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_5760_Библиотеки_им_академика_М_И_Перельмана.pdf
32
Chapter 1 · Fundamental Principles
gain will lead to color overow and may obscure atheroscle-
1
rotic lesions in arteries or thrombus in a partially occluded
vein (. Fig. 1.40b). is phenomenon contributes to the
unreliability of planimetric stenosis grading, which is based
on the cross- sectional vessel area depicted in color duplex
scans.
1.1.4.6 Doppler Angle
e evaluation of an artery or vein in the B-mode image combined with the spectral Doppler information obtained with
a dened ultrasound beam and known angle of insonation
enables reliable calculation of blood ow velocity from the
Doppler shi. To ensure reliable Doppler angle correction
for accurate measurement, a long stretch of the vessel of
interest should be displayed in
the longitudinal plane
(parallel vessel walls along the width of the monitor). e
smaller the angle between the Doppler beam and the direction of the owing blood, the higher the Doppler frequency
shi of the reected echo and the higher the sensitivity for
ow detection. At angles of 60–90°, the Doppler frequency
shi decreases and ow velocity measurement becomes progressively more unreliable, and no ow signals are depicted
when the Doppler beam is perpendicular to the vessel wall
(
. Fig.1.40c–e). Conversely, perpendicular insonation of a
reecting surface provides optimal B-mode information
With a linear-array transducer, the examiner can use beam
steering (lateral beam deection) for electronically changing
the direction of the beam to achieve a good Doppler angle
relative to the direction of blood ow. Technically, the beam
can only be steered either right or le by a maximum of
20°, resulting in a Doppler angle of ≥70°. erefore, beam
steering can improve the angle, but does not achieve the best
Doppler angle of <60° (for minimizing measurement errors)
when interrogating a vessel segment running parallel to the
skin surface (. Fig.1.40e).
In the clinical setting, the examiner must nd a compro-
mise between optimal B-mode imaging and optimal ow
evaluation
. e usual procedure is to rst try and achieve
a perpendicular beam angle for morphologic evaluation of
the vessel wall and to then optimize the angle for spectral
Doppler interrogation using either electronic beam steering
or manual manipulation of the transducer (curved-array or
sector probe). Another measure is to place the Doppler gate
at the edge of the scan eld.
In the color ow image of a vessel coursing parallel to the
skin surface (e.g., carotid or femoral artery) examined with
a curved-array transducer held perpendicular to the body
surface, the colors representing the blood ow information
may change in brightness (luminosity) without this reecting dierent ow velocities. Under these conditions, ow
near the margin of the image is displayed in brighter colors
due to higher Doppler shis and in increasingly darker colors toward the center, where the frequency shi decreases
as the insonation angle increases (and no color as the angle
approaches 90°). is phenomenon is relevant only in the
velocity mode and not in the power mode, as the latter is virtually independent of the Doppler angle.
1.1.4.7 Physical Limitations ofColor
Duplex Ultrasound
As a result of the vast amount of information to be processed,
color duplex scanning has a much poorer spatial and temporal resolution than pure B-mode imaging. Axial resolution
is proportional to the wavelength in B-mode imaging, while
it is dependent on the number of sample volumes placed
along the color Doppler scan line in the color duplex mode.
e use of smaller sample volumes improves axial resolution
but at the expense of sensitivity and accuracy in Doppler
shi evaluation as the signal-to-noise ratio deteriorates.
Lateral resolution in color duplex imaging is determined by
the number of color Doppler lines processed per centimeter. e frame rate decreases as the number of Doppler lines
increases, compromising temporal resolution, in particular
at greater scan depths.
As a result of these limitations, the axial resolution of
color duplex ultrasound is on the order of 0.4–1.0mm with
a lateral resolution of only 1.0–2.0mm, which is four to ten
times lower than the B-mode scan resolution (Widder 1999).
In duplex ultrasound, it takes roughly 50–200ms to create one color image, depending on the depth of the target
vessel and size of the color box. is corresponds to a frame
rate
or sampling rate of 5Hz. When a low PRF is selected,
.
the speed at which the color Doppler lines sweep the sector
is similar to or slightly below the mean ow velocity in arteries. erefore, a single color ow frame may simultaneously
depict systolic ow and early diastolic ow (e.g., displayed
in red and blue, respectively, see . Fig.1.37). Due to the low
temporal resolution, however, the color coding does not fully
reect the pulsatile character of ow.
Slow ow produces smaller Doppler frequency shis,
which have to be extracted from short echo pulse packets for
each scan line consisting of a number of individual pulses.
e scan lines must be processed successively.
ough the insonation angle should ideally be as small as
possible for optimal velocity measurement, this is not always
practical because there will be a longer delay when the color
box is tilted
compromise, in particular when examining vessels deeper in
the body. Tilting the color box by 20° and 30° prolongs the
round trip time by 13% and 31%, respectively.
Color duplex imaging, like all diagnostic ultrasound
techniques, is impaired by
ing
caused by bowel gas or calcied structures (bone or calcied plaques in the lumen). e examiner can circumvent
such interfering structures by moving the transducer, but this
oen increases the distance between the transducer and the
target anatomy, and hence the round trip time.
A strong reector in the beam path can act like a mirror and generate a phantom image in another area of the
scan. Such mirror images can be identied by angling the
transducer, which will make the mirror artifacts disappear
or appear in a dierent location. When the ultrasound beam
strikes interfaces of high acoustic impedance at a right angle,
reverberations (repeat echoes) may occur with the ultrasound pulses being reected to and fro, resulting in a kind
(beam steering). is is why one must nd a
scattering and acoustic shadow-
1.1 · Technical Principles ofDiagnostic Ultrasound
33
1
of ping-pong eect. Slight angulation of the transducer prevents reverberations but will also reduce reection from the
interface and thus degrade image quality.
