Targeted delivery of doxorubicin into tumor cells via MMP-sensitive PEG hydrogel-coated magnetic iron oxide nanoparticles (MIONPs)

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

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Colloids and Surfaces B: Biointerfaces journal homepage: www.elsevier.com/locate/colsurfb

Targeted delivery of doxorubicin into tumor cells via MMP-sensitive PEG hydrogel-coated magnetic iron oxide nanoparticles (MIONPs) Caner Nazli a , Gozde S. Demirer b , Yasemin Yar a , H. Yagci Acar a,c , Seda Kizilel a,b,∗ a b c

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

a r t i c l e

i n f o

Article history: Received 4 February 2014 Received in revised form 29 July 2014 Accepted 31 July 2014 Available online 10 August 2014 Keywords: Integrin targeting Protease responsive PEG hydrogel coating Controlled release Doxorubicin Magnetic iron oxide nanoparticles Targeted nanocarrier

a b s t r a c t Targeting tumors with nano-scale delivery systems shows promise to improve the therapeutic effects of chemotherapeutic drugs. However, the limited specificity of current nano-scale systems for cancer tissues prevents realization of their full clinical potential. Here, we demonstrate an effective approach to creating as targeted nanocarriers for drug delivery: MIONPs coated with integrin-targeted and matrixmetalloproteinase (MMP)—sensitive PEG hydrogel scaffolds. The functional PEG hydrogel coating has been designed for active loading as well as triggered intra-cellular release of the cancer therapeutic agent doxorubicin (DOX). Our study demonstrated that coated nanocarriers could be taken into cancer cells 11 times more efficiently than uncoated ones. Furthermore, confocal laser scanning microscopy images revealed that these targeted nanocarriers could efficiently deliver and release DOX into the nuclei of HeLa cells within 2 h. Coating MIONPs with multifunctional PEG hydrogel could be a promising alternative to existing vehicles for targeted delivery of DOX into tumor tissue. © 2014 Elsevier B.V. All rights reserved.

1. Introduction Chemotherapy is commonly employed to treat cancer patients [1]. However, non-specificity of chemotherapeutic agents and resistance of cancer cells to drugs compromise the efficiency of chemotherapy [1–4]. Previous studies have used various types of drug-loaded nanoparticles, and a few of such systems are commercially available [5]. For clinical use, nanoparticles should demonstrate a couple of important properties including excellent in vivo stability, high drug-loading capacity, and long circulation time in the bloodstream [4,6]. Responsiveness to local stimulus such as pH, temperature, or protease would add additional benefits for drug delivery vehicle design. These properties can be used to achieve specific accumulation of nanoparticles at the cancer sites and/or to deliver therapeutic drug in the cancerous tissue in response to a stimulus [4]. MIONPs, which are superparamagnetic, have demonstrated great potential for simultaneous imaging and targeted delivery of therapeutic agents into tumor sites [7–9]. MIONPs may be coated with various materials. When polymers

∗ Corresponding author at: Koc¸ University, Chemical and Biological Engineering, Istanbul 34450, Turkey. Tel.: +90 212 338 1836; fax: +90 212 338 1548. E-mail address: [email protected] (S. Kizilel). http://dx.doi.org/10.1016/j.colsurfb.2014.07.049 0927-7765/© 2014 Elsevier B.V. All rights reserved.

are considered for coatings, end-grafting, direct adsorption, or selfassembly of linear polymers via hydrophobic interactions can be utilized [10–16]. Dextran has been widely used as a coating material due to its biocompatability and high affinity to iron oxide surfaces [17–20]. For example, a contrast agent based on MIONP was coated with 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 during the 1990s. However, the product did not meet all expectations of magnetic particle imaging and was withdrawn by the company at the end of 2008. PEG, a well-studied polymer due to its biocompatible and nontoxic nature, has been conjugated to MIONPs to increase circulation time in blood and to prevent agglomeration and degradation by macrophages in blood [4,25]. Despite a PEG coating, previously developed MIONPs may still have low colloidal stability and lack stimuli responsiveness for enhanced drug release at the specific site [26,27]. Responsive design of the MIONP coating is also beneficial to accomplish high contrast in medical imaging [28]. Hence, advanced strategies are needed both to enhance therapeutic efficacy and minimize side effects of cancer drugs. The on-demand delivery of chemotherapeutics at a targeted site to improve therapeutic efficacy is crucial for translating research into clinical application [29–32]. To achieve this end goal, many studies have developed

