysulfonamide shields the TAT by electrostatic interaction. Only PEG is exposed to the outside which could make the carrier long circulating; b) when the system experiences a decrease in pH (near tumor) polysulfonamide loses charge and detaches, exposing TAT for interaction with tumor cells. Reproduced with permission from reference [68].

2.3. Ligand exposure by pop up mechanism

This mechanism is described in Section 4.

3. Endosomal pH (pHendo) Targeting

Accelerated anticancer drug release from L-histidine-based polymeric micelles could be triggered by an early endosomal pH of 6.0 [62]. The primary objective of this strategy is to create drug-loaded micelles that destabilizes at an early endosomal pH of 6.0 such that drug release at both the tumor extracellular pH (pHe) and the lysosomal pH of 5.0 can be minimized. This system is shown to be effective for cytosolic high dose drug delivery with minimal drug loss during circulation and in extracellular domain. It is also assumed that the endosolytic activity at early endosomal pH may minimize the leakage of digestive lysosomal enzymes [70].

3.1. Receptor-mediated endocytosis and pHendo targeting

When a mixed PHSM/f micelle with up to 40 wt% PLLA-b-PEG was used, a fraction of this micelle, depending on PLLA-b-PEG content, was destabilized by tumor extracellular pH and the remaining was internalized by folate receptor mediated endocytosis. To eliminate this micelle destabilization before internalization, a micelle system consisted of poly(His-co-phenylalanine (Phe))-b-PEG and PLLA-b-PEG-folate was designed [62]. In particular, the pH sensitivity of the micelle is controlled by the His/Phe block composition, and was fine tuned to target early endosomal pH through blending with PLLA-b-PEG by using anticancer drug, DOX. Since pKb of poly(His-co-Phe (16 mole%))-b-PEG was around 6.3, this block copolymer was blended with 20 wt % of PLLA-b-PEG-folate for targeting endosomal pH of 6.0. This micelle (as denoted ‘EndoPHSM/f’) releases minimal drug above pH 6.0 and demonstrated the triggered release at pH 6.0, indicating tuned micelle destabilization at early endosomal pH. When EndoPHSM/f was internalized into tumor cells via folate receptor-mediated endocytosis, this system effectively killed tumor cells through a focal high dose of DOX in the cytosol, resulting from active internalization, accelerated DOX release triggered by endosomal pH, and disruption of endosomal membrane. Currently, In vivo studies using this micelle system are in progress, and the results will be reported shortly.

3.2. Overcoming MDR

Most antitumor agents are not very effective in tumor chemotherapy: as anticipated that may be because of MDR in various tumor cells. Additionally, premature drug release in blood leads to systemic side effects, and fails to locally concentrate at the site of action. Especially for drug resistant cells, slow drug release kinetics in tumors may decrease the drug efficacy at the site [53]. Consequently, this issue has provoked a strong interest to create a new drug delivery system for the intracellular focal high dose targeting based on triggered release mechanism to achieve maximal therapeutic efficacy [71, 72].

After identification of Pgp as a major cause of MDR in in vitro cell studies and the discovery that verapamil modulates Pgp, tremendous efforts have been made for the identification of more effective Pgp modulators, which have high binding affinity to Pgp without any disruption to normal biological functions and with minimal pharmacokinetic interference [73-78]. Numerous Pgp modulators have been synthesized by chemical modification using existing drugs such as verapamil, cyclosporine A, glibenclamide and many other compounds that are in various stages of clinical studies: phase I, II and even III [73-78]. However, none of these chemical entities has been proven to be effective in the phase III studies yet. MDR, in clinical settings, is multifaceted and only the Pgp inhibition does not seem to be as effective as anticipated [79]. As an alternative to Pgp modulators, various drug carriers have been extensively tested to overcome MDR in vitro and in vivo [80-89]. Nevertheless, none of the approach has been proven to be effective in clinical MDR tumor treatments so far.

EndoPHSM/f micelles are shown to be highly effective in treatment of MDR tumor cells [61]. The endosomal pH triggering anticancer drug release and the endosomal escaping activity of polyHis allows cytosolic delivery of anticancer drug, by avoiding drug sequestration mechanism in MDR cells [43, 61] and through bypassing MDR proteins expression on cellular membrane via folate receptor mediated endocytosis [90]. The EndoPHSM/f demonstrated a similar degree of cytotoxicity against MDR tumor cells (MCF-7/DOXR with Pgp overexpression) when compared with free DOX against the drug-sensitive tumor cells. Investigation on A2780/AD (ovarian carcinoma drug-resistant tumor) xenografts in nude mice for in vivo efficacy demonstrated the tumor regression in mice treated by EndoPHSM/f was extremely promising and superior to PHSM/f (unpublished data).

