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

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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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