Добавил:
Sekretar
kiopkiopkiop18@yandex.ru
t.me/Prokururor I Вовсе не секретарь, но почту проверяю
Опубликованный материал нарушает ваши авторские права? Сообщите нам.
Вуз:
Предмет:
Файл:Ординатура / Хирургия / Библиотека им академика М.И. Перельмана / Книга_5772_Библиотеки_им_академика_М_И_Перельмана.pdf
1 Physical and Technical Principles of Color Doppler Sonography
R. Kubale and G. Hetzel
Historical Development
Duplex sonography and its offshoot, color duplex sonography,
are relativelynew methods that combine the pulsed echo technique of sectional image formation with the Doppler evaluation of blood flow. It is based on B-mode ultrasound, which
permits an accurate morphological description of gynecological and obstetric findings as a result of increasingly higher resolution and faster frame rates. The first attempts to measure
blood flow date back to Satomura
1970s that Pourcelot
Doppler to investigate cerebrovascular diseases. Since 1977,
1
11
was able to use continuous wave (CW)
12
. It was not until the early
B-Mode Sonography
Physical Principles of Echo Production
The ultrasound wave is a density and pressure wave that propagates through a medium, its behavior essentially conforming
to the laws of sound and optics. Sound attenuation, sound reflection, and resonance phenomena can be utilized as means of
identifying defects in the insonated medium. The pulse-echo
principle is most commonly used in medicine and will be discussed below in greater detail.
Acoustic impedance. The echoes that are displayed in an ultrasound image are based on reflections that occur when the
sound wave travels through media of different acoustic impedance. Acoustic impedance (Z) is a material-dependent (i.e.
tissue-dependent) quantity that is the product of the density of
the material (
being, we shall assume that this velocity is a material- and
tissue-independent constant (equation 1)
Processes and Approaches, p. 25).
ρ) and the velocity of sound in it (c). For the time
(see New Technical
authors have describe d Doppler ultrasound applications in the
abdomen and especially in obstetrics and gynecology
During the past 15 years, these applications have become
widely adopted and established through the use of color-flow
techniques and ongoing refinements in instrumentation
Below we shall explore the physical and technical principles that are essential for optimally utilizing and interpreting
B-mode and color Doppler sonography and avoiding misinterpretations.
basic frequency (harmonic energy). These frequencies, which
create unwanted signals in the conventional B-mode image
(see Principle of Echo Detection and Scanning Techniques,
p. 3), can be utilized for specific “harmonic” imaging. This technique, called tissue harmonic imaging (THI), makes it possible
to examine even patients who are technically difficult to scan
(see New Techniques of Signal Acquisition and Processing,
p. 25).
The acoustic impedance of tissues has a value similar to that
of water. The exact value varies with the composition of the
tissue (Table 1.
sues are markedly different from the impedances of air and
bone.
Table 1.1 Impedance values and degree of reflection in biological tissues
Medium Sound
1). Note that the acoustic impedances of soft tis-
Reflection
)
factor R
(relative to
H
velocity
c (m/s)
Specific
density
ρ (g/cm
Impedance
Z (g cm
3
)
–2s–1
3, 13, 16
10
.
O)
2
.
Z = c ⫻ ρ (1)
Because the sound wave consists of alternating high and low
pressure peaks, the tissue that transmits the wave is transiently compressed, and the sound velocity changes: the
“peaks” at a higher pressure move faster than the “valleys” at a
low pressure. This phase-dependent change in sound velocity
distorts and steepens the original sine wave, analogously to a
2
water wave breaking on the shore. As a result, the reflected
wave acquires a component that is a multiple of the emitted
Water 1496 0.997 1.491 0.000
Fat 1476 0.928 1.370 0.042
Muscle 1568 1.058 1.660 0.054
Liver tissue 1570 1.055 1.660 0.054
Bone 3360 1.850 6.200 0.614
Air 331 0.0012 0.3972 0.999
Impedance (Z) is defined as the product of the specific density (ρ) of the tissue
and the sound velocity (c). An incident ultrasoundbeam is partially reflected at
interfaces between tissues of different impedance. The degree of the reflection is measured by the reflection factor R in relation to water.
Reflection. Reflection occurs at every interface between two
media that have different acoustic impedances. The reflected
portion of the ultrasound wave (the echo) increases with the
magnitude of the acoustic impedance difference. The amount
of reflection that occurs at the interface between two different
tissues with impedances Z
reflection factor R (Table 1.
R =
Z
Z
tissue1
tissue1
– Z
+ Z
tissue2
tissue2
and Z
tissue1
1) as given by equation (2):
is described by the
tissue2
(2)
If the impedance mismatch between the tissues is small, the
transmitted ultrasound wave will retain enough energy to produce additional echoes in deeper tissue layers. But when a
large impedance mismatch exists, such as tissue–air or tissue–
bone interfaces, almost all of the incident sound energy is reflected. Objects located behind such an interface cannot be
visualized with ultrasound.
