Добавил:
Sekretar
kiopkiopkiop18@yandex.ru
t.me/Prokururor I Вовсе не секретарь, но почту проверяю
Опубликованный материал нарушает ваши авторские права? Сообщите нам.
Вуз:
Предмет:
Файл:Ординатура / Хирургия / Библиотека им академика М.И. Перельмана / Книга_5817_Библиотеки_им_академика_М_И_Перельмана-1.pdf
1 Physical and Technical Fundamentals of Ultrasound and Doppler Ultrasound
1
the velocity of sound within the material. Table 1.1
provides an overview of the densities of various tissues
and the sound velocities within these tissues.
The distance between two maximal compressions is
the wavelength λ. The number of oscillations of a
molecule in unit time is the frequency f, which is
measured in hertz (Hz). Their depth of penetration and
resolution capacity are listed in
single oscillation per second, i.e., 1 Hz = 1/s.
Table 1.2. One hertz is a
1.2
igure
F
provides an overview of the frequency range of sound
waves and their applications. Ultrasound waves in the
frequency range of 2−30 MHz are used for diagnostic
medical imaging.
The velocity of a wave is determined by the product
of its wavelength and frequency:
c=λ × f
In medical diagnostic imaging the velocity of sound is
assumed to be independent of tissue, namely a constant quantity of 1540 m/s.
The displacement of an individual molecule from its
resting point in diagnostic imaging amounts to about
−6
2×10
waves is typically 0.6 × 10
mm (0.000002mm). The pressure of sound
5
Pa. At this pressure each
oscillating molecule accelerates appreciably to the
5
order of 10
times gravitational acceleration.
Table 1.1 Densities, sound velocities, and attenuation of
various substances
Substance Density
Fat 0.97 1470 0.5
Bone marrow 0.97 1700
Muscle 1.04 1568 2
Liver 1.055 1540 0.7
Brain 1.02 1530 1
Bone (compact) 1.7 3600 4−10
Water (20°) 0.9982 1492 0.002
Air (sea level) 0.0013 331
Table 1.2 Values for diagnostic ultrasound
Transmitting
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)
(g/cm
Depth of
penetration (cm)
Sound
2
)
velocity
(m/s)
Lateral
resolution
(mm)
Attenuation
(db/
MHz cm)
Axial resolution
(mm)
Fig. 1.2 Frequency ranges of
sound waves and their applications.
12
10 Hz
9
10 Hz
20 kHz
16 Hz
0 Hz
Ultrasound
Sound
Infrasound
Upper frequency limit for mechanical waves, as the wavelengths
now become smaller than the distance between molecules
Medical diagnosis 2–30 MHz
Materials testing 50 kHz – 1MHz
Ultrasonic cleaners 25 kHz – 50 kHz
Sonar equipment, naval positioning 16 Hz – 16 kHz
Seismological waves
The Generation of Ultrasound
Ultrasonic waves can be generated in a variety of ways,
the mechanical method being the simplest. Sounding a
4
tuning fork of less than 2 mm generates ultrasonic
waves with a frequency up to 200 kHz. For engineering
use such as cutting, drilling, or milling of very small
parts (e.g., to divide semiconductors) ultrasound is
generated by magnetostriction. In this method a ferromagnetic substance is placed in a magnetic field with
alternating polarity, generating frequencies of up to
about 50 kHz at considerable sonic intensities.
In 1880 the Curies discovered the piezoelectric ef-
fect, which is used for the generation of ultrasonic
waves in medical diagnostic imaging. If pressure is exerted on an ionic crystal and an elastic deformation in a
defined direction is then imposed on the crystal, the
internal charge shifts. Electric potentials result on its
surface, negative on one side, positive on the other. The
voltage increases as the pressure increases. If, con-
versely,an electric potential is applied to the surface of
a piezoelectric crystal, the latter will become elon-
gated or shortened according to the direction of the
voltage. If an alternating potential is applied, the crystal begins to oscillate. Quartz and tourmaline are especially active piezoelectric substances. The piezoelectric
materials used to build current ultrasonic probes are
made of scintered ceramic such as barium titanate.
igures
F
image of such a piezoelectric ceramic and a schematic
view of the multilayered arrangement used in its structure.
