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Colloids and Surfaces B: Biointerfaces 122 (2014) 674–683

Contents lists available at ScienceDirect

Colloids and Surfaces B: Biointerfaces

jo ur nal ho me p ag e: www.elsev ier .com/ locate /co lsur fb

argeted delivery of doxorubicin into tumor cells via MMP-sensitiveEG hydrogel-coated magnetic iron oxide nanoparticles (MIONPs)

aner Nazli a, Gozde S. Demirerb, Yasemin Yara, H. Yagci Acara,c, Seda Kizilel a,b,∗

Koc University, Material Science and Engineering, Istanbul 34450, TurkeyKoc University, Chemical and Biological Engineering, Istanbul 34450, TurkeyKoc University, Chemistry Department, Istanbul 34450, Turkey

r t i c l e i n f o

rticle history:eceived 4 February 2014eceived in revised form 29 July 2014ccepted 31 July 2014vailable online 10 August 2014

eywords:

a b s t r a c t

Targeting tumors with nano-scale delivery systems shows promise to improve the therapeutic effectsof chemotherapeutic drugs. However, the limited specificity of current nano-scale systems for cancertissues prevents realization of their full clinical potential. Here, we demonstrate an effective approach tocreating as targeted nanocarriers for drug delivery: MIONPs coated with integrin-targeted and matrix-metalloproteinase (MMP)—sensitive PEG hydrogel scaffolds. The functional PEG hydrogel coating hasbeen designed for active loading as well as triggered intra-cellular release of the cancer therapeutic agent

ntegrin targetingrotease responsive PEG hydrogel coatingontrolled releaseoxorubicinagnetic iron oxide nanoparticles

argeted nanocarrier

doxorubicin (DOX). Our study demonstrated that coated nanocarriers could be taken into cancer cells11 times more efficiently than uncoated ones. Furthermore, confocal laser scanning microscopy imagesrevealed that these targeted nanocarriers could efficiently deliver and release DOX into the nuclei of HeLacells within 2 h. Coating MIONPs with multifunctional PEG hydrogel could be a promising alternative toexisting vehicles for targeted delivery of DOX into tumor tissue.

© 2014 Elsevier B.V. All rights reserved.

. Introduction

Chemotherapy is commonly employed to treat cancer patients1]. However, non-specificity of chemotherapeutic agents andesistance of cancer cells to drugs compromise the efficiency ofhemotherapy [1–4]. Previous studies have used various typesf drug-loaded nanoparticles, and a few of such systems areommercially available [5]. For clinical use, nanoparticles shouldemonstrate a couple of important properties including excellent

n vivo stability, high drug-loading capacity, and long circulationime in the bloodstream [4,6]. Responsiveness to local stimulusuch as pH, temperature, or protease would add additional ben-fits for drug delivery vehicle design. These properties can be usedo achieve specific accumulation of nanoparticles at the cancerites and/or to deliver therapeutic drug in the cancerous tissuen response to a stimulus [4]. MIONPs, which are superparamag-

etic, have demonstrated great potential for simultaneous imagingnd targeted delivery of therapeutic agents into tumor sites [7–9].IONPs may be coated with various materials. When polymers

∗ Corresponding author at: Koc University, Chemical and Biological Engineering,stanbul 34450, Turkey. Tel.: +90 212 338 1836; fax: +90 212 338 1548.

E-mail address: skizilel@ku.edu.tr (S. Kizilel).

ttp://dx.doi.org/10.1016/j.colsurfb.2014.07.049927-7765/© 2014 Elsevier B.V. All rights reserved.

are considered for coatings, end-grafting, direct adsorption, or self-assembly of linear polymers via hydrophobic interactions can beutilized [10–16]. Dextran has been widely used as a coating materialdue to its biocompatability and high affinity to iron oxide surfaces[17–20]. For example, a contrast agent based on MIONP was coatedwith carboxydextran to prevent aggregation of particles.[21–24]This agent, commercialized as Resovist® (Bayer Schering Pharma,Berlin-Wedding, Germany), was used for imaging liver lesions dur-ing the 1990s. However, the product did not meet all expectationsof magnetic particle imaging and was withdrawn by the companyat the end of 2008.

PEG, a well-studied polymer due to its biocompatible and non-toxic nature, has been conjugated to MIONPs to increase circulationtime in blood and to prevent agglomeration and degradation bymacrophages in blood [4,25]. Despite a PEG coating, previouslydeveloped MIONPs may still have low colloidal stability and lackstimuli responsiveness for enhanced drug release at the specific site[26,27]. Responsive design of the MIONP coating is also beneficial toaccomplish high contrast in medical imaging [28]. Hence, advancedstrategies are needed both to enhance therapeutic efficacy and

minimize side effects of cancer drugs. The on-demand deliveryof chemotherapeutics at a targeted site to improve therapeuticefficacy is crucial for translating research into clinical application[29–32]. To achieve this end goal, many studies have developed

es B: Biointerfaces 122 (2014) 674–683 675

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staddWioasGcai�wft(ofzmiSti

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Scheme 1. Illustration of intra-cellular delivery of cancer drug DOX into cancer cellsby targeted nanocarriers. RGDS ligand of the targeted nanocarriers binds to �v�3

C. Nazli et al. / Colloids and Surfac

rug-delivery systems using materials such as inorganic nanoparti-les [33,34], liposomes [35,36], biomolecules [37], micelles [38,39],arbon materials [40,41], and polymeric nanoparticles [42,43].

