Gum Arabic-Coated Magnetic Nanoparticles for Potential Application
in Simultaneous Magnetic Targeting and Tumor Imaging
Lei Zhang,1,2,3Faquan Yu,3Adam J. Cole,3Beata Chertok,3Allan E. David,3
Jingkang Wang,1,2and Victor C. Yang1,2,3,4
Received 13 July 2009; accepted 30 September 2009; published online 20 October 2009
Abstract. Magnetic iron oxide nanoparticles (MNP) coated with gum arabic (GA), a biocompatible
phytochemical glycoprotein widely used in the food industry, were successfully synthesized and
characterized. GA-coated MNP (GA-MNP) displayed a narrow hydrodynamic particle size distribution
averaging about 100 nm; a GA content of 15.6% by dry weight; a saturation magnetization of 93.1 emu/g
Fe; and a superparamagnetic behavior essential for most magnetic-mediated applications. The GA
coating offers two major benefits: it both enhances colloidal stability and provides reactive functional
groups suitable for coupling of bioactive compounds. In vitro results showed that GA-MNP possessed a
superior stability upon storage in aqueous media when compared to commercial MNP products currently
used in magnetic resonance imaging (MRI). In addition, significant cellular uptake of GA-MNP was
evaluated in 9L glioma cells by electron spin resonance (ESR) spectroscopy, fluorescence microscopy,
and MRI analyses. Based on these findings, it was hypothesized that GA-MNP might be utilized as a
MRI-visible drug carrier in achieving both magnetic tumor targeting and intracellular drug delivery.
Indeed, preliminary in vivo investigations validate this clinical potential. MRI visually confirmed the
accumulation of GA-MNP at the tumor site following intravenous administration to rats harboring 9L
glioma tumors under the application of an external magnetic field. ESR spectroscopy quantitatively
revealed a 12-fold increase in GA-MNP accumulation in excised tumors when compared to contralateral
normal brain. Overall, the results presented show promise that GA-MNP could potentially be employed
to achieve simultaneous tumor imaging and targeted intra-tumoral drug delivery.
KEY WORDS: brain tumor; drug delivery; gum arabic; magnetic nanoparticle; magnetic targeting.
Recent advancements in nanotechnology have shed light
on utilizing magnetic iron oxide nanoparticles (MNP) as drug
carriers. With their unique response to an external magnetic
field, MNP can be both passively and actively targeted to
tumors. Passively, the size of nano-carriers enables them to
extravasate into tumor interstitium through the hyperperme-
able vasculature characteristic of solid tumors, a phenomenon
known as the enhanced permeability and retention (EPR)
effect (1). Actively, the magnetic properties of MNP allow
their targeting to and accumulation at the tumor site with the
aid of an external magnetic field, a phenomenon typically
referred to as magnetic targeting (2). In addition, MNP
reduce both T1and T2/T2* relaxation times, rendering them
“visible” in magnetic resonance imaging (MRI)—a non-
invasive, high spatial resolution clinical modality that has
been used extensively in tumor diagnosis (3).
Polymer-coated MNP, having core–shell structure, typi-
cally consists of a mono- or multi-crystalline magnetic, iron
oxide (magnetite or maghemite) core coated with a biocom-
patible polymer shell. In one respect, the coating is important
because it can provide functional groups for conjugation of
drug molecules and/or targeting ligands. In another, the
coating can alter the surface properties of “bare” MNP,
altering in vivo stability and circulation half-life. With respect
to both stability and half-life, nanoparticle aggregation is of
greatest concern as aggregated MNP are no longer nano-
sized. Larger particles are rapidly identified by the reticu-
loendothelial system (RES), cleared by the liver and spleen,
and are unable to extravasate tumor vasculature, resulting in
minimal tumor accumulation via the EPR effect. Normally,
iron oxide cores alone aggregate quickly at physiological pH
and in ionic environments due to van der Waals attraction
between particles (4). Polymer coatings, however, create
steric hindrance to such forces in addition to providing a
resistance to RES clearance—the most common route of
MNP elimination in vivo. Hydrophilic polymers are normally
preferred as they provide a strong steric barrier to opsonin
adsorption, thereby improving circulation half-life (5).
1Tianjin Key Laboratory for Modern Drug Delivery and High
Efficiency, Tianjin University, Tianjin 300072, China.
