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SPECIAL ISSUE - REVIEW
The use of microbubbles in Doppler ultrasound studies
Piero Tortoli Æ Francesco Guidi Æ Riccardo Mori ÆHendrik J. Vos
Received: 7 March 2008 / Accepted: 28 October 2008 / Published online: 11 November 2008
� International Federation for Medical and Biological Engineering 2008
Abstract Ultrasound contrast agents (UCAs) are widely
used in Doppler studies, either for simple echo enhance-
ment purposes, or to increase the low signal-to-clutter ratio
typical of microcirculation investigations. Common to all
Doppler techniques, which are briefly reviewed in this
paper, is the basic assumption that possible phase and
amplitude changes in received echoes are only associated
with UCA microbubble movements due to the drag force of
blood. Actually, when UCAs are insonified, phenomena
such as rupture, displacement due to radiation force, and
acoustically driven deflation might influence the results of
Doppler investigations. In this paper, we investigate the
possible Doppler effects of such phenomena by means of a
numerical simulation model and a special acousto-optical
set-up which allows analysis of the behavior of individual
microbubbles over relatively long time intervals. It is thus
found that all phenomena produce evident Doppler effects
in vitro, but that bubble displacement and deflation in
particular, are not expected to significantly interfere with
clinical measurements in standard conditions.
Keywords Ultrasound contrast agents � Doppler �Primary radiation force � Microbubble rupture �Microbubble deflation
1 Introduction
Medical Doppler ultrasound (US) studies are traditionally
committed to characterize blood flow in arteries and veins,
but emerging applications also look upon the assessment of
myocardial and cerebral perfusion, as well as the detection
of plaque neovascularization at the carotid artery level.
Such studies are typically based on the measurement of
phase changes involved by blood movement in backscat-
tered echoes. Because of these phase changes, in fact,
sampling of US echoes in subsequent pulse repetition
intervals (PRI) yields a time-varying ‘‘Doppler’’ signal,
whose frequency results proportional to blood velocity
[30]. In some cases, the measurements are made difficult
by the low energy of echoes backscattered by red blood
cells (RBC), yielding poor signal-to-noise ratio. This hap-
pens more frequently in transcranial Doppler examinations,
in the analysis of deep vessels, and in coronary flow reserve
evaluation obtained through the trans-thoracic approach.
Further difficulties may be due to the movements of tissue
surrounding blood vessels, which produce an undesired
Doppler signal, typically called clutter. The clutter may be
higher, both in amplitude and frequency, than the signal
produced by slowly moving blood, as is typical in the
microcirculation.
Such limitations can be overcome by intravenously
injecting ultrasound contrast agents (UCAs) consisting of
micron-sized bubbles containing low-soluble gas and
phospholipids, sugars or polymer coating [54]. The large
difference between the acoustic impedance of plasma and
the encapsulated gas generates much stronger echoes than
those produced by RBCs [2, 20], especially for micro-
bubbles having a diameter close to resonance [54].
Impressive effects in spectral and color Doppler have been
reported after several animal experiments [35] and human
P. Tortoli (&) � F. Guidi � R. Mori
Electronic and Telecommunications Department,
University of Florence, Florence, Italy
e-mail: piero.tortoli@unifi.it
URL: http://orione.det.unifi.it
H. J. Vos
Biomedical Engineering, Thorax Center,
Erasmus Medical Center,
Rotterdam, The Netherlands
123
Med Biol Eng Comput (2009) 47:827–838
DOI 10.1007/s11517-008-0423-y
studies concerning, e.g., transcranial vessels [33] mitral
valve regurgitation, aortic stenosis and pulmonary venous
flow [53].
The key issue of basic contrast enhanced Doppler
methods is that they do not need any extra processing, and
standard US instruments can be used. The main drawback
is represented by color saturation and blooming [36], which
can be limited only by keeping the transmission power and
the receive gain as low as possible. On the other hand,
major benefits of UCA use in Doppler investigations have
been obtained by working at higher Mechanical Index
(MI), because this yields nonlinear microbubble oscillation
with rich harmonic (and subharmonic) content of corre-
sponding backscattered echoes. For example, by
transmitting at one frequency, and selectively receiving at
twice that frequency, an increased contrast between flow
and surrounding tissues has been obtained. This principle is
exploited in harmonic (spectral/color/power) Doppler
methods [13, 14], which have been successfully applied to
coronary arteries and myocardium [5, 43] and parenchyma
of abdominal organs [45].
