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Файл:Ординатура / Хирургия / Библиотека им академика М.И. Перельмана / Книга_5760_Библиотеки_им_академика_М_И_Перельмана.pdf
1.1.2.4 Blood Flow Measurement–17
1.1.3 Physical Principles ofColor-Coded Duplex Ultrasound–20
1.1.3.1 Velocity Mode–20
1.1.3.2 Power Doppler Mode–23
1.1.3.3 B-Flow Mode (Brightness Flow)–24
1.1.3.4 Intravascular Ultrasound–25
1.1.3.5 Three-Dimensional/Four-Dimensional Ultrasound–26
1.1.4 Factors Aecting (Color) Duplex Imaging– Pitfalls–26
1.1.4.1 Scattering andAcoustic Shadowing–26
1.1.4.2 Mirror Artifact–26
1.1.4.3 Maximum Flow Velocity Detectable– Pulse Repetition
Frequency–26
1.1.4.4 Minimum Flow Velocity Detectable– Wall Filter, Frame Rate–30
1.1.4.5 Transmit andReceive Gain–30
1.1.4.6 Doppler Angle–32
1.1.4.7 Physical Limitations ofColor Duplex Ultrasound–32
1.1.5 Ultrasound Contrast Agents–33
1.1.5.1 Approved Ultrasound Contrast Agents andUses–33
1.1.5.2 Mechanisms ofAction–34
1.1.5.3 Ultrasound Techniques Using Contrast Agents–35
1.1.5.3.1 Contrast-Enhanced Duplex Ultrasound–35
1.1.5.3.2 Contrast Harmonic Imaging–35
1.1.5.3.3 Stimulated Acoustic Emission Imaging–35
1.1.5.4 Summary ofTechnical Aspects andClinical Indications–35
1.1.6 Safety ofDiagnostic Ultrasound–36
1.1.6.1 Thermal Eects–36
1.1.6.2 Mechanical Eects–36
1.1.6.3 Specic Risks ofIndividual Ultrasound Techniques–36
1.1.6.3.1 B-Mode–36
1.1.6.3.2 M-Mode–36
1.1.6.3.3 CW Doppler–36
1.1.6.3.4 PW Doppler–37
1.1.6.3.5 Color Doppler–37
1.1.6.4 Conclusion–37
1.2 Hemodynamic Principles–37
1.2.1 Laminar Flow–37
1.2.2 Flow Proles andPerfusion Regulation–40
1.2.2.1 Low-Resistance Flow–40
1.2.2.2 High-Resistance Flow–40
1.2.2.3 Perfusion Regulation–42
1.2.3 Stenosis Grading andBlood Flow Measurement–42
1.2.3.1 Poststenotic Parameters–47
1.2.3.1.1 Acceleration Time– Resistive Index–47
1.3 Machine Settings–47
1.1 · Technical Principles ofDiagnostic Ultrasound
1.1 Technical Principles ofDiagnostic
Ultrasound
1.1.1 Gray-Scale Ultrasonography (B-Mode)
1.1.1.1 Historical Milestones
e potential for using the reection of ultrasound in the
visualization of the internal organs of the human body was
recognized about 80years ago. e rst attempts at using
ultrasound in medical diagnosis were made in the late 1930s
by the Austrian neurologist K.T. Dussik. He developed
what he referred to as hyperphonography, a sonographic
transmission technique for the visualization of the cerebral
ventricles. Also in the 1940s, American scientists began
experimenting with ultrasound reection to examine biological objects. Among the early pioneers were Ludwig and
Struthers, who used this new technique for detecting gallstones. Other important milestones in the history of diagnostic ultrasound were the development of B-mode imaging
by Howry and Bliss and the introduction of the echo pulse
method by Leksell in Sweden, which he used to determine
the position of midline brain structures in the intact skull,
thus marking the start of echoencephalography. In 1954,
Edler and Herz presented the rst description of M-mode
echocardiography.
e Japanese physicist Satomura is credited with implementing the rst medical applications of the Doppler principle. He and his colleagues investigated the use of Doppler
frequency shis to evaluate moving cardiac structures and to
measure the velocity of red blood cells. e advent of the rst
real- time scanner, developed by Krause and Soldner, completely changed the practice of medical ultrasound scanning
and marks yet another important step in the success story of
diagnostic ultrasound.
Modern ultrasound oers excellent image quality and
diagnostic capabilities, with its outstanding position among
radiologic imaging techniques being due to its versatility, low
cost, exibility, and safety.
is chapter introduces the physical and technical fundamentals of medical ultrasound and outlines the range of
techniques available today, which will help readers to make
optimal use of the diagnostic capabilities of ultrasound and
choose the best technique for the intended application.