The major approach in enhancing specificity is designingrug-delivery vehicles that specifically recognize tumor cells. Asn example, targeting �v�3 integrin is a popular approach inumor-targeted drug-delivery with nanoparticles. Transmembraneeceptor �v�3 integrin has angiogenic, migration features, and isverexpressed in tumor cells [4]. �v�3 integrin initiates angiogene-is and regulates cell migration, which are crucial for gaining accesso nutrients and spreading of cancer cells [44].

Tumor tissues also secrete high amounts of proteolytic enzymes,uch as MMPs, which degrade the basement membrane and natu-al extracellular matrix (ECM), opening up more space for tumorrowth [44,45]. Elevated expression of MMP-2 and MMP-9 wasbserved in some malignant tumor tissues such as breast cancer46,47], colon cancer [48,49], brain tumors [50,51], and other malig-ancies [52]. For example, in a study by Lizasa et al. [53], lung canceratients had plasma MMP-9 concentrations of 71.0 ± 60.2 ng/ml,hereas in healthy control patients 36.3 ± 13.2 ng/ml was mea-

ured. In another study by Tutton et al. [54], significant increases inMP-2 and MMP-9 were observed in protein expression in colorec-

al cancer specimens compared to normal colonic tissue. MMP-2nd MMP-9 total protein concentrations in normal colonic tissueere measured as 4.2 ±0.4 ng/mg and 6.5 ± 3.5 ng/mg, respectively.

n the tumor environment, however, total protein concentrationsf MMP-2 and MMP-9 were increased to 180.4 ± 99.3 ng/ml and83.4 ± 73.9 ng/ml, respectively. Due to these high concentrationsf MMPs measured in tumor tissue, proteolytically degradableeptide sequences attracted significant attention, and subsequentesearch used MMP cleavable peptide conjugated drugs [55,56] andolymeric shells [57].

In this study, we designed a multifunctional MIONP that mayerve as a contrast agent in MRI, bind and carry chemotherapeu-ic drug, target overexpressed receptors on cancer cell surface,nd be sensitive to MMP. These MIONPs have proteolyticallyegradable PEG hydrogel coatings with integrin-binding RGDSomains, and are loaded with doxorubicin (DOX; Scheme 1).e hypothesized that MMP-sensitive domains within the coat-

ng can be effectively cleaved as a result of high concentrationf MMPs in the cancer cell environment, and this coating mayllow for the release of loaded cancer drug DOX into cancer cellspecifically. We synthesized acrylate-PEG-RGDS and acrylate-PEG-GGPQG↓IWGQGK–PEG-acrylate conjugates, and included theseonjugates in the PEG diacrylate prepolymer solution. We call thiscrylate mixture a functional prepolymer solution. The proteolyt-cally degradable sequence, GGGPQG↓IWGQGK, is derived from1(I) collagen and can be degraded by several MMPs [58]. Next,e suspended MIONPs with immobilized photoinitiator in the

unctional prepolymer solution, followed by surface-initiated pho-opolymerization, to obtain a functional coating around MIONPsScheme 2). We investigated the synthesis and characterizationf functional PEG hydrogel-coated MIONPs using Fourier Trans-orm Infrared (FT-IR) spectroscopy, dynamic light scattering (DLS),eta potential, and scanning electron microscopy (SEM). Confocalicroscope images demonstrated that intra-cellular DOX release

nto cancer cells could only be achieved with the targeted MIONPs.ignificantly improved internalization of MIONPs and DOX reten-ion were accompanied by decreased HeLa cell viability within 48 hn the in vitro studies.

The method presented here extends our previous work onhe PEG hydrogel-coated MIONPs via surface-initiated photopoly-

erization [11]. The technique presented here, which coatsIONPs with a multifunctional PEG hydrogel, has been used

or receptor-targeted drug-delivery into cancer cells and specificnloading of the chemotherapeutic agent. This coating can have a

integrin, and nanocarrier can internalize into cancer cells via receptor-mediatedendocytosis. After degradation of MMP-sensitive domains in PEG hydrogel coatingof targeted nanocarriers, loaded DOX is released into cytoplasm of cancer cells.

variety of applications when multiple functionality of MIONP isdesirable.