1.3 and 1.4 show an electronmiscroscopic
Physical Effects
Basic Concepts
Fig. 1.3 Electronmiscroscopic photograph of a piezoelectric
ceramic.
Fig. 1.4 Diagram of the multilayer structure of a piezoelectric 컄
ceramic.
Physical Effects
Reflection and Refraction
The propagation of sound in tissue follows the laws of
optical waves. At the interface between two tissues of
different densities there is a sudden change in impedance. Impedance is defined as the resistance Z to
sonic waves, calculated by the product of sound velocity × density. At such an acoustic interface, part of the
incident sound wave is reflected (reflection), while
another part is refracted and continues into the tissue
1.5). When the incidence of the
(transmission) (
sound wave on an interface is vertical, the reflection
gradient (R) is calculated by the formula:
Fig.
30 µm crystal layer
+
–
+
–
3 µm electrode
+
–
Reflection
Refraction
α
α
S
Fig. 1.5 Reflection and transmission of a sound wave at an
acoustic interface (S = transducer; Z = impedance).
Tissue1
Z1
Tissue 2
Z2
Acoustic
interface
5
1 Physical and Technical Fundamentals of Ultrasound and Doppler Ultrasound
1
The reflection gradient at the transition from liver
tissue (Z
can be calculated to be 0.0008. This means that less
than a one hundred thousandth of the propagated
sound energy is reflected at this boundary interface. At
the transition from fatty tissue (Z
(Z
2
lated. Over 99 % of propagated sound energy is reflected here. Thus, all the ultrasound is reflected and
behind this interface no more ultrasonic energy is
available to penetrate further into the tissue. For this
reason segments of gas-filled intestine and lungs cannot be reached by ultrasound, and the layer of air between an ultrasonic transducer and skin must be eliminated by contact gel.
= 1.66×105) to renal tissue (Z2= 1.63 × 105)
1
= 1.42 × 105) to air
= 43) a reflection gradient of 0.9987 can be calcu-
1
Scattering
As a rule the surfaces separating tissues are not smooth
but rough or diffuse. Sound waves are not reflected
directionally at diffuse interfaces but are scattered in
F
ig.
the form of a spherical wave (
smaller than the wavelength cause mainly scatter.
Tissue structures exceeding the wavelength cause reflection. Scatter echoes generate the typical textured
pattern of parenchymatous organs.
1.6). Tissue structures
Interference
When two or more sound waves arrive at the same
time, two waves in different phases may meet (i. e., the
compression phase of one wave may coincide with the
rarefaction phase of the other), and they may weaken
each other. Similarly, waves in the same phase may
coincide and reinforce each other. This phenomenon is
called interference, and the spatial distribution of
areas of weakening and reinforcement are called interference patterns. Interference patterns significantly
determine the appearance of an ultrasound image.
Diffraction
When the straight propagation of a sound wave is
impeded, for example, by the edge of an object, the
waveis bent (diffracted) around such an edge. Thus the
sound waves reach the space behind the object, which
1.7).
ig.
would normally lie in its shadow (
F
Absorption
The energy of a sound wave diminishes in the direction
of itspropagation. The internal friction of the oscillating
molecules is transformed into heat, i.e., the sound wave
is absorbed. The values of attenuation listed in
may be roughly averagedto 1 dB/(MHz cm). Absorption
is dependent on frequency.On the one hand, high ultrasound frequencies are useful because their short
wavelengths provide precise localized resolution; on
the other hand, it is also necessary to examine organs
lying deep below the surface, and for this examination
lower ultrasound frequencies with their longer
wavelengths are more suited because they attenuate
less. At a frequency of 10 MHz the attenuation is 10dB/
cm; at a frequencyof 3 MHz the attenuation is 3 dB/cm.
Assuming an initial output of 100 dB, the depth of penetration at a frequency of 10 MHz can be calculated to be
5 cm (10 cm total path) and at a 3 MHz frequency the
penetration would be 17 cm (33 cm total path).