1.1.1.2 Sound Waves
When a molecule is activated to vibrate around its equilibrium position, the vibration is transmitted to a neighbor in
the medium and from there to the next molecule and so on.
In this way, kinetic energy is propagated from one molecule
to the next, spreading through the medium in a sine wave
pattern. is pattern of the spreading of kinetic energy is
known as a continuous wave or an acoustic wave (sound
wave). A sound wave alternately compresses (positive pressure) and expands (negative pressure) the medium it travels
through (. Fig. 1.1). Particles can vibrate parallel or perpendicular to the direction of energy propagation, giving
3
Elongation
Compression Expansion
. Fig. 1.1 Diagram of the propagation of a longitudinal wave illus-
trating cyclic compression and expansion (Courtesy of Hitachi Ltd.,
which also provided the historical material presented in 7 Sect. 1.1.1.1)
. Table 1.1 Typical sound velocities, densities, and attenu-
ation values in some important biological tissues and other
media in the body
Medium Sound
velocity (m/s)
Fat 1470 0.97 0.5
Bone marrow 1700 0.97 –
Muscle 1568 1.04 2
Liver 1540 1.055 0.7
Brain 1530 1.02 1
Bone (compact) 3600 1.7 4–10
Water (20°C) 1492 0.9982 0.002
Air 331 0.0013 –
Density
(g/cm2)
Attenuation
(dB/MHz cm)
rise to longitudinal waves (along the direction of travel) and
transverse waves (perpendicular to the direction of travel).
Particles excited in the ultrasound range vibrate around their
resting positions at a rate of 20,000 to one billion times per
second.
In gases and liquids, only longitudinal wave propagation
is possible, as the shear forces necessary for the spread of
transverse vibration are absent. In physical terms, biological
tissues can be viewed as viscous uids, which is why the eect
of transverse waves is negligible. In such a medium the speed
of sound increases with density, which in turn is dened by
the force of molecular cohesion (
. Table 1.1). e average
speed of sound in biological tissues is approx. 1540m/s.
Waves can be described with reference to several properties. Wavelength λ is the distance between two consecutive
1
4
Cf=×
l
Chapter 1 · Fundamental Principles
1
. Table 1.2 Commonly used transmit frequencies and result-
ing properties of the ultrasound beam
Transmit
frequency
(MHz)
2 0.78 25 3 0.8
3.5 0.44 14 1.7 0.5
5 0.31 10 1.2 0.35
7.5 0.21 6.7 0.8 0.25
10 0.16 5 0.6 0.2
15 0.1 3.3 0.4 0.15
Wave length
(mm)
Penetration
depth
(cm)
Lateral
resolution
(mm)
Axial
resolution
(mm)
on one side and positive on the other. As the degree of stress
increases, so does the voltage. Conversely, when a positive
or negative voltage is applied to the surface of a piezoelectric crystal, the material expands or contracts, depending on
the direction of the current. When an alternating current is
applied, the piezoelectric crystal is activated and begins to
vibrate. Materials possessing strong piezoelectric properties
are quartz and tourmaline. State-of-the-art transducers use
semicrystalline polymers such as polyvinylidene uoride
(PVDF).
1.1.1.4 Physical Factors Aecting
theUltrasound Scan
An ultrasound image is created by processing the echoes
returning to the transducer from various depths of the
body upon emission of an ultrasound pulse of a specic
The following relationships exist between these parameters:
the higher the transmit frequency (and therefore the shorter
the wavelength), the higher the resolution– but the lower the
penetration depth
frequency (. Fig.1.2a). A two-dimensional (2D) image is
generated from adjacent ultrasound lines. Two-dimensional
morphologic images are acquired by applying short pulses
of energy using only a small number of wavelengths to
optimize spatial resolution. e round trip time is the time
delay between the emission of an ultrasound pulse and the
. Table 1.3 Parameters dening a sound wave
Property Denition
return of the reected echo and is a function of the distance
between the transducer and reector. Reection occurs at
the boundaries between media that dier in their sound
propagation properties, or acoustic impedance. Hence,
Period Duration of a complete vibration
Wavelength Spatial extension of a period
Frequency Number of periods per second
Amplitude Measure of sound energy
an ultrasound image does not represent tissue structures
directly but rather interfaces between tissues of dierent
acoustic impedance.
