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Файл:Ординатура / Хирургия / Библиотека им академика М.И. Перельмана / Книга_5814_Библиотеки_им_академика_М_И_Перельмана
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138 Diagnostic ultrasound: imaging and blood ow measurements
(a)
(b)
Figure 4.44 Portable scanners from three different manufacturers: (a) Fuji
Sonosite (courtesy of Sonosite), (b) pocket scanner with a 3 MHz element-phased
array (courtesy of Siemens), and (c) Vscan phased array system with color Doppler
(courtesy of GE).
139Chapter four: Gray-scale ultrasonic imaging
(c)
Figure 4.44 (Continued) Portable scanners from three different manufacturers: (a)
Fuji Sonosite (courtesy of Sonosite), (b) pocket scanner with a 3 MHz elementphased array (courtesy of Siemens), and (c) Vscan phased array system with color
Doppler (courtesy of GE).
Another advancement is the elimination of the cable that connects
an ultrasonic probe to the imaging console by adopting wireless technology. Figure4.45 shows a wireless portable scanner developed by Siemens
where the wireless radio frequency (RF) signal is transmitted at 8.7 GHz.
Figure 4.45 A wireless scanner capable of color Doppler. (Courtesy of Siemens.)
140 Diagnostic ultrasound: imaging and blood ow measurements
(a)
(b)
Although ultrasound computed tomographical scanners have only
found limited success in breast imaging, a new breast scanner very similar
to x-ray mammography has recently been introduced. A unique feature
of this scanner is the concave curved linear array probe that conforms
better to the breast curvature. The linear array is mechanically scanned
linearly to acquire multiple slices of B-mode images for later 3D reconstruction. It was approved by the FDA for dense breast scanning as a complementary tool for x-ray mammography. Figure 4.46(a) shows how the
Figure 4.46 (a) Automated breast ultrasonic scanner and (b) corresponding coronal image. (Courtesy of GE.)
141Chapter four: Gray-scale ultrasonic imaging
linear mechanical scanning of the linear array is performed on a patient
and (b) a coronal B-mode breast image acquired by such a scanner, which
shows a lesion at the lower right-hand corner. A coronal view, which is not
possible with a conventional ultrasound scanner, can be obtained following 3D reconstruction.
References and Further Reading Materials
Flax SW and O’Donnell M. Phase-aberration correction using signals from point
reectors and diffuse scatterers: Basic principles. IEEE Trans Ultrasonics
Ferroelect Freq Cont 1988; 35: 768–778.
Greenleaf JF. Computerized tomography with ultrasound. Proc IEEE 1983; 71:
330–337.
Karaman M and O’Donnell M. Synthetic aperture imaging for small scale systems.
IEEE Trans Ultrasonics Ferroelect Freq Cont 1995; 42: 429–442.
Nock L and Trahey GE. Phase aberration correction in medical ultrasound using
speckle brightness as a quality factor. J Acoust Soc Am 1989; 85: 1819–1826.
O’Donnell M. Coded excitation system for improving the penetration of real-time
phased array imaging system. IEEE Trans Ultrasonics Ferroelect Freq Cont
1992; 39: 341–351.
Ophir J and Maklad NF. Digital scan converters in diagnostic ultrasound imaging.
Proceedings of the IEEE 1979; 67: 654–664.
Shung KK, Smith MB, and Tui BWN. Principles of medical imaging. San Diego:
Academic Press, 1992.
Shung KK and Thieme GA. Ultrasonic scattering by biological tissues. Boca Raton,
FL: CRC Press, 1993.
Thomenius KE. Evolution of ultrasound beamformers. In Levy M, Schneider SC,
and McVoy BR (eds.), Proceedings of the 1996 IEEE Ultrasonics Symposium,
New York, 1996, pp. 1615–1622.
