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Файл:Ординатура / Хирургия / Библиотека им академика М.И. Перельмана / Книга_3786_Библиотеки_им_академика_М_И_Перельмана
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Part II
https://t.me/med1917
Vein Testing
Ultrasound Physics
https://t.me/med1917
Frank R. Miele
5
Contents
5.1 Introduction ................................................. 61
Ultrasound Physics and Ultrasound
5.2
Basics ............................................................ 62
5.2.1 Sound Wave Parameters ................................ 62
5.2.2 Transducers, Absorption, and Resolution ..... 65
5.2.3 Linear Phased Arrays .................................... 66
5.3
Ultrasound Imaging .................................... 66
5.3.1 Operating Frequency ..................................... 66
5.3.2 Transmit Power ............................................. 67
Transmit Focus .............................................. 67
5.3.3
5.3.4 Signal and Gain ............................................. 67
5.3.5 Signal and Compression ................................ 67
5.3.6 Signal and Ambient Light ............................. 68
5.3.7 Harmonic Imaging ........................................ 68
5.3.8 Compound Imaging ....................................... 69
5.3.9 Artifacts ......................................................... 69
5.4
Spectral Doppler .......................................... 71
5.4.1 Doppler Theory and Basics ........................... 71
5.4.2 Doppler Shift Detection ................................ 72
Angle Correction, Velocity, Direction,
5.4.3
and Acceleration ............................................ 72
5.4.4 Optimal Operating Frequency ....................... 72
5.4.5 Wall Filters .................................................... 73
5.5 Color Doppler .............................................. 73
5.5.1 Color Doppler Basics .................................... 73
5.5.2 Optimizing Operating Frequency .................. 74
5.5.3 Color Priority and Color Gain ....................... 74
Color Scales and the PRF .............................. 75
5.5.4
5.5.5 Decreasing the Color Scales .......................... 75
5.5.6 Color Wall Filters .......................................... 76
5.5.7 Color Power Doppler .................................... 77
Conclusion ............................................................... 77
F.R. Miele, MSEE
Pegasus Lectures, Inc., Forney, TX, USA
e-mail: fmiele@pegasuslectures.com
Abstract
Success in phlebology is predicated on the
quality and accuracy of the ultrasound imaging and Doppler data pre-intervention, during
intervention, and post-intervention. In order to
avoid errors, image quality must be optimized,
Doppler must be performed and interpreted
correctly, and artifacts must be recognized and
minimized. These requirements are met often
through manipulation of scanning technique as
well as changing instrument control settings.
Taken together, these requirements designate
that both foundational physics and ultrasound
instrumentation must be well understood to
ensure high-quality patient care. As such, the
starting point for phlebology is a treatment in
ultrasound physics and instrumentation.
5.1 Introduction
Success in phlebology is predicated on the quality and accuracy of the ultrasound imaging and
Doppler data pre-intervention, during intervention, and post-intervention. In order to avoid
errors, image quality must be optimized, Doppler
must be performed and interpreted correctly,
and artifacts must be recognized and minimized.
These requirements are met often through manipulation of scanning technique as well as changing instrument control settings. Taken together,
these requirements designate that both foundational physics and ultrasound instrumentation
must be well understood to ensure high-quality
E. Mowatt-Larssen et al. (eds.), Phlebology, Vein Surgery and Ultrasonography,
DOI 10.1007/978-3-319-01812-6_5, © Springer International Publishing Switzerland 2014
61
62
1
a
b
l
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F.R. Miele
patient care. As such, the starting point for phlebology is a treatment in ultrasound physics and
instrumentation.
5.2 Ultrasound Physics
Ultrasound Basics
and
5.2.1 Sound Wave Parameters
Understanding ultrasound physics and instrumentation requires a brief review of sound and
wave parameters. Sound is a mechanical wave,
implying a physical interaction between the
ultrasound (the wave) and tissue (the medium).
It is precisely the mechanical interaction which
allows for ultrasound image generation. As the
sound waves propagate through the body,
the physical interactions result in changes to the
wave characteristics which are then processed to
produce an image. Since varying the initial signal characteristics affects the characteristics of
the returning signal, we start by discussing the
basic wave parameters. By alternately compressing and rarefying the particles of the medium,
sound waves propagate through the medium.
The number of compression rarefaction pairs per
second is referred to as the frequency (Fig. 5.1).
