Ultra Magnetic Liposomes for MR Imaging, Targeting, and Hyperthermia

Article pubs.acs.org/Langmuir

Ultra Magnetic Liposomes for MR Imaging, Targeting, and Hyperthermia Gael̈ le Béalle,† Riccardo Di Corato,‡ Jelena Kolosnjaj-Tabi,‡,§ Vincent Dupuis,† Olivier Clément,§ Florence Gazeau,‡ Claire Wilhelm,‡ and Christine Ménager*,† †

Université Pierre et Marie Curie, UPMC-Univ Paris 06, Laboratoire PECSA-UMR 7195-CNRS-ESPCI, 4 place Jussieu, 75252 Paris cedex 05, France ‡ Laboratoire Matièreset Systèmes Complexes (MSC), UMR 7057 CNRS/Université Paris - Diderot, 10 rue Alice Domon et Léonie Duquet, 75205 Paris cedex 13, France § Université Paris Descartes, INSERM U970, Paris Cardiovascular Research Center − PARCC, 56 rue Leblanc, 75015 Paris, France S Supporting Information *

ABSTRACT: Magnetic liposomes offer opportunities as theranostic systems. The prerequisite for efficient imaging, tissue targeting or hyperthermia is high magnetic load of these vesicles. Here we describe the preparation of Ultra Magnetic Liposomes (UMLs), which may encapsulate iron oxide nanoparticles in a volume fraction of up to 30%. This remarkable magnetic charge provides UMLs with high magnetic mobilities, MRI relaxivities, and heating capacities for magnetic hyperthermia. Moreover, these UMLs are rapidly and efficiently internalized by cultured tumor cells and, when they are administered to mice, they can be vectorized to tumors by an external magnet.

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Néel relaxations) induced by an alternating magnetic field generate heat, essential to trigger drug release9,10 or to selectively destroy the adjacent cells.11,12 Few works recently demonstrated that hyperthermia could possibly be used on humans to induce a local heat (thermotherapy) in tumors, after local administration.13 In recent years, much research focused on magnetic carriers such as liposomes or polymersomes for drug encapsulation, allowing protection and delivery of the active substance. The latter may be stored either in the bilayer14−18 or in the aqueous core of vesicles.2,19−23 However, an efficient method that allows a high encapsulation efficacy of nanoparticles to improve their physical properties is yet to be found. In this paper, we suggest a new strategy to prepare Ultra Magnetic Liposomes (UMLs) suitable for systemic delivery characterized by an outstanding loading potential for magnetic nanoparticles. A high amount of MNPs in liposomes is critical for efficient MRI detection, magnetic targeting and heating of the sites of interest while minimizing the injected dose. We demonstrate here that the magnetic properties, MR relaxivities, magnetophoretic mobilities and heating capacity of UMLs make them unique nanoplatforms for controlled drug targeting and hyperthermia. Herein, not only do we evidence their excellent internalization by tumor cells in

INTRODUCTION During the last decades, significant headway has been made in cancer surgery, chemotherapy and radiotherapy. However, most therapeutic molecules cause systemic side effects and consequently limit the administered dose and the treatment efficiency. Recent advances in the chemistry of new materials for drug delivery have led to the development of triggerable systems for on-demand drug delivery.1,2 These materials could enhance therapeutic efficacy and reduce systemic toxicity by exploiting a stimuli-responsive release of the delivered drug. One way to achieve the synthesis of such triggerable systems, which could be used in both imaging and therapy (commonly designed as theranostics), is to combine synthetic nanoparticles with tunable physical properties and stimulus-sensitive materials. Magnetic nanoparticles (MNPs) based on iron oxide (maghemite γ-Fe2O3 or magnetite Fe3O4) are widely used in nanomedicine, both for their biocompatibility and superparamagnetic properties. These particles have overall sizes typically ranging between 5 and 20 nm and behave as a single magnetic domain, enabling magnetic resonance imaging (MRI) detection in vivo.3 Recently, many efforts have been made for the development of new classes of nanocontainers, able to encapsulate MNPs.4−7 Thus, the collective properties of these preparations, such as vesicles, can also allow the in vivo accumulation on a specific site of interest. This targeting can be achieved by the application of a static magnetic field in the vicinity of tumors or tissues.2,8 Once the targeting accomplished, the relaxation processes of nanoparticles (Brownian/ © 2012 American Chemical Society