Acoustic impedance describes the
frequency-dependent resistance that an ultrasound beam
encounters as it passes through a tissue. It is equal to the
speed of sound propagation multiplied by the density of the
tissue. e greater the dierence in impedance, the greater
points of maximum compression, and frequency f is the
number of vibrations of a molecule per unit time, given in
hertz (Hz). One hertz corresponds to one cycle per second,
or 1Hz=1/s. e frequency range of diagnostic ultrasound
is 2–30MHz. e speed of a sound wave, C, is the product of
wavelength and frequency:
the reection of the ultrasound wave (and therefore the
greater the strength of the echo or signal) and the smaller its
transmission into deeper tissue (. Fig.1.2a). Other physical processes besides reection and scattering that aect
the ultrasound scan are refraction, interference, diraction,
attenuation, and absorption.
e wavelengths occurring in diagnostic ultrasound are
determined by the frequency emitted by the transducer (carrier frequency) and range from 0.78 to 0.15mm over the
2–10MHz frequency range typically used in vascular imaging (. Table1.2). e properties dening a sound wave are
summarized in . Table1.3.
1.1.1.3 Generating Ultrasound Waves
In most ultrasonic transducers for medical imaging, the
piezoelectric eect discovered by Pierre and Jacques Curie
in 1880 is used to generate ultrasound waves. When mechanical stress is applied to piezoelectric materials such as ionic
crystals, they experience an elastic deformation which results
in a shi in internal charge distribution. In this way, electric
voltages are generated at the surfaces– which are negative
1.1.1.4.1 Reection andRefraction
e propagation of sound waves in biological tissues is governed by the laws of wave optics. Tissues vary in density and
hence dier in acoustic impedance. Impedance Z is the product of the density of a medium and the speed of sound in it.
At an acoustic interface in the body, an incident ultrasound
beam is partially reected and partially refracted. Refraction
means that the wave passes through the interface, changing
its direction of travel (. Fig.1.2b). e dierence in acoustic impedance between the two tissues forming the interface
determines how much of the beam is reected and how
much is transmitted: the greater the dierence, the greater
the amount of energy that is reected back; the smaller the
dierence, the greater the amount of energy that is transmitted. Medical ultrasound thus functions like a sonar, exploiting dierences in acoustic impedance between two adjacent
tissues rather than absolute acoustic properties.
è
ø
12
12
1.1 · Technical Principles ofDiagnostic Ultrasound
Beam perpendicular to interface
Reflection
a'
5
Interface
Refraction
a'
1
Reflection
Stronger when the impedance
mismatch is large
Sonar and medical ultrasound rely on:
difference in acoustic impedance between two
tissues/media
a
. Fig. 1.2 a Generation of an ultrasound image: reection– transmission. b Interaction of ultrasound with interfaces in the body according to
the laws of wave optics (for details see text) (Courtesy of Hitachi Ltd.)
Stronger when the impedance
mismatch is small
e reection gradient, R, is given by the following equa-
tion for incident angles perpendicular to an interface:
2
-
ö
÷
+
R
ZZ
æ
=
ç
ZZ
Transmission
Impedance mismatch
Z1 Z2
b
scattered when it strikes an object that is much smaller than
its wavelength, and it is reected when it strikes an object
much larger than its wavelength. Scattering gives rise to the
characteristic echotexture of parenchymal organs in ultrasound images.
Since structures perpendicular to the beam are rare in
For an ultrasound beam striking the interface between liver
tissue (Z1= 1.66 ×105) and renal tissue (Z2= 1.63× 105),
the equation yields a reection gradient of R= 0.000008,
meaning that this boundary reects less than one hundred
thousandth of the incident energy. In contrast, nearly all of
the incident energy (over 99%) is reected from the inter-
5
face between fatty tissue and air (Z1=1.42× 10
, Z2=43,
R=0.9987), leaving virtually no ultrasound energy to travel
deeper into the tissue. is is why the lungs or bowel loops
containing air cannot be examined by ultrasonography and
also why it is necessary to eliminate air intervening between
the ultrasound probe and the skin surface by applying ultrasound gel.
e echoes reected back from an interface between
media of dierent acoustic impedance are available for image
generation only if the interface is relatively perpendicular to
the ultrasound beam (angles of incident and reected beam).
For this reason, structures such as vessel walls perpendicular
to the beam appear fairly bright compared to vessel walls tangential to the beam since most echo pulses are reected back
to the transducer by the former. Reection occurs at the surfaces of particles that are larger than the wavelength, while
scattering predominates when they are smaller.
clinical ultrasound examinations, an ultrasound image is
chiey generated from a mixture of reected and scattered
echoes. Aggregations of tissue cells scatter the beam diusely
in all directions. erefore, a structure appears bright and is
clearly dened when it is perpendicular to the ultrasound
beam because the image information is mainly derived from
reected echoes; its visualization is weaker and less bright
when the ultrasound beam strikes tangentially and only diffusely reected echoes are available to generate the image,
although impedance is identical in both cases.