Thus, a 5 MHz wave compresses and rarefies the
molecules of a medium five million times per
second. The distance between the compressions
is referred to as the wavelength.
With respect to ultrasound, the wavelength is
an extremely important parameter as the wavelength affects both the type of reflection that
occurs as well as the axial resolution of the
image. As depicted in Fig. 5.2a, b, within the
same medium, higher frequencies result in
shorter wavelengths while lower frequencies
result in longer wavelengths.
Acoustic Power and Intensity
The power of the wave relates to the pressure
developed within the tissue. In order to produce
the sound waves, a piezoelectric transducer is
excited by an electrical signal, referred to as the
transmit voltage. Higher transmit voltages produce higher acoustic pressure fields, increasing
the strength of the reflected signal as well as
increasing the maximum depth of penetration. Of
course, because of the physical interaction with
the medium, a higher acoustic pressure potentially increases the risk of causing tissue damage
(referred to as a bioeffect). The parameter that is
measured to assess both the ability to improve
signal strength and the risk of inducing a bioeffect is the intensity. The intensity is a measure of
the distribution of the power per unit area, as
Fig. 5.1 Sound waves
(longitudinal wave) propagate
through compression and
rarefactions of the medium
Fig. 5.2 (a) Higher
frequency and shorter
wavelength. (b) Lower
frequency and longer
wavelength
Compression
Rarefaction
Equilibrium
Frequency =
l
Period
l
=
2 cm = 13 µs
1 cm
2 cm
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indicated by Eq. 5.1. Higher intensities increase
sensitivity and penetration but also increase the
risk of causing thermal bioeffect (an increase in
temperature that can cause protein denaturing
leading to cell death) and mechanical bioeffects
referred to as cavitation (a rapid phase transition
which creates cavities that can implode within
the tissue).
63
Intensity Power/Area
Propagation Velocity
(5.1)
6.5 µs
and Roundtrip Time
Another important wave parameter is the propagation velocity. The propagation velocity is
the speed at which a wave travels through a
medium and is determined by the properties of
the medium. In essence, more dense tissues tend
to be stiffer (higher bulk modulus) which results
in higher propagation velocities. Ultrasound
machines start with a simplified assumption that
sound travels at 1,540 m/s (1.54 mm/μs). Based
on this assumption and simple application of the
distance equation, the time for sound to travel
1 cm in the body is calculated to be approximately 6.5 μs. Since ultrasound imaging is based
Fig. 5.3 Roundtrip effect and imaging travel time per
centimeter (Reprinted from “Essentials of Ultrasound
Physics: the Board Review Book,” by F. R. Miele, p. 327.
Copyright 2008 by Miele Enterprises, LLC. Reprinted
with permission)
on signal reflection, the roundtrip effect must be
taken into account, which implies that 13 μs are
required for each 1 cm of imaging depth (2 cm
roundtrip travel) (Fig. 5.3). Of course, when the
actual speed of sound varies significantly enough
from the assumed 1,540 m/s, error exists with the
displayed image (portraying structures too shallow when higher than the assumed speed and too
deep when slower than the presumed speed).
highly specular and hence highly angle dependent. For specular reflection, the ideal angle is an
incident angle of 0° (which implies that the beam
direction is perpendicular to the reflecting structure) (Fig.
5.4).
As the angle increases (the beam direction
relative to the reflective surface becomes more
acute), an increasing percentage of the reflected
energy is lost, decreasing the ability to visualize
Reflection
Reflection from tissues is dependent on geometric and acoustic properties. Geometrically, when
the reflecting surface is large and smooth with
respect to the wavelength, a very angle- dependent
type of reflection, specular reflection, occurs.
When the surface is rough relative to the wavelength, the much less angle-sensitive (back) scattering occurs. Finally, when the reflecting
structures are small relative to the wavelength (as
occurs with red blood cells), Rayleigh scattering
occurs. Note that reflection from a needle is
the reflecting structure. Eventually, at large
enough incident angles, the reflecting structure
will not be at all visualized. This change in visualization as a result of angle-dependent reflection
is referred to as anisotropy. Practically speaking,
this fact has ramifications when dealing with
ultrasound-guided needle procedures. The ideal
incident angle of 0° is generally not possible. By
using a shallower entrance approach, the incident
angle remains smaller and the needle is generally
better visualized. The incident angle can be
somewhat manipulated by rocking the transducer
64
%−
21
()
()
ab c
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F.R. Miele
qi=0°
Fig. 5.4 Incident angle (Reprinted from “Ultrasound Physics & Instrumentation,” by F.R. Miele, p. 147. Copyright
2006 by Miele Enterprises, LLC. Reprinted with permission)
qi=20° qi=40°
(what is sometimes referred to as heel-to-toe
rocking) and/or by changing the electronic
B-mode steering (for those systems that offer this
capability). Note that the ability to adequately
visualize tendons and nerves is also very angle
dependent.