This approach, although forcing the US system to be
equipped with wideband transducer and receiver, is now
made commercially available by a number of companies
because of its exceptional performance in terms of signal-
to-clutter ratio and, consequently, of its ability to detect
even the smallest vessels [14].
The inherent limitation of harmonic Doppler methods is
that severe bandwidth restrictions are imposed by the need
to avoid any overlap between the transmitter and receiver
bands. Optimal contrast is thus obtained at the expense of a
degraded resolution.
This compromise is overcome by pulse inversion
methods [40], in which multiple pulses of alternating
polarity are subsequently transmitted over consecutive
PRIs. By properly combining the corresponding echoes,
only the nonlinear components associated to microbubble
backscattering are enhanced. A high contrast is thus
obtained by using the full transducer bandwidth (i.e.,
with high resolution), although the useful Doppler
bandwidth is reduced by a factor of 2 [40]. This
approach has been exploited for real-time perfusion
imaging of the myocardium [63] and in combination with
special radiofrequency and Doppler filtering, in order to
differentiate low-velocity microbubbles associated with
perfusion from the higher-velocity microbubbles in larger
vessels [12]. Further contrast enhancements can be
obtained through the generalization of the pulse inversion
method [11], i.e., by applying phase shifts other than half
a period to the transmitted pulses of a color Doppler
sequence, or by means of contrast specific sequences
which use a combination of changes in pulse phase and
amplitude [56].
It can be observed that in many applications UCAs are
used mainly in detecting blood movements, similar to
power Doppler, more than in estimating its velocity. This is
partially true in the so-called intermittent imaging mode, in
which bubbles are first disrupted [44] with an US burst of
high amplitude, and the changing level of energy in sub-
sequent power Doppler images is tracked to estimate the
velocity at which fresh microbubbles refill the microves-
sels. There are also emerging applications like the
detection of carotid plaque neovascularization [31, 34]
which exploit UCAs for vasa vasorum imaging. However,
since blood is only a vehicle in such applications, and there
is no direct interest in detecting its movement, strictly
speaking, these cannot be considered Doppler methods.
Common to all the above Doppler techniques are the
basic assumptions that possible phase and amplitude
changes in received echoes are due only to the bubble
displacements, and that these displacements are the same
as those of RBCs. Actually, when microbubbles are in-
sonified, there can be phenomena other than the non-linear
oscillations [61], which might influence the results of
Doppler investigations. In this paper, we consider in detail
three such phenomena: displacement due to radiation force,
rupture and acoustically driven deflation.
2 UCA-driven ambiguous Doppler effects
Displacement of microbubbles due to primary radiation
(also termed ‘‘Bjerknes’’) force [4], is a well-known phe-
nomenon [24, 28]. In general terms, radiation force induces
a roughly additive bubble displacement in the US wave
propagation direction [19, 65], which mainly depends on
the bubble composition and US intensity and frequency.
However, this effect is also influenced by other factors such
as the fluid drag and the possible proximity of other bub-
bles, or of a vessel wall. A microbubble driven near its
resonance frequency, in particular, experiences a large net
radiation force and can be appreciably displaced; the phase
of backscattered echoes is correspondingly affected by
such an extra movement.
The rapid rupture of a phospholipids coated microbub-
ble typically occurs when, because of high-intensity US
excitation, its instantaneous radius increases to an exces-
sive degree [15]. The accumulated energy is higher than
that dissipated, and subsequent oscillation becomes unsta-
ble. This phenomenon is influenced by the amplitude and
central frequency of the driving signal as well as by bubble
size and shell composition. The bubble may split into
smaller bubbles (‘‘fragmentation’’), which still maintain the
coating, and in some cases coalesce again [6, 8, 16, 18, 57].
Instantaneous rupture of so-called hard shelled bubbles
occurs when the gas core escapes through a shell defect
828 Med Biol Eng Comput (2009) 47:827–838
123
(‘‘sonic cracking’’) [6, 9, 18, 25]. Whichever the rupture
mechanism, it produces strong decorrelation between
subsequent pulse-echoes [16].
The deflation (or dissolution) of a gas bubble is a rela-
tively slow process through which the gas diffuses from the
core to the surrounding liquid, causing a volume reduction.
The phenomenon can happen naturally (passive dissolu-
tion), due to the overpressure inside the bubble coming
from the surface tension, or can be accelerated by the wall
motion induced by an external driving force (acoustically
driven deflation) [32, 38]. In general, the loss of gas is
accompanied by a deviation from the spherical shape or by
the shedding of the shell material [7, 58, 59]. The gradual
bubble volume reduction and the change of mechanical
properties correspondingly influence both amplitude and
phase of backscattered echoes [16].
