Diagnostic Ultrasound

2009 Pages·2018·227.066 MB·English

by  Carol M. Rumack, Stephanie R. Wilson, J. William Charboneau, Deborah Levine

DIAGNOSTIC
ULTRASOUND
DIAGNOSTIC
ULTRASOUND
5TH EDITION
CAROL M. RUMACK, MD, FACR
Vice Chair of Education and Professional Development
Professor of Radiology and Pediatrics
Associate Dean for GME
University of Colorado School of Medicine
Denver, Colorado
DEBORAH LEVINE, MD, FACR
Co-Chief of Ultrasound
Director of OB/Gyn Ultrasound
Vice Chair of Academic Affairs
Department of Radiology
Beth Israel Deaconess Medical Center
Professor of Radiology
Harvard Medical School
Boston, Massachusetts
PART ONE: Physics
CHAPTER
Physics of Ultrasound
1
Christopher R.B. Merritt
SUMMARY OF KEY POINTS
• Quality imaging requires an understanding of basic capabilities of conventional gray-scale
acoustic principles. imaging.
• Image interpretation requires recognition and • Knowledge of mechanical and thermal bioeffects of
understanding of common artifacts. ultrasound is necessary for prudent use.
• Special modes of operation, including harmonic imaging, • High-intensity focused ultrasound has potential therapeutic
compounding, elastography, and Doppler, expand the applications.
CHAPTER OUTLINE
BASIC ACOUSTICS Two-Dimensional Arrays Interpretation of the Doppler Spectrum
Wavelength and Frequency Transducer Selection Interpretation of Color Doppler
Propagation of Sound IMAGE DISPLAY AND STORAGE Other Technical Considerations
Distance Measurement SPECIAL IMAGING MODES Doppler Frequency
Acoustic Impedance Tissue Harmonic Imaging Wall Filters
Reflection Spatial Compounding Spectral Broadening
Refraction Three-Dimensional Ultrasound Aliasing
Attenuation Ultrasound Elastography Doppler Angle
INSTRUMENTATION Strain Elastography Sample Volume Size
Transmitter Shear Wave Elastography Doppler Gain
Transducer IMAGE QUALITY OPERATING MODES: CLINICAL
Receiver Spatial Resolution IMPLICATIONS
Image Display IMAGING PITFALLS Bioeffects and User Concerns
Mechanical Sector Scanners Shadowing and Enhancement THERAPEUTIC APPLICATIONS:
Arrays DOPPLER SONOGRAPHY HIGH-INTENSITY FOCUSED
Linear Arrays Doppler Signal Processing and Display ULTRASOUND
Curved Arrays Doppler Instrumentation
Phased Arrays Power Doppler
All diagnostic ultrasound applications are based on the detec- empower ultrasound with its unique diagnostic capabilities. The
tion and display of acoustic energy reflected from interfaces user must understand the fundamentals of the interactions of
within the body. These interactions provide the information acoustic energy with tissue and the methods and instruments
needed to generate high-resolution, gray-scale images of the used to produce and optimize the ultrasound display. With this
body, as well as display information related to blood flow. Its knowledge the user can collect the maximum information from
unique imaging attributes have made ultrasound an important each examination, avoiding pitfalls and errors in diagnosis that
and versatile medical imaging tool. However, expensive state- may result from the omission of information or the misinterpreta-
of-the-art instrumentation does not guarantee the production tion of artifacts.1
of high-quality studies of diagnostic value. Gaining maximum Ultrasound imaging and Doppler ultrasound are based on
benefit from this complex technology requires a combination the scattering of sound energy by interfaces of materials with
of skills, including knowledge of the physical principles that different properties through interactions governed by acoustic
1
2 PART I Physics
physics. The amplitude of reflected energy is used to generate is the hertz (Hz); 1 Hz = 1 cycle per second. High frequencies
ultrasound images, and frequency shifts in the backscattered are expressed in kilohertz (kHz; 1 kHz = 1000 Hz) or megahertz
ultrasound provide information relating to moving targets such (MHz; 1 MHz = 1,000,000 Hz).
