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12
Chapter 1 · Fundamental Principles
unchanged because uid reects and attenuates only little of
1
the ultrasound energy. Time gain compensation therefore
amplies echoes returning from behind a low-attenuation
region more than necessary. Acoustic enhancement can be
exploited diagnostically in distinguishing a uid-lled lesion
such as a cyst from a solid mass (. Fig.1.14).
1.1.1.9.4 Side Lobes
A transducer transmits not only the main beam (also called
the main lobe) but also some weaker beams, or side lobes,
on either side of the primary beam in the near eld. When
a side lobe strikes a strong reector, the obliquely deected
echoes are misrepresented in the resulting image because
they are processed as if they had originated from the main
1.1.1.9.3 Edge Eect
e edge eect is a form of acoustic shadowing that is
observed at the margins of curved, uid-lled spaces such
as cysts and is assumed to be caused by a combination
beam (. Fig.1.16). Modern ultrasound systems use various
techniques, such as delay time calculation or suppression of
echoes not returning along a path perpendicular to the trans-
ducer face, to minimize side lobe eects.
of refraction and reection. When a parallel ultrasound
beam passes through the lateral border of such a space,
sound is diverted into the surrounding tissue. As a result,
no ultrasound signal penetrates beyond the diverting
structure, and hence no diagnostic information is obtained
from that area. is phenomenon also explains the incomplete display of the margins of certain structures such as
the fetal head or a blood vessel depicted in cross-section
(. Fig.1.15).
1.1.1.9.5 Reverberation Artifact
is type of artifact is also known as multiple reection
artifact and occurs when ultrasound is reected back to
the transducer from a strongly reective surface in the near
eld. Part of the returning echo is properly processed by the
transducer, while another part is reected back into the body.
Sound can thus bounce back and forth between the reec-
tor and the transducer face (ping-pong eect). e resulting
. Fig. 1.13 Posterior shadowing occurs when a large impedance
mismatch or object with high sound absorption is encountered
( Courtesy of Hitachi Ltd.)
. Fig. 1.15 Edge eects are
caused by a combination of
refraction and reection when a
parallel ultrasound beam passes
through the lateral border of a
curved, uid-lled space. Right
section: Transverse image of an
artery showing the eect of an
ultrasound beam tangentially hitting the arterial wall. The beam is
refracted, giving rise to an acoustic shadow posteriorly, where
no ultrasound energy is present
that can be reected (Courtesy of
Hitachi Ltd.)
. Fig. 1.14 Acoustic enhancement occurs behind low-attenuating
areas (Courtesy of Hitachi Ltd.)
A
X
X
c
2c
1.1 · Technical Principles ofDiagnostic Ultrasound
. Fig. 1.16 Diagram of side lobe artifact (Courtesy of Hitachi Ltd.)
. Fig. 1.17 Diagram of reverberation artifact. This is the repeat
reection of an ultrasound beam hitting a strong reector near the
transducer. In this situation, sound will bounce back and forth between
the reector and the transducer. Echoes from multiple reections
return to the transducer later than the direct echoes and are misrepresented in the image as a copy of the original object at a greater depth
(Courtesy of Hitachi Ltd.)
reverberation artifact is seen in the display as several equidistant echoes decreasing in brightness with depth. is
artifact typically arises when there is a large acoustic impedance mismatch near the transducer (so tissue/air interface)
(. Fig.1.17).
1.1.1.9.6 Geometric Distortion
In processing returning echoes and creating an image, the
ultrasound system relies on certain assumptions, for example,
that ultrasound travels in a straight line or at a constant speed
in the body. In fact, however, an ultrasound beam can be
deected from its straight path, and the speed of sound varies
slightly with the tissue. As a result, the ultrasound image may
not reect the exact anatomic location of a feature.
