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Published online before print June 11, 2007, 10.1148/radiol.2442060912
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(Radiology 2007;244:439-448.)
© RSNA, 2007

Experimental Studies

Magnetic Targeting of Nanometric Magnetic Fluid–loaded Liposomes to Specific Brain Intravascular Areas: A Dynamic Imaging Study in Mice1

Charlotte Rivière, PhD, Marie-Sophie Martina, PhD, Yutaka Tomita, MD, PhD, Claire Wilhelm, PhD, Alexy Tran Dinh, MS, Christine Ménager, PhD, Elisabeth Pinard, PhD, Sylviane Lesieur, PhD, Florence Gazeau, PhD, and Jacques Seylaz, PhD

1 From the Laboratoire Matière et Systèmes Complexes, CNRS UMR 7057, Université Paris 7-Denis Diderot, 140, rue de Lourmel, 75015 Paris, France (C.R., C.W., F.G.); Equipe Physico-Chimie des Systèmes Polyphasés, Chatenay-Malabry, France (M.S.M., S.L.); Centre de Recherche Cardiovasculaire Lariboisière, Paris, France (Y.T., A.T.D., E.P., J.S.); and Laboratoire des Liquides Ioniques et Interfaces Chargées, Paris, France (C.R., C.M.). Received May 25, 2006; revision requested July 26; revision received September 19; accepted October 26; final version accepted December 14. Supported by Institut National de la Santé et de la Recherche Medicale (INSERM), Ministère de l'Education Nationale de l'Enseignement Superieur et de la Recherche (ACI Nanosciences 145), and Direction Générale de l'Armement (DGA). Y.T. supported by INSERM. Address correspondence to F.G. (e-mail: floga{at}ccr.jussieu.fr).


   ABSTRACT
TOP
 ABSTRACT
INTRODUCTION
 MATERIALS AND METHODS
 RESULTS
 DISCUSSION
 ADVANCES IN KNOWLEDGE
 References
 
Purpose: To prospectively determine, by using dynamic imaging, whether a magnet placed over a specific area of the mouse brain could target systemically administered rhodamine-labeled magnetic fluid–loaded liposomes (MFLs) to that brain region.

Materials and Methods: Experiments were performed with a French Ministry of Agriculture permit and regional ethics committee authorization. In seven anesthetized C57BL/6 mice, a closed cranial window was implanted above the left parieto-occipital cortex. A laser-scanning confocal fluorescence microscope (LSCFM) was used to track the intravenously injected rhodamine-labeled MFLs within this cortical area, through the cranial window. The MFLs were video monitored for 2 minutes every 15 minutes for 1 hour after injection. A magnet was placed on the cranial window implanted in four mice, while no magnet was placed in three (control) mice. After dynamic in vivo imaging, static in vivo imaging was performed with a different LSCFM. Ex vivo fluorescence histologic analysis was then performed. Paired Student t testing was used to compare the cerebral blood flow and two-dimensional flow values before and 1 hour after MFL injection. For image analysis, intergroup comparisons were performed by using an independent t test.

Results: In vivo video monitoring through the window revealed that the rhodamine-labeled MFLs accumulated in the mouse brain microvasculature exposed to the magnet—first within superficial brain venules and then within intracerebral venules—with no significant change in blood flow (P > .05). MFLs accumulated neither in the arterioles of the mice with a magnet nor in the arterioles of the control mice. Static in vivo imaging findings confirmed the microvascular localization of the rhodamine-labeled MFLs, and histologic findings specified their accumulation on the side of the magnet only.

Conclusion: Real-time in vivo imaging of rhodamine-labeled MFLs in the mouse brain cortex revealed that these nanosystems can be magnetically targeted, through microvessels, to selected brain areas.

Supplemental material: http://radiology.rsnajnls.org/cgi/content/full/2442060912/DC1

© RSNA, 2007


   INTRODUCTION
TOP
 ABSTRACT
 INTRODUCTION
MATERIALS AND METHODS
 RESULTS
 DISCUSSION
 ADVANCES IN KNOWLEDGE
 References
 
Various strategies to concentrate drugs within targeted tissues and thereby avoid the adverse effects of systemic administration have been developed. Drug delivery to the brain is especially challenging because the blood-brain barrier is nearly impermeable. One approach is to encapsulate drugs within nanostructures such as liposomes or nanospheres (1–5). Liposomes are particularly attractive as drug delivery systems because they are phospholipid-based biocompatible structures of adjustable size. They can coencapsulate both hydrophobic (within the phospholipid bilayer) and hydrophilic (inside the vesicle) molecules, such as those of therapeutic chemicals, genes, and contrast agents (6–13). The stealthiness of liposomes coated with polyethylene glycol makes them promising as passive drug delivery devices (13–16). Liposomes encapsulating doxorubicin (Caelyx; Schering Plough, Kenilworth, NJ) are currently being tested in a phase II clinical trial and have been found to prolong the survival of patients with malignant glioma (17,18).

Active targeting can also be performed by using magnetic forces (19,20). This strategy has been tested in patients with advanced stages of cancer (21–24) and is well tolerated. Most magnetic targeting studies have been focused on aqueous suspensions of micrometric magnetic particles as drug carriers, but the plasma half-life of such particles is relatively short, restricting their use through the systemic route.