4. Double Targeting, pHe and pHendo: a universal approach for drug-sensitive and resistant tumors

In an attempt to address the issues related to MDR and tumor heterogeneity simultaneously, one approach is to use tumor cell non-specific interactions that can be activated by tumor microclimate such as pH (pHe targeting) along with triggered release in the endosomes.

Polymeric micelles with pH-induced ligand repositioning on the micelle surface was developed [60]. As shown in Figure 3, the mixed micelle consisted of polyHis-b-PEG and PLLA-b-PEG-b-polyHis (Mw 1,000)-biotin, which is multifunctional; the shorter polyHis block in PLLA-b-PEG-b-polyHis-biotin is located at the interface of the hydrophobic core of PLLA and polyHis and the hydrophilic PEG shell, due to the high water solubility of neighboring PEG and biotin. Biotin was selected in this study to demonstrate the proof of concept. The interfacial polyHis caused PEG chain bending and the biotin burying in the PEG shell, derived from the polyHis-b-PEG block copolymer. As a result, the micelle is stable above pH 7.2 and hides the conjugated biotins. However, as the pH is lowered below pH 7.2, the degree of ionization of polyHis increases. The interfacial short polyHis (Mw 1,000) becomes ionized first and at the critical degree of ionization its hydrophobic interaction with the core phase weakens. As a result, the PEG-b-polyHis-biotin portion expands, exposing biotin out of the PEG shell. The pH 7.0 seems to be the critical point for this expansion as demonstrated by pH-dependent turbidity of the micelle solution containing avidin, which is a tetrameric protein with four biotin-binding sites. Furthermore, when the solution pH was decreased, the relative transparency of the solution is gradually reduced to 10 % between the pH range of 6.8−6.0. This is attributed to the ionization of the polyHis block located in the core and subsequent micelle destabilization by ionized polyHis escaping from the micelle. This process might cause reduced transparency, presumably through certain degree of aggregation of the remaining PLLA-b-PEG block copolymer. In summary, when the environmental pH for the micelle is lowered slightly (pH ∼7.0; tumor acidic pH), biotin is exposed on the micellar surface and can interact with cells, which facilitates biotin receptor-mediated endocytosis. Whereas, when the pH is lowered further (pH<6.5), the micelle destabilizes, resulting in disruption of the endosomal membrane and enhanced cytosolic drug release. This intelligent micelle clearly presented that the hidden biotin at pH 7.4 is exposed at pH 7.0 by the pop-up mechanism, which enhanced the cell cytotoxicity of the DOX loaded micelle at tumor acidic pH.

Figure 3.

Figure 3

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Schematic diagram depicting the central concept of pH induced biotin repositioning on the micelle. While above pH 7.0, biotin that is anchored on the micelle core via a pH-sensitive molecular chain actuator (polyHis) is shielded by PEG shell of the micelle; biotin is exposed on the micelle surface (6.5 < pH < 7.0) and can interact with cells, which facilitates biotin receptor-mediated endocytosis. When the pH is further lowered (pH < 6.5), the micelle destabilizes, resulting in enhanced drug release and disrupting cell membranes such as endosomal membrane. Reproduced with permission from reference [60].

To replace biotin with TAT, a micelle system was constructed with polyHis-b-PEG and PLLA-b-PEG-b-polyHis (Mw 2,000)-TAT [45]. TAT is a non-specific cell penetrating peptide, which has the strong capability to translocate the polymeric micelles into cells.

The pH-dependent micelle uptake by cells is represented in Figure 4. At pH 7.4, micelle uptake was minimized. At pH 7.0, the uptake showed a 30-fold increase compared to pH 7.4, which probably is due to partial TAT expression on micellar surface. At pH 6.8, 70-fold increased micelle cellular uptake was shown as compared to pH 7.4. This observation indicates that at tumor pHe the TAT peptide is exposed on the micellar surface and interacts with cells, which facilitates macropinocytosis. This nanosystem proved to be effective for various in vivo solid tumors including drug-sensitive and drug-resistant phenotypes and is anticipated to replace cumbersome and selective antibody or ligand-based targeting technology.

Figure 4.