13
1.1.2 Basic Physics ofDoppler Ultrasound
In 1842, the Austrian physicist and mathematician Christian
Johann Doppler described what is now called the Doppler
eect or Doppler shi. is phenomenon refers to the change
in frequency of a wave resulting from relative movement
between the source of the wave and an observer. A familiar
example is an ambulance siren: although the emitted frequency remains the same, the siren has a higher pitch when
the ambulance is approaching and a lower pitch when the
vehicle is receding. e pitch changes abruptly at the moment
the ambulance passes the observer. us, the pitch of the
siren perceived by the human ear depends on the direction
of motion relative to the observer and remains consistently
high while the vehicle is approaching and consistently low
while it is receding. is is dierent from the intensity of the
sound, or the loudness of the siren, which increases gradually
as the vehicle approaches and then decreases gradually aer
the vehicle has passed the observer. e Doppler eect occurs
when the source or the observer is moving toward or away
from the other or when both are moving relative to each other.
For a vehicle traveling at a speed of 100km/h, the dierence
in pitch due to the Doppler eect is almost two whole tones.
Compared to the emitted frequency, the received frequency is higher when the source and receiver approach each
other and lower during the recession (. Fig.1.18a).
is dierence in frequency, occurring when the source
and/or receiver of a sound wave move relative to each other,
is known as the Doppler eect or Doppler shi.
In diagnostic ultrasound, the Doppler eect is used to
calculate blood ow velocity from the dierence in frequency
between the emitted and reected waves; this was rst reported
by Satomura in 1959. e signals reected by moving red blood
cells have a dierent frequency than the emitted beam. In this
case, the transducer transmitting and receiving the signals is
stationary and the frequency shi is caused by the motion of
the reector (red blood cells). In this situation, the Doppler
shi occurs twice– when the ultrasound beam emitting from
the stationary transducer strikes the red blood cells and when
the blood cells backscatter the signal, now acting as a moving source with the transducer becoming a stationary receiver.
e Doppler shi frequency depends on the frequency of the
transmitted ultrasound waves, the velocity of the moving red
blood cells, and the angle at which the Doppler beam intersects the vessel. is angle is known as the Doppler angle.
e Doppler eect can be used to calculate blood ow
velocity because the Doppler shi frequency depends on the
direction of blood ow and is proportional to the speed of the
moving red blood cells. e shi is detected by the Doppler
probe. e direction of blood ow relative to the transducer
determines whether the returning echoes have a higher or
lower frequency, and the ow velocity determines the magnitude of the frequency shi (. Fig.1.18b). is relationship
is expressed in the Doppler equation:
Fv
××
os
FFF
=-=
dr
0
0
a
1
14
Chapter 1 · Fundamental Principles
1
Doppler Effect
Stationary source Moving source
∆F = Fr – F0 = 2 · F0 · v
+ Df
f
f
0
. Fig. 1.18a, b Doppler eect. a Dependence of the Doppler shift (change in frequency between source and receiver) on the velocity of the
moving source and its direction of motion relative to the reector. b Diagram of Doppler interrogation of a vessel with laminar blood ow. The
arrows in the vessel are vectors representing dierent ow velocities. Blood ow is fastest in the center and decreases toward the wall. The drawing illustrates the eect of the angle of incidence on the Doppler measurement. In the equation for calculating the Doppler shift, this angle is represented by the cosine function. The Doppler shift increases with the acuity of the angle (cosine of 90°=0) (T, transmitter; R, receiver; F0, emitted
frequency; Fr, reected frequency)
f
0
Df »
f
0
– Df
0
v
c
Fd Doppler frequency shi
F0 emitted frequency
Fr reected frequency
ν mean ow velocity of the reecting red blood cells
f
0
ba
. Table 1.5 Dependence of the Doppler shift frequency (Df)
on the angle of insonation
Parameter Values
Vessel
cos α
c
T
R
0
F
r
F
α
c speed of sound in so tissue (about 1540m/s)
α angle between ultrasound beam and direction of blood
ow
In the transcutaneous measurement of blood ow by Doppler ultrasound, angle correction is necessary to calculate the
ow velocity because the Doppler beam cannot be aligned
Angle α 0° 30° 45° 60° 90°
Cos α 1 0.866 0.707 0.5 0
Df (MHz) 7.79 6.75 5.51 3.90 0
Percentage error 0 13 29 50 100
parallel to the direction of ow. e transformation with
representation of the dierent velocity vectors is expressed
mathematically as a cosine function of the angle between the
sound beam and the blood vessel (cos α).