as blood. To produce, detect, and process ultrasound data, users In nature, acoustic frequencies span a range from less than
must manage numerous variables, many under their direct control. 1 Hz to more than 100,000 Hz (100 kHz). Human hearing is
To do this, operators must understand the methods used to limited to the lower part of this range, extending from 20 to
generate ultrasound data and the theory and operation of the 20,000 Hz. Ultrasound differs from audible sound only in its
instruments that detect, display, and store the acoustic information frequency, and it is 500 to 1000 times higher than the sound we
generated in clinical examinations. normally hear. Sound frequencies used for diagnostic applications
This chapter provides an overview of the fundamentals of typically range from 2 to 15 MHz, although frequencies as high
acoustics, the physics of ultrasound imaging and flow detection, as 50 to 60 MHz are under investigation for certain specialized
and ultrasound instrumentation with emphasis on points most imaging applications. In general, the frequencies used for ultra-
relevant to clinical practice. A discussion of the therapeutic sound imaging are higher than those used for Doppler. Regardless
application of high-intensity focused ultrasound concludes the of the frequency, the same basic principles of acoustics apply.
chapter.
Propagation of Sound
In most clinical applications of ultrasound, brief bursts or pulses
BASIC ACOUSTICS of energy are transmitted into the body and propagated through
tissue. Acoustic pressure waves can travel in a direction perpen-
Wavelength and Frequency
dicular to the direction of the particles being displaced (transverse
Sound is the result of mechanical energy traveling through matter waves), but in tissue and fluids, sound propagation is primarily
as a wave producing alternating compression and rarefaction. along the direction of particle movement (longitudinal waves).
Pressure waves are propagated by limited physical displacement Longitudinal waves are important in conventional ultrasound
of the material through which the sound is being transmitted. imaging and Doppler, while transverse waves are measured in
A plot of these changes in pressure is a sinusoidal waveform shear wave elastography. The speed at which pressure waves
(Fig. 1.1), in which the Y axis indicates the pressure at a given move through tissue varies greatly and is affected by the physical
point and the X axis indicates time. Changes in pressure with properties of the tissue. Propagation velocity is largely determined
time define the basic units of measurement for sound. The distance by the resistance of the medium to compression, which in turn
between corresponding points on the time-pressure curve is is influenced by the density of the medium and its stiffness or
defined as the wavelength (λ), and the time (T) to complete a elasticity. Propagation velocity is increased by increasing stiffness
single cycle is called the period. The number of complete cycles and reduced by decreasing density. In the body, propagation
in a unit of time is the frequency (f) of the sound. Frequency velocity of longitudinal waves may be regarded as constant for
and period are inversely related. If the period (T) is expressed a given tissue and is not affected by the frequency or wavelength
in seconds, f = 1/T, or f = T × sec−1. The unit of acoustic frequency of the sound. This is in contrast to transverse (shear) waves for
FIG. 1.1 Sound Waves. Sound is transmitted mechanically at the molecular level. In the resting state the pressure is uniform throughout the

medium. Sound is propagated as a series of alternating pressure waves producing compression and rarefaction of the conducting medium. The
time for a pressure wave to pass a given point is the period, T. The frequency of the wave is 1/T. The wavelength, λ, is the distance between
corresponding points on the time-pressure curve.
CHAPTER 1 Physics of Ultrasound 3
Air 330
Fat 1450
Water 1480
Soft tissue
1540
(average)
Liver 1550
Kidney 1560
Blood 1570
Muscle 1580
Bone 4080
1400 1500 1600 1700 1800
Propagation velocity (meters/second)
FIG. 1.2 Propagation Velocity. In the body, propagation velocity of
FIG. 1.3 Propagation Velocity Artifact. When sound passes through

sound is determined by the physical properties of tissue. As shown,
a lesion containing fat, echo return is delayed because fat has a propaga-
this varies considerably. Medical ultrasound devices base their measure-
tion velocity of 1450 m/sec, which is less than the liver. Because the
ments on an assumed average propagation velocity of soft tissue of
ultrasound scanner assumes that sound is being propagated at the
1540 m/sec.
average velocity of 1540 m/sec, the delay in echo return is interpreted
as indicating a deeper target. Therefore the final image shows a mis-
registration artifact in which the diaphragm and other structures deep
to the fatty lesion are shown in a deeper position than expected (simulated
image).
which the velocity is determined by Young modulus, a measure
of tissue stiffness or elasticity.