Wagner RF, Smith SW, Sandrik JM, and Lopez H. Statistics of speckle in ultrasound
B-scans. IEEE Trans Soncis Ultrasonics 1983; 30: 156–163.
chapter ve
c
2cos
Doppler ow measurements
As was discussed earlier in Chapter 2, the Doppler effect provides a
unique capability for ultrasound to measure blood ow (Evans and
McDicken, 2000; Jensen, 1996). Upon insonication by an ultrasound
beam, the echoes scattered by blood carry information about the velocity of blood ow. Blood ow measurements are frequently performed
in a clinical environment to assess the state of blood vessels and functions of an organ. Ultrasonic Doppler instruments allow a measurement
of instantaneous blood ow velocity. Combined with pulse-echo instruments, instantaneous ow rate in a blood vessel as a function of time and
cardiac output can be measured noninvasively with ultrasound. At present, very few clinical options are available to do so. Figure5.1 shows an
ultrasound beam of frequency f insonifying a blood vessel, making an
angle of θ relative to the velocity v. Here it is assumed that blood ows in a
vessel with a uniform velocity v. The returned echoes are Doppler shifted.
The Doppler shift frequency fd is related to the ultrasound frequency f by
Equation (2.43):
v
f
=
d
θ
f
where c is the sound velocity in blood and may be assumed to be 1540 m/s.
The Doppler-shifted frequencies happen to be in the audio range for blood
ow velocities in the human body for an ultrasound frequency between
1 to 15 MHz.
Conventionally, two different approaches have been used for ultrasonic Doppler ow measurements: continuous-wave (CW) and pulsedwave (PW) Doppler.
5.1 Nondirectional CW owmeters
A CW system is shown in Figure 5.2. A probe consisting of two piezoelectric elements, one for transmitting the ultrasound signal and one
for receiving echoes returned from blood, is excited by an oscillator.
The Doppler-shifted echoes are amplied, demodulated, and band pass
ltered to remove the carrier frequency and other spurious signals.
Suppose that the ultrasound signal generated by the oscillator is given by
143
144 Diagnostic ultrasound: imaging and blood ow measurements
Bt
dd
1
2
()
shift
am
R
elemen
n
Transducer
Skin surface
Blood vessel
v
θ
Doppler equatio
2vcosθ
=
f
d
c
f
Figure 5.1 An ultrasound beam is incident upon a blood vessel and makes an
angle of θ relative to the direction of blood ow.
Acos(ωt), where A denotes signal amplitude and ω (the angular frequency)
= 2πf. The demodulated signal would be
gAtB tA
(, )cos()cos[()]
dd
{cos[(2)]cos
t
}ωω =ω ω+ω= ω+ω+ ω
d
where the echoes are represented by Bcos[(ω + ωd)t] and ωd = 2πfd. The
magnitude of constant B is determined by the scattering strength of
blood. Much work has been done to better understand the relationship
between the Doppler power generated by blood and hematological
Master
Transducer
eceiving
t
Transmitting
element
oscillator
Amplifier
Speaker
Doppler sound Average Doppler
Coherent demodulation
Demodulator
Filter
Zero crossing
counter
Figure 5.2 Block diagram of a CW Doppler owmeter.
Spectrum
analyzer
Spectrasonogr
Time domain Frequency domain
Transmitted signal
f
0
t
Received signal
f
0
f
c
t
Demodulated signal
f
t
fc = Frequency of clutter signal e.g., from blood vessel wall
f
c
f
d
145Chapter ve: Doppler ow measurements
f
d
f
f
Figure 5.3 Doppler signals in the time and frequency domain showing the effect
of demodulation.
and hemodynamic factors (Shung et al., 1992; Mo and Cobbold, 1993).
Doppler power from blood has been found to be related to ow disturbance, hematocrit, and the degree of red blood aggregation, which is in
turn affected by the concentration of plasma proteins such as brinogen and local shear rate. The output of the demodulator contains both
the ultrasound carrier frequency and the Doppler shift, illustrated in
Figure 5.3, where the signals in the time and frequency domains are
shown, respectively, on the left and right. The carrier signal can be
readily removed by band pass ltering by setting the cutoff frequency
of the band pass lter at the high end to be much lower than the carrier
frequency. A problem in ultrasonic Doppler blood ow measurement is
that the blood vessels that produce large reected echoes are slow moving as well. In Doppler terminology, these large, slow-moving echoes are
called clutter signals, shown in Figure5.3 as fc. The cutoff frequency of
the band pass lter at the low end has to be designed to minimize the
interference of these clutter signals. The design of this band pass lter
in the low-frequency region, which performs the function of high pass,
146 Diagnostic ultrasound: imaging and blood ow measurements
∫
0
High pass or wall filter response
f
c
Figure 5.4 Clutter rejection lter or wall lter is used to suppress large echoes produced
by slow-moving blood vessel walls. fc = frequency of clutter signal (e.g., from blood
vessel wall).
f
d
f
also called clutter rejection lter, has been problematic due to the fact
that the magnitude of clutter signals is several orders higher than those
from blood and may mask those from slow-moving blood (Figure5.4).