2School of Chemical Engineering, Tianjin University, Tianjin 300072,
3Department of Pharmaceutical Sciences, College of Pharmacy,
University of Michigan, 428 Church Street, Ann Arbor( Michigan
4To whom correspondence should be addressed. (e-mail: vcyang@
The AAPS Journal, Vol. 11, No. 4, December 2009 (#2009)
1550-7416/09/0400-0693/0#2009 American Association of Pharmaceutical Scientists
Gum arabic (GA) is a nontoxic, hydrophilic, phytochem-
ical glycoprotein polymer widely used as a stabilizer in the
food and pharmaceutical industries. It is a heterogeneous
polymer comprising three main components: low-protein
content arabinogalactan (90%); high-protein content arabi-
nogalactan (10%); and high-protein content glycoproteins
(<1%) (6,7). Although not fully understood, the most widely
accepted structure of GA can be described as a number of
arabinogalactan units link to a polypeptide chain. Recently,
GA has been used to functionalize and stabilize nano-
particles. GA molecules contain charged groups (amine and
carboxyl) that can physically adsorb onto the surface of a
nanoparticle (6,8). Through its highly branched polysacchar-
ide structures, GA causes steric repulsion between nano-
particles to improve colloidal stability. Indeed, it has been
demonstrated that GA-coated gold nanoparticles were stable
in human serum albumin solution and strong ionic environ-
ments (6). GA has also enhanced nanoparticle stability in vivo
(9). Furthermore, GA contains abundant carboxyl groups that
can be easily activated and readily linked with other biomole-
cules (10). As an example, doxorubicin was conjugated to a
GA-coated MNP via a pH-sensitive hydrazone bond (11).
In this work, we developed a facile method to prepare a
GA-coated MNP (GA-MNP). Results showed that GA-MNP
exhibits good stability in a physiologically simulated environ-
ment. In addition, we showed that GA can be readily
conjugated with bioactive agents using the fluorophore rhod-
amine B as a model compound. In vitro studies revealed a
significant cellular uptake of rhodamine-linked GA-MNP in
9L glioma cells. Preliminary in vivo investigations in rats
bearing 9L glioma tumors showed, for the first time, that GA-
MNP could potentially be used as a MRI-visible drug carrier,
magnetically targetable to tumors.
MATERIALS AND METHODS
Unless otherwise stated, FeCl2·4H2O (>99%),
FeCl3·6H2O (>99%), NH3·H2O (NH3 content, 28∼30%),
GA, 2-[N-morpholino] ethane sulfonic acid (MES), 1-ethyl-3-
(3-dimethylaminopropyl) carbodiimide hydrochloride (EDC),
and all the other chemicals and solvents were purchased from
Sigma-Aldrich (St. Louis, MO). Reagents for cell culture
were purchased from Gibco Invitrogen (Carlsbad, CA).
Dialysis bags were obtained from Spectrum Laboratories,
Inc. (Rancho Dominguez, CA). Water was deionized (dH2O)
on a Milli-Q water purification system (Millipore, Billerica,
MA). Commercial MNP products used for in vitro stability
comparisons, including fluidMAG-D (starch coating), fluid-
MAG-CMX (carboxymethyl dextran coating), fluidMAG-
Heparin (heparin coating), and fluidMAG-DEAE (dextran
diethylaminoethyl coating) were provided by Chemicell®
(Berlin, Germany). These MNP products all possess a
hydrodynamic diameter of approximately 100 nm.
Preparation of GA-MNP
Magnetite (Fe3O4) MNP were synthesized using a
previously reported co-precipitation method (12). In brief, a
solution containing ferric chloride and ferrous chloride was
added dropwise to a 1.5 M NaOH solution under vigorous
mechanical stirring and nitrogen gas protection at room
temperature. The reaction temperature was gradually
increased to 75°C and held for 1 h under stirring and nitrogen
gas protection. The synthesized MNP product were separated
using a magnetic separator and washed five times with dH2O.
GA-MNP were prepared by adding 1 ml of the above
MNP (10 mg Fe/ml) to 5 ml of 10% GA solution. The GA/
MNP mixture was vortexed for 1 min and sonicated for
10 min to form a transparent colloidal solution. GA-MNP
were then purified using a magnetic separator and washed
five times with dH2O.