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drug-delivery systems using materials such as inorganic nanoparticles [33,34], liposomes [35,36], biomolecules [37], micelles [38,39], carbon materials [40,41], and polymeric nanoparticles [42,43]. The major approach in enhancing specificity is designing drug-delivery vehicles that specifically recognize tumor cells. As an example, targeting ␣v ␤3 integrin is a popular approach in tumor-targeted drug-delivery with nanoparticles. Transmembrane receptor ␣v ␤3 integrin has angiogenic, migration features, and is overexpressed in tumor cells [4]. ␣v ␤3 integrin initiates angiogenesis and regulates cell migration, which are crucial for gaining access to nutrients and spreading of cancer cells [44]. Tumor tissues also secrete high amounts of proteolytic enzymes, such as MMPs, which degrade the basement membrane and natural extracellular matrix (ECM), opening up more space for tumor growth [44,45]. Elevated expression of MMP-2 and MMP-9 was observed in some malignant tumor tissues such as breast cancer [46,47], colon cancer [48,49], brain tumors [50,51], and other malignancies [52]. For example, in a study by Lizasa et al. [53], lung cancer patients had plasma MMP-9 concentrations of 71.0 ± 60.2 ng/ml, whereas in healthy control patients 36.3 ± 13.2 ng/ml was measured. In another study by Tutton et al. [54], significant increases in MMP-2 and MMP-9 were observed in protein expression in colorectal cancer specimens compared to normal colonic tissue. MMP-2 and MMP-9 total protein concentrations in normal colonic tissue were measured as 4.2 ±0.4 ng/mg and 6.5 ± 3.5 ng/mg, respectively. In the tumor environment, however, total protein concentrations of MMP-2 and MMP-9 were increased to 180.4 ± 99.3 ng/ml and 483.4 ± 73.9 ng/ml, respectively. Due to these high concentrations of MMPs measured in tumor tissue, proteolytically degradable peptide sequences attracted significant attention, and subsequent research used MMP cleavable peptide conjugated drugs [55,56] and polymeric shells [57]. In this study, we designed a multifunctional MIONP that may serve as a contrast agent in MRI, bind and carry chemotherapeutic drug, target overexpressed receptors on cancer cell surface, and be sensitive to MMP. These MIONPs have proteolytically degradable PEG hydrogel coatings with integrin-binding RGDS domains, and are loaded with doxorubicin (DOX; Scheme 1). We hypothesized that MMP-sensitive domains within the coating can be effectively cleaved as a result of high concentration of MMPs in the cancer cell environment, and this coating may allow for the release of loaded cancer drug DOX into cancer cells specifically. We synthesized acrylate-PEG-RGDS and acrylate-PEGGGGPQG↓IWGQGK–PEG-acrylate conjugates, and included these conjugates in the PEG diacrylate prepolymer solution. We call this acrylate mixture a functional prepolymer solution. The proteolytically degradable sequence, GGGPQG↓IWGQGK, is derived from ␣1(I) collagen and can be degraded by several MMPs [58]. Next, we suspended MIONPs with immobilized photoinitiator in the functional prepolymer solution, followed by surface-initiated photopolymerization, to obtain a functional coating around MIONPs (Scheme 2). We investigated the synthesis and characterization of functional PEG hydrogel-coated MIONPs using Fourier Transform Infrared (FT-IR) spectroscopy, dynamic light scattering (DLS), zeta potential, and scanning electron microscopy (SEM). Confocal microscope images demonstrated that intra-cellular DOX release into cancer cells could only be achieved with the targeted MIONPs. Significantly improved internalization of MIONPs and DOX retention were accompanied by decreased HeLa cell viability within 48 h in the in vitro studies. The method presented here extends our previous work on the PEG hydrogel-coated MIONPs via surface-initiated photopolymerization [11]. The technique presented here, which coats MIONPs with a multifunctional PEG hydrogel, has been used for receptor-targeted drug-delivery into cancer cells and specific unloading of the chemotherapeutic agent. This coating can have a

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Scheme 1. Illustration of intra-cellular delivery of cancer drug DOX into cancer cells by targeted nanocarriers. RGDS ligand of the targeted nanocarriers binds to ␣v ␤3 integrin, and nanocarrier can internalize into cancer cells via receptor-mediated endocytosis. After degradation of MMP-sensitive domains in PEG hydrogel coating of targeted nanocarriers, loaded DOX is released into cytoplasm of cancer cells.