B-Mode Sonography
Principles of Ultrasound Instrumentation
Principle of Echo Detection and Scanning
Techniques
Ultrasound pulses are generated and processed by means of
piezoelectric elements that are assembled into a transducer. In
principle, a transducer is both a transmitter and a receiver of
sound. An electric pulse (pulse length 0.5–2 wave trains of 1
duration) excites the active elements of the transducer. The resulting mechanical vibration propagates through the tissue.
The mechanical vibration is reflected from a target object, and
the returning sound wave induces the transducer elements to
generate an electric signal. The time (t) between the emission
of the pulse and the reception of the echo is a measure of the
distance (z) of the transducer elements from the reflecting object (pulse-echo principle)
(equation 4):
z =1/2 ct (4)
µs
Refraction. Reflection is strongly angle-dependent and is subject to the same laws as optical refraction. If the sound strikes
the interface between different tissues at a perpendicular
angle, most of the reflected wave will return to the transducer.
But if the sound strikes the interface at an oblique angle, only
part of the reflected wave will reach the transducer, and the
wave will deviate from its original direction as it continues on
through the second medium (refraction).
Scattering. If the sound beam encounters a rough surface or
small reflectors whose diameters are significantly smaller than
the wavelength of the sound, the beam will be reflected or
“scattered” in various directions. This occurs not just at interfaces but everywhere in the tissue. Because of scattering, homogeneous soft tissue is depicted as an interference pattern
with a coarse or fine texture, depending on the transducer
frequency (Speckle).
Absorption and attenuation. In addition to losses due to reflection and scattering, some of the ultrasound energy is absorbed
in a frequency-dependent fashion as it overcomes cohesion
and relaxation forces in the medium. The amplitude of the
sound pressure declines exponentially with the thickness of
the medium traversed—i.e., it is attenuated. Attenuation refers
to the relationship of the initial sound pressure to its pressure
after the sound has propagated for a given distance. It depends
on the distance traveled by the sound pulse, the frequency of
the transducer, and a material-specific constant. The following
simplified formula
(equation 3)
is valid for soft tissues and for
frequencies from 0.2 MHz to 100 MHz:
where c is the velocity of sound and t is the time from pulse
transmission to echo reception.
B-mode image. A two-dimensional sectional image is produced when a number of adjacent ultrasound beams are transmitted and received in the same plane. In the B-mode technique (for “brightness mode”), the echoes are displayed at a
brightness that is proportional to their amplitude. The brightness-modulated signals that are received (ultrasound lines)
are temporarily stored in a matrix. The contents of this matrix
are transferred to a monitor at correlative sites and assembled
into a geometrically correct ultrasound sectional image
(Fig. 1.
1).
Physical and Technical Principles
Attenuation (dB) = frequency (MHz)⫻
distance traveled (cm) (3)
Fig. 1.1 Production of an ultrasound image, illustrated for a linear
array transducer. Transverse abdominal scan through the liver, kidneys, pancreas, and spine. Whena pulse is transmitted, portions of the
ultrasound energy are reflected from the organ surfaces and from in-
terfaces within the organ parenchyma. These echoes are assigned to
specific locations based on the transit time for a known sound velocity.
The sum of the brightness-modulated points yields a sectional image
in which the liver, aorta, pancreas, and kidneys can be evaluated.
3
Physical and Technical Principles of Color Doppler Sonography
The first ultrasound imagers were based on the manual
movement of a single transducer over the patient (compound
scan). These were initially replaced by mechanical transducers
in which an element was rotated to produce a fan-shaped
beam pattern.
Transducer arrays. Most scanners today employ transducer arrays consisting of individual elements arranged in a closely
spaced row. Each array has a certain number of elements, in
some cases more than 196.The elements are fired (excited) in
groups, each of which transmits and receives an ultrasound
line. Starting on one side of the array, the edge line is acquired
first. Adding an element on one side of the group and turning
off an element on the other side, the active group is shifted
across the array. As the next group transmits and receives, the
tissue is scanned in a sequential fashion. As this process is con-
a Linear b Convex
1
B
A
AB AB
B
A
tinued, a sectional image is produced. If the elements in the
array are arranged in a straight row, the transducer is classified
as a linear array.If the elements are arranged along a curve, the
unit is called a curved or convex array (Fig. 1.
2).
Electronic beam steering. A phased array scanner can scan the
tissue in a sector- or trapezoid-shaped pattern. All of the elements in the array are active (group width m equals the number of elements n). A sector-shaped scan is produced by elec-
tronic b eam steering (Fig. 1.
3). The wavefront is steered by im-
posing time delays on the excitation pulses that are applied to
the elements in the array. In this process the individual narrow
elements (of width approximately 0.5
λ; λ = wavelength) be-
have like “Huygen point emitters,” meaning that each element
is the source of a spherical wave in relation to the scan plane.
All of the spherical waves add together to form a wavefront
that is angled away from the array as determined by the time
delay (Fig. 1.
3). To accomplish electronic steering in an array,
the elements must be spaced sufficiently close together, i.e.,
there must be a sufficiently fine element pitch or high element
density. The individual elements must be so narrow that their
directional characteristic is broad enough to allow adequate
lateral signal transmission. The returning echoes must not be
extinguished by the width of an individual element.
Recent technical developments include transducers with a
trapezoidal beam pattern. The “trapezoid scan” is produced by
sweeping a linear-array image field and represents the special
case of a sector scan with a virtual pivot point located behind
the transducer surface (Fig. 1.