Scattering contributes to the attenuation (loss of energy) of
the ultrasound beam as it travels through the body and in turn
depends on the transmitted frequency. A higher transmit frequency results in greater attenuation and limits the penetration
depth of the ultrasound pulse. e emitted intensity decreases
exponentially with distance and is inuenced by an attenuation
coecient that varies with the type of tissue through which the
beam travels in the human body (fat, muscle, blood). In the
human body, it ranges from 0.3 to 0.6dB/MHz cm. e energy
is converted into absorption heat.
Higher carrier frequencies result in a lower penetration
depth because attenuation loss is greater. e increasing
attenuation can be compensated for to some extent by adjusting amplication (depth-dependent gain) (. Fig. 1.3b).
1.1.1.4.2 Scattering andAttenuation
e interface between tissues of dierent acoustic impedance
is typically not smooth but rough. A sound wave interacting with a rough surface will be scattered in all directions
in the form of a spherical wave rather than along one path
(. Fig. 1.3a). An incident ultrasound wave is also mostly
Using transducers with a wide frequency range results in the
predominance of lower frequencies with greater penetration
depths because attenuation of higher frequencies is more
pronounced.
In addition to scattering and reection, there is refraction
at the interface between dierent media. Refraction in the
6
zct= /2
Chapter 1 · Fundamental Principles
1
. Fig. 1.3a, b Scattering and attenuation of sound waves. a Scattering: Most ultrasound beams do not strike reecting structures in the body at
a right angle, which is why the incident beam is scattered in all directions. As a result, only a small proportion of the emitted energy is backscattered to the transducer and available for generating the ultrasound image. An ultrasound beam reected from an interface between two tissues
with the same dierence in acoustic impedance will yield much stronger echoes than a beam scattered at that interface (resulting in poorer visualization) (Modied from Widder and Görtler 2004). b Attenuation reduces the amplitude of the reected ultrasound beam with echoes returning
from structures deeper in the body being attenuated more strongly. To create a uniform image from all signals despite their dierent amplitudes,
time gain compensation (TGC) is used, which changes the receive gain over time, applying greater amplication to echoes returing from deeper
in the body (using a set of sliding knobs or paddles)
direction of the normal to the interface occurs when there
is an increase in sound velocity in the next medium, and
refraction away from the normal occurs when the velocity
decreases. Refraction may lead to misinterpretation of the
location and size of the structure visualized.
Time Gain Compensation
Attenuation
50% 100%
2
4
8
ba
Amplification
0
2
4
8
Noise
Depth [cm]Depth [cm]
60
dB
1dB/mHz cm. e attenuation values for a selection of biological tissues are given in
. Table1.1. e rate of absorption
depends not only on the tissue type but also on the emitted
ultrasound frequency, with higher frequencies attenuating
more quickly. Lower ultrasound frequencies, with long wave-
13.3
26.6
53.3
Time
[µs]
lengths, thus allow the examination of deeper structures,
1.1.1.4.3 Interference
When two or more sound waves superimpose, they can be
out of phase (i.e., one wave’s compression phase coincides
with the other’s expansion phase), thus cancelling each other
out (destructive interference), or they can be in phase (i.e.,
the compression and expansion phases line up), thus reinforcing each other (constructive interference). e spatial
while high ultrasound frequencies are desirable for the better
spatial resolution they aord. For an ultrasound frequency of
10MHz, for instance, the attenuation is 10dB/cm as opposed
to only 3dB/cm for 3MHz. Assuming an output of 100dB,
the penetration depth would be 5cm for 10MHz and 17cm
for 3MHz (corresponding to a total path length of 10 and
34cm, respectively).
distribution of areas of constructive and destructive interference is known as the interference pattern. Such interference
1.1.1.5 Generating anUltrasound Image
patterns are largely responsible for the visual appearance of
an ultrasound image.
Interferences of sound waves can change the amplitude
and thus the brightness of an image despite an identical
acoustic impedance in the boundary zone. Depending on the
momentary phase of the wave, the amplitude is either amplied or diminished.
1.1.1.5.1 Pulse-Echo Technique
Nearly all diagnostic ultrasound techniques rely on pulsed
excitation signals. An ultrasound beam is generated by
applying short electrical pulses of about 1s to the piezoelectric crystal in the transducer, which converts the electrical
energy into mechanical vibrations. e transducer is then
switched to receive mode. e ultrasound wave passes into
1.1.1.4.4 Diraction
Diraction is the ability of a sound wave to bend around
the corners of an obstacle in its path and to spread into the
shadow region behind the obstacle.
the body, is reected from tissue interfaces, and returns to
the transducer in the form of an echo. e incoming echoes
are then converted back into electrical signals. e time, t,
between transmission and reception of the pulse is measured in order to calculate the length of the path traveled,
1.1.1.4.5 Attenuation andAbsorption
e intensity of an ultrasound wave diminishes as it propagates through the body. is loss of energy is known as attenuation and is caused by dierent processes, one of which is
which is the product of ultrasound velocity, c, along the
path and t. Dividing the product by the factor 2 yields z,
the distance of the reecting structure from the ultrasound
probe.
absorption– the conversion of ultrasound energy into heat.
