Magnetic Resonance in Medicine 63:224–229 (2010)
Enhancement of Gas-Filled Microbubble R2* by Iron
Oxide Nanoparticles for MRI
April M. Chow,1,2 Kannie W. Y. Chan,1,2 Jerry S. Cheung,1,2 and Ed X. Wu1,2*
Gas-filled microbubbles have the potential to become a unique
intravascular MR contrast agent due to their magnetic susceptibility effect, biocompatibility, and localized manipulation via
ultrasound cavitation. However, microbubble susceptibility
effect is relatively weak when compared with other intravascular MR susceptibility contrast agents. In this study, enhancement of microbubble susceptibility effect by entrapping
monocrystalline iron oxide nanoparticles (MIONs) into polymeric microbubbles was investigated at 7 T in vitro. Apparent
T2 enhancement (DR2*) induced by microbubbles was measured
to be 79.2 6 17.5 sec21 and 301.2 6 16.8 sec21 for MION-free
and MION-entrapped polymeric microbubbles at 5% volume
fraction, respectively. DR2* and apparent transverse relaxivities
(r2*) for MION-entrapped polymeric microbubbles and MIONentrapped solid microspheres (without gas core) were also
compared, showing the synergistic effect of the gas core with
MIONs. This is the first experimental demonstration of microbubble susceptibility enhancement for MRI application. This
study indicates that gas-filled polymeric microbubble susceptibility effect can be substantially increased by incorporating
iron oxide nanoparticles into microbubble shells. With such an
approach, microbubbles can potentially be visualized with
higher sensitivity and lower concentrations by MRI. Magn
C 2009 Wiley-Liss, Inc.
Reson Med 63:224–229, 2010. V
Key words: MRI; contrast agent; microbubbles; susceptibility;
iron oxide nanoparticles
both low- and high-molecular-weight therapeutic compounds, and enhancing high-intensity focused ultrasound therapy by increasing the local heating rate (3–5).
Microbubbles can potentially be used as an intravascular MR susceptibility contrast agent in vivo due to the
induction of large local magnetic susceptibility difference by the gas-liquid interface. Moreover, microbubbles
can be locally cavitated and destroyed by focused ultrasound (2); hence, the MR signals can be temporally and
spatially manipulated because microbubble disappearances will diminish the susceptibility effect. Early experiment with AlbunexV (Molecular Biosystems Inc., San
Diego, CA), an ultrasound contrast agent consisting of
air-filled microbubbles with human albumin shell, illustrated the potential of air-filled microbubbles as a MR
susceptibility contrast agent (6). Feasibility of microbubbles as an MR pressure sensor, based on the susceptibility change caused by pressure-induced microbubble size
change, has been explored through theoretical and phantom studies (7,8). Linear relationship between apparent T2
(R2* ¼ 1/T2*) and volume fraction for OptisonV (Amersham
Health, Princeton, NJ) microbubbles of human albumin
shells with perfluorocarbon as core gas, was first reported
by our group at 7 T (9). Recently, we investigated the susceptibility effect of lipid-based microbubble SonoVueV
(Bracco Diagnostics, Milan, Italy) and air-filled custommade albumin-coated microbubbles (10). R2* dependency
on microbubble volume fraction was also reported for
LevovistV (Schering AG, Berlin, Germany), air-filled
microbubbles with palmitic acid shells, through an in
vitro phantom study at 1.5 T (11). Magnetic susceptibility enhancement induced by a gas-liquid interface
was demonstrated recently by simulations and MR
experiments using air-filled cylinders in water (12),
consolidating the feasibility of gas-filled microbubbles
as an MR susceptibility contrast agent. However,
microbubble susceptibility effect is relatively weak
when compared with other intravascular MR susceptibility contrast agents. The dosage used in the in vivo
experiments reported so far substantially exceeded
the maximum ultrasound dosage recommended for a
10-min human myocardial study (9,10).