2). Only linear arrays that have a
high element density can produce a trapezoid scan (see New
Developments in Transducer Technology, p.23).
d Trapezeiumc Sector
Fig. 1.2 Functional principles of different transducer types.
a In a linear array, the elements are fired in offset groups at times A
and B, producing a rectangular beam pattern.
b In a convex array, the transducer has a curved surface resulting in a
fan-shaped beam.
c Sector scan with a phased-array transducer. All the elements are ac-
4
tive.
d Trapezoid scan. This unit is operated like a phased array with a “virtual pivot point” (VD) located behind the transducer.
VD
Transmitter
Delay circuits Pulses at
τ
1
τ
4
τ
8
τ
n
t0−τ
time t
1
t0−τ
t
Array
1
4
8
n
n
Effective
aperture A
ϕ sound
direction
c=τ
– c
n
c=sound
velocity
Wavefront
s
0
Fig. 1.3 Phased array with electronic beam steering. The phased
τ
array transducer usesa variable electronic time delay (
τ
direction of the beam. If the time delay
is greater than τn, the excita-
i
) to changethe
i
tion pulses in the transducer will reach element n of the array first,
before reaching element 1. The superimposed wave segments will
combined to form a wavefront that is steered away from the transducer axis by the angle
, depending on the time delay.
B-Mode Sonography
Transducer Selection: Penetration Depth, Beam
Pattern, and Resolution
The selection of a transducer or scanning technique (see Principle of Echo Detection and Scanning Techniques, p. 4) is
guided by the nature of the desired information and by anatomical considerations. Important decision-making criteria
are the size of the region under study, the necessary scanning
depth, and the necessary resolution.
Field-of-view width and penetration depth. Since the linear
array produces a wide near field, this type of array is preferred
for scanning superficial regions. The convex array has a
broader far field with good near-field resolution, making it an
ideal compromise for abdominal scanning. The phased array
has a sector format that is particularly useful for difficult imaging conditions, as in echocardiography for scanning deeper
structures. The trapezoid scan has advantages similar to the
convex scan in terms of good near-field resolution and far-field
width. An additional advantage, as in the phased array, is a
small, straight transducer face that is easier to couple to the
body surface for numerous applications.
Penetration depth and transmission frequency. The penetration depth is the maximum distance between the sound transmitter (transducer) and the deepest site in the tissue from
which a reflection (echo) can still be acquired. Owing to the
frequency-dependent absorption of ultrasound in tissue, the
penetration depth is not only tissue-dependent but also
frequency-dependent. For this reason, the selected transmission frequency is a tradeoff between penetration depth and
desired resolution and thus depends on the nature of the study.
High frequencies are well suited for the high-resolution imaging of superficial structures, while low frequencies are better
for imaging deeper structures. The maximum penetration
depth is inversely proportional to the transmission frequency.
ab
Pulse length
Resolution
Slice thickness
Fig. 1.4 Definition of axial and lateral resolution.
a Two interfaces or point reflectors located at different depths along
the beam axis can still be recognized as separate structures if the distance between them is greater than one-half the spatial length of the
emitted pulse. Despite depth-dependent and frequency-dependent
tissue absorption with an increasing shift of the echoes toward lower
frequencies, the axial resolution of the image remains fairly constant.
b Lateral resolution depends on the varying width of the ultrasound
beam with increasing depth. Two adjacent point reflectors in the
beam can still be identified as separate points if the distance between
them is greater than the “6-dB beam width” at the depth of interest.
This parameter is useful for describing the spatial resolution of two
points in the transducer planeand for estimating the effect of point re-
flectors or interfaces that are directly adjacent to the imaged surface
in the longitudinal direction (scan plane–thickness phenomenon).
Scan plane
Physical and Technical Principles
Resolution. Resolution defines the smallest distance between
two point reflectors that can still be recognized as separate
points. Axial resolution denotes resolution along the axis of the
beam. Lateral resolution describes resolution at right angles to
the beam axis (Fig.1.
4).
Axial resolution. Axial resolution is determined chiefly by the
spatial length of the emitted ultrasound pulse. Short pulse
lengths combine d with a broadband transducer (see below)
provide the best axial resolution. Typical values for axial resolution are 2
λ (λ = wavelength). Consequently, the axial resolu-
tion is usually better than the lateral resolution and increases
with the ultrasound frequency.
Lateral resolution and beam pattern. Lateral resolution is one
of the basic parameters for describing the quality of image formation and affects the diagnostic performance of the system as
a whole. Generally it is in the range of 3–5 wavelengths and depends on the ultrasound beam pattern.
Figure 1.
5 shows a beam pattern in which all of the intensi-
ties are normalized to the maximum sound pressure. It is easier to analyze resolution in this type of diagram. Owing to interference effects, even a flat transducer produces a con-
Beam
Near field
Focal zone
Far field
Fig. 1.5 Near field, far field, and focal zone of an ultrasound beam (the
area insonated by an emitted ultrasound pulse). The near field has an
inhomogeneous interference structure that greatly limits the image
quality. Resolution is best in the focal zone, where the beam is constricted. As the beam diverges again in the far field, there is an increasing deterioration of lateral resolution.