Fig. 1.2 shows typical longitudinal propagation velocities for propagation velocity of sound for the tissue is known. For example,
a variety of materials. In the body the propagation velocity of if the time from the transmission of a pulse until the return of
sound is assumed to be 1540 meters per second (m/sec). This an echo is 0.000145 seconds and the velocity of sound is 1540 m/
value is the average of measurements obtained from normal soft sec, the distance that the sound has traveled must be 22.33 cm
tissue.2,3 Although this value represents most soft tissues, such (1540 m/sec × 100 cm/m × 0.000145 sec = 22.33 cm). Because
tissues as aerated lung and fat have propagation velocities sig- the time measured includes the time for sound to travel to the
nificantly less than 1540 m/sec, whereas tissues such as bone interface and then return along the same path to the transducer,
have greater velocities. Because a few normal tissues have propaga- the distance from the transducer to the reflecting interface is
tion values significantly different from the average value assumed 22.33 cm/2 = 11.165 cm. By rapidly repeating this process, a
by the ultrasound scanner, the display of such tissues may be two-dimensional (2-D) map of reflecting interfaces is created to
subject to measurement errors or artifacts (Fig. 1.3). The propaga- form the ultrasound image. The accuracy of this measurement
tion velocity of sound (c) is related to frequency and wavelength is therefore highly influenced by how closely the presumed velocity
by the following simple equation: of sound corresponds to the true velocity in the tissue being
observed (see Figs. 1.2 and 1.3), as well as by the important
c= fλ
assumption that the sound pulse travels in a straight path to and
Thus a frequency of 5 MHz can be shown to have a wavelength from the reflecting interface.
of 0.308 mm in tissue: λ = c/f = 1540 m/sec × 5,000,000 sec−1 =
Acoustic Impedance
0.000308 m = 0.308 mm. Wavelength is an important determinant
of spatial resolution in ultrasound imaging, and selection of Current diagnostic ultrasound scanners rely on the detection
transducer frequency for a given application is a key user decision. and display of reflected sound or echoes. Imaging based on
transmission of ultrasound is also possible, but this is not used
Distance Measurement
clinically at present. To produce an echo, a reflecting interface
Propagation velocity is a particularly important value in clinical must be present. Sound passing through a totally homogeneous
ultrasound and is critical in determining the distance of a reflect- medium encounters no interfaces to reflect sound, and the
ing interface from the transducer. Much of the information used medium appears anechoic or cystic. The junction of tissues or
to generate an ultrasound scan is based on the precise measure- materials with different physical properties produces an acoustic
ment of time and employs the principles of echo-ranging (Fig. interface. These interfaces are responsible for the reflection of
1.4). If an ultrasound pulse is transmitted into the body and the variable amounts of the incident sound energy. Thus when
time until an echo returns is measured, it is simple to calculate ultrasound passes from one tissue to another or encounters a
the depth of the interface that generated the echo, provided the vessel wall or circulating blood cells, some of the incident sound
4 PART I Physics
× ms
×
FIG. 1.4 Ultrasound Ranging. The information used to position an echo for display is based on the precise measurement of time. Here the

time for an echo to travel from the transducer to the target and return to the transducer is 0.145 ms (0.000145 seconds). Multiplying the velocity
of sound in tissue (1540 m/sec) by the time shows that the sound returning from the target has traveled 22.33 cm. Therefore the target lies half
this distance, or 11.165 cm, from the transducer. By rapidly repeating this process, a two-dimensional map of reflecting interfaces is created to
form the ultrasound image.
energy is reflected. The amount of reflection or backscatter is
Examples of Specular Reflectors
determined by the difference in the acoustic impedances of the
materials forming the interface.