Microbubbles are generally composed of a shell of biocompatible materials, such as proteins, lipids, or polymers, with filling gas. The microbubble shell can be stiff
(denatured proteins or polymers) or flexible (phospholipids). The shell thickness ranges from 10 nm to 200 nm,
with thinner shells typically used for protein and
lipid microbubbles, while thicker shells are used for
polymeric microbubbles (PMBs) (13). The effective
microbubble magnetic susceptibility can be manipulated
by changing the shell thickness and the magnetic susceptibility of the shell or filling gas (14). As the type of
R
R
R
Gas-filled microbubbles were originally developed as an
intravascular contrast agent in ultrasound imaging to
enhance acoustic backscattering. Recently, gas-filled
microbubbles have been employed in therapeutic applications due to their unique cavitation and sonoporation
properties (1,2). Local microbubble cavitation by spatially focused ultrasound can be applied in achieving
site-specific release of incorporated drugs or genes inside
microbubbles. Microbubble-mediated sonoporation can
dramatically increase cell permeability and intracellular
uptake, with no apparent tissue damage and toxicity.
Furthermore, the unique microbubble cavitation phenomenon has been exploited to achieve several therapeutic interventions, like sonothrombolysis, transient opening of the blood-brain barrier potentially for delivery of
1
Laboratory of Biomedical Imaging and Signal Processing, The University of
Hong Kong, Pokfulam, Hong Kong.
2
Laboratory of Biomedical Engineering, Department of Electrical and
Electronic Engineering, The University of Hong Kong, Pokfulam, Hong Kong.
Grant sponsor: Hong Kong Research Grant Council; Grant number: CERG
HKU 7642/06M.
*Correspondence to: Ed X. Wu, Ph.D., Laboratory of Biomedical Imaging
and Signal Processing, Department of Electrical and Electronic
Engineering, The University of Hong Kong, Pokfulam, Hong Kong. E-mail:
ewu@eee.hku.hk
Received 23 January 2009; revised 21 July 2009; accepted 27 July 2009.
DOI 10.1002/mrm.22184
Published online 1 December 2009 in Wiley InterScience (www.interscience.
wiley.com).
C 2009 Wiley-Liss, Inc.
V
224
R
Enhancement of Gas-filled Microbubble R*2
filling gas is chosen mainly to improve the microbubble
stability in vivo, modifying the magnetic susceptibility
of the shell is a preferred means to increase the microbubble magnetic susceptibility. Theoretical study has
indicated that, by embedding or coating magnetic nanoparticles, the magnetic susceptibility of the shell can be
increased (8,14), thus enhancing the microbubble susceptibility effect and alleviating the dosage requirement for
MRI applications.
We hypothesize that by incorporating iron oxide nanoparticles into PMBs, the magnetic susceptibility of the
shell could be substantially increased, which would lead
to a stronger microbubble magnetic susceptibility and
higher R2*. In this study, we aim to experimentally demonstrate that microbubble R2* can be enhanced by entrapping monocrystalline iron oxide nanoparticles (MIONs)
into PMB shells.
MATERIALS AND METHODS
All MRI measurements were acquired on a 7-T MRI scanner with a maximum gradient of 360 mT/m (70/16 PharmaScan; Bruker Biospin GmbH, Germany). Microbubble
phantom study was performed with 38-mm quadrature
resonator for radiofrequency transmission and receiving.
Synthesis of PMBs
MIONs were obtained from the MGH Center of Molecular
Imaging Research, MA, and used in this study without
modification (15,16). They are ultrasmall superparamagnetic iron oxide nanoparticles coated with dextran. The
particles’ hydrodynamic diameter as determined by
unimodal analysis was 29.4 nm, and with 4- to 5-nm
iron oxide cores (17).