All the aforementioned phenomena thus determine a
change of amplitude and/or phase in the backscattered
echoes, which might be interpreted as Doppler effects. In
the next sections, for each phenomenon an experimental
procedure is described and shown capable of emphasizing
the related Doppler effects in individual microbubbles.
The amplitude of these effects and their possible influence
on clinical investigations are discussed in the final
section.
2.1 Doppler effects due to radiation force
Radiation force is caused by the pressure gradient acting
upon the vibrating bubble surface during the passing of an
US wave. The magnitude of the corresponding bubble
displacement is related to the parameters of the transmitted
US beam, the physical properties of the bubbles, and the
characteristics of the fluid in which they are suspended. In
order to investigate the displacement phenomenon, a suit-
able set-up has been realized in which the bubbles are
insonified while freely floating in a still fluid.
2.1.1 Experimental set-up
The experimental set-up, shown in Fig. 1a [69], is based on
a water tank in which small concentrations of microbubbles
are suspended. All experiments were made in water at
room temperature. A single-element US probe insonifies
the bubble suspension, under control of the Bubble
Behavior Testing (BBT) system, i.e., an electronic board
which can perform standard and unconventional transmit
and receive functions.
Two different types of bubbles were used: the experi-
mental BR14 (Bracco Research SA, Geneva, Switzerland)
UCA, containing perfluorobutane in a phospholipids shell
which was used at a 1:150.000 dilution, and thermoplastic
F-04E microspheres (Matsumoto Yusi-Seiyaku Co. Ltd,
Osaka, Japan) containing hydrocarbon gas (C3H8 and
C4H10) in a polymer shell [62], employed at a concen-
tration of 10 lg/liter.
The BBT board was connected to two different single-
element focussed US transducers. The first was a 3 MHz
centre frequency, 70% fractional bandwidth transducer
with 75 mm focal depth (Vermon M3, Tours, France), and
the other was a 6 MHz, 75% fractional bandwidth trans-
ducer with 25 mm focal depth (Imasonic, Voray sur
l’Ognon, France).
Pulses of programmable frequency, length and pressure
were produced by the BBT system to generate US fields
with different characteristics. The combination of a sensi-
tive transducer and a low noise receiver allowed a high
signal-to-noise ratio to be obtained even from single bub-
bles. The received echo-signals were digitized through 14-
bit resolution A/D converters, and then coherently quad-
rature-demodulated, low pass filtered and stored in a
circular buffer memory. For each transmitted pulse, 128
complex signal samples were taken to cover a suitable
range around the transducer focus. The raw samples were
available for storage in a PC file, allowing the possibility of
post-processing.
A digital processing unit provided the needed compu-
tational power to perform real-time elaboration and data
dispatching to a PC for presentation purposes, through a
USB 2.0 port. Two different displays are available: the
classic M-mode display, showing the time evolution of
received signal power, and the Multigate Spectral Doppler
(MSD) display, which shows the instantaneous distribution
of all Doppler spectra produced over the investigated range
depths.
Before each experimental test, the microbubble sus-
pension was prepared by mixing distilled water with
bubbles, then keeping the solution still for approximately
5 min.
2.1.2 Simulation model
A study was performed in order to compare experimental
bubble displacements to trajectories predicted by a
numerical simulation model. The model used in this sec-
tion accounts for the dominant forces that simultaneously
act on the bubble through Newton’s Second Law, and
estimates the bubble displacement after an US pulse.
The primary radiation force acting on a freely moving
and oscillating, highly compressible bubble, smaller than
the wave length [4], is well modeled by the following
equation
F~US ¼ �V tð Þr~p z; tð Þ
(see Table 1 for a description of the symbols), in which the
instantaneous volume, V(t), is estimated on the basis of the
Med Biol Eng Comput (2009) 47:827–838 829
123
radial excursion predicted by a Herring type differential
equation (see below).
A second major contribution comes from the viscosity
of the surrounding medium and can be modeled by a quasi-
static drag force [22, 66, 71] as
F~D ¼ �1
4pCDRlu~rRe;
where ur is the velocity of the bubble relative to the fluid
and Re the Reynolds number. In this equation the terms
related to the drag coefficient Cd were empirically esti-
mated by different researchers [23, 41, 52, 70] and we
assumed final values as reported in [69].