Body tissues roughly attenuate ultrasound energy at a rate of
()()
1.1 · Technical Principles ofDiagnostic Ultrasound
Amplitude
7
1
Round trip time
. Fig. 1.4 In A-mode scanning, the amplitudes of the reected
echoes are displayed unidimensionally, representing the distances of
the reecting boundaries in the tissue from the transducer (Courtesy of
Hitachi Ltd.)
If the time difference is 0.13ms, for instance, the reflecting structure in the body is 10 cm from the ultrasound
probe. Current ultrasound systems generate and transmit
3000–5000 ultrasound pulses per second and simultaneously receive and process returning echoes to generate an
image.
1.1.1.5.2 Time Gain Compensation
Echoes returning from deeper within the body are weaker
than those arising from structures closer to the transducer.
Since the distance they have to travel is longer, they experience greater attenuation. To compensate for these differences and to display the signals returning from equally
reective boundaries with a similar brightness– regardless
of the distance traveled– the incoming echoes are amplied in a depth-dependent manner. is method of variable amplication of echoes as a function of their round
trip time is known as time gain compensation (TGC),
depth-gain compensation, or swept gain (. Fig.1.3b). e
user can set the gains for signals returning from dierent
depths.
1.1.1.5.3 A-Mode
A-mode or amplitude mode is the simplest and oldest technique of diagnostic ultrasound. e amplitudes of the pulses
returning to the transducer are displayed as spikes along
a vertical baseline on a cathode ray oscilloscope with the
position of a spike representing the distance between the
reecting boundary and the transducer (. Fig. 1.4). is
technique provides one-dimensional information and can
be used to make precise length and depth measurements. Its
use is now restricted to specialized applications including the
measurement of corneal thickness in ophthalmology and the
noninvasive evaluation of the paranasal sinuses in othorhinolaryngology.
. Fig. 1.5 In B-mode scanning, the echoes reected from boundaries
between tissues of dierent acoustic impedance are displayed twodimensionally as bright/dark spots with brightness levels representing
the intensity of the reected echoes (Courtesy of Hitachi Ltd.)
1.1.1.5.4 B-Mode
B-mode or brightness mode scans dier from A-mode displays in that the amplitudes of the returning echoes are displayed on a monitor as dots of varying brightness rather than
as spikes (. Fig.1.5). e brightness of the dots represents
the strength of the echoes. Most modern ultrasound systems
can display 256 levels of brightness (gray scales). e human
eye in comparison can distinguish only about 20 gray levels
in an image. e dots representing the echoes returning to
the transducer aer emission of a pulse are arranged along
a straight line (beam line or scan line). Aer all echoes from
preceding pulses have returned, pulses to generate successive scan lines are transmitted. Once all echoes have been
detected and processed, the complete 2D B-mode image is
displayed.
Suppose that we wish to generate a complete B-mode
image with a penetration depth of 15cm, a width of the scan
area, x, of 5cm, and a line spacing, ∆x, of 1mm. Using the
pulse-echo technique, generation of one scan line takes about
0.2ms. With the known ultrasound speed of 1540ms in living tissue, the total scan time, T, can be calculated as:
Tzxcx=
2/D
In our example, the total scan time is 10 ms, corresponding to a frame rate of 100Hz. is means that 100 complete
images can be generated per second, which is fast enough to
allow real-time imaging.
1.1.1.5.5 M-Mode
M-mode or motion mode (also known as time-motion or
TM-mode) diers from B-mode imaging in that the ultrasound beam is stationary and emitted repeatedly to obtain
echoes from moving reectors in the beam path at dierent times. e M-mode information is displayed along a
time axis with the resulting tracing depicting the movement
of a structure such as a cardiac valve in a wavelike manner
(. Fig.1.6). As with B-mode imaging, using the pulse-echo
8
Chapter 1 · Fundamental Principles
1
Time
Imaging depth
. Fig. 1.6 In M-mode scanning, the temporal changes in returning
echoes are displayed, representing the motion of reecting interfaces
toward and away from the transducer over time (Courtesy of Hitachi Ltd.)