(or Δf) is proportional to the velocity of blood ow,
F
d
cos α, and the carrier frequency of the ultrasound beam.
For angles of about 90°, the cosine function yields values around 0, at which there is no Doppler frequency shi,
and the Doppler shi increases as the angle decreases (with a
maximum cosα of 1 at α=0°).
e blood ow velocity is calculated by solving the
Doppler shi equation for V:
calculation. At angles around 90°, a Doppler shi is no longer
detectable and the ow direction cannot be determined. is
is reected in the color duplex scan by the absence of colorcoded ow signals although ow is present.
. Table1.5 lists the Doppler shi frequencies for dier-
ent angles of incidence, illustrating how the percentage error
in calculating blood ow velocity increases with the Doppler
angle. e values were calculated for a transmitted frequency
of 6MHz and a blood ow velocity of 1ms/1.
It is apparent from the examples listed in . Table1.5 that
no Doppler shi is detectable at a 90° angle of incidence. e
VFF
=-
()
×
r
0
c
×
2cos
a
F
0
reason is that when the ultrasound beam is perpendicular
to the direction of blood ow, there is no relative movement
between the Doppler probe and red blood cells. Velocity
is formula allows calculation of the blood ow velocity
from the measured Doppler frequency shi at a given transmit frequency and angle of incidence. e accuracy of the
calculation increases with the acuity of the angle. Ideally, the
Doppler angle should be kept at or below 60° to minimize
errors in the calculation of ow velocity. At angles above 60°,
even minor errors in determining the Doppler angle (which
are unavoidable in the clinical setting, especially when
curved vessels are interrogated) unduly distort the velocity
measurement is most accurate when the Doppler beam
is aligned parallel to the blood ow. If this is not possible,
accurate velocity estimates can only be made if the Doppler
angle is measured using angle correction. e Doppler angle
is measured by placing the angle correction cursor parallel to
the direction of ow in the B-mode image. For precise calcu-
lation, a correction factor of 1/cosα is used.
. Table1.6 lists
the correction factors for dierent Doppler angles and the
overestimation or underestimation of blood ow velocities
1.1 · Technical Principles ofDiagnostic Ultrasound
. Table 1.6 Relationship between Doppler angle and error in
blood ow velocity calculation
15
1
Angle α Correction factor
1/cos α
30° 1.15 ±3%
45° 1.41 ±6%
60° 2.00 ±9%
70° 2.92 ±14%
75° 3.86 ±21%
80° 5.76 ±30%
Error in calculated
blood ow velocity
resulting from cursor misplacement. e data in . Table1.6
illustrate how the error in calculating blood ow velocities
increases with the Doppler angle. e examiner must therefore try to minimize the insonation angle for Doppler interrogation.
Doppler shi frequencies are extracted by the demodulator of the ultrasound system based on a comparison of
the returning Doppler-shied signal and the transmitted
frequency. e Doppler shi frequencies occurring in medical imaging are in the audible range and can be output to a
loudspeaker. Information about the direction of ow relative
to the transducer can also be extracted from the Doppler signal; this, however, requires more sophisticated demodulation
techniques. Blood ow toward the transducer produces a
positive frequency shi, and blood ow away from the transducer a negative shi.
Blood ow velocity varies across the vessel lumen. Blood
cells move faster in the center and slower near the wall due
to friction, giving rise to a laminar ow prole. Other factors aecting the ow prole include the pulsatility of blood
ow and the elasticity of the vessel wall or changes in ow
resulting from bends in the vessel, branching, and narrowing. e Doppler signal derived from owing blood thus
contains a range of frequencies, which can be extracted
using a mathematical algorithm called fast Fourier transform (FFT). is spectral analysis enables changes in blood
ow velocity to be displayed over time. In the resulting
Doppler spectrum or waveform, the magnitudes of positive and negative shis are displayed above and below the
baseline, respectively. e distribution of frequency shis
or velocities at any given point in time is encoded in the
brightness of the pixels.