variety of applications when multiple functionality of MIONP is desirable. 2. Materials and methods 2.1. Materials Eosin Y (91%), 1-vinyl-2-pyrrolidinone (VP; 99%), poly (ethylene glycol) diacrylate (PEGDA) (MW: 575 Da), Woodward’s reagent K (WRK) (2-ethyl-5-phenylisox-azolium-3 -sulfonate, 95%), phosphate-buffered saline (PBS), Triton® X-100, hydroxyethyl piperazineethanesulfonic acid (HEPES), and trypsin-ethylenediaminetetraacetic acid were obtained from Sigma–Aldrich (St. Louis, MO). FeCl3 ·6H2 O and FeCl2 ·4H2 O were obtained from Fluka (Sigma–Aldrich, Taufkirchen, Germany). Ammonium hydroxide (26% NH3 in water, w/w) was obtained from Riedel-de Haën (Hanover, Germany). APTMS was obtained from Gelest, Inc. (Morrisville, PA). Triethanolamine (TEOA) (99.5%), sodium bicarbonate, and 37% formaldehyde were purchased from Merck (Whitehouse Station, NJ). Hydrochloric acid (HCl) and sodium hydroxide (NaOH) were purchased from Riedel-de Haën. RGDS (MW: 433 Da) and GGGPQG↓IWGQGK (MW: 1141 Da) peptides were purchased from Elim Biopharmaceuticals, Inc. (Hayward, CA). Acrylate PEG N-hydroxysuccinimide (acr-PEG-NHS) (MW: 3.4 kDa) was obtained from Nektar (Huntsville, AL). Potassium ferrocyanide was obtained from Labkim (Istanbul, Turkey). Doxorubicin (DOX) was purchased from MP Biomedicals (Solon, OH). DAPI (4 ,6-Diamidino-2-Phenylindole, Dihydrochloride) was purchased from Invitrogen Co. (Carlsbad, CA). Collagenase type 1 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-glutamine were purchased form Gibco (Life Technologies Products, Grand Island, NY). Penicillin/streptomycin was obtained from Invitrogen Co (Carlsbad, CA). Donor horse serum was obtained from Biological

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Scheme 2. (A) Schematic representation of MIONPs coating within multifunctional PEG hydrogel. (B) Synthesis of acryl-PEG-RGDS and acryl-PEG-GGGPQG↓IWGQGK-PEGacryl conjugates.

Industries (Rainbow Scientific, Inc., Windsor, CT). The Amicon Ultra centrifugal filter device (NMWL: 3000) was purchased from EMD Millipore. The dialysis cassette (MW cutoff: 3.5 and 20 kDa) was obtained from Pierce (Micro BCA Protein Assay; Thermo Fisher Scientific Inc., Rockford, IL). Teflon syringe filters (0.2 ␮m pore size) were purchased from Nalgene (Rochester, NY). Petri dishes were obtained from Nunc (Roskilde, Denmark). Cell culture plates (96- and 24-well) were purchased from Greiner Bio-One GmbH (Solingen, Germany). Ninety-six-well LUMITRAC® 600 white immunology plates were purchased from USA Scientific, Inc. (Ocala, FL).

2.2. Synthesis and coating of MIONPs MIONPs were coated above within PEG hydrogel via surfaceinitiated photopolymerization as previously described [11,59,60]. In brief, EY was covalently bound to the surface amine groups of APTMS-coated MIONPs (Scheme 2). EY bound MIONPs were suspended in prepolymer solution composed of 6.25% PEGDA (w/v) (MW: 575 Da), TEOA (141 mM), and VP (12.5 mM) in

deionized water. Acr-PEG-RGDS and acr-PEG-GGGPQG↓IWGQGKPEG-acr (acr-PEG-DEG-PEG-acr) conjugates were also synthesized and characterized as described above (Scheme 2). To purify the degradable conjugate, a dialysis membrane cassette with molecular weight cut off (MWCO) of 7 kDa was used to obtain conjugate with diacrylate functionality. To obtain a PEG hydrogel coating with an integrin binding RGDS, acr-PEG-RGDS (1.48 × 10−5 M) and acrPEG-DEG-PEG-acr (1.0 × 10−5 M) were added into the prepolymer solution before the photopolymerization reaction. The mixtures were exposed to 514 nm green light at a flux of 2.5 mW/cm2 using an argon ion laser (Coherent Inc., Santa Clara, CA) for 60 s. To remove unreacted components, the polymerization mixture was washed using the Amicon Ultra centrifugal filter, providing clean MIONPs 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 coated MIONPs (non-targeted nanocarriers), and multifunctional PEG