Table
1.1
Diffuse
interface
S
S
Tissue 1
Z1
Tissue 2
Z2
6
Fig. 1.6 Spherical scattering of a sound wave at a diffuse interface (S = transducer; Z = impedance).
Fig. 1.7 Diffraction of a sound wave at a diffracting edge
(S = transducer).
Generating the Image
Generating the Image
Pulse-Echo Procedure
Almost all ultrasound procedures used in medical diagnosis are based on the so-called pulse-echo procedure. A brief electrical impulse is applied to the piezoelectric crystal of the transducer, and this impulse is
transformed into an ultrasonic pulse by the piezoelectric crystal. The duration of the pulse is about 1 µs. The
transducer is then switched to act as a receiver. The
sound wave penetrates the tissue and is reflected from
the internal boundary interface. The part of the sonic
pulse returning to the piezoelectric element is called
an echo; it elicits an electric impulse in the element.
The time differences between the emission of the ultrasonic pulse and the reception of each echo are then
measured. The product of the ultrasound velocity c and
the time difference t gives the distance z covered by
the ultrasonic pulse. Dividing this number into 2 gives
the exact position of the reflecting structure with respect to the transducer:
Z = ct/2
For a time difference of, for example, 0.13ms calculation shows a distance of 10cm between transducer and
reflector. Depending on the construction of the transducer, 3000−5000 ultrasonic pulses a second may be
emitted. At the same time echoes are captured, calculated, and transformed into images.
Time Gain Compensation
Because some of their sound is absorbed, energy
(amplitude) from echoes traveling over a longer period
of time from deeper tissues is lower than that of echoes
with shorter travel times. Thus, interfaces having equal
reflection gradients will emit signals of different
amplitudes according to their depth. To equalize the
representation of signals, those received at the element after a longer travel time are proportionally
amplified. Such amplification is known as time gain
compensation (TGC), or depth gain compensation
(DGC), and can be controlled by the user of the ultrasound apparatus.
Basic Concepts
A-Mode
The simplest and also the oldest imaging procedure
used in diagnosis is the A-mode procedure. The amplitudes of the electric signals generated at the transducer are displayed on a cathode ray oscilloscope, proportional to the distances of the boundary surfaces in
the tissue being examined (Fig. 1.8). Because it represents amplitude, this procedure is known as amplitude
mode, or A-mode. A-mode procedures provide only
one-dimensional information. Today they are still used
in ophthalmology (to determine the thickness of the
cornea) and otorhinolaryngology (for noninvasive examination of the nasal sinuses).
Fig. 1.8 A-mode scan. The ampli-
tude of the electric signals
generated at the transducer are displayed on a cathode ray oscilloscope.
S = transducer
RV = right ventricle
LV = left ventricle
AO = aorta
LA = left atrium
LV
S
RV
AO
LA
7
1 Physical and Technical Fundamentals of Ultrasound and Doppler Ultrasound
1
B-Mode
In contrast to A-mode, in this procedure amplitudes
are represented not as deflections (peaks) but as bright
spots on a monitor. The brightness steps of these spots
are proportional to the amplitudes of the electric signals and hence those of the echoes. The stronger the
signal, the brighter is the point on the image. This procedure is known as brightness mode, or B-mode. In
current ultrasonic systems about 256 degrees of
brightness can be represented in shades of gray (grayscale display), though the human eye can only distinguish 80 shades of gray. The individual brightness
spots are arranged in a straight line, and the sound
beam is now shifted before sending a new pulse. If the
new lines acquired by this procedure are displayed
corresponding to their localization, a two-dimensional
Fig.
cross-sectional image is obtained (
needed to display a line at a depth of, for example,
15cm (cf. Pulse-Echo Procedure). For the display x of a
width of 5 cm, an interval between lines of ∆x of 1 mm,
and the known ultrasound velocity c of 1540ms
scan time T of 10 ms can be obtained by the formula:
T = (2zx)/(C∆x)
1.9). 0.2 ms are
−1
, the
M-Mode
In contrast to B-mode display, in M-mode (motion
mode), the sound beam is not shifted, but is kept in a
fixed position over the organ to be examined. The individual lines are again displayed side by side on a time
axis. By this means it becomes possible to visualize
movements as they occur in the imaged tissue
ig.