2. Materials and methods

2.1. Materials

Eosin Y (91%), 1-vinyl-2-pyrrolidinone (VP; 99%), poly (eth-ylene glycol) diacrylate (PEGDA) (MW: 575 Da), Woodward’sreagent K (WRK) (2-ethyl-5-phenylisox-azolium-3′-sulfonate,95%), phosphate-buffered saline (PBS), Triton® X-100,hydroxyethyl piperazineethanesulfonic acid (HEPES), andtrypsin-ethylenediaminetetraacetic acid were obtained fromSigma–Aldrich (St. Louis, MO). FeCl3·6H2O and FeCl2·4H2O wereobtained from Fluka (Sigma–Aldrich, Taufkirchen, Germany).Ammonium hydroxide (26% NH3 in water, w/w) was obtainedfrom Riedel-de Haën (Hanover, Germany). APTMS was obtainedfrom Gelest, Inc. (Morrisville, PA). Triethanolamine (TEOA) (99.5%),sodium bicarbonate, and 37% formaldehyde were purchasedfrom Merck (Whitehouse Station, NJ). Hydrochloric acid (HCl)and sodium hydroxide (NaOH) were purchased from Riedel-deHaën. RGDS (MW: 433 Da) and GGGPQG↓IWGQGK (MW: 1141 Da)peptides were purchased from Elim Biopharmaceuticals, Inc. (Hay-ward, CA). Acrylate PEG N-hydroxysuccinimide (acr-PEG-NHS)(MW: 3.4 kDa) was obtained from Nektar (Huntsville, AL). Potas-sium ferrocyanide was obtained from Labkim (Istanbul, Turkey).Doxorubicin (DOX) was purchased from MP Biomedicals (Solon,OH). DAPI (4′,6-Diamidino-2-Phenylindole, Dihydrochloride) waspurchased from Invitrogen Co. (Carlsbad, CA). Collagenase type1 was purchased from Worthington Biochemical Corporation(Lakewood, NJ). Dulbecco’s modified Eagle medium (liquid, with

4.5 g/L d-glucose and sodium pyruvate) (DMEM) and l-glutaminewere purchased form Gibco (Life Technologies Products, GrandIsland, NY). Penicillin/streptomycin was obtained from InvitrogenCo (Carlsbad, CA). Donor horse serum was obtained from Biological

676 C. Nazli et al. / Colloids and Surfaces B: Biointerfaces 122 (2014) 674–683

S l PEG ha

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2

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cheme 2. (A) Schematic representation of MIONPs coating within multifunctionacryl conjugates.

ndustries (Rainbow Scientific, Inc., Windsor, CT). The Amiconltra centrifugal filter device (NMWL: 3000) was purchased fromMD Millipore. The dialysis cassette (MW cutoff: 3.5 and 20 kDa)as obtained from Pierce (Micro BCA Protein Assay; Thermo

isher Scientific Inc., Rockford, IL). Teflon syringe filters (0.2 �more size) were purchased from Nalgene (Rochester, NY). Petriishes were obtained from Nunc (Roskilde, Denmark). Cell culturelates (96- and 24-well) were purchased from Greiner Bio-OnembH (Solingen, Germany). Ninety-six-well LUMITRAC® 600hite immunology plates were purchased from USA Scientific, Inc.

Ocala, FL).

.2. Synthesis and coating of MIONPs

MIONPs were coated above within PEG hydrogel via surface-nitiated photopolymerization as previously described [11,59,60].

n brief, EY was covalently bound to the surface amine groupsf APTMS-coated MIONPs (Scheme 2). EY bound MIONPs wereuspended in prepolymer solution composed of 6.25% PEGDAw/v) (MW: 575 Da), TEOA (141 mM), and VP (12.5 mM) in

ydrogel. (B) Synthesis of acryl-PEG-RGDS and acryl-PEG-GGGPQG↓IWGQGK-PEG-

deionized water. Acr-PEG-RGDS and acr-PEG-GGGPQG↓IWGQGK-PEG-acr (acr-PEG-DEG-PEG-acr) conjugates were also synthesizedand characterized as described above (Scheme 2). To purify thedegradable conjugate, a dialysis membrane cassette with molec-ular weight cut off (MWCO) of 7 kDa was used to obtain conjugatewith diacrylate functionality. To obtain a PEG hydrogel coating withan integrin binding RGDS, acr-PEG-RGDS (1.48 × 10−5 M) and acr-PEG-DEG-PEG-acr (1.0 × 10−5 M) were added into the prepolymersolution before the photopolymerization reaction. The mixtureswere exposed to 514 nm green light at a flux of 2.5 mW/cm2 usingan argon ion laser (Coherent Inc., Santa Clara, CA) for 60 s. Toremove unreacted components, the polymerization mixture waswashed using the Amicon Ultra centrifugal filter, providing cleanMIONPs with polymerized multifunctional PEG hydrogel coating.These MIONPs were stored at 4 ◦C in deionized water.

2.3. Characterization of MIONPs

APTMS-coated MIONPs (MIONPs), only PEG hydrogel coatedMIONPs (non-targeted nanocarriers), and multifunctional PEG

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ydrogel-coated MIONPs (targeted nanocarriers) were character-zed in terms of structure, size, morphology, and surface charge.T-IR spectroscopy (Nicolet iS10 FT-IR spectrometer; Fisher Scien-ific, Vienna, Austria) was used to confirm the presence of PEGDAhains and to characterize functional groups of the polymericoating around nanoparticles. Core diameter, and morphology ofon-targeted and targeted nanocarriers were analyzed via SEMZeiss Ultra Plus, Bruker; Optronik Ltd. Sti, Ankara, Turkey). Theydrodynamic diameter and zeta potential were characterizedith a Zetasizer instrument at room temperature (Malvern Zeta-

izer nano S ZEN 1600, wavelength 633 nm; Malvern Instruments,alvern, UK).