Received: November 21, 2011 Revised: July 13, 2012 Published: July 16, 2012 11834

dx.doi.org/10.1021/la3024716 | Langmuir 2012, 28, 11834−11842

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Transmission Electron Microscopy (TEM). Liposomes were diluted 2000 times after synthesis and magnetically sorted during 2 days. A droplet was then deposited on a carbon coated copper grid and dried. Liposomes were characterized by a JEOL 100 CX Transmission Electronic at 60 keV. Cryogenic Transmission Microscopy (Cryo-TEM). Vitrified UMLs were prepared by deposition on mesh copper grids (TED PELLA). Film thickness was reduced using a blotting paper before quenching into liquid ethane with a vitrification system (LEICA EM CPC). Grids were stored in liquid nitrogen. Imaging was performed at 200 kV using a JEOL 2100 HC microscope with a LaB6 filament and a GIF TRIDIEM postcolumn energy filter. Images were recorded on a GATAN camera 2K × 2K. Magnetophoresis. Magnetophoretic mobilities of individual liposomes and magnetically labeled cells were measured in magnetic fields and field gradients of 180mT − 195 T/m and 145 mT − 17 T/ m created by a magnetized nickel rod (50 μm in diameter) and a permanent magnet respectively. From the balance between the viscous and magnetic forces experienced by the moving magnetic objects in the field gradient, a magnetic load is deduced and expressed as a MNPs volume fraction inside liposomes or as an iron content per cell (see a detailed description of the experimental setups in previous papers26,27). For each condition, three independent measurements were performed, each involving the tracking of 100 individual liposomes or cells and leading to the average magnetic load for each of them (as well as the distribution among the population). Error is expressed as the standard deviation between the averaged values for the three measurements. Squid Magnetometry. Magnetization measurements were performed on a Quantum Design Ltd. MPMS squid magnetometer. Samples were prepared as follows: 50 μL of solution was introduced in a PMMA cell sealed with a PMMA cap and nonmagnetic glue. The sample cells were fixed inside plastic straws, tightened to the positioning rod and inserted in the instrument at 250 K under a zero magnetic field to freeze the solutions. Field Cooled (FC) and Zero Field Cooled (ZFC) magnetizations were recorded in the range 5 K − 250 at 2 K/min for a 50 Oe probing field. Corrections for the diamagnetic response of sample cell and solvent were applied to the data. Hyperthermia Measurements. A laboratory-made device was used.8 It consists of a resonant RLC circuit, using a 16 mm coil producing an alternating magnetic field with a frequency ranging from 300 kHz to 1.1 MHz and with amplitudes up to 27 kA/m. Temperature was probed with a fluorooptic fiber thermometer and recorded every 0.7 s. The magnetic samples (Vs = 300 μL) are introduced in an eppendorf placed into the copper coil. The latter has a variable capacity in the range 10 pF−4 nF and a self-inductance of 25 μH. The coil was cooled with circulating water. Temperature of the water was controlled to obtain an equilibrium temperature of 30 ± 0.5 °C in the samples. Specific loss power (SLP) was calculated from the initial linear rise of temperature as a function of time (dT/dt):

vitro, but we also confirm magnetic vectorization to subcutaneous tumors in mice and we show that they can be efficiently detected by MRI both in vitro and in vivo.

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EXPERIMENTAL SECTION

Synthesis of MNPs. Magnetic fluid consists of an aqueous suspension of nanoparticles of maghemite (γ-Fe2O3) synthesized by alkaline coprecipitation of FeCl2 (0.9 mol) and FeCl3 (1.5 mol) salts, according to Massart’s procedure.24 Superparamagnetic maghemite grains were produced by oxidizing 1.3 mol of magnetite with 1.3 mol of iron nitrate under boiling. After magnetic decantation, volumes of 2 L of distilled water and 360 mL of HNO3 20% were added and the mixture was stirred for 10 min. Maghemite nanoparticles were washed several times with acetone (3 × 1 L) and ether (2 × 500 mL) and suspended in water. Aftewards, nanoparticles were sorted by size, by adding HNO3 (0.45 M) to the suspension followed by magnetic decantation. This operation was repeated with the deposit until suitable size was obtained. Sodium citrate at a molar ratio nFe/nCit = 0.13 was added to the sorted nanoparticles and the mixture was heated at 80 °C for 30 min to promote absorption of citrate anions onto their surface. Citrated nanoparticles were precipitated in acetone and suspended in water. At the end of the synthesis, volume fraction and average size of maghemite grains were determined by fitting the magnetization curve of MNPs using the Langevin’s Law. Two nanoparticles sizes were selected: 7 nm (d0 = 7 nm, polydispersity index σ = 0.35, volume fraction of nanoparticles in the suspension φ = 3.1%, specific susceptibility χ/ϕ of 8.2 (Supporting Information)) and 9 nm (d0 = 9 nm, polydispersity index σ = 0.35, volume fraction of nanoparticles in the suspension φ = 1.9%, specific susceptibility χ/ϕ of 15.5 (SI)). RX diffractograms, TEM micrographs and magnetization curves of the MNPs are available in Supporting Information (Figures SI1, SI2, SI3, Supporting Information). Preparation of UMLs. Chloroform solutions of 1,2-dipalmitoylsn-glycero-3-phosphocholine (DPPC), 1,2-distearoyl-sn-glycero-3phosphocholine (DSPC), 1,2-distearoyl-sn-glycero-3-phosphoethanolamine-n-[(carboxy(polyethyleneglycol)2000](ammonium salt) (DSPE-PEG2000) and L-α-phosphatidylethanolamine-n-(lissaminerhodaminesulfonyl B) (ammonium salt) (Rhod-PE) were purchased from Avanti Polar lipids, Inc. Chloroform and diethyl ether were supplied by Carlo Erba reagents and VWR. Both sizes of MNPs were tested and the amount of lipids was kept constant (2.5 mg/mL for each preparation). UMLs were prepared by the reverse phase evaporation method, established by Skoza et al.25 and modified as follows: a mixture of DPPC/DSPC (90/10 mol %, 250 μL) or DPPC/ DSPC/Rhod-PE/DSPE-PEG 2000 (84/10/1/5 mol %, 315 μL) for in vivo experiments was dissolved in 3 mL of diethyl ether. Chloroform is added to solubilize lipids (typically 900 μL). Afterward 1 mL of MNPs dispersed in water (or in a buffer 0.108 M NaCl, 0.02 M sodium citrate and 0.01 M HEPES, pH = 7.4 for in vivo experiments) was introduced before sonication at room temperature for 20 min to produce a waterin-oil emulsion. Preparation was immediately transferred to a 50 mL round-bottom flask and remaining organic solvent evaporated with a rotavapor R-210 (Buchi) at 25 °C until the gel phase disappeared. Afterward, liposomes were filtrated through a 450 nm filter. Purification of liposomes from nonencapsulated maghemite MNPs was performed by magnetic sorting using a strong magnet purchased from Calamit (Fe−Nd−B 150 × 100 × 25 mm). The operation was repeated three times every 2 h and liposomes were finally separated from the supernatant and recovered. Magnetization curve of liposomes is available in Supporting Information (Figure SI4). Quasi-elastic Light Scattering (QELS). Diameters were determined by using a Zetasizer Nano ZS (Malvern, UK) at 90° scattering angle. Samples of vesicles were diluted in the appropriate buffer (sodium citrate 5 mM or buffer 0.108 M NaCl, 0.02 M sodium citrate and 0.01 M HEPES) to optimize the signal registered and kept at 25 °C during measurement. Diameters were deduced from the Stokes−Einstein law for spherical particles (d = kBT/3πηD where D is the translational diffusion coefficient, kB is the Boltzmann constant and η is the dispersant viscosity).

