Biodegradable polymeric vesicles containing magnetic nanoparticles, quantum dots and anticancer drugs for drug delivery and imaging

Biomaterials 35 (2014) 3885e3894

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Biodegradable polymeric vesicles containing magnetic nanoparticles, quantum dots and anticancer drugs for drug delivery and imaging Fei Ye a, Åsa Barrefelt a,1, Heba Asem a,1, Manuchehr Abedi-Valugerdi a, Ibrahim El-Serafi a, Maryam Saghafian a, Khalid Abu-Salah b, **, Salman Alrokayan b, Mamoun Muhammed c, Moustapha Hassan a, d, * a

Division of Experimental Cancer Medicine, Department of Laboratory Medicine (LABMED), Karolinska Institutet, SE-141 86 Stockholm, Sweden King Abdullah Institute for Nanotechnology, King Saud University, Riyadh 11451, P.O. Box 2455, Saudi Arabia Division of Functional Materials, School of Information and Communication Technology, Royal Institute of Technology (KTH), SE-164 40 Stockholm, Sweden d Clinical Research Center, Karolinska University Hospital - Huddinge, SE-141 86 Stockholm, Sweden b c

a r t i c l e i n f o

a b s t r a c t

Article history: Received 15 December 2013 Accepted 16 January 2014 Available online 1 February 2014

We have developed biodegradable polymeric vesicles as a nanocarrier system for multimodal bioimaging and anticancer drug delivery. The poly(lactic-co-glycolic acid) (PLGA) vesicles were fabricated by encapsulating inorganic imaging agents of superparamagnetic iron oxide nanoparticles (SPION), manganese-doped zinc sulfide (Mn:ZnS) quantum dots (QDs) and the anticancer drug busulfan into PLGA nanoparticles via an emulsion-evaporation method. T2* -weighted magnetic resonance imaging (MRI) of PLGAeSPIONeMn:ZnS phantoms exhibited enhanced negative contrast with r2* relaxivity of approximately 523 s
1 mM
1 Fe. Murine macrophage (J774A) cellular uptake of PLGA vesicles started fluorescence imaging at 2 h and reached maximum intensity at 24 h incubation. The drug delivery ability of PLGA vesicles was demonstrated in vitro by release of busulfan. PLGA vesicle degradation was studied in vitro, showing that approximately 32% was degraded into lactic and glycolic acid over a period of 5 weeks. The biodistribution of PLGA vesicles was investigated in vivo by MRI in a rat model. Change of contrast in the liver could be visualized by MRI after 7 min and maximal signal loss detected after 4 h post-injection of PLGA vesicles. Histological studies showed that the presence of PLGA vesicles in organs was shifted from the lungs to the liver and spleen over time. Ó 2014 Elsevier Ltd. All rights reserved.

Keywords: Biodegradable polymer Multifunctional nanoparticles Anticancer drug delivery Busulfan Fluorescence imaging Magnetic resonance imaging

1. Introduction Functional nanoparticles used in the treatment of cancer attract extensive attention due to their intrinsic physical properties, long blood circulation time, specific targeting capability, enhanced intracellular uptake and manipulation of molecular behavior on the nanometer scale [1e4]. So far, studies on the treatment of cancer and other diseases using nanomaterials have mainly focused on therapy by locally produced cytotoxic heat [5,6], targeted drug delivery [7,8] or a combination of these two strategies [9e11]. In particular, several nano-structured particles based on polymers

* Corresponding author. Division of Experimental Cancer Medicine, Department of Laboratory Medicine (LABMED), Karolinska Institutet, SE-141 86 Stockholm, Sweden. Tel.: þ46 8 5858 3862; fax: þ46 8 5858 3800. ** Corresponding author. Tel.: þ966 996 4675956. E-mail addresses: [email protected] (K. Abu-Salah), [email protected], [email protected] (M. Hassan). 1 These authors contributed equally. 0142-9612/$ e see front matter Ó 2014 Elsevier Ltd. All rights reserved. http://dx.doi.org/10.1016/j.biomaterials.2014.01.041

[12], lipids [13], inorganic materials [14] or natural materials [15] have been investigated for therapeutic purposes, mainly as drug delivery vehicles. With appropriate encapsulation, drugs are more stable in a physiological environment and the kinetics of the drugs can be more carefully controlled [16]. Furthermore, targeted drug delivery can be developed to improve chemotherapy in cancer treatment, not only by reducing the adverse effects in non-target organs but also by enhancing the therapeutic efficacy in the targeted organ [17]. A colloidal system based on biodegradable polyester nanoparticles (such as polylactic acid (PLA)), polylactic-co-glycolic acid (PLGA) and polycaprolactone (PCL) nanoparticles represents one of the most promising candidates for in vivo diagnosis and treatment for cancers, from preclinical development to clinical translation [1,18]. The use of these amphiphilic polymers results in the formation of nanoparticles with a hydrophobic core and a hydrophilic shell. The coreeshell structure allows them to encapsulate and carry poorly water-soluble drugs [19] and to release these drugs at a sustained rate in the optimal range of drug concentration [20]. They

