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Файл:Ординатура / Хирургия / Библиотека им академика М.И. Перельмана / Книга_5786_Библиотеки_им_академика_М_И_Перельмана.pdf
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Near field Far field
Transducer Lens BW
F
2r
Side lobe
Main lobe
Fig. 1.10 Schematic illustration of an ultrasound beam emitted from a transducer with a
circular surface with a diameter 2r. The focal distance is set at F and the beam width (BW) is
the effect of the focusing.
The solution is to use high frequency if we are looking at structures close to
the transducer and low frequency if we are looking at structures further away.
Let us go back to the ultrasound beam. Ideally, one would like the ultrasound
beam to be thin and round, and shoot into the tissue along a straight line, hitting
structures which cause echoes that return to the transducer along the same line.
Then only structures in the thin path of the beam would cause echoes. We have
learned that this is not so. The beam has a near field where we may manipulate
the beam and a far field where the beam diverges due to diffraction, where it is
not possible to manipulate the beam. The beam has a main lobe and side lobes
(Fig. 1.11). The side lobes may be considered ‘skirts’ around the main lobe body.
When such a complex beam is shot into the tissue, all the structures hit by the
Physics and instrumentation
Fig. 1.11 Schematic sketch of an ultrasound beam, demonstrating the main lobe and the
side lobes.
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main lobe and structures which in reality are located on the side of the main lobe,
but within the side lobes, will cause echoes to be returned to the transducer and
be displayed along the centre of the imaginary line through the centre of the main
lobe. This will cause a ‘smear-out’ effect of the image.
The presence of side lobes in addition to the main lobe reduces the quality of our
image. Structures outside the main lobe will be picked up by the side lobes and on the
final image they will be displayed along the centre line through the main lobe.
Improving the overall beam quality is accomplished through the focusing process which may be achieved in various complex ways. One technique, dynamic
focusing, may help us understand the principle of focusing. One submitted pulse
may cause many returned echoes which hit the transducer surface over a time
period, depending on how far the echoes have travelled on their way down to the
various reflectors and then back. Focusing is a process that may be done on the way
out and on the return of echoes. Focusing on the returned echoes is always done.
Since we know when a pulse has been transmitted, we may focus the returned
echoes by changing the focus level in the tissue at certain time intervals following
the transmission of the pulse. This will cause echoes, which originate from a depth
Ultrasound in obstetrics and gynaecology
of, for example, 2, 4, 6, 8 and 10 cm, to be focused separately on return. Thus, the
focusing process will affect the area between 2 and 10 cm, in the example above.
Additionally, we may focus our area of interest especially on the way out to
obtain the highest image quality possible in the specific area where we are looking. Arrows along the side of the image indicate the manually set foci. Optimum
quality is usually achieved employing two to three foci in the area of interest.
The process of directing the focus of our beam to the area we are looking is one
of the most important manual adjustments we make during ultrasound scanning.
Unfortunately, focusing is one of those manual adjustments which are most often
forgotten, a practice that exemplifies a lack of technical understanding of the person
performing the scanning.
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RESOLUTION
To be able to interpret our image, define discrete structures and make precise
measurements on an ultrasound image, we need to understand the basic principles of resolution.
Resolution is defined as the smallest distance we can have between two structures
and still be able to distinguish them as two separate structures.
On a two-dimensional ultrasound image, we have an axial plane, a lateral plane
and an elevation plane (Fig. 1.12). The resolution in these three different planes
is determined by various physical laws that we have to understand to optimize
the adjustment of our machine settings, select the best transducer for our purpose
and make measurements as precise as possible.
The axial resolution may be called the range resolution or the radial resolution.
The resolution in the axial plane is the best of the three. The axial resolution is
mainly determined by the length of the transmitted pulse. A ‘pulse’ always consists of a few oscillations in spite of effective damping factors. The absolute length
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Transducer
Axial plane
Azimuth plane
Elevation
plane
5 MHz
2.5 MHz
5 MHz
2.5 MHz
5 MHz
2.5 MHz
Fig. 1.12 The three planes on an ultrasound image: the axial, the azimuth and the elevation
plane.
Physics and instrumentation
Fig. 1.13 In the upper part, a 5 MHz pulse is shown travelling towards a target, which may
be a blood vessel. The pulse is short enough to be able to hit the anterior and posterior
walls separately, thus two separate echoes will be reflected and make two separate dots on
the screen when they hit the transducer. Below, the 2.5 MHz pulse is longer and the echoes
from the anterior and posterior walls of the vessel will overlap, so only one large dot will be
displayed on our screen. The 2.5 MHz pulse was not able to resolve the two vessel walls as
two separate structures.
of a pulse may therefore be reduced by increasing the ultrasound frequency. The
principle of the axial resolution is demonstrated in Figure 1.13. The total pulse
length of a 5 MHz pulse is typically shorter than that of the 2.5 MHz pulse.