1.1.2.1 Continuous Wave Doppler Ultrasound
Continuous wave (CW) Doppler (. Fig. 1.19) uses two
transducer elements– one continuously transmitting and the
other continuously receiving ultrasound. Blood ow velocity
is calculated from the frequency shi of the signal reected
by the moving red blood cells.
CW Doppler systems may be directional or nondirectional. Nondirectional systems cannot discriminate between
CW Doppler
R
T
f
. Fig. 1.19 Diagram of continuous wave (CW) Doppler ultrasound.
Ultrasound pulses are continuously emitted by the transmitter (T), and
frequency-shifted signals reected by red blood cells moving at dierent velocities (V) are picked up by the receiver (R)
f’
V
positive and negative ow directions. In a directional system, information on the ow direction is extracted from the
phase shi. As ultrasound is continuously transmitted and
received, CW Doppler cannot assign the returning Doppler
signal to a specic depth. Hence, the returning signal contains ow information from all vessels along the beam path.
With arteries and veins oen lying close together, the CW
Doppler signal simultaneously represents arterial and venous
ow. When performed with a high transmit frequency, CW
Doppler allows sensitive examination of supercial vessels.
e advantage of CW Doppler lies in the detection of
high ow velocities without aliasing, which is accomplished
by the use of separate transmit and receive crystals for the
simultaneous emission and reception of ultrasound signals.
1.1.2.2 Pulsed Wave Doppler Ultrasound/
Duplex Ultrasound
Pulsed wave (PW) Doppler (. Fig.1.20) is similar to conventional B-mode scanning in that the same piezoelectric
elements alternately emit ultrasound pulses and receive the
incoming echoes.
e depth from which a returning signal originates can
be determined by calculating the round trip time (based on
knowledge of the speed of sound in tissue) as follows: a short
pulse is emitted, and the system is switched o for some time
before the receive mode is switched on. In this way, only
echoes arriving at the transducer face with the system in the
receive mode are processed, ignoring echoes arriving during the o-mode. e time during which the transducer is
in the receive mode is the range gate. By changing the range
gate, the operator can dene the sample volume or Doppler
window. A typical sample volume encompasses the entire
16
Chapter 1 · Fundamental Principles
diameter of the target vessel. e number of pulses emit-
1
ted per second is the pulse repetition frequency (PRF). e
maximum PRF that can be used decreases with the depth of
the vessel interrogated, as it then takes longer for the echoes
to return to the transducer.
Sound waves travel through the human body at a fairly
constant speed of approx. 1540m/s. Hence, the round trip
time varies with the distance between the reector and the
transmitter, and the operator can dene a scan depth using
a time lter. An electronic gate then opens briey, allowing
only signals from this site to pass, while discarding all echoes
coming in earlier or later. It is thus possible to selectively
record Doppler signals from the specied depth. e combination of PW Doppler with real-time gray-scale imaging
is the basis for duplex ultrasonography. PW Doppler has
the advantage of providing axial resolution (discrimination of vessels along the ultrasound beam), but is limited by
the fact that it fails to adequately record high-velocity signals (depending on the transmit frequency and penetration
depth). Using a single crystal for transmitting and receiving
signals requires a delay between pulses for the processing of
PW Doppler
returning echoes. e longer the pulse delay, the lower the
peak ow velocity that can be detected.
Duplex ultrasound combines 2D real-time imaging with
pulsed Doppler and thus provides ow information from a
sample volume at a dened depth. Duplex scanning enables
calculation of blood ow velocity from the Doppler frequency
R
+
T
shi as the angle of incidence between the ultrasound beam
and the vessel axis can be measured in the B-mode image.
1.1.2.3 Frequency Processing
f
f’
In a blood vessel, blood components move with dierent
velocities, which are represented in the Doppler spectrum by
a range of frequencies with dierent amplitudes reecting the
distribution of ow velocities in the vessel. e spectrum is
analyzed using fast Fourier transform (FFT), which breaks
down the waveform into a series of sinusoidal waveforms.
. Fig. 1.20 Diagram of pulsed wave (PW) Doppler ultrasound. The
transducer alternately emits short ultrasound pulses (T, transmitter)
and records the reected echoes at dened intervals (R, receiver)
For the individual frequency values, the corresponding
amplitudes are calculated and displayed in dierent shades
of gray (. Fig.1.21).