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can be further functionalized with polyethylene glycol (PEG) to avoid nonspecific absorption by proteins and fast clearance by the immune system [21,22] as well as equipped with targeting ligands for delivery of drugs to specific pathological sites [12]. Another important aspect is the encapsulation of inorganic nanoparticles together with anticancer drugs into the core of polymeric nanoparticles for localized contrast enhancement in different medical visualization techniques, which makes them superior in physical and chemical properties to commercial imaging agents [23,24], thereby providing precise diagnosis and evaluation of therapeutic efficacy. In the present investigation, we report the development of a multifunctional polymeric drug delivery system aiming to deliver anti-cancer drugs and to enable in vivo and in vitro imaging in order to study the cell uptake as well as biodistribution of polymeric nanoparticles by magnetic resonance imaging (MRI) and histopathology. This drug delivery system consisted of PLGA nanoparticles encapsulating two hydrophobic inorganic nanocrystals, superparamagnetic iron oxide nanoparticles (SPION) for in vivo MRI [25] and cadmium-free manganese-doped zinc sulfide (Mn:ZnS) quantum dots (QDs) for fluorescence in vitro imaging. The PLGA vesicles (i.e., PLGAeSPIONeMn:ZnS) were also loaded with the chemotherapeutic drug busulfan [26], which is used in high doses as a conditioning agent prior to stem cell transplantation. Our aim was to construct a drug delivery system able to efficiently entrap and release lipophilic anticancer drugs and track cellular uptake in vitro as well as the biodistribution in vivo via noninvasive MRI. Such a vehicle is of significant value for diagnosis and therapy of cancer, where simultaneous drug delivery and therapeutic efficacy follow up is needed. 2. Materials and methods 2.1. Chemicals Poly(lactic-co-glycolic acid) (PLGA) with the brand name PURASORBÒ PDLG 5002A (molecular weight ca. 15 kDa), terminated with carboxylic acid and having a ratio of 50/50 for DL-lactide/glycolide, was obtained from Purac Biomaterials, Gorinchem, the Netherlands. Sodium oleate, ferric chloride hexahydrate (FeCl3$6H2O), n-hexane, octyl ether, dichloromethane, manganese chloride (MnCl2), stearic acid (SA), tetramethylammonium hydroxide (TMAOH), zinc acetate dihydrate (ZnAc2), sulfur, oleylamine (OLA), octadecene (ODE), 1-dodecanethiol, and PVA were purchased from Sigma Aldrich, Munich, Germany and used without any further purification. 2.2. Synthesis of SPION, Mn:ZnS QDs, and PLGAeSPIONeMn:ZnS nanoparticles Monodisperse SPION were synthesized by thermal decomposition of a Fe-oleate complex in octyl ether at approximately 297  C in the presence of oleic acid according to a previously reported method [27]. The Fe3O4 nanocrystals were stabilized with oleic acid and dispersed in dichloromethane at a concentration of 9.1 mg/ mL Fe. Mn:ZnS QDs were synthesized by a nucleation-doping strategy [28]. First, manganese stearate (MnSt2) was prepared by dropwise addition of methanolic MnCl2 solution into a mixture of SA and TMAOH in methanol [29]. A mixture of MnSt2 and 1-dodecanethiol in ODE was then degassed at 100  C for 15 min, followed by the addition of sulfur and ZnAc2 in sequence at 250  C. The Mn:ZnS nanoparticles thus obtained were washed against acetone and finally re-dispersed in dichloromethane. Dichloromethane solutions of PLGA, SPION and Mn:ZnS were mixed with PVA aqueous solution (1:20 oil to water ratio) using a probe-type sonicator to form an emulsion, which was agitated overnight to evaporate the organic solvent and washed against de-ionized (DI) water (15 MU cm) to collect PLGAeSPIONeMn:ZnS nanoparticles. The PBS suspension of these particles was deposited on a copper grid and positively stained for TEM examination using a 2% aqueous solution of phosphotungstic acid (H3PW12O40). 2.3. Characterization of nanoparticles The morphology and elemental composition of SPION, Mn:ZnS, and PLGAe SPIONeMn:ZnS nanoparticles were characterized by JEM-2100F field emission transmission electron microscope (FE-TEM) operating at an accelerating voltage of 200 kV. The hydrodynamic size of the particles was measured by dynamic light scattering (DLS) (DelsaÔNano particle size analyzer, Beckman Coulter, Brea, CA, USA). The magnetization measurements were performed using a vibrating sample magnetometer (VSM-NUOVO MOLSPIN, Newcastle-upon-Tyne, UK). The optical