.4. In vitro degradation kinetics of targeted nanocarriers andelease of DOX from targeted nanocarriers

Targeted nanocarriers were suspended in HEPES buffer (0.1 MEPES + 0.15 M NaCl + 5 mM CaCl2 at pH 7.2). Targeted nanocarriers

0.1 mg/mL) were suspended in HEPES buffer (0.1 M HEPES + 0.15 MaCl + 5 mM CaCl2 at pH 7.2) in the presence and absence of col-

agenase type I enzyme (0.1 mg/mL at 37 ◦C) for 4 days. Duringhis treatment, hydrodynamic diameters of nanocarriers were mea-ured at specific time points within the first 72 h.

The aqueous suspensions of non-targeted and targeted nanocar-iers (1 mg/mL) were suspended separately in an aqueous solutionf DOX (0.1 mg/mL) at a nanocarrier-to-DOX weight ratio of 10.ach mixture was sonicated for 3 h and incubated for 48 h at roomemperature. To remove unloaded DOX, samples were rinsed threeimes with Amicon Ultra centrifugal filter. Concentration of DOXn the eluted solution was determined with the absorbance of freeOX using NanoDrop 1000 spectrophotometer (Fisher Scientific,K) at 490 nm. The percentage of DOX loading efficiency (DLE) wasalculated using the following equation:

DLE = amount of DOX in coated nanoparticlesamount of DOX in buffer medium

× 100

DOX-loaded targeted nanocarriers were suspended in HEPESuffer in the presence and absence of collagenase type 1 enzyme0.1 mg/mL), separately. Next, dialysis was carried out using dial-sis cassette membranes (MWCO = 20 kDa) in HEPES buffer in theresence and absence of collagenase type 1 enzyme at 37 ◦C, and atpecific time points 1 mL solution was withdrawn from the bufferedium and replaced with 1 mL fresh buffer. Concentration of DOX

n the withdrawn medium was measured at each time point.

.5. Evaluation of cytotoxicity of DOX-loaded nanocarriers

Cytotoxicity assays were carried out with HeLa cells and mouseT3 fibroblast cells to investigate the influence of non-targeted andargeted nanocarriers on the viability of cancer and healthy cellines. HeLa cells (5 × 104 cells/mL in a 96-well plate) were culturedor 24 h in DMEM (pH 7.4, with 10% donor horse serum (DHS), 1%enicillin/streptomycin, and l-glutamine). Suspensions with non-argeted and targeted nanocarriers in HEPES buffer were added toeLa cells and mouse 3T3 fibroblast cell culture mediums, sepa-

ately. Viabilities were measured after 24, 48 and 72 h of incubationsing CellTiter-Glo Luminescent viability assay (Promega, Madison,I) with Fluoroskan Ascent luminescence reader (Fisher Scientific,

ienna, Austria). DOX concentrations of 1, 5, and 10 �g/mL wereested for each nanocarrier group.

.6. Internalization of DOX-free nanocarriers

Internalization of bare MIONPs, non-targeted and targetedanocarriers by HeLa cells were investigated via Prussian bluetaining (PBS). HeLa cells (5 × 104 cells/mL) were cultured for 24 h

iointerfaces 122 (2014) 674–683 677

in 24-well plates on glass coverslips. After the attachment of HeLacells to coverslips, nanoparticles at a concentration of 0.1 mg/mLin HEPES buffer were added to each well. At the end of 24 h, cellsattached to coverslips were washed three times with PBS and fixedwith 4% formaldehyde in PBS for 20 min at room temperature. Cellswere incubated with 0.1% Triton X-100 in PBS for 5 min and thenrinsed with PBS. Fixed HeLa cells were stained with 5% potassiumferrocyanide in 10% HCl for 15 min. Prussian blue stained cells werevisualized via a camera equipped with optical microscopy (EclipseNi-U Optical Microscope, Nikon, Istanbul, Turkey).

Cellular uptake of nanocarriers was also characterized viainductively coupled plasma optical emission spectrophotometer(ICP-OES) (Spectro Genesis; Spectro Analytical Instruments GmbH,Kleve, Germany). HeLa cells were grown as described previously.Next, MIONPs, non-targeted and targeted nanocarriers were trans-ferred to cell culture medium at a concentration of 0.1 mg/mL after24 h of incubation. As a control group, HEPES buffer was added intocell culture medium without nanocarriers. Following 24 h of incu-bation, cells in all groups were washed three times with PBS. Next,cells were lysed with 3 M HCl + 0.25% Triton X-100 for 2 h, then thecell lysates were incubated in a mixture of 65% nitric acid (0.25 mL)and 30% hydrogen peroxide (0.25 mL) for 24 h at 90 ◦C, and ICP-OESanalysis was applied for all samples. To calculate the amount ofiron (Fe) per cell in each sample, two groups of samples were used,where cell numbers were calculated with CellTiter-Glo Lumines-cent viability assay. Reported values were averages of at least threeexperiments.