Characterization of GA-MNP
Phase composition of lyophilized magnetic nanoparticles
powder was analyzed with a Rigaku Rotoflex 200B 12KW
rotating anode X-ray diffractometer (RIGAKU, Inc., Tokyo,
Japan) equipped with Cu-Kα radiation (λ=1.54056 Å). Mor-
phology of GA-MNP was obtained on a JEOL 3010 high-
resolution transmission electron microscope (HR-TEM,
JEOL, Ltd., Tokyo, Japan). Samples were prepared by
dropping diluted particle suspensions on copper grids coated
with formvar film (Ted Pella, Inc., Redding, CA) followed by
drying at room temperature. Hydrodynamic particle size was
measured by photon correlation spectroscopy using a PSS
Nicomp 380 ZLS particle sizing system (Nicomp, Inc., Santa
Barbara, CA). Iron concentration of all MNP samples was
determined by inductively coupled plasma optical emission
spectroscopy using an Optima 2000 DV spectrometer (Perkin-
Elmer, Inc., Boston, MA). Magnetization measurements were
performed using an MPMS-XL superconducting quantum inter-
ference device (SQUID) magnetometer (Quantum Design Inc.,
San Diego, CA). To measure relative GA content, GA-MNP
were lyophilized and analyzed by thermogravimetric analysis
(TGA) using a TGA-7 instrument (Perkin-Elmer, Inc.).
GA-MNP Stability and Comparison with Other Commercial
In vitro stability of GA-MNP and other MNP products
was assessed in a physiological simulated buffer of Dulbecco’s
phosphate-buffered saline (PBS, with 0.901 mM Ca2+and
0.493 mM Mg2+) containing 10% fetal bovine serum (FBS).
Briefly, MNP samples were added to the above solution and
shaken to homogeneity (final iron concentration, 50 μg/ml).
Immediately afterwards, aliquots containing 0.2 ml of the
sample were loaded onto a 96-well plate, and the turbidity in
the wells was measured at 400 nm using a micro-plate reader
(Powerwave X340, BioTek Instruments, Inc., Winooski, VT) at
5-min intervals for 1 h (13). In addition, 3 ml of each test sample
was added to a 20-ml glass scintillation vial for visual
observation up through 24 h.
Cell Viability Evaluation
Rat 9L glioma cells (Brain Tumor Research Center,
University of California, San Francisco) were cultured at
37°C under a humidified atmosphere of 5% CO2in Dulbec-
co’s modified Eagle’s medium (DMEM) supplemented with
10% FBS, 100 IU/ml penicillin, 100 μg/ml streptomycin, and
694Zhang et al.
0.29 mg of L-glutamine. The cells were seeded on a 96-well
plate at ∼104cells/well and incubated with fresh media for
24 h. Following incubation, the culture media were replaced
with fresh media containing GA-MNP at various
concentrations (0.01–20 mg/ml), and cells were further
incubated at 37°C for 4 h. The GA-MNP solution was then
replaced with fresh media, and cell viability was measured
after 24 h using a standard (3-(4,5-dimethylthiazol-2-yl)-2,5-
diphenyltetrazolium bromide (MTT) assay (10).
Cellular Uptake of GA-MNP
RhB-GA was prepared by conjugating rhodamine B
(RhB) with GA through the formation of an amide bond as
previously described (10). Briefly, the GA solution was
prepared by dissolving 2 g GA in 20 ml 0.1 M MES, whereas
a RhB solution was prepared by dissolving 40 mg rhodamine
B in 1 ml 0.1 M MES. The two solutions were then mixed,
followed by the addition of 25 mg/ml EDC solution. The
mixture was stirred for 2 h at room temperature. The RhB-
GA conjugates were then purified by dialysis using dH2O for
3 days. The RhB-GA-MNP product were then prepared
analogously to GA-MNP as described above.
To carry out cell uptake studies, rat 9L glioma cells were
cultured in a 12-well plate and incubated with 80 μl of either
GA-MNP or the RhB-GA-MNP (5 mg Fe per milliliter) in
serum-free DMEM or DMEM with 10% FBS (120 μl) at
37°C for 2 h. After incubation, the cells were washed five
times with PBS and subjected to fluorescence and in vitro MR
To quantitatively compare the degree of cellular uptake
between GA-MNP and a commercial starch-coated MNP
(i.e., fluidMAG-D), rat 9L glioma cells were cultured in a 12-
well plate and incubated with the two different MNP samples
(at the same iron concentration ranging from 0.02 to 2.0 mg
Fe per milliliter) in 1 ml DMEM containing 10% FBS at 37°C
for 2 h. Following incubation, the cells were washed five times
with PBS, and cellular iron was measured by ESR spectro-
scopy according to a previously established protocol (2). ESR
spectra of samples were acquired using an EMX ESR
spectrometer (Bruker Instruments Inc., Billerica, MA). Cells
cultured without MNP were used as a control.