F
1.10). At a depth of 15 cm of the displayed tissue
(
the scan time is 0.2 ms, i. e., an image can be generated
about 5000 times a second. Thus, even very rapid
movements, for example, of the heart valves can be
displayed. Since the scales of the image and of the time
axis are known, movements, velocities, and acceleration can be accurately measured. Hence M-mode is of
great value in echocardiography.
From this may be calculated a repetition frequency for
the image of 100 Hz, i. e., 100 separate images per second. It is therefore practicable to generate an image in
real time.
S
RV
AO
LV
LA
S
8
Fig. 1.9 B-mode scan. Generation of a two-dimensional cross-sectional image. S = transducer, RV = right ventricle, LV = left ventricle, AO = aorta, LA = left atrium
Fig. 1.10 M-mode scan. The course of
movements in the insonated field is
displayed as it develops over time.
S = transducer
RV = right ventricle
LV = left ventricle
AO = aorta
LA = left atrium
The Sound Field
S
t
RV
AO
The Sound Field
Resolution
Resolution describes the smallest interval between
two structures that allows them to be displayed on the
monitor as two distinct objects. Resolution is
measured in millimeters. Axial resolution in the direction of the sound beam must be distinguished from
lateral resolution, which occurs in a plane at right an-
gles to the axis of the sound beam (Table
olution is determined by the pulse length of the ultrasound beam. It usually measures one or more
wavelengths. Higher resolution can be attained by
using higher ultrasound frequencies with their shorter
wavelengths. However, as demonstrated when discussing absorption (p. 6), the penetration of ultrasound diminishes with increasing frequency. Hence it
is not possible to avoid the use of lower frequencies to
display deeper tissue structures. Lateral resolution is
1.2). Axial res-
LV
LA
determined by the width of the ultrasound beam, i.e.,
it is proportional to the diameter of the sound beam.
F
e 1.11 is a schematic representation of a simple ul-
igur
trasonic transducer. The sound field is composed of a
narrow focused near field and a divergent far field. Precise scanning is only possible in the focused near field.
To improve the resolving power of the ultrasonic beam
it must be focused.
Focusing
An ultrasonic beam can be focused in various ways.
The simplest is the use of an acoustic lens (
since essentially the physical laws of wave optics are
valid for the spread of sound (cf. Reflection and Calcu-
p.
lation,
5). Hence the ultrasonic beam is maximally
focused at a fixed point. This point is known as the
focal point. The element may be given a concave
Fig. 1.12),
Basic Concepts
Fig. 1.11 Diagram of the sound
field of a simple ultrasonic trans-
ducer.
Near field Far field
Acoustic lens
Focal point
9
1 Physical and Technical Fundamentals of Ultrasound and Doppler Ultrasound
Fig. 1.13 Generation of a concave wave front by time delay of
individual crystal elements in an
array.
Array of elements
10
1
shape. This is known as internal focusing and is used
in the single element systems of mechanical sector
scanners.
The ultrasound beam is focused electronically in
order to reach as variable a depth of focus as possible.
Current ultrasonic transducers are built in the form of
arrays, consisting of a greater number of individual elements arranged next to each other. The total number of
individual elements amounts to between 60 and 256,
according to the type of transducer. Several of these individual crystal elements are grouped together to send
and receive ultrasound rays. A concave wave front can
be generated by regulating the timing of individual
Fig.