Oscillator
5 MHz
2640 Hz
5 MHz
10 cm/s
30 cm/s
50 cm/s
60 cm/s
. Fig. 1.21 Function of a Doppler transducer. Ultrasound waves are emitted by an oscillator and reected by red blood cells moving through
the vessel at dierent velocities. The signal is reected with a shifted frequency, or Doppler shift, which depends on the speed and relative direction of the moving reectors. The received Doppler signal is composed of a range of frequencies, which have to be sorted by fast Fourier transform (FFT) before they can be displayed over time in the form of a Doppler frequency spectrum or waveform (Diagram courtesy of GE Healthcare)
Analyzer
5.00044 MHz
5.00132 MHz
5.0022 MHz
5.00264 MHz
2200 Hz
1320 Hz
440 Hz
Frequency
3
2
1
0
s
Doppler frequency
1.1 · Technical Principles ofDiagnostic Ultrasound
Amplitude
17
1
Wall filter
a
b
. Fig. 1.22 a Three-dimensional Doppler frequency spectrum showing the distribution of individual Doppler shifts (amplitudes), ow direc-
tions (above and below the time axis), and ow velocities (computed from Doppler frequency shifts). The heights of the boxes correspond to the
amplitudes of the respective Doppler frequencies. A Doppler frequency spectrum represents amplitudes by dierent levels of brightness. In colorcoded duplex ultrasound, the averaged ow velocity at a given point in time (black boxes) is displayed in color according to the ow direction
and superimposed on the two-dimensional gray-scale image in real time (According to P.M.Klews, in Wolf and Fobbe 1993). b Doppler frequency
spectrum of the supercial femoral artery (left section). The histogram plotted on the vertical axis on the left represents the distribution of the different Doppler frequency shifts during systole. In the Doppler waveform, this distribution is represented by dierent levels of brightness (laminar
ow). The right section shows the corresponding distribution during systole in the common carotid artery, which has less pulsatile ow
Time
According to Fourier’s theorem, any periodic waveform can be reconstructed from its component waveforms.
Conversely, in spectral analysis, a complex waveform of a
given frequency (Doppler shi frequency) is decomposed
into its frequency components. In this case, the FFT yields
the amplitudes of the individual frequencies of the respective
sine and cosine functions, which together make up the waveform. e individual frequencies thus separated are continuously displayed over time in the Doppler frequency spectrum
(spectral waveform). e Doppler spectrum contains the following information on blood ow (. Fig.1.22a):
5 e vertical axis representing dierent ow velocities as
Doppler frequency shis
5 e horizontal axis representing the time course of the
frequency shis
5 Density of points, or color intensity, on the vertical axis rep-
resenting the number of red blood cells moving at a certain
velocity (may also be plotted in the form of a histogram)
Flow toward and away from the transducer is processed
simultaneously and respectively represented above and
below the baseline (zero ow velocity line).
Alternatively, some ultrasound devices display the magnitudes of the dierent velocity components in a separate
power spectrum. is is done by measuring the signal intensities of the individual Doppler frequencies at a specic time
in the cardiac cycle and displaying the spectral distribution
in a histogram (. Fig.1.22b; . Table1.7).
1.1.2.4 Blood Flow Measurement
e most important parameters for evaluating and quantifying blood ow that can be derived from the Doppler frequency spectrum are:
5 Peak systolic frequency (mainly relevant for quantifying
stenosis)
5 Peak end-diastolic frequency (stenosis, ow character)
5 Averaged blood ow velocity
18
QmLmean flow velocity cm s
(/ )/
()
min =×60
Chapter 1 · Fundamental Principles
1
. Table 1.7 Spectral displays
Type of spectrum Information displayed
color coding may change as a result of a change in the
ow direction relative to the sector-shaped ultrasound
beam. In this case, the area of transition between red
and blue is black (while it is yellow in aliasing). Black
Power spectrum Display of the power, or strength, of
individual frequencies
Frequency spectrum Display of shifted frequencies or
blood ow velocities over time
Usual mode of display Frequency spectrum
indicates that no Doppler frequency shi information is
obtained because the ultrasound beam is at a 90° angle
to the vessel axis.