In Vivo Study
All animal experiments were conducted according to
protocols approved by the University of Michigan Committee
on Use and Care of Animals.
9L glioma cells were grown to confluency, harvested, and
resuspended in serum-free DMEM at a concentration of ∼104
cells per microliter. Ten microliters of cell suspension was
implanted in the right forebrains of Fisher 344 rats (body
weight, 200 g) at a depth of 3–4 mm beneath the skull through
a 1-mm diameter burr hole. The surgical field was cleaned
with 70% ethanol, and the burr hole was sealed with bone
wax (Ethicon Inc., Summerfield, NJ) to prevent extracerebral
extension of the tumor (13,14).
MRI experiments were performed using a 12-cm hori-
zontal bore, 7 Tesla Varian Unity Inova imaging system
(Varian, PaloAlto, CA) according to a previously established
procedure (2). Animals were anesthetized with 1.5% isoflur-
ane/air mixture and imaged using a 35-mm diameter quad-
rature RF head coil (USA Instruments Inc., OH). MRI of the
rat brain was initiated 10 days after tumor cell implantation.
Axial sections of the rat brain were acquired with a T2-
weighted fast spin echo sequence using the following param-
eters: repetition time (TR)=4 s, echo time (TE) = 60 ms, field
of view = 30×30 over 128×128 matrix, slice thickness = 1 mm,
slice separation = 1.5 mm, four signal averages per phase
encoding step. T2-weighted images were inspected to deter-
mine which slice corresponded to the best cross-sectional
visualization of the tumor. A single gradient echo (GE) scan
was acquired at this optimized position to provide qualitative
information on GA-MNP accumulation at the tumor site. GE
images were acquired with the following parameters: TR=
275.13 ms, TE=15 ms, field of view = 30×30 over 128×128
matrix, slice thickness = 1 mm.
For magnetic targeting, animals were anesthetized
through inhalation of a 1.5% isoflurane/air mixture. Rats
were then placed ventrally on a platform with the head
positioned between the poles of an electromagnet. GA-MNP
(in PBS) were injected at a dose of 12 mg Fe per kilogram
through the tail vein and retained in the magnetic field for
30 min. Animals dosed with GA-MNP, but not subject to
magnetic targeting, were used as controls. Animals were
imaged with MRI before the administration of GA-MNP and
after magnetic targeting (30 min after GA-MNP administra-
tion). To quantitatively examine the accumulation of GA-
MNP in the tumor and normal brain tissue, animals were
killed immediately following targeting. Brain tumors were
carefully dissected from the right hemisphere. Tumor and left
hemisphere tissues were analyzed by ESR as previously
RESULTS AND DISCUSSION
Figure 1 showed the X-ray diffraction (XRD) pattern of
GA-MNP with characteristic peaks of 2θ at 30.1°, 35.4°, 43.0°,
53.5°, 56.9°, and 62.6°, corresponding to the indices (220),
(311), (400), (422), (511), and (440), respectively, of the
magnetite crystal. In agreement with the results reported in
Fig. 1. X-ray diffraction pattern of GA-MNP lyophilized powder.
The XRD pattern confirmed that GA-MNP contained a magnetite
695Gam Arabic MNPs for Tumor Targeting and Imaging
the literature, XRD analysis confirmed that GA-MNP were
composed of pure magnetite with spinal crystal structure (16).
The morphology of GA-MNP was examined using
transmission electron microscopy (TEM). TEM image in
Fig. 2 visually showed the core–shell structure of GA-MNP.
The cores consisted of a number of spherical magnetite nano-
crystals with an average size of 14±3.8 nm coated with a shell of
GA polymer. The multi-nanocrystal cores were formed prob-
ably because the GA molecules, much larger than the nano-
crystals, could adsorb onto several nanocrystals through their
carboxyl groups and hold them together. The relative GA
content of GA-MNP by mass was found to be 15.6% by TGA.
