Block Copolymer Based Nanoparticles for Theranostic Intervention of

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Block Copolymer Based Nanoparticles for Theranostic Intervention of Cervical Cancer: Synthesis, Pharmacokinetics, and in Vitro/in Vivo Evaluation in HeLa Xenograft Models Shweta Dumoga,†,‡ Yogesh Rai,† Anant Narayan Bhatt,† Anjani Kumar Tiwari,† Surendra Singh,‡ Anil K. Mishra,*,† and Dipti Kakkar*,† †

Institute of Nuclear Medicine and Allied Sciences, Defence Research and Development Organization, Brig. S.K. Mazumdar Road, Timarpur, Delhi 110054, India ‡ Department of Chemistry, University of Delhi, Delhi 110007, India S Supporting Information *

ABSTRACT: Polymer-based nanoparticles have proven to be viable carriers of therapeutic agents. In this study, we have developed nanoparticles (NPs) from polypeptide-polyethylene glycol based triblock and diblock copolymers. The synthesized block copolymers poly(ethylene glycol)-b-poly(glutamic acid)-b-poly(ethylene glycol) (GEG) and poly(ethylene glycol)-bpoly(glutamic acid) (EG) conjugated with folic acid for targeting specificity (EGFA) have been used to encapsulate methotrexate (MTX) to form M-GEG and M-EGFA NPs aimed at passive and active targeting of cervical carcinoma. In-vitro SRB cytotoxicity and hemolysis assays revealed that these NPs were cytocompatible to healthy human cells and hemocompatible to human RBCs. Cellular uptake by FACS demonstrated their prompt internalization by human cervical carcinoma (HeLa) cells and points toward an apoptotic mechanism of cell kill as confirmed by AO/EB staining as well as histological analysis of explanted HeLa tumors. Pharmacokinetics and biodistribution studies were performed in New Zealand albino rabbits and HeLa xenografted Athymic mice models, respectively, by radiolabeling these NPs with 99mTc. Passive tumor accumulation and active targeting of MTX-loaded polymeric nanoparticles to folate expressing cells were confirmed by intravenous administration of these 99mTclabeled M-GEG and M-EGFA NPs in HeLa tumor bearing nude mice and clearly visualized by whole-body gamma-SPECT images of these mice. Survival studies of these xenografted mice established the antiproliferative effect of these MTX-loaded NPs while corroborating the targeting effect of folic acid. These studies proved that the M-GEG NPs and M-EGFA NPs could be effective alternatives to conventional chemotherapy along with simultaneous diagnostic abilities and thus potentially viable theranostic options for human cervical carcinoma. KEYWORDS: block copolymers, MTX, folate targeting, HeLa cells, theranostic nanoparticles, cervical cancer

1. INTRODUCTION Cancer has become a major bottleneck in the progress of medical sciences which is facing numerous challenges in its attempt to completely eradicate this disease. Invasive cervical cancer is the second most common malignancy among women globally, with almost 50% mortality rate and developing countries accounting to nearly one-third of the global cervical cancer deaths. Though chemotherapy using small molecular antitumor agents like cisplatin, docetaxel, methotrexate, doxorubicin, etc. is the applied treatment methodology in © XXXX American Chemical Society

cancer therapy, it is associated with severe adverse side effects, primarily due to nonspecific drug distribution as well as rapid clearance from the circulation.1,2 It is a fundamental challenge of all anticancer therapies to deliver sufficient therapeutic agent into cancer cells while sparing normal cells and tissues. Nanotechnology-based novel carrier systems like liposomes, Received: April 9, 2017 Accepted: June 13, 2017 Published: June 13, 2017 A

DOI: 10.1021/acsami.7b04982 ACS Appl. Mater. Interfaces XXXX, XXX, XXX−XXX

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ACS Applied Materials & Interfaces

collateral damage to negative tissues but also enhance the cellular uptake of the nanocarriers and thus contribute to treatment efficacy. Folate receptor is one such biomarker that is highly overexpressed on human cervical cancer cells by almost 36% as compared to other FR (+ve) cell lines like SKOV-3, MCF-7, and MDA-MB-231.19 Its structural analog and FDA approved antimetabolite, Methotrexate, is a chemotherapeutic agent that forms an integral part of different anticancer regimes effectively used for the treatment of various human malignancies, including acute lymphoblastic leukemia, osteosarcoma, Hodgkin’s disease, cervical cancer, choriocarcinoma, and related trophoblastic tumors.20 However, its high dose regimen can potentially cause significant toxicity, including hepatotoxicity, neurological damage, and memory loss (by crossing the blood brain barrier), interstitial pneumonitis, nephrotoxicity, and at times acute renal failure. Nonetheless, encapsulation of MTX in polymeric nanoparticles could prove to be an effective strategy to overcome some of the undesirable side effects encountered while using the drug in its native form. Though passive membrane permeability of Methotrexate is minimal, due to its polar surface (owing to presence of two carboxylate moieties in its structure) it circumvents this obstacle by taking the help of multiple cellular uptake mechanisms for cellular internalization. These include influx transporters vis-a-vis reduced folate carriers (RFCs) (KD = 4.3 μM), proton-coupled FA transporter, and membrane-associated FA receptors (KD > 100 nM).21 Despite this, huge amounts of the drug needs to be administered in the case of conventional chemotherapy. This is because on cell internalization, the majority (almost 75%) of the free MTX reportedly gets effluxed out of cancerous cells (in the mono, di, or triglutamate form) due to ATP-driven efflux pumps (ATP-binding cassette (ABC) transporters) in cellular membrane like ABC B1(pgP), ABC C (MRP(1−4)), BCRP (ABC G2), etc. which expel out xenobiotics to counteract drug accumulation in cells (Figure 1). This is also one of the reasons for cells developing resistance to the drug. As the polyglutamylation progresses beyond four glutamate derivatives (tetraglutamates), the membrane transporters like ABCG2, MRP, etc. become ineffective, resulting in