absorbance and fluorescence intensity of Mn:ZnS and PLGAeSPIONeMn:ZnS nanoparticles were measured by Lambda 900 UVeViseNIR spectrometer (Perkin Elmer, Waltham, MA, USA) and LS 55 Fluorescence spectrometer (Perkin Elmer, Waltham, MA, USA), respectively. Electron paramagnetic resonance (EPR) measurement of Mn:ZnS was done on a Bruker ELEXSYS EPR spectrometer (X-band, 9e 10 GHz) at 113 K (Bruker, Billerica, MA, USA). Concentrations of iron, manganese and zinc in samples were measured by Thermo Scientific iCAP 6500 inductively coupled plasma atomic emission spectroscopy (ICP-AES) (Thermo Fisher Scientific, Kungens Kurva, Sweden). 2.4. In vitro phantom magnetic resonance imaging Phantoms (10 mL) of PLGAeSPIONeMn:ZnS nanoparticles with 0.05 mM, 0.1 mM, 0.2 mM, 0.3 mM, 0.4 mM, 0.5 mM, and 1 mM of iron were made by mixing the nanoparticle suspension with agarose gel (3 wt%) in DI water at 85  C and letting it cool down naturally overnight in 50 mL Eppendorf centrifuge tubes. The phantoms were placed in the extremity coil of a 3 T MRI scanner (Siemens Trio, Siemens, Erlangen, Germany). A gradient echo T2* sequence with a fixed repetition time (TR) of 2000 ms and 12 TEs of 2e22.9 ms was used for MR imaging to obtain T2* -weighted images. Circular ROIs (region of interest) were placed manually on the images and the negative logarithmic values of the signal intensities at different TEs were plotted versus the respective TE values. The T2* relaxation time was calculated as the slope of a semi-log plot of the signal intensities versus the TEs. In phantoms with a high concentration of iron oxide, the calculations were based on fewer TEs excluding those TEs where full transaxial relaxation had already occurred. 2.5. In vitro cellular uptake and fluorescence imaging To evaluate the effects of cellular uptake for PLGAeSPIONeMn:ZnS nanoparticles, we used the murine J774A macrophage cell line (European Type Tissue Culture Collection, CAMR, Salisbury, UK). These cells were obtained as a kind gift from Professor Carmen Fernandez, Department of Immunology, Wenner-Gren Institute, Stockholm University, Stockholm, Sweden. First, the J774A cells were cultured in DMEM supplemented with 10% heat-inactivated fetal bovine serum, penicillin (100 U/mL) and streptomycin (100 mg/mL) (InvitrogenÔ, Life Technologies, Carlsbad, CA, USA) in a 50 cm2 tissue culture flask (Costar, Corning, NY, USA). The cultures were maintained at 37  C in a humidified atmosphere containing 5% carbon dioxide. J774A cells were then cultured in 8-chamber polystyrene vessel tissue culture treated glasses at a density of 5  105 cells/chamber, at 37  C, for 12 h in an atmosphere containing 5% carbon dioxide to allow cell attachment. Thereafter, the cell culture medium was aspirated from each chamber and substituted with the medium alone (negative control) or the same medium containing PLGAeSPIONe Mn:ZnS nanoparticles at concentrations of 1000, 100, 50, 25, 12.5, 6.25 and/or 3 mg/ mL. Chambers were then incubated at 37  C for 1, 2, 4 and 24 h in an atmosphere containing 5% carbon dioxide. The uptake experiment was terminated at each time point by aspirating the test samples, removing the chamber and washing the cell monolayers with ice-cold PBS three times. Each slide was then fixed with methanoleacetone (1:1, v/v), followed by examination under a Nikon Eclipse i80 fluorescence microscope (Nikon, Tokyo, Japan) at a wavelength of 520 nm. 2.6. In vitro drug release For in vitro busulfan release experiments, 30 mg busulfan was dissolved in dichloromethane solution containing PLGA, SPION and Mn:ZnS QDs, and then emulsified with PVA at a total volume of 6 mL. After evaporation of organic solvent and centrifugation to wash off unloaded drugs, PLGA vesicles containing drugs were transferred into a cellulose permeable membrane bag with a molecular weight cutoff (MWCO) of 12e14 kDa to dialyze against PBS solution at 37  C. Entrapment efficiency of busulfan in PLGAeSPIONeMn:ZnS nanoparticles was calculated as [(mass of the total drug
mass of free drug)  100%/mass of total drug]. Three parallel release experiments were conducted and samples were taken at specific time points. Concentrations of busulfan released in dialysis media, left in dialysis bag or left in centrifuged supernatant were measured by gas chromatography (SCION 436-GC; Bruker, Billerica, MA, USA) with electron capture detector (ECD) according to a method reported previously by Hassan et al. [30]. The release percentage of loaded busulfan is averaged from the three parallel experiments with error bars representing standard deviation. 2.7. In vitro degradation of PLGA vesicles The synthesized PLGA vesicles were placed in a cellulose permeable membrane bag (MWCO 12e14 kDa) and dialyzed against 1 L PBS (pH 7.4) at 37  C. At predetermined time intervals, a 5 mL aliquot of PBS solution was withdrawn and fresh PBS was added into the dialysis solution. The concentration of lactic acid released in PBS was measured by high-performance liquid chromatography (HPLC). The HPLC system consisted of a Gilson autoinjector (100 mL loop), an LKB HPLC pump 2150 (Pharmacia Inc., Sweden), an LDC analytical spectromonitor 3200 UV detector (Riviera Beach, FL, USA) and a CSW 32 chromatography station integrator. Separation was performed on a Zorbax SB-CN column (4.6 mm  150 mm; 5 mm) from Agilent Technologies (Santa Clara, CA, USA), and the column was maintained at room temperature during analysis. The mobile phase was composed of NH4H2PO4 (0.1 M,

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(70%). Further dehydration and paraffinization of the selected tissue were performed in a vacuum infiltration processor before they were embedded in paraffin according to RENI trimming guidelines [31]. Sections (4 mm) were mounted on superfrost glass slides and stained with Perls’ Prussian blue staining for iron. In addition, immunohistochemistry was performed with a primary antibody against CD68 macrophage marker (MCA 341, AbD Serotec, Kidlington, UK) in the lung, liver, and spleen sections. Biotinylated rabbit anti-mouse IgG (E0354, Dako, Stockholm, Sweden) was used as a secondary antibody. The sections were then treated by DAB peroxidase substrate (SK-4100, Vector, Burlingame, CA, USA). CD68 brown signal was developed using VECTASTAINÒ ABC kit (PK6100, Vector, Burlingame, CA, USA). To produce counterstaining, immunohistochemistry was performed followed by additional Prussian blue staining. The slides were cover slipped and mounted with Pertex mounting medium (177070, Histolab, Spånga, Sweden).

3. Results 3.1. Synthesis and characterization of PLGAeSPIONeMn:ZnS nanoparticles Fig. 1. Schematic diagram showing the composition of PLGAeSPIONeMn:ZnS nanoparticles and their multiple applications.

pH 2.5) and flow rate was set at 0.8 mL/min with a running time of 6 min. Injection volume was 50 mL. Detection of lactic acid was performed at 210 nm compared to a calibration curve for lactic and glycolic acids. Lactic acid eluted after 4.0 min. 2.8. In vivo magnetic resonance imaging The animal study was approved by the Stockholm Southern Animal Research Ethics Committee and was performed in accordance with Swedish Animal Welfare law. Two male Sprague Dawley rats (500 g, Charles River Laboratories, Germany) were anaesthetized using an intraperitoneal (I.P.) injection of 60 mg/kg sodium pentobarbital (APL, Kungens Kurva, Sweden). The rats were put head first in an extremity coil and imaged in the coronal plane at 3 T in a Siemens Trio MRI scanner (Siemens, Erlangen, Germany), pre- and post-intravenous (I.V.) injections of 0.85 mL of PLGAeSPIONeMn:ZnS nanoparticles suspended in the aforementioned DMEMfetal bovine serum-penicillin-streptomycin media ([Fe] ¼ 0.25 mg/mL). Imaging was performed using a T2* -weighted sequence before and after injection at 4, 8, 12 and 30 min as well as at 1, 2, 3, 4 and 24 h. The images were evaluated using a PACS workstation (Sectra, Linköping, Sweden) and R*2 ð1=T2* Þ was determined for liver, spleen, brain and kidneys by measuring the signal intensity in circular ROIs on the images at different TEs and then calculating T2* using several TEs. 2.9. Histological examination and tissue distribution Rats were injected intravenously with 0.85 mL of PLGAeSPIONeMn:ZnS nanoparticles suspended in the aforementioned media ([Fe] ¼ 0.25 mg/mL). The animals were sacrificed at 30 min, 4 h and 24 h post-injection by injecting an overdose of pentobarbital. Various organs such as liver, spleen, and lungs were dissected and fixed in paraformaldehyde (PFA, 4%) for 48 h followed by ethanol