In vitro waterbath experiments in which two precision
pumps generated dierent ow proles demonstrated good
Levels of brightness or color represent the density of a given
frequency in the frequency band
correlation (r= 0.98) between the volume ow rates measured by duplex ultrasound and volumetry (Schäberle and
Seitz 1991;
. Fig.1.24).
Even in deeper vessels, highly reproducible measure-
ments can be obtained by performing Doppler inter-
5 Intensity-weighted mean blood ow velocity (which is
the basis for calculation of the volume ow rate)
5 Variance (spectral broadening due to ow disturbances)
rogations at angles as close to 0° as possible to minimize
the eects of errors in angle setting. Repeated ultrasound
measurement of ow in the superior mesenteric artery
performed in 28 fasting subjects in the morning revealed a
Based on these parameters, the following quantities can be
calculated:
5 Angle-corrected peak systolic velocity (PSV) and end-
diastolic velocity (EDV) can be calculated from the
Doppler waveform. Mean ow velocity is calculated on
the basis of the signal intensities.
5 e volume ow rate is calculated from the intensity-
weighted mean blood ow velocity and the vascular
cross-sectional area using the following equation:
day-to-day variation of 11% in peak systolic velocity (PSV)
and of 9.7% in end- diastolic velocity (EDV) (
. Fig. 1.25).
Repeated diameter measurement using the leading-edge
method showed a day- to- day variation of 2.2% (Schäberle
and Seitz 1991).
Another source of error that can lead to over- or underestimation of average ow velocity is to use inadequate transmit or receive gain settings (. Fig.1.26).
e main uncertainty in determining the volume ow
rate, however, arises from the measurement of the vessel
diameter and the resulting inaccuracy in calculating the
.
crosssectional area
- ccm2)
(
cross-sectional area (. Fig.1.27). In B-mode images, vessel
walls appear thicker than their true anatomic size. is is due
Quantitative evaluation of blood ow requires estimation of
the Doppler angle to calculate angle-corrected blood ow
velocity. e Doppler shi alone does not provide this information. To minimize errors in the calculation of blood ow
velocity and other parameters, the angle should be as small as
possible and not exceed 60°.
At a Doppler angle of 60°, an error of ±5° in the estimated
angle of insonation will lead to a 20% error in the calculated
velocity. e magnitude of the error increases disproportionately with the angle of insonation (
. Fig.1.23).
Various measures are available to optimize the angle of
insonation for spectral Doppler interrogation and measurement of blood ow velocity:
5 Use of a unilateral waterpath (linear-array transducer).
5 Electronic beam steering: Successive ring of the
elements in a linear-array transducer produces an
to the so-called blooming eect resulting from the strong
reection of the ultrasound beam at the interface between
blood and the vessel wall (. Fig.1.28b).
In summary, sonographic determination of volume ow
rates is prone to the following pitfalls:
5 Determination of average ow velocity
5 Doppler angle error
5 Uncertainty in the calculation of the vessel cross-
sectional area
5 Inaccuracy in vessel diameter measurement (bloom-
ing eect)
5 Assumption of a circular vessel cross-section.
5 Variation in the cross-sectional area during the
cardiac cycle
5 Respiratory variation in vascular cross-sectional area
(veins)
ultrasound wave that is emitted from the transducer
at a specic angle (to steer the color box and make the
insonation angle as small as possible).