polymer-drug conjugates, nanogels, nanospheres, polymeric micelles, and solid lipid nanoparticles offer potential solutions to meet these challenges.3 Of these, polymeric nanocarrier systems have attracted much attention due to their favorable characteristics like possibility of diverse composition, flexible methods of synthesis, ease of functionalization, and variation in surface chemistry that can be achieved by suitable alterations in the constituent polymers as per the intended application.4 These properties can influence their fate in the physiological environment and their drug-delivery features in particular. Polymer-based nanoparticles have offered some of the most promising opportunities for in vivo diagnosis and treatment of many diseases as evident from the plethora of formulations approved for clinical use or already in advanced stages of clinical trials. These include polymeric nanomedicines like PEG asparaginase (Oncaspar), PEG-b-poly(D,L-Lactic acid) (Genexol-PM), PEG-b-polyaspartate(NK 105/Nanocarrier), PEG-bpoly(glutamic acid) (NC 6004 Nanoplatin), PEG-encapsulated irrinotecan (NKTR 102), PEG-cyclodextrin (CRLX 101/ Cerulean), Poly(hydroxypropyl methacrylate) (AP 5346/ Prolindac), and polyacetal-b-poly(1-hydroxy methyl (ethylene hydroxy methyl formal) (XMT 1001).5,6 However, among the anticancer nanomedicines in clinical trials, only the docetaxelloaded PEG−PLGA nanoparticles (BIND-014, BIND Therapeutics) is an investigational nanomedicine in phase II clinical trial stage that caters to cervical cancer in addition to several others like prostrate, metastatic NSCLC, head and neck, cholangiocarcinoma, and bladder cancer.7 Since these nanoparticles target the prostrate-specific membrane antigen (PSMA) for delivery of docetaxel, as on date there is no globally accepted system for cervical cancer either in the market or in the clinical trial stage. As seen above, block copolymers are major role players here owing to their favorable attributes of self-assembly in aqueous media into particles at a nanoscale.8,9 Most of them predominantly have PEG/PEO as the nontoxic hydrophilic block.10,11 Use of PEG not only offers improved water solubility but also provides longer circulation time. This is due to its ability to mask the antigenic determinants of proteins, abrogating their immunogenicity.12 Incorporation of a synthetic polypeptide like polyglutamic acid as a constituent of the block copolymer along with PEG offers additional benefits owing to their attractive features of long-term biodegradability, stable secondary structures formed by them, accompanied by their inherent biocompatibility, as compared to traditional synthetic polymers.13−15 Nanoparticles formulated from such copolymeric systems can serve as effective delivery vehicles for chemotherapeutic agents due to their ability to carry large amounts of therapeutic cargos within their compartments. These passively targeted nanoparticles rely on the EPR effect and gain from the leaky tumor vasculature and impaired lymphatic drainage of tumor tissues to gain retention in the tumor microenvironment.16 Additionally, they offer the option to functionalize them with specific ligands intended to be recognized and specifically bind to receptors ubiquitously expressed by most malignancies on their cell surface. Taking advantage of this possibility, tumor specificity can be achieved by the active targeting approach.17,18 These include a myriad of targeting moieties vis-a-vis folate receptor, glycoproteins, epidermal growth factor (EGFR), human estrogen receptor (HER), aptamers, CD44, PSMA, etc. When linked to a therapeutic drug or its carrier, these ligands can be exploited to carry the nonselective drug specifically into the cancer cell. The active targeting of cancer cells can not only avoid the unwanted

Figure 1. Schematic showing contribution of ABC transporters in cellular efflux of methotrexate. Free methotrexate is effluxed out of the cell by MRP1−4, whereas its di- and triglutamates are extruded by ABC G2. (DHFR, dihydrofolate reductase; RFC, reduced folate carrier; MRP, multidrug resistance protein; FPGS, folylpolyglutamate synthetase; and γ-GHase, γ-glutamyl hydrolase.) B


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Figure 2. Schematic representation of actively targeted M-EGFA and passively targeted M-GEG polymeric nanoparticles for theranostic intervention of Cervical cancer.

a slow exit of the higher chain polyglutamates from the cells as they find it difficult to traverse the cellular lipid bilayers. This leads to retention of the polyglutamates (which account for only 10−25% of the internalized dose) inside the cells to carry out anticancer activity due to their similar affinity to DHFR as MTX. However, once encapsulated inside the polymeric nanoparticles, uptake of MTX would not only decrease in normal cells due to shielding offered by the polymeric shell but would concomitantly enhance in the tumor cells with the aid of membrane-associated FRα (via RME), which are up-regulated in malignant cells, thus bypassing the cell-transport mechanism and releasing an overwhelming concentrated drug cargo deep within the cytoplasm.22 The present work was therefore initiated to design a system that could meet at least some of the challenges faced by drug delivery systems (DDS). In this work, we have designed two appropriate nanoscaled carrier systems for effective entrapment of the chemotherapeutic drug MTX, for improved systemic circulation and intracellular delivery. A PGA-b-PEG-b-PGA triblock copolymer has been synthesized to evaluate its potential as a passively targeted system (formulated as MGEG NPs) and its folate conjugated diblock counterpart (PGAb-PEG-FA) (formulated as M-EGFA NPs) for active targeting of the drug-loaded system to folate receptors on tumor cells using the conjugated folic acid on the surface of nanoparticles (Figure 2). This research paper entails the complete biological evaluation of these drug-loaded nanoformulations in terms of cytotoxicity, intracellular uptake (both qualitatively and quantitatively), and mode of cell kill (apoptosis) in human carcinoma HeLa cells. The hemolysis and antitumor efficacy of these MTX-loaded polymeric nanoparticles have been investigated in HeLa xenografts in Athymic mice while in vivo biodistribution and pharmacokinetics have been performed with 99mTc-labeled nanoparticles. Tumors excised from these models were then studied for histopathological changes in order to comprehend the pathological alterations that may be encountered due to differential delivery of drug (MTX) by the passive or active approach. Eventually, longevity studies in these

xenografts brought about the differentiation pattern of both the approaches as discussed in the article.

2. MATERIALS AND METHODS 2.1. Materials. All the chemicals and reagents were of analytical grade and used without any further purification unless otherwise stated. Poly glutamic acid (PGA) (Mw ∼ 6449), poly ethylene glycol diacrylate (PEGDA) (Mw 6000), poly ethylene glycol methacrylate (PEGMA) (Mw 500), folic acid, 97%(FA), N,N′-dicyclohexylcarbodiimide, ≥ 99.0%, (DCC), 4-(dimethylamino)pyridine (DMAP), and N,N-dimethylformamide, ≥ 99%, (DMF) were purchased from Aldrich. Triethyl amine (TEA), methanol (MeOH), methotrexate (99%), rhodamine 123 (dye content 99%), sulphorhodamine B, trichloro acetic acid (TCA), tris(hydroxymethyl)aminomethane (Tris base), bisbenzimide trihydrochloride (Hoechst 33342), Dulbecco’s modified eagle’s medium (DMEM), cellulose dialysis membrane (12 kDa), acridine orange (AO), and ethidium bromide (EB) were procured from Sigma-Aldrich (St. Louis, MO) and used as received. 2.2. Cell Culture and Animal Models. Human cervical carcinoma cell line (HeLa cells) were procured from American Type Culture Collection (ATCC). HeLa cells were cultured in FA-deficient DMEM supplemented with 10% fetal bovine serum (FBS) and 1% (v/ v) penicillin-streptomycin. The cells were cultivated in a humidified atmosphere in an incubator (Thermo Scientific) at 37 °C with 5% CO2 for 24 h. Female balb/c athymic mice aged 5−6 weeks, weighing 23 ± 2g were obtained from our experimental animal facility (EAF) and provided with standard food and water ad libitum. They were carefully housed at temperature and humidity conditions of 21 ± 2 °C and 50 ± 5%, respectively, with 250−300 lx light input. Tumor models were prepared by subcutaneous injection of 1 × 106 tumorigenic cells in a volume of 100 μL, in the right hind leg.23,24 All animal experiments and study protocols were approved by the institutional animal ethical committee (IAEC). 2.3. Characterization. 1H NMR and 13C NMR spectra were recorded on a Bruker Avance 400 MHz spectrometer in D2O at 298 K using 15−20 mg compound in WILMAD NMR tubes (5 mm diameter). Functional groups of the block copolymers were identified using a Thermo scientific Nicolet 8700 Infrared microscopy (ATR) spectrometer providing additional evidence for the synthesis. Molecular weight distributions (polydispersity index, PDI = Mw/Mn) of the block copolymers were determined by gel permeation C