SLP =

CVs dT m dt

(1) −1

−1

with Cwater = 4185 J L K the volume specific heat capacity of the sample, Vs is the sample volume and m is the mass of magnetic material in the sample. Relaxivity Measurements and Magnetic Resonance Imaging (MRI). Magnetic resonance imaging was performed on a BrukerBioSpec 40 cm bore actively shielded 4.7 T scanner equipped with a cryocooled probe (CryoProbe). The scanner was interfaced to ParaVision software for preclinical MRI research. Tumors were examined with three-dimensional (3D) susceptibility weighted gradient echo sequence TR/TE = 20/5 ms, flip angle of 25°, resolution of 50 × 50 × 50 μm, two averages, and field of view of 1.5 × 1.5 × 0.60 cm. The scan time was 11 min. For relaxometry experiments, T1 and T2 were measured on spin echo sequences on the same apparatus. Calibration curves were prepared at 0.06, 0.13, 0.25, 0.5, and 1 mM of iron determined by FAAS and samples were immobilized in a 0.3% (w/w water) agarose 11835


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Figure 1. (a,b) TEM and (c) cryo-TEM micrographs of UMLs prepared by REV process. At low magnification a large number of dense vesicles are observed with diameters 200 nm in average. MNPs are trapped inside unilamellar vesicles (c) and dipole−dipole interaction can occur as exemplified by magnification (b). PEG-Rhodamine was acquired with λexc at 561 nm and λem at 607 ± 18 nm. Stability in Human Plasma. UMLs-PEG containing 9 nm nanoparticles have been synthesized by the reverse phase evaporation process described above and diluted in human plasma (Sigma) at a concentration of 0.1 M of iron. After short stirring, the solution was kept at 37 °C for 30 min and DLS curve was registered at the same temperature. Tumor Inoculation Cell Culture and Tumor-Bearing Mice. The murine carcinoma TC-1 cell line was kindly provided by Dr. Eric Tartour from the European Hospital G. Pompidou. Cells were cultured and propagated in stock T-flasks at 37 °C, in 95% relative humidity, and in 5% CO2 in RPMI1640 medium supplemented with 10% fetal calf serum,1% sodium pyruvate, 100 units/mL penicillin, and 100 mg/mL streptomycin. Cell concentrations were determined by counting trypsinized cells with a hemocytometer. Tumor inoculum was prepared as a single-cell suspension of 1 × 107 cells/mL in phosphate saline buffer (PBS). Each 7-week-old NMRI nude mouse (N = 6, weighting 20 ± 1 g, provided by Janvier, France) received two 50 μL inocula (5 × 105 cells) administered by subcutaneous injection above the anterior left and right paw in the clavicular region. Mice were housed in polypropylene cages and were provided with food and water ad libitum. Fourteen days following tumor inoculation tumors reached the diameter of approximately 10 mm and mice were ready for further treatment. Animals were handled according to the European Community guidelines for the care and use of laboratory animals (European Directive 86/609/EEC on the protection of Animals used for Experimental and other scientific purposes and its amendment (2003/65/EC)). The Institutional Animal Care and Use Committee of Paris Cardiovascular Research Center (PARCC) approved animal protocols. Proof of Concept Study of the in vivo Magnetic Vectorization of UMLs. Mice where anesthetized with 1.5% isofluorane while a magnet was placed over the right tumor and magnetic liposomes (100 μL; 0.1 M of iron, N = 4 mice) or, in the negative control, the equivalent dose of free maghemite nanoparticles (N = 2 mice) were injected into the right retro-orbital venous sinus. After 1hthe magnet was removed and the animals underwent MRI. Subsequently the mice were euthanized and pieces of livers, spleens and tumors were fixed with pH 7.4 phosphate-buffered 10% formalin and embedded in paraffin. Five-micrometer sections were then stained with Prussian blue and Nuclear Red.