A good axial resolution requires a short pulse. Several factors may contribute to a
short pulse: one of them is the wavelength. A high frequency (i.e. short wavelength)
will make the pulse relatively short and improve the axial resolution.
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The lateral resolution affects measurements across the azimuth plane, which is
perpendicular to the direction of the beam. The lateral resolution is governed by
different physical laws from the axial resolution and is poorer than the axial resolution. Among the factors that affect the lateral resolution are the quality of the
beam and the size of the side lobes (see Fig. 1.11). In the process of optimizing
the beam quality, the aim is to have a thin main lobe and small side lobes.
The lateral resolution perpendicular to the direction of the beam is poorer than the
axial resolution. Measurements made in the axial direction are more precise than
those made across the image perpendicular to the beam.
MEASUREMENT
Generally, when we measure a distance in the axial direction, we put one electronic
calliper on an echo and move the next calliper to another echo to assess the distance
between the two. However, we are not actually measuring the distance between the
two, but rather the time it takes for a pulse to travel from the transducer to the structure closest to the transducer and to the structure further away. So when we measure
Ultrasound in obstetrics and gynaecology
a distance, our calculations are based on time rather than on a physical distance.
We have to take into account that of the two, axial resolution is better than lateral resolution. If we measure in the plane perpendicular to the beam, the beam
quality will influence our measurement. A relatively thick beam will make the
endpoint of a structure appear blurred and make the distance between two points
appear slightly larger than in reality. This phenomenon has a consequence for the
measurements across the screen, for example the occipitofrontal diameter of the
skull and even the femur length.
8
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TIME GAIN COMPENSATION
When a pulse propagates through the tissue, it will gradually lose its energy. This
loss is caused mainly by power absorption and to a smaller extent by reflection, scattering and geometric spread. This process takes place as the pulse travels away from
the transducer and as the echo is on its way back to the transducer. The absorption
of ultrasound energy increases with increasing frequency. The attenuation causes
the reflected echoes from structures deep in the tissue to be weaker than those
emerging from nearby structures. If we do not compensate for this phenomenon,
our image will appear imbalanced (Fig. 1.14). The speed of sound in the human tis-
sue is constant; the echoes emerging from the deeper areas arrive later than those
from the upper structures. Thus, we may compensate for the loss of power from
the late-arriving echoes by inserting a time variable gain in the receiver amplifier.
This is called time gain compensation (TGC). The basic TGC is preset in modern
machines, but we may have to adjust manually to fine-tune our image. Usually it
is possible to make an overall adjustment of the gain as well as adjustments affecting the local area ranging from the near to the far field of the image. The setting of
the TGC also affects our measurements and it is an important part of the training
to learn how to set it correctly. A TGC adjusted too high will produce blurry edges
and measurement of distance between structures will be longer than in real life.
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Fig. 1.14 A section through the planum biparietale. The area close to the transducer
is correctly adjusted while the distal area has hardly visible low-energy echoes as a
consequence of the insufficient compensation for the attenuation of sound emerging from
the deeper sections of the tissue.
The fine-tuning of our image using the TGC is one of the most important
adjustments we make. The grey-scale level of the image ought to appear well balanced from the upper to the lower part of the image. The adjustment must aim
at achieving the full register of grey tones between the black areas and the white
highlights. The setting of the TGC has an influence on our measurements.
Physics and instrumentation
ARTIFACTS
Artifacts in ultrasound imaging may be distortions or any form of incorrect appearance affecting an image and giving misleading information as we try to interpret from
the image. The imaging process using ultrasound technology may cause numerous
artifacts that we have to be aware of. Some of the main artifacts are as follows:
Edge shadows
•
Attenuation shadows
•
Enhancement
•
Reverberations.
•
EDGE SHADOWS
In obstetrics, edge shadows are mainly observed during scanning of the fetal
head. When the sound enters a round structure such as the fetal head it emerges
from tissue with a velocity of 1540 m/s through the bone of the fetal skull that
has a sound velocity of ≈3000 m/s. The sound will then be refracted and leave
a shadow-like impression on both sides of the fetal skull (Fig. 1.15).
ATTENUATION SHADOWS
Bone absorbs ultrasound and the echo amplitude will then be reduced behind
ossified structures. This is frequently observed during fetal heart scanning when
the image of the heart may be in the shadow of the ribs or the vertebrae. In
gynaecology, dense structures such as myomas may to a lesser degree reduce the
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Fig. 1.15 Edge shadows. On both sides of the fetal skull, the ultrasound beam is refracted
and then leaves a shadow below.
amplitude of the sound. Such shadows may give us information about the structure
that is causing the shadow.