PMBs with MIONs incorporated or entrapped in shells
were produced by adapting a double-emulsion procedure
(18) (Fig. 1). In brief, 0.5 g poly(D,L-lactide-co-glycolic
acid) 50:50 (Sigma, St. Louis, MO) was dissolved in
10 mL of ethyl acetate (Sigma). One milliliter of MION
solution (1.164 mg Fe/mL) was added to the polymer
solution and sonicated using an ultrasound probe for
30 sec. The water/oil emulsion was then added to a 5%
poly(vinyl alcohol) (Sigma) solution and homogenized for
5 min. The double (water/oil)/water emulsion was then
added into a 2% isopropyl alcohol (Sigma) solution and
stirred at room temperature for 1 h. The capsules were
collected by centrifugation, washed once with deionized
water, and centrifuged at 15 C for 5 min at 3000 g, and
the supernatant was discarded. The capsules were then
washed three times with hexane (Sigma). The capsules
were frozen in a 80 C freezer and lyophilized using a
freeze dryer to fully dry the capsules and sublime the
encapsulated water. The whole procedure took about
2 days. MION-free PMBs were synthesized with the same
procedure, except using deionized water instead of the
MION solution. Six batches of MION-free and MIONentrapped PMBs were produced for the following MRI
experiment. To demonstrate the synergistic effect of the
gas core with MIONs, two batches of MION-entrapped
solid microspheres were also prepared with the same procedures without performing lyophilization for compari-
225
FIG. 1. Flow diagram representing the adapted double-emulsion
method for synthesizing iron oxide nanoparticles entrapped in
PMBs.
son. Note that this was to compare the MION-containing
particles with (MION-entrapped PMBs) and without the
gas core (MION-entrapped solid microspheres).
Characterization of PMBs
PMB phantoms were prepared by adding 2 mL saline to
50 mg of the lyophilized powder, resulting approximately
5% microbubble volume fraction. The microbubbles were
then placed in separate 4-cm-long, 1-cm-diameter, 2-mL
cylindrical phantom tubes. Fresh microbubble vials were
used to make the phantoms. MION-entrapped solid microsphere phantoms of 5% volume fraction were also prepared using 2-mL cylindrical tubes. Each phantom tube
was slowly warmed to room temperature and gently
mixed for 2 min outside the magnet prior to MR measurements. To ensure uniform suspension of microbubbles or
microspheres, the phantom was then continuously stirred
by rotation inside the magnet. It was then arrested in horizontal position immediately prior to MRI data acquisition.
Microbubbles were observed to migrate upward due to
the buoyant force, as expected (9). Initially, there was a
uniform suspension of microbubbles. Microbubbles
started to migrate upward; therefore, in the final state
the microbubbles aggregated in the upper part of the
tube. As a result, the microbubble concentration at the
middle of the tube gradually decreased to zero. During
this process, apparent T2 enhancement (DR2*) was continuously measured by acquiring multiecho gradientecho (GE) signals without phase encoding for 2 min from
a 1-mm axial slice at middle of the phantom tube. The
measurement was repeated six times for each microbubble phantom. The parameters were pulse repetition
time ¼ 1000 ms, eight echo times ¼ 3.5 to 28 ms with
3.5-ms increment, flip angle ¼ 30 , field of view in
frequency-encoding direction ¼ 51.2 mm, acquisition
matrix ¼ 128 128, and number of excitations ¼ 1. R2*
values were computed by monoexponential fitting of the
peak magnitudes of the multi-echo GE signals. Microbubble-induced DR2* was then calculated as the difference
between the R2* in the initial state and that in the final
state as the phantom changed from the uniform microbubble suspension to a microbubble-free state. For
MION-entrapped solid microspheres, they were observed
to settle to the bottom. Similar to PMBs, during the
226
Chow et al.
settling process of MION-entrapped solid microspheres,
DR2* was continuously measured by acquiring multiecho
GE signals without phase encoding for 2 min from a 1mm axial slice at the middle of the phantom tube, using
the same sequence described above. Note that the microspheres at the middle of the tube gradually changed
from a well-suspended state to microsphere-free one during this process.
Using the same microbubble phantoms, R2* maps of
the microbubble-free suspending solutions (i.e., after the
completion of microbubble flotation) were acquired with
multiecho GE imaging sequences. The parameters were
pulse repetition time ¼ 1000 ms, eight echo times ¼ 5 to
40 ms with 5-ms increment, flip angle ¼ 30 , field of
view ¼ 51.2 mm 51.2 mm, slice thickness ¼ 1 mm, acquisition matrix ¼ 128 128, spatial resolution ¼ 0.4 0.4 1 mm3 and number of excitations ¼ 1. To confirm
the entrapment of MIONs inside PMBs, cavitation was
performed to destroy the microbubbles by placing the
phantoms in a sonicator bath with ultrasound of frequency 40 kHz (BransonicV 1510; Branson Ultrasonic,
Danbury, CT). After the microbubble cavitation, their R2*
maps of the suspending solutions were acquired again,
using an identical protocol. Changes in the R2* values of
the MION-free and MION-entrapped PMB suspending
solutions after cavitation were determined. In addition,
MION apparent transverse relaxivity (r2*) was determined. In brief, randomly uniform suspensions were
prepared for MION concentrations of 0, 10.0, 20.0, and
30.0 lg Fe/mL with the addition of saline. They were
then placed in separate 2-mL cylindrical phantom tubes
for R2* mapping, using identical protocol.