5 Manual manipulation of the transducer (sector and
curved-array transducers): A curved-array transducer
with a small footprint enables a wide range of motion
including angulation for optimization of the Doppler
angle. However, the examiner must be aware that the
e uncertainty in sonographic vessel diameter measurement
can be minimized and systematized by using the leading-toleading-edge (LTL) method and low gain settings. With the
LTL method, the diameter is measured from the reection
of the nearest outer wall to that of the opposite inner wall
. Fig.1.28a). In vitro experiments found a greater accuracy
(
for diameters below 13mm and showed the overestimation
1
Error in calculating flow
1.1 · Technical Principles ofDiagnostic Ultrasound
Cosine function
0
-1
0° 90°
100
80
60
velocity
40
20
0
01020304050607080
a
Angle between Doppler beam and vessel
180° 270° 360°
10°
5°
90°
%
100
90
80
70
60
50
40
30
20
10
Error in calculating volume flow rate
bc
Error > 5°
10 20 30 40 50 60 70 80
Doppler angle of incidence
< 5°
90°
19
1
5°
d
. Fig. 1.23 a The Doppler equation incorporates the angle between the ultrasound beam and the owing blood in the form of the cosine func-
tion (cosα), with the shift being highest when the beam strikes the vessel tangentially (cosine of 0°=1) and lowest when the beam is perpendicular to the direction of blood ow (cosine of 90°=0). The larger the Doppler angle, the greater the resulting error in the velocity calculation in case
of inaccurate placement of the angle correction cursor (graphically shown for errors of 5° and 10°). Such errors are unavoidable, particularly when
aligning the cursor with the vessel wall in curved vessel segments. b The graph illustrates the angle-dependent error in ow measurement for a
misalignment of ±5°. Overestimation of the Doppler angle results in greater error in the velocity calculation than underestimation. c Error in blood
ow velocity calculation resulting from misalignment of the angle correction cursor in vessels running obliquely through the scan plane. Alignment of the angle correction cursor is more dicult if a blood vessel passes obliquely through the scan plane in the B-mode image (left drawing).
An oblique course is suggested if only a short segment of a long straight vessel is depicted. In such a case, the transducer should be turned to
obtain a B-mode scan visualizing a long straight vessel segment (right drawing) for optimal positioning of the angle correction cursor. d Uncertainty concerning the Doppler angle of insonation in a tortuous vessel. In a curved vessel segment, the angle of insonation varies through a range
of 5°–65° over a short stretch, making it dicult to accurately determine the Doppler angle for calculating ow velocity. Left color ow image and
corresponding Doppler waveform: Velocity measurement in a very tortuous internal carotid artery (ICA). With the sample volume positioned in
the curved segment (to conrm or rule out clinically suspected kinking stenosis), a maximum peak systolic velocity (PSV) of 88cm/s and a peak
end-diastolic velocity (EDV) of 24cm/s were calculated with a Doppler angle of 5° (top drawing). Right color ow image and waveform: With an
assumed Doppler angle of 65°, a PSV of 191cm/s and an EDV of 41cm/s were calculated in the curvature of the vessel (bottom drawing)
of diameters to be less severe than the underestimation
reported for the inner-to-inner-edge method (ITI) (Smith
1984). Moreover, use of the LTL method systematizes the
unavoidable measurement error, thereby improving the
reproducibility of measurements.
Diameter variations during the cardiac cycle can be taken
into account by measuring both systolic and diastolic diameters (in the time-motion mode) and considering them in
the ow volume calculation with dierent weightings (1/3
systole +2/3 diastole).
Other parameters that characterize blood ow are the
pulsatility index (PI) and the resistive index (RI) according to
Pourcelot. ese indices have the advantage that they are not
dependent on the Doppler angle of insonation. e resistive
65°
indices, in particular the Pourcelot index, reect wall elasticity as well as the peripheral resistance of the organ supplied
(. Fig.1.28c, d).
e Pourcelot index increases with peripheral resistance,
while end-diastolic velocity (EDV) decreases. Stenosis or
occlusion in peripheral arteries with triphasic ow alters
the Doppler waveform and hence the Pourcelot index. It can
thus serve as a semiquantitative parameter for estimating
the degree of stenosis. In an artery supplying a parenchymal
organ, a relevant decrease in the Pourcelot index between
the prestenotic and the poststenotic segment can be interpreted as indicating hemodynamically signicant stenosis,
for instance, when examining a patient with suspected renal
artery stenosis.
20
p
cm/
cm/s
Error in calculating volume flow rate
Vessel diameter
2mm
Chapter 1 · Fundamental Principles
V
1
d
s
70
60
50
40
30
20
10
0
0102030405060 cm/s
. Fig. 1.24 In vitro ow measurement by duplex ultrasound.