Ultrasound in obstetrics and gynaecology
ENHANCEMENT
Enhancement is the opposite of attenuation shadow. The phenomenon may be
seen behind cysts (Fig. 1.16). This artifact may also be used to characterize the
structure causing the enhanced area.
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Fig. 1.16 Simple cyst demonstrating the enhancement artifact. The sound that is passing
through the cyst is not attenuated in the same degree as the sound passing through the tissue
on the right and left side of the cyst. Therefore, the amplitude of the sound immediately
below the cyst is higher than on the sides and consequently it looks as if the area below the
cyst has been enhanced by selectively turning up the time gain compensation.
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Object Image
Main echo
Reverberation
Reverberation
Main echo
A
B
ImageObject
REVERBERATIONS
The artifact referred to as reverberation or multiple reflections is common and
may distort the image in several ways. The basic principle of making an image
using sound is to send a pulse, wait for the pulse to return as an echo and then a
dot is put on the screen corresponding to the time the pulse has taken to travel
on its way down to the reflecting structure and back again. A pulse may also
be reflected back and forth between interfaces before returning to the transducer. The extra travel time this process takes will cause the false echoes to arrive
later than echoes emerging directly from the original structure so that then several lines on the image may present themselves as copies of the original (Fig.
1.17). Such reverberations may easily be recognized. Layers of fat may also cause
reflections and reverberations in the image that presents itself as a diffuse cloud
of noise and is thus not so easy to recognize as artifacts. The lower mechanical
impedance of sound in fat (sound velocity ≈1420 m/s) and muscle tissue (sound
velocity ≈1560 m/s) may cause reverberations.
Reverberations may be complex in their appearance and not always easy to
detect. Using a curved array transducer may reduce the effect of reverberations.
The echoes are scattered out of the field, which causes the curved array to have
a good near-field view.
Physics and instrumentation
Fig. 1.17 Two examples of reverberations. In the upper panel (A) the main echo
schematically is represented by a blood vessel. A ‘copy’ of these echoes may be found as
reverberations at a lower level. Fatty tissue may also cause reverberations which may show
up as a diffuse haze (B).
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References
1. Bom N, Lancée CT, v Zwieten G, Kloster
FE, Roland J. Multiscan echocardiography
I. Technical description. Circulation
1973;48(5):1066–1074
2. Desch CH, Sproule DO, Dawson WJ. The
detection of crack in steel by means of
supersonic waves. J Iron Steel Inst 1946;
153:319
3. Donald I, Wicar WA, Brown TG.
Investigation of abdominal masses by pulsed
ultrasound. Lancet 1958;1:1188
4. Firestone FA.The supersonic reflectoscope,
an instrument for inspecting the interior of
the solid parts by means of sound waves.
J Acoustic Soc America 1946;17:314
5. Edler H, Hertz CH. The use of ultrasonic
reflectoscope for the continuous recording
of movements of heart walls. Kgl Fysiograph
Ultrasound in obstetrics and gynaecology
Saellskap Lund Förh 1954;40:23
6. Edler I. Ultrasound cardiography. The
diagnostic use of ultrasound in heart disease.
Acta Med Scand 1955;308(suppl):32
7. Howry, DH, Bliss WR. Ultrasonic
visualisation of soft tissue structures of the
body. J Lab Clin Med 1952;40:579
8. Jago JR, Whittingham TA, Heslop R. The
influence of ultrasound scanner beam width
on femur length measurements. Ultrasound
Med Biol 1994;20(8):699–703
9. Kratochwil A. Ultraschalldiagnostik in
Geburtshilfe und Gynäkologie. Georg
Thieme Verlag, Stuttgart, 1968
10. Krause W, Soldner R.
Ultraschallbildverfahren (B-Scan) mit hoher
Bildfrequenz für medizinische Diagnostik.
Elektromedica 1967;4:1
11. Sundén B. On the diagnostic value of
ultrasound in obstetrics and gynæcology.
Acta Obstet Gynaecol Scand 6(suppl):114
12. Wild JJ, Reid JM. Application of echoranging techniques to the determination
of structure of biological tissues. Science
1952;28:226–230
Further reading
Angelsen B. Ultrasound Imaging. Waves, Signals, and Signal Processing. Basic Principles, Wave
Generation, Propagation, and Beam forming in Homogenous Tissue. Vol I. Emantec, Trondheim,
Norway, 2000. www.ultrasoundbook.com
Angelsen B. Ultrasound Imaging. Waves, Signals, and Signal Processing. Propagation and Scattering in
Homogenous, Nonlinear Tissue with Contrast Agent. Imaging and Doppler Measurement. Vol II.