PLGAeSPIONeMn:ZnS nanoparticles were synthesized by an emulsioneevaporation process, and the composition of the particles is shown in Fig. 1. First, hydrophobic SPION and Mn:ZnS QDs to function as imaging agents were fabricated separately. Next, PLGA vesicles entrapping the payloads of imaging agents and busulfan were prepared by an oil-in-water (O/W) emulsion method followed by solvent evaporation of the volatile organic phase at room temperature. Specifically, the SPION, Mn:ZnS QDs and busulfan were incorporated into the hydrophobic domain of PLGA molecules via hydrophobic interaction, and the PLGA vesicles were then formed in the presence of polyvinyl alcohol (PVA) emulsifier. After evaporation of the organic solvent in the emulsion, the PLGA vesicles entrapping SPION and Mn:ZnS QDs with busulfan were washed using de-ionized water and re-dispersed in phosphate buffer solution (PBS) or in Dulbecco’s modified Eagle medium (DMEM). The morphology and crystal structure of SPION, Mn:ZnS, and PLGA vesicles were examined by transmission electron microscopy (TEM). According to the TEM images of nanoparticles (Fig. 2a, b), the average diameter was 10.7 nm (standard deviation s z 8%) for SPION, 3.1 nm (s z 10%) for Mn:ZnS QDs and 93 nm (s z 20%) for PLGAeSPIONeMn:ZnS vesicles. High resolution TEM images (Fig. 2c, d) show the single crystalline nature of SPION and Mn:ZnS QDs, respectively. The loading of these particles in PLGA vesicles can be seen clearly in Fig. 2e. Images of unstained (Fig. S1a) and large area (Fig. S1b) PLGA vesicles as well as the EDS spectrum (Fig. S2) can be found in Supplementary Data.

Fig. 2. Field emission transmission electron microscope (FE-TEM) images of PLGA nanoparticles. (a) SPION and (b) Mn:ZnS QDs with inset of electron diffraction patterns, and high resolution image of a single (c) SPION and (d) Mn:ZnS QD; (e) TEM image of positively stained PLGAeSPIONeMn:ZnS nanoparticles.

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Fig. 3. The magnetic property of SPION and the elemental analysis of PLGAeSPIONeMn:ZnS. (a) Field-dependent magnetization measurement of SPION at room temperature. (b) EPR spectra of Mn:ZnS QDs. (c) Line scan for elemental analysis and (d) UVeVis absorbance and fluorescence spectra of PLGAeSPIONeMn:ZnS nanoparticles.

The magnetic property of SPION was examined on fielddependent magnetization and no hysteresis appeared (Fig. 3a), demonstrating the superparamagnetic property desired for T2 and T2* MRI application. The optical property of Mn:ZnS QDs was also studied. Fig. S3a shows the UVeVisible absorbance and fluorescence spectra. The characteristic peak of fluorescence emission at 594 nm is due to doping Mn2þ ions into a ZnS lattice (3 wt% of Mn to Zn according to ICP results). In order to understand the location of Mn ions in the ZnS host matrix, we recorded the electron paramagnetic resonance (EPR) spectrum of Mn:ZnS nanocrystals (Fig. 3b). The observed six-line spectrum resulted from the hyperfine interaction of the unpaired electrons with 55Mn nuclear spin (I ¼ 5/2), which provided evidence for well-dispersed doping of Mn2þ ions without any clustering. The hyperfine coupling constant of Mn:ZnS, A ¼ 68.8 G, determined from the EPR spectrum, is in close agreement with the literature data for Mn2þ ions in the tetrahedral sites of cubic zinc blende lattice [32,33]. The elemental composition of PLGAe SPIONeMn:ZnS vesicles was analyzed by energy dispersive spectroscopy (EDS) and inductively coupled plasma (ICP) techniques. A typical EDS line scan on one PLGA vesicle (Fig. 3c) shows the spectral counts corresponding to the elements of Fe, O, Zn, S, and Mn. Originated from the loaded Mn:ZnS QDs, the fluorescence emission peak of PLGAeSPIONeMn:ZnS is centered at 599 nm (Fig. 3d) which is located in the observable window of fluorescence microscopy and facilitates further imaging application.

3.2. In vitro phantom magnetic resonance imaging An aqueous suspension of PLGAeSPIONeMn:ZnS vesicles was used to prepare phantoms for MRI measurement. The hydrodynamic size distribution of these vesicles was studied and the results

Fig. 4. Magnetic resonance imaging of phantoms. (a) T2* -weighted MR phantom images of PLGAeSPIONeMn:ZnS at different TEs (TR ¼ 1200 ms; TE ¼ 2 ms, 3.9 ms, 5.8 ms). (b) Proton transverse relaxation rate ðR*2 ¼ 1=T2* Þ of phantom samples versus iron concentration.

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Fig. 5. Superimposed light and fluorescence microscopy images of Quantum Dots (QD) labeled PLGAeSPIONeMn:ZnS nanoparticles in the J774A cells. (a, b) untreated control J774A at 4 h and 24 h, (c, d) J774A incubated with PLGAeSPIONeMn:ZnS nanoparticles for 4 h and 24 h (e, f) The corresponding fluorescence imaging of J774A incubated with PLGAe SPIONeMn:ZnS-QD nanoparticles for 4 h and 24 h.

are shown in Fig. S4. The MRI contrasting effect of PLGAeSPIONe Mn:ZnS was then evaluated by T2* -weighted MR images using a 3 T instrument. With increasing concentrations of PLGA vesicles, the signal intensity of MRI decreased owing to the increase in loaded SPION (Fig. 4a). For PLGA vesicle phantoms with the same concentrations of iron oxide, the increase in echo time (TE; in Fig. 4a and Fig. S5a) and flip angle (Fig. S5b) also induce decreased signal intensity for MRI. It is noted that the signal intensity decreases more rapidly with increasing TE for phantoms containing higher concentrations of iron oxide. These characteristics allow the application of PLGAeSPIONeMn:ZnS as a negative contrast agent for MRI. In Fig. 4b, the transverse relaxivity (r2* ) is obtained as the slope of linear fitting for the relaxation rate at different iron concentrations, which is 523 s
1 mM
1 Fe for PLGAeSPIONeMn:ZnS vesicles.