An alternative system based on the coupling of liposome and nanoparticle technologies was developed recently (10). This system consists of superparamagnetic maghemite nanocrystals encapsulated in large unilamellar liposomes that are sterically stabilized by surface-grafted polyethylene glycol chains to form magnetic fluid–loaded liposomes (MFLs). These hybrid nanoscale particles (<200 nm in total size) have the magnetic properties of iron oxide nanocrystals combined with the physicochemical properties of liposomes, which are essential for active organ targeting through the systemic route. These particles also have good physical and biological stability, no cellular toxicity, stealthiness, and mobility under a magnetic field gradient. Magnetic targeting of MFLs to superficial solid tumors implanted in mice has already been achieved (25). The specific accumulation of MFLs in tumors exposed to an external magnet has been monitored with magnetic resonance (MR) imaging. Thus, the purpose of our study was to prospectively determine, by using dynamic imaging, whether a magnet placed over a specific area of the mouse brain could target systemically administered rhodamine-labeled MFLs to that brain region.


   MATERIALS AND METHODS
TOP
 ABSTRACT
 INTRODUCTION
 MATERIALS AND METHODS
RESULTS
 DISCUSSION
 ADVANCES IN KNOWLEDGE
 References
 
Synthesis of Magnetic Nanoparticles
Monodisperse maghemite anionic nanocrystals were synthesized (C.M.) as previously described (26,27). Final adjustment of both the aqueous medium (a 7.4-pH buffer consisting of 108 mmol/L sodium chloride, 20 mmol/L sodium citrate, and 10 mmol/L 4-[2-hydroxyethyl]-1-piperazineethanesulfonic acid) and the maghemite concentration was achieved by means of ultrafiltration (MACROSEP filter, 50-kDa cutoff; Fisher Scientific Labosi, Elancourt, France). The resulting nanoparticles consisted of monocrystalline ferrimagnetic monodomains of maghemite with a crystal diameter of 8 nm, as determined by using magnetization measurements (C.M.).

Preparation of Rhodamine-labeled MFLs
Rhodamine-labeled MFLs were prepared as described by Martina et al (10) and Lesieur et al (26). The lipid content (expressed as a percentage) of the films—comprising 94 mol of egg phosphatidylcholine, 5 mol of 1.2-diacyl-SN-glycero-3-phosphoethanolamine-N-[methoxy(polyethylene glycol)-2000], and 1 mol of rhodamine-polyethylene—was determined by means of weighing (M.S.M.). Nonentrapped maghemite particles were removed by means of gel exclusion chromatography performed with a Sephacryl S1000 Superfine column (Amersham Pharmacia Biotech, Piscataway, NJ) and a 1-mL syringe (Terumo, Elkton, Md).

Unimodal analysis performed by using quasielastic light scattering with a nanosizer (model N4 MD; Coultronics, Margency, France) at 25°C with a 90° scattering angle (for a lipid concentration of 0.5 x 10–2 mol) (M.S.M.) revealed the mean hydrodynamic diameter to be 180 nm ± 40 (standard deviation).

It has been verified (10) that the intrinsic magnetic properties of the initial suspension of maghemite nanoparticles are preserved after encapsulation within liposomes. Thus, the rhodamine-labeled MFLs exhibited superparamagnetic behavior, with no magnetization in a zero field and magnetization following a Langevin function in a magnetic field.

Magnetic Sorting
To select the rhodamine-labeled MFLs that contained the highest number of encapsulated magnetic nanoparticles, the MFLs were exposed to a 0.1-T, 50 x 10-mm rectangular neodymium magnet (Calamit, Paris, France) for 4 hours. The rhodamine-labeled MFL suspension (500 µL) was placed in a 500-µL Eppendorf tube, to which a bar magnet was attached. The MFLs with a high load of magnetic fluid were retained by the magnetic field gradient of the attached magnet and separated from the MFLs with a low load, as determined by two authors (C.R., M.S.M.) in consensus. After elution of the supernatant, the magnetic fraction was resuspended in half the initial volume.

The initial preparation of superparamagnetic rhodamine-labeled MFLs contained 49 mmol/L iron (measured with flame spectroscopy) (C.M.) and 20 mmol/L lipid (determined by means of weighing) (M.S.M.). Magnetic separation yielded a concentrated solution composed of 59 mmol/L iron and 16 mmol/L lipid, and the selected rhodamine-labeled MFLs contained a high iron load (estimated nanoparticle volume fraction of 1%–4%, corresponding to 60–300 magnetic nanoparticles per liposome). No substantial change in the mean size of the MFLs occurred after magnetic separation (180 nm ± 40 before separation, 182 nm ± 34 afterward).

Magnetic Force Calibration
For magnetic targeting (Appendix E1; http://radiology.rsnajnls.org/cgi/content/full/2442060912/DC1), a cylindrical 3-mm-diameter, 30-mm-long neodymium magnet (Calamit) was used to create a 0.1-T magnetic field in the upper part of the mouse brain. In our setup, the magnetic force (Fm) to which the rhodamine-labeled MFLs were subjected was mainly created by the vertical magnetic gradient (z-axis direction, of unit vector ez: dB/dz):

Formula
where m is the magnetic moment of the liposome in the 0.1-T magnetic field, e is the unit vector of the z-axis direction, and dB/dz is the variation of B with z (gradient of B over z). The magnetic gradient within the observation window along the z-axis direction, (gradB)z, was calibrated with commercially available micrometric magnetic beads (Dynabeads M-280 Tosylactivated; Dynal Biotech, Lake Success, NY) and calculated to be 42 T/m.

The global magnetic moment of each liposome, together with the magnetic force acting on the liposome, is directly proportional to the number of encapsulated magnetic nanoparticles. This setup allows one to apply a force, F, on each rhodamine-labeled MFL on the order of 10–15 · N, where F = m · gradB, m = mv · V · N, mv is the volumic magnetization of nanoparticles, V is the nanoparticle volume, N is the number of particles per liposome, and mv = 2.17 x 105 A/m—the volumic magnetization of nanoparticles at 0.1 T (10,28).