To compare the effect of MIONs in MION-entrapped
PMBs and solid microspheres (two batches each from
the MR experiments), their absolute iron contents were
determined with inductively coupled plasma mass
spectrometry using Agilent 7500a inductively coupled
plasma mass spectrometry (Agilent Technologies,
Santa Clara, CA). Particles were incubated in concentrated nitric acid at 80 C for 4 h and then placed in
3% nitric acid with 5 parts per billion as internal
standard. The r2* relaxivities of these two types of particles were computed as the R2* in the initial uniform
suspended state, normalized with their respective iron
contents.
R
RESULTS
Characterization of Microbubble Suspensions
The light micrographs of a representative batch of
MION-entrapped PMBs are depicted in Fig. 2a. The estimated size distribution was from 1 to 25 lm, with mean
diameter 9.77 lm for MION-entrapped PMBs, as estimated from light microscopic analysis (Fig. 2b). Similar
size distributions were observed for MION-free PMBs
and MION-entrapped solid microspheres.
Measurement of DR2* of MION-free and MION-entrapped
PMB Suspensions
Figure 3a shows the typical multiecho GE signals from a
MION-free PMB suspension phantom at 5% volume frac-
FIG. 2. a: Representative light microscopy of MION-entrapped
PMBs. b: Histogram showing diameter distribution for a representative batch of MION-entrapped PMBs. The mean diameter of
MION-entrapped PMBs is 9.77 lm.
tion in its initial well-suspended state. Figure 3b shows
the multiecho GE signals in the suspension’s final microbubble-free state. R2* of MION-free PMBs was plotted
against time in Fig. 3c. Figure 3d-f is the corresponding
figure for a MION-entrapped PMB suspension phantom
at 5% volume fraction. DR2* induced by microbubbles
was calculated as the R2* decrease between the initial
and final state. Note that when the lyophilized powder
was suspended in saline, initial bursting of PMBs could
occur (19). For MION-entrapped PMBs, such bursting
caused some MIONs to dissolve into the suspending solution. This accounted for the difference in the R2* in the
final microbubble-free state for MION-free and MIONentrapped PMBs in Fig. 3c and f. Figure 4 shows the
individual microbubble-induced DR2* values in the six
batches of MION-free and MION-entrapped PMBs at 5%
volume fraction, each measured with six repeated measurements. The average DR2* was 79.2 17.5 sec1 and
301.2 16.8 sec1 for MION-free and MION-entrapped
PMBs (mean standard error), respectively.
The R2* values of the suspending solutions (i.e.,
microbubble-free suspension after the completion of
Enhancement of Gas-filled Microbubble R*2
227
FIG. 3. The typical multiecho GE signals from a MION-free PMB phantom (5% volume fraction) in (a) its initial well-suspended state
and (b) its final microbubble-free state with the monoexponential fitting in dotted lines. c: R2* vs time. The multiecho GE signals from a
MION-entrapped PMB phantom (5% volume fraction) in (d) its initial well-suspended state and (e) its final microbubble-free state. f: R2*
vs time.
microbubble upward flotation) measured before and after
ultrasound cavitation are summarized in Table 1. Two-tail
paired Student’s t test showed no significant difference in
R2* before and after cavitation for MION-free PMBs. However, suspending solution R2* increased substantially after
cavitation for MION-entrapped PMBs (P < 0.01). This
indicated that MIONs in MION-entrapped PMBs were
released into the suspending solutions after the microbubble destruction by ultrasound cavitation.
For MION-entrapped solid microspheres, the average
DR2* was 82.5 3.6 sec1 (mean standard error).