Transducer
Center line
Lateral cross
sections throught the beam
Lateral
extent of a
sound beam
Intensity at
center line
Axial
5
Physical and Technical Principles of Color Doppler Sonography
stricted beam pattern with a natural focus. The lateral resolution is best within the focal zone of the transducer. The depth of
the focal zone is always frequency-dependent. The focal zone Z
of a flat transducer (with a natural focus) is defined by equation
(5):
Zf= kA2/λ (5)
where A is the aperture, or the active transducer length used
for transmission and reception, and
λ is the frequency-depend-
ent wavelength.
The boundary between the focal zone of the transducer and
the near field, which is marked by inhomogeneities (extinction
phenomena), is defined as twice the value of the lateral resolution at the focus. This near-fieldportion of the image is difficult
or impossible to interpret because of artifacts. The far field
begins at twice the value of the lateral resolution at the focus.
The focus can be moved toward the transducer, thereby improving resolution, by the use of acoustic lenses, such as ones
made of silicone rubber, by curving the transducer face (mechanical focusing), or by electronic focusing. For a given
frequency, it is advantageous to use the largest possible aper-
1
ture with suitable focusing. The advantages of a large aperture
are greater penetration depth and better lateral resolution. The
aperture can be reduced during reception by modifying the
number of active elements (dynamic aperture) to minimize interference in the near field.
Resolution in the z axis. A third, frequently neglected, parameter is the resolution in the plane perpendicular to the trans-
a
Transmit focusing
Array elements
5
τ
3
3
τ
1
1
τ
0
0
1
3
s
5
s=path in the tissue
Delay line
Focus
´
´
´
b
Addition
Receive focusing
5
τ
3
3
τ
1
1
τ
0
0
1
3
5
Fig. 1.6 Transmit and receive focusing.
a In transmit focusing, the excitation pulses to the array elements are
delayed such that the outer crystals are fired before the inner ones, allowing the beam to be focused at different depths.
b In receive focusing, echoes returning from the point designated as
the focus arrive at element 0 first, where the return path is the shortest. If the echo-induced signals from that element are delayed by time
τ
, they can be time- and phase-matched to the signals from elements
0
5 and 5′. If all of the echo-induced signals are delayed according to the
6
differences in their return times, all the elements will yield coherent
signal components, so that adding the delayed signals will give an optimum echo display for the focal point.
ducer (z-axis resolution). This parameter is described by the
slice thickness from which reflectors and scatterers located
outside the transducer plane can generate an echo in the trans-
f
ducer plane (see Artifacts and Pitfalls, p. 8). The slice thickness
focus of a given transducer is dependent on the transmission
frequency.
Signal Processing: Focusing, Preprocessing, and
Postprocessing
Focusing. Every type of transducer that is to be connected to an
ultrasound imaging system requires special front-end circuitry
that contains a transducer-specific interface. This circuitry includes all the components that link the transducer elements to
the processing channels. Most of the front-end electronics is
for transmit and receive focusing. As noted above, the beam of
a flat transducer converges to a natural focus. While this focus
can be modified by acoustic lenses, true variable adjustment of
the depth of the focal zone can be achieved only by electronic
means. In all types of arrays, focusing in the scan plane is accomplishing by means of variable electronic delay circuitry. Focusing can be done in both the transmit and receive modes
(Fig. 1.
6).
Receive focusing. In receive focusing, the echo signals that return to the array first are delayed by the difference in arrival
times from the more peripheral signals. The goal is to ensure a
coherent display of all signals that are received from the variable focal point.
Transmit focusing. The transmitted beam can be focused by
delaying the excitation pulses in the array. Multizone focusing
can be achieved only by scanning sequentially at different focal
depths, since the focus of the wavefront can no longer be
changed after the pulse has left the array. While transmit focusing allows the user to define the number and location of the
focal zones, multizone focusing results in a slower frame rate.
For example, three focal zones triple the time required to
generate an image, and so the frame rate is reduced by a factor
of 3.
Dynamic focusing. Other focusing possibilities are available in
the receive mode. By varying the delay times during reception
´
´
´
of the echo signals, it is possible to keep the focus at the depth
from which echoes are momentarily being received. This
process, called dynamic focusing, does not reduce the frame
rate and operates automatically throughout echo reception,
without user input, according to an optimized algorithm. The
electronic focusing described above is a feature of all electronic
transducers.
Time gain compensation. Preprocessing refers to the part of
the signal processing circuit in which scan data are manipulated before they reach the scan converter. It begins with the
removal of unwanted effects caused by the sound attenuation
and center frequency changes that occur with increasing
depth. Echoes from deeper objects are attenuated more than
shallower echoes on their way to the transducer. Time gain
compensation (TGC) is a preprocessing technique that compensates for this attenuation by amplifying the echoes from
B-Mode Sonography
more distant reflectors as a function of their depth. Otherwise
the echo signals from deeper sources would appear increasingly dark. Generally the TGC is set manually, because “homogeneous image appearance” and “region of interest” are not
quantities that can be preprogrammed. One way to optimize
the setting is to take a subcostal transverse scan of the liver, for
example, and then adjust the TGC curve until the brightness of
the liver appears homogeneous. If scanning must be done
through fluid (e.g., ascites) or through the uterus, the TGC
should be reduced accordingly to avoid tissue effects of over-
whelming brightness.