Compromise:
resolution – penetration depth
d – smallest distance between two
structures that is resolved
1 MHz
z
Penetration depth, z
1 MHz
a
Frequency
Frequency
1/d
Resolution, 1/d
10 MHz
10 MHz
technique, it takes 0.2ms to generate a scan line with a penetration depth of 15cm. is results in a high frame rate (up
to about 5000 frames per second), aording a high temporal resolution, which is useful in evaluating rapidly moving
structures such as cardiac valves or vessel walls. M-mode is
used for echocardiography, allowing very precise measure-
b
ment of the cardiac chambers and walls and quantitative
evaluation of cardiac motion.
1.1.1.6 Resolution
Image resolution, which is given in millimeters, is dened as
the smallest distance between two structures that is necessary
. Fig. 1.7a, b Parameters aecting axial and lateral resolution.
aRelationship between axial resolution and transmit frequency (wavelength): axial resolution increases with transmit frequency (but at the
cost of penetration depth). b Eect of beam width on lateral resolution
(Courtesy of Hitachi Ltd.)
to represent them as separate entities on a monitor. When
applied to ultrasound scans, resolution describes the spatial
discrimination between two structures diering in acoustic
impedance. A distinction is made between axial resolution
(resolution in the direction of sound propagation) and lateral
resolution.
Axial resolution is determined by the length of the
excitation pulse and is typically one or a few wavelengths.
A higher- frequency transducer emits shorter wavelengths,
resulting in better axial resolution. Attenuation, however,
also increases with frequency, limiting the maximum depth
from which echoes can be received. Hence, relatively low
transmit frequencies are indispensible for imaging structures deeper in the body. e examiner must therefore
strike a balance between spatial resolution and imaging
depth (. Fig. 1.7a). Axial resolution depends on wave-
length alone and improves as the wavelength decreases
(or the frequency increases), ranging from 0.2 to 1mm
(. Table1.2).
Lateral resolution is the ability to separate two closely
spaced echoes that lie in a plane perpendicular to the direction of the sound wave. It is also inuenced by the transmit
frequency, and hence wavelength, but is mainly determined
by the focusing capabilities of the ultrasound system and the
resulting beam properties.
Lateral resolution is determined by the width of the ultra-
sound beam and is best when the beam is narrow (
. Fig.1.7b).
e beam prole changes along the beam path, consisting of
a well-focused, narrow near eld and a divergent far eld.
e ultrasound beam can be focused to improve image quality. In this way, optimal resolution can be accomplished in a
small target zone, while resolution outside this zone is much
poorer. e slow speed of sound in human tissue (1540m/s)
and the aim of achieving a high frame rate (real-time imaging) limit the number of scan lines per image. In order to
relate the echoes to a specic depth, it is necessary to wait for
the arrival of the returning echo from the respective depth of
the preceding pulse before emission of the next ultrasound
pulse. e transmitted or received pulse is focused in a longitudinal direction relative to the transducer, and focusing of
the returning pulse in the scan plane is optimized in smaller
steps (dynamically, almost continuously with the arrival time
of the pulse).
e achievable resolution is determined by the wave-
length of the ultrasound beam. It is ½ λ (wavelength) for axial
+ Multiple zone focusing
1.1 · Technical Principles ofDiagnostic Ultrasound
resolution and much poorer for lateral resolution with a value
of 4 λ. Consequently, a high transmit frequency is desirable
to achieve good axial and lateral resolution (. Table1.2). On
the other hand, due to attenuation, lower transmit frequencies are necessary to achieve greater penetration depth. When
deeper vessels are scanned, a compromise must be found at
the expense of spatial discrimination of the vessel structures
of interest (poorer spatial resolution resulting from a lower
transmit frequency) (. Fig.1.7a).
e depth of a reector in the body (encoded in the
B-mode image) is calculated from the round trip time,
which increases with depth, as does attenuation. erefore,
echo signals arriving from deeper within the body are progressively more strongly amplied in order to visualize them
with the same intensity in the resulting image (see 7 Sect.
1.1.1.5.2
). Overall gain and depth gain are adjusted according to the distance of the vessel of interest from the body
surface. e gain is crucial for the amplitude or intensity
of the signal, and along with output energy and signal-tonoise limit, it must be set properly when assessing vascular
structures.
1.1.1.7 Beam Focusing
ere are several techniques for focusing an ultrasound
beam. e simplest option is to use an acoustic lens, which
has the same eect as a glass lens for visible light. A concave
acoustic lens placed in front of the transducer provides weak
focusing at a xed depth. e site of maximum focusing is
referred to as the focal point or focal zone. Alternatively, the
crystal in the transducer can be made concave, providing
internal focusing. is technique is used in single-element
mechanical sector scanners.
More exible beam forming, with a variable depth of the
focal point, is accomplished using electronic beam focusing.