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ACS Applied Materials & Interfaces chromatography (GPC) using a Viscotek TDA305 system (Malvern Instruments). The thermal stability of the block copolymers and their respective precursors were analyzed by thermogravimetric analysis (Pyris Diamond TGA, PerkinElmer) over a temperature range of 50−900 °C, under a nitrogen atmosphere at a heating rate of 10 °C/min. The change in other thermal characteristics on conjugation was verified by differential thermal scanning (DSC Q20, TA Instruments). 2.4. Synthesis of Block Copolymers. Synthesis of PGA-b-PEGb-PGA Triblock Copolymer. A simple Michael addition approach was used to synthesize the PGA-b-PEG-b-PGA triblock copolymer.25 Briefly, poly(ethylene glycol) diacrylate (PEGDA) (0.01163 mmol, 1 equiv) was dissolved in anhydrous MeOH(10 mL) and added to a stirred solution of poly glutamic acid (PGA) (0.0233 mmol, 2 equiv) in anhydrous MeOH(10 mL) and triethylamine (TEA) (0.0233 mmol, 1 equiv) at constant stirring under a nitrogen atmosphere. After 72 h at RT, the solvent (methanol) and TEA were evaporated, and the residue was collected and analyzed by chemical characterization. To verify the progress of the reaction, the ninhydrin test (with nbutanol:acetic acid:water solvent system for peptides) was performed to confirm the conjugation between 1°-NH2 of PGA and acrylate moiety of PEGDA (details in the Supporting Information). Synthesis of PGA-b-PEG-FA Diblock Copolymer. The PGA-b-PEG diblock copolymer was synthesized by the same Michael addition approach by using poly(ethylene glycol) monoacrylate (PEGMA) (MW = 500) in place of poly ethylene glycol diacrylate (PEGDA), while the rest of the procedure was the same as for the preparation of the triblock copolymer. Briefly, a methanolic solution of poly ethylene glycol monoacrylate (PEGMA) (0.01163 mmol) was added to a methanolic solution of poly glutamic acid (PGA) (0.0233 mmol, in 10 mL anhydrous MeOH) and triethylamine (TEA) (0.0233 mmol) at RT and kept for stirring up to 72 h. The completion of the reaction was monitored in situ by the ninhydrin test (as mentioned above). The reaction mixture was then concentrated by evaporation of the solvent and TEA, and the obtained residue was precipitated in diethyl ether, air-dried, and the compound was analyzed by chemical characterization. Folic acid was conjugated to the obtained PGA-b-PEG diblock copolymer by the Steglich esterification method.26 For prior activation of the carboxylic acid group, DCC (0.03597 mmol) and DMAP (0.01439 mmol) were codissolved in DMF (5 mL) and added to folic acid dihydrate (0.01439 mmol) in a reaction flask, under stirring at RT with protection from light. PGA-b-PEG diblock copolymer (0.01439 mmol) was dissolved in DMF (5 mL) and stirred for 45 min and added to the reaction flask dropwise under nitrogen atmosphere and left for stirring at RT. After 72 h of continuous stirring, the reaction mixture was concentrated by evaporation of solvent (DMF) under reduced pressure and the final product was recovered by precipitation in cold diethyl ether. 2.5. Preparation of MTX-Loaded Polymeric Nanoparticles (NPs). Nanoparticle formation of triblock and diblock copolymer was carried out using the nanoprecipitation technique while utilizing the dropping method.27 Dialysis was performed to remove the organic solvent. This easy and simple technique is based on the intrafacial deposition due to the displacement of a solvent with the nonsolvent and gives good reliable results.28 Both the block copolymers were dissolved in dimethyl sulfoxide (DMSO) (1 mL) in various weight ratios, and the concentration of block copolymers varied from 0.1 mg/ mL to 1.0 mg/mL. Distilled water (9 mL) was then added to the polymer solution at high speed under vigorous stirring at RT. The mixed solution was dialyzed against distilled water for 24 h using a dialysis bag (MWCO = 3000 Da) to remove the organic solvent. The distilled water was replenished at an interval of every 2−3 h to ensure removal of the organic solvent. Complete removal of DMSO was verified by measurement of UV−vis absorbance of the dialysate from time-to-time (Figure S3). For this study, methotrexate (MTX) was used as the model anticancer drug to generate a proof of concept for this theranostic platform. The preparation procedure of MTX-loaded polymeric nanoparticles was similar to that of the unloaded polymeric

nanoparticles but protected from light exposure. MTX and block copolymers were added in various weight ratios. MTX was dissolved in DMSO (1 mL) along with the copolymers. The concentrations of MTX was varied from 0.1 to 1.0 mg/mL, while that of the copolymers was kept at 1.0 mg/mL. The mixed solution was dialyzed against distilled water under dark conditions for 24 h to remove the unloaded drug as well as the undesired organic solvent. Rhodamine-labeled M-GEG and M-EGFA NPs were prepared by covalently conjugating rhodamine with methotrexate(M-Rh), as per the procedure elaborated in the Supporting Information. The conjugated M-Rh was then encapsulated in GEG and EGFA by the same procedure as used to encapsulate MTX alone. Determination of Drug Content in the Nanoparticles. To quantify the amount of drug encapsulated in the polymeric nanoparticles, they were added to 2 mL DMSO in order to rupture the nanoparticles and release the entrapped drug. The obtained solution was analyzed by a UV−vis spectrophotometer (Lambda 365 UV/vis, PerkinElmer). The characteristic absorbance of MTX was recorded at 303 nm and compared with a standard curve for the drug generated in DMSO, containing MTX concentrations varying from 0− 100 μg/mL. Finally the drug-entrapment efficiency and drug-loading content were calculated from the following formula:

drug loading (%DL) amount of drug encapsulated in NPs = × 100 amount of drug added entrapment efficiency (%EE) amount of drug encapsulated in NPs = × 100 amount of drug loaded NPs 2.6. Nanoparticle Characterization. Both M-GEG and M-EGFA NPs were characterized by physicochemical methods. Dynamic light scattering (DLS) and zeta potential: the mean diameter of both the drug-loaded polymeric nanoparticles in the dispersion state was determined by a NanoSizer 3000 ZS90 (Malvern Instruments, Malvern, UK) at a fixed angle of 90 deg. Samples were diluted 10fold with 0.45 μm filtered dust-free distilled water to give the recommended scattering intensity. Each value obtained was the mean of three measurements of 120 s each. The hydrodynamic diameter was determined by using the Stokes−Einstein equation and calculated from the autocorrelation function of the intensity of light scattered from particles, assuming a spherical form for the particles.29 The particle charge was quantified as the zeta-potential by laser Doppler anemometry, using a NanoSizer 3000 (Malvern). Samples were prepared by dispersing the particles in PBS buffer (pH 7.4 and ionic strength 10 mM). Zeta-potentials were determined by the device, according to Smoluchowski’s equation from the mean electrophoretic mobility.30 The results are the mean of three determinations. Scanning Electron Microscopy (SEM) and Atomic Force Microscopy (AFM). Surface morphology of both the drug-loaded nanoparticles were determined by Zeiss Supra55, scanning electron microscope. Samples were coated on silicon wafers and then dried at RT (23 ± 2 °C). The dried samples were finally mounted on stubs and coated with gold (∼20 nm thickness) using a Sputter Coater JFC-1100 (JEOL, Japan) and then observed under the microscope. The same samples were used for further verification of morphology analysis by atomic force microscopy (ScanAsyst, Bruker) carried out in the tapping mode. 2.7. In Vitro Drug Release and Kinetics. The in vitro drug release profiles of the MTX-loaded polymeric nanoparticles were monitored by the dialysis method at three different pH conditions. Phosphate-buffered saline (PBS pH 7.4, 0.15 M), acetate buffer pH 5.5, and acetate buffer pH 4.5 were used as the dialysis media, respectively. A 2 mL aliquot of the drug-loaded nanoparticles (1 mg/ mL) was placed in an end-sealed dialysis bag (5 mL Float-a-lyser MWCO 3.5−5 kDa) and kept in a dialysis medium containing 1000 mL of the three buffers, respectively, at 37 °C with a stirring speed of 200 rpm up to a time point until the absorbance value of the drug D


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out with Hoechst dye (1 μg/mL; 10 min) followed by washing with PBS. The cellular internalization of the drug and drug-loaded NPs was then viewed by an inverted fluorescence microscope (Olympus BC 60 Fluorescence Microscope, Japan). Quantitative Cell Uptake. Flow cytometry was used to quantify the cellular uptake of rhodamine-labeled M-GEG and M-EGFA NPs in HeLa cells. HeLa cells, seeded in a PD-35 at a cell density of 0.075 × 106 cells/well were incubated at 37 °C and 5% CO2 for 24 h. After 24 h, rhodamine-labeled M-Rh, M-Rh-GEG, and M-Rh-EGFA NPs (5, 10, 15, and 20 μM) were added to the medium under dark conditions. The mean fluorescence intensity (MFI) of M-Rh in the cells was analyzed by FACS (LSR-II Flow Cytometer, BD, Mountainview) at an excitation wavelength of 511 nm and emission wavelength of 530 nm. Untreated cells were taken as negative control and the autofluorescence intensity obtained from them was subtracted from the fluorescence intensity of the treated cells. 2.11. Apoptosis Study. Acridine orange/ethidium bromide (AO/ EB)-based staining was used to observe the cell kill process, which is one of the economic and convenient methods to detect apoptosis in tumor cells.36 Briefly, 3000 HeLa cells/well were seeded in a 96-well plate. Twenty-four hours post seeding, cells were treated with the MRh, M-Rh-GEG, and M-Rh-EGFA NPs (15 and 20 μM concentration each) under dark conditions. At respective time points (24, 48, and 72 h), cells were centrifuged at 100g, 10 min, followed by acridine orange and ethidium bromide staining (1:1; 100 μg/mL). Their images were then captured under a fluorescence microscope (Nikon Eclipse, TE300). 2.12. Radiolabeling Studies and in Vivo Evaluation. Radiolabeling. The drug-loaded nanoformulations, M-GEG and M-EGFA were radiolabeled with 99mTc radionuclide by the direct labeling method as per our previous work37 (details in the Supporting Information). Both the nanoparticles were labeled with 99mTc in a reproducible manner with high labeling efficiency (>99%). Maximum labeling efficiency of NPs was obtained after optimizing all the labeling parameters (vis-a-vis pH, concentration of reducing agent, temperature, and incubation time). Both 99mTc-M-GEG NPs and 99mTc-MEGFA NPs remained stable with >98% labeling up to 24 h in human serum. Pharmacokinetics. Hereafter called the technetium-labeled nanoformalations, 99mTc-M-GEG and 99mTc-M-EGFA were used for pharmacokinetic evaluation. Their blood clearance was analyzed in healthy New Zealand albino rabbits weighing 2.5−3 kg (n = 3). Threehundred microliters (10 MBq) of 99mTc-M-GEG and 99mTc-M-EGFA were administered intravenously through the dorsal ear vein in rabbits. Blood containing radiolabeled nanoformulations was withdrawn from the other ear vein at different time intervals (5 min to 24 h), and their radioactive counts were measured in a well type gamma counter. Persistence of activity in circulation was calculated as the percentage of the injected dose remaining in blood, assuming the total blood volume of the animal to be 7% of its body weight. In Vitro Serum Stability. With the use of the above blood samples, plasma was separated out by centrifugation and plasma proteins were precipitated by addition of 10% trichloroacetic acid (TCA). The remaining serum was incubated at 37 °C up to 24 h, and the stability of the 99mTc-labeled NPs in serum was evaluated by measuring their respective radioactivities in a gamma counter. Biodistribution and SPECT Imaging. HeLa tumor xenografted athymic mice were segregated into two groups of 12 mice each. Onehundred microliters (3.7 MBq) of 99mTc-M-GEG and 99mTc-M-EGFA were injected in these tumor models through their tail vein (i.v.), respectively. At set time intervals of 1 h, 2 h, 4 h, and 24 h post injection the mice (three for each time point) were first imaged using a planar gamma camera in order to visualize the in vivo uptake of 99mTclabeled nanoparticles in HeLa xenografts. Gamma images were captured using HAWK-EYE Symbia T2 gamma camera (Seimens, USA), equipped with a low-energy (140 kev) parallel collimator, and the counts were observed for the tumor region, the region of interest (ROI). After gaps of 0.5 h, the mice were quickly euthanized at the designed time points and blood was sampled through the ophthalmic artery. Their heart, lung, liver, spleen, kidney, stomach, intestine,

released into the outside medium tapered off to negligible. A 2 mL sample was withdrawn from the dialysis medium at regular time points and replaced with an equal volume of fresh medium (buffer) to maintain sink conditions. These samples were subsequently analyzed by UV spectroscopy (Lambda 365 UV/vis, PerkinElmer). Finally, the cumulative release percentage (%) of the drug from the nanoparticles at each time point was calculated as ⎡ ⎧ A ⎫⎤ cumulative release percentage (%) = ⎢1 − ⎨ t ⎬⎥× 100 ⎢⎣ ⎩ A 0 ⎭⎥⎦ where At is the amount of drug released at time t and A0 is the initial amount of drug added to the NPs at the time of loading. These experiments were conducted in triplicate. 2.8. Hemolysis Assay in Human Blood. The hemolytic activity of the polymers was investigated according to previous literature with minor alterations.31,32 Fresh human blood from healthy volunteers was collected in heparinized tubes and centrifuged at 3000 rpm for 10 min to isolate red blood cells (RBCs) in the form of a pellet. The pellet obtained was washed with cold PBS pH 7.4, centrifuged at 3000 rpm for 10 min and resuspended in the same buffer. This process was carried out thrice to ensure complete removal of serum. Further, nanoparticle suspensions of different concentrations, also prepared in PBS, were added to the erythrocyte suspension (2:8 ratio) and incubated for up to 24 h at 37 °C in an incubator. The release of hemoglobin was determined after centrifugation (3000 rpm for 10 min) by photometric analysis of the supernatant at 540 nm at intervals of 1 h, 2 h, 4 h, and 24 h, respectively. Complete hemolysis was achieved by using 1% (v/v) Triton X-100 which yielded the 100% positive control value while PBS gave the negative control value. Less than 5% hemolysis was regarded as a nontoxic level in our experiments. This value is as per ASTM 75633 and has also been reported in other literature as well.34 The experiment was performed in triplicate. The hemolysis percentage of RBCs was calculated using the following formula: hemolysis (%) =