Dynamic frequency filtering. Another preprocessing step is dynamic frequency filtering. The frequency-dependent attenuation of the sound, with a depth-dependent spectral shift
toward lower frequencies, makes this a necessary step. Concurrent dynamic frequency filtering compensates for this shift
effect and improves the signal-to-noise ratio in deeper regions.
Multifrequency and broadband technique. For these techniques to provide useful image improvement, the transducer
must have the capacity for broadband excitation and broadband signal processing. Other advantages of broadband transducers are a shortened pulse length for better axial resolution
(see Fig. 1.
this way a single transducer can be used to image superficial
structures at high resolution as well as structures located at
greater depths. The multifrequency technique and broadband
technique significantly expand the performance range of
transducers and simplify the examination.
Edge enhancement. Other steps are demodulation and lowpass filtering and the ability to modify the shape of the echo
signal in preprocessing. Edge enhancement, for example, can
be used to steepen the slope of a change in signal level, thus
providing better differentiation of echoes across an interface.
(Edge enhancement is accomplished by high-pass filtering of
the demodulated echo signals and is applied only to line-ofsight data.) Tissues appear to have a more fine-grained texture,
and interfaces perpendicular to the beam direction (e.g., vessel
wall boundaries) are more clearly defined. The effect, then, is
an apparent improvement of axial resolution.
Dynamic range. The dynamic range makes it possible to match
the contrast level of the ultrasound image to diagnostic requirements. From the large range of echo intensities that can
be handled by modern systems, the user can select the dynamic range from the smallest to the brightest visible B-mode
echoes that appears best for the current application. A smaller
dynamic range is useful for defining contours, measuring the
cranial biparietal diameter, and vascular imaging. A higher dynamic range is useful for evaluating the structure of tissues like
the placenta, as this will define even the faintest tissue echoes.
Scan converter. The function of the scan converter is to convert
the data acquired by various scanning techniques (e.g., in polar
coordinates) into a correct spatial display pattern (e.g., in cartesian coordinates) on the viewing screen. Because data readout from the scan converter follows a television format (video)
or computer format (e.g., SVGA), the ultrasound information
4) and the capacity for multifrequency excitation. In
requires intermediate (digital) storage on or before reaching
the scan converter.
Correlation of the B-mode image. “Correlation” of the B-mode
image is an averaging process that is carried out in the image
memory. The time-averaging process can be structured in such
a way that when new information is entered into memory, a
portion of the stored (old) information from the previous
image is retained in memory, and the new value is added to it.
The effect is similar to “afterglow” (persistence) on the monitor
screen. Correlation can be useful in slowly changing image
sequences (e.g., in the abdomen, peripheral vessels, etc.),
where it improves the signal-to-noise ratio. With fast-moving
processes (e.g., fetal cardiac imaging), setting the correlation
too high can cause a smearing of structures with loss of image
quality.
Postprocessing. In postprocessing, the gray-scale depiction of
the image information on the monitor can be manipulated by
altering the data transfer characteristics between the scan con-
verter and the monitor. This may be done, for example, to enhance or suppress low-level echoes in the final image.
Analysis of B-Mode Information and Artifacts
Optimizing the Examination and Analyzing the
B-Mode Information
Homogeneous display of echo texture. After selecting the
transducer that is appropriate for the desired information and
the region under study (see Transducer Selection: Penetration
Depth, Beam Pattern, and Resolution, p. 5), the user adjusts the
transmission power output, focal depth, receiver gain, and TGC
curve to obtain a homogeneous display of echo texture within
the area of interest. It is good practice to adjust these parameters in, say, a subcostal oblique scan of the liver. If fluid is present in the near field (e.g., ascites), the TGC should be reduced to
avoid effects of too much brightness.
The principal criteria for evaluating organs such as the liver,
spleen, kidneys, uterus, and ovaries are size, shape, delineation, and changes in the basic echo pattern. The change may be
hypoechoic, hyperechoic, or isoechoic. These pattern changes
are helpful in detecting lesions such as hepatic, uterine and
ovarian tumors or morphological changes in the placenta.
Artifacts and Pitfalls
Ultrasound image formation is based on certain physical assumptions, some idealized, such as those of constant sound
velocity and attenuation or the assumption of straight-line
sound propagation in the body. Deviations from these assumptions lead to artifacts. Some of these phenomena are diagnostically useful and furnish additional information on tissue properties, while others can lead to misinterpretation.
“Echo enhancement.“ Acoustic enhancement behind cysts is a
diagnostically useful artifact. With the assumption of constant
attenuation, the TGC curve produces a depth-dependent signal
amplification in order to achieve uniform echo amplitudes
Physical and Technical Principles
7
Physical and Technical Principles of Color Doppler Sonography
from different depths (see Signal Processing: Focusing, Preprocessing, and Postprocessing, p. 6). However, the low degree of
absorption and reflection in fluids causes too much signal gain,
resulting in enhancement of echo brightness behind the fluid
(Fig. 1.