A third contribution comes from the ‘‘added mass’’
force, the inertia due to the bubble surrounding fluid, which
has been widely derived in the literature [1, 10, 26, 46, 47,
49, 51, 52, 55],
F~AM ¼ �1
2q V
ou~
otþ u~r
dV
dt
� �
Contributions of gravity, buoyancy, and additional
added mass [52], are neglected as they are assumed to be
significantly less than the radiation force and fluid drag.
The bubble is modeled as an encapsulated gaseous
sphere immersed in a viscous and compressible fluid,
accounting for the acoustical approximation of limited US
velocity.
The shell is assumed to be solid, incompressible and
viscoelastic accordingly to the Kelvin–Voigt model [17,
Fig. 1 Complete experimental
set-up including the BBT
system combined with: (a) the
phantom used to test the bubble
displacement and rupture; (b)
the phantom linked to the
optical system used in deflation
studies. BBT is Bubble
Behavior Testing system, TX/
RX is transmitting/receiving
system, Synch is
synchronization signal
Table 1 Variables and parameter values as used in the simulation
Symbol Name Unit and value
F Force N
V Bubble volume m3
p Acoustic pressure Pa
p0 Ambient pressure 101 kPa
z Direction of US propagation
and displacement
–
Re Reynolds number –
R Bubble radius m
l Fluid dynamic viscosity 0.9 9 10-3 Pas (water)
4 9 10-3 Pas (blood)
ur Bubble relative velocity ms-1
CD Drag coefficient –
q Density of medium 998 kg 9 m-3 (water)
1,060 kg 9 m-3 (blood)
c Speed of sound 1,480 m/s
c Polytropic gas exponent –
Gs Shell shear modulus 18 MPa (F04-E)
32 MPa (BR14)
ds Shell thickness 2.7% of R0 (F04-E)
4 nm (BR14)
lsh Shell viscosity 0.23 Pas (F04E)
0.19 Pas (BR14)
dth Thermal damping –
f Driving frequency Hz
Blood values from [27]. Shell and gas values from [69]
830 Med Biol Eng Comput (2009) 47:827–838
123
39]. The shell mass is conserved during vibration, and the
thickness is assumed much smaller than the radius R [69].
Bubble volume-time oscillation is evaluated considering
spherical oscillation and a radius evolution predicted by a
model based on a modified Herring equation [68], extended
with terms describing the influence of bubble shell [17, 39].
q R €Rþ 3
2_R2
� �¼ p0
R0
R
� �3c
1� 3cc
_R
� �
� 4l _R
R� 12Gsds
R0
R
� �31
R0
� 1
R
� �
� 12lshds
R0
R
� �2 _R
R2� 2pf dthqR _R� ðp0 þ pÞ
The thermodynamic system follows the polytropic
relation pVc = constant with c calculated according to
Hoff [37]. The thermal damping is added in a linear form
assuming no extra contribution from the shell. For
simplicity, surface tension of both the polymer and
phospholipids coating is set to zero [39, 50].
2.1.3 Results
When a bubble population is insonified, each bubble pro-
duces an echo pulse whose amplitude is converted, through
the M-mode display palette, into a light spot. During sub-
sequent PRIs, the light-spot amplitude and position changes
according to the bubble depth, producing a light-trace.
In Fig. 2, different traces show the instantaneous posi-
tions of individual bubbles in time. Each trace clearly
shows that the corresponding bubble moves away from the
transducer surface. The instantaneous velocity is propor-
tional to the local trace slope and reaches peak values when
the bubble is in the transducer focal zone. The steepest and
brightest traces can be related to bubbles with a diameter
closer to resonance size. This is consistent with the reso-
nance dependency of the US radiation force and the higher
scattering cross-section of a resonating bubble.
The same phenomenon can be observed through the
MSD-display, typically used in hemodynamic studies to
detect the blood velocity profiles within human arteries.
For each investigated depth, this display shows the Doppler
spectra obtained from the echoes of all bubbles located at
that depth.
Figure 3 shows an example of MSD display obtained
when a few bubbles were intercepted along the beam axis.