Knowing the magnetic force, we estimated the magnetophoretic mobility of a single MFL to be 0.6 µm/sec by using the balance of magnetic and viscous forces. When two liposomes are close, they also experience a dipole-dipole interaction force (Fdipdip) on the order of 10–12 · N, where Fdipdip = µ0m2/r4, µ0 is the permeability, and r is the distance between two liposomes—that is, 180 nm minimum.

Animal Preparation
Experiments were performed by using seven anesthetized male C57BL/6 mice (Centre d'Elevage R. Janvier, l'Arbresle, France) weighing 22–25 g, with French Ministry of Agriculture permit number 02934 for experimentation on vertebrate animals and regional ethics committee authorization for animal experimentation in these specific procedures. Anesthesia was induced with 1.8%–2.0% isoflurane and maintained with 1.4%–1.6% isoflurane in a 25% O2, 75% N2O mixture. A custom-built system was used to secure the mouse's head and stabilize the experimental apparatus under the confocal microscope while delivering anesthetic gas.

Implantation of Closed Cranial Window
All surgical procedures were performed under an operating microscope (model S21; Carl Zeiss, Oberkochen, Germany). The protocol is described in detail elsewhere (29). Briefly, 1 week before the experiments, a 3-mm-diameter hole was drilled in the mouse's skull, above the left parieto-occipital cortex (Y.T.). The center of the cranial hole was located 2.5 mm posterior to the bregma and 2.5 mm lateral to the midline (30). The dura mater was left intact. A 140-µm-thick circular quartz coverslip was sealed to the bone with dental cement (Ionosit; Dental Material Gesellschaft, Hamburg, Germany) to make the preparation waterproof. No bleeding or inflammation occurred, as determined by two authors (Y.T., J.S.) in consensus.

In Vivo Experiments
Seven mice were randomly separated into two groups: Four mice were exposed to the 30-mm-long external magnet used for magnetic force calibration, which was placed in contact with the window, while the other three mice served as magnet-free control animals (Fig 1). A catheter was inserted into the left tail vein to inject fluorescent tracers (Y.T.).


Figure 1
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  		Figure 1: Experimental setup in which a closed cranial window was implanted above the cortex in seven mice. Four animals were exposed to a magnet, and three served as controls. Time schedules are shown on the left for the control group and on the right for the magnet group. The magnet was withdrawn during visualization to create the necessary working space.

 
Dynamic in vivo imaging.—Dynamic in vivo imaging of the pial and cortical microcirculations was performed in real time by using a dynamic laser-scanning confocal fluorescence microscope. The anesthetized mouse was placed under the microscope. The mouse head was secured in a device that closely fit the confocal microscope stage. The cranial window was located horizontally below the lens. Images were obtained to a depth of up to 200 µm within the cortex.

A confocal laser-scanning unit (Viewscan DVC-250; BioRad, Hercule, Calif) attached to a fluorescence microscope (Optiphot-2D; Nikon France SAS, Champigny sur Marne, France) was used for rapid two-dimensional image acquisition (50 frames per second), as previously described (31) for the measurement of fluorescent red blood cell velocity and flow. Appropriate filters were used for fluorescein isothiocyanate (FITC) and rhodamine. The light source of the confocal microscope was an argon-krypton laser operating at wavelengths of 488 and 568 nm. The objectives were x10 and x4 dry lenses (Nikon France SAS). The numerical apertures were 0.40 and 0.55, and the working distances were 1.8 and 2.0 mm, respectively. Images from a single plane were visualized with a silicon-intensified target camera (model C-2400; Hamamatsu Photonics, Shizuoka, Japan) and digitally video recorded (DSR-25 DVCAM; Sony, Saitama, Japan). One advantage to using this dynamic confocal microscope, as compared with using the classic confocal microscope, is the shorter radiation time due to the larger width of the pinhole. The shorter radiation period makes repeated dynamic imaging in the same animal possible.

A tracer dose (10 µL) of FITC-dextran (Sigma, Saint-Quentin Fallavier, France) (molecular weight, 71 200 Da; 20 mg/mL) was first injected to visualize the microvascular network (J.S.). We distinguished meningeal microvessels from cortical microvessels by changing the focal length of the microscope. Arterioles were characterized on the basis of the high velocity of circulating fluorescent tracers, which made it impossible to determine the direction of the blood flow. A 50-µL bolus of FITC-dextran was then injected to measure the local blood flow (J.S.). Finally, the rhodamine-labeled MFLs (about 1013 MFLs per 100 µL) were injected (M.S.M.) and tracked through the closed cranial window. The distribution of the MFLs at a depth of up to 200 µm beneath the brain surface was video monitored for 2 minutes every 15 minutes for 1 hour. The images were stored on a personal computer with an 8-bit frame-grabber card (model LG-3; Scion, Frederick, Md).

In the mice with implanted magnets (magnet group), during injection of the rhodamine-labeled MFLs and during each subsequent 15-minute period within 1 hour, the magnet was placed in direct contact with the closed cranial window (C.R.). The magnet-to–brain surface distance was 140 µm, which corresponded to the thickness of the quartz coverslip. For reasons related to the working space, the magnet had to be removed during image acquisition—that is, for 2 minutes.

Static in vivo imaging.—At the end of the four dynamic image acquisitions, 1 hour after the tracer injection, static in vivo high-spatial-resolution imaging of both the rhodamine-labeled MFLs and the FITC-dextran–labeled vessels was performed through the cranial window by using a classic laser-scanning confocal fluorescence microscope (model MRC 600; BioRad). The image sampling rate (0.5 frame per second) of this microscope is much slower than that of the Viewscan confocal microscope system. Images were obtained at a depth of up to 200 µm beneath the brain surface (J.S.) just after removal of the magnet. The magnet had been placed above the window for a total of 52 minutes before image acquisition.