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hydrogel-coated MIONPs (targeted nanocarriers) were characterized in terms of structure, size, morphology, and surface charge. FT-IR spectroscopy (Nicolet iS10 FT-IR spectrometer; Fisher Scientific, Vienna, Austria) was used to confirm the presence of PEGDA chains and to characterize functional groups of the polymeric coating around nanoparticles. Core diameter, and morphology of non-targeted and targeted nanocarriers were analyzed via SEM (Zeiss Ultra Plus, Bruker; Optronik Ltd. Sti, Ankara, Turkey). The hydrodynamic diameter and zeta potential were characterized with a Zetasizer instrument at room temperature (Malvern Zetasizer nano S ZEN 1600, wavelength 633 nm; Malvern Instruments, Malvern, UK). 2.4. In vitro degradation kinetics of targeted nanocarriers and release of DOX from targeted nanocarriers Targeted nanocarriers were suspended in HEPES buffer (0.1 M HEPES + 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 M NaCl + 5 mM CaCl2 at pH 7.2) in the presence and absence of collagenase type I enzyme (0.1 mg/mL at 37 ◦ C) for 4 days. During this treatment, hydrodynamic diameters of nanocarriers were measured at specific time points within the first 72 h. The aqueous suspensions of non-targeted and targeted nanocarriers (1 mg/mL) were suspended separately in an aqueous solution of DOX (0.1 mg/mL) at a nanocarrier-to-DOX weight ratio of 10. Each mixture was sonicated for 3 h and incubated for 48 h at room temperature. To remove unloaded DOX, samples were rinsed three times with Amicon Ultra centrifugal filter. Concentration of DOX in the eluted solution was determined with the absorbance of free DOX using NanoDrop 1000 spectrophotometer (Fisher Scientific, UK) at 490 nm. The percentage of DOX loading efficiency (DLE) was calculated using the following equation: % DLE =

amount of DOX in coated nanoparticles × 100 amount of DOX in buffer medium

DOX-loaded targeted nanocarriers were suspended in HEPES buffer in the presence and absence of collagenase type 1 enzyme (0.1 mg/mL), separately. Next, dialysis was carried out using dialysis cassette membranes (MWCO = 20 kDa) in HEPES buffer in the presence and absence of collagenase type 1 enzyme at 37 ◦ C, and at specific time points 1 mL solution was withdrawn from the buffer medium and replaced with 1 mL fresh buffer. Concentration of DOX in the withdrawn medium was measured at each time point. 2.5. Evaluation of cytotoxicity of DOX-loaded nanocarriers Cytotoxicity assays were carried out with HeLa cells and mouse 3T3 fibroblast cells to investigate the influence of non-targeted and targeted nanocarriers on the viability of cancer and healthy cell lines. HeLa cells (5 × 104 cells/mL in a 96-well plate) were cultured for 24 h in DMEM (pH 7.4, with 10% donor horse serum (DHS), 1% penicillin/streptomycin, and l-glutamine). Suspensions with nontargeted and targeted nanocarriers in HEPES buffer were added to HeLa cells and mouse 3T3 fibroblast cell culture mediums, separately. Viabilities were measured after 24, 48 and 72 h of incubation using CellTiter-Glo Luminescent viability assay (Promega, Madison, WI) with Fluoroskan Ascent luminescence reader (Fisher Scientific, Vienna, Austria). DOX concentrations of 1, 5, and 10 ␮g/mL were tested for each nanocarrier group. 2.6. Internalization of DOX-free nanocarriers Internalization of bare MIONPs, non-targeted and targeted nanocarriers by HeLa cells were investigated via Prussian blue staining (PBS). HeLa cells (5 × 104 cells/mL) were cultured for 24 h

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in 24-well plates on glass coverslips. After the attachment of HeLa cells to coverslips, nanoparticles at a concentration of 0.1 mg/mL in HEPES buffer were added to each well. At the end of 24 h, cells attached to coverslips were washed three times with PBS and fixed with 4% formaldehyde in PBS for 20 min at room temperature. Cells were incubated with 0.1% Triton X-100 in PBS for 5 min and then rinsed with PBS. Fixed HeLa cells were stained with 5% potassium ferrocyanide in 10% HCl for 15 min. Prussian blue stained cells were visualized via a camera equipped with optical microscopy (Eclipse Ni-U Optical Microscope, Nikon, Istanbul, Turkey). Cellular uptake of nanocarriers was also characterized via inductively 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 transferred to cell culture medium at a concentration of 0.1 mg/mL after 24 h of incubation. As a control group, HEPES buffer was added into cell culture medium without nanocarriers. Following 24 h of incubation, 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 the cell 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-OES analysis was applied for all samples. To calculate the amount of iron (Fe) per cell in each sample, two groups of samples were used, where cell numbers were calculated with CellTiter-Glo Luminescent viability assay. Reported values were averages of at least three experiments. 2.7. Internalization of DOX-loaded nanocarriers HeLa cells were seeded on coverslips in 24-well plates at a density of 5 × 104 cells/mL for 24 h as described above. DOX concentration of 10 ␮g/mL was loaded into nanocarriers, and these nanocarriers were added to cell culture medium. After 2 h of incubation, treated cells were washed with PBS three times. Next, cell nucleus was stained with DAPI, according to the standard protocol provided by the supplier. Cellular internalization of DOX was visualized by camera-equipped laser scanning confocal microscopy (Eclipse 90-i, Nikon). To quantify the Fe and DOX amount per cell for each sample, ICP-OES and NanoDrop 1000 spectrophotometer were used, respectively. 2.8. Statistical analysis The results of all data sets were evaluated using one-way analysis of variance. The results are shown as the mean value (plus or minus standard deviation) of the triplicate samples unless stated otherwise. Variances between datasets were considered statistically significant for P-values less than 0.05. 3. Results and discussion 3.1. Nanocarrier synthesis PEG hydrogel coating on an eosin-immobilized MIONP surface was conducted via surface-initiated photopolymerization as described previously. Functional PEG-MIONPs were coated with PEG hydrogel prepolymer solution that included acr-PEG-RGDS and acr-PEG-GGGPQG↓IWGQGK-PEG-acr conjugates. The presence of PEG hydrogel, RGDS and GGGPQG↓IWGQGK peptides around MIONP structure were confirmed with FT-IR (Fig. S1). FTIR bands at 1728 cm−1 indicated C O stretching of PEG acrylate. In the FT-IR spectra of non-targeted nanocarriers, C C stretching bands of acrylate groups at 1640 cm−1 disappeared, which indicated the reaction of the double bonds during photopolymerization reaction. After the addition of acr-PEG-RGDS and