7).
“Shadowing by a strong reflector.“ Another potentially useful
artifact is the distal shadowing that occurs behind strong reflectors. It may be caused by substances such as bone, stones,
and air, which have an acoustic impedance very different from
that of adjacent soft tissues (Table 1.
impedance mismatch, much of the ultrasound energy is reflected back toward the transducer. If the “reflector” is insonated at an approximate right angle, it produces a very
bright echo (Fig. 1.
the impression of a renal stone or similar lesion. However, if
the reflector is insonated at a very oblique angle relative to the
scan plane, most of the energy is reflected away.Shadowing occurs, but the reflector is not represented by a bright echo in the
B-mode image.
Other shadowing effects are seen at the edge of circular objects. If the beam intercepts an oblique interface at a tangential
angle, deflection of the beam may occur, producing a distal
1
acoustic shadow. The interfaces are only partially defined.
Reverberations. Reverberations are artifacts caused by multiple reflections between two highly reflective interfaces. The
prolonged transit time leads to faulty spatial encoding, creating a cascade of parallel echo bands spaced at uniform intervals. Reverberation artifacts can be eliminated by insonating
the reflecting surface at an oblique angle.
Comet-tail artifact. A comet-tail or resonance artifact is also
caused by multiple reflections. When reflective interfaces are
located very close together with weakly attenuating tissue be-
7). The associated distal shadow reinforces
1). Because of this large
tween them, they will generate many closely spaced echoes
that form a long, bright stripe located distal to the actual reflector. This can occur with minute air bubbles, for example, or
when structures with a high sound velocity, such as cholesterol
crystals or metallic foreign bodies, are encountered by the
beam, setting up multiple internal reflections.
Mirror-image artifacts. Another special case is the mirrorimage artifact. When an object is located in front of a strong reflector, it will be “insonated” by that reflective interface. The
returning echo is reflected back again and, because of its prolonged transit time, forms a second, virtual image of the object
behind the strong reflector (Fig. 1.
Slice-thickness artifacts. Slice-thickness artifacts appear as
areas situated at right angles to the scan plane that appear to
fill in the echo-free space within small, fluid-filled objects. The
artifact can be eliminated by changing the beam direction or
using a different probe.
Side-lobe artifacts. Signals from reflectors located outside the
scan plane can produce “ghost echoes” due to side-lobe effects.
Side lobes are sound pressure peaks that are located outside
the main beam and are directed obliquely outward from it.
These artifacts have become rare, however, as a result of refinements in equipment.
Refraction and diffraction artifacts. Refraction and diffraction
artifacts occur at interfaces between media that have very
different ultrasound velocities. An example is the interface between bone (higher sound velocity) and silicone or other plastics (lower sound velocity).
7).
Fig. 1.7 B-mode artifacts: examples from abdominal
imaging.
a Hepatic cyst with posterior acoustic enhancement
due to overamplification of the ultrasound signals.
b Gallbladder with a stone (bright echo with a posterior acoustic shadow) and edge shadows (arrows).
c Comet-tail artifact caused by multiple reflections
from air.
d Hepatic calcification with a mirror-image artifact
projected on the opposite side of the diaphragm.
a
b
8
c
d
Duplex and Color Doppler Sonography
Physical Principles of Motion Detection
Duplex and Color Doppler Sonography
The techniques of Doppler shift acquisition are classified as
one-dimensional or two-dimensional.
Doppler effect. The oldest approach to motion detection is the
Doppler technique. It is based on the Doppler effect, first described mathematically by the physicist Christian Johann
Doppler (1803–1853). He observed that the light emitted from
stars that were moving toward the earth was shifted toward
the blue end of the spectrum, i.e., toward higher frequencies
and shorter wavelengths (“blue shift”). Conversely, a red shift
was observed when the earth and star were moving apart. The
same principle occurs in acoustics. The sound of an automobile
engine changes its pitch when it moves relative to an observer.
The change is most noticeable at the moment when the vehicle
is passing the observer.
Frequency shift. When the Doppler effect is applied in medical
diagnostics, the moving objects are red blood cells that are
moving toward or away from the transducer at various velocities (Fig.1.
from the moving blood cells show a frequency shift of
8). The echo signals reflected back to the transducer
∆f rela-
tive to the original transmitted frequency f. This shift depends
on the magnitude and direction of the blood flow velocity and
is described mathematically by the equation (6):
∆f =(2fv cos Θ)/c (6)
where
transmitted frequency (f) and the frequency of the reflected
signal. It is proportional to the velocity v of the blood cells and
to the cosine of the angle
beam and the vessel axis (Fig. 1.
to the sound velocity in the medium being scanned (approximately 1540m/s) and is usually within the audible range of
500–20,000 Hz in medical examinations. The factor 2 in the
equation relates to the fact that the Doppler effect is operative
twice in the pulse-echo technique: once when the pulse travels
from the transducer to the moving blood cells, and again when
the reflected signal returns to the receiver.