SPIONeMn:ZnS nanoparticles in the J774A murine macrophage cell line, in virtue of the intrinsic fluorescence property of loaded Mn:ZnS QDs. Fig. 5a, b shows the superimposed optical and fluorescence imaging of non-treated J774A cells at 4 h and 24 h incubation, respectively. In comparison, the cells treated with PLGAe SPIONeMn:ZnS nanoparticles exhibited much stronger fluorescence intensity in the cell plasma area at the same incubation time of 4 h (Fig. 5c) and 24 h (Fig. 5d). The corresponding fluorescence imaging of the treated cells seen in Fig. 5e, f shows that the uptake of PLGA vesicles can greatly enhance visualization of the cells compared with non-treated cells with very low fluorescence intensity (data not shown). Maximum fluorescence intensity for treated cells was achieved after 24 h incubation due to continuous localization of PLGA vesicles in the cells, demonstrating their high uptake efficiency.

3.3. In vitro cellular uptake and fluorescence imaging

3.4. In vitro drug release and degradation of PLGA vesicles

Fig. S6(aec) in Supplementary Data shows the fluorescence images of Mn:ZnS QDs with different emission wavelengths using a set of filters. We then evaluated the cell uptake effect of PLGAe

The drug release kinetics of busulfan-loaded PLGAeSPIONe Mn:ZnS vesicles was studied at pH 7.4 using a dialysis method; the release profile versus time is demonstrated in Fig. 6. The initial

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liver. The signal continues to decrease and the imaged liver becomes very dark, especially on the images taken with longer TEs. Minimal signal intensity appears at 4 h post-injection and remains low until 24 h post-injection. In Fig. 9, the relaxation rate of the liver is plotted against post-injection time. It quickly increases from 64 s
1 pre-injection to 152 s
1 8 min post-injection, with its highest value being 161 s
1 4 h post-injection, after which it slowly decreases to 137 s
1 24 h post-injection. Fig. 10 shows the T2* -weighted MR images of kidney slices. In these, a smaller signal decrease can be observed over the 24 h post-injection period compared with liver slices. However, no change in signal intensity was calculated in the brain and testes. 3.6. Histological examination and tissue distribution

Fig. 6. Profile of busulfan release from PLGAeSPIONeMn:ZnS nanoparticles in PBS solution at pH 7.4. Data represent average values for n ¼ 3, and the error bars indicate standard deviation.

concentration of busulfan in a mixture of dichloromethane and PBS solution was 5 mg/mL, and the entrapment efficiency of busulfan in PLGA vesicles was calculated to be 89  2%. The percentage of release, represented on the vertical axis, was calculated by dividing the amount of busulfan diffused into dialysis media by the total amount of busulfan loaded into the PLGA vesicles. We found that around 70e80% of busulfan was released after 5 h of dialysis. The degradation of the drug carrier was also tested and Fig. 7 shows the percentage of lactic acid degraded from PLGA vesicles at pH 7.4 and 37  C for a period of 5 weeks. At 2 weeks, about 12% lactic acid was degraded and released. The degradation was found to follow a quasilinear trend and around 32% lactic acid was degraded at 5 weeks. 3.5. In vivo magnetic resonance imaging The in vivo MRI of PLGA vesicles was tested in a rat model. Fig. 8 shows a coronal view of liver slices, where the liver appears white on T2* -weighted images before injection of PLGA vesicles. There was a rapid signal decrease in the liver and spleen post-injection. Just after 7 min post-injection the liver imaging becomes darker, indicating the fast accumulation of iron-containing PLGA vesicles in the

The pattern of PLGA vesicle biodistribution and uptake were studied histologically in the lung (Fig. 11aec) and liver (Fig. 11def). We employed Prussian blue staining to detect PLGA vesicles through their iron content. As shown in Fig.11a and d, clustered PLGA vesicles were observed in the lungs and liver respectively as early as 30 min post-injection. The number of clustered PLGA vesicles in pulmonary macrophages reached their maximum at 4 h post-injection (Fig.11b) and then decreased until 24 h post-injection, when PLGA vesicles could rarely be observed (Fig. 11c). PLGA vesicles were also found in liver tissue at 30 min post-injection (Fig. 11d), but the amount of vesicles remained similar at 4 h and 24 h post-injection with only a minor decline (Fig. 11e and f, respectively). We confirmed the phagocytosis of PLGA vesicles by macrophages through applying immunohistological staining for macrophages followed by additional Prussian blue staining in liver (Fig. 12aec) and spleen (Fig. 12def) sections. In the liver, clusters of PLGA vesicles were found in the sinusoids at 30 min post-injection, mostly associated with macrophages (Fig. 12a). At 4 h post-injection (Fig. 12b), cluster frequency was similar to that seen at 30 min post-injection. Meanwhile, the PLGA vesicle clusters were clearly less frequent at 24 h post-injection, as seen in Fig. 12c. In the spleen, fewer PLGA vesicles were observed from 30 min post-injection (Fig. 12d) and increased in number at 4 h to reach a maximum at 24 h post-injection (Fig. 12e and f, respectively). The frequency of vesicles at 4 h post-injection in the spleen was higher compared to that seen at 30 min postinjection; however, the number of clusters increased significantly at 24 h post-injection, which indicates an increased amount of PLGA vesicles accumulated in the spleen over time. 4. Discussion

Fig. 7. Release profile of lactic acid by degradation of PLGAeSPIONeMn:ZnS nanoparticles in PBS solution during a period of 5 weeks (pH 7.4, T ¼ 37  C). Data were collected from three parallel experiment and error bars indicate standard deviation.