A lter with a very steep slope or a method that carries out some forms
of echo cancellation may be used (Jensen, 1996). The signal after bandpass ltering can be processed in different ways. It may be heard with
a speaker since the Doppler shift is in the audible range. Alternatively,
a zero-crossing counter can be used to estimate the mean Doppler frequency, or a spectrum analyzer can be used to display the spectrum.
The zero-crossing counter estimates the number of zero-crossings of a
signal. The number of zero-crossings, N, and the mean frequency, fm, of
a signal are given, respectively, by
∞
2
fPfdf
()
∫
N
0
2
=
∞
Pfdf
0
()
(5.1)
∞
fP fdf
()
∫
0
f
=
m
∞
Pfdf
∫
(5.2)
()
where P(f) is the probability density function at frequency f. For a pure
sinusoidal signal of frequency fm, N = 2fm. Complication arises if the signal is not sinusoidal, as in the case of Doppler ow measurements, where
147Chapter ve: Doppler ow measurements
=−
r
Cross-sectional view
of the vessel
R
r
Figure 5.5 Laminar blood ow in an artery has a parabolic ow prole.
v(r)
v
max
the blood ow is not uniform. For blood ow in a vessel, the velocity is
related to radial distance r, shown in Figure5.5, by the following equation
(Nichols and O’Rourke, 1990):
n
r
(5.3)
R
where v
vr v
() 1
max
is the peak velocity, n is an index indicating the nature of ow,
max
and R is the radius of the blood vessel. For parabolic ow, n = 2 and N =
1.15 f
, where f
max
is the maximal Doppler frequency.
max
The spectrum is usually displayed in the format shown in Figure5.6(a),
where the vertical axis indicates Doppler frequency or velocity, the horizontal axis indicates time, and the gray scale indicates the intensity of the
Doppler signal at that frequency or velocity. Figure5.6(b) illustrates how
one vertical line in the Doppler sonogram is calculated and the complete
Doppler sonogram obtained. At each instant of time, the vertical line displayed represents the Doppler spectrum calculated at that time within a
5 to 10 ms time window. For a timescale of many seconds, the Doppler
spectrum obtained within a very short window at a certain time can be
represented by a line. From the Doppler spectrum, the mean frequency
or other frequencies, e.g., median frequency, where the Doppler power
spectrum is split into two equal halves, and mode frequency, where the
Doppler power is the highest, can be readily estimated.
Doppler owmeters have been used to noninvasively assess vascular
disorders. Flow disturbances near a stenosis cause the Doppler spectrum
to broaden because blood ow velocity uctuates. There is, however, a
caveat that must be recognized to avoid misdiagnoses: transit time spectral broadening. This is illustrated in Figure5.7(a) for a single scatterer
traversing an ultrasound beam at velocity v. A nite time is needed for the
scatterer to traverse the beam. In the time domain, there is a nite time
duration, which is dened by Δt (–3 dB time duration from peak value),
148 Diagnostic ultrasound: imaging and blood ow measurements
(a)
Echo amplitudefromamovingstructure
city
(b)
t
1
An FFTisperformedonthewindowedwaveformandacolorscaleisusedto
Figure 5.6 (a) Spectrasonogram of CW Doppler signals produced by a mitral
valve regurgitation jet. The top image is a B-mode apical four-chamber view of
the heart. The dotted line indicates the direction of the Doppler beam. (Courtesy
of GE Medical Systems.) (b) How a spectrasonogram is calculated.
4μs
denotethesignalamplitude
Time
Frequency or velo
t
1
Time
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