2.7. Internalization of DOX-loaded nanocarriers

HeLa cells were seeded on coverslips in 24-well plates at adensity of 5 × 104 cells/mL for 24 h as described above. DOX con-centration of 10 �g/mL was loaded into nanocarriers, and thesenanocarriers were added to cell culture medium. After 2 h of incu-bation, treated cells were washed with PBS three times. Next, cellnucleus was stained with DAPI, according to the standard proto-col provided by the supplier. Cellular internalization of DOX wasvisualized by camera-equipped laser scanning confocal microscopy(Eclipse 90-i, Nikon). To quantify the Fe and DOX amount per cellfor each sample, ICP-OES and NanoDrop 1000 spectrophotometerwere used, respectively.

2.8. Statistical analysis

The results of all data sets were evaluated using one-way anal-ysis of variance. The results are shown as the mean value (plus orminus standard deviation) of the triplicate samples unless statedotherwise. Variances between datasets were considered statisti-cally significant for P-values less than 0.05.

3. Results and discussion

3.1. Nanocarrier synthesis

PEG hydrogel coating on an eosin-immobilized MIONP sur-face was conducted via surface-initiated photopolymerization asdescribed previously. Functional PEG-MIONPs were coated withPEG hydrogel prepolymer solution that included acr-PEG-RGDSand acr-PEG-GGGPQG↓IWGQGK-PEG-acr conjugates. The pres-ence of PEG hydrogel, RGDS and GGGPQG↓IWGQGK peptidesaround MIONP structure were confirmed with FT-IR (Fig. S1). FT-IR bands at 1728 cm−1 indicated C O stretching of PEG acrylate.

In the FT-IR spectra of non-targeted nanocarriers, C C stretch-ing bands of acrylate groups at 1640 cm−1 disappeared, whichindicated the reaction of the double bonds during photopoly-merization reaction. After the addition of acr-PEG-RGDS and

678 C. Nazli et al. / Colloids and Surfaces B: B

F(

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o

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ig. 1. Scanning electron microscope (SEM) image of (A) non-targeted nanocarriers,B) targeted nanocarriers (scale bar:100 nm).

cr-PEG-GGGPQG↓IWGQGK-PEG-acr conjugates into the prepoly-er, a new peak at 1670 cm−1appeared, which corresponded to the

mide I (C O stretching) peak coming from the peptide sequence.ll of these features observed in the FT-IR confirmed the presencef functional PEG hydrogel around MIONPs and the absence of anyesidual acrylate (Fig. S1).

Supplementary Fig. S1 related to this article can be found, in thenline version, at http://dx.doi.org/10.1016/j.colsurfb.2014.07.049.

The size and surface charge of nanocarriers are importanthysicochemical parameters for designing targeted drug-deliveryehicles. Size and morphology of nanocarriers before and afterydrogel coating were characterized using SEM and DLS. Fig. 1emonstrated that the particles have spherical and ellipsoid shapes.llipsoidal shapes most probably occurred as a result of fusing ofwo or more MIONPs during polymerization reaction. SEM resultsemonstrated that non-targeted nanocarriers had an averageiameter of 28.5 ± 4.8 nm, where this value was 186.5 ± 15.8 nmor targeted nanocarriers (Fig. 1). These results were similar to theydrodynamic diameters measured with dynamic light scatteringDLS): 25.3 ± 2.9 and 209.9 ± 24.3 nm for non-targeted and targetedanocarriers, respectively (Table 1). The diameter of MIONPs withnly APTMS groups on the surface was 7.3 ± 0.7 nm. Larger diam-ters were measured with PEG hydrogel coating. This finding wasost probably due to the formation and swelling of PEG hydrogel

round MIONPs. Even larger sizes measured for targeted nanocar-iers might be due to the longer PEG chains used for RGDS andegradable peptide conjugates, which might have promoted accessf integrin receptors and MMP to their ligands.

Tumor blood vessels consist of a high number of reproducingndothelial cells, where the absence of pericyte and abnormal base-

ent membrane formation results in higher vascular permeability

4]. The size of endothelial pores varies from 10 to 1000 nm, whicheans that theoretically it may be possible for nanocarriers within

able 1ydrodynamic diameter and zeta potential values of bare MIONPs, non-targetednd targeted nanocarriers. Data are mean ± standard deviation (SD) (n = 3).

Sample name Hydrodynamic size Zeta potential

Bare MIONPs 7.3 ± 0.7 14 ± 1.49Non-targeted nanocarriers 25.3 ± 2.9 −4.18 ± 0.45Targeted nanocarriers 209.9 ± 24.3 −3.37 ± 0.28

iointerfaces 122 (2014) 674–683

the size range of 20–200 nm to enter and localize at the intersti-tial space in the tumor environment [4]. In addition, surface chargeof nanocarriers can influence cellular uptake by tumor cells [61].Macrophages can easily eliminate cationic and highly negativelycharged nanoparticles due to non-specific internalization [61]. Zetapotential values for MIONPs, non-targeted and targeted nanocar-riers varied from −4.18 ± 0.45 mV to 14 ± 1.49 mV in HEPES buffer(Table 1). Targeted nanocarriers had slightly negative zeta potential(−3.37 ± 0.27 mV) due to the lone pair electrons of oxygen atomsin PEG chains (Table 1). This negative zeta potential may preventnon-specific internalization of nanocarriers by macrophages duringcirculation in blood because non-specific uptake can be minimizedwithin the zeta potentials range of −9 to −3 mV [61]. This valueof zeta potential may also facilitate deep penetration and inter-nalization of nanocarriers into the tumor site specifically [62,63] ifnanocarriers can be tagged with tumor-specific molecules and/ordirected to the tumor site magnetically as in the case of MIONPs[64].