Image Analysis
The microvascular network was visualized through the cranial window owing to the green fluorescent signal emitted by the FITC-labeled plasma. Red fluorescent signals from the rhodamine-labeled MFLs were differentiated as small circulating spots, small static spots, or large static areas of intense fluorescent emission. The small circulating spots and small static spots were due to individual rhodamine-labeled MFLs or small aggregates of MFLs, while the large static areas of intense fluorescence reflected larger MFL accumulation in the vessels. The amount of rhodamine-labeled MFL accumulation was calculated from time-lapse frames as follows (C.R.):

1. Images obtained with the x4 objective were used to provide an overview of the cranial window and to roughly analyze rhodamine-labeled MFL accumulation. An intensity threshold (typically on the order of 150–254 on a 256 gray-scale image) was adjusted to analyze the accumulation in terms of the total area covered by rhodamine-labeled MFLs. This area was then normalized to the microscope field of view and expressed as follows: An = (100 · At)/FOVt, where An is the normalized area (expressed as a percentage); At, the total area above the intensity threshold; and FOVt, the total field of view. Only regions in which the surface area above the intensity threshold was larger than 100 square pixels (1500 µm2) were considered.

2. Images obtained with the x10 objective were used to analyze individual rhodamine-labeled MFLs or small MFL aggregates in the intracerebral microvessels 15 and 60 minutes after injection. Small fluorescent spots—that is, areas smaller than 20 square pixels (50 µm2)—within the intracerebral vessels in the reconstituted windows were counted and then normalized to the field of view. The number of spots per 0.025 mm2 was documented.

Cerebral Blood Flow Measurements
To determine whether the rhodamine-labeled MFLs modified the cerebral circulation, we measured the local blood flow semiquantitatively by using laser Doppler flowmetry (Laser-Doppler Flowmeter, model MBF 3D; Moor Instruments, Axminster, England). The laser probe was placed in the center of the cranial window, and two authors (Y.T., M.S.M.) in consensus performed the measurements before and after in vivo imaging. In addition, a two-dimensional flow mapping method was used to measure the cerebral blood flow when the mouse was placed under the dynamic confocal microscope (Appendix E2; http://radiology.rsnajnls.org/cgi/content/full/2442060912/DC1) (Y.T.). This method has been described in detail elsewhere (32,33). Briefly, the bolus injection of FITC-dextran enabled determination of the regional mean transit time of blood. The microflow (1 divided by the mean transit time of blood) was tentatively deduced, displayed on a two-dimensional map, and plotted as a frequency distribution (histogram). The local microflow was measured at the start and end of the dynamic experiment. Consequently, a bolus of FITC-dextran was injected twice (at each measurement time).

Histologic Analysis
Histologic analysis was conducted to confirm the microvascular localization of the rhodamine-labeled MFLs and to determine if MFLs were present outside the area of the window and in the depth of the parenchyma. To allow the MFLs to stay within the vessel to study their vascular distribution, no fixative was used intracardially. At the end of the experiments, the mice were sacrificed by means of isoflurane overdose (A.T.D.). The brain was carefully removed (A.T.D.) and then immersed in 4% paraformaldehyde in 0.1 mol/L phosphate buffer (pH, 7.4) at 4°C for 4 days and in 20% sucrose in phosphate buffer (for cryopreservation) for 2 days. A cryomicrotome (model CM3050S; Leica, Wetzlar, Germany) was used to cut 45-µm-thick coronal sections around the area corresponding to the window (A.T.D.). Three authors (J.S., E.P., and M.S.M., with 10 years, 10 years, and 1 year of experience, respectively) in consensus analyzed the sections by using the laser-scanning confocal fluorescence microscope (model MRC 600) with a x10 objective and appropriate zooming and a microscope (model CLSM 510; Carl Zeiss) with a Plan Apochromat 63x/1.4NO oil objective (M.S.M.) coupled to LSM 5 Image Browser software (Carl Zeiss).

Statistical Analyses
Data are expressed as means ± standard deviations. The paired Student t test was used to compare cerebral blood flow and two-dimensional flow values before and 1 hour after rhodamine-labeled MFL injection. For image analysis, an independent t test was used to perform comparisons between each group (Origin 6.0; Originlab, Northampton, Mass). P < .05 was considered to indicate significance.


   RESULTS
TOP
 ABSTRACT
 INTRODUCTION
 MATERIALS AND METHODS
 RESULTS
DISCUSSION
 ADVANCES IN KNOWLEDGE
 References
 
Dynamic in Vivo Imaging
Circulating rhodamine-labeled MFLs were visualized in all vessels just after they were injected and during every 2-minute video monitoring period within 1 hour. In the magnet-free control animals, rhodamine-labeled MFL accumulation was negligible (mean accumulation, 0.3% ± 0.1 of area under the cranial window 60 minutes after injection), and few static MFL spots were seen (means of 0.1 spot per 0.025 mm2 ± 0.1 and 0.2 spot per 0.025 mm2 ± 0.1 15 and 60 minutes after injection, respectively) (Fig 2).


Figure 2
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  		Figure 2: Mean changes in the time dependency of overall rhodamine-labeled MFL accumulation within the field of view. Top: Graph shows mean changes in the area above an intensity threshold (normalized by using the total observed area) as a function of time after MFL injection. Three mice were control animals, and four were exposed to a magnet. Middle and bottom: Typical examples of dynamic confocal video-recorded images of fluorescent rhodamine-labeled MFLs in the brains of live mice from control (middle) and magnet (bottom) groups, viewed through a closed cranial window. The fields of view before (time zero) and up to 60 minutes after intravenous injection of the MFLs in a control mouse (middle) and magnet-exposed mouse (bottom) are shown. Note the MFL accumulation under the magnet. Scale bar = 500 µm.