Percentage Reflection
and Acoustic Properties
The percentage of the sound wave energy that
reflects at a boundary between two media is
determined by the disparity between the acoustic
properties of the adjoining tissues (Eq.
5.2). This
property is referred to as the acoustic impedance.
The acoustic impedance of a medium is directly
proportional to the density and the propagation
velocity of the medium. At a boundary between
two mediums, when a large mismatch (difference) in acoustic impedances exists, the percentage of the sound energy reflected is high. Of
course, as more energy reflects from a particular
point, there is less energy to insonify the inferior
tissues (lower transmission). When the impedance mismatch is smaller, less reflection (and
hence more transmission) occurs. It is important
to note that higher impedance mediums do not
imply high reflection percentages. As specified in
the equation, the amount of reflection is related to
the difference in impedances such that transitioning from a high to a low impedance, or from a
low to a high impedance, results in large reflections. When the sound wave transitions from a
high impedance medium to another high impedance medium, very little reflection occurs.
Similarly, when the sound wave propagates from
a medium with a low acoustic impedance to
another medium with a low impedance, very little reflection occurs. This fact explains why fresh
thrombus comprised primarily of red blood cells
of the same acoustic properties may not be visualized whereas older clot, especially when
calcified, generally presents with a very echogenic appearance.
ZZ ZZ
()
/+Reflection=
21
2
(5.2)
Application of the reflection equation is useful
in predicting and understanding the acoustic
appearance of various tissue structures. Relative
to the acoustic impedance of muscle, the acoustic
impedance of tendon and bone is quite high,
resulting in very strong reflection (hyperechoic
images). In comparison, fat has a lower acoustic
impedance than muscle and therefore tends to
produce lower acoustic echoes. Acoustically,
blood appears relatively “homogenous” such that
Overall
Frequency (MHz)
Sensitivity (dB)
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65
the blood pool itself typically produces very little
reflection (hypoechoic). Note that when fluidfilled cystic structure is insonified, the boundary
of the fluid represents a very large acoustic
impedance mismatch and hence produces a
strong echo; however, when the fluid itself is
homogenous, no reflection occurs, and hence, no
internal echoes are visualized (anechoic). At the
other extreme from fluids, metallic structures
such as needles and surgical clips represent very
high impedances. Relative to the lower impedance of the surrounding tissue, the high impedance of these metallic objects results in very large
reflection percentages and, hence, very strong
bright echoes.
For ultrasound-guided needle procedures,
visualization of nerves is obviously very important. Sonographically, nerves commonly have a
honeycomb appearance in cross section, which
translates to a striated appearance longitudinally.
The explanation for this appearance, of course,
is explained by the reflection equation (Eq. 5.2)
in conjunction with the properties of nerves. A
nerve consists of a bundle of fascicles. The bundle is encased with the epineurium, and each fascicle is encased by perineural connective tissue.
The nerve fascicles generally appear hypoechoic
surrounded by more hyperechoic perineural connective tissue. The epineurium is also echogenic.
5.2.2 Transducers, Absorption,
and
Resolution
Ultrasound transducers are created from piezoelectric materials. Piezoelectric materials convert
mechanical stress (such as occurs with mechanical vibration from sound waves) into electrical
energy (voltage). These materials also exhibit
“reciprocity” which means that the same crystal can convert electrical energy into mechanical energy. Transducers operate over a range
of frequencies, referred to as the transducer
bandwidth as depicted in Fig. 5.5. This bandwidth allows for imaging at different frequencies without swapping transducers as well as
allowing for different operating frequencies for
different ultrasound modalities (i.e., imaging at
transducer bandwidth
Lower
frequency
Fig. 5.5 Bandwidth (Reprinted from “Essentials of
Ultrasound Physics: the Board Review Book,” by
F.R. Miele, p. 327. Copyright 2008 by Miele Enterprises,
LLC. Reprinted with permission)
MHz while performing Doppler at 3.0 MHz).