Diaphragm
Acoustic impedance (Z) is determined by product of the
Vessel wall
density (ρ) of the medium propagating the sound and the Wall of urine-filled bladder
propagation velocity (c) of sound in that medium (Z = ρc). Endometrial stripe
Interfaces with large acoustic impedance differences, such as
interfaces of tissue with air or bone, reflect almost all the incident
energy. Interfaces composed of substances with smaller differences
in acoustic impedance, such as a muscle and fat interface, reflect
only part of the incident energy, permitting the remainder to on the angle of insonation (exposure to ultrasound waves).
continue onward. As with propagation velocity, acoustic imped- Specular reflectors will return echoes to the transducer only if
ance is determined by the properties of the tissues involved and the sound beam is perpendicular to the interface. If the interface
is independent of frequency. is not at a near 90-degree angle to the sound beam, it will be
reflected away from the transducer, and the echo will not be
Reflection
detected (see Fig. 1.5A).
The way ultrasound is reflected when it strikes an acoustic Most echoes in the body do not arise from specular reflectors
interface is determined by the size and surface features of the but rather from much smaller interfaces within solid organs. In
interface (Fig. 1.5). If large and relatively smooth, the interface this case the acoustic interfaces involve structures with individual
reflects sound much as a mirror reflects light. Such interfaces dimensions much smaller than the wavelength of the incident
are called specular reflectors because they behave as “mirrors sound. The echoes from these interfaces are scattered in all
for sound.” The amount of energy reflected by an acoustic interface directions. Such reflectors are called diffuse reflectors and account
can be expressed as a fraction of the incident energy; this is for the echoes that form the characteristic echo patterns seen in
termed the reflection coefficient (R). If a specular reflector is solid organs and tissues (see Fig. 1.5B). The constructive and
perpendicular to the incident sound beam, the amount of energy destructive interference of sound scattered by diffuse reflectors
reflected is determined by the following relationship: results in the production of ultrasound speckle, a feature of
tissue texture of sonograms of solid organs (Fig. 1.6). For some
R=(Z −Z )2 (Z +Z )2
2 1 2 1 diagnostic applications, the nature of the reflecting structures
where Z and Z are the acoustic impedances of the media forming creates important conflicts. For example, most vessel walls behave
1 2
the interface. as specular reflectors that require insonation at a 90-degree angle
Because ultrasound scanners only detect reflections that return for best imaging, whereas Doppler imaging requires an angle of
to the transducer, display of specular interfaces is highly dependent less than 90 degrees between the sound beam and the vessel.
CHAPTER 1 Physics of Ultrasound 5
A B
FIG. 1.5 Specular and Diffuse Reflectors. (A) Specular reflector. The diaphragm is a large and relatively smooth surface that reflects sound

like a mirror reflects light. Thus sound striking the diaphragm at nearly a 90-degree angle is reflected directly back to the transducer, resulting in
a strong echo. Sound striking the diaphragm obliquely is reflected away from the transducer, and an echo is not displayed (yellow arrow).
(B) Diffuse reflector. In contrast to the diaphragm, the liver parenchyma consists of acoustic interfaces that are small compared to the wavelength
of sound used for imaging. These interfaces scatter sound in all directions, and only a portion of the energy returns to the transducer to produce
the image.
propagation velocities of sound in the media forming the interface
(Fig. 1.7). Refraction is important because it is one cause of
misregistration of a structure in an ultrasound image (Fig. 1.8).
When an ultrasound scanner detects an echo, it assumes that
the source of the echo is along a fixed line of sight from the
transducer. If the sound has been refracted, the echo detected
may be coming from a different depth or location than the image
shown in the display. If this is suspected, increasing the scan
angle so that it is perpendicular to the interface minimizes the
artifact.