solution. Magnetic nanoparticles of 7 nm (polydispersity index σ = 0.35), 9 nm (polydispersity index σ = 0.35) and UMLs encapsulating each size of MNPs were tested. Image processing and analysis were made with the open source software OsiriX (3.9.2. version). Cell Incubation. Human breastadenocarcinoma MCF-7 cells (ATCC) were cultured in adhesion in DMEM medium supplemented with L-glutamine (0.002 M), penicillin (50 IU/mL), streptomycin (50 IU/mL), 10% fetal calfserum and maintained at 37 °C and 5% CO2 in humidified atmosphere. Cell labeling with nanoparticles was performed by adding filter-sterilized suspension of 9 nm citrate coated MNPs in serum free RPMI culture medium, at a final concentration of iron of 1.5 mM (or 84 μg/mL of iron). Cells were incubated for 30 min at 37 °C, washed 3 times, followed by a 1 h chase at 37 °C in nanoparticles free RPMI culture medium to let the cells internalize MNPs still on the cell surface. Concerning magnetic liposomes, cell labeling was performed by adding UMLs suspended into RPMI culture medium at an equivalent iron concentration for 30 min at 37 °C. After incubation, cells were washed 3 times and incubated again for a chase period of 1 h in liposome free RPMI culture medium. After the whole process, cells were detached using trypsin. For electron microscopy visualization, 1 million cells were centrifugated, fixed with 2.5% glutaraldehyde in 0.1 M sodium cacodylate buffer, postfixed with 1% osmium tetroxide containing 1.5% potassium cyanoferrate, gradually dehydrated in ethanol, and embedded in Epon. Ultrathin sections of 70 nm were examined. For magnetic cell quantification (cell magnetophoresis), detached cells were resuspended at a cell density of 0.5 million cells per mL in RPMI culture medium and introduced in a 1 mm thick Hellma chamber. Cytotoxicity Assay. The cytotoxicity of the UMLs has been analyzed by AlamarBlue assay (Life Technologies). MCF-7 cells have been incubated with MNPs and two kinds of UMLs, differing for the use of DSPE-PEG2000 in the preparation (UMLs and UMLs-PEG). In the case of UMLs, the effect of the permanent magnet has (50 T/ m) been also evaluated. After the incubation step (see previous section), the culture medium was replaced with the AlamarBlue solution, following the protocol provided by manufacturer. The resulting fluorescence has been analyzed by a microplate reader (BMG FluoStar Galaxy), with a excitation wavelength of 550 nm and by collecting the fluorescence at 590 nm. All the experimental points reported have been reproduced in triplicate. Confocal Microscope Analysis. Confocal microscope images has been acquired by using a Andor Technology with Olympus JX81/ BX61 Device/Yokogawa CSU Device spinning disk microscope (Andor Technology plc, Belfast, Northern Ireland), equipped with a 60×Plan-ApoN oil objective lens (60×/1.42 oil, Olympus). MCF-7 cells have been incubated with UMLs-PEG (or UMLs-PEG-Rhodamine) as reported in the previous section. A chase period of 12 h after the incubation was introduced, in order to drive the liposome localization mainly in the lysosomal compartment. The cells were stained with Lysotracker Green DND-26 (150 nM for 2 h, λexc = 488 nm, λem = 525 ± 15 nm), Hoechst 33342 (5 μg/mL for 20 min, λexc = 405 nm, λem = 465 ± 30 nm) and CellMask DeepRed (7.5 μg/mL for 5 min, λexc = 640 nm, λem = 660 ± 20 nm). The red signal from UMLs-

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RESULTS AND DISCUSSION Ultra Magnetic Liposomes (UMLs), Ultra Magnetic stealth Liposomes (UMLs-PEG) and Ultra Magnetic stealth fluorescent Liposomes (UMLs-PEG-Rhodamine) are prepared by reverse phase evaporation process (REV) with constant and very low amount of lipids (2.5 mg). Maghemite nanoparticles (9 or 7 nm) coated with citrate ligands and dispersed in water or in a buffer were used for the preparation of liposomes. At the end of the synthesis process, a high amount of spherical ultra 11836