 
With application of the magnet, no static fluorescence was observed in the arterioles, while the concentration of rhodamine-labeled MFLs in the capillaries and venules increased gradually in a two-phase manner. First, the area above a defined threshold of fluorescence intensity increased sharply for the first 15 minutes after injection, reached a plateau at 30 minutes, and remained stable from 30 to 60 minutes (mean accumulation, 3.6% ± 0.6 of area under the cranial window, corresponding roughly to a 15-fold increase in the normalized fluorescent area within the meningeal vessels compared with the area of fluorescence in the magnet-free control animals [Fig 2]) (P < .05). From 30 to 60 minutes, large areas (up to 5 mm2) of MFL accumulation were seen in the meningeal venules (10–100 µm in diameter). Second, between 15 and 30 minutes after the injection, small individual spots of MFL fluorescence or microdomains (<50 µm2) appeared in the cortical venules (Fig 3b). The mean number of these small spots on the venule walls increased from 0.6 spot per 0.025 mm2 ± 0.4 at 15 minutes to 1.7 spots per 0.025 mm2 ± 1.0 at 60 minutes (P < .05). This corresponds roughly to a 10-fold increase in the number of small spots in the cortical vessels in the magnet group compared with the number of small spots in the magnet-free control group.


Figure 3A
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  		Figure 3a: Dynamic in vivo images of rhodamine-labeled MFLs in the meningeal and intracerebral vessels of (a) a control mouse and (b) a mouse exposed to the magnet (x10 objective). Note the accumulation of small MFL aggregates in b. In all images, only immobile rhodamine-labeled MFLs were quantified. The mean number of small fluorescent spots counted in the intracerebral vessels in each reconstituted window and normalized to the field of view is shown for both groups. The control group comprised three mice, and the magnet group comprised four. The numbers of spots per 10 000 square pixels (0.025 mm2) are given.

 

Figure 3B
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  		Figure 3b: Dynamic in vivo images of rhodamine-labeled MFLs in the meningeal and intracerebral vessels of (a) a control mouse and (b) a mouse exposed to the magnet (x10 objective). Note the accumulation of small MFL aggregates in b. In all images, only immobile rhodamine-labeled MFLs were quantified. The mean number of small fluorescent spots counted in the intracerebral vessels in each reconstituted window and normalized to the field of view is shown for both groups. The control group comprised three mice, and the magnet group comprised four. The numbers of spots per 10 000 square pixels (0.025 mm2) are given.

 
No reversibility of rhodamine-labeled MFL accumulation occurred in the cortical venules during the image acquisitions. In contrast, in the meningeal venules, a part of the MFL accumulation was transient during the first video monitoring period (ie, after magnet removal at 15 minutes) but not thereafter.

Static in Vivo Imaging
Sixty minutes after injection of the rhodamine-labeled MFLs, the FITC-labeled microvascular network was visible in boththe control and the magnet groups (Fig 4, A1, B1). Rhodamine-labeled MFLs were present in the intracerebral venules and capillaries in both groups (Fig 4, A2, B2). No MFL accumulation could be detected in the control group (Fig 4, A2, A3), but in the magnet group, accumulation in the venules was clearly visible and formed regions of intense rhodamine fluorescence (Fig 4, B2, B3) not only in the meningeal venules but also in the superficial cortical venules (inset in Fig 4, B3).


Figure 4
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  		Figure 4: In vivo static confocal images of rhodamine-labeled MFLs in brain vessels 60 minutes after injection. Left: The vascular network is labeled with FITC-dextran, a green fluorescent tracer, in the control (A1) and magnet (B1) groups. Middle: Sixty minutes after injection, only diffuse rhodamine fluorescence due to the labeled MFLs is seen within the vascular network in the control group (A2), whereas in the magnet group (B2), MFL accumulation in the cerebral venules is clearly visible. Right: Consequently, the merged images show only the FITC-induced green fluorescence of the vascular network in the control group (A3) but intense red fluorescent regions in the green fluorescent microvessels in the magnet group (B3). Rhodamine-labeled MFLs have accumulated not only as large areas in vessels but also as small aggregates of a few individual MFLs (inset in B3).

 
Cerebral Blood Flow
There was no significant difference between the local blood flow before and that after injection of the rhodamine-labeled MFLs in either group (P > .05), as determined with laser Doppler flowmetry (P > .05) and two-dimensional flow measurement (P > .05) (Table, Fig 5).


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  		Mean Microflow Values before and after Intravenous Injection of Rhodamine-labeled MFLs

 

Figure 5
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  		Figure 5: Two-dimensional blood flow analysis after injection of FITC-dextran. The two-dimensional map (left) and histogram (right) of each flow measurement are shown for a control mouse (top) and a mouse exposed to a magnet (bottom). Blood flow was not significantly modified after 60 minutes (t60), with or without the magnet (P > .05). Scale bar = 400 µm. Freq. = frequency, MTT = mean transit time, t0 = before MFL injection.

 
Histologic Analysis
Examination of the histologic slices revealed a vascular network labeled with FITC-dextran–induced green fluorescence and rhodamine-labeled MFL–induced red fluorescence; these findings confirmed and expanded the dynamic and static in vivo imaging data. In the hemisphere that was equipped with the cranial window and exposed to the magnetic field gradient, rhodamine-labeled MFLs were present in the intracerebral venules and capillaries to a larger extent than they were in the contralateral hemisphere venules and capillaries (Fig 6). The presence of MFLs in the brain tissue could not be ascertained because intravascular rhodamine-labeled MFLs were not washed out by a fixative in these experiments.