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the size range of 20–200 nm to enter and localize at the interstitial space in the tumor environment [4]. In addition, surface charge of nanocarriers can influence cellular uptake by tumor cells [61]. Macrophages can easily eliminate cationic and highly negatively charged nanoparticles due to non-specific internalization [61]. Zeta potential values for MIONPs, non-targeted and targeted nanocarriers 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 atoms in PEG chains (Table 1). This negative zeta potential may prevent non-specific internalization of nanocarriers by macrophages during circulation in blood because non-specific uptake can be minimized within the zeta potentials range of −9 to −3 mV [61]. This value of zeta potential may also facilitate deep penetration and internalization of nanocarriers into the tumor site specifically [62,63] if nanocarriers can be tagged with tumor-specific molecules and/or directed to the tumor site magnetically as in the case of MIONPs [64]. 3.2. In vitro evaluation of MMP-sensitive degradation of PEG hydrogel coating and DOX release Fig. 1. Scanning electron microscope (SEM) image of (A) non-targeted nanocarriers, (B) targeted nanocarriers (scale bar:100 nm).

acr-PEG-GGGPQG↓IWGQGK-PEG-acr conjugates into the prepolymer, a new peak at 1670 cm−1 appeared, which corresponded to the amide I (C O stretching) peak coming from the peptide sequence. All of these features observed in the FT-IR confirmed the presence of functional PEG hydrogel around MIONPs and the absence of any residual acrylate (Fig. S1). Supplementary Fig. S1 related to this article can be found, in the online version, at http://dx.doi.org/10.1016/j.colsurfb.2014.07.049. The size and surface charge of nanocarriers are important physicochemical parameters for designing targeted drug-delivery vehicles. Size and morphology of nanocarriers before and after hydrogel coating were characterized using SEM and DLS. Fig. 1 demonstrated that the particles have spherical and ellipsoid shapes. Ellipsoidal shapes most probably occurred as a result of fusing of two or more MIONPs during polymerization reaction. SEM results demonstrated that non-targeted nanocarriers had an average diameter of 28.5 ± 4.8 nm, where this value was 186.5 ± 15.8 nm for targeted nanocarriers (Fig. 1). These results were similar to the hydrodynamic diameters measured with dynamic light scattering (DLS): 25.3 ± 2.9 and 209.9 ± 24.3 nm for non-targeted and targeted nanocarriers, respectively (Table 1). The diameter of MIONPs with only APTMS groups on the surface was 7.3 ± 0.7 nm. Larger diameters were measured with PEG hydrogel coating. This finding was most probably due to the formation and swelling of PEG hydrogel around MIONPs. Even larger sizes measured for targeted nanocarriers might be due to the longer PEG chains used for RGDS and degradable peptide conjugates, which might have promoted access of integrin receptors and MMP to their ligands. Tumor blood vessels consist of a high number of reproducing endothelial cells, where the absence of pericyte and abnormal basement membrane formation results in higher vascular permeability [4]. The size of endothelial pores varies from 10 to 1000 nm, which means that theoretically it may be possible for nanocarriers within Table 1 Hydrodynamic diameter and zeta potential values of bare MIONPs, non-targeted and targeted nanocarriers. Data are mean ± standard deviation (SD) (n = 3). Sample name