Each bubble is represented as a light-spot, whose bright-
ness corresponds to the related Doppler power, the vertical
position corresponds to the bubble position, and the hori-
zontal position corresponds to the bubble mean Doppler
frequency. Due to the specific set-up, the bubble dis-
placement results roughly parallel to the beam axis, and the
mean frequency can be directly converted to mean velocity
through the Doppler equation with zero angle. In this
example, obtained by insonifying BR14 microbubbles with
2 MHz 4-cycle pulses repeated at 950 Hz and a 500 kPa of
peak negative pressure (PNP), a maximum Doppler shift of
20 Hz was measured corresponding to an average velocity
of about 7.5 mm/s. However, it should be kept in mind that
each bubble only actually moves during the US excitation,
while remaining more or less still during the remaining part
of the PRI. Considering the selected pulse repetition fre-
quency (PRF) and burst length, while neglecting transitory
effects, a peak velocity of about 4 m/s, reached during
excitation, can be estimated for this bubble.
An accurate estimate of the peak velocity yielded from
the radiation force can be obtained by integrating the light-
spot in the MSD display over a long time. This integration
involves a population of bubbles, surely including resonant
bubbles aligned at the US beam axis. The latter are
expected to experience the maximum radiation force, and
thus the maximum peak velocity, while all the other bub-
bles, off-axis or not resonating, move slower.
Figure 4a shows an integral MSD display obtained after
exciting polymer F-04E microspheres for 8 s. The super-
imposed trace corresponds to the mean velocities predicted
by our model for a resonant F-04E micro-sphere that
moves along the beam axis.
As expected, at each depth, the experimental measured
velocities are actually distributed between zero and the
Fig. 2 M-mode display of
thermoplastic microspheres
insonified by the Imasonic
transducer with 8-cycle 8 MHz
pulses at PRF = 5 kHz
(1.5 MPa PNP)
Med Biol Eng Comput (2009) 47:827–838 831
123
maximum value predicted by the model. Around the
transducer focal depth (%80 mm) the mean velocity of
resonant bubbles is about 12 mm/s.
The agreement between the simulated and experimental
results confirm the validity of our model [69]. This has
encouraged us to estimate the peak velocity that resonating
bubbles can achieve during US excitation, directly through
the model. Figure 4b shows the peak velocities estimated for
F-04E and BR14 bubbles excited with 2.5 MHz 10-cycle
pulses within a range of pressures. It is clearly shown that
instantaneous velocities in the range of meters/s are achieved
even at pressures of a few hundred kPa. We expect that the
model will start to fail for BR14 at elevated acoustic pres-
sures, when disruption and/or deflation dominate the bubble
dynamics (see next sections). Therefore, the displacement of
BR14 at pressures over 300 kPa (MI 0.2) is only an indica-
tion of the velocity that an ‘‘undamaged’’ bubble could reach.
2.2 Doppler effects due to bubble rupture
2.2.1 Methods
The experimental set-up described in the previous section
was also used to observe the Doppler effects associated
with the presumed destruction of single microbubbles.
Driving parameters (central frequency, number of
cycles, pressure amplitude) were set in such a way to create
conditions which may determine rupture events. When
pushed by the primary radiation force, bubbles are trans-
ferred to deeper regions of the acoustic field. They
experience pressure that could induce rupture phenomena
especially (but not exclusively) in the focal region.
2.2.2 Results
A typical ‘‘event’’ is shown in Fig. 5a through the M-mode
display. The white trace corresponds to the bubble path,
whose increasing slope and brightness reflect the experi-
enced increasing US pressure. At time t = 7.3 s, when the
bubble is at a depth of about 75 mm, i.e., close to the
transducer focal region, a modification can be observed: the
trace changes its slope and the brightness simultaneously
shows a brief increase, followed by a rapid reduction. The
trace then completely disappears at 7.35 s.
In the corresponding real-time MSD-display, the spectra
computed through 128-point fast-Fourier transforms (FFTs)
at depths between 70 and 120 mm highlight the different
Doppler shift components coming from each traveling scat-
terer. Figure 5b, in particular, reports the MSD display
obtained from the echo-data collected over an interval of
Fig. 3 BR14 microbubbles
excited with 2 MHz, 4-cycle
pulses (500 kPa PNP). a M-
mode display. b MSD display
frozen at time t = 17.5 s
Fig. 4 a Experimental velocity
profile of a population of F-04E
microspheres excited by the
Vermon transducer with
2.5 MHz, 10 cycle pulses at
1 kHz of PRF (300 kPa PNP).