Figure 6A
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  		Figure 6a: Images of histologic sections. (a) On a reconstructed brain section (magnification, x10) from a mouse exposed to a magnet, rhodamine-labeled MFLs are clearly visible in the cortical venules in the hemisphere exposed to the magnetic field gradient (right), whereas the MFLs are hardly distinguishable in the contralateral hemisphere (left). (b) Higher-magnification (x63) view of the sections in a. Left (control group) and right (magnet group) images represent the sum of 40 images.

 

Figure 6B
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  		Figure 6b: Images of histologic sections. (a) On a reconstructed brain section (magnification, x10) from a mouse exposed to a magnet, rhodamine-labeled MFLs are clearly visible in the cortical venules in the hemisphere exposed to the magnetic field gradient (right), whereas the MFLs are hardly distinguishable in the contralateral hemisphere (left). (b) Higher-magnification (x63) view of the sections in a. Left (control group) and right (magnet group) images represent the sum of 40 images.

 
The presence of numerous spots with intense rhodamine fluorescence in the hemisphere that was equipped with the cranial window and exposed to the magnetic field gradient indicates that rhodamine-labeled MFLs accumulated in this hemisphere owing to a magnetic attraction.


   DISCUSSION
TOP
 ABSTRACT
 INTRODUCTION
 MATERIALS AND METHODS
 RESULTS
 DISCUSSION
ADVANCES IN KNOWLEDGE
 References
 
By using in vivo dynamic imaging, we found that an externally applied magnetic force led to the significant accumulation of optimized rhodamine-labeled MFLs in the cerebral venules (P < .05). The normalized fluorescent area within the meningeal and cortical vessels increased 15- and 10-fold, respectively, without significantly affecting local blood flow, as measured with laser Doppler flowmetry and two-dimensional flow measurement.

MFL Properties and Mechanism of Magnetic Targeting
Circulating rhodamine-labeled MFLs were visualized in the brain microvasculature of meningeal and cortical areas examined through a cranial window for at least 1 hour after the MFLs were systemically injected, regardless of magnet application. Thus, these nanosized (180-nm) PEGylated MFLs, with combined properties of slow clearance and efficient recirculation, appear to be good candidates for passive drug targeting and to fulfill major requirements for efficient active targeting, such as magnetic targeting.

High concentrations (estimated nanoparticle volume fraction, 1%–4%) of superparamagnetic nanoparticles were encapsulated in rhodamine-labeled MFLs and thus led to the magnetic mobility of the MFLs, on the order of 0.6 µm/sec for a single rhodamine-labeled MFL exposed to the magnetic field gradient described herein (10). Circulating liposomes also experience dipole-dipole interaction forces owing to already immobilized liposomes; these forces may generate a magnetic attraction that exceeds the simple attraction of the external magnet by three orders of magnitude. In comparison, the mean velocity of erythrocytes through capillaries is about 0.5 mm/sec under the same experimental conditions (31).

Moreover, low levels of nonspecific interactions between nonfunctionalized PEGylated liposomes and endothelial cells have been reported (34–36). Although many other factors may play a role, we suggest that the accumulation of rhodamine-labeled MFLs in meningeal and cortical venules is due first to reduced blood flow in the direct proximity of the venule walls, which allows the magnetically driven liposomes to rest on the venule walls, and second to dipole-dipole interactions between liposomes in the presence of the magnetic field. This hypothesis of a two-step kinetic process for rhodamine-labeled MFL magnetic targeting is in accordance with our dynamic imaging findings.

Because the dipolar interaction force is directly proportional to the number of accumulated magnetic liposomes, liposomes appear to accumulate in a sort of chain reaction: The higher the number of trapped rhodamine-labeled MFLs, the deeper the MFLs are stopped. In our study, the MFLs were trapped in cortical venules within 15–30 minutes of the magnet exposure.

Limitations
The results of this study in mice provide proof of principle, but many obstacles must be overcome before clinical use can be envisaged. Magnetic targeting could lead to substantial clinical applications for brain treatment if rhodamine-labeled MFLs are able to cross the blood-brain barrier. Complementary analysis of this part of the process is necessary to address this key issue. Moreover, although no significant change in local blood flow was measured after injection of the rhodamine-labeled MFLs or during magnetic targeting, the likelihood that greater differences would have occurred with use of a larger number of animals cannot be ruled out.

Practical applications: MFLs appear to enable spatial and temporal control of drug delivery to selected areas of the cerebrovascular network. This type of biophysical targeting is particularly suited for the brain, where drug delivery is crucial. Targeting could be possible in many human brain regions with use of a magnetic needle or wire placed close to the target lesion after craniotomy (37).

Moreover, in addition to allowing spatial drug targeting by means of magnetic guidance, the properties of the encapsulated magnetic material, together with rhodamine labeling of the lipid bilayer, also enable monitoring of the rhodamine-labeled MFLs independently with two important and complementary noninvasive imaging modalities in the research environment: optical imaging and MR imaging. MR imaging is a powerful imaging technique that is particularly well suited for high-spatial-resolution brain mapping because it allows noninvasive follow-up of brain therapy (38). MFLs have high contrast properties (10,25) and could be of interest for many long-term therapeutic protocols, including hyperthermia treatment (39).

Use of an optical imaging setup enabled real-time dynamic in vivo analysis of the rhodamine-labeled MFL magnetic targeting process in animal models. With the development of optical setups for noninvasive imaging (40–42), together with fluorescent biotracers (43,44), rhodamine-labeled MFLs appear to be versatile medication and contrast agent carriers.


