Cardiovascular Vincent
Dousset,
MD
a
Felix
W. Wehrli,
Popliteal Artery MR Imaging-US
terms:
Arteries,
Arteries,
popliteal,
namics,
924.12984,
MR
studies,
ies,
924.12984
parative
extremities,
924.12984 924.1299
924.1299 a
a
a
Blood
Magnetic
924.91 Blood, flow Blood vessels,
a
vessels,
resonance
US (MR),
a
dystudcom-
studies
Radiology
1991;
a
Adehine
Louie,
MD
John
a
Listerud,
179:437-441
T
HE
feasibility
of measuring
hemo-
without
dynamic parameters by means magnetic resonance (MR) imaging has been demonstrated for a variety
of time-of-flight 14) methods.
(1-9) However,
none
of which
planar nique,
Fourier which,
stantial
was
is based
accuracy (5,7,15),
on the
flow-encoding in our view,
advantages
niques. The
objective
to assess
oven of the
the
of
and phase (10only a few
studies have assessed the any one of these methods
against
bi-
other
an accepted
standard,
em realized
in the
form
sional embodiments spin-echo detection
repro-
duplex was and
lat-
of one-dirnen-
by means by Feinberg
of and
Mark (17) and a gradient-echo implementation by Hennig et al (13). A second objective of the study was to the
technique’s
potential
detailed information velocity distribution,
not obtainable with Doppler the purpose of this validation, normal popliteal artery was
US. For the chosen.
The rationale for this choice was as follows: (a) The popliteal artery is readily accessible to duplex scanning. (b) It provides a large mange of both forward and reverse velocities covering the entire spectrum of physiobogic flow velocities. (c) It provides a straight vessel segment where flow is
expected tumbed. required parallel
to be relatively (d) No angular because the to the z axis.
encoding
undiscorrection is vessel is nearby
first
Market
St.
Ste
420,
Philadelphia,
PA
19104.
RSNA,
1991
METHODS
Subjects Popliteal artery flow dynamics were studied in 11 male and female volunteers (age range, 21-53 years; mean, 37 years)
(13),
where
the
in a plane
parallel
of the
velocity-encoding
moment
vent
velocity
olution
aliasing.
of 2 cm/sec
locity-encoding
spins
With per
steps
a velocity
pixel for
and
each
res-
64 ye-
velocity
di-
rection (128 total), a range of velocities from 0 to 128 cm/sec is covered. Temporal resolution was achieved by electrocardiographically triggering the pulse sequence. The intersequence delay (temporal resolution) chosen was 43 msec. Further, the pulse sequence, which is illustrated
pensated
in
on the
Figure
2, is velocity
com-
section-selection/readout
axis.
The protocol used consisted lowing steps: A series of axial bocalizer images were obtained
by means
Re-
ceived October 30, 1990; revision requested December 28; revision received January 17, 1991; accepted January 22. Address reprint requests to F.W.W. C
AND
vascular
gradient is governed by its amplitude and the spacing between the two lobes of opposite polarity. Spacing and step size thus determine the velocity resolution, which needs to be appropriately selected to pre-
of a transmit/receive
to determine
velocity-encoding
MATERIALS
peripheral
to the flow direction and subsequently velocity-sensitized by means of stepping a bipolar gradient in the direction of flow (Fig 1). Velocity is encoded by stepping a bipolam balanced gradient (18), which takes the place of the spatial phase-encoding gradient and imparts a first moment on the spins, as originally suggested by Moran (16), followed by the collection of a gradient echo. In contrast to Moran’s approach, however, velocity is encoded in one spatial direction only to achieve clinically practical imaging times. The
coil Hospi3440
known
Imaging
are excited
study
and
MR
velocity
tech-
present
accuracy
any
abnormalities, recruited in the Department of Radiology in the investigators’ institution.