According to the inductively coupled plasma mass spectrometry measurements, the absolute iron contents of
MION-entrapped PMBs and MION-entrapped solid
microspheres at 5% volume fractions were 28 lg Fe/mL
and 29 lg Fe/mL respectively. Their mean r2* relaxivities
were then estimated as 21.0 sec1/(lg Fe mL1) and 7.2
sec1/(lg Fe mL1). As for free MIONs in saline suspension, DR2* was observed to increase linearly with MION
concentration as expected. Their r2* relaxivity was estimated to be 2.5 sec1/(lg Fe mL1) (422.2 mM1 s1).
DISCUSSION
In this study, entrapping iron oxide nanoparticles in
PMBs was demonstrated for the first time to enhance the
FIG. 4. Individual DR2* measurements in six batches of MION-free
and six batches of MION-entrapped PMB suspensions (5% volume fraction). The error bars represent standard deviation.
228
Chow et al.
Table 1
Measurements of the Suspending (i.e., Microbubble-Free) Solution R2* Values for MION-Free and MION-Entrapped PMBs Before and
After Ultrasound Cavitation (Mean Standard Error, Number of Batches ¼ 6)
Microbubble-free suspending solutions R2* (sec1), mean SE
MION-free PMBs
MION-entrapped PMBs
Before cavitation
After cavitation
Change
2.9 0.2
38.2 6.2
3.2 0.3
92.7 10.1
0.3 0.4 (ns)
54.5 9.5 (**)
Two-tail paired Student’s t test was performed with ** for P < 0.01 and ns for insignificance.
microbubble MR susceptibility effect. Microbubbleinduced DR2* of MION-entrapped PMBs was found to be
significantly higher than that of MION-free PMBs. With
the increasing availability of high-field MRI systems in
both clinical and research settings, microbubbles offer
the promise as a viable and unique contrast agent since
their MR susceptibility effect can be substantially
enhanced through this approach.
Microbubble-induced DR2* depends on the microbubbles’ radius, volume fraction, overall magnetic susceptibility difference between the microbubble and the blood
plasma (Dv), and amplitude of the static field (7,8). Given
the possible microbubble toxicity when the dose exceeds
1 mL/kg (14) and the filtering of microbubbles larger
than 10 lm by the lung capillary bed (9), microbubble
volume fraction and radius cannot be freely chosen.
Therefore, Dv enhancement is a preferred way to increase
microbubble-induced DR2*. In theory, Dv is determined
by (14)
"
#
9ð1 þ vs ÞDvgs b3
3
Dvsp þ
Dv ¼
3 þ vs
ð2vs þ 3Þðvs þ 3Þ þ 2v2s b3
[1]
where Dvsp ¼ (vs vp)/l0, Dvgs ¼ (vg vs)/l0, l0 is the
permeability of free space. b Denotes the ratio of inner
radius of the microbubble to outer radius. vp, vs And vg
are the magnetic susceptibility of suspending solution,
shell, and filling gas, respectively. By embedding or coating magnetic nanoparticles, vs and therefore Dv can be
enhanced (14), as demonstrated experimentally in the
current study. Note that vs increases monotonically with
the radius, density, and magnetic susceptibility of the
nanoparticles embedded in the shell (14).
Microbubble-induced DR2* for MION-entrapped PMBs
(301.2 16.8 sec1) was significantly higher than that of
MION-free PMBs (79.2 17.5 sec1), demonstrating that
microbubble susceptibility can be enhanced by the
MIONs entrapped in shells, as suggested by the earlier
theoretical study (14). Furthermore, the r2* relaxivity of
MION-entrapped PMBs (21.0 sec1/(lg Fe mL1)) was
observed to be much higher than that of MIONentrapped solid microspheres (7.2 sec1/(lg Fe mL1))
and free MION suspension (2.5 sec1/(lg Fe mL1)).
These relaxivity results underscore the importance of
both gas core and entrapped MIONs in achieving strong
overall microbubble susceptibility effect. Note that the
higher r2* relaxivity of MION-entrapped solid microspheres as compared to free MIONs is expected because,
for a given concentration, the susceptibility effect of
MIONs is stronger when they are compartmentalized.