3.2. In vitro evaluation of MMP-sensitive degradation of PEGhydrogel coating and DOX release

Degradation profiles for MMP-sensitive coating of targetednanocarriers were investigated via hydrodynamic size measure-ments using DLS. Targeted nanocarriers were incubated in HEPESbuffers with collagenase type 1 enzyme at 37 ◦C, and hydro-dynamic diameter was measured at different time points for 3 days.Targeted nanocarriers demonstrated a sharp decrease in hydro-dynamic diameter from 229.4 ± 12.86 nm to 59 ± 3.25 nm after oneday incubation in collagenase type 1 containing buffer. Incubationof nanocarriers for three days in collagenase buffer led to furtherdecrease in hydrodynamic size down to 30.86 ± 5.03 nm. Targetednanocarriers incubated in the absence of collagenase type 1 enzymehad slightly increased hydrodynamic diameters during the incuba-tion period due to swelling of the PEG hydrogel coating (Fig. 2A).These results confirmed that targeted nanocarriers were sensitiveto collagenase type 1 enzyme, and were degraded only in its pres-ence.

DOX (DOX.HCl) was loaded into non-targeted and targetednanocarriers in HEPES buffer to examine drug release profilesin the presence and absence of collagenase type 1 enzyme at37 ◦C. The interaction of positively charged DOX with negativelycharged materials such as 2,3-dimethylmaleic anhydride nanogels[65] and Laponite® nanodisks [66] were previously demonstrated.Drug-loading efficiency for nanocarriers in both groups was high,where 97.3% and 98.3% of DOX could be loaded into non-targetedand targeted nanocarriers, respectively. High loading efficiency ofnanocarriers may be attributed to the interaction of the slightlynegative PEG coating with the positively charged DOX. Nanocar-riers (10.2 �g/mL) contained 1 �g/mL DOX when DOX was loadedinto non-targeted nanocarriers. As shown in Fig. 2B, DOX releasefrom targeted nanocarriers improved in the presence of collagen-ase enzyme. After 10 h, release of DOX from targeted nanocarriers inthe presence of collagenase slightly exceeded the release observedin the absence of collagenase. Major differences in release rateswere observed after 3 days and at the end of 4 days incubation:60% of DOX was released from nanocarriers in the presence ofcollagenase, while 36% of loaded DOX was released from nanocar-riers in the absence of collagenase. The higher release of DOXobtained in the presence of collagenase was attributed to thespecific degradation of the coating and disassembling of the aggre-gates in response to collagenase, and hence faster diffusion of the

drug from nanocarriers. Even though the size of the coating wasdecreased from 229.4 ± 12.86 nm to 59 ± 3.25 nm at the end of day1, DOX release continued for 4 days. Despite the high collagenaseenzyme used for the degradation of the carrier (0.1 mg/mL), the

C. Nazli et al. / Colloids and Surfaces B: B

Fig. 2. (A) Degradation kinetics of targeted nanocarriers in the presence and absenceof 0.1 mg/mL of collagenase type 1. (B) DOX release versus time from DOX-loadedtargeted nanocarriers in the presence and absence of collagenase type 1. Filled redtriangle: DOX release from DOX-loaded targeted nanocarriers in HEPES buffer, filledblue squares: DOX-loaded targeted nanocarriers in HEPES buffer in the presence of0.1 mg/mL of collagenase type 1. Data are mean ± SD (n = 3). (For interpretation oftt

rtannncwtahlc

3

tinD(l

he references to color in figure legend, the reader is referred to the web version ofhe article.)

elease of DOX from the carrier in the presence of collagenase is 1.6imes greater than the release observed in the absence of collagen-se. This limited release could have occurred due to the presence ofondegradable PEGDA monomer used in the prepolymer. That is,o matter how high the collagenase concentration was, we wouldot observe complete degradation of the coating. The other reasonould be associated with the presence of crosslink density gradientsithin the coating and high crosslink densities at points closer to

he nanoparticle surface. This crosslink density gradient occurs as result of surface-initiated photopolymerization technique usedere, and might have resulted in higher drug-loading at stiffer

ocations close to the nanoparticle surface within the membraneoating [67].