1.1.5 Ultrasound Contrast Agents
Color duplex imaging with a high-resolution transducer usually allows adequate visualization of the peripheral arteries in
the gray-scale mode; the evaluation of blood ow, however,
may be impaired, either by the presence of sclerotic vascular
lesions or by scattering due to edema or other localized so
tissue changes. In these situations, a microbubble contrast
agent, or echo enhancer, can be used to improve blood ow
imaging. For vascular ultrasound examinations in the clinical
setting, however, a contrast agent is rarely needed as there are
only a few situations (e.g., identication of a suitable recipient vessel for crural bypass graing) in which color duplex
and spectral Doppler imaging are degraded by poor imaging
conditions and do not provide the information required for
treatment planning.
Ultrasound contrast agents are gas-lled microbubbles
and enhance the contrast between blood and surrounding
tissue by
partment
administration. Microbubbles can thus enable or improve
the identication and evaluation of vessels with slow or low
blood ow, which may be dicult to identify with conventional ultrasound techniques. While this is an advantage for
vascular applications, ultrasound contrast agents are mainly
used to evaluate organ and lesion perfusion, to characterize
focal liver lesions for example. Contrast-enhanced ultrasound (CEUS) plays virtually no role in the routine clinical
examination of patients with angiologic or vascular surgical
conditions. Notable exceptions are CEUS examinations of
small peripheral vessels and below-the-knee arteries with
slow, postocclusive ow, transcranial duplex scanning, the
search for endoleaks aer stenting of the aorta, demonstration of intraplaque neovascularization in the assessment of
plaque vulnerability, and evaluation of inammatory activity
in patients with vasculitis. With the use of contrast agents,
ultrasound to some extent gives up the crucial advantages it
normally has over other imaging modalities– low cost, short
examination time, and noninvasiveness. Moreover, in those
body regions where diagnostic improvement is achieved
through the administration of an echo enhancer, competing
imaging modalities are used because they provide high accuracy and oen enable better documentation of the ndings.
and properties are commercially available from various manufacturers. Basically, the contrast agents consist of microbubbles composed of a gas core encapsulated by a thin shell
or stabilized by a carrier medium.
the
producing strong reections in the vascular com-
during their lifetime of 3–5min aer intravenous
Ultrasound contrast agents with dierent compositions
ese two factors– the shell and the gas core– determine
stability of ultrasound microbubbles:
5 e microbubble gas core is stabilized by a shell
that contains surface stabilizers (palmitic acid,
phospholipids) or substances that form a capsule on the
molecular level (albumins, polymers).
5 e stability of the core can be improved by using a
heavier gas instead of air (sulfur hexauoride, peruoropropane). Such gases have lower diusivity, higher
physical density, and a lower saturation constant, reducing the solubility of the microbubbles.
e more stable microbubbles used in newer ultrasound
contrast agent preparations have certain advantages, such as
a longer blood half-life due to lower spontaneous solubility.
Bubbles containing a high-molecular-weight gas can pass
the pulmonary circulation because they are more resistant to
destruction when exposed to changing pressures.
1.1.5.1 Approved Ultrasound Contrast
Agents andUses
While a variety of microbubble contrast agents diering in
core and shell composition were investigated around the turn
of the century, only a few microbubble preparations were
ultimately approved for clinical use.
Levovist was the rst ultrasound contrast agent approved
in Europe but has since been taken o the European market.
e preparation consists of a suspension of galactose-based
air bubbles coated with a stabilizing palmitic acid layer.
SonoVue is approved in Europe as an echo enhancer for
vascular applications. e microbubbles have a mean diameter of 5–10μm, contain sulfur hexauoride, and are stabilized by phospholipids. Being eliminated from the body by
exhalation, SonoVue is not nephrotoxic.
e microbubbles are injected as a single bolus of
1–2.4mL (at a rate of 1mL/s) or as a smaller bolus of 0.5–
1.0mL followed by continuous infusion of 1mL/min over
a few minutes. Aer bolus injection, enhancement of the
arterial lumen begins aer 10–30s and peaks aer 30–60s,
followed by a gradual decrease in intensity over 3–8 min
(imaging window). e imaging window can be extended by
infusion of the microbubble preparation. Depending on the
infusion rate, a 15–20dB increase in intensity is observed
aer 1min. Excessive enhancement with appearance of ow
signals outside the vascular space (color blooming) immediately aer bolus administration can be counteracted by
adjusting transmit gain. e slow linear decrease in intensity
following the initial peak aer bolus injection ensures adequate enhancement for several minutes, which is long enough
for most vascular applications. Dynamic contrast-enhanced
ultrasound (CEUS) with generation of time-intensive curves
(TIC) allows estimation of blood volume and regional blood
ow in the target vasculature or a vascular segment of interest (Dietrich etal. 2012).
e intravascular half-life of the microbubbles primarily
depends on their inherent stability and the acoustic energy
applied, which in turn is determined by the ultrasound system’s output and attenuation of the ultrasound beam while
passing through the tissue.