A sample − A −vecontrol A+vecontrol − A −vecontrol

× 100

2.9. In Vitro Cytotoxicity Assay. The pharmacological activity of MTX-loaded GEG and EGFA NPs was studied against HeLa cells with MTX as a positive control and saline as a negative control. Their in vitro cytotoxicity was evaluated using the colorimetric SRB assay, which is based upon the quantitative staining of cellular proteins by the sulforhodamine B protein dye.35 HeLa cells were seeded into 96 well plates (0.005 × 106 per well). After 24 h post cell seeding, the cells were incubated with different concentrations (5, 10, 15, and 20 μM) of MTX, M-GEG NPs, and M-EGFA NPs, respectively. After 4, 8, 24, and 48 h of incubation, the media was removed and cell growth terminated. Cells were fixed with TCA (10%; 45 min), and unbound TCA was washed with deionized water. All these steps were carried out gently by ensuring minimum disturbance to the cells/plates. They were then stained with SRB (0.4%; 45 min at RT) followed by washing with 1% acetic acid and kept for air drying. Protein-bound SRB was solubilized in 10 mM Tris base (pH = 10), and absorbance was measured at 564 nm using a microplate reader (Multiwell Plate Reader, Biotech Instruments). 2.10. Cellular Uptake by Fluorescence Microscopy. The cellular uptake and intracellular release behavior were verified by fluorescence microscopy and flow cytometry. For this, methotrexate was conjugated to a fluorescence active dye, rhodamine-123, and this MTX-rhodamine conjugate (M-Rh) was encapsulated into polymeric nanoparticles (M-Rh-GEG and M-Rh-EGFA NPs) and then evaluated by fluorescence microscopy. Qualitative Cell Internalization. Human cervical carcinoma Hela cells were seeded in PD-35 containing coverslips with a density of 0.075 × 106 cells/PD and cultured for 24 h at 37 °C and 5% CO2. Twenty-four hours post seeding 5, 10, 15, and 20 μM of M-Rh, M-RhGEG, and M-Rh-EGFA NPs were added to the medium. At the respective time points (4, 8, 24, and 48 h), nuclear staining was carried E


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ACS Applied Materials & Interfaces Scheme 1. Scheme of Synthesis of (A) GEG and (B) EGFA

tumor muscle, and contralateral muscle were harvested and weighed. Remnant radioactivity in each organ was measured using a gamma counter and calculated as a percentage of injected dose per gram organ weight. The radioactivity in the tail (point of injection) was also measured to calculate the absolute radioactivity of the particular tissue/organ, and decay correction was applied to all the radioactivity measurements. 2.13. Histological Analysis. Tumor-bearing mice models for histology analysis were established with nude mice (20−25 g, female) by the same method as described earlier.38 Once the tumor volume reached up to 100 mm3 in volume, MTX, M-GEG, and M-EGFA NPs (15 mg/m2 BSA) were injected every 3 days into the mice via tail vein.39,40 For histological analysis, tumor tissues were excised from the mice on the 1st, 6th, 14th, and 29th day post injection. The excised tumors were fixed with 4% paraformaldehyde solution and stored in 70% ethanol. Tissues were also excised from the same mice but from the opposite flanks to be used as controls. The sliced tumor tissues (5 μm thickness) were stained by Hematoxylin and Eosin (H&E) to ascertain the histological alterations and were then observed under an optical microscope (Bx51, Olympus). 2.14. Survival Study and Relative Tumor Volume. A human cervical carcinoma xenograft was generated by subcutaneously injecting a HeLa cell suspension (100 μL containing 2 × 106 cells) into the right flank of the female nude balb/c mice with an average body weight of 23 ± 2 g. Once the tumors reached a volume of 100 mm3, the mice were randomly divided into the following 4 treatment groups with six mice each: control (saline), MTX, M-GEG NPs, and M-EGFA NPs. Treatments were administered intravenously every 3 days for 30 days via the tail vein injection of 0.9% NaCl (saline), MTX, M-GEG, and M-EGFA NPs. Both free drug and drug-loaded micelles were injected at a dose of 15 mg of MTX equivalent per m2 BSA. To quantify the antitumor efficacy, the tumor sizes were measured using a vernier calliper and quantified according to the following equation: volume = (length × width2)/2. The relative tumor volume was obtained by the following formula: relative tumor volume (%) = (tumor volume on that day/initial value before the first dosing) × 100%. Between the 14th and 29th day, the number of mice in the control group became insufficient to carry out analysis due to early death, so the tumor growth study was ended, while the life span study ended on day 35 of treatment.

polyglutamic acid-block-polyethylene glycol-block-poly glutamic acid [PGA-b-PEG-b-PGA (GEG)] in an ABA arrangement and diblock copolymer system based on poly ethylene glycol mono acrylate (PEGMA) and PGA to form a PEG-b-PGA (EG) diblock copolymer and further attached with a folic acid biomarker (via ester linkage by Steglich esterification) to form PEG-b-PGA-FA (EGFA). Subsequently, the antifolate methotrexate was encapsulated into these polymeric NPs as a model drug for in vitro and in vivo evaluation in HeLa cells. These MTX-loaded PEG−PGA-based nanoparticles were of particular interest as they exhibited specific tumor uptake and prominent antitumor activity to validate their potential as new nanomedicines. 3.1. Synthesis of Block Copolymers. Biodegradable poly glutamic acid (PGA) and biocompatible PEG were conjugated by the Michael addition reaction of the terminal amine functionality of PGA and acrylate moieties of PEG at room temperature (Scheme 1). The PGA-b-PEG diblock copolymer (EGFA) was synthesized by the conjugate addition of terminal amine of PGA to the acrylate moiety of PEG mono acrylate, while the folate functionalization was incorporated into it by the esterification of the terminal hydroxyl group of PEG moiety with carboxylic group of folic acid via Steglich esterification protocol. The addition reaction used here was a simple yet versatile approach to synthesize the PEG-b-poly(glutamic acid) block copolymers as compared to the widely reported ringopening polymerization of N-carboxy anhydride of L-glutamic acid initiated by amine-terminated PEG where the chain length of the polypeptide can vary from reaction to reaction and will have to be optimized for each polymerization reaction.41−43 The synthesized GEG triblock copolymer and the folic acid functionalized EGFA diblock copolymers were chemically characterized by NMR (1H, 13C), IR spectroscopy, GPC, and UV−vis and thermally analyzed by TGA and DSC (Figures S1 and S2). The data from these characterizations verified the successful synthesis of GEG and EGFA block copolymers. NP Formation-Characterization and Drug Loading. Nanoparticle formation of GEG and EGFA NPs as well as the physical incorporation of MTX into them was achieved by the nanoprecipitation method followed by dialysis. Both the block copolymers could self-assemble into core−shell nano scale particles in aqueous solution and displayed very low polydispersity of less than 0.15−0.16 with particle diameters of 105 and 90 nm, revealing a limited distribution of sizes and