For in vivo applications of magnetic nanoparticles, such
as magnetic targeting and MRI, superparamagnetic behavior
is of paramount importance. Superparamagnetic particles do
not remain magnetized in the absence of an external magnetic
field. Thus, they do not agglomerate without exposure to a
field. MNP only exhibit superparamagentic behavior, though,
below a certain size threshold, namely the size of a single
magnetic domain. The domain size for GA-MNP has been
determined to be below the apparent limit of 25 nm (17).
Indeed, negligible hysteresis was observed in magnetization
experiments (Fig. 3), confirming that GA-MNP possessed
superparamagnetic behavior. Overall, the synthesized GA-
MNP showed negligible coercivity (Hc) and remnant magnet-
ization (Mr), as well as a saturation magnetization value of
commercial products obtained from Chemicell® (94 emu/g Fe)
(15). Saturation magnetization was slightly lower than that of the
bulk magnetite (127 emu/g Fe), probably due to the fact that the
surface nonmagnetic layer (also known as “dead layer”) of GA-
MNP accounted for a higher composition fraction when the
hand, the saturation magnetization of GA-MNP was greater
than that of the FDA-approved contrast agent Feridex®
(70 emu/g Fe), possibly attributed to the smaller magnetite
crystal size of Feridex® (∼4.8 nm) (19). It has been demon-
strated that larger magnetite crystals yield stronger saturation
Particle size measurements (Fig. 4) showed that the
synthesized GA-MNP possessed a mean hydrodynamic
diameter of 118±12 nm in PBS buffer. It is important to note
that the hydrodynamic size of MNP is a key factor in
governing the success of magnetic targeting and reduced in
vivo clearance. It has been reported that attractive magnetic
forces on MNP smaller than 50 nm are not sufficient to
overcome forces from Brownian motion, resulting in poor
MNP accumulation at the target site when subject to magnetic
targeting (21). Thus, from a targeting standpoint, MNP with
large core structures and hydrodynamic sizes are preferred.
MNP over a certain size (e.g., >300 nm), however, are rapidly
cleared from the body via the RES, resulting in a significantly
shortened circulation half-life (22). Moreover, the size cutoff
of tumor vasculature permeation is on the order of several
hundred nanometers (1). Larger MNP and aggregates may
not be able to extravasate into the interstitial space of the
tumor, resulting in minimal tumor accumulation. In our
previous studies, we demonstrated that starch-coated MNP
with a mean size of 100 nm were capable of extravasation and
accumulation in tumor with the aid of magnetic targeting (2).
It is of little doubt that a 100 nm GA-MNP would yield
similar results. Further studies, though, are needed to validate
Fig. 2. TEM image of GA coated MNP (scale bar, 100 nm). The
image showed the core–shell structure of GA-MNP. The cores
consisted of magnetite nanocrystals coated by a GA polymer shell
(indicated by arrows)
Fig. 3. Magnetization of GA-MNP measured by SQUID. The figure
showed the induced magnetization of GA-MNP corresponding to
both increasing and decreasing external magnetic field. The lack of
hysteresis loop confirmed that the prepared GA-MNP possessed a
Fig. 4. Hydrodynamic particle size of GA-MNP in PBS buffer. GA-
MNP showed a narrow particle size distribution with mean of 118 ±
12 nm. The observed size was similar to commercial MNP products
696 Zhang et al.
Stability Evaluation and Comparison
MNP are typically administrated intravenously. There-
fore, stability prior to, during, and after administration is of
great importance in determining the fate of MNP. Among
various methods, turbidity is a sensitive and straightforward
means to monitor the stability of a colloidal suspension (4).
Through this method, we discovered that all MNP samples
investigated, including our GA-MNP product, remained
stable in water for a period of several months. When mixed
with physiologically simulated buffer, however, fluidMAG-
CMX, fluidMAG-Heparin, and fluidMAG-DEAE samples
became turbid immediately after suspension, suggesting rapid
agglomeration (Fig. 5). In contrast, the turbidity of GA-MNP
and fluidMAG-D samples remained unchanged over the
course of 1 h. Indeed, GA-MNP suspension was clear even
after 24 h of storage. The good stability of the GA-coated
MNP could be due to dynamic motion of the highly branched
polysaccharide chains in GA that promote strong steric
resistance between individual particles.