.3. In vitro cytotoxicity and cellular uptake of nanocarriers

Cytotoxicity and tumor targeting properties of non-targeted andargeted nanocarriers were investigated with HeLa cells. Cells werencubated with free-DOX, DOX-free nanocarriers and DOX-loaded

anocarriers. Cells were treated with 1, 5, and 10 �g/mL of freeOX or an equivalent dose of DOX-loaded nanocarriers for 3 days

Fig. 3). As shown in Fig. 3A, after 24 h of drug exposure, DOX-oaded non-targeted or targeted nanocarriers reduced the viability

iointerfaces 122 (2014) 674–683 679

of cells slightly, at the highest drug concentration of 10 �g/mL. After48 and 72 h, cells treated with DOX-loaded targeted nanocarriershad viability of about 10%, whereas the viability of DOX-loadednon-targeted nanocarriers remained above 80% at the highest DOXconcentration of 10 �g/mL (Fig. 3B). After 3 days incubation, HeLacells treated with DOX-loaded targeted nanocarriers had decreasedviability of 60% and 50% at 1 and 5 �g/mL DOX concentrations,respectively (Fig. 3C). Interestingly, free DOX at 1 and 5 �g/mL didnot reduce the viability of cells below 80%. These results demon-strated that DOX-loaded targeted nanocarriers showed dose andtime dependent activity and were efficient to decrease the via-bility of HeLa cells. This decreased viability could be explainedby the mechanism of nanocarrier internalization by endocytosisand therefore better localization of the drug in HeLa cells. Tar-geted nanocarriers were most probably taken into HeLa cells viareceptor-specific endocytosis, hence cells did not have sufficienttime to develop resistance toward the cancer drug. However, freeDOX in the medium allows faster recognition by cancer cells, andhence more time for cancer cells to develop resistance toward thedrug. As a control experiment, HeLa cells was treated with nanocar-riers without DOX. Results showed that metabolism of HeLa cellswas not significantly influenced by DOX-free nanocarriers at thesame nanocarrier concentration, which indicated cytocompatibil-ity of the PEG-coated MIONPs (Fig. S2).

Supplementary Fig. S2 related to this article can be found, in theonline version, at http://dx.doi.org/10.1016/j.colsurfb.2014.07.049.

The toxicity of DOX-loaded targeted nanocarriers was alsotested against healthy cell lines from mouse fibroblast cells. Resultsshowed that free DOX samples at 5 and 10 �g/mL concentrationsnegatively influenced healthy cell survival (Fig. S3). When DOX-loaded non-targeted or targeted nanocarriers were used, fibroblastcell viabilities of 90% and 60% were measured at the highest DOXconcentration, respectively (Fig. S3). This viability result could beexplained by the differences in the interaction mechanism of DOXand fibroblasts when DOX was introduced freely or was loadedwithin nanocarriers. Cell internalization was not as efficient for thecase of fibroblasts as it was for HeLa cells, due to the fewer num-ber of integrin receptors on fibroblasts. Also, collagenase enzymeconcentration is not high around fibroblasts; a high concentrationis necessary for degradation of coating and hence release of DOXfrom nanocarriers. Free DOX could easily interact with healthy cellsand can negatively influence cell survival, while DOX loaded withinnanocarriers may not interact with cells quickly. These results sug-gest that targeted nanocarriers would be efficient to minimizeinteraction of a therapeutic drug with healthy cells, and that highinternalization of targeted nanocarriers into tumor cells followedby minimized cancer cell survival due to targeted drug-deliverywould be achieved with this system. However, with the use of non-targeted systems such as MMP-sensitive peptide conjugated DOX,it may be challenging to protect healthy tissues from the undesir-able side effects of the drug for a cancer patient [55]. As a controlexperiment, fibroblast cells were treated with nanocarriers with-out DOX. Results showed that fibroblast cell metabolism was notsignificantly influenced by DOX-free nanocarriers at the same con-centration, so a major source for cytotoxicity is the DOX by itself(Fig. S4).

Supplementary Figs. S3 and S4 related to this article can befound, in the online version, at http://dx.doi.org/10.1016/j.colsurfb.2014.07.049.

To investigate cellular uptake of nanocarriers, ICP-OES mea-surements and Prussian blue staining were carried out using theHeLa cell line. Fig. 4A shows Fe amounts measured quantitatively

with ICP-OES per cell incubated with MIONPs, non-targeted andtargeted nanocarriers separately. Results demonstrated 11-foldhigher uptake by HeLa cells with targeted nanocarriers com-pared to MIONPs. Non-targeted MIONP internalization was also

680 C. Nazli et al. / Colloids and Surfaces B: Biointerfaces 122 (2014) 674–683

F and

c

lsnscto

ig. 3. HeLa cell viability in the presence of free DOX, DOX-loaded non-targetedoncentrations of 1, 5 and 10 �g/mL DOX. Data are mean ± SD (n = 4).

ow, based on the Fe content. This Fe content is not unusualince PEG usually limits opsonization and non-specific uptake ofanocarriers. This result was further confirmed with Prussian blue

taining, which showed that the intensity of blue color from theells treated with targeted nanocarriers was 12.5 times higherhan the intensity obtained from those treated with nontargetedr bare MIONPs (Fig. 4B and Fig. S5). This result is consistent

targeted nanocarriers after (A) 24, (B) 48 and, (C) 72 h of incubation at altered

with our previous study, confirmed further that incorporation ofcollagenase-sensitive peptide sequence into the coating did notcompromise cellular uptake of targeted nanocarriers, and that the

presence of RGDS within the coating was essential to enhanceuptake by HeLa cells [11].