3. RESULTS AND DISCUSSION Despite the exploitation of polymeric nanomaterials in biomedical applications, the biofunctionality of PEG-polypeptide-based NPs has not been explored as delivery vehicles for MTX, and their therapeutic properties have not been concurrently studied for cervical cancer applications. Herein we have designed a triblock copolymeric system based on F


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Figure 3. Physico-chemical characterization of (A) GEG NPs and (B) EGFA NPs: (i) particle size distribution, (ii) zeta potential, (iii) UV−vis overlay spectra of drug loaded and unloaded systems in PBS 7.4 along with solution images. Respective scanning electron microscopy (SEM) and the atomic force microscopy (AFM) images have been shown as the insets in graphs (i) and (ii).

3.2. In Vitro Drug Release and Kinetics. The behavior of the encapsulated drug in the physiological environment can be estimated by studying its in vitro release profile under simulated physiological conditions. The in vitro release profile of MTX was investigated by the dialysis−bag diffusion technique. In order to evaluate the pH-responsive release behavior of MTX from the triblock and diblock copolymer nanoparticles, the drug release was studied at three different pH values: pH 7.4 to simulate the blood pH, pH 5.5 to simulate the acidic pH of the tumor microenvironment, and pH 4.5 to simulate the late endosomal pH.45 The NPs that enter the in vivo system by the intravenous mode encounter a pH gradient from pH 7.4 in the blood circulation to a slightly lower pH of 7.1−6.5 in the tumor microenvironment and extra cellular matrix, which is lower than that of normal tissues. Once the NPs enter the tumor cells, the pH values further drop to pH 6.0−5.0 in the endosomes and pH 5.0−4.0 in the lysosomal compartments of tumor cells while the redox potential in the tumor cytosol presents a further reducing environment in comparison to the extracellular components.46,47 It is notable that a grossly sustained release profile is presented for both the MTX-loaded GEG and EGFA NPs at all the three pH conditions studied (Figure 4). The drug release process seems to get significantly accelerated when the pH of

validating a unimodal distribution (Figure 3). The GEG and EGFA NPs showed zeta potential in the range of −12.8 and −17.5 mV respectively, indicating the stability of the particles in solution (Figure 3). The stability of the drug-loaded nanoparticles upon aging and dilution was further verified by monitoring their particle sizes over a period of 7 days and variation of concentration from 2 to 0.5 mg/mL (Figure S4). No significant change was observed in the hydrodynamic diameters either on dilution or upon aging, thus indicating the desirable stability of these nanoparticles. These dimensional characteristics could prove beneficial in enhancing the EPR effect and improving the drug accumulation of these prepared nanoparticles at the tumor site.44 The morphology of drug-loaded polymeric NPs was visualized by SEM and AFM measurements which indicated that the shape of these NPs was spherical for M-GEG and MEGFA respectively. The loading of MTX into these nanoparticles was verified with the help of UV−vis spectroscopy, and the drug loading content was calculated from these UV results. The percentage entrapment efficiency (% EE) was found to be 93.7% and 95.62% while the drug-loading content (% DL) was found to be 38.2% and 46.88% for M-GEG and MEGFA NPs, respectively. G


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Figure 4. Time-dependent release of MTX from (A) M-GEG NPs and M-EGFA NPs at three different pH values (7.4, 5.5, and 4.5) compared to free MTX at 37 °C. Inset shows the best-fit release kinetics models (Krosmeyer-Peppas) for both the NPs. The values represent the mean ± SD (n = 3).

predicted with the help of the Korsmeyer-Peppas diffusion model, which gave the best fit to the experimental drug release data (Figure 4 inset).49 The value of the diffusional exponent, n, derived from this fit predicted a Fickian diffusion of the drug from GEG and EGFA NPs at a pH of 7.4 (n = 0.437 and 0.382, respectively), indicating a diffusion-controlled drug release due to the usual molecular diffusion of the drug owing to a chemical potential gradient.50 However, at pH values of 5.5 and 4.5, it predicted a non-Fickian diffusion through anomalous transport with n values increasing with a decrease in pH from 5.5 (n = 0.708 and 0.629, respectively) to 4.5 (n = 0.844 and 0.719, respectively).51 Several simultaneous processes are considered to take place in this drug transport mechanism involving stresses and erosion as well as disentanglement of the polymer chains, while in the previous case the approximation is based on the assumption that the polymer chains are impermeable segments immersed in a solution.52 Purely polymer relaxation is believed to be the release mechanism for n values > 0.85, where a case-II transport is proposed as the mechanism of release. Similar analysis of release kinetics has been done by Peppas et al. for block copolymeric systems.53 These results clearly indicated that pH had an effect on the drug release mechanism, while analysis of these release data gave us a fair understanding of the release mechanism itself.

the release medium is reduced from 7.4 to lower pH values. This may be attributed to the increased ionization of MTX as the pH of the medium approaches closer to the pKa values of the drug (5.58, 4.3, and 3.08 at 37 °C) which would in turn result in increased solubility of the drug in the outside medium. This accelerating drug release with decreasing pH of the surrounding medium would be a highly desirable feature of a drug-loaded nanovehicle. Release Kinetics. The mechanism of release of MTX from the MTX-loaded nanoparticles could be elucidated through mathematical modeling of controlled release systems. Therefore, the data obtained from the above in vitro drug release studies was modeled to various kinetic equations, vis-a-vis the zero-order, first-order, Higuchi, and Korsmeyer-Peppas (Power Law) models, to better understand the underlying mechanism of release. The zero- and first-order equations did not show any correlation with the data, indicating that the drug release was independent of time and concentration (Figure S7). This led us to believe that the geometry and shape of the NPs, their aspect ratio, and particle size distribution have a role to play in governing the drug release behavior.48 Different release mechanisms have been proposed for different geometries such as slabs, spheres, cylinders, discs (tablets), etc. The spherical shape of these NPs was validated from the SEM images after which the release mechanism of MTX could be H


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Figure 5. Hemolysis plot of (A) M-EGFA NPs and M-GEG NPs (@ 5 mg/mL). p < 0.01; student’s t test; n = 3 experimental repeats per treatment. (B and C) show actual images at 1 h and 4 h incubation, respectively (negative control is phosphate buffer solution (PBS); positive control is 2% (v/ v) Triton X-100).