The superimposed black line
reports the simulated resonant
bubble velocities;
b instantaneous peak velocities
(see text) of resonant polymer
and BR14 bubbles excited as in
a, at varying pressures
832 Med Biol Eng Comput (2009) 47:827–838
123
approximately 0.1 s around t = 7.3 s, i.e., around the
‘‘event’’. This display suggests that the bubble is accelerating
while approaching the depth around 76.5 mm. Here, the
wideband spectrum suggests that there is a strong decorre-
lation, probably due to the bubble rupture. Then, there are still
some decreasing spectral contributions up to a depth of about
78 mm, where the bubble spectrum seems to disappear.
A second example, related to BR14 bubbles, is reported
in Fig. 6. Figure 6a reports a portion of the M-mode display
in which three traces are recognizable. The brighter trace, at
about 95 mm depth, shows an abrupt change at approxi-
mately 18.8 s. The backscattered signal amplitude changes
with some fluctuations before the ‘‘critical event’’. For these
phospholipids shelled UCA, similar amplitude variations
were frequently observed before the bubble disappearance.
Figure 6b shows the MSD display frozen around the rupture
event. Again, the bubble rupture yields a significant tran-
sient Doppler signal characterized by broad spectral content.
2.3 Doppler effects due to acoustically induced bubble
deflation
Deflation involves a gradual bubble volume reduction,
which needs to be monitored over several PRI. At the same
time, backscattered echoes are expected to be characterized
by amplitude and phase closely related to the changing
bubble size. The set-up has thus been modified in order to
add the capability of measuring the bubble diameter while
continuing to acquire the corresponding echoes.
2.3.1 Methods
In this investigation, ultra low concentrations of either
BR14 or Definity (Lantheus Medical Imaging Inc., North
Billerica, MA, USA) microbubbles were pushed by a syr-
inge into a 200 lm diameter fiber immersed in water. The
bubbles were insonified using a focused transducer (Ver-
mon M3) connected to the BBT system, which also
captured the echoes. Simultaneously, the bubble in the fiber
was optically observed through a long-distance dry
microscope objective (Olympus LMPLFL 109, 0.13 NA,
Tokyo, Japan). The microscope was an upright Olympus
BX-FM with 49 extra zoom. Its image was captured by a
commercial digital video camera (MotionPro 10k, Redlake,
San Diego, CA, USA) capable of collecting up to 10,000
frame/s, and having 4 GB circular memory storage capa-
bility. The final resolution was up to 3.3 pixels/lm. Frame
synchronization to the BBT system was applied in order to
Fig. 5 F-04E thermoplastic
bubble rupture event as
observed through: a M-mode
display and b MSD display. The
bubbles were excited with 4-
cycle 4 MHz pulses at 1 kHz
PRF (MI = 0.4). Some intact
bubbles are also observable in
the region between 90 and
120 mm
Fig. 6 BR14 bubble rupture
event at 95 mm depth as
observed through: a M-mode
display and b MSD display. The
bubbles were excited with
Hanning windowed 4-cycle
2 MHz pulses at 240 Hz PRF
and MI = 0.7. Two traces
produced by intact bubbles can
be seen at 60 and 85 mm,
respectively
Med Biol Eng Comput (2009) 47:827–838 833
123
maintain the correlation between each frame and the US
echo received at the corresponding PRI. The large amount
of memory available for both the BBT system and the
camera was adequate enough to record an acoustically
driven deflation over an interval of several seconds. When
desired, the recorded frames and the US echo-signals could
be stored into data files.
Before starting each acquisition, one bubble was care-
fully positioned in the focal point through the syringe. We
were careful to isolate the bubble in a region of ±1 mm
along the fiber, which was large enough to guarantee that
no echoes could be produced by nearby scatterers.
The transmitted pulses were Hanning windowed bursts
of either 9 or 15 sinusoidal cycles at 3 or 2.25 MHz. The
PRF was set at 250 Hz, low enough to avoid interference
from static reflections in the water tank. Transmitted pulses
could produce PNP in the range 50–600 kPa.
When a deflation phenomenon was detected, both video
and acoustic raw data were recorded until a stable bubble
state was reached. Video frames were post-processed in
Matlab (Mathworks Inc., Natick, MA, USA) to estimate the
instantaneous bubble radius. The received echoes were
quadrature-demodulated and Doppler processed in the end
through standard spectral analysis and autocorrelation
methods [42].
2.3.2 Results
During these experiments, we observed the behavior of
about 100 microbubbles.