Clinical CEUS examinations can be performed using
conventional ultrasound techniques; however, the acoustic
1
Transmitted
Received
Second harmonic
0
1
2
34
Chapter 1 · Fundamental Principles
energy will rapidly destroy the microbubbles. e life span
of the bubbles is longer when contrast-specic ultrasound
modes such as low-mechanical (MI) index imaging are
employed. is means that the examination is performed
with a lower output power or decreased MI.With this technique, dynamic real-time imaging can be performed over
several minutes aer administration of the microbubbles.
For most indications in vascular ultrasound, it is usually sufcient to inject an echo enhancer bolus of 1.2–2.4mL, followed by a 10-mL saline ush (0.9%).
Ultrasound microbubbles do not diuse from the blood
into surrounding tissues. ey do not leave the vascular system unless blood escapes through a hole in the vessel wall
(e.g., an endoleak). Hence, they have no nephrotoxic eects
and a low overall rate of adverse events. Life-threatening anaphylactic reactions have been reported to occur in less than
0.002% of cases.
ere are several contraindications to the use of SonoVue
as an echo enhancer:
5 Severe pulmonary hypertension, uncontrolled arterial
hypertension, acute lung failure
5 Acute coronary syndrome, severe cardiac insuciency,
malignant arrhythmia
5 Acute respiratory distress syndrome, e.g., bronchial
asthma
5 Pregnancy and breastfeeding (safety remains to be
proven during pregnancy and lactation)
5 Patient age below 18
5 Known intolerance of sulfur hexauoride
e preparation of a patient for an ultrasound examination
with use of an echo enhancer includes obtaining written
informed consent and placing a venous line.
1.1.5.2 Mechanisms ofAction
Ultrasound contrast agents act by increasing the proportion of scattered and reected ultrasound pulses from the
blood, thereby improving both the Doppler signal and the
signal-to- noise ratio (SNR). How strongly the microbubbles
enhance reection depends on their diameter (factor of 6),
the transmit frequency used (factor of 4), and their compressibility. Most microbubbles used as echoenhancers contain gas and enhance backscatter because of the
mismatch between their gas cores and the liquid component of blood
. e intensity of backscatter is determined by
acoustic
the microbubble concentration in the blood and the reection capacity of the individual bubbles, which is a function
of their scatter cross-section. However, the maximum bubble
size is limited by the fact that they must pass the lungs (bubble
size <8μm). It has been shown that backscatter enhances the
echo signal intensity of blood by 15–25dB (Kaps and Seidel
1999). Low ultrasound beam power induces linear oscillation of the microbubbles, transforming the bubbles into
small “ultrasound transmitters.” is is another mechanism
by which microbubbles enhance the ultrasound signal. e
resonance frequency is inversely proportional to the bubble
diameter.
ultrasound pulse
Power
f
. Fig. 1.41 The Doppler-shifted signal returning to the transducer
contains both the fundamental and harmonics, which can be extracted
by comparison with the emitted pulse
nonliner
ultrasound signal
f
f
f
3
At higher power, the microbubbles oscillate in a nonlinear
fashion, producing strong signals at fundamental and nonfundamental frequencies. e nonfundamental or
frequencies
are multiples of the transmitted frequency. e
harmonic
second harmonic has the highest energy and is therefore most
relevant for diagnostic ultrasound (. Fig.1.41). Because the
resonance frequencies of the 2–7-μm gas bubbles are within
the range of the transmitted frequencies of 2.5–10 MHz
typically used in diagnostic ultrasound, the bubble vibrations
produce an additional signal amplication. e corresponding frequencies are received and processed along with the
Doppler-shied frequencies. is resonance behavior of the
microbubbles
improves the SNR by a further 30–35dB, provided that the bandwidth of the ultrasound system extends
over a sucient range of frequencies to enable the generated
harmonics to be detected (Correas etal. 1997).
Short pulses of high energy can be applied to make the
microbubbles burst, producing ultrasound signals that are
detected with high sensitivity. In contrast to the enhancing
mechanisms outlined in the preceding sections, bursting is
independent of blood ow and the resulting signals merely
show the distribution of the collapsed microbubbles at the
time of imaging.
Both SonoVue and Levovist are taken up and eliminated
by the reticuloendothelial system of the liver. In the liver, the
microbubbles can be made to burst by exposing them to a
high-energy beam. In this way, they can contribute to the
sonoscintigraphic identication of liver metastases, which
do not have a reticuloendothelial system. Alternatively, focal
liver lesions may be dierentiated on the basis of their blood
supply (predominantly portal venous versus arterial) and
dierences in contrast agent arrival times aer bolus injection. Because of their selective uptake, the microbubbles also
have the potential to be used as vehicles for the targeted local
1.1 · Technical Principles ofDiagnostic Ultrasound
delivery of chemotherapy. Other echo enhancer preparations
consist of suspensions (some of which contain human albumin) and bubbles stabilized for specic needs.
1.1.5.3 Ultrasound Techniques Using
Contrast Agents
1.1.5.3.1 Contrast-Enhanced Duplex Ultrasound
e reection of ultrasound by microbubbles present in
the blood selectively enhances the vascular system, thus
improving the delineation of arteries and veins from surrounding tissue (in color duplex and power Doppler).
Blood ow velocity is not aected by the microbubbles, and
therefore spectral Doppler analysis can be performed for
quantication of blood ow velocity (CW/PW Doppler) in
the same way as without contrast medium but with lower
gain. When a frequency-based Dopppler technique is used,
however, the eect of contrast enhancement can only be
exploited in large and medium-sized vessels as the signals
returning from slow microcirculatory ow in small vessels
cannot be adequately separated from the echoes produced
by moving tissue.