Hydrodynamic size

Zeta potential

Bare MIONPs Non-targeted nanocarriers Targeted nanocarriers

7.3 ± 0.7 25.3 ± 2.9 209.9 ± 24.3

14 ± 1.49 −4.18 ± 0.45 −3.37 ± 0.28

Degradation profiles for MMP-sensitive coating of targeted nanocarriers were investigated via hydrodynamic size measurements using DLS. Targeted nanocarriers were incubated in HEPES buffers with collagenase type 1 enzyme at 37 ◦ C, and hydrodynamic diameter was measured at different time points for 3 days. Targeted nanocarriers demonstrated a sharp decrease in hydrodynamic diameter from 229.4 ± 12.86 nm to 59 ± 3.25 nm after one day incubation in collagenase type 1 containing buffer. Incubation of nanocarriers for three days in collagenase buffer led to further decrease in hydrodynamic size down to 30.86 ± 5.03 nm. Targeted nanocarriers incubated in the absence of collagenase type 1 enzyme had slightly increased hydrodynamic diameters during the incubation period due to swelling of the PEG hydrogel coating (Fig. 2A). These results confirmed that targeted nanocarriers were sensitive to collagenase type 1 enzyme, and were degraded only in its presence. DOX (DOX.HCl) was loaded into non-targeted and targeted nanocarriers in HEPES buffer to examine drug release profiles in the presence and absence of collagenase type 1 enzyme at 37 ◦ C. The interaction of positively charged DOX with negatively charged 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-targeted and targeted nanocarriers, respectively. High loading efficiency of nanocarriers may be attributed to the interaction of the slightly negative PEG coating with the positively charged DOX. Nanocarriers (10.2 ␮g/mL) contained 1 ␮g/mL DOX when DOX was loaded into non-targeted nanocarriers. As shown in Fig. 2B, DOX release from targeted nanocarriers improved in the presence of collagenase enzyme. After 10 h, release of DOX from targeted nanocarriers in the presence of collagenase slightly exceeded the release observed in the absence of collagenase. Major differences in release rates were observed after 3 days and at the end of 4 days incubation: 60% of DOX was released from nanocarriers in the presence of collagenase, while 36% of loaded DOX was released from nanocarriers in the absence of collagenase. The higher release of DOX obtained in the presence of collagenase was attributed to the specific degradation of the coating and disassembling of the aggregates in response to collagenase, and hence faster diffusion of the drug from nanocarriers. Even though the size of the coating was decreased from 229.4 ± 12.86 nm to 59 ± 3.25 nm at the end of day 1, DOX release continued for 4 days. Despite the high collagenase enzyme used for the degradation of the carrier (0.1 mg/mL), the

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Fig. 2. (A) Degradation kinetics of targeted nanocarriers in the presence and absence of 0.1 mg/mL of collagenase type 1. (B) DOX release versus time from DOX-loaded targeted nanocarriers in the presence and absence of collagenase type 1. Filled red triangle: DOX release from DOX-loaded targeted nanocarriers in HEPES buffer, filled blue squares: DOX-loaded targeted nanocarriers in HEPES buffer in the presence of 0.1 mg/mL of collagenase type 1. Data are mean ± SD (n = 3). (For interpretation of the references to color in figure legend, the reader is referred to the web version of the article.)

release of DOX from the carrier in the presence of collagenase is 1.6 times greater than the release observed in the absence of collagenase. This limited release could have occurred due to the presence of nondegradable PEGDA monomer used in the prepolymer. That is, no matter how high the collagenase concentration was, we would not observe complete degradation of the coating. The other reason could be associated with the presence of crosslink density gradients within the coating and high crosslink densities at points closer to the nanoparticle surface. This crosslink density gradient occurs as a result of surface-initiated photopolymerization technique used here, and might have resulted in higher drug-loading at stiffer locations close to the nanoparticle surface within the membrane coating [67]. 3.3. In vitro cytotoxicity and cellular uptake of nanocarriers Cytotoxicity and tumor targeting properties of non-targeted and targeted nanocarriers were investigated with HeLa cells. Cells were incubated with free-DOX, DOX-free nanocarriers and DOX-loaded nanocarriers. Cells were treated with 1, 5, and 10 ␮g/mL of free DOX or an equivalent dose of DOX-loaded nanocarriers for 3 days (Fig. 3). As shown in Fig. 3A, after 24 h of drug exposure, DOXloaded non-targeted or targeted nanocarriers reduced the viability