In general, the bubbles did not present any deflation,
either in absence of US excitation, or when they had an
initial diameter much larger than the resonant one. For
smaller bubbles (e.g., no more than 5 lm diameter for
BR14 insonified at 3 MHz) the typical dissolution profile
[37] started with a slow diameter reduction, followed by a
faster reduction when the bubble radius was on the order of
its resonant size. While becoming smaller than the resonant
size, the size reduction slowed down again and was ulti-
mately stopped, after which the diameter remained
constant despite of the US excitation.
The deflation rate actually depended on the excitation
conditions. For example, BR14 excited with 5-cycle,
3 MHz, 210 kPa (PNP) pulses, showed a typical rate of
about 1 lm/s. When the diameter approached the resonant
diameter, the deflation rate was always much faster.
During the deflation process, the backscattered echo
amplitude changed according to the instantaneous diameter.
Two examples are reported in Fig. 7. Figure 7a shows
the experimental results obtained when a Definity bubble
deflated from 2.3 lm to approximately 1.5 lm, corre-
sponding to about 300 pulses. The radius changed with a
sigmoid shape, similar to that observed by Borden et al. [8]
for bubbles deflating with an intact shell. Figure 7b reports
on the deflation of a BR14 bubble. The deflation rate
increases as the diameter reaches about 2.2 lm, and a
narrow amplitude peak can be observed, immediately
before the rate becomes maximum.
In both cases, the phase of the echo, arbitrarily set to
zero when the measured bubble diameter was largest,
clearly decreases. In Fig. 7b, the phase change is over -p/
2, and the minimum value is reached when the bubble is
totally deflated and also the received echo amplitude is
very low.
The echo change in terms of both amplitude and phase
yields a decorrelation which can be evidenced through
Doppler processing. Figure 8a shows the Doppler spectra
obtained for the above BR14 bubble through two 128-point
FFTs (i.e., each covering a time interval of about 0.1 s)
evaluated before and during the fast deflation phase,
Fig. 7 Bubble radius (R0),
normalized amplitude (A) and
phase (Dh), as measured during
deflation. a Definity
microbubble excited with 15-
cycle 2.25 MHz Hanning
windowed pulses at 250 Hz
PRF (50 kPa PNP); b BR14
microbubble excited with 9-
cycle 3 MHz Hanning
windowed pulses at 1 kHz PRF
(600 kPa PNP)
834 Med Biol Eng Comput (2009) 47:827–838
123
respectively. The FFT related to the time interval over
which the bubble diameter rapidly changes, evidences an
appreciable spectral broadening.
Figure 8b shows the results of a 16-point autocorrelation
algorithm applied on the full acquisition interval. The
estimated mean Doppler frequency is about zero every-
where, except when the phase of the echo signal rapidly
changes, i.e., when the diameter is closer to resonance.
3 Discussion
Doppler US studies are traditionally committed to detect-
ing blood movement through analysis of phase changes in
echoes backscattered by RBC. When contrast microbub-
bles are intravenously injected, it is assumed that they are
dragged by blood flow and move at the same velocity
amplitude and direction as the RBCs [48], and that any
phase change in the echo can be attributed to the dis-
placement of the scatterers. The goal of this paper was to
investigate how valid such an assumption is in the presence
of phenomena which are specifically associated with UCA.
We have chosen an experimental approach in which single
bubbles were taken into consideration through acoustical
and optical methods.
US excitation pushes the microbubbles, and the velocity
vector is shifted accordingly toward the US wave propa-
gation direction. For resonant, or nearly resonant bubbles,
two different types of velocities are actually involved. One
can be very high (m/s), is reached in a few microseconds
and is maintained as long as the US pulse excites the
bubble. The other is the ‘‘apparent’’ mean velocity, i.e., the
one detected with pulsed Doppler methods, is lower by a
factor corresponding to the pulse duty cycle (which is
typically about 0.01).
In vitro experiments [65] have shown that the displace-
ment phenomenon can produce perceptible contributions
during Doppler analysis, especially in terms of spectral
broadening. However, in clinical applications this behavior
is strongly restrained by blood viscosity, as shown in
Fig. 4b, and by the maximum allowed US intensity due to
safety regulations.
Secondary radiation forces between oscillating bubbles
might, in theory, corrupt the Doppler signal as the bubbles
are mutually attracted and repulsed due to ultrasound.
However, this force and resulting displacement of bubbles
will have random distribution of strength and direction, so
on average the effect should be zero, although it might lead
to a broader Doppler spectrum.
The strong decorrelation of echoes associated to single
rupturing bubbles has already been observed [16]. Our
experiments, although preliminary, clearly show that
breaking bubbles yield echoes with different signatures
which could perhaps be associated to different rupture
mechanisms. For example, when observing lipid shelled
bubbles through the M-Mode display, we frequently
detected quick (PRI-to-PRI) brightness changes in traces.