Array transducers consist of multiple crystal elements placed
side by side. Depending on the scanner type, the number of
individual elements ranges from 60 to 256. Variable numbers of elements can be activated simultaneously to form
an ultrasound beam. If the elements forming the beam are
excited at slightly dierent times, a concave wavefront is
generated, causing the beam to converge at the focal point.
e site of the focal point can be manipulated by varying the
number of active elements and the pattern of excitation of
individual elements. e user can thus adjust the beam to
achieve maximum lateral resolution at the anatomic site of
interest. Modern scanners use multiple zone focusing, which
reduces the frame rate, as several consecutive beams with
dierent focal points are transmitted to generate a scan line.
A technique known as dynamic focusing allows the focus of
the beam to be altered during reception by imposing variable
delays on signals from dierent depths. With this technique,
the reception focus can be optimized without compromising
the frame rate. Groups of 8–128 elements are used for focusing the beam (
. Fig.1.8).
9
+ Variable focal zone
. Fig. 1.8 Beam focusing in modern array probes. By delaying the
ring of the central element after the ring of the outer elements a
curved wavefront is produced, resulting in a focused beam (Courtesy
of Hitachi Ltd.)
Lateral resolution is limited by the proximity of the
transducer elements activated to emit an ultrasound pulse.
Resolution along the longitudinal axis can be improved by
exciting only a limited number of elements at a time and not
the whole array. A more focused beam is achieved by later
excitation of the transducer elements in the center. Dynamic
focusing is accomplished by applying small time delays to the
excitation pulses driving the individual transducer elements.
Resolution in the third direction, or slice thickness, depends
on the position in the image.
1.1.1.8 Types ofTransducers
1.1.1.8.1 Principle ofOperation
Most electronic ultrasound transducers used today contain a
number of individual piezoelectric elements for transmitting
and receiving ultrasonic waves. To create a complete image,
the ultrasound beam has to pass through adjacent areas of
tissue. Parallel ultrasound beams are generated by varying
the groups of elements within the array that are simultaneously active. A group of elements is excited to generate the
rst scan line. e next adjacent scan line is formed by shiing the group of active elements along the transducer array,
one element position from the rst group– for example, elements 1–5 produce the rst beam, 2–6 the second, 3–7 the
third, and so on (
. Fig. 1.9). e second ultrasound beam
generated in this way is said to be shied by the width of one
element. e number of scan lines used to generate an image
can be increased by varying the number of elements activated simultaneously to generate each beam. For instance, if
the second beam is generated using the same group of elements as for the rst beam plus one additional element on
the le side (and no element on the right side is switched
o), then the axis of the second beam is shied by half an
element width relative to the rst beam. e third beam is
generated by removing one element on the right side without
1
10
Scan direc
Element group
Chapter 1 · Fundamental Principles
adding an element to the le side. In this way, the number of
1
lines scanned to produce an image is doubled. A higher line
density is desirable for improving image quality; however, it
also reduces frame rate.
1.1.1.8.2 Linear Arrays
In a linear array transducer, the individual crystal elements
are arranged in a straight row (. Fig.1.10) and can be pulsed
to generate adjacent parallel ultrasonic beams, producing a
rectangular image with nearly constant resolution over the
entire scan depth. A linear array is made up of 60–196 elements, with an element width of 1–4 λ, and operates at a
frequency of 5–13 MHz. An acoustic lens can be used for
focusing perpendicular to the direction of beam propagation.
1.1.1.8.3 Curved or Convex Arrays
A curved or convex array transducer is a linear array, with the
individual elements arranged along a curved line, to produce
tion
a sector image (
. Fig.1.11). As the lines fan out with increas-
ing distance from the transducer, lateral resolution decreases
. Fig. 1.9 Emission and reception of a series of parallel ultrasound
beams by successive excitation of groups of transducer elements for
generation of an ultrasound image (Courtesy of Hitachi Ltd.)
with depth. A typical curvilinear array consists of at least
96 elements and has a radius of 25–80mm and a frequency
range of 3–7MHz. Most curvilinear scanners produce sector
images ranging in size from 60° to 90°.
. Fig. 1.10 Diagram of a linear array with the crystal elements
arranged in a straight row (Courtesy of Hitachi Ltd.)
1.1.1.8.4 Sector Scanners
Sector scanners have a smaller radius (<25mm) than curved
arrays and also have a small footprint, resulting in a narrow
near eld. With a beam-steering angle >90°, these probes are
especially useful where access is dicult, such as in the imaging of the heart through the intercostal spaces (echocardiography), or for endoluminal applications such as transvaginal
ultrasound.