elements in the group (
converges to a focal point. A sound beam generated in
this way is maximally focused at this point and pro-
1.13). Such a wave front
Scanning Procedures
Principle of Operation
The principle of operation of modern scanners is
shown in Figure
dividual elements in an array are grouped together to
generate an ultrasound beam. Such a group generates
an ultrasound pulse and also receives the returning
echo signals. When an element on the left side is
switched on, while an element on the right side is
switched off, the new group of elements generates
another ultrasound ray that has been shifted by the
width of one element. The linear density can also be increased by varying the breadth of the group: First a
group of elements generates an ultrasound pulse, then
one element, for example, on the left side, is switched
off, while on the right none is switched on. The axis of
the next ultrasound beam is now shifted relative to the
previous group by the width of half an element. If, now,
an element on the right side is switched off, while
none is activated on the left, the number of lines of
sight has been increased by a factor of 2. On the other
hand, the time required to scan a complete image increases, i. e., the image repetition frequency is halved.
1.14. As described above, several in-
vides high lateral resolution. The focal length can be
shifted by varying the width of a group and the timing
of the elements. In this way the user can set the focus
and with it the point of maximal resolution to the area
of diagnostic interest. In modern ultrasound systems it
is possible to set several such transmission foci simultaneously. It should be noted that this requires that
each sound beam must be repeated for each focal
length, and this reduces the repetition frequency of the
image. The focusing of the received beam has a special
feature: Since the echoes from deeper regions at the
transducer arrive later than those from closer regions
it is possible in practice to let the receiving focus shift
to the deep area. This procedure is known as dynamic
receiving focusing. The groups used for focusing vary
in width between 8 and 128 elements.
Array of elements
Direction of scan
Fig. 1.14 Mode of operation of a modern ultrasound scanner.
Scanning Procedures
Various scanning devices are described in what fol-
lows.
Linear Array Scanner
In a linear array scanner, also known as a parallel scanner, the crystal elements are arranged in a straight line
ig.
F
1.15). The individual ultrasound beams run in par-
(
allel lines, generating a rectangular cross-sectional
image. The resolution is more or less equally good over
the whole depth being displayed. The number of elements in a linear array scanner varies between 60 and
196, with a width of each element of 1−4λ. The
frequency range of linear array transducers lies between 5 MHz and 13 MHz. An acoustic lens is generally
used to focus a linear array scanner in a plane trans-
verse to the direction of the sound beam.
Curved or Convex Array Scanner
Curved or convex array scanners are effectively a
special type of linear array scanner. The mode of operation corresponds to that of a linear array scanner. They
differ in the curved arrangement of their elements,
1.16).
which generates a pie-shaped sound field (
Since the density of the lines of sight diminishes at a
distance from the transducer, lateral resolution is reduced with increasing depth. Typically a convex array
scanner has more than 96 elements. Its radius varies
between 25 mm and 80 mm and the frequency ranges
between 3 MHz and 7 MHz. Most often the sound field
extends over an angle of 60−90°.
Fig.
Crystal elements
Basic Concepts
Fig. 1.15 Linear array or parallel scanner. The crystal elements
are arranged side by side in a straight line.
Crystal elements
Sector Scanner
This differs from the curved array scanner by having a
smaller radius (쏝25 mm). This results in a smaller application surface, a narrow near field, and an angle of
departure of 쏜90°. These properties are utilized when
imaging through a small sound window such as the intercostal spaces in echocardiography or in transvaginal
sonography.
Phased Array Scanner
The arrangement of elements in a phased array scanner is the same as that in a linear array scanner.
However, instead of a group of elements, all the elements take part in the generation of a scanned line of
sight.
왘 It is possible to generate a wavefront running at an
angle to the surface of the transducer surface by
Fig. 1.16 Curved or convex array scanner. The crystal arrangement is curved, generating a pie-shaped sound field.
using a time-delayed firing of the elements
F
1.17).
ig.
(
왘 A pie-shaped sound field is generated by changing
the settings. Phase array transducers have a small
contact area of 12−20 mm using between 64 and
128 elements.
왘 The sector angle is between 80° and 90°. The
frequency range is between 2 MHz and 7 MHz.
These transducers are very costly because of their
complex electronics and are used chiefly in cardiology
and transcranial sonography.