Microbubble contrast agents have been used to evaluate
slow ow in peripheral vessels, including postocclusive ow,
and carotid artery stenosis, and studies have shown that echo
enhancers improve the accuracy of color duplex ultrasound
in identifying pseudo-occlusion (Fürst et al. 1999; Ferrer
et al. 2000). In the vertebral territory, ultrasound contrast
agents can improve evaluation in patients with poor scanning conditions or with a hypoplastic vertebral artery.
Echo enhancers have also been advocated to improve the
SNR in the examination of deep vessels such as the pelvic
or renal arteries. Using state-of-the-art ultrasound equipment, however, these vascular areas rarely pose diagnostic
problems. A study of duplex ultrasound with echo enhancer
administration in renal artery stenosis found an increase in
diagnostic yield from 64% to 84%, while the improvement
in the sensitivity and specicity for identifying high-grade
stenosis was negligible (Claudon etal. 2000).
1.1.5.3.2 Contrast Harmonic Imaging
is ultrasound technique can improve vascular imaging by
selectively displaying the harmonics specic to microbubbles.
A broadband transducer is used to detect the harmonics generated by the contrast microbubbles, especially the second
harmonic. e backscatter from microbubbles allows better
separation of the echoes from stationary tissues, thus oering
advantages in the sonographic evaluation of slow ow.
1.1.5.3.3 Stimulated Acoustic Emission Imaging
Contrast agent microbubbles are destroyed when exposed
to high ultrasound energy (high mechanical index)
(. Fig.1.42). e bursting bubbles emit transient ultrasound
signals, which can be detected with a broadband transducer.
e system registers the variation in signals from pulse to
pulse. is information is displayed along with the spatial
information of the B-mode image to show the contrast agent
35
. Fig. 1.42 Ultrasound techniques using contrast agents. The
amount of ultrasound energy applied determines how the microbubbles interact with the ultrasound beam, giving rise to dierent
techniques of image generation. Exposure of microbubbles to low
energy (I) causes an increase in reection and backscatter. Intermediate ultrasound energy levels (II) induce nonlinear oscillation of the
microbubbles, resulting in the emission of second harmonics, which
selectively enhance the blood signal (contrast harmonic imaging).
High ultrasound energy (III) destroys the microbubbles, releasing frequencies that provide information on the distribution of the contrast
bubbles in the circulation (stimulated acoustic emission)
distribution in the macro- and microcirculation at a given
point in time.
1.1.5.4 Summary ofTechnical Aspects
andClinical Indications
In color duplex ultrasound, the administration of echo
enhancers can lead to excessive enhancement of the color
Doppler signal with color overow obscuring perivascular
structures and parts of the vessel wall, in particular when
larger vessels are examined. is eect can be remedied by
lowering the receive gain. On the other hand, the strong
reection produced by the contrast bubbles can attenuate structures farther away from the transducer than the
enhanced vessel.
A Doppler waveform obtained aer echo enhancer
administration will show spectral broadening with almost
complete lling-in of the systolic window. Again, the eect
can be counteracted by lowering the receive gain.
Basically, all sonographic techniques exploiting the different eects of ultrasound contrast agents in the vascular
compartment aim at improving sensitivity to blood ow
phenomena. However, with the sophisticated ultrasound
technology available today, most vascular applications do not
require use of a contrast agent. Moreover, in those rare cases
where an adequate diagnostic evaluation with color duplex
imaging is not possible, the use of a microbubble contrast
agent is oen limited as well and oers no benets. Despite
these cautionary remarks, there are a few situations in which
contrast microbubbles improve diagnostic yield. e most
important indications for vascular CEUS are:
1
36
Chapter 1 · Fundamental Principles
5 Transcranial duplex imaging
1
5 Evaluation for renovascular disease
5 Vein mapping for identication of a suitable recipient
segment before crural bypass gra surgery
5 Search for endoleaks in patients with an aortic stent.
the vibrating bubbles. is phenomenon is known as
microstreaming and can generate very high pressures with
disruption of cell membranes. In the other form of cavita-
inertial cavitation, existing bubbles or cavitation nuclei
tion,
expand during the low-pressure phase and then collapse
violently. Microbubble collapse is a highly localized process
It is expected, however, that ultrasound contrast agents will
be used increasingly in the evaluation of microcirculation,
e.g., for identifying plaque neovascularization (Seidel etal.
2006; Claudon etal. 2008).
occurring on the order of microseconds. Collapsing bubbles
can produce extremely high temperatures and pressures but
these dissipate rapidly. Bursting bubbles therefore have the
potential to destroy cells and tissues. ere is scientic evi-
dence to suggest that inertial cavitation is a threshold phe-
nomenon and will only occur if microbubbles already present
1.1.6 Safety ofDiagnostic Ultrasound
in the acoustic eld are exposed to excessive acoustic pressures
and frequencies. Pressure below the cavitation threshold will
Ever since the early 1960s, when this technique was rst used
for diagnostic imaging, the potential biological hazards of
medical ultrasound have been discussed. Ultrasound traveling through the human body can have two eects known
to cause changes in biological systems. Firstly there are the
thermal eects resulting from the conversion of ultrasound
energy into heat and secondly there are the mechanical
eects arising from pressure changes associated with the
propagation of sound waves in a medium.
never by itself lead to cavitation, not even during extremely
long exposure to ultrasound. Inertial cavitation induced by
diagnostic ultrasound procedures therefore remains a mere
theoretical possibility and has never been reported invivo.