Attenuation
As the acoustic energy moves through a uniform medium, work
is performed and energy is ultimately transferred to the transmit-
ting medium as heat. The capacity to perform work is determined
by the quantity of acoustic energy produced. Acoustic power,
FIG. 1.6 Ultrasound Speckle. Close inspection of an ultrasound expressed in watts (W) or milliwatts (mW), describes the amount
image of the breast containing a small cyst reveals it to be composed
of acoustic energy produced in a unit of time. Although measure-
of numerous areas of varying intensity (speckle). Speckle results from
ment of power provides an indication of the energy as it relates
the constructive (red) and destructive (green) interaction of the acoustic
fields (yellow rings) generated by the scattering of ultrasound from small to time, it does not take into account the spatial distribution of
tissue reflectors. This interference pattern gives ultrasound images their the energy. Intensity (I) is used to describe the spatial distribution
characteristic grainy appearance and may reduce contrast. Ultrasound of power and is calculated by dividing the power by the area
speckle is the basis of the texture displayed in ultrasound images of
over which the power is distributed, as follows:
solid tissues.
I(W/cm2)=Power(W) Area(cm2)
Refraction
The attenuation of sound energy as it passes through tissue
When sound passes from a tissue with one acoustic propagation is of great clinical importance because it influences the depth in
velocity to a tissue with a higher or lower sound velocity, there tissue from which useful information can be obtained. This in
is a change in the direction of the sound wave. This change in turn affects transducer selection and a number of operator-
direction of propagation is called refraction and is governed by controlled instrument settings, including time (or depth) gain
Snell law: compensation, power output attenuation, and system gain levels.
Attenuation is measured in relative rather than absolute units.
sinθ sinθ =c c
1 2 1 2
The decibel (dB) notation is generally used to compare different
where θ is the angle of incidence of the sound approaching levels of ultrasound power or intensity. This value is 10 times
1
the interface, θ is the angle of refraction, and c and c are the the log of the ratio of the power or intensity values being
2 1 2 10
6 PART I Physics
θ = 20°
1
Tissue A
c = 1540 m/sec
1
Tissue B
c = 1450 m/sec
2
A B
θ = 1188.88°
2
FIG. 1.7 Refraction. When sound passes from tissue A with propaga-

tion velocity (c) to tissue B with a different propagation velocity (c),
1 2
there is a change in the direction of the sound wave because of refraction.
The degree of change is related to the ratio of the propagating velocities
of the media forming the interface (sinθ1/sinθ2 = c1/c2).
compared. For example, if the intensity measured at one point
in tissues is 10 mW/cm2 and at a deeper point is 0.01 mW/cm2,
the difference in intensity is as follows:
(10)(log 0.0110)=(10)(log 0.001)=(10)(−log 1000)
10 10 10
=(10)(−33)=−30dB
As it passes through tissue, sound loses energy, and the pressure C
waves decrease in amplitude as they travel farther from their
source. Contributing to the attenuation of sound are the transfer FIG. 1.8 Refraction Artifact. (A) and (B) Production of an artifact
by refraction of sound in a transverse scan of the mid abdomen. The
of energy to tissue, resulting in heating (absorption), and the
direct sound path properly depicts the location of the object. (B) A
removal of energy by reflection and scattering. Attenuation is
“ghost image” (red) produced by refraction at the edge of the rectus
therefore the result of the combined effects of absorption, scat- abdominis muscle. The transmitted and reflected sound travels along
tering, and reflection. Attenuation depends on the insonating the path of the black arrows. The scanner assumes the returning signal
frequency as well as the nature of the attenuating medium. High is from a straight line (red arrow) and displays the structure at the
incorrect location. (C) Axial transabdominal image of the uterus showing
frequencies are attenuated more rapidly than lower frequencies,
a small gestational sac (A) and what appears to be a second sac (B)
and transducer frequency is a major determinant of the useful
due to refraction artifact.
depth from which information can be obtained with ultrasound.
Attenuation determines the efficiency with which ultrasound
CHAPTER 1 Physics of Ultrasound 7
PRF determines the time interval between ultrasound pulses
Water 0.00
and is important in determining the depth from which unambigu-
Blood 0.18 ous data can be obtained both in imaging and Doppler modes.
The ultrasound pulses must be spaced with enough time between
Fat 0.63
the pulses to permit the sound to travel to the depth of interest
Soft tissue
0.70 and return before the next pulse is sent. For imaging, PRFs from
(average)
1 to 10 kHz are used, resulting in an interval of 0.1 to 1 ms
Liver 0.94
between pulses. Thus a PRF of 5 kHz permits an echo to travel
Kidney 1.00 and return from a depth of 15.4 cm before the next pulse is sent.