phenotypic changes and genetic drift.55 Another advantage is that HeLa cells overexpress FA receptors on their surface to a larger extent, which can be leveraged to target folate-capped nanoparticles to these cells, and thus obtain enhanced specificity toward cervical cancer. In Vitro Cytotoxicity and Cellular Uptake Analysis. The sulforhodamine B assay was used to verify the anticancer activity of the MTX-loaded polymeric nanoparticles, owing to its practical advantages over tetrazolium assays vis-a-vis sensitivity over MTT assay under stress conditions. The SRB protein assay provides higher accuracy as it stains only the recently lysed cells baring the cell debris, and thus drug sensitivity data are not affected unlike other prevalent methods.56 Viability of HeLa cells was examined after incubation with a series of free MTX and MTX-loaded nanoparticles (M-GEG and M-EGFA NPs) for up to 48 h. The discontinuous form of the protocol did not give very encouraging results initially. Therefore, the continuous form was applied to test, and time points of 4 h, 8 h, 24 h, and 48 h were chosen keeping in mind the observations that had been made in uptake studies as well as release kinetics. A plot of percentage viability against the concentration revealed a time and dose-dependent cytotoxicity of the free drug MTX and the drug-loaded nanoparticles (Figure 6). Since MTX is a cytotoxic anticancer drug, the antiproliferative activity of HeLa cells could

3.3. Hemolysis Assay. Hemolysis is an important aspect that may limit the usage of biomaterials. These MTX-loaded polymeric nanoparticles are intended to have an intravenous mode of administration, thus making it imperative to evaluate their blood compatibility prior to injection. Therefore, a hemolysis assay was performed on M-GEG and M-EGFA NPs in order to quantify the RBCs membrane damaging properties of these nanoparticles. No apparent hemolysis was observed in either of the samples, over a wide range of concentrations 0.1− 5 mg/mL, over a period of 1 h to 24 h (Figure 5). Around 24 h, MTX started showing greater hemolytic activity, while the hemolytic effect of the nanoparticles remained insignificant (hemolysis percentage < 2%) up to 24 h. The result of this study confirmed that exposure to these MTX-loaded nanoparticles would not result in rupture of red blood cells, and they could be employed as safe drug delivery carriers without concerns of hemotoxicity. 3.4. In Vitro Evaluation in HeLa Cells. Human cancer cell lines like Hela, SiHa, C33A, Ca Ski, ME-180, etc. can serve as fundamental models to evaluate the efficacy of the therapeutic agents in cancer theranostics, both by means of in vitro studies (as monolayer cultures) as well as for in vivo studies (as xenografts in mice).54 HeLa, the first cultured human cell line for cervical carcinoma, has been derived from a high-grade, high-stage aggressive cancer that had acquired the necessary I


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based GEG NPs, however, showed lower cytotoxicity both at 24 h and 48 h than the native drug. A possible reason for this could be the slower release of MTX from these GEG NPs as seen in the previous section. The FA-decorated EGFA NPs nonetheless showed similar cytotoxicity as native MTX toward HeLa cells after 24 h incubation. This could be explained by the higher cellular uptake of EGFA through FR-receptor-mediated endocytosis, resulting in higher intracellular concentration of MTX, thus eliciting similar cytotoxicity as the native drug. This was also corroborated by the in vitro drug release results. At 48 h, a noticeable decrease in cell viability was observed in case of M-EGFA, resulting in an almost 30% increase in cytotoxicity as compared to that observed at 24 h incubation. This indicated the higher cell-kill effect of EGFA NPs as compared to the GEG NPs and native drug. Since these cytotoxic effects were observed at a 20 μM concentration, this same concentration was further used to evaluate their cell internalization by fluorescence microscopy. HeLa cells were incubated with rhodamine-labeled M-GEG and M-EGFA NPs in order to clearly visualize the cell internalization of these NPs both qualitatively and quantitatively. After a 4 h incubation, fluorescence could be detected mainly in the cytosol for all the three compounds and the intensity of the fluorescence signals continued to increase over the next 24 h. Even though M-Rh showed relatively higher fluorescence in the cytoplasm up to 24 h, its cell uptake and count of cells incubated with M-Rh alone seemed to decrease drastically as evident in the 48 h images (Figure 7). This could be attributed to the efflux of the native drug over the course of time by ABC membrane transporter proteins as also reported in the previous literature.57 The M-Rh fluorescence in cells treated with M-Rh-GEG NPs was relatively lower than in the case of free drug and M-Rh-EGFA, most likely due to the slower cellular uptake of M-Rh-GEG NPs that were stealthed by a dense layer of PEG shells and slowed the release of drugs from the compact triblock copolymer NPs. Despite this, it was

Figure 6. In vitro cytotoxicity of MTX, M- GEG NPs, and M-EGFA NPs toward HeLa cells after 24 and 48 h exposure at different MTX concentrations (5, 10, 15, and 20 μM). The cell viability was evaluated by sulpho-rhodamine B (SRB assay). Data are presented as mean ± SD (n = 3) (p < 0.05, student’s t test, n = 3).

be ascribed to the anticancer activity of MTX because the NPs alone did not affect the growth of HeLa cells as seen in separate experiments (Figure S8, panels a and b). Lowest viability of cells was observed initially at 24 h in the presence of MTX alone than in the presence of MTX entrapped in nanoparticles, perhaps owing to the different mechanisms by which the free drug and the nanoparticles enter the normal/tumor cells. The small MTX molecule is easily taken up by the cells primarily by the reduced folate carrier (RFCs) (KD ∼ 4.3 mM) as well as by the folate receptors (FRs) (KD ∼ 200 nM) and can readily interact with the DHFR present in the cytosol, thus causing its cell kill effect. However, there happens to be a time lag between uptake of the drug-loaded NPs and the availability of the free drug in the cytosol after its release from the NPs and therefore a delayed cell-kill effect by drug-loaded NPs. At 48 h, however, not much decrease in the cell viability was observed on incubation with MTX alone, perhaps owing to the efflux of the drug from the cells over the course of time by the ABC membrane transporter proteins.21 The triblock copolymer

Figure 7. Cellular internalization of MTX, M-GEG NPs, and M-EGFA NPs in Hela cells after 4, 8, 24, and 72 h incubation, examined by fluorescence microscopy. The nuclei were stained blue with Hoechst dye, while the dye-tagged NPs showing green fluorescence accumulated in the cytoplasm as well as the nuclei in the case of M-EGFA NPs- treated cells. Scale bar = 10 μm. J