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of cells slightly, at the highest drug concentration of 10 ␮g/mL. After 48 and 72 h, cells treated with DOX-loaded targeted nanocarriers had viability of about 10%, whereas the viability of DOX-loaded non-targeted nanocarriers remained above 80% at the highest DOX concentration of 10 ␮g/mL (Fig. 3B). After 3 days incubation, HeLa cells treated with DOX-loaded targeted nanocarriers had decreased viability of 60% and 50% at 1 and 5 ␮g/mL DOX concentrations, respectively (Fig. 3C). Interestingly, free DOX at 1 and 5 ␮g/mL did not reduce the viability of cells below 80%. These results demonstrated that DOX-loaded targeted nanocarriers showed dose and time dependent activity and were efficient to decrease the viability of HeLa cells. This decreased viability could be explained by the mechanism of nanocarrier internalization by endocytosis and therefore better localization of the drug in HeLa cells. Targeted nanocarriers were most probably taken into HeLa cells via receptor-specific endocytosis, hence cells did not have sufficient time to develop resistance toward the cancer drug. However, free DOX in the medium allows faster recognition by cancer cells, and hence more time for cancer cells to develop resistance toward the drug. As a control experiment, HeLa cells was treated with nanocarriers without DOX. Results showed that metabolism of HeLa cells was not significantly influenced by DOX-free nanocarriers at the same nanocarrier concentration, which indicated cytocompatibility of the PEG-coated MIONPs (Fig. S2). Supplementary Fig. S2 related to this article can be found, in the online version, at http://dx.doi.org/10.1016/j.colsurfb.2014.07.049. The toxicity of DOX-loaded targeted nanocarriers was also tested against healthy cell lines from mouse fibroblast cells. Results showed that free DOX samples at 5 and 10 ␮g/mL concentrations negatively influenced healthy cell survival (Fig. S3). When DOXloaded non-targeted or targeted nanocarriers were used, fibroblast cell viabilities of 90% and 60% were measured at the highest DOX concentration, respectively (Fig. S3). This viability result could be explained by the differences in the interaction mechanism of DOX and fibroblasts when DOX was introduced freely or was loaded within nanocarriers. Cell internalization was not as efficient for the case of fibroblasts as it was for HeLa cells, due to the fewer number of integrin receptors on fibroblasts. Also, collagenase enzyme concentration is not high around fibroblasts; a high concentration is necessary for degradation of coating and hence release of DOX from nanocarriers. Free DOX could easily interact with healthy cells and can negatively influence cell survival, while DOX loaded within nanocarriers may not interact with cells quickly. These results suggest that targeted nanocarriers would be efficient to minimize interaction of a therapeutic drug with healthy cells, and that high internalization of targeted nanocarriers into tumor cells followed by minimized cancer cell survival due to targeted drug-delivery would be achieved with this system. However, with the use of nontargeted systems such as MMP-sensitive peptide conjugated DOX, it may be challenging to protect healthy tissues from the undesirable side effects of the drug for a cancer patient [55]. As a control experiment, fibroblast cells were treated with nanocarriers without DOX. Results showed that fibroblast cell metabolism was not significantly influenced by DOX-free nanocarriers at the same concentration, so a major source for cytotoxicity is the DOX by itself (Fig. S4). Supplementary Figs. S3 and S4 related to this article can be found, in the online version, at http://dx.doi.org/10.1016/j.colsurfb. 2014.07.049. To investigate cellular uptake of nanocarriers, ICP-OES measurements and Prussian blue staining were carried out using the HeLa cell line. Fig. 4A shows Fe amounts measured quantitatively with ICP-OES per cell incubated with MIONPs, non-targeted and targeted nanocarriers separately. Results demonstrated 11-fold higher uptake by HeLa cells with targeted nanocarriers compared to MIONPs. Non-targeted MIONP internalization was also

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Fig. 3. HeLa cell viability in the presence of free DOX, DOX-loaded non-targeted and targeted nanocarriers after (A) 24, (B) 48 and, (C) 72 h of incubation at altered concentrations of 1, 5 and 10 ␮g/mL DOX. Data are mean ± SD (n = 4).

low, based on the Fe content. This Fe content is not unusual since PEG usually limits opsonization and non-specific uptake of nanocarriers. This result was further confirmed with Prussian blue staining, which showed that the intensity of blue color from the cells treated with targeted nanocarriers was 12.5 times higher than the intensity obtained from those treated with nontargeted or bare MIONPs (Fig. 4B and Fig. S5). This result is consistent

with our previous study, confirmed further that incorporation of collagenase-sensitive peptide sequence into the coating did not compromise cellular uptake of targeted nanocarriers, and that the presence of RGDS within the coating was essential to enhance uptake by HeLa cells [11]. Supplementary Fig. S5 related to this article can be found, in the online version, at http://dx.doi.org/10.1016/j.colsurfb.2014.07.049.