Supplementary Fig. S5 related to this article can be found, in theonline version, at http://dx.doi.org/10.1016/j.colsurfb.2014.07.049.

C. Nazli et al. / Colloids and Surfaces B: Biointerfaces 122 (2014) 674–683 681

Fig. 4. Internalization of bare MIONPs, non-targeted and targeted nanocarriers by HeLa cells after 24 h of incubation. Characterization by (A) inductively coupled plasmao blue sc

3

itnR(H(cFmcflrssFa

ptical emission spectrometry (ICP-OES), data are mean ± SD (n = 3). (B) Prussian

itation of this figure, the reader is referred to the web version of the article.)

.4. Characterization of intra-cellular DOX

Free DOX and DOX-loaded nanocarriers within HeLa cells weremaged using laser scanning confocal microscopy. HeLa cells werereated for 2 h with free DOX or DOX-loaded nanocarriers. HeLa celluclei were stained using DAPI, and staining were observed as blue.ed luminescence from DOX was observed within cell cytoplasmFig. 5). Results demonstrated that the highest uptake of DOX byeLa cells was achieved with DOX-loaded targeted nanocarriers

Fig. 5). Fig. 5 shows that nanocarriers without a targeting ligandould not be internalized into cells, a result consistent with thee content/cell measured by ICP-OES. This ICP-OES measurementay suggest that internalization of multifunctional PEG hydrogel-

oated MIONPs into cancer cells was specific because DOXuorescence could not be observed for the group where nanocar-iers were coated within PEG hydrogel only. In addition, the result

howed that non-targeted PEG hydrogel coating prevented non-pecific interaction between nanocarriers and cells, as expected.luorescence intensities in confocal images were also quantified,nd 1.65 times higher red fluorescence was observed for the cells

taining (scale bar = 25 �m). (For interpretation of the references to color near the

treated with DOX-loaded targeted nanocarriers compared to thecells treated with free DOX (Fig. S6A). In contrast, no fluorescencewas detectable in samples treated with DOX-loaded non-targetednanocarriers. To quantify intra-cellular DOX in each sample, cellswere lysed, and concentration of DOX in HeLa cell lysates wascompared. Results demonstrated that targeted nanocarriers were2.2 times more efficient in terms of DOX uptake into HeLa cellscompared to free DOX (Fig. S6B). In addition, Fe concentration ofall cells incubated with DOX-loaded nanocarriers was measuredvia ICP-OES after 2 h incubation at identical DOX concentration.Cells incubated with DOX-loaded targeted nanocarriers contained6.2 ± 0.025 pg Fe/cell, while cells treated with DOX-loaded non-targeted nanocarriers had only 0.12 ± 0.005 pg Fe/cell, which wassimilar to the value measured for control cells not incubated witha nanocarrier (0.12 ± 0.027; Fig. S6C). These results demonstratedthat MMP-sensitive and functional PEG hydrogel coating of MIONPs

could significantly enhance internalization of nanoparticles anddelivery of DOX into HeLa cells compared to control groups.

Supplementary Fig. S6 related to this article can be found, in theonline version, at http://dx.doi.org/10.1016/j.colsurfb.2014.07.049.

682 C. Nazli et al. / Colloids and Surfaces B: Biointerfaces 122 (2014) 674–683

F in HeLL (red),

i e web

4

cdniHdnTdtcdp

ig. 5. Laser scanning confocal microscope characterization of intra-cellular DOX

eft column:cell nuclei staining by DAPI (blue), middle column: DOX fluorescence

nterpretation of the references to color in figure legend, the reader is referred to th

. Conclusion

Targeted nanocarriers were designed for receptor-specific intra-ellular uptake of MIONPs, and for triggered release of the cancerrug DOX into HeLa cells. MIONPs coated with this strategy didot compromise cell viability in the absence of DOX, where signif-

cantly improved intra-cellular uptake of MIONPs was observed byeLa cells. Furthermore, DOX-loading within targeted nanocarriersecreased cancer cell viability within 2 days, while non-targetedanocarriers did not have a major influence on cancer cell viability.he approach used here demonstrated that targeted intra-cellularelivery of a cancer drug DOX into cancer cells was possible

hrough receptor-targeted nanocarriers, and that this approachould significantly reduce the systemic side effects of toxic cancerrugs. The multifunctional nanocarrier system developed here isromising to impact in several areas: to improve the efficiency of

a cells that were treated with free DOX, non-targeted and targeted nanocarriers.right column: overlay images of right and middle column (scale bar = 10 �m). (For

version of the article.)

existing chemotherapeutic approaches, to reduce the side effectsof the drug, and to minimize drug resistance of cancer tissue. TheMMP-sensitive and functional PEG hydrogel coating of MIONPsalso have potential for simultaneous imaging of tumor-targeteddrug-delivery and triggered drug release into the tumor site.

Acknowledgements

This study was supported by FP7-Marie Curie-IRG-DMIOL239471 to SK. SEM images were obtained at Koc University Sur-face Science Center (KUYTAM). The authors would like to thankIbrahim Hocaoglu for providing help with ICP-OES measurements.

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