Comparison of mean ow velocity determined by duplex ultrasound
(Vd) and volumetry (Vp). Dierent ow proles were generated by
two precision pumps (I and II). The mean axis shift of 3.75cm/s with
shift of the zero line was due to a software error and was corrected
by the manufacturer following these experiments. Vp=mean actual
ow velocity calculated from volumetrically determined ow rate/
cross-sectional area of the tube; Vd=mean ow velocity determined by
duplex ultrasound (mean of ve individual measurements) (Schäberle
and Seitz 1991)
y = 1.13x + 3.48
Pump I:
1st measurement
2nd measurement
Pump II:
y = x + 3.9
r = 0.99
r = 0.97
V
. Fig. 1.26 Spectral Doppler waveform from the superior mesenteric
artery (bottom) obtained with adequate settings and the corresponding curve of mean ow velocities over time automatically computed by
the ultrasound machine (top). The blood ow velocity averaged over
three cardiac cycles is 31cm/s
%
100
80
60
40
0.2
20
0
02 46 8101
1.0 mm Error
0.5
200
150
100
80
. Fig. 1.25 Peak systolic velocities (PSV) and end-diastolic velocities
(EDV) measured in the superior mesenteric artery of fasting subjects
on two successive days (n=28)
PSV
cm/s
40
30
20
10
. Fig. 1.27 Errors in volume ow rate calculation resulting from
dierent measurement accuracies in determining vessel diameter
(forerrors ranging from 0.2 to 1.0mm)
1.1.3 Physical Principles ofColor-Coded
Duplex Ultrasound
1.1.3.1 Velocity Mode
Color duplex ultrasound combines the presentation of twodimensional (2D) morphologic information with superimposed ow data of a dened area displayed in color. e
frame rate is much lower for the color-coded 2D display
EDV
of ow information than for the conventional (black-andwhite) display because it takes much longer to compute the
2D distribution of ow.
In conventional duplex ultrasound, a small gate (sample
volume) is dened in the real-time gray-scale image for
1.1 · Technical Principles ofDiagnostic Ultrasound
D
(Ieading
edge)
21
1
a
kHz
A
Pulsatility index (PI):
kHz
A
Pourcelot index:
. Fig. 1.28a–d Vessel diameter measurement and resistive indices. a Diagram illustrating the three main methods of sonographic vessel diam-
eter measurement (left part of drawing): The diameter can be measured from the outer wall to the outer wall (outer-to- outer-edge method, OTO,
blue arrows), from the inner wall to the inner wall (inner-to-inner-edge method, ITI, white arrows), or from the outer wall reection closest to the
transducer to the reection from the opposite inner wall (leading-edge-to-leading-edge method, LTL, downward arrows). The LTL method minimizes and standardizes overestimation of vessel diameters due to blooming. b Measurement of the diameter of the superior mesenteric artery
(MS) using the leading-edge method: The images illustrate how the vessel wall is overemphasized as a result of the blooming eect. Systolicdiastolic variation in diameter: the gray-scale scan on the left coincidentally depicts the maximum systolic extension of 7.8mm, while the timemotion display shows the variation in diameter from 7.8mm in systole to 6.9mm in diastole. c Diagrams of resistive indices. The Pourcelot index
is calculated from peak systolic (PSV) and end-diastolic velocities (EDVs), while the pulsatility index (PI) can only be calculated when the system’s
software allows calculation of time-averaged velocity (TAV). d The Pourcelot index, which is typically used in the spectral Doppler evaluation of
parenchymal organ blood ow, is dependent on the patient’s heart rate. In subjects with tachycardia, EDVs are cut o, resulting in a lower Pourcelot index than in patients with bradycardia and the same peripheral resistance. The variation in heart rate and the resulting eect on the Pourcelot
index is illustrated here by the waveform from the external carotid artery (ECA) obtained in a patient with occlusion of the internal carotid artery
(ICA) and absolute arrhythmia: EDV decreases with the length of diastole. In this patient, arrhythmia results in Pourcelot indices that dier by more
than 10% (0.89 and 0.79, calculated from EDVs of 6.7cm/s and 12.5cm/s) (see . Fig. 5.25)
mean
A–B
mean
mean
B
A–B
A
b
dc