Microbubble-induced DR2* can be manipulated by
ultrasound through microbubble cavitation. In this study,
the microbubble susceptibility effect of MION-free PMBs
decreased due to the microbubble destruction by ultrasound cavitation. In contrast, for MION-entrapped PMBs,
the MIONs released and the microbubble fragments
formed after microbubble destruction still exhibited susceptibility effect. A preliminary experiment was performed in three out of the six MION-entrapped PMB
microbubble phantoms. The R2* measured in the initial
well-suspended state decreased from 451.3 64.5 sec1
to 257.4 20.4 sec1 after cavitation. However, the R2*
reduction (193.9 sec1) here was still much higher than
the MION-free PMB DR2* (79.2 sec1), suggesting that
microbubble cavitation can be detected with increased
sensitivity when using the MION-entrapped PMBs proposed in this study.
As shown in Table 1, an R2* increase of 54.5 9.5
sec1 in the microbubble-free suspending solutions for
MION-entrapped PMBs was observed after cavitation.
Given the r2* of 2.5 sec1/(lg Fe mL1) determined in
this study for free MIONs in saline, the amount of
MIONs released into the suspending solution after
microbubble cavitation was estimated to be 21.8 lg Fe/
mL. Taking into account the microbubble volume fraction (5%) and approximate microbubble density (1.16 108 microbubbles/mL), the MION loading was estimated
to be 0.19 pg Fe per MION-entrapped PMB, assuming
that all released MIONs were free and uniformly distributed in suspending solution after cavitation. On the
other hand, from the inductively coupled plasma mass
spectrometry measurements, the MION loading was estimated to be 0.25 pg Fe per MION-entrapped PMB. Note
that the above MION loading estimated from MR measurements was lower likely because the MIONs in the
polymeric fragments after cavitation might not be free;
they could be still bound with shell fragments and settle
to bottom of the phantom tube.
Using the microbubble fabrication procedure demonstrated in this study, drugs or genes may also be incorporated into PMBs (18,19) and delivered to specific sites
under MRI monitoring. Furthermore, gadolinium chelates may be encapsulated into PMBs using this fabrication procedure. In this case, the magnetic susceptibility
of intact microbubbles could be enhanced, decreasing
the signal intensity in T2*-weighted imaging. Upon cavitation, however, the entrapped gadolinium chelates
could be released and in contact with surrounding water
molecules, thus shortening the T1 relaxation time and
producing increased signal in T1-weighted imaging.
Such dark-to-bright change can be a potentially attractive
Enhancement of Gas-filled Microbubble R*2
way to monitor microbubble-based drug delivery and
therapeutic applications.
In this study, only PMBs were investigated for DR2*
enhancement by incorporating MIONs. In fact, a slightly
modified synthesizing procedure to embed MIONs into
albumin-coated microbubbles has been explored in a
pilot experiment in our laboratory. The microbubble susceptibility effect was comparably enhanced (data not
shown). Note that similar procedure of entrapping iron
oxide nanoparticles may be also applicable in enhancing
lipid microbubble susceptibility effect.
In general, gas-filled microbubbles have relatively
short in vivo lifetime ( several minutes), primarily as a
result of their destruction in the alveoli due to gaseous
exchange (13). Such limited lifetime can be a challenge
in various microbubble applications. Nevertheless, the
microbubble fabrication technology is advancing and
microbubble stability could be improved substantially,
for example, by adding surfactant molecules, using the
multiphase mixing technique (20). Moreover, the large
size of the PMBs synthesized may produce adverse
effects in vivo. Optimization of size distribution for passage through the lung capillary bed may be sought in
future in vivo studies.
CONCLUSIONS
In this study, we experimentally demonstrated for the first
time that gas-filled PMB susceptibility effect can be substantially enhanced by incorporating iron oxide nanoparticles into microbubble shells. With such an approach,
microbubbles can be monitored by MRI with higher sensitivity or lower concentrations, which may lead to the practical use of microbubbles as an intravascular MRI contrast
agent and MRI guidance in various microbubble-based
drug delivery and therapeutic applications.
ACKNOWLEDGMENTS
We thank Prof. Danny Chan in the Department of Biochemistry and Prof. W. T. Wong in the Department of
Chemistry of the University of Hong Kong for technical
assistance.