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magnetic vesicles is generated. The objects diameters were measured by Dynamic Light Scattering (DLS) and plotted with the intensity as a function of diameter. Vesicles of 217 nm (polydispersity index σ = 0.27) and 238 nm (polydispersity index σ = 0.18) in diameter for UMLs and UMLs-PEG respectively are produced (curves are available in Supporting Information, Figure SI5). Such sizes are at the limit for in vivo applications, liposomes 100 nm in diameter being preferred to avoid fast clearance or uptake by the reticuloendothelial system. However the UMLs have been designed here to exhibit the best magnetic responsivity for magnetic targeting and MRI and hyperthermia applications, justifying to be placed at the limit of 200 nm in size (in particular the magnetic force on the liposome then increases to a factor close to 10 from 100 to 200 nm). Electron microscopy micrographs (Figure 1a) show electron-dense objects and high magnification view confirms the presence of nanoparticles trapped and dispersed in the aqueous core of the vesicles (inset Figure 1a). The magnetic responsivity of vesicles is so strong that they even form dimers or small chains when submitted to the magnetized tweezers used for TEM grids’ manipulation (Figure 1b). Furthermore, the extremely low lipid concentration used in this synthesis results in the formation of unilamellar liposomes, avoiding a complementary extrusion process as confirmed by cryo-TEM observations (Figure 1c). We presume here that the low lipid/ nanoparticle ratio applied in the synthesis protocol controls the formation of these original vesicles. This hypothesis also relies on previous studies revealing that this ratio controls the morphology of the objects prepared.23 To confirm this hypothesis, UMLs were prepared with different ratios lipid/ nanoparticle with an amount of lipids kept constant (2.5 mg). Nanoparticles were diluted at volume fractions of 3.1, 0.78, 0.39, and 0.31% (313, 78, 39 and 3 mg/mL) for each synthesis. For a volume fraction of 3.1% (313 mg/mL, MNPs not diluted), TEM micrographs shows spherical objects (Figure SI6a, Supporting Information), well dispersed and densely packed with MNPs. When decreasing the concentration of iron, liposomes are still visible but appear more aggregated, less spherical and less filled with nanoparticles (Figure SI6.b,c). For highly diluted MNPs (3 mg/mL), liposomes are mostly empty and nanoparticles are outside the vesicles (Figure SI6d). Thus, for a lipid concentration of 2.5 mg/mL, the concentration of nanoparticles clearly influences the resulting iron content of vesicles. This observation completes the previous study carried out by Wijaya et al.23 who used solutions 10 times more concentrated in lipids. They evidenced that for a fixed concentration of lipids (23−27 mg/mL), decreasing the concentration of MNPs from 105 to 12 mg/mL drives to bicelles/MNPs-loaded vesicles, empty vesicles/MNPs-loaded vesicles and finally MNPs-loaded vesicles. Thus, the presence of MNPs influenced the phase diagram of lipids. In our case, for a concentration of lipids of 2.5 mg/mL, decreasing the concentration of MNPs from 313 mg/mL to 3 mg/mL always leads to liposomes, from densely packed liposomes to mixture of loaded and empty vesicles and finally empty vesicles. Magnetophoresis experiments were carried out to quantify the magnetic mobility and magnetization of UMLs encapsulating 7 or 9 nm MNPs. In the vicinity of a magnetized nickel wire, vesicles immediately started to accumulate on the wire forming a large brown ring after a few seconds (Figure 2a). The analysis of UMLs displacements yields velocities of 41 ± 5 μm/ s for UMLs encapsulating 9 nm nanoparticles and 55 ± 6 μm/s for liposomes trapping 7 nm MNPs (Figure 2b). From these

Figure 2. (a) UMLs rapidly accumulate on a magnetized nickel wire (optical micrographs before and after 10 min of experiment). (b) Magnetic migration of individual liposomes was recorded every 0.1 s in an observation window located 150 μm apart from the magnetic wire (as delimited by an open rectangle in a). The corresponding magnetic field gradient in this window is 190 T/m. (c) Average of the measured magnetic velocities and equivalent magnetic volume fractions for the UMLs encapsulating 7 and 9 nm MNPs. (d) Field Cooled and Zero Field Cooled susceptibilities measured for UMLs and MNPs.

velocities, the magnetic moment of each type of UML as well as the corresponding volume fraction of MNPs can be estimated to 24 ± 3% and 33 ± 4% for 9 and 7 nm nanoparticles, 11837