   ADVANCES IN KNOWLEDGE
TOP
 ABSTRACT
 INTRODUCTION
 MATERIALS AND METHODS
 RESULTS
 DISCUSSION
 ADVANCES IN KNOWLEDGE
References
 


   ACKNOWLEDGMENTS
 
The authors thank Valerie Nicolas (Plateau Technique-Imagerie Cellulaire IFR141) for technical assistance and performing the confocal studies, Delphine Talbot for nanoparticle characterization, and Nathalie Kubis for help and fruitful discussions.


   FOOTNOTES
 

Abbreviations: FITC = fluorescein isothiocyanate • MFL = magnetic fluid–loaded liposome

Author contributions: Guarantors of integrity of entire study, all authors; study concepts/study design or data acquisition or data analysis/interpretation, all authors; manuscript drafting or manuscript revision for important intellectual content, all authors; manuscript final version approval, all authors; literature research, C.R., M.S.M., C.W., S.L.; experimental studies, all authors; statistical analysis, C.R., M.S.M., Y.T., S.L.; and manuscript editing, C.R., M.S.M., Y.T., E.P., S.L., F.G., J.S.

Authors stated no financial relationship to disclose.


   References
TOP
 ABSTRACT
 INTRODUCTION
 MATERIALS AND METHODS
 RESULTS
 DISCUSSION
 ADVANCES IN KNOWLEDGE
 References
 

  1. Vinogradov SV, Bronich TK, Kabanov AV. Nanosized cationic hydrogels for drug delivery: preparation, properties and interactions with cells. Adv Drug Deliv Rev 2002;54:135–147.[CrossRef][Medline]
  2. Vinogradov SV, Batrakova EV, Kabanov AV. Nanogels for oligonucleotide delivery to the brain. Bioconjug Chem 2004;15:50–60.[CrossRef][Medline]
  3. Pardridge WM. Vector-mediated drug delivery to the brain. Adv Drug Deliv Rev 1999;36:299–321.[CrossRef][Medline]
  4. Begley DJ. Delivery of therapeutic agents to the central nervous system: the problems and the possibilities. Pharmacol Ther 2004;104:29–45.[CrossRef][Medline]
  5. Garcia-Garcia E, Andrieux K, Gil S, Couvreur P. Colloidal carriers and blood-brain barrier (BBB) translocation: a way to deliver drugs to the brain? Int J Pharm 2005;298:274–292.[CrossRef][Medline]
  6. Luciani A, Olivier JC, Clement O, et al. Glucose-receptor MR imaging of tumors: study in mice with PEGylated paramagnetic niosomes. Radiology 2004;231:135–142.[Abstract/Free Full Text]
  7. Torchilin VP. Polymeric contrast agents for medical imaging. Curr Pharm Biotechnol 2000;1:183–215.[CrossRef][Medline]
  8. Tilcock C. Delivery of contrast agents for magnetic resonance imaging, computed tomography, nuclear medicine and ultrasound. Adv Drug Deliv Rev 1999;37:33–51.[CrossRef][Medline]
  9. Weinmann HJ, Ebert W, Misselwitz B, Schmitt-Willich H. Tissue-specific MR contrast agents. Eur J Radiol 2003;46:33–44.[CrossRef][Medline]
  10. Martina MS, Fortin JP, Ménager C, et al. Generation of superparamagnetic liposomes revealed as highly efficient MRI contrast agents for in vivo imaging. J Am Chem Soc 2005;127:10676–10685.[CrossRef][Medline]
  11. Lasic DD, Templeton NS. Liposomes in gene therapy. Adv Drug Deliv Rev 1996;20:221–266.[CrossRef]
  12. Lasic DD. Novel applications of liposomes. Trends Biotechnol 1998;16:307–321.[CrossRef][Medline]
  13. Marjan JM, Allen TM. Long circulating liposomes: past, present and future. Biotechnol Adv 1996;14:151–175.[CrossRef][Medline]
  14. Allen TM, Hansen C, Martin F, Redemann C, Yau-Young A. Liposomes containing synthetic lipid derivatives of poly(ethylene glycol) show prolonged circulation half-lives in vivo. Biochim Biophys Acta 1991;1066:29–36.[Medline]
  15. Needham D, Hristova K, McIntosh TJ, Dewhirst MW, Wu N, Lasic DD. Polymer-grafted liposomes: physical basis for the "stealth" property. J Liposome Res 1992;2:411–430.
  16. Papahadjopoulos D, Allen T, Gabizon A, et al. Sterically stabilized liposomes: improvements in pharmacokinetics and antitumor therapeutic efficacy. Proc Natl Acad Sci U S A 1991;88:11460–11464.[Abstract/Free Full Text]
  17. Fabel K, Dietrich J, Hau P, et al. Long-term stabilization in patients with malignant glioma after treatment with liposomal doxorubicin. Cancer 2001;92:1936–1942.[CrossRef][Medline]
  18. Hau P, Fabel K, Baumgart U, et al. Pegylated liposomal doxorubicin-efficacy in patients with recurrent high-grade glioma. Cancer 2004;100:1199–1207.[CrossRef][Medline]
  19. Lubbe AS, Alexiou C, Bergemann C. Clinical applications of magnetic drug targeting. J Surg Res 2001;95:200–206.[CrossRef][Medline]
  20. Hafeli UO. Magnetically modulated therapeutic systems. Int J Pharm 2004;277:19–24.[CrossRef][Medline]
  21. Lubbe AS, Bergemann C, Huhnt W, et al. Preclinical experiences with magnetic drug targeting: tolerance and efficacy. Cancer Res 1996;56:4694–4701.[Abstract/Free Full Text]