5
Higher
frequency
High-frequency ultrasound generally has the
advantage of superior axial (depth), lateral (side
by side), and elevation (slice thickness) resolution. The disadvantage of high-frequency imaging is significantly less penetration as a result of
increased absorption (the conversion of sound
energy into heat within the tissue). Absorption
rates increase exponentially with increasing
transmit frequency and increasing depth, implying that using even slightly higher frequencies
can significantly decrease penetration and signal
strength at deeper imaging depths. Broadband
transducers offer the flexibility without having
to change out transducers of transmitting with
lower frequency for improved sensitivity and
penetration at greater depths as well as transmitting at higher frequencies for improved resolution for shallower imaging.
Transducers can be constructed from single
crystals (referred to as pencil probes, pedofs, or
Doppler only) or multiple elements to form
phased arrays. Phased array transducers consist
of different form factors including:
Sectors, which bring the image to a point designed
for rib access and used generally in cardiac
imaging although also used for transcranial
and, sometimes, abdominal imaging
Curved linears, which produce broad near field
by a curved surface, generally used on the
abdomen and for invasive transducers such as
transrectal and transvaginal
Linears, which have a large flat surface ideal for
contact with relatively flat surfaces such as the
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Time
Fig. 5.6 Linear transducer imaging formats (Reprinted from “Ultrasound Physics & Instrumentation,” by F.R. Miele,
pp. 277–279. Copyright 2006 by Miele Enterprises, LLC. Reprinted with permission)
neck, arms, and legs. Linear transducers have
three different imaging formats including
unsteered, steered, and trapezoidal (Fig. 5.6)
Time
frame frequency, commonly referred to as the
frame rate. The frame rate is one of the predominant determinants of the ability to resolve changes
in time, or the temporal resolution. Higher frame
rates imply better temporal resolution. Higher
5.2.3 Linear Phased Arrays
frame rates can be achieved by minimizing the
time required to produce the frame. Clearly nar-
Since venous imaging is performed primarily
with linear phased arrays, we will emphasize the
functionality of linear arrays. Linear images are
rowing the image and decreasing the depth setting decrease the frame time and improve
temporal resolution.
produced by activating a group of elements,
referred to as the aperture, to produce a single
beam. Generally the same aperture is then used to
5.3 Ultrasound Imaging
receive over time the returning echoes. Based on
the depth setting, the system determines the
5.3.1 Operating Frequency
required receive time, referred to as the pulse
repetition period (PRP) before transmitting the
next line. Recall that 13 μs is required for each
cm of imaging depth such that the PRP is simply
the imaging depth multiplied by 13
μs/cm. Once
the required “listen” time has transpired, the
system activates another group of elements to
produce the next acoustic beam and the process is
repeated until the user-specified region of the
patient is scanned. The frequency at which the
lines can be transmitted, the pulse repetition
frequency (PRF) is equal to the reciprocal of the
pulse repetition period (PRP). Once the
entire region is scanned, the system goes back to
the beginning and repeats the process, creating
the next frame of data. The time required to
acoustically produce a frame is simply calculated
as the time required to produce a line (the PRP)
multiplied by the number of lines that constitute
the frame. The reciprocal of the frame time is the
Changing the transmitted wave parameters
can have significant impact on the image quality and clinical value of the data acquired. As
already mentioned, most modern systems offer
broad bandwidth transducers that allow the user
to select the operating frequency. Typically, the
frequency range for linear transducers is between
about 2.5
MHz and about 15 MHz. This range is
too broad for any one transducer and is generally
covered by either two or three different transducers (i.e., a specific vendor may offer broadband linear transducers with frequency ranges
of 2.5–7, 5–12, and 8–15 MHz). For imaging,
the general rule for operating frequency selection is to choose the highest frequency that still
results in adequate penetration. Following this
rule ensures that resolution is optimized while
still presenting adequate signal above noise
thresholds. Practically speaking, there can only
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be general and not specific rules about appropriate frequency selection for conventional imaging,
since attenuation with depth varies from patient
to patient.
5.3.2 Transmit Power
For deeper imaging, as occurs with deeper vessels on large patients, the transmit power should
be increased. For the superficial venous system,
the transmit power can generally be reduced from
the maximum without compromise of image
quality.