1.1.1.8.5 Phased Arrays
In a phased-array transducer, the elements are also arranged
in a linear array. e dierence is that all elements are excited
to generate a scan line. However, time delays are introduced
between pulsing consecutive elements to produce a wavefront that is no longer perpendicular to the transducer face
. Fig. 1.12). By choosing appropriate delays between the
(
excitation of individual elements, it is possible to direct the
beam at a desired angle. Using this method, the beam can be
steered through a range of angles to produce a sector image.
Phased- array transducers use a smaller array of elements
(64–128), resulting in a small footprint of 12–20mm. e
beam covers a sector of 80–90° with frequency ranging from
2 to 7MHz. Since they require complex electronic circuitry,
phased-array devices are expensive and are used mainly for
cardiac and transcranial imaging.
. Fig. 1.11 Diagram of a curved array with the crystal elements
arranged along a curved line (Courtesy of Hitachi Ltd.)
1.1.1.8.6 Mechanical Sector Scanners
Compared with electronic phased arrays, mechanical systems are fairly simple regarding the control of transducer elements and signal processing. ere are basically two designs
of mechanical devices: the rotating wheel transducer and the
wobbler transducer.
W2
W1
1.1 · Technical Principles ofDiagnostic Ultrasound
T2
T1
. Table 1.4 Overview of ultrasound artifacts
Underlying mechanism Type of artifact
11
1
. Fig. 1.12 Generation of a pie-shaped image by the successive
excitation of groups of elements in a phased-array probe (Courtesy of
Hitachi Ltd.)
5 Rotating wheel transducer. is type usually comprises
three to ve transducer elements mounted 120–72°
apart on a wheel. A motor housed in the handle turns
the wheel at a constant rate in one direction. One of the
crystal elements at a time is activated as it rotates past
an acoustically transparent window. e active element
scans a sector-shaped region. en the next crystal
rotates past the window, generating a second image.
5 Wobbler transducer. In this type of mechanical sec-
tor scanner, a single crystal oscillates about a pivotal
point, producing a beam that covers a sector of 60–100°.
Since the wobbler transducer consists of a single crystal
element, no complex adjustment is required. Another
advantage it has over the rotating wheel transducer is
that the sector angle is variable. Both mechanical devices
are limited, however, by the fact that only a single element is used to produce the ultrasound beam and thus
only xed beam focusing is possible.
1.1.1.8.7 Annular Phased Arrays
An annular phased array is an oscillating transducer combining features of mechanical and electronic devices. Instead of
a single element, the transducer consists of several concentric
rings (annuli). Each ring can be excited separately, allowing
variable focusing in two dimensions.
1.1.1.8.8 Disadvantages ofMechanical
Transducers
Regardless of their design, mechanical probes are subject to
wear and require maintenance. Moreover, they are relatively
slow, not allowing rapid switching between dierent scan
Nonuniform ultrasound
propagation in the
human body
Nonuniform ultrasound
attenuation
Ultrasound beam
characteristics
Structural artifacts Speckles
Structures with misregistered location
Refraction artifact
Reverberation artifact
Mirror artifact
Acoustic shadowing
Edge artifact
Acoustic enhancement
Side lobe artifact
Line distortion
Falsely perceived sediment
modes (B-mode, M-mode, Doppler). Real-time display of
B-mode/M-mode or B-mode/Doppler information is generally not possible.
1.1.1.9 Ultrasound Artifacts
Artifacts play a much greater role in diagnostic ultrasound
compared with other imaging modalities such as computed
tomography (CT) or magnetic resonance imaging (MRI).
One fundamental issue is that several simplifying assumptions are made, namely that parameters such as the speed
of sound in tissues, the propagation of ultrasound, and the
attenuation are constant. Another important source of artifacts in the ultrasound image is the use of inadequate instrument settings. At the same time, however, some common
artifacts can be exploited to advantage because they may
provide additional diagnostic information on tissue composition. Oen, artifacts can be identied by moving the transducer: artifacs will change position or disappear while actual
tissue structures will not.
. Table1.4 provides an overview of ultrasound artifacts
and their underlying causes. e artifacts that are most relevant to vascular applications are described in more detail in
the following sections.
1.1.1.9.1 Posterior Shadowing
Acoustic shadowing is the occurrence of hypoechoic areas
behind certain objects due to loss of energy; it is one of the
most commonly encountered ultrasound artifacts. ese
artifacts can occur deep to a strong reector such as air, which
is dicult to penetrate by the ultrasound beam because of
a strong acoustic mismatch, or behind highly attenuating
structures such as bone or calculi, which absorb much of the
ultrasound energy (. Fig.1.13).
1.1.1.9.2 Acoustic Enhancement
Acoustic enhancement is an increase in brightness behind
a low-attenuating area, in particular uid-lled spaces such
as cysts. An ultrasound beam passing through uid is nearly