However, one must also be aware that invivo evidence
of potential bioeects of ultrasound is very dicult, if not
impossible, to obtain: cavitation can occur anywhere in the
body, and the damage it produces may be very local, involving only a few cells. Modern ultrasound equipment incorporates safety mechanisms allowing the user to limit the
1.1.6.1 Thermal Eects
Exposure to diagnostic ultrasound can increase tissue temperature because sound energy is absorbed and converted
average acoustic output, thus avoiding peak pressures that
could theoretically lead to cavitation or other mechanical
bioeects.
to heat. e ability to absorb energy varies with the tissue;
it is low in body uids (amniotic uid, blood, urine) and
high in bones. Adult bones absorb 60–80% of the incident
1.1.6.3 Specic Risks ofIndividual Ultrasound
Techniques
ultrasound energy. In addition, absorption in the body is also
aected by technical parameters, most notably the output
frequency of the transducer. Higher frequencies are absorbed
more rapidly. A temperature increase of 2.5°C can severely
damage biological tissues, while an increase of 1°C is generally considered harmless. Experimental evidence suggests
that the thermal eects of diagnostic ultrasound procedures
pose no health hazard.
1.1.6.3.1 B-Mode
B-mode imaging is generally performed at very low acoustic
output, resulting in intensities below 10mW/cm2. e individual pulses are very short (<1ms) and are emitted at a PRF
of less than 5kHz to achieve high resolution. As the energy
transmitted into the body is dissipated over a large volume,
the resulting rise in tissue temperature is so small that it
remains below the limit of detection. Diagnostic B-mode
1.1.6.2 Mechanical Eects
Most of the mechanical eects of ultrasound that are poten-
ultrasound is considered absolutely safe in terms of potential
hazards to patients.
tially harmful to living tissues are related to the formation,
growth, and possible collapse of tiny gas bubbles in the
ultrasound eld, a process known as cavitation. Recall that
ultrasound propagates through tissue in waves of alternating
high and low pressure. Bubble formation or the expansion of
existing bubbles occurs when the negative pressure is large
enough. e occurrence of cavitation and its eects depend
on the frequency and intensity of the transmitted ultrasound
1.1.6.3.2 M-Mode
is technique uses higher energies and may theoretically
cause tissue heating. In this mode, a stationary beam is
emitted repeatedly to evaluate moving structures. e scan
volume is smaller than in B-mode imaging, but the PRF is
much lower (only approx. 1kHz). M-mode sonography is
also considered safe.
waves as well as on the focus of the acoustic eld. Two types
of cavitation are commonly described: stable and inertial (or
transient).
Stable cavitation refers to the continuous oscillation of
gas-lled bodies in response to the alternating positive and
negative pressures to which they are exposed in an ultrasound eld. Such cyclic expansions and contractions result
in an increased ow in the uid-like medium surrounding
1.1.6.3.3 CW Doppler
As with M-mode techniques, the scan volume is small and
there is continuous exposure. e power output can reach
up to 100mW, and some procedures have the potential to
produce biologically signicant temperature rises. Possible
mechanical eects are much less of a concern than in B-mode
or M-mode scanning, despite the higher acoustic output. e
PP
()
()
12
r
××
×
h
p
1.2 · Hemodynamic Principles
37
1
transmit power depends on the depth of the target anatomy.
It is the operator’s responsibility to keep the overall examination time as short as is consistent with achieving diagnostically useful results.
1.1.6.3.4 PW Doppler
Again, the exposed volume is relatively small and the PRF
is high. e individual pulses are oen twice as long as
with B-mode or M-mode techniques. Taken together, the
machine settings used in PW Doppler applications can result
in considerable exposure, and the risk of tissue heating is far
greater. Conversely, mechanical eects are negligible because
the intensity of the emitted pulses is the same as in B-mode
and M-mode imaging.
1.1.6.3.5 Color Doppler
e acoustic output in color Doppler imaging is intermediate
between that of B-mode and PW Doppler. Mechanical eects
are negligible. e emitted ultrasound pulses are distributed
over a relatively large tissue volume. Temperature rises are
higher than with B-mode imaging but lower than with PW
Doppler techniques.
1.1.6.4 Conclusion
Current uses of diagnostic ultrasound expose the body to
intensities that do not exceed 100mW/cm2. ere is no evidence that these intensities damage living tissues.
Nevertheless, the examiner should always seek to minimize exposure by limiting both the power output and the
duration of scanning to what is absolutely necessary to obtain
the desired diagnostic information. is approach is known
as the ALARA principle (as low as reasonably achievable)
and applies to all diagnostic imaging modalities. Doppler
ultrasound, which uses higher intensities, should not be
employed during the rst three months of pregnancy.
1.2 Hemodynamic Principles
1.2.1 Laminar Flow
Although blood ow is subject to specic conditions due to
the solid components in plasma and the elasticity of the vessel wall, it basically follows the laws of ow dynamics. ese
laws govern the ow of a uid in tubes and apply to watery or
oily solutions of a constant viscosity (Newtonian uid) and
assume that ow velocity under these conditions is primarily a function of the pressure dierence that exists between
the two ends of the tube. ese ideal conditions for continuous laminar ow are typically not met in a living organism
because various factors such as elasticity of the vessel wall,
pulsatility resulting from cardiac activity, curving of vessels,
and branching aect blood ow, resulting in changing velocity distributions in the moving layers of the blood.
Moreover, blood is not a watery or oily solution of constant viscosity but a suspension of solid blood cells in plasma.