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Fig. 4. Internalization of bare MIONPs, non-targeted and targeted nanocarriers by HeLa cells after 24 h of incubation. Characterization by (A) inductively coupled plasma optical emission spectrometry (ICP-OES), data are mean ± SD (n = 3). (B) Prussian blue staining (scale bar = 25 ␮m). (For interpretation of the references to color near the citation of this figure, the reader is referred to the web version of the article.)

3.4. Characterization of intra-cellular DOX Free DOX and DOX-loaded nanocarriers within HeLa cells were imaged using laser scanning confocal microscopy. HeLa cells were treated for 2 h with free DOX or DOX-loaded nanocarriers. HeLa cell nuclei were stained using DAPI, and staining were observed as blue. Red luminescence from DOX was observed within cell cytoplasm (Fig. 5). Results demonstrated that the highest uptake of DOX by HeLa cells was achieved with DOX-loaded targeted nanocarriers (Fig. 5). Fig. 5 shows that nanocarriers without a targeting ligand could not be internalized into cells, a result consistent with the Fe content/cell measured by ICP-OES. This ICP-OES measurement may suggest that internalization of multifunctional PEG hydrogelcoated MIONPs into cancer cells was specific because DOX fluorescence could not be observed for the group where nanocarriers were coated within PEG hydrogel only. In addition, the result showed that non-targeted PEG hydrogel coating prevented nonspecific interaction between nanocarriers and cells, as expected. Fluorescence intensities in confocal images were also quantified, and 1.65 times higher red fluorescence was observed for the cells

treated with DOX-loaded targeted nanocarriers compared to the cells treated with free DOX (Fig. S6A). In contrast, no fluorescence was detectable in samples treated with DOX-loaded non-targeted nanocarriers. To quantify intra-cellular DOX in each sample, cells were lysed, and concentration of DOX in HeLa cell lysates was compared. Results demonstrated that targeted nanocarriers were 2.2 times more efficient in terms of DOX uptake into HeLa cells compared to free DOX (Fig. S6B). In addition, Fe concentration of all cells incubated with DOX-loaded nanocarriers was measured via ICP-OES after 2 h incubation at identical DOX concentration. Cells incubated with DOX-loaded targeted nanocarriers contained 6.2 ± 0.025 pg Fe/cell, while cells treated with DOX-loaded nontargeted nanocarriers had only 0.12 ± 0.005 pg Fe/cell, which was similar to the value measured for control cells not incubated with a nanocarrier (0.12 ± 0.027; Fig. S6C). These results demonstrated that MMP-sensitive and functional PEG hydrogel coating of MIONPs could significantly enhance internalization of nanoparticles and delivery of DOX into HeLa cells compared to control groups. Supplementary Fig. S6 related to this article can be found, in the online version, at http://dx.doi.org/10.1016/j.colsurfb.2014.07.049.

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Fig. 5. Laser scanning confocal microscope characterization of intra-cellular DOX in HeLa cells that were treated with free DOX, non-targeted and targeted nanocarriers. Left column:cell nuclei staining by DAPI (blue), middle column: DOX fluorescence (red), right column: overlay images of right and middle column (scale bar = 10 ␮m). (For interpretation of the references to color in figure legend, the reader is referred to the web version of the article.)

4. Conclusion Targeted nanocarriers were designed for receptor-specific intracellular uptake of MIONPs, and for triggered release of the cancer drug DOX into HeLa cells. MIONPs coated with this strategy did not compromise cell viability in the absence of DOX, where significantly improved intra-cellular uptake of MIONPs was observed by HeLa cells. Furthermore, DOX-loading within targeted nanocarriers decreased cancer cell viability within 2 days, while non-targeted nanocarriers did not have a major influence on cancer cell viability. The approach used here demonstrated that targeted intra-cellular delivery of a cancer drug DOX into cancer cells was possible through receptor-targeted nanocarriers, and that this approach could significantly reduce the systemic side effects of toxic cancer drugs. The multifunctional nanocarrier system developed here is promising to impact in several areas: to improve the efficiency of

existing chemotherapeutic approaches, to reduce the side effects of the drug, and to minimize drug resistance of cancer tissue. The MMP-sensitive and functional PEG hydrogel coating of MIONPs also have potential for simultaneous imaging of tumor-targeted drug-delivery and triggered drug release into the tumor site. Acknowledgements This study was supported by FP7-Marie Curie-IRG-DMIOL 239471 to SK. SEM images were obtained at Koc University Surface Science Center (KUYTAM). The authors would like to thank Ibrahim Hocaoglu for providing help with ICP-OES measurements. References [1] M.M. Gottesman, T. Fojo, S.E. Bates, Nat. Rev. Cancer 2 (2002) 48.

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