Emantec, Trondheim, Norway, 2000. www.ultrasoundbook.com
Hatle L, Angelsen B (eds). Doppler ultrasound in cardiology. Physical principles and clinical
applications. Lea and Febiger, Philadelphia, 1986
Kremkau FW. Diagnostic ultrasound. Principles and Instruments, 7th edn. WB Saunders,
Philadelphia, 2006
Maulik D (ed). Doppler ultrasound in obstetrics and gynecology. Springer, New York, 1997
Woo J A short history of the development of ultrasound in obstetrics and gynecology.
www.ob-ultrasound.net/history1.html
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Biological effects and safety aspects
Francis A Duck
ABSTRACT
Full exploitation of diagnostic ultrasound requires careful consideration of
potential risks. Ultrasound causes small increases in tissue temperature. Whilst
commonly of only fractions of a degree, some conditions can give temperature
increases which could approach 10°C, particularly at exposed bone during
Doppler modes. Safety thresholds are derived from studies into thermal
teratology. Tissues can also be damaged mechanically by gas body activation,
although this mechanism appears to be of very minor concern for most obstetric
applications. Another bioeffects mechanism is radiation pressure, whose
presence is demonstrated by acoustic streaming. Epidemiological studies have
yet to demonstrate unequivocally any causal relationship between exposure to
ultrasound in utero and developmental changes, although all published studies
relate to earlier, low-intensity exposure regimens. There is yet insufficient
understanding of the interaction between ultrasound and the developing embryo
and fetus at all stages in pregnancy, and this lack of detailed knowledge still
advises care and prudence in the use of ultrasound in obstetrics. On-screen safety
indices may assist clinical users to make improved safety judgements.
KEYWORDS
Epidemiology, exposure, gas body activation, mechanical index, non-thermal
effects, regulations, thermal effects, thermal index, ultrasound safety.
INTRODUCTION
Diagnostic ultrasound has an enviable reputation for safety, and the lack of evidence of significant hazard and consequent risk has been one of the key factors
which has established it as the pre-eminent imaging method in obstetrics. Whilst
the severe biological effects associated with x-radiation became abundantly clear
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very early, ultrasound gives no obvious evidence of tissue damage until very high
intensities are used. However, it is now appreciated that even diagnostic levels of
ultrasound can cause small but potentially significant tissue responses. Therefore,
both the design and clinical use of equipment which uses ultrasound for diagnosis
must be subject to the general rule that the diagnostic benefit must be sufficient
to outweigh the potential for harm – a risk/benefit judgement has to be made.
Until about a decade ago, manufacturers designed ultrasound scanners for
particular applications – for example for cardiology, ophthalmology, obstetric or
vascular scanning. Regulation in the USA restricted output intensity from obstetric scanners to be about eight times lower than the highest available. These limits served also to constrain output from equipment available in other countries.
In the early 1990s regulations in the USA were relaxed, in part to allow Doppler
modes to be used in obstetrics, allowing the highest output to be used for all
applications. Manufacturers now sell equipment for obstetric use that can operate
at levels previously reserved only for peripheral vascular applications. They are
required also to display values of safety indices, which reflect the changing output of the machine as it is used clinically for different applications and patients.
Ultrasound in obstetrics and gynaecology
These values of the mechanical index (MI) and thermal index (TI) are intended
to allow users to make a risk/benefit judgement. In order to do this, clinicians
and other users of the equipment must know of the potential hazards inherent
in using ultrasound, and be advised about the interpretation of the safety indices.
This chapter is intended to introduce the reader to these issues. More detailed
information may be found in other publications
2,10,11
and in a series of safety
tutorial articles which are available on the web page of the European Federation
of Societies for Ultrasound in Medicine and Biology (www.efsumb.org/ecmus.
htm) and the International Society for Ultrasound in Obstetrics and Gynecology
(www.isuog.org).
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ACOUSTIC OUTPUT OF DIAGNOSTIC ULTRASOUND SCANNERS
Exposure to ultrasound at sufficiently high levels is capable of causing lethal
damage to tissues. Knowledge of acoustic output serves to ensure that diagnostic
exposures are limited to levels that may be used safely. Broadly, two aspects of the
ultrasound beam are measured, which guide answers to two questions: how much
energy is in the beam and how big are the pulses of ultrasound?
The energy may be described in terms of total acoustic power (energy per
second) or acoustic intensity (power through a specific area). Both of these are
related to the temperature rise in tissue. Commonly the spatial-peak temporalaverage intensity is quoted rather than the acoustic power (I
square centimetre). The size of the pulse is usually measured by its peak rarefactional pressure, pr, in megapascals (MPa). One megapascal is approximately equal
to 10 atmospheres. This quantity is related to the potential for gas body activation or acoustic cavitation. Normally, tables giving I
and pr present the highest
spta
values reached anywhere, and these are found typically near to the focus of the
ultrasound beam.
, in milliwatts per
spta