Blood viscosity is primarily dependent on the hematocrit level
and is only constant when hematocrit is below 10, increasing
exponentially at higher levels. Other factors aecting blood
viscosity are plasma viscosity and vessel diameter. In the
terminal capillary bed, viscosity is additionally inuenced
by the deformation of red blood cells. Despite these specic
features of blood ow, some basic hemodynamic terms and
laws are useful and will make it easier to understand normal and abnormal ow in arteries and veins. In addition,
invitro experiments and invivo blood ow measurements
using duplex scanning have provided new insights into the
ow behavior in specic vessels under normal and abnormal
conditions as well as under the inuence of pharmacologic
agents.
Laminar ow is characterized by a constant velocity over
time. Flow in a tube is brought about by a pressure dierence between the two ends of the tube. e pressure dierence (P
– P2) is proportional to the volume ow rate. e
1
volume ow rate (I) is proportional to the tube diameter (r)
and inversely proportional to its length (l) and the viscosity
of the uid (η). Mathematically, this relationship is expressed
in the
Hagen–Poiseuille law:
4
-
12
××
8ph
l
××
-
=
R
PP r
I
=
By analogy with Ohm’s law, ow resistance can be calculated
from the Hagen–Poiseuille equation:
8
l
R
=
4
It follows that resistance is proportional to the length of the
tube (l) and the viscosity of the liquid (η). Overall resistance
is most strongly aected by the radius (r) of the tube, which
appears in the equation raised to the fourth power. is
means that decreasing the vessel radius by one half, for example, increases ow resistance by a factor of 16. Peripheral
resistance in the vascular system is regulated according to
demand, primarily by the tone of the arterioles, and aects
the pulsatility of blood ow in the large arteries supplying
these territories. erefore, it is also reected in spectral
Doppler tracings from these arteries.
e ow prole of continuous ow is determined by
inertial and frictional forces. Friction produces a laminar
or, in the 3D model, parabolic ow prole. Flow is fastest in
the center of a vessel and decreases toward the wall, where it
approximates zero.
In color duplex images, this decrease in blood ow
velocity from midstream to the vessel wall is indicated by
brighter colors in the center and darker colors near the wall
(. Fig.1.43a). e following factors determine the shape of
the ow prole of blood:
5 Velocity
5 Viscosity (internal friction)
5 Adhesion of the blood to the vessel wall (external friction)
5 Cohesion (forces that occur between adjacent molecules
of like composition).
38
Chapter 1 · Fundamental Principles
1
. Fig. 1.43 a Typical triphasic Doppler waveform of the popliteal artery. In the color ow image, laminar ow is characterized by brighter color-
ing in the center with darker colors representing slow ow near the wall. Red indicates blood ow toward the transducer. A vein closer to the
transducer is displayed in blue, indicating blood ow away from the transducer. The triphasic waveform consists of a steep upslope (A) to peak
systolic velocity (PSV) (B), a deceleration phase (C), a short phase of early diastolic backward ow (D), and forward ow from the middle to the end
of diastole (E). The magnitude and duration of diastolic forward ow (E) depend on peripheral resistance (sympathetic tone) and the thrust generated by the compliant aorta (windkessel eect). Blood ow toward the transducer is displayed above the baseline, and ow away from it below
the line. The Doppler angle of insonation is 59°, and PSV is 85cm/s. The dierent intensities of the individual pixels in the Doppler waveform
reect the number of red blood cells moving at a given velocity. The amplitude can also be represented in the form of a histogram. b Use of a low
pulse repetition frequency (PRF) to ensure good color lling of arteries with slow ow below the knee. In the rst color image (left), the anterior
tibial artery (blue, ow toward the periphery, away from the transducer) is depicted with central aliasing (color change from blue to yellow to red).
This example illustrates a laminar ow prole with fast ow in the center of the artery and lower ow velocities near the wall due to friction. The
second color image depicts early diastolic blood ow at the same site, which is due to the fast image generation. In the center, arterial reux due
to high peripheral resistance is seen as a superimposed wave (red, toward transducer) while ow toward the periphery predominates nearer the
walls (blue, away from transducer). The view illustrates true ow reversal relative to the ultrasound beam rather than aliasing. True ow reversal is
characterized by a color change from blue to black to red. The third color ow image depicts blood ow (blue, away from transducer) toward the
periphery without aliasing in mid-diastole. While each of the three color ow images depicts peripheral arterial blood ow at a specic time during the cardiac cycle, the corresponding Doppler waveform (right) displays the ow changes over time
Blood diers from Newtonian uid in that its viscosity is
not an inherent property that only varies with temperature but is mainly determined by the hematocrit level and
other factors such as plasma viscosity (which in turn is predominantly dependent on the brinogen concentration),
red blood cell deformability, and the degree of shearing.
In an artery or vein with laminar ow, shear stress, like
thrust, is weakest in the center and strongest near the wall
(. Fig.1.43b).