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respectively (Figure 2c). This corresponds to a content of ∼100 mol of iron per mol of lipids which is much higher than data reported until now (0,53 mol Fe/mol lipids,2 10 mol Fe/mol lipids,28,29 19 mol Fe/mol lipids30). This high volume fraction of MNPs inside liposomes leads to interparticle magnetic dipole−dipole interactions, which can be evidenced by SQUID experiments. These interactions clearly influence the temperature dependence of FC and ZFC susceptibilities (Figure 2d). In agreement with their magnetic anisotropy energies, the 7 and 9 nm diluted MNPs in colloidal suspension display a superparamagnetic blocking temperature (corresponding to the maximum of ZFC magnetization) of Tb ≈ 60 K and Tb ≈ 90 K respectively. By contrast, the same MNPs encapsulated in UMLs show a distinct magnetic behavior: the maximum of ZFC magnetization is shifted to higher temperatures (∼125 and 190 K for the UMLs-7 nm and UMLs-9 nm respectively), while a flattening of the FC curve below this “blocking” temperature is observed. These features clearly indicate strong dipole−dipole interactions between MNPs which are generally observed for uniformly dispersed particles at very large volume fractions above 10%31,32 or for cell internalized MNPs.33 Estimation of the dipole−dipole interaction energy from the experimental shift of Tb drives to a volume fraction of ∼20% for the 9 nm MNPs and ∼30% for the 7 nm MNPs in good agreement with the direct quantification from velocity measurements.34 This magnetic content heightens the heating capacity of UMLs in magnetic hyperthermia. Interestingly, MNPs liposomal encapsulation at these very high local concentrations substantially increases their heating capacity under alternating magnetic field. This phenomenon was investigated using a homemade device previously described.8 Magnetic samples (Vs = 300 μL) were placed in a coil producing an alternating magnetic field characterized by a frequency of 700 kHz and an amplitude of 27 kA m−1. Temperature was measured every 0.7 s for 25 min and finally plotted as a function of time. The ΔT, defined as the difference between the maximum reached temperature and the initial temperature (30 ± 0.5 °C), was measured for each sample. For 7 and 9 nm MNPs, ΔT obtained are 34 ± 0.5 °C and 46.1 ± 0.5 °C respectively and for UMLs encapsulating 7 and 9 nm, ΔT are 14.9 ± 0.5 °C and 40.7 ± 0.5 °C respectively (Figure 3a). After hyperthermia, liposomes were observed in TEM to check that no leakage of nanoparticles occurred upon heating. Pictures revealed that liposomes are slightly aggregated but still filled with nanoparticles (Figure SI7). From this experiment, the initial linear rise in temperature versus time dependence, dT/dt can be measured. Using the eq 1, it is therefore possible to determine the specific loss power (SLP or specific absorption rate) defined as the thermal power dissipation divided by the mass of magnetic crystal. Liposomes encapsulating 9 and 7 nm MNPs respectively show values of 438 W/g and 164 W/g compared to 270 W/g and 108 W/g for 9 and 7 nm MNPs uniformly dispersed in colloidal suspension (Figure 3a). For what concerns hyperthermia performance, these data highlight the influence of nanoparticle size as previously observed8 but also underline the noticeable effect of their local confinement in liposomes. To go further into their characterization, we investigated the influence of the high loading capacity on relaxivities. These parameters govern the efficiency of the UMLs as MRI contrast agents. Experimentally, whatever the size of MNPs (7 or 9 nm) encapsulated into the liposomes, the transversal relaxivity r2

Figure 3. (a) SLP values and (b) relaxivity measurements for 7 and 9 nm MNPs alone and UMLs encapsulating MNP.

(spin−spin relaxation process) is found to be enhanced after encapsulation into liposomes (Figure 3b) confirming the efficiency of these systems as T2 contrast agents. Actually, the core of densely packed MNPs generates high r2 values due to the resulting high magnetic moment of the UMLs. It is wellknown that the higher the local iron concentration, the better the r2 value is. Here, we also point out the role of the particle diameter considering that UMLs-9 nm (volume fraction ≈ 20%) exhibit higher r2 than UMLs-7 nm (volume fraction ≈ 30%).35 Besides, the longitudinal relaxivity r1 in UMLs is lowered when MNPs are encapsulated as it was observed before in the case of magnetic LUVs prepared by spontaneous swelling,2 in clusters of MNPs and polymer36 or MNPs internalized in cells.37 This effect could be attributed to a saturation of the longitudinal relaxing effect of the MNPs, which are highly confined in liposome compartments in comparison to dispersed free MNPs. The lipid bilayer is also likely to play a role as a barrier to the exchange of water molecules between the interior and the exterior of the liposomes. As a result of the simultaneous decrease of r1 and increase of r2, UMLs demonstrate a better MRI efficiency than free nanoparticles with r2/r1 ratio, which jumps from 28.2 to 45.7 for 7 nm MNPs and from 53.7 to 84.8 for 9 nm MNPs when they get trapped (Figure 3b). Our relaxivity experiments agree with previous studies; however the very high loading of MNPs induce a striking increase if compared to experiments carried out on conventional magnetic liposomes prepared with maghemite nanoparticles. In the case of UMLs, this r2/r1 ratio appears to be much higher than the values obtained in other studies: Martina et al.2 obtained a r2/r1 ratio of 6.2 (with 7 nm maghemite nanoparticles, magnetic field of 0.47 T), Pauser et al.38 prepared magnetic stealth liposomes with a r2/r1 ratio of 11838