Muscle
1.30 Transducer
(parallel)
Muscle 3.30 A transducer is any device that converts one form of energy to
(transverse)
another. In ultrasound the transducer converts electric energy
Bone 5.00 to mechanical energy, and vice versa. In diagnostic ultrasound
systems the transducer serves two functions: (1) converting the
Air 10.00
electric energy provided by the transmitter to the acoustic pulses
0 2 4 6 8 10 directed into the patient and (2) serving as the receiver of reflected
Attenuation (dB/cm/MHz) echoes, converting weak pressure changes into electric signals
for processing.
FIG. 1.9 Attenuation. As sound passes through tissue, it loses energy

through the transfer of energy to tissue by heating, reflection, and Ultrasound transducers use piezoelectricity, a principle
scattering. Attenuation is determined by the insonating frequency and discovered by Pierre and Jacques Curie in 1880.5 Piezoelectric
the nature of the attenuating medium. Attenuation values for normal materials have the unique ability to respond to the action of an
tissues show considerable variation. Attenuation also increases in propor-
electric field by changing shape. They also have the property of
tion to insonating frequency, resulting in less penetration at higher
generating electric potentials when compressed. Changing the
frequencies.
polarity of a voltage applied to the transducer changes the thick-
ness of the transducer, which expands and contracts as the polarity
penetrates a specific tissue and varies considerably in normal changes. This results in the generation of mechanical pressure
tissues (Fig. 1.9). waves that can be transmitted into the body. The piezoelectric
effect also results in the generation of small potentials across
INSTRUMENTATION the transducer when the transducer is struck by returning echoes.
Positive pressures cause a small polarity to develop across the
Ultrasound scanners are complex and sophisticated imaging transducer; negative pressure during the rarefaction portion of
devices, but all consist of the following basic components to the acoustic wave produces the opposite polarity across the
perform key functions: transducer. These tiny polarity changes and the associated voltages
• Transmitter or pulser to energize the transducer are the source of all the information processed to generate an
• Ultrasound transducer ultrasound image or Doppler display.
• Receiver and processor to detect and amplify the backscattered When stimulated by the application of a voltage difference
energy and manipulate the reflected signals for display across its thickness, the transducer vibrates. The frequency of
• Display that presents the ultrasound image or data in a form vibration is determined by the transducer material. When the
suitable for analysis and interpretation transducer is electrically stimulated, a range or band of frequen-
• Method to record or store the ultrasound image cies results. The preferential frequency produced by a transducer
is determined by the propagation speed of the transducer material
Transmitter
and its thickness. In the pulsed wave operating modes used for
Most clinical applications use pulsed ultrasound, in which brief most clinical ultrasound applications, the ultrasound pulses
bursts of acoustic energy are transmitted into the body. The contain additional frequencies that are both higher and lower
source of these pulses, the ultrasound transducer, is energized than the preferential frequency. The range of frequencies produced
by application of precisely timed, high-amplitude voltage. The by a given transducer is termed its bandwidth. Generally, the
maximum voltage that may be applied to the transducer is limited shorter the pulse of ultrasound produced by the transducer, the
by federal regulations that restrict the acoustic output of diagnostic greater is the bandwidth.
scanners. Most scanners provide a control that permits attenuation Most modern digital ultrasound systems employ broad-
of the output voltage. Because the use of maximum output results bandwidth technology. Ultrasound bandwidth refers to the range
in higher exposure of the patient to ultrasound energy, prudent of frequencies produced and detected by the ultrasound system.
use dictates use of the output attenuation controls to reduce This is important because each tissue in the body has a charac-
power levels to the lowest levels consistent with the diagnostic teristic response to ultrasound of a given frequency, and different
problem.4 tissues respond differently to different frequencies. The range of
The transmitter also controls the rate of pulses emitted by frequencies arising from a tissue exposed to ultrasound is referred
the transducer, or the pulse repetition frequency (PRF). The to as the frequency spectrum bandwidth of the tissue, or tissue