The aim of the present investigation was to construct a biodegradable drug delivery system that is easy to image using MRI in order to study its biodistribution in vivo. The nano-carrier was also constructed to be utilized in cellular uptake as well as for histopathological studies, as illustrated in Fig. 1. To optimize the particle size and morphology of PLGA vesicles for drug delivery application, we investigated the influence of organic solvents (oil phase) and surfactants (emulsifier) on formation of these vesicles during an emulsion-evaporation process. We found that a water-immiscible solvent (e.g., dichloromethane) is more effective than water-miscible (e.g., tetrahydrofuran, ethanol, acetone) or partially miscible (e.g., ethyl acetate) solvents in forming separated and spherical polymeric vesicles. Non-ionic surfactant (e.g., PVA) was found to be superior to ionic surfactants (e.g., cetyltrimethylammonium bromide, sodium dodecyl sulfate) in forming high quality monodisperse PLGAeSPIONe Mn:ZnS nanoparticles in a water-dichloromethane emulsion system; however, the previously reported method [34] using Pluronic F127 was not reproduced in the current work.

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Fig. 8. T2* -weighted in vivo MR images of rat before and after intravenous injection of PLGAeSPIONeMn:ZnS nanoparticles at a dose of 0.42 mg Fe/kg. Images were taken preinjection and at 7 min, 2 h 40 min, 4 h, and 24 h post-injection with different TEs (2 ms, 3.9 ms, and 5.8 ms). The regions of liver in the coronal planes are encircled by white dashed lines.

For manganese-doped QDs, once Mn2þ ions are doped inside a semiconductor host the strong electronic interaction between d states of Mn2þ and sep states of the host generates intermediate energy states, through which the host generated exciton relaxes (4T1/6A1) resulting in emission centered at 595 nm [35,36]. Based on the hyperfine coupling constant of Mn:ZnS QDs, if Mn2þ ions

Fig. 9. Variation of relaxation rate (R*2 , s
1) of PLGAeSPIONeMn:ZnS nanoparticles in the liver at different time points pre- and post-injection.

locate on the surface of a host particle a much higher value of hyperfine splitting is expected [37], which is not valid here. However, additional peaks at the lower and higher magnetic field were observed (marked with arrows in Fig. 3b) indicating the presence of a small fraction of Mn2þ ions on the surface of ZnS nanoparticles. The encapsulated SPION are responsible for MRI contrast enhancement. They are well known for their capability to shorten the transverse (spinespin) relaxation time, resulting in a decrease in MR signal intensity [38]. The transverse relaxivity (r2* ), i.e. the changes in relaxation rate ðR*2 ¼ 1=T2* Þ per unit concentration, is the assessment for efficacy of MRI T2* contrast enhancement. The value of transverse relaxivity for PLGAeSPIONeMn:ZnS vesicles, 523 s
1 mM
1 Fe, is much higher than that seen in previous reports for single SPION [39], clusters [34] or assembly [40] of SPION, indicating the high efficiency of PLGAeSPIONeMn:ZnS for MR T2* imaging. This can be attributed to the synergetic effect obtained when SPION are in intimate contact, thereby enhancing the local magnetic field. To investigate the interaction between PLGA vesicles and cells by fluorescence imaging, we used nontoxic zinc sulfide QDs. By doping a small amount of manganese into ZnS QDs, the fluorescence emission peak was tuned from violet blue (approximately 400 nm) to the red end of the visible spectrum (approximately 595 nm), fitting with the observable range of emission filters for fluorescence microscopy. The increase in fluorescence intensity of PLGA vesicle-treated cells until 24 h incubation provides the possibility of long-term imaging for tracking and observation purposes. These results indicate the capacity of PLGA vesicles for delivery of chemotherapeutic agents into the cells as well as their usefulness as an optical imaging agent.

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Fig. 10. T2* -weighted in vivo MRI of coronal slices taken at pre-injection and at 7 min, 45 min, 1 h, 3 h, 4 h, and 24 h post-injection of PLGAeSPIONeMn:ZnS nanoparticles. The two kidneys are encircled by white dashed lines.

Simple entrapment of payload in PLGA vesicles allows loading and delivery of lipophilic drugs for specific cancer treatment. In this study, we have demonstrated the validity of using PLGA vesicles for loading and release of the anticancer drug, busulfan. Currently, its main application is for conditioning prior to bone marrow transplantation [26,41]. Unlike other systems, where the drug release shows a burst model at the beginning of the process, our results showed a relatively constant release rate during the first 2 h, after which drug release continued at a lower rate for 5e6 h to release a total of 70e80% of the loaded drugs. This behavior might be related to the amphiphilic nature of PLGA nanoparticles and coating effect of PVA layers. Preferably, a steady zero-order release profile needs to be built by employing surface coating or other precise controls of the drug release rate. This would prevent side effects due to a drug

release burst from polymers as well as maintain the drug plasma level during conditioning to ensure success of the following engraftment. To investigate the stability of PLGA vesicles used as a drug carrier, we tested the degradation of PLGA vesicles by measuring the concentration of lactic acid, one of the degradation products. This relatively slow degradation behavior of PLGA vesicles suggests their potential for use in long-term drug release. We further investigated the effects of PLGAeSPIONeMn:ZnS vesicles on in vivo MR imaging using a rat model. T2* -weighted MR imaging of the rat was performed at different time points after intravenous tail injection of PLGA vesicles. In a comparison of contrast efficiency, our PLGA vesicles provided higher relaxation rate by using a dose of iron 20 times lower than the previously reported dextran-coated SPION [42]. The rapid decrease of signal

Fig. 11. Tissue distribution of nanoparticles in rat (aec) lungs and (def) liver visualized by Perls’ Prussian blue iron staining. Histological sections were extracted at (a, d) 30 min, (b, e) 4 h, and (c, f) 24 h after intravenous injection of PLGAeSPIONeMn:ZnS nanoparticles. The scale bars represent 20 mm.