Cellular Uptake of GA-MNP
MNP and GA have been shown to be nontoxic and
biocompatible by other researchers (6,23). Consistent with
their findings, our MTT assay results indicated that GA-MNP
were not cytotoxic to 9L glioma cells in DMEM with and
without FBS even at an iron concentration as high as 20 mg
Fe per milliliter (data not shown). Fluorescence microscopy
demonstrated internalization of RhB-GA-MNP into 9L
glioma cells 2 h after incubation (Fig. 6a). It is speculated
that cellular uptake is attributed to endocytosis due to the
small size of GA-MNP, as observed by other investigators
studying other nanoparticles (24). No fluorescence signal was
found inside the nucleus, suggesting that RhB-GA-MNP
remains in the cytosol or endosome. This result was some-
what anticipated because MNP of ∼100 nm would have
difficulty penetrating through the narrow channels of nuclear
pores (diameter <40 nm) (10).
Consistent with results reported in the literature, in vitro
GE MR images further confirmed cellular uptake of GA-
MNP (19). Pronounced hypointensity was observed in cells
treated with GA-MNP when compared to the control,
indicating the presence of GA-MNP inside the cells (see
Fig. 6b). Furthermore, Fig. 6c revealed an iron concentration-
dependent drop in signal intensity on GE images for GA-
MNP, implicating higher cellular uptake of GA-MNP at
Quantitative measurements of cellular uptake showed
that both GA-MNP and starch-MNP exhibit a concentration-
dependent uptake of nanoparticles by 9L glioma cells (Fig. 7).
The GA-MNP product, however, consistently displayed a two
to threefold higher degree of uptake than starch-MNP when
compared at the same iron concentrations. This result
suggested that a GA-coated surface might possess a higher
affinity toward tumor cells than that of starch.
Brain Tumor Targeting and Imaging
Similar to findings recently reported by our laboratory
(2), GE MR images qualitatively revealed accumulation GA-
MNP in the brains of animals subjected to magnetic targeting.
Fig. 5. Physiologically simulated stability of GA-MNP and several
commercial MNP products including: heparin- (Heparine); GA-
(GAM); DEAE- (DEAE); CMDX- (CMDX); and starch-coated
(Starch) MNP in PBS (containing 10% FBS) medium. Turbidity
changes were measured as a function of time
Fig. 6. a Fluorescence microscopy of cellular uptake of RhB-GA-
MNP (scale bar, 10 μm). b In vitro gradient echo MR images of 9L
glioma cells, from top: 1 control cells, 2 cells treated by GA-MNP in
DMEM with 10% FBS, and 3 cells treated by GA-MNP in DMEM
without serum. c Gradient echo MR images of GA-MNP at iron
concentrations of 0, 5, 10, 20, and 50 ppm (from top to bottom)
Fig. 7. Cellular uptake of GA-MNP and starch-MNP by 9L glioma
cells after 2 h of incubation (n=3). *At all concentrations tested (from
0.02 to 2 mg/ml), cellular uptake of GA-MNP was higher than starch-
MNP (unpaired t test, p<0.05)
697Gam Arabic MNPs for Tumor Targeting and Imaging
As shown in Fig. 8a, tumor was clearly visible on a T2-weighted
image. When compared with the GE baseline image shown in
Fig. 8c clearly demonstrated the accumulation of GA-MNP at
the tumor site. In contrast, no detectable hypointensity was
foundinotherpartsofthe brain.SimilarMRIstudies werealso
conducted on brain tumor-harboring animals not subjected to
magnetic targeting, as shown in Fig. 8d–f; no obvious signal
change was observed in the GE image (Fig. 8f). The MR
images qualitatively proved that magnetic targeting improved
brain tumor accumulation.
Quantitative tissue analysis by ESR further confirmed
the selective accumulation of GA-MNP in tumor after
magnetic targeting. For targeted animals (Fig. 8g), GA-MNP
accumulation in the tumor tissue (63.8±14.6 nmol Fe per
gram tissue, n=4) was 12-fold higher than that found in the
contralateral brain tissue (5.3±3.5 nmol Fe per gram tissue,
n=4). For animals not targeted, GA-MNP accumulation in
tumors (8.4±5 nmol Fe per gram tissue, n=3) was threefold
higher than that found in the contralateral brain (2.9±
2.1 nmol Fe per gram tissue, n=3). Results from the control
group indicated that the enhanced permeability of the tumor
vasculature alone resulted in nanoparticle accumulation in
tumors. With the application of the external magnetic field,
though, tumor levels of MNP increased eightfold, suggesting
that magnetic targeting further enhanced nanoparticle accu-
mulation. Based on our recent findings (15), enhanced tumor
selectivity might be related to pathological alteration of blood
flow dynamics in the brain tumor. For instance, the linear
blood flow rate (0.31 cm/s) in normal brain vessels is
estimated to be about fourfold higher than that (0.074 cm/s)
in brain tumors (15). A slower flow rate gives GA-MNP
longer residence time in the microvasculature of the tumor.