All MR data were obtained on a 1 .5-T clinical imager (Signa; GE Medical Systems, Milwaukee) with use of biplanar
techhas sub-
ultrasound (US). The method first suggested by Moran (16)
for providing on intmaluminal
the Department of Radiology, University of Pennsylvania,
PhD
of
ducibility of the Fourier flow-encoding technique by calibrating it
highlight
I From tal of the
MD,
Hemodynamics: Correlation’
Temporally resolved velocity measurements in the pophiteal arteries of 11 healthy subjects were obtained by means of magnetic resonance (MR) imaging with use of the Fourier flow-encoding technique. Excellent agreement with corresponding Doppler ultrasonography (US) data (r .97, slope 0.99, intercept -1.5 cm/sec) was demonstrated over the entire velocity range from 50 to -20 cm/sec. The method was rapid and its implementation straightforward. Further, MR imaging was shown to provide the intraluminal velocity distribution relevant for the determination of true flow rates, not obtainable with Doppler US. Index
PhD
Radiology
the
level
image
of the folspin-echo bilaterally
extremity at which
would
be
the ob-
tamed. The site of measurement selected was an arbitrary but easily reproducible landmark-the intercondylar space-to ensure that the site of analysis was the same in all subjects. The parameters of the velocity-encoding pulse sequence were as follows: repetition time, 43 msec (mini-
437
mum
intersequence
delay);
echo
time,
30
msec; flip angle, 30#{176}; field of view, 16 cm; section thickness, 5 mm; number of signals averaged, one; velocity resolution, 20 mm/sec per pixel; and number of velocity encodings, 128. Under these conditions, 15-25
temporal
commodated, heart rate. velocity
increments
could
placement. Peak forward and revelocities of the three wave compoderived from the two modalities
were
then
identity
correlated, was
encoding
was
the
Doppler
of
2 mm-
duplex
subjected to The Doppon an ATL 9
instrument
(Advanced
Technology Laboratories, Bothell, Wash) with use of a 5-MHz transducer. The cursor was placed at the location of interest (ie, at the center of the vessel) with the US beam at an angle of 45#{176}-60#{176} (19-21).
The velocity
measured
regres-
Figure 3 shows a typical set of MR data, consisting of a transaxial bocalizem image (Fig 3a) with which the velocity analysis was performed, as
(Fig 3b-3d). phase-dependent velocity
From
waveform
meimages
an analysis flow images, was
with multiple flow (A, B, and
(Fig
4c). This
to result of the
from pressure
behavthe
opgradi-
ent built up along the wave propagation front following ventricular contraction, and vascular compliance (20,21). The correlation of the maximum central velocity data from the two modalities (peak forward and reverse velocities for the peaks labeled A, B, and C) is illustrated in Figure 5. A correlation coefficient of .97, a slope of 0.99, and an intercept of -1.5 cm! sec (suggesting that MR overesti-
RESULTS
of temporally flow-encoding
(19,22)
ion is known posing effects
sion.
well as a series solved Fourier
US
Mark
by linear
oscillation and reverse
C peaks)
ac-
equivalent
The same volunteers were Doppler (duplex) US analysis. ler waveforms were obtained
damped forward
and the line of
determined
depending on the subject’s The total imaging time for the
128 heartbeats or approximately utes per extremity.
Ultra
be
cursor verse nents
of the the
determined
(Fig 4). Note the close correspondence of the MR-derived waveform with the Doppler velocity waveform, which exhibits the characteristic RF
was the maxi-
-411)
A V
mum (central) velocity, representing the fastest-moving red blood cells, which is also the quantity measured with MR imaging.
TE
-5
____
a-
Vessels
Data
Analysis
The Fourier tamed
maximum central velocity in the flow-encoding images was obby measuring the maximum excur-
sion
of the
in
each
resolved was
velocity data, ated no more tion
C F E
bolus
temporally correction
of the
direction). Doppler
SIIca
Imaging
made
of the
images. to the
No angular tllpate
MR-derived
as the vessel typically devithan 10% from the omienta-
velocity-encoding
gradient
(z
US velocities were read off the velocity waveform by means of
Figure 1. Principle ing technique. Spins
of flow-encoding imagare excited in a plane
perpendicular
velocity
signal
is read
to
the
out in the same
vector,
and
3.