REFERENCES
1. Ferrara KW. Driving delivery vehicles with ultrasound. Adv Drug
Deliv Rev 2008;60:1097–1102.
229
2. Unger EC, Hersh E, Vannan M, Matsunaga TO, McCreery T. Local
drug and gene delivery through microbubbles. Prog Cardiovasc Dis
2001;44:45–54.
3. Culp WC, Porter TR, Lowery J, Xie F, Roberson PK, Marky L. Intracranial clot lysis with intravenous microbubbles and transcranial
ultrasound in swine. Stroke 2004;35:2407–2411.
4. Hynynen K, McDannold N, Sheikov NA, Jolesz FA, Vykhodtseva N.
Local and reversible blood-brain barrier disruption by noninvasive
focused ultrasound at frequencies suitable for trans-skull sonications.
Neuroimage 2005;24:12–20.
5. Kaneko Y, Maruyama T, Takegami K, Watanabe T, Mitsui H, Hanajiri
K, Nagawa H, Matsumoto Y. Use of a microbubble agent to increase
the effects of high intensity focused ultrasound on liver tissue. Eur
Radiol 2005;15:1415–1420.
6. Moseley ME, Wendland MF, Rampil I, Barnhart J. Microbubbles: a
novel MR susceptibility contrast agent. In: Proceedings of the 10th
Annual Meeting of the ISMRM, San Francisco, CA, 1991. p 1020.
7. Alexander AL, McCreery TT, Barrette TR, Gmitro AF, Unger EC.
Microbubbles as novel pressure-sensitive MR contrast agents. Magn
Reson Med 1996;35:801–806.
8. Dharmakumar R, Plewes DB, Wright GA. On the parameters affecting
the sensitivity of MR measures of pressure with microbubbles. Magn
Reson Med 2002;47:264–273.
9. Wong KK, Huang I, Kim YR, Tang H, Yang ES, Kwong KK, Wu EX.
In vivo study of microbubbles as an MR susceptibility contrast agent.
Magn Reson Med 2004;52:445–452.
10. Cheung JS, Chow AM, Guo H, Wu EX. Microbubbles as a novel contrast agent for brain MRI. Neuroimage 2009;46:658–664.
11. Ueguchi T, Tanaka Y, Hamada S, Kawamoto R, Ogata Y, Matsumoto
M, Nakamura H, Johkoh T. Air microbubbles as MR susceptibility
contrast agent at 1.5 tesla. Magn Reson Med Sci 2006;5:147–150.
12. De Guio F, Benoit-Cattin H, Davenel A. Signal decay due to susceptibility-induced intravoxel dephasing on multiple air-filled cylinders: MRI simulations and experiments. MAGMA 2008;21:
261–271.
13. Quaia E. Microbubble ultrasound contrast agents: an update. Eur
Radiol 2007;17:1995–2008.
14. Dharmakumar R, Plewes DB, Wright GA. A novel microbubble construct for intracardiac or intravascular MR manometry: a theoretical
study. Phys Med Biol 2005;50:4745–4762.
15. Wu EX, Wong KK, Andrassy M, Tang H. High-resolution in vivo
CBV mapping with MRI in wild-type mice. Magn Reson Med 2003;
49:765–770.
16. Wu EX, Tang H, Jensen JH. Applications of ultrasmall superparamagnetic iron oxide contrast agents in the MR study of animal models.
NMR Biomed 2004;17:478–483.
17. Shen T, Weissleder R, Papisov M, Bogdanov A Jr, Brady TJ. Monocrystalline iron oxide nanocompounds (MION): physicochemical
properties. Magn Reson Med 1993;29:599–604.
18. El-Sherif DM, Wheatley MA. Development of a novel method for synthesis of a polymeric ultrasound contrast agent. J Biomed Mater Res
A 2003;66:347–355.
19. Patil Y, Panyam J. Polymeric nanoparticles for siRNA delivery and
gene silencing. Int J Pharm 2009;367:195–203.
20. Dressaire E, Bee R, Bell DC, Lips A, Stone HA. Interfacial polygonal nanopatterning of stable microbubbles. Science 2008;320:
1198–1201.