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13.8 (with SPIO nanoparticles, magnetic field of 2.35 T) and Bulte et al.30 obtained a r2/r1 ratio of 80 (with 16 nm maghemite nanoparticles, magnetic field of 1.5 T). Our evolution of relaxivity profile is in favor of an efficient detection of UMLs by T2-weighted spin echo sequences and furthermore by T2*-weighted gradient echo sequences which are sensitive to local changes in susceptibility, as used during our in vivo imaging of UMLs accumulation in tumors. Among the two UMLs designed here, the one encapsulating 9 nm nanoparticles show the best potential for future MRI (higher relaxivity r2 value) and therapeutic hyperthermia (higher heating power) applications. We chose to further test this 9 nm loaded UMLs both in vitro for cell internalization and in vivo for magnetic targeting. We first assessed the ability of a tumor cell line (human breast adenocarcinoma MCF-7) cells to internalize the UMLs. For experiments on cells, two formulations were compared: UMLs and UMLs containing 5% mol of DSPE-PEG2000 (UMLs-PEG). A small quantity of PEG-grafted lipids inserted in the bilayer lengthen twice liposomes circulation time by avoiding fast opsonization and sequestration by the reticuloendothelial system (RES).39 In our experiment, tumor cells were incubated with dispersed MNPs or with UMLs or with UMLs-PEG in a zero magnetic. Electron microscopy micrographs undoubtedly revealed that incorporation of nanoparticles within the cells is more efficient when nanoparticles are vectorized by liposomes. First, intracellular vesicles (namely endosomes) filled with MNPs increase in number when nanoparticles are trapped in the UMLs (Figure 4b1) compared to MNPs alone (Figure 4a1). When liposomes were modified with a PEG coating (UMLs-PEG) we observed a decrease in the liposomes uptake. This behavior should be explained by an improved colloidal stability of UMLs-PEG in the cell culture condition and/or by a partial inhibition of the cellular uptake.40 Interestingly, vesicles shell is lost once internalized in the cell, likely owing to membrane fusion during uptake. However, rare circular structures are still visible in the endosomes. Complementary pictures are available in Supporting Information (Figure SI8). The magnetic liposomes internalization process has subsequently been also investigated by confocal microscopy. After the incubation with UMLs-PEG, the lysosomes were stained with a specific marker (Lysotracker Green DND-26). The treated cells showed a high amount of large lysosomes in the cytoplasm and there is a good colocalization with the black spot detected in the corresponding bright-field, ascribable to the UMLs-PEG internalization (Figure SI9, Supporting Information). As a further confirmation of the cellular localization, MCF-7 cells were incubated with UMLs prepared with a red-emitting rhodamine-lipid (UMLs-PEG-Rhodamine). The confocal microscope analysis revealed also in this case a colocalization between the green signal from Lysotracker and the red signal from UMLs (Figure 4a2, b2, c2). The cellular uptake for each condition was then quantified by magnetophoresis (Figure 5a). The derived iron contents per cell are in the order of 4 pg per cell (4.1 ± 0.4) for 9 nm interacting with the cells alone, 28 pg per cell (28.4 ± 1.3) for UMLs and 16 pg (16.4 ± 2) for UMLs-PEG. It is therefore possible to achieve a very high cellular magnetic content for a short incubation time (30 min with 84 μg/mL of iron). These data are comparable to other works, but with a cellular uptake necessitating much longer incubation times: Soenen et al.41 accumulated 47 pg of iron in 3T3 fibroblasts after 4 h incubation with 100 μg/mL of iron, Bulte et al.42 found 10 pg

Figure 4. 1. TEM micrographs of tumoral MCF-7 cells incubated with 9 nm MNPs (a1), UMLs (b1), and UMLs-PEG (c1). 2. Confocal microscope characterization of MCF-7 cells incubated with UMLsPEG-Rhodamine. The cells were analyzed for red fluorescence of rhodamine (a2) and green signal of Lysotracker Green DND-26 (b2). The cells have been counterstained with Hoechst 33342 (blue, for nuclei) and CellMask Deep Red (magenta, for plasma membranes). The overall merge of the fluorescence channels is reported in panel c2. The scale bar corresponds to 10 μm.