According to the continuity law, a decrease in the cross-
sectional area
in the course of a vessel segment leads to an
increase in mean ow velocity. Blood ow through a vessel
segment with an abrupt change in caliber becomes attened
(plug ow) upon entering the narrower vessel segment. In
plug ow, inertial forces are stronger than frictional forces,
resulting in the same ow velocity of all uid layers in the
vessel except for a thin layer near the wall. So-called turbulent ow results when the inertial forces become even stronger than the frictional forces, which bring order to the course
of ow. Turbulent ow is characterized by an irregular ow
pattern with ow in dierent directions. e typical parabolic ow pattern develops aer a certain stretch along which
=+
1.2 · Hemodynamic Principles
A
Continuity equation
1
A
2
39
V
2
=
V
1
A = Vascular cross-sectional area
V = Mean flow velocity
Stenosis
Marginal zone of turbulent flow
1
v
1
a
Outer wall
Lateral angle α
b
A
1
Separation
v
2
point
A
2
Inner wall
Flow divider
Separation
zone
v
3
Reattachment
point
A
3
c
PressureFlow velocity
Turbulent flow
Distance
Distance
Artery
Additional
loss due to
turbulent
flow
Additional
loss due to
turbulent
flow
. Fig. 1.44 a When a uid such as blood enters a narrower lumen, parabolic ow changes into plug ow and returns to its original prole only
after having traveled some distance under the inuence of shear stress. According to the continuity equation, ow velocity increases in proportion to the decrease in diameter. b Flow in a vessel branching. Thrust and shearing are highest at the inner wall of the branching. Separation
occurs at the outer wall, where thrust is rather low. Physiologic ow separation occurs in the carotid bulb (see . Fig. 5.49 (Atlas)). c Diagram of
ow in a vessel segment with higher-grade stenosis and corresponding curves (solid blue lines) representing the eects of the stenosis on pressure (top) and ow velocity (bottom) in the stenotic and poststenotic segment. The law of conservation of energy predicts that static energy
(blood pressure) is converted to kinetic energy (ow velocity) (Bernoulli equation). It thus follows that the intrastenotic increase in ow velocity
results in a proportional drop in pressure (neglecting other factors such as blood viscosity and systolic-diastolic ow variation). The actual pressure and ow velocity measured in a poststenotic vessel segment (dotted lines) are lower than theoretically predicted because the equation does
not consider losses resulting from turbulence and friction
frictional forces predominate. Physiologically, this occurs
when blood leaves the le ventricle and enters the ascending aorta. e sharp velocity gradient between ow in the
center and the thin boundary layer near the wall in plug ow
is associated with strong shear stress.
e law of conservation of energy states that the total
amount of energy
in a closed system remains constant.
Applied to blood ow, this means that the total energy in
a stenotic vessel is the same before and aer the stenosis
(unless there is loss of energy from the system) and that there
is an inversely proportional relationship between static and
dynamic components (. Fig.1.44):
It follows from this law that increasing ow velocity within
a stenosis (E
intravascular pressure (E
) is associated with decreasing tangential
kinetic
). e reverse applies to the
static
poststenotic segment: increasing pressure results in turbulent ow with slow ow components near the wall and can
promote intramural hematoma formation.
An abrupt decrease in the cross-sectional area in a stenotic vessel segment and the resulting increase in ow velocity are associated with progressive disturbance of laminar
ow, which will nally become turbulent. Turbulent ow
above a critical velocity is characterized on color duplex
ultrasound by a mosaic of colors reecting the dierent ow
directions. e transition from laminar to turbulent ow
EEE
totalstatic kinetic
can be calculated by means of the dimensionless Reynolds
40
××
Chapter 1 · Fundamental Principles
number, which depends on mean ow velocity (v), vessel
1
diameter (d), density of the uid (p), and viscosity (η):
vd p
Re =
h
to changes in pressure. e pressure amplitude generated by
le ventricular activity is smoothed out by the
of the aorta and other large elastic, or conducting, arteries
compliance
(windkessel eect), resulting in a more steady ow. Another
factor aecting the ow prole is the peripheral resistance.
Flow is highly pulsatile in the extremity arteries because
Data from invitro model experiments show blood ow to
be fairly laminar for Reynolds numbers up to 2000 and to
become increasingly turbulent as the number exceeds 2000.
Turbulent blood ow is characterized by a pattern of ran-
dom ow directions, seen in color ow images as color shis
(indicating retrograde ow components) or a mosaic of colors.
In turbulent ow, part of the kinetic energy is converted
into acoustic energy, producing a characteristic bruit, which
can be detected by auscultation.
In vessels with pulsatile ow, which normally is laminar, turbulent ow may occur at specic phases of the cardiac cycle under physiologic conditions. is phenomenon
depends on the ow prole (high pulsatility) and pulse rate.
A sudden increase in the vessel diameter results in a
longer ow prole and greater velocity gradient across the
peripheral resistance is high at rest, giving rise to the characteristic triphasic Doppler waveform. An increase in the
peripheral blood demand leads to dilatation of the arterioles,
and the resulting decrease in peripheral resistance changes
the Doppler waveform. Peripheral resistance may decrease
under normal (muscle activity) or abnormal conditions
(local inammation, postocclusive ischemia, tumor perfusion). A decrease in pheripheral resistance leads to an
increase in the diastolic ow component. Moreover, the
character of the waveform is aected by central regulatory
processes (increase in heart rate, blood pressure) and vessel
wall elasticity (diabetes mellitus).
As peripheral resistance is a crucial factor inuencing
blood ow and hence the Doppler waveform, a distinction is
made between low-resistance ow and high-resistance ow
(
. Fig.1.45).
vessel lumen. If the dierence between a narrow and wide
(poststenotic) segment exceeds a certain value, ow separa-
tion
and eddy currents will occur near the wall. Flow separation near the wall is also observed in branching blood vessels
(. Fig.1.44).