5.3.3 Transmit Focus
A third important system setting which affects
sensitivity is the transmit focus or foci. When
imaging superficially with only one focus, the
focus should be set at the depth of vein or artery
being assessed. For deep veins when inadequate
sensitivity is a potential issue, the focus should
be set just slightly deeper than the vein of interest since, for deeper imaging depths, the highest
beam intensity is generally shallower than the
focal depth. Using multiple foci improves the
lateral resolution and sensitivity over the focal
range and in general is very useful for venous
imaging. The downside of using multiple foci
is that image acquisition slows, decreasing the
frame rate, hence, degrading temporal resolution.
5.3.4 Signal and Gain
The visualization of grayscale signals on the
screen is the result of a complex interaction
between multiple systems and environmental
settings including the receiver gain, time gain
compensation (TGC), compression settings,
monitor contrast and brightness, and the ambient lighting of the room. Before discussing the
approach to setting the system controls, it is
imperative to first discuss the ultrasound signal relative to the light perception of the
human eye. First, there is a need for signal
amplification since even the highest amplitude
signals reflecting from the body are too small
to be adequately visualized. The system control, known as “receiver gain,” is basically a
multiplier that increases the amplitude of all
received signals. TGC is also receiver gain but
differs in how the gain is applied. Unlike the
receiver gain that multiplies all signals, TGC
is depth dependent. Signals that are received
from deeper depths arrive later in time and
generally require more amplification than signals from shallower depths. TGC allows for
the image to be subdivided into zones based
on depth, and then different amplification
applied to each of these zones to compensate
for the varying attenuation.
5.3.5 Signal and Compression
Reflected ultrasound signals span a very large
range of intensities (referred to as the “signal
dynamic range”). Reflections from mediums
such as nerves, tendons, bones, and calcifications
tend to be of very high amplitude as a result of
the large acoustic impedance mismatches relative to the surrounding tissues. Conversely,
reflections from within blood, fluids, and fresh
thrombus tend to be of low amplitude as a result
of the relative acoustic impedance homogeneity.
The dynamic range of these reflected signals
often spans 80 or even 100 dB (a factor in amplitude of 10,000:1 and 100,000:1). The human eye
for a given ambient light is capable of detecting
less than 36
The result is that the dynamic range of reflected
signals from diagnostic ultrasound far exceeds
the dynamic range of the human eye. This fact
illustrates why compression must be employed
in displaying ultrasound images. As the name
suggests, compression reduces the dynamic
range by mathematically remapping the wider
range of signal amplitudes to a smaller range of
signal amplitudes. The problem is, that by compressing the dynamic range of reflected signals,
it is more than conceivable that some information may be lost and an important differentiation
between tissue types may be missed (i.e., an
dB (fewer than 64 shades of gray).
68
Decreased c
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F.R. Miele
Fig. 5.7 Distortion of wave
producing harmonic energy
(Reprinted from “Ultrasound
Physics & Instrumentation,”
by F. R. Miele, p. 664.
Copyright 2006 by Miele
Enterprises, LLC. Reprinted
with permission)
Transmintted wave
Nonlinear response
inability to differentiate a mass from surrounding
normal tissue or an inability to visualize a thrombus from the surrounding blood pool).
In light of the potential to not visualize important signals, ultrasound systems are designed
with multiple compression maps. The names of
the compression function vary relative to system
manufacturers, but this function is generally
called compression, grayscale, dynamic range, or
post-processing. Regardless of the function
name, the various compression settings map the
signal strengths to differing brightness levels.
Since the maps compress the signals nonlinearly,
each setting changes the appearance of the image
grayscale, potentially masking or unmasking a
thrombus or mass visualized or not visualized at
a different setting.
5.3.6 Signal and Ambient
Light
Ultimately, the image is presented on a display
monitor. The monitor should be calibrated relative to the ambient light. In brighter ambient
light, the monitor brightness must be increased so
that low-level signals are displayed with an
appropriate brightness so as to be presented
above the visual threshold.
Compression
Increased c
5.3.7 Harmonic Imaging
The imaging technique discussed up to this point
has assumed conventional, fundamental imaging.
With fundamental imaging, the transmit frequency
band and the receive frequency band are the same.