Figure 5. (a) Internalization efficiency and (b) cell viability after incubation of MNPs and corresponding UMLs [Fe] = 1.5 mM for 30 min.

of iron with oligodendrocyte precursor cells after 1−2 days of incubation with 25 μg/mL of iron, Martina et al.43 achieved 19 pg of iron with PC3 tumor cells, after 4 h of incubation with 1 11839


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Figure 6. Ultra Magnetic liposomes in vivo. MR scans of murine tumors (a−d) before the injection (a) and after injection of free iron oxide nanoparticles following magnetic targeting (b), the tumor prior injection (c) and after injection of Ultra Magnetic Liposomes following magnetic targeting revealing several hypointense dots within the tumor (d). Sagittal coverage was amplified in all images by overlapping three 50 μm slices with minimum intensity projection (MinIP) volume-rendering technique with OsiriX image processing software. Optical microscopy micrographs (e, f) of the zone of interest marked on figure (d) of the tumor. Histological sections (thickness 5 μm) were stained with Pearls and Nuclear Red staining (e, f), blue arrows indicate characteristic blue colored iron-loaded cells.

mg/mL of iron or Jendelova et al.44 reported 17.5 pg of iron in bone marrow stroma cells after 48−72 h of incubation with 110 μg/mL of iron. Finally, the effect on cell toxicity of the UMLs in the conditions used for in vitro experiments has been investigated by AlamarBlue assay. After the incubation, the cells did not show any morphological change and the metabolic activity was not affected by the internalization of liposomes UMLs or polyethylene coated UMLs-PEG (Figure 5b). In addition, the incubation in presence of a permanent magnetic field gradient did not produce a different behavior in comparison with negative control MCF-7 cells and MNPsdoped cells. To establish the feasibility of magnetic tumor targeting by our liposomes in vivo, a proof of concept study was performed on six mice. This experiment was only performed with UMLsPEG to maximize their circulation time and the probability to

come across the tumor. Before injecting liposomes to the mice, the stability of liposomes in human plasma was checked. After 30 min at 37 °C, UMLs-PEG (0.1 mM of iron) did not show any aggregation or nanoparticles leakage (Figure SI10, Supporting Information). A magnet was placed over the tumor of six mice and UMLs were injected intravenously to four animals or, in the control group of two mice, magnetic nanoparticles were injected alone. High resolution MRI was performed prior and post injection of the suspension of iron oxide nanoparticles (Figure 6a and b) or UMLs (Figure 6c and d). While the free nanoparticles did not give any hypointensity signal inside the tumors after magnetic targeting (Figure 6b), the superparamagnetic particles that were loaded within the liposomes seemed to disrupt the magnetic field homogeneity in the tissue (Figure 6d). These foci of signal intensity loss in our T2* weighted gradient echo sequence are visible as black dots 11840


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in the MR scan (Figure 6d) of the tumor that was formerly adjacent to the magnet. Note that these hypo intense dots were present in tumors adjacent to the magnet of all UML treated mice (see Supporting Information, Figure SI11). In contrast, collateral tumors where no exterior magnetic field was applied showed little if any dark spots (figure not shown). In order to confirm that localized MR signal loss may be attributed to iron nanoparticles vectorized by UMLs, the scans were compared to histological sections of tumors after Pearls staining (Figure 6e and f). Optical microscopy (Figure 6e and f) revealed characteristic blue spots. Prussian blue staining therefore validated the presence of intracytoplasmatic iron in cells inside the tumor, which was more prominent in tumors adjacent to the magnet in UMLs-treated mice. These findings demonstrate the efficiency of UMLs to infiltrate the tumor with an enhanced accumulation when a magnet is applied to target the UMLs and show that our liposomes can be readily detected by high resolution MRI.

CONCLUSIONS In summary, we present the preparation of Ultra Magnetic Liposomes by a modified reverse phase evaporation process, which allows the incorporation of a high amount of nanoparticles in the aqueous core. The magnetic load and morphology of vesicles were confirmed by microscopy techniques and consolidated by magnetic experiments. These vesicles represent promising vectors as they greatly penetrate tumor cells and can be magnetically vectorized to solid tumors in vivo as evidenced by MRI and corroborated by histological analysis. Besides, UML’s heating capacity enables their use for hyperthermia, which aims to destroy tumor cells or could be used for triggered drug release due to the thermosensitive properties of UMLs’ membrane. Moreover, such ultra magnetic liposomes reveal to be very interesting tools to mimic and monitor tumor growth pressure in vivo. ASSOCIATED CONTENT

S Supporting Information *

Supplementary figures. This material is available free of charge via the Internet at http://pubs.acs.org.

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REFERENCES

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AUTHOR INFORMATION

Corresponding Author

*E-mail: [email protected]. Tel: +33 1 44 27 30 47. Fax: +33 1 44 27 32 28. Notes

The authors declare no competing financial interest.

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ACKNOWLEDGMENTS We are highly grateful to Delphine Talbot for the preparation of different magnetic nanoparticles used for our experiments and Aude Michel and Sophie Neveu for TEM pictures. We kindly thank Laëtitia Pidial for the TC1 cells culture, Gwennhael Autret for her assistance on MRI apparatus and Jean-Pierre Lechaire, Ghislaine Frébourg and Géraldine Toutirais from the Electron Microscopy Service of the Institut de Biologie Intégrative (IFR 83), Université Pierre et Marie Curie, for their expertise and practical assistance on Cryo-TEM and electron microscopy visualization. This work was supported by European projects ENCITE (European Network for Cell Imaging and Tracking Expertise, grant agreement number 201842). 11841


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Ultra Magnetic Liposomes for MR Imaging, Targeting, and Hyperthermia

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