Figure 4

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pH-dependent cellular uptake of the micelles in cultured MCF-7 tumor cells caused by TAT exposure: pH 7.4 (•), pH 7.0 (■), pH 6.8 (▲). Particle size and size distributions of micelle suspensions (in PBS) at pH 7.4 were measured by dynamic light scattering. Each data point represents an average with standard deviation (n=3). Reproduced with permission from reference [45].

5. Virus-mimetic (VM) nanogel

Viruses infect specific cells of the host organisms, replicate, destruct the cells and spread from one to another cell causing diseases [91, 92]. They circulate long in the blood and as disease progress they become more pathogenic [93]. Drug delivery vehicles often mimic only few aspects of virus such as size, surface modifications for longer residence in the body before their clearance [94-97], which has attracted many investigators who design delivery vehicles particularly for anticancer agents [98-100]. In particular, the nanosystem including virus-like infectious properties can be spotlighted. This system has a capsid-like protein capsule, able to infect specific cells, injects toxin, destroy infected cells, and migrates to neighboring pathologic cells by repeated cell cycles.

The virus-like infectious nanogel consists of a hydrophobic core (poly(His-co-Phe)) and two layers of hydrophilic shells (PEG and bovine serum albumin (BSA)) (Figure 5) [46]. One end of PEG is linked to the core forming block and other to BSA, which forms a capsid-like outer shell. The structure of core and inner shell was formed by oil-in-water emulsion method [101-103]. At high pH the core of this nanogel is rigid; however, the core swells by the ionization of polyHis at low pH. When these nanogels are exposed to early endosomal pH of 6.4 [45], the size grew abruptly, reaching 355 nm. The size changes by cycling pH between 7.4 and 6.4 are also reversible [46]. This reversible swelling/deswelling by pH of the core is closely linked to the release rate of incorporated DOX (Figure 6). The nanogels release a significant amount of DOX at endosomal pH (e.g., pH 6.4), while reducing DOX release rate at cytosolic [48] or extracellular pH (e.g., pH 7.4−6.8). Furthermore, due to the known proton buffering effect [58] of polyHis and observed substantial nanogel volumetric expansion within cell endosomes, nanogels are proposed to be able to physically disrupt endosomal membranes. This allows the VM nanogels and anticancer drug (released from the nanogels) to transfer from the endosomes to the cytosol, where the VM nanogels rapidly shrink back to their original size with the new, more neutral local pH, and thereby reduce the drug release rate. Free drug released by endosomal pH stimulus and will be in the cytosol, then diffuses into the nucleus, and finally to the pharmacological target site. The drug action on the cells will induce apoptosis and eventually disintegrates, which in turn releases the nanogels from the cell for subsequent infection and action in neighboring cells (Figure 7). This nanogel demonstrated repeated infectious cycles in cultures of drug-resistant tumor cells. Consequently, this appraoch is thought to have a high potential for maximizing drug efficacy in treating tumors, inflamed tissues, and other diseases due to sequential cytotoxic action. Further in vitro evaluations and in vivo investigations are required to confirm their potential.

Figure 5.

Figure 5

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Schematic presentation of the virus-like nanogel. See the text for more details.

Figure 6.

Figure 6

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Release rate of DOX from DOX-loaded virus-like infectious nanogels. 150 μg of DOX was encapsulated into 1 mg of DOX-loaded VM-nanogels. The pH of the solution is stepwise adjusted to pH 7.4, pH 6.8 and pH 6.4 at one-hour intervals. Each data point represents an average with standard deviation (n=3). Reproduced with permission from reference [46].

Figure 7.

Figure 7

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Virus-like nanogels infect cells selectively depending on specific interactions, kill the host cells, and migrate to neighboring cells as virus does to repeat the cycle.

6. Conclusion

Tumor extracellular pH- and/or endosomal pH-responsive micelles, TAT shield/deshield nanosystem, virus-like infectious nanogels, and pop-up micelles are the examples of the novel anticancer drug delivery systems for overcoming limitations of conventional drug delivery systems. These systems increase target drug accumulation at tumor sites or at intercellular cytosolic compartments in tumor cells with less drug distribution to normal tissues and organs. In particular, the pH-sensitive micelles or nanogels presented here are the unique delivery systems in the treatment of MDR. The constituting polymer components also have no apparent cytotoxicity and systemic toxicity in small animal model. These outlooks motivate us to carry out a preclinical study, although more detailed studies are need to further prove the hypotheses. In the future, these technologies will be extended by using various anticancer drugs and molecular tools for active internalization to achieve a general platform for solid cancer chemotherapy.

8. Acknowledgments

The works covered in this review article are supported by NIH CA 101850 and 122356.

Footnotes

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