∆f represents the detectable Doppler shift between the
Θ between the Doppler sampling
8). It is inversely proportional
One-dimensional or spectral techniques. In the one-dimensional techniques, the vessel is scanned by one beam and the
flow velocities are analyzed along the path of that beam. The
information acquired is displayed as a time-dependent Doppler spectrum. The spectral sampling region in one-dimensional
techniques may cover the full depth of the beam (continuous-
wave or CW Doppler) or only a specific, selected site along the
path of the b eam (pulsed-wave or PW Doppler) (Fig. 1.
9).
Two-dimensional techniques. Two-dimensional techniques
yield a spatial distribution of Doppler information that is acquired from multiple sampling sites. Multiple beams are transmitted, the received signals are selectively analyzed according
to their depth, and the result is superimposed on the B-mode
image in color-encoded form. The size of the analyzed area
may cover the entire B-mode image, or it may be confined to a
window called the “color box.“
The image may depict the mean velocity in the volume element and the flow direction (color duplex sonography, CDS),
or it may represent the sum of the squares of the amplitudes as
a measure of the total number of moving red cells or the intensity of the blood flow (power Doppler imaging). Other techniques are based on the velocity-related position change in
characteristic echoes (time-domain correlation or CVI)
1
or on
the digital amplitude subtraction of two pulses emitted at
different times (bi-flow technique)
18
. The latter technique
yields a display with high spatial resolution, similar to digital
subtraction angiography. At present, both of these techniques
can be used only for superficial scanning under favorable insonation conditions. They have little or no practical importance in gynecology, so they need not be discussed here.
Technical Principles and Equipment Settings
One-Dimensional Techniques of Motion Detection
(CW and PW Doppler)
Physical and Technical Principles
Transducer
Θ
Fig. 1.8 Principle of Doppler sonography.
v
Doppler effect
f=2(f/c)v cos
∆
f=Doppler frequency
∆
f=transmission frequency
c=sound velocity
v=flow velocity
=beam-vessel angle
Θ
CW Doppler. Continuous wave(CW) Doppler employstwo separate crystals, one of which continuously transmits sound
waves while the other continuously receives the reflected
echoes (Fig. 1.
9a). Because of its continuous operation, CW
Doppler cannot provide depth discrimination or spatial localization of an echo. The advantage of the technique is that even
Θ
very high flow velocities can be definitively analyzed (see
below).
PW Doppler. Pulsed wave (PW) Doppler makes it possible to
sample flow velocities of selected areas. The same crystals are
used for transmitting and receiving, and short pulses are transmitted into the body as in B-mode imaging. A designated time
interval is allowed for the pulse to travel to the desired Doppler
9
sampling site and return to the transducer, at which point the
gate for echo reception is briefly opened (Fig. 1.
9b). The ex-
Physical and Technical Principles of Color Doppler Sonography
1
a. CW Doppler
TM: Transmitter
R: Receiver
Transducer
TM
b. PW Doppler
Transmit time: T
Reception time: t
Transducer
Amplitude
Time
R
R
T
t
R
Transmit
T
Receive
PRF
2T t
max
1
=
T
Amplitude
Time
Fig. 1.10 Principle of the fast Fourier transform (FFT). Just as the
human ear is able to break down a musical chord into its separate
notes (e.g., G, B, and D), the FFT is designed to resolve a frequency
Amplitude
++
Time
Amplitude
Frequency
Amplitude
Time
spectrum into its constituent frequencies. The result is displayed as an
t
R
t
amplitude–frequency distribution (bottom graph).
10
Fig. 1.9 One-dimensional techniques: principles of CW and PW
Doppler.
a CW Doppler. The transducer contains two crystals, each of which
transmits and receives continuously. This instrument can detect even
high Doppler frequency shifts with no aliasing or other artifacts. The
disadvantage is that all the vessels insonated by the beam yield one
mixed signal, makingit impossible to assign thesignal to a specific ves-
sel.
b PW Doppler. The same piezoelectric crystals are used to transmit
and receive, and are driven in pulses. The echoes are received at a
selected time interval (range gate), which depends on the transit time
T that the signal takes to reach the transducer from the desired depth.
The reception time t
determines the size of the sample volume that
R
will be analyzed.
aminer selects the size and depth of the range gate on the basis
of visual guidance in the B-mode or color Doppler image.
Demodulation. The first step in obtaining velocity information
from the echo signal is demodulation, which separates the
Doppler signal (Doppler frequency
∆f) from the reference or
transmitted frequency. Demodulation is carried out in a phase
detector by combining the echo signal with a reference signal,
followed by rectification and low-pass filtering.
Spectral analysis. The echo returning from the blood to the
transducer contains a mixture of frequencies that depend on
the velocity distribution of the individual blood cells. In the
next step, therefore, the frequencies contained in the Doppler
signal are analyzed according to their distribution and amplitudes. This spectral analysis is accomplished by a mathematical algorithm called the fast Fourier transform (FFT). Just as the
human ear can resolve a musical chord into its constituent
notes (Fig. 1.
10), the FFT processor can break down the de-
modulated signal into its underlying (harmonic) Doppler components and their amplitudes.