Furthermore, attraction of GA-MNP to an external magnetic
field further enhances retention and accumulation of GA-
MNP at the tumor. Overall, the selectivity of GA-MNP to the
tumor confirms its in vivo stability. Since GA-MNP were able
to extravasate the tumor, as evidenced visually by the clear
hypointense MR images, it appeared that GA-MNP did not
aggregate during in vivo application.
MNP were successfully stabilized and functionalized
with nontoxic, phytochemical GA polymer via a facile
procedure. Results showed that GA-MNP product possessed
several major advantages over existing commercial MNP
products: (1) improved stability under physiological condi-
tions; (2) reactive functional groups for easy linking with
biofunctional compounds or other molecules to improve their
performance; and (3) a high level of cellular uptake by tumor
cells. In vivo magnetic targeting studies revealed that GA-
MNP accumulated in the brain tumor 12-fold higher than the
normal brain. The observed selectivity in tumors was visibly
monitored by MRI. In conclusion, GA-MNP showed great
promise as a tool for achieving simultaneous imaging of and
magnetically targeted drug delivery to virtually all types of
solid tumors, including the brain model presented. To that
end, further animal investigations are currently underway in
This work was supported in part by NIH R01 grants
CA114612, NS066945, and the Hartwell Foundation Biomed-
ical Research Award. In addition, this work was also partially
sponsored by the World Class University (WCU) program
through the Korea Science and Engineering Foundation
funded by the Ministry of Education, Science and Technology
(R31-2008-000-10103-01). Lei Zhang was a recipient of the
Chinese Program of Introducing Talents of Discipline to
Universities, no. B06006. Victor C. Yang is currently a
Principal Investigator in the Department of Molecular
Medicine and Biopharmaceutical Sciences, College of Med-
icine/College of Pharmacy, Seoul National University, South
Korea. Beata Chertok was the recipient of Fred W. Lyons Jr.
and Rackham Pre-Doctoral Fellowships. Adam Cole was a
recipient of a NIH Pharmacological Sciences and Bio-related
Chemistry Training Program (GM007767 from NIGMS)
grant and is currently an American Foundation of Pharma-
ceutical Education (AFPE) Pre-Doctoral Fellow.
Fig. 8. In vivo magnetic targeting of 9L glioma bearing rats. a–c Representative MR images of the targeted animals. d–f Representative MR
images of the non-targeted animals. The arrows in a–f showed the position of tumors. g GA-MNP content in excised brain tumors and normal
brain tissues of targeted (n=4) and non-targeted rats (n=3) in the ESR studies. *The plot showed a statistically significant difference comparing
the targeted tumors to the targeted normal brain tissues and non-targeted tumors and normal brain tissues (unpaired t test, p<0.001)
698Zhang et al.
1. Arruebo M, Fernandez-Pacheco R, Ibarra MR, Santamaria J.
Magnetic nanoparticles for drug delivery. Nano Today. 2007;2:22–32.
2. Chertok B, Moffat BA, David AE, Yu FQ, Bergemann C, Ross
BD, et al. Iron oxide nanoparticles as a drug delivery vehicle for
MRI monitored magnetic targeting of brain tumors. Biomaterials.
3. Wang YXJ, Hussain SM, Krestin GP. Superparamagnetic iron
oxide contrast agents: physicochemical characteristics and appli-
cations in MR imaging. Eur Radiol. 2001;11:2319–31.
4. Petri-Fink A, Steitz B, Finka A, Salaklang J, Hofmann H. Effect of
cell media on polymer coated superparamagnetic iron oxide nano-
particles (SPIONs): colloidal stability, cytotoxicity, and cellular
uptake studies. Eur J Pharm Biopharm. 2008;68:129–37.
5. Mornet S, Vasseur S, Grasset F, Duguet E. Magnetic nano-
particle design for medical diagnosis and therapy. J Mater Chem.