Flow separation gives rise to recirculation zones, in
which relative stasis of ow, in conjunction with shear stress,
induces platelet aggregation with release and adhesion of
procoagulative agents, which in turn can trigger local atherogenic processes. is is a possible mechanism contributing to
the preferred occurrence of atherosclerotic lesions in dividing and branching vessel segments.
e most notable example of ow separation can be
encountered when imaging the origin of the internal carotid
artery (ICA), where it occurs as a result of both widening
in the bulb area and branching (see . Fig. 5.49 (Atlas)). e
1.2.2.1 Low-Resistance Flow
Arteries supplying parenchymal organs and the brain are
characterized by a fairly steady blood ow as a result of low
peripheral resistance
. In these arteries, a moderate systolic
rise is followed by a steady ow that persists throughout diastole. is ow prole is typical of the renal, hepatic, splenic,
internal carotid, and vertebral arteries.
Continuous diastolic ow in the arteries supplying parenchymal organs is necessary to ensure constant perfusion of
these organs. is is accomplished by the lower peripheral
resistance of these vascular beds and the windkessel function
of the large conducting arteries including the aorta, which
jointly produce a more continuous ow than would result
from the action of the le ventricle and aortic valve alone
(. Fig.1.45).
high wall pressure in this area, in conjunction with slow ow
in separation zones, contributes to the preferred development of carotid bulb plaque on the wall opposite the external carotid artery origin (. Figs.1.44b, 5.49 (Atlas), and 5.56
(Atlas)). Flow separation can also occur downstream of a stenosis, where the vascular cross-sectional area increases again
(see . Fig.1.46a, b), and increased wall pressure can cause
poststenotic dilatation.
Poststenotic vascular dilatation or even aneurysm is rare
in atherosclerotic stenosis; it is more common in patients
with other vascular conditions not causing wall sclerosis
such as compression syndromes (see . Figs. 3.101, 2.105, and
2.106 (all Atlas)) or bromuscular dysplasia.
1.2.2.2 High-Resistance Flow
High peripheral resistance results in a more pulsatile ow
with a steep systolic upslope during the acceleration phase,
followed by deceleration and a signicant reux in early diastole, short backward ow in mid-diastole, and typically zero
ow in late diastole. is pattern is referred to as triphasic
ow
.
e systolic pulse wave is in part reected by the high
peripheral resistance and thus moves backward through
the arterial system until the ow is again redirected toward
the periphery by the inux of blood during the next cardiac
cycle. is ow component is small due to the high peripheral resistance.
As a result of the high pressure in the arterioles supplied
1.2.2 Flow Proles andPerfusion Regulation
by the limb arteries, signicant blood ow in these vessels
occurs only during systole when systemic pressure is higher
Unlike laminar ow, pulsatile ow changes periodically over
time. Phases of acceleration and deceleration vary in relation
than peripheral pressure. e pressure during diastole is too
low to produce blood ow toward the periphery.
1.2 · Hemodynamic Principles
LowerPeripheral resistanceHigher
Arteriolar diameter Arteriolar diameter
Autoregulation
Adaptive regulation
+ p
Pressure pulse generated by the heart
Reflected pressure wave
- p
D pa
41
1
Normotensive
Rest
c
. Fig. 1.45 a Eect of peripheral resistance on the Doppler waveform. Pulsatility increases with peripheral resistance. b Flow pulse curves
resulting from superimposition of the pressure pulse generated by the heart and the pressure waves reected by the distal vascular bed in arteries with high peripheral resistance (left) and low peripheral resistance (right). For an identical pressure pulse generated by the heart (above the
baseline), the pulse curve resulting from interaction with the reected pulse (below the baseline) is highly pulsatile when peripheral resistance is
high and less pulsatile when peripheral resistance is low. The pressure gradient (∆p) calculated by subtracting the reected pulse wave from the
pulse wave emitted by the heart is directly proportional to blood ow velocity (according to the Hagen–Poiseuille law). c Changes in arteriolar
diameter in response to blood pressure changes (autoregulation, left) and in response to exercise (adaptive regulation, right). d Simplied schematic illustration of changes in blood pressure (with ow velocity changing in proportion), peripheral resistance, and total area of the vascular
bed through the circulatory system (arteries, capillaries, veins)
Blood pressure
Exercise
Cross-sectional area
High Low
b
Aorta
Pressure
Resistance
d
Peripheral resistance
Arteries
Systolic
Diastolic
Capillaries
Arterioles
Vena cava
Veins
Venules
High-resistance ow occurs in the arteries supplying the
muscles and the skin, such as the arteries of the arms and
legs, and the external carotid artery. e ratio of skin to muscle supply determines the amount of diastolic forward ow.
When peripheral demand increases (muscle activity, inammation), the arterioles dilate to reduce local vascular resistance, and forward ow, primarily during diastole, increases.
Transitions between these two ow patterns may
occur under normal and abnormal conditions. Besides, there
are vessels with mixed patterns. An example is the superior
mesenteric artery, which has pulsatile ow like a limb artery
but also has a signicant end-diastolic ow component. e
amount of end-diastolic ow is regulated adaptively and
increases with demand aer ingestion of food. Adaptive
adjustment of diastolic ow through arteriolar vasodilation
occurs in all arteries with high-resistance ow. Other factors
inuencing late diastolic ow include systemic factors such
as sympathetic tone and the windkessel function of the aorta.
is is why loss of aortic compliance and of vascular elasticity
results in more pulsatile ow.
An arteriovenous (AV) stula turns high-resistance ow
into low-resistance ow. A change from low-resistance to
high-resistance ow in a transplant renal artery is an important diagnostic criterion in the diagnosis of gra rejection.