The frequency resolution and temporal resolution of the
spectrum depend on the number of flow samples (generally
between 64 and 256) that are used in computing the FFT algorithm. The greater the number of samples that go into the computation, the higher the frequency resolution. With 128-point
FFT, the information from 128 consecutive Doppler signals
(transmitted pulses) is broken down into the underlying
frequencies. With the next transmitted pulse, the calculation
results in a consecutive complete spectrum with 128 frequencies.
Doppler spectrum. The result of the analysis is displayed either
as an amplitude–frequency curve for a given point in time
(Fig. 1.
10) or as a time-varying frequency spectrum (Doppler
spectrum) that covers one or more cardiac cycles. In the latter
display,which is commonly used in medicine, the amplitude of
the Doppler frequencies at each point in time is represented by
the brightness of the pixels (Fig. 1.
11).
The Doppler spectrum forms the basis for a more detailed
analysis, making it possible to determine and quantitate flow
velocities and volumes and also to calculate resistance values
that can demonstrate prestenotic or poststenotic changes in
the vascular bed (see Analysis of Doppler Information and Artifacts, p. 19).
Pulse repetition frequency. One disadvantage of pulsed Doppler flowmetry is that the transit time of the pulses (T) in the abdomen ranges from approximately 0.03 to 0.26 ms, depending
on the scanning depth (e.g. 3–20 cm). This limits the shortest
possible time interval between two successive pulse transmis-
Duplex and Color Doppler Sonography
Fig. 1.11 Time-varying frequency spectrum (Doppler spectrum). The
spectrum was sampled from an artery over two cardiac cycles and
plotted by FFT analysis. The horizontal axis is time, and the vertical axis
is the Doppler frequencyshift or, after angle correction,the flow velocity. The brightness of the image points in the spectrum is proportional
to the amplitude of the particular Doppler shift. In the first cardiac
cycle the peak velocities (bright line at the top of the trace) and the
mean velocities (dark line at the center) are shown as curves overtime
course.
Relevant instrument settings:
Doppler frequency: 5.5 MHz
Pulse repetition frequency: 520 Hz
Wall filter: 50 Hz
Angle correction: 63⬚
sequent scan at the same location (Fig.1.
12). The result is a cor-
relation vector that is characterized by its amplitude and phase
difference. The phase difference at each image point in the
color line supplies the mean frequency shift, and this value is
directly proportional to the mean flow velocity. The variance
can also be determined and displayed to indicate turbulence.
Kasai and Nameka first used the mathematical process of
autocorrelation to calculate blood flow in 1982.An autocorrelation A(t) of the function g(t) is described by
+⬁
A(t)=
a
RF echo Digitized signal I/Q v
g′(τ) g(τ + t) dt (8)
∫
–⬁
Processing of demodulated signals
Quadrature
phase
detector
Flow analyzer
Fast Fourier transform
Autocorrelation
Autoregression
equation (8):
Flow
ana-
lyzer
Image
me-
mory
Physical and Technical Principles
sions. Hence, the rate at which the separate pulses are transmitted (pulse repetition frequency, PRF, equation 7)
cannot
exceed 1/T (e.g., 25.6 or 3.8 kHz) without causing errors of
depth discrimination.
PRF =
1
T
c
=
(7)
2s
In equation (7), s is the depth (m), PRF is the pulse repetition
frequency (Hz), and c is the sound velocity (1540 m/s).
Two-Dimensional Techniques of Motion Detection
(Color and Power Doppler)
Color Duplex Sonography
Whereas pulsed spectral Doppler displays the velocity distribution over time at a designated sampling site, color duplex
scanning analyzes the flow velocities at a number of sites that
are distributed over all or part of the sectional image. Colorflow imaging, unlike spectral Doppler, uses autocorrelation
rather than Fourier analysis.
Autocorrelation. Recall that in FFT analysis each Doppler signal
is sampled from the gate up to 256 times for each time interval
in the spectrum, and the corresponding number of frequencies
are calculated for that location. By contrast, autocorrelation in
color Doppler imaging compares (correlates) all the demodulated Doppler signals along a color scan line that are sampled at
a given time interval after transmission with those of the sub-
b
Autocorrelation method
I(Q)
T
=1/PRF
Time
1. First scan
sequence
2. Second scan
sequence
2 ... 16
Sampling
pairs
∆Θ
∆Θ
1
2
Mean
phase
difference
≈
∆Θ
mean
flow
velocity v
Fig. 1.12 Two-dimensional Doppler techniques. Principle of signal
analysis (illustrated for autocorrelation).
a Common principle of signal processing in all autocorrelation, autoregression, and FFT methods. The high-frequency signal is demodulated in the quadrature phase detector. This demodulation is done
twice, with a phase shift of 90⬚ (I, Q). After it has been filtered by suita-
ble wall filters, the signal (which is digitized in modern systems) is fed
to the actual analyzer. A fast Fourier transform, autocorrelation, or au-
toregression is carried out, depending on the method used. The result
is usually stored as color information in image memory.
b In the most commonly used autocorrelation method, pairs of successive samplings are tested for their phase difference in a flow analyzer. This difference is a measure of the flow velocity v. Time averaging the phase differences in the selected scan sequence yields the
mean velocity v, which is color-encoded and superimposed on the Bmode image. (After Haerten 1993.
5
)
11