Since the latter part of the 1990s, second harmonic
imaging has revolutionized ultrasound. For harmonic imaging, the system transmits at a lower frequency band, referred to as the fundamental, and
then receives and processes signals at twice the
fundamental signal frequency, referred to as the
second harmonic. In order to operate over such a
wide frequency range, the transducers used for harmonic imaging require very broad bandwidth.
When processing the returning echoes, the system
applies a filter to “look at” just the second harmonic
bandwidth, explicitly attempting to eliminate signal with frequencies in the transmit bandwidth.
Harmonic signals are generated in tissue
because of a nonlinear response of the tissue to
the compression and rarefaction induced by the
sound waves. During compression, the density
of the tissue molecules increases, resulting in a
slight increase in propagation speed. During rarefaction, the density decreases such that there
is a slight decrease in the propagation velocity. In essence, the sound wave is distorted as it
propagates through the tissue (Fig.
5.7) as some
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fundamental energy is converted into harmonic
energy within the sound wave. Although the generated harmonics are not limited to second harmonics, as of now, ultrasound systems are not
making use of progressively weaker higher order
harmonics. The generation of harmonic energy is
very nonlinear with the beam intensity (related to
the parameter called the mechanical index [MI]).
Slightly decreasing the beam intensity (a lower
MI) results in significantly less harmonic signal
generation. As a result, harmonic imaging is very
sensitive to where the beam focus is placed in the
image as well as to the overall transmit power
used when imaging. The best harmonic imaging occurs with higher transmit power and near
within the focal region (the area about the depth
of the focus).
The primary benefit of harmonic imaging is
a reduction of clutter artifacts. Since harmonic
beams are inherently narrower than beams generated using fundamental imaging, lateral resolution is improved. Additionally, because the beam
intensity is generally low in the near field before
the beam has converged to the focus, harmonic
generation is generally low in the near field.
Since most imaging artifacts result from strong
reflectors in the near field (referred to as clutter
signals), lower-level harmonic signals in the near
field result in fewer imaging artifacts, improving image quality. For deep imaging in which
attenuation dominates, harmonic imaging is not
recommended as harmonic generation drops precipitously with lower beam intensities.
amplitude increasing by a factor “n.” For example, adding nine signals, each with an amplitude
V, would result in a combined signal with
of 0.1
an amplitude of 0.9 V. In contrast, the noise
(which exists in every image) is random from
frame to frame. When the noise from each of the
nine frames is averaged, the noise amplitude
also grows, but at a slower rate. In fact, the noise
generally grows at the rate of the square root of
the number of frames used in the average (√n).
Using the same example, imagine that the noise
level is 1 nanovolt (nV); averaging nine frames
together would result in the noise becoming bigger by the square root of nine, or 3 × 1 nV = 3 nV.
Therefore, the signal grew nine times larger
whereas the noise grew only three times as large,
implying that the signal-to-noise ratio increased
by a factor of 9, 3, or 3 times. In other words,
averaging improves the signal-to-noise ratio by
the square root of the number of samples in the
average.
As already mentioned, in addition to
averaging, compound image varies the image
steering from frame to frame. Recall that specular reflection is very angle dependent. Combining
this fact with the fact that most imaging artifacts
are caused by specular reflection, it should be
clear that compound imaging results in artifacts
tend to “average out.” The net result is that compound imaging generally results in images with
better signal-to-noise (improved sensitivity) and
fewer artifacts.
5.3.8 Compound Imaging
Compound imaging (also known as Sono CT or
Crossbeam Technology) generally results in
improved image quality. There are two primary
components to compound imaging: frame
averaging and varying steering angles between
averaged frames. By averaging frames, the signal-to-noise ratio is improved as long as any
changes in the region of interest occur slowly
relative to the frame acquisition rate. In cases
where the signal is changing slowly, the signal
in each of the “n” frames are in phase such that
adding together the frames results in the signal
5.3.9 Artifacts
Ultrasound imaging is based on a series of
assumptions. Whenever one or more of these
assumptions are violated, artifacts exist. Some of
the many assumptions made by ultrasound
include:
• The speed of sound is 1,540 m/s (equivalent to
13 μs for sound to travel 2 cm, or image 1 cm).
• Sound travels only down to a structure and back
and does not reverberate between structures.
• Sound travels in a straight line such that the
beam path is “straight.”
• Attenuation is never so significant so that deeper
structures are still adequately insonified.
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