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Fig. 12. Combination of Immunohistochemical (IHC) and Prussian blue staining on liver and spleen. (aec) liver and (def) spleen sections removed at (a, d) 30 min, (b, e) 4 h, and (c, f) 24 h post-injection of PLGAeSPIONeMn:ZnS nanoparticles. The scale bars represent 20 mm.

intensity or faster T2* relaxation in the liver compared to other organs is due to the PLGAeSPIONeMn:ZnS nanoparticles being taken up by hepatic macrophages [43]. Other organs had fewer macrophage cells able to take up the nanoparticles, and the signal intensity therefore exhibited less or no change. We used the histological method to study the biodistribution of PLGA vesicles in the rat in order to understand the biological fate of these vesicles in lung, liver and spleen. The result of histological studies on lung and liver is consistent with that of in vivo MRI for liver slices, where the signal decreased rapidly at the beginning and reached its lowest level at 4 h post-injection due to a large amount of PLGA accumulated in the liver. Alternatively, the biodistribution and uptake of PLGA vesicles can also be studied in macrophages treated with immunohistological staining. The distribution of vesicles was found to gradually shift from liver to spleen, which is also seen in previous reports on nanoparticle biodistribution. Panagi et al. have reported that the dose affects the biodistribution of PLGA particles between blood and the mononuclear phagocyte system, while no dose effect on the biodistribution was observed when stealth poly(Lactide-co-glycolide)-monomethoxypoly(ethyleneglycol) (PLGAmPEG) was used [44]. In the present investigation, the PLGA particles were rapidly cleared from the lungs to the liver and within 24 h were found in spleen. No sign of higher uptake into the lungs or distribution to the brain tissues was observed as was reported with PLGAePEGePLGA [22]. These results show that the present construction of PLGA differs from other nanoparticles and suggest that the present vesicles might be used as drug carriers in therapy as well as for monitoring the cellular uptake in order to follow treatment efficacy. 5. Conclusions In summary, we have synthesized PLGAeSPIONeMn:ZnS vesicles as a multifunctional drug delivery system consisting of a biodegradable polymeric shell containing a payload of multiple imaging agents and an anti-cancer drug. The MRI phantom of PLGA vesicles exhibits high r2* relaxivity and greatly enhanced T2* -weighted MR imaging contrast. J774A macrophage cells showed high uptake of PLGA vesicles labeled with quantum dots, and the uptake was improved over time. The quantum dots enhanced the

fluorescence visualization of cell uptake. PLGA vesicles have also been demonstrated to have high entrapment efficiency for the lipophilic drug busulfan as well as sustained drug release. The in vivo MRI and histological studies in a rat model elucidated the biodistribution of PLGA vesicles. The biodistribution to different tissues was confirmed using immunohistochemistry and Prussian blue staining. Altogether, these results strongly suggest further development of PLGA as a multifunctional diagnostic and therapeutic tool for cancer treatment. Further investigation of a PLGA drug delivery system for cancer treatment is ongoing, including organ targeting, controlling of drug release and in vivo evaluation of the treatment efficacy of chemotherapy in tumor models. Acknowledgments The authors thank Dr. T. Astlind (Dept. of Biophysics, Arrhenius laboratories, Stockholm University) for EPR measurement. Financial support from The Swedish Cancer Society, The Swedish Childhood Cancer Foundation and the NPST program of King Saud University (Project No: 11-NAN1462-02) is also acknowledged. Appendix A. Supplementary data Supplementary data related to this article can be found at http:// dx.doi.org/10.1016/j.biomaterials.2014.01.041 References [1] Hrkach J, Von Hoff D, Ali MM, Andrianova E, Auer J, Campbell T, et al. Preclinical development and clinical translation of a PSMA-targeted docetaxel nanoparticle with a differentiated pharmacological profile. Sci Transl Med 2012;4:128ra39. [2] Kam NWS, O’Connell M, Wisdom JA, Dai H. Carbon nanotubes as multifunctional biological transporters and near-infrared agents for selective cancer cell destruction. Proc Natl Acad Sci U S A 2005;102:11600e5. [3] Medina SH, El-Sayed MEH. Dendrimers as carriers for delivery of chemotherapeutic agents. Chem Rev 2009;109:3141e57. [4] Yang F, Jin C, Subedi S, Lee CL, Wang Q, Jiang Y, et al. Emerging inorganic nanomaterials for pancreatic cancer diagnosis and treatment. Cancer Treat Rev 2012;38:566e79. [5] Huang X, El-Sayed IH, Qian W, El-Sayed MA. Cancer cell imaging and photothermal therapy in the near-infrared region by using gold nanorods. J Am Chem Soc 2006;128:2115e20.

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[6] Kawano T, Niidome Y, Mori T, Katayama Y, Niidome T. PNIPAM gel-coated gold nanorods for targeted delivery responding to a near-infrared laser. Bioconjug Chem 2009;20:209e12. [7] Scott RC, Crabbe D, Krynska B, Ansari R, Kiani MF. Aiming for the heart: targeted delivery of drugs to diseased cardiac tissue. Expert Opin Drug Deliv 2008;5:459e70. [8] Shi M, Ho K, Keating A, Shoichet MS. Doxorubicin-conjugated immunonanoparticles for intracellular anticancer drug delivery. Adv Funct Mater 2009;19:1689e96. [9] Chen R, Zheng X, Qian H, Wang X, Wang J, Jiang X. Combined near-IR photothermal therapy and chemotherapy using gold-nanorod/chitosan hybrid nanospheres to enhance the antitumor effect. Biomater Sci 2013;1:285e93. [10] Kumar CSSR, Mohammad F. Magnetic nanomaterials for hyperthermia-based therapy and controlled drug delivery. Adv Drug Deliv Rev 2011;63:789e808. [11] You J, Zhang R, Zhang G, Zhong M, Liu Y, Van Pelt CS, et al. Photothermalchemotherapy with doxorubicin-loaded hollow gold nanospheres: a platform for near-infrared light-trigged drug release. J Control Release 2012;158:319e28. [12] Nicolas J, Mura S, Brambilla D, Mackiewicz N, Couvreur P. Design, functionalization strategies and biomedical applications of targeted biodegradable/ biocompatible polymer-based nanocarriers for drug delivery. Chem Soc Rev 2013;42:1147e235. [13] Arias JL, Clares B, Morales ME, Gallardo V, Ruiz MA. Lipid-based drug delivery systems for cancer treatment. Curr Drug Targets 2011;12:1151e65. [14] Vallet-Regí M, Balas F, Arcos D. Mesoporous materials for drug delivery. Angew Chem Int Ed 2007;46:7548e58. [15] Huang S, Fu X. Naturally derived materials-based cell and drug delivery systems in skin regeneration. J Control Release 2010;142:149e59. [16] Lasic DD. Doxorubicin in sterically stabilized liposomes. Nature 1996;380: 561e2. [17] Brannon-Peppas L, Blanchette JO. Nanoparticle and targeted systems for cancer therapy. Adv Drug Deliv Rev 2004;56:1649e59. [18] Danhier F, Ansorena E, Silva JM, Coco R, Le Breton A, Préat V. PLGA-based nanoparticles: an overview of biomedical applications. J Control Release 2012;161:505e22. [19] Hevus I, Modgil A, Daniels J, Kohut A, Sun C, Stafslien S, et al. Invertible micellar polymer assemblies for delivery of poorly water-soluble drugs. Biomacromolecules 2012;13:2537e45. [20] Seremeta KP, Chiappetta DA, Sosnik A. Poly(ε-caprolactone), EudragitÒ RS 100 and poly(ε-caprolactone)/EudragitÒ RS 100 blend submicron particles for the sustained release of the antiretroviral efavirenz. Colloids Surf B Biointerfaces 2013;102:441e9. [21] Gu B, Xie C, Zhu J, He W, Lu W. Folate-PEG-CKK2-DTPA, a potential carrier for lymph-metastasized tumor targeting. Pharm Res 2010;27:933e42. [22] Song Z, Feng R, Sun M, Guo C, Gao Y, Li L, et al. Curcumin-loaded PLGA-PEGPLGA triblock copolymeric micelles: preparation, pharmacokinetics and distribution in vivo. J Colloid Interface Sci 2011;354:116e23. [23] Jun YW, Lee JH, Cheon J. Chemical design of nanoparticle probes for highperformance magnetic resonance imaging. Angew Chem Int Ed 2008;47: 5122e35. [24] Resch-Genger U, Grabolle M, Cavaliere-Jaricot S, Nitschke R, Nann T. Quantum dots versus organic dyes as fluorescent labels. Nat Meth 2008;5:763e75. [25] Lee D. Highly effective T2 MR contrast agent based on heparinized superparamagnetic iron oxide nanoparticles. Macromol Res 2011;19:843e7.