  22. Lubbe AS, Bergemann C, Riess H, et al. Clinical experiences with magnetic drug targeting: a phase I study with 4'-epidoxorubicin in 14 patients with advanced solid tumors. Cancer Res 1996;56:4686–4693.[Abstract/Free Full Text]
  23. Alexiou C, Arnold W, Klein RJ, et al. Locoregional cancer treatment with magnetic drug targeting. Cancer Res 2000;60:6641–6648.[Abstract/Free Full Text]
  24. Alexiou C, Jurgons R, Schmid RJ, et al. Magnetic drug targeting: biodistribution of the magnetic carrier and the chemotherapeutic agent mitoxantrone after locoregional cancer treatment. J Drug Target 2003;11:139–149.[CrossRef][Medline]
  25. Fortin-Ripoche JP, Martina MS, Gazeau F, et al. Magnetic targeting of magnetoliposomes to solid tumors with MR imaging monitoring in mice: feasibility. Radiology 2006;239:415–424.[Abstract/Free Full Text]
  26. Lesieur S, Grabielle-Madelmont C, Menager C, et al. Evidence of surfactant-induced formation of transient pores in lipid bilayers by using magnetic-fluid-loaded liposomes. J Am Chem Soc 2003;125:5266–5267.[CrossRef][Medline]
  27. Massart R. Preparation of aqueous magnetic liquids in alkaline and acidic media. IEEE Trans Magn 1981;17:1247–1248.[CrossRef]
  28. Wilhelm C, Gazeau F, Bacri JC. Magnetophoresis and ferromagnetic resonance of magnetically labeled cells. Eur Biophys J 2002;31:118–125.[CrossRef][Medline]
  29. Tomita Y, Kubis N, Calando Y, et al. Long-term in vivo investigation of mouse cerebral microcirculation by fluorescence confocal microscopy in the area of focal ischemia. J Cereb Blood Flow Metab 2005;25:858–867.[CrossRef][Medline]
  30. Franklin KBJ, Paxinos G. The mouse brain in stereotaxic coordinates. San Diego, Calif: Academic Press, 1997.
  31. Seylaz J, Charbonne R, Nanri K, et al. Dynamic in vivo measurement of erythrocyte velocity and flow in capillaries and of microvessel diameter in the rat brain by confocal laser microscopy. J Cereb Blood Flow Metab 1999;19:863–870.[CrossRef][Medline]
  32. Schiszler I, Tomita M, Fukuuchi Y, Tanahashi N, Inoue K. New optical method for analyzing cortical blood flow heterogeneity in small animals: validation of the method. Am J Physiol Heart Circ Physiol 2000;279:H1291–H1298.[Abstract/Free Full Text]
  33. Tomita Y, Tomita M, Schiszler I, et al. Moment analysis of microflow histogram in focal ischemic lesion to evaluate microvascular derangement after small pial arterial occlusion in rats. J Cereb Blood Flow Metab 2002;22:663–669.[Medline]
  34. Gosk S, Vermehren C, Storm G, Moos T. Targeting antitransferrin receptor antibody (OX26) and OX26-conjugated liposomes to brain capillary endothelial cells using in situ perfusion. J Cereb Blood Flow Metab 2004;24:1193–1204.[Medline]
  35. Mulder WJ, Strijkers GJ, Habets JW, et al. MR molecular imaging and fluorescence microscopy for identification of activated tumor endothelium using a bimodal lipidic nanoparticle. FASEB J 2005;19:2008–2010.[Abstract/Free Full Text]
  36. Cansell M, Parisel C, Jozefonvicz J, Letourneur D. Liposomes coated with chemically modified dextran interact with human endothelial cells. J Biomed Mater Res 1999;44:140–148.[CrossRef][Medline]
  37. Iacob G, Rotariu O, Strachan NJ, Hafeli UO. Magnetizable needles and wires: modeling an efficient way to target magnetic microspheres in vivo. Biorheology 2004;41:599–612.[Medline]
  38. Henson JW, Gaviani P, Gonzalez RG. MRI in treatment of adult gliomas. Lancet Oncol 2005;6:167–175.[CrossRef][Medline]
  39. Pardoe H, Clark PR, St Pierre TG, Moroz P, Jones SK. A magnetic resonance imaging based method for measurement of tissue iron concentration in liver arterially embolized with ferrimagnetic particles designed for magnetic hyperthermia treatment of tumors. Magn Reson Imaging 2003;21:483–488.[CrossRef][Medline]
  40. Laemmel E, Genet M, Le Goualher G, Perchant A, Le Gargasson JF, Vicaut E. Fibered confocal fluorescence microscopy (Cell-viZio) facilitates extended imaging in the field of microcirculation: a comparison with intravital microscopy. J Vasc Res 2004;41:400–411.[CrossRef][Medline]
  41. Mehta AD, Jung JC, Flusberg BA, Schnitzer MJ. Fiber optic in vivo imaging in the mammalian nervous system. Curr Opin Neurobiol 2004;14:617–628.[CrossRef][Medline]
  42. Al-Gubory KH. Fibered confocal fluorescence microscopy for imaging apoptotic DNA fragmentation at the single-cell level in vivo. Exp Cell Res 2005;310:474–481.[CrossRef][Medline]
  43. Licha K, Olbrich C. Optical imaging in drug discovery and diagnostic applications. Adv Drug Deliv Rev 2005;57:1087–1108.[CrossRef][Medline]
  44. Weissleder R, Ntziachristos V. Shedding light onto live molecular targets. Nat Med 2003;9:123–128.[CrossRef][Medline]




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