11
1 Physical and Technical Fundamentals of Ultrasound and Doppler Ultrasound
Mechanical Sector Scanners
1
T2 T1
W1
W2
Fig. 1.17 Phased array scanner. By using time delay the
generated sound waves are made to run at an angle to the
transducer surface (T = point in time; W = sound wave front).
Single element
In contrast to the electronic phased array scanner, a
mechanical sector scanner requires little investment in
its regulation and signal processing. The resulting
favorable price−performance ratio has allowed mechanical transducers to compete with electronic transducers in low-cost systems. Mechanical sector scanners are divided into rotary transducers and wobbler
transducers.
Rotary Principle
Figure 1.18 illustrates schematically the operation of a
rotor transducer. Three to five single elements are usually arranged on a rotor at equal angular distances. A
motor in the handle turns the rotor. The element rotating past the sound window is activated and covers a
sector-shaped sound field. The next element then rotates past the window and generates a second image.
12
Single element
Fig. 1.18 Diagram of the operation of a rotary transducer.
Three to five individual elements mounted on a rotor at equal
intervals generate a pie-shaped sound field as they rotate past
the sound window.
Fig. 1.19 Wobbler transducer. A single element defines a sector between 60° and 100° by angulating to and fro.
Ultrasound Artifacts
Wobbler Principle
In a wobbler transducer a single element angulates
back and forth, covering a sector of between 60° and
F
ig.
100° (
considerably less energy is required to regulate it. In
addition, the sector angle can be adjusted, in contrast
to a rotary transducer. Because both types of transducer use only one element at a time, their focus is al-
ways fixed.
1.19). Since only a single element is used,
Annular Phased Array Transducer
The annular phased array transducer works on the
wobbler principle, and combines a mechanical and
electronic scanner. It consists of an array of ring ele-
Ultrasound Artifacts
Ultrasonic imaging is plagued with far more artifacts
(imaging errors) than other diagnostic imaging procedures such as computed tomography (CT) or magnetic resonance imaging (MRI) scans, because the assumed values of parameters, such as sound velocity,
straight line sound propagation, attenuation, often de-
viate from actual values. Artifacts may also be due to
inadequate instrument settings. However, experience
has shown that some of the important artifacts described below may be diagnostically useful and can
provide additional information about the properties of
the examined tissue.
ments arranged concentrically inside each other instead of single elements. Each individual ring can be
regulated separately. This allows variable focusing in
two dimensions (cf. Focusing, p.
9).
Disadvantages of Mechanical Scanners
Regardless of their principle, mechanical sector scanners are subject to wear and require maintenance.
Moreover, rapid switching between scan modes (Bmode, M-mode, Doppler) is impossible because of inertia. In general, real-time display of B-mode/M-mode
or B-mode/Doppler cannot be performed.
wave leaving, for example, a cyst retains nearly as
much energy as it had when entering it. The area behind the cyst therefore appears brighter than the surrounding tissue (
in differential diagnosis.
Fig. 1.21). Such an artifact can be used
Basic Concepts
Distal Acoustic Shadowing
One of the most commonly found artifacts is distal
acoustic shadowing. When meeting strong reflectors,
i.e., structures with a resistance to sound waves that
greatly deviate from that of the surrounding tissues
(e.g., air) or structures that strongly attenuate sound
energy such as bone or calculi, the greater part of the
ultrasound energy is reflected or absorbed (cf. Reflection and Calculation, p.
tor obviously less energy is available than in the surrounding tissue. This phenomenon is called an acoustic
shadow (
Fig. 1.20).
5). Behind such a strong reflec-
Dorsal Sound Amplification
With optimal TGC (cf. p. 7) the phenomenon of dorsal
sound amplification may be observed behind areas
offering weak attenuation. In hollow fluid-filled spaces
ultrasound is subject to less reflection and absorption
than in the surrounding tissues. Thus, an ultrasonic
Fig. 1.20 Distal acoustic shadowing. Mode of generation (left)
and simplified diagram of the ultrasound display (right).
13
Fig. 1.21 Dorsal sound amplification. Mode of generation
(left) and simplified diagram of the ultrasound display (right).