[26] Hassan M. The role of busulfan in bone marrow transplantation. Med Oncol 1999;16:166e76. [27] Park J, An K, Hwang Y, Park JG, Noh HJ, Kim JY, et al. Ultra-large-scale syntheses of monodisperse nanocrystals. Nat Mater 2004;3:891e5. [28] Zhang W, Li Y, Zhang H, Zhou X, Zhong X. Facile synthesis of highly luminescent Mn-doped ZnS nanocrystals. Inorg Chem 2011;50:10432e8. [29] Pradhan N, Peng X. Efficient and color-tunable Mn-doped ZnSe nanocrystal emitters: control of optical performance via greener synthetic chemistry. J Am Chem Soc 2007;129:3339e47. [30] Hassan M, Ehrsson H. Gas chromatographic determination of busulfan in plasma with electron-capture detection. J Chromatogr B Biomed Sci Appl 1983;277:374e80. [31] Morawietz G, Ruehl-Fehlert C, Kittel B, Bube A, Keane K, Halm S, et al. Revised guides for organ sampling and trimming in rats and mice e Part 3. A joint publication of the RITA and NACAD groups. Exp Toxicol Pathol 2004;55:433e 49. [32] Chin PTK, Stouwdam JW, Janssen RAJ. Highly luminescent ultranarrow Mn doped ZnSe nanowires. Nano Lett 2009;9:745e50. [33] Srivastava BB, Jana S, Karan NS, Paria S, Jana NR, Sarma DD, et al. Highly luminescent Mn-doped ZnS nanocrystals: gram-scale synthesis. J Phys Chem Lett 2010;1:1454e8. [34] Kim J, Lee JE, Lee SH, Yu JH, Lee JH, Park TG, et al. Designed fabrication of a multifunctional polymer nanomedical platform for simultaneous cancertargeted imaging and magnetically guided drug delivery. Adv Mater 2008;20:478e83. [35] Bhargava RN, Gallagher D, Hong X, Nurmikko A. Optical properties of manganese-doped nanocrystals of ZnS. Phys Rev Lett 1994;72:416e9. [36] Erwin SC, Zu L, Haftel MI, Efros AL, Kennedy TA, Norris DJ. Doping semiconductor nanocrystals. Nature 2005;436:91e4. [37] Nag A, Sapra S, Nagamani C, Sharma A, Pradhan N, Bhat SV, et al. A study of Mn2þ doping in CdS nanocrystals. Chem Mater 2007;19:3252e9. [38] Weissleder R, Moore A, Mahmood U, Bhorade R, Benveniste H, Chiocca EA, et al. In vivo magnetic resonance imaging of transgene expression. Nat Med 2000;6:351e4. [39] Ye F, Laurent S, Fornara A, Astolfi L, Qin J, Roch A, et al. Uniform mesoporous silica coated iron oxide nanoparticles as a highly efficient, nontoxic MRI T2 contrast agent with tunable proton relaxivities. Contrast Media Mol Imaging 2012;7:460e8. [40] Lee JE, Lee N, Kim H, Kim J, Choi SH, Kim JH, et al. Uniform mesoporous dyedoped silica nanoparticles decorated with multiple magnetite nanocrystals for simultaneous enhanced magnetic resonance imaging, fluorescence imaging, and drug delivery. J Am Chem Soc 2009;132:552e7. [41] Hamerschlak N, Lima M, Kerbauy F. Hematopoietic stem cell transplantation in elderly patients with myelodysplastic syndrome and acute myelogenous leukemia: use of busulfan/fludarabine for conditioning. In: Hayat MA, editor. Stem cells and cancer stem cells. Netherlands: Springer; 2013. pp. 263e9. [42] Kumar M, Yigit M, Dai G, Moore A, Medarova Z. Image-guided breast tumor therapy using a small interfering RNA nanodrug. Cancer Res 2010;70:7553e61. [43] Muthiah M, Park I-K, Cho C-S. Surface modification of iron oxide nanoparticles by biocompatible polymers for tissue imaging and targeting. Biotechnol Adv 2013;31:1224e36. [44] Panagi Z, Beletsi A, Evangelatos G, Livaniou E, Ithakissios DS, Avgoustakis K. Effect of dose on the biodistribution and pharmacokinetics of PLGA and PLGAmPEG nanoparticles. Int J Pharm 2001;221:143e52.