6. Kattumuri V, Katti K, Bhaskaran S, Boote EJ, Casteel SW, Fent
GM, et al. Gum arabic as a phytochemical construct for the
stabilization of gold nanoparticles: in vivo pharmacokinetics and
X-ray-contrast-imaging studies. Small. 2007;3:333–41.
7. Islam AM, Phillips GO, Sljivo A, Snowden MJ, Williams PA. A
review of recent developments on the regulatory, structural and
8. Jayme ML, Dunstan DE, Gee ML. Zeta potentials of gum arabic
stabilised oil in water emulsions. Food Hydrocolloid. 1999;13:459–65.
9. Kannan R, Rahing V, Cutler C, Pandrapragada R, Katti KK,
Kattumuri V, et al. Nanocompatible chemistry toward fabrication
of target-specific gold nanoparticles. J Am Chem Soc.
10. Fan XB, Tan J, Zhang GL, Zhang FB. Isolation of carbon
nanohorn assemblies and their potential for intracellular deliv-
ery. Nanotechnology. 2007;18:195103. doi:10.1088/0957-4484/18/
11. Banerjee SS, Chen DH. Multifunctional pH-sensitive magnetic
nanoparticles for simultaneous imaging, sensing and targeted
intracellular anticancer drug delivery. Nanotechnology.
12. Kim DK, Zhang Y, Voit W, Rao KV, Muhammed M. Synthesis
and characterization of surfactant-coated superparamagnetic
monodispersed iron oxide nanoparticles. J Magn Magn Mater.
13. Huang YZ, Chen JL, Chen XJ, Gao JQ, Liang WQ. PEGylated
synthetic surfactant vesicles (Niosomes): novel carriers for
oligonucleotides. J Mater Sci Mater Med. 2008;19:607–14.
14. Ross BD, Zhao YJ, Neal ER, Stegman LD, Ercolani M, Ben-
Yoseph O, et al. Contributions of cell kill and posttreatment
tumor growth rates to the repopulation of intracerebral 9L
tumors after chemotherapy: An MRI study. Proc Natl Acad Sci
U S A. 1998;95:7012–7.
15. Chertok B, David AE, Huang YZ, Yang VC. Glioma selectivity
of magnetically targeted nanoparticles: a role of abnormal tumor
hydrodynamics. J Control Release. 2007;122:315–23.
16. Hong J, Xu DM, Yu JH, Gong PJ, Ma HJ, Yao SD. Facile
synthesis of polymer-enveloped ultrasmall superparamagnetic
iron oxide for magnetic resonance imaging. Nanotechnology.
17. Sato T, Iijima T, Seki M, Inagaki N. Magnetic properties of
ultrafine ferrite particles. J Magn Magn Mater. 1987;65:252–6.
18. Willard MA, Kurihara LK, Carpenter EE, Calvin S, Harris VG.
Chemically prepared magnetic nanoparticles. Int Mater Rev.
19. Lee HY, Lee SH, Xu CJ, Xie J, Lee JH, Wu B, et al. Synthesis
and characterization of PVP-coated large core iron oxide nano-
particles as an MRI contrast agent. Nanotechnology.
20. Jun YW, Huh YM, Choi JS, Lee JH, Song HT, Kim S, et al.
Nanoscale size effect of magnetic nanocrystals and their utiliza-
tion for cancer diagnosis via magnetic resonance imaging. J Am
Chem Soc. 2005;127:5732–3.
21. Yavuz CT, Mayo JT, Yu WW, Prakash A, Falkner JC, Yean S, et al.
Low-field magnetic separation of monodisperse Fe3O4nanocrystals.
22. Corot C, Robert P, Idee JM, Port M. Recent advances in iron
oxide nanocrystal technology for medical imaging. Adv Drug
Deliver Rev. 2006;58:1471–504.
23. Weissleder R, Stark DD, Engelstad BL, Bacon BR, Compton CC,
White DL, et al. Superparamagnetic iron oxide: pharmacokinetics
and toxicity. Am J Roentgenol. 1989;152:167–73.
24. Lu CW, Hung Y, Hsiao JK, Yao M, Chung TH, Lin YS, et al.
Bifunctional magnetic silica nanoparticles for highly efficient
human stem cell labeling. Nano Lett. 2007;7:149–54.
699 Gam Arabic MNPs for Tumor Targeting and Imaging