This seems to be indicative of fragmentation and/or coa-
lescence. In experiments with the polymer-coated F-04E,
we sometimes saw division of a single trace into multiple
traces, each with different brightness and slope, which
could also indicate fragmentation. When a trace shows a
brighter spike, followed by a gradual brightness decrease
and slope change, it is suggestive of possible sonic
cracking.
In the frequency domain, bubble rupture always deter-
mines an instantaneous wideband Doppler spectrum.
In this study, deflation has been shown to yield dramatic
amplitude and phase changes, which in principle, could be
erroneously interpreted in Doppler terms. Actually, the
FFT of echoes produced by a bubble during its rapid
deflation has shown an evident spectral broadening. More
evident effects are obtained if the same bubble echoes are
analyzed through autocorrelation, the technique which is
Fig. 8 Results of Doppler
analysis of echoes from the
bubble observed in Fig. 7b.
a 128-point, Hanning weighted
FFT spectra evaluated at time
2.2 and 2.9 s; b mean Doppler
frequency shift versus time,
evaluated through 16-point
autocorrelation
Med Biol Eng Comput (2009) 47:827–838 835
123
most used in color Doppler systems. Here, the detected
phase changes can correspond to Doppler shifts in the
range of 30–40% of the PRF. Hence Doppler effects are
remarkable, but only during a small part of a process which
is typically slow and requires a large number of pulses to
be transmitted on the same bubble.
Although the observation of single bubbles based on
optical and/or acoustical approaches is the best means of
performing detailed investigations of each phenomenon, it
is best to remember that clinical applications employ full
populations of UCA containing a range of diameters cor-
responding to a wide range of resonance frequencies. This
has practical consequences on the extent at which all the
observed phenomena produce Doppler effects. Since only
resonant bubbles are appreciably displaced by radiation
force [19, 69], their effects on Doppler spectrum can be
masked by the contributions from all other, non resonant
bubbles. In fact, it was shown by Tortoli et al. [66] that, in
standard clinical conditions, the bubble displacements were
not sufficient enough to interfere with Doppler measure-
ments. Similar considerations may apply to the deflation
phenomenon which, although not equally selective, seems
to happen more likely for bubbles with a radius in the range
of a resonant one.
Other phenomena typical of UCA, such as compression-
only behavior and thresholding effect, could also, in prin-
ciple, yield modifications in the detected Doppler signals.
Compression-only behavior [21], where the bubble shows
higher compression than expansion during oscillation, will
result in echoes having a major increase of higher har-
monics and a reduction of fundamental frequency
scattering. Thresholding [29] is the effect where the
acoustic pressure has to exceed a certain acoustic level
before the coated bubble starts to oscillate. However, it was
suggested that both compression-only and threshold
behavior in their selves are repeatable processes [29, 50],
and hence the echo of a still bubble will be equal for
multiple Doppler US pulses. Therefore, no spurious
Doppler contributions are expected from such repeatable
processes.
More evident Doppler effects in a full microbubble
population have been reported for the bubble rupture
phenomenon. Tienmann [64] observed that in Harmonic
Power Doppler Imaging the rupture leads to the appearance
of distributed pixels with random noise while in color
Doppler there are pixels with different colors, in contrast to
the monochromatic background. This finding is consistent
with our results for single bubbles as shown in Fig. 8b.
Bevan [3] recently noticed that the frequently observed
‘‘flash’’ in Doppler imaging comes from decorrelation
rather than as a direct consequence of any stimulated
acoustic emission. In [67], the different effects of micro-
bubble destruction (wideband noise) and displacement
(spectral broadening) in Doppler measurements for UCA
populations were highlighted.
Finally, the fact that phenomena such as bubble dis-
placement and deflation are not expected to produce
evident Doppler effects in vivo can be considered either
good news or bad news, depending on the investigation
targets. As discussed before, in standard clinical investi-
gations it is certainly good that these phenomena do not
interfere with the Doppler exam. However, in new UCA
applications such as targeted drug delivery [60], the
capability of displacing and deflating contrast bubbles can
be very useful. It can be derived from the results presented
in this study that such standard transmission signals are not
ideal in emphasizing such phenomena. Hence, more studies
are needed to obtain an optimal control of inherent
mechanisms in order to increase the efficacy of contrast
agents in such applications.
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