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Figure 8. Quantification of M-Rh fluorescence intensity in HeLa cells after 4 h, 24 h, 48 h, and 72 h incubation with MTX, M-GEG NPs, and MEGFA NPs (20 μM of MTX in each NPs) by FACS analysis (p 99%, and they were observed to be stable in serum up to 24 h (labeling efficiency >98%) (Figure S11). The 99mTc-labeled MTX and MTX-encapsulated GEG and EGFA NPs were intravenously injected into healthy rabbits to generate their blood clearance profile. A comparative evaluation revealed a gradual clearance of these nanoparticles from systemic circulation over a period of 24h, p.i., exhibiting a biexponential pattern. The blood circulation time of the MTXloaded nanoparticles was significantly extended as compared to that of free MTX, with a 1.4 times longer half-life (t1/2). The MEGFA NPs showed elimination half-lives of t1/2 (fast) of 48 min and t1/2 (slow) of 25 h, while M-GEG NPs showed a protracted clearance with t1/2 (fast) of 24 min and t1/2 (slow) of 15 h in contrast to the rather somewhat faster clearance of free MTX with t1/2 (fast) as 24 min and t1/2 (slow) as 15 h (Figure 10). The decreased clearance of the drug entrapped in the nanoparticles as compared to the free MTX may be explained by the enhanced in vivo circulation time of polymeric nanoparticles based on PEG-b-PGA nanosized vehicles, as also reported for other polymeric nanoparticles.61,62 99m Tc-GEG NPs and 99mTc-EGFA NPs were injected into HeLa xenografts in nude mice (in the tail vein) for both qualitative and quantitative estimation of the in vivo and ex vivo biodistribution of the 99mTc-labeled nanoparticles with respect to the tumor. The gamma imaging technique helped us to clearly trace the whole body radioactive uptake of the nanoparticles after predetermined time intervals of 1, 2, 4, and 24 h, respectively (Figure 10). The MTX-loaded nanoparticles exhibited a stronger radioactive uptake of 99m Tc-labeled nanoparticles at the tumor site. Maximum tumor uptake was observed after 4 h, p.i., over the period of incubation, as seen in Figure 10. Nevertheless, the M-EGFA NPs showed a reasonably higher uptake at the tumor site as compared to the M-GEG NPs. This higher accumulation in the case of the M-EGFA NPs as compared to the M-GEG NPs could be attributed to the tumor targeting effect of FA ligands to the FRs on the tumor cell surface, indicating a strong directional ability of surface biomarkers (folic acid) toward the tumor site. This was also in agreement with the in vitro qualitative and quantitative cellular localization studies. For further confirmation, the quantitative distribution of these nanoparticles in various tissues and organs was evaluated by measurement of the 99mTc radioactivity in the respective organs after the gamma imaging study. Figure 10 shows the ex vivo distribution of the 99mTc-labeled nanoparticles in all the major organs of the HeLa xenografts. The relatively high concentrations of GEG and EGFA in the kidney point toward their renal mode of excretion. These results supported by the relatively shorter systemic circulation of free MTX suggest that it may be more susceptible to capture by RES and account for its high uptake in the liver and spleen (Figure S12). Histology. The apoptotic effect of these NPs were verified by histology analysis of the H&E stained section of tumors explanted from MTX, M-GEG NPs, and M-EGFA NPs treated mice (Figure 11). Significant damage was visualized in these tumor sections with increasing time as compared to the control. In fact, the tumor cells of the control group showed more or less intact structures with more chromatin and binucleolates, suggesting that the tumor cells were in the rapid growth phase.63 In contrast, the tumor cells treated with MTX- and MTX-loaded NPs exhibited degenerative changes with the passage of time. For example, the images show that on the tenth day the tumor cells in the control group has increased in N


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4. CONCLUSION The success of nanotherapeutic delivery is a combined effect of three key mechanistic elements: (i) specific cellular binding, (ii) intracellular uptake of drug-carrying nanoparticles by the targeted cells, and (iii) controlled release of encapsulated drug molecules in an active form.66 The conveniently fabricated MTX-loaded PEG- and PGA-based nanoparticles prepared in this study have shown promising results to fulfill these basic requirements of a nanotherapeutic by the sustained and targeted drug release at the tumor site, owing to the significant effect of folic acid which acts as a homing device and directs the MTX-loaded nanoparticles toward the tumor. Of the prepared NPs, the MTX-loaded EGFA NPs exemplified targeted drug release at the tumor site while the MTX-loaded GEG NPs had the ability of sustained drug release and thus potential for longterm action on the tumor cells, while sparing toxicity to healthy tissues. Furthermore, high intracellular drug release resulting in superior antitumor activity could be ascribed to the EPR effect of polymer-based nanocarriers, prolonged systemic circulation, effective tumor distribution, sufficient cell internalization in a compounding manner, favorable intracellular drug release, and desired cell-kill by induction of apoptosis. Overall, these results indicate that these theranostic nanoparticles are not only effective delivery vehicles for tumor− specific delivery of anticancer therapeutics such as MTX (by intravenous injection) aimed at FR(+ve) tumors but also have an added feature of simultaneous diagnosis. Since many FR(+ve) tumors often show poor clinical prognosis, these NPs offer a viable therapeutic approach for their eventual translation to clinics.

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ASSOCIATED CONTENT

S Supporting Information *

The Supporting Information is available free of charge on the ACS Publications website at DOI: 10.1021/acsami.7b04982. Synthesis and characterization of block copolymers, preparation of polymeric nanoparticles and optimization of their particle size, release kinetics models, synthesis and chemical characterization of M-Rh, calibration curve of MTX, cytotoxicity study of GEG and EGFA NPs, and detailed radiolabeling procedure of nanoparticles with 99m Tc (PDF)

Figure 12. In vivo evaluation of antitumor efficacy of M-EGFA NPs, M-GEG NPs, and free MTX. (A) Representative images of HeLa tumors xenografted in nude mice models and (B) relative tumor volume. Error bars indicate mean ± SD (n = 5), p < 0.05; student’s t test. (C) Kaplan Meyer plot of survival rates of HeLa tumor bearing mice (6 mice per group).

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AUTHOR INFORMATION

Corresponding Authors

*E-mail: [email protected]; [email protected]. *Address: Division of Cyclotron and Radiopharmaceutical Sciences, Institute of Nuclear Medicine and Allied Sciences, Defence Research and Development Organization, Delhi110054, India. E-mail: [email protected].

An evaluation of the life-span of these Hela xenografts further indicated the MTX-induced toxicity as a reflection of the treatment efficiency of these MTX-loaded NPs. All the mice in the control group died within 30 days postinjection. In contrast, all the mice treated with M-GEG and M-EGFA NPs survived well during this duration of evaluation and were mobile and active. It was inferred that there was a substantial constructive increase in the lifetime/average survival time and quality of life of the NPs treated mice. As evidenced in Figure 12, these PGA−PEG-based nanoparticles can significantly improve the antitumor effectiveness of MTX by encapsulation and extend the life-span of tumor-bearing mice. No significant loss in body weight was observed in the cases of M-EGFA NPs and M-GEG NPs injected athymic tumor mice that could be associated with obvious side effects on other body tissues.

ORCID

Anil K. Mishra: 0000-0003-2523-9045 Author Contributions

S.D. synthesized the block copolymers and carried out in vivo evaluation of nanoparticles, Y.R. and A.B. performed the in vitro experiments, A.K.T. assisted in manuscript preparation and analysis, S.S. and A.K.M. helped in chemical characterization, D.K. planned all the experiments, conceptualized the manuscript, and supervised the entire project. All authors have approved the final version of the manuscript. O


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ACS Applied Materials & Interfaces Notes

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The authors declare no competing financial interest.

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ACKNOWLEDGMENTS We thank the Director, Institute of Nuclear Medicine and Allied Sciences, for providing necessary facilities and infrastructure for carrying out this research work. Fellowship provided to S.D. by CSIR, India, is duly acknowledged. We are thankful to SSPL, DRDO, and USIC facility of Delhi University for instrumentation support.

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