(a) Transaxial
spin-echo
localizer
tion
and
readout
compensated.
the
direction.
echo
time,
gradients
RF
GFE
are
first-moment
radio frequency, flow-encoded gradient. =
TE
=
C.
b.
a. Figure
Figure 2. Flow-encoding pulse sequence. Instead of spatial phase encoding, velocity is encoded by means of a bipolar gradient (G) coincident with the section-selection orientation. This gradient is stepped to encode velocity. Note further that the section-selec-
15-25
image,
showing the flow bolus (white arrow) from popliteal set of temporally resolved velocity-encoding images tion of the velocity during cardiac cycle. Annotations R wave (in percent).
together
with
flow-encoding
artery (black arrow). (b-d) in normal subject showing pertain to the phase delay
image, Typical evolufrom the
d.
438
a
Radiology
May
1991
0
10
20
30
40
50
60
70
80
9(
% PHASE b.
C.
Figure
4. (a) Velocity waveform, derived from the temporally resolved MR images of Figure 3b-3d, plotted as a percentage of the delay from the R wave. Note the tripolar waveform characteristic of the damped oscillation resulting from the vascular resistance to the pressure wave and the vessel’s compliance. Peaks labeled A, B, and C are the first forward, first reverse, and second reverse maximum velocities, mespectively. LPA = left popliteal artery. (b) Duplex localizer image with cursor indicating site of velocity measurement. (C) Doppler US waveform.
ple the entire RR interval. Finally, the temporal resolution achievable with MR imaging may be inferior that achievable with US, although
MR is capable
of correctly
delineat-
ing the complex bidirectional form. On the positive side,
MR imaging
as well as the Pouncebot index (also denoted “resistive index”), defined as (VA VB)/ VA, where VA and VB represent the velocities pertaining to peaks A and B, respectively. The latter is a measure of vascular impedance (20,23). The remarkable degree of mutual agreement observed is indicative of the high accuracy of MR imaging in hemodynamic analysis. :60.40.20
0 Maximum
20 Velocity
40
60
80
1001201s0
us
Figure 5. Correlation MR imaging velocity
DISCUSSION
of Doppler US and data in 11 subjects. Ve-
locities compared are the maximum central velocities. Note the three clusters of data points pertaining to the A, B, and C peaks, respectively (see Fig 4a). The correlation coefficient was .97 with a slope of 0.99 and an intercept of -1.5 cm/sec. indicating the excellent accuracy of the MR measurements. The high-velocity A peak (arrow) found in one of the test subjects with both modalities is abnormal.
Hemodynamic niphenal vascular flow encoding
mates
flow
found.
velocity
by
less
than
5%)
The data is summarized in the Table, which lists means and standard deviations for the three peaks as denived by means of the two modalities,
Volume
179
a
Number
2
distinguishes
ages
itself
in
several respects from Doppler US. It has, of course, the disadvantage of not being real time and therefore of being sensitive to variations in heart rate. A further limitation (which can be circumvented by using retrospective gating) (24) is the failure to sam-
information
of Figure
6, which
were
ac-
quired to illustrate the significant changes in velocity profile occurring during different phases of the cardiac cycle. Whereas the profile is Poiseuibhan during most of the cardiac cycle, this is not the case during the occurnence of strong acceleration or decel-
flow of the pewith MR means of
assessing the hemodynamic panametens of the peripheral arterial system. The method is found to be accurate oven a wide range of velocities and its implementation straightforward. It thus lends itself as an add-on to a morphologic imaging examination, with only insignificantly prolonged total examination time.
MR imaging
were
analysis system is a viable
wavehowever,
not obtainable with Doppler US in that it allows the complete distribution of velocities across the vessel lumen. This becomes evident in the higherresolution Fourier flow-encoding im-
eration, when vail, causing
(cm/sec)
provides
to
behavior
the flow
inertial forces to be pluglike.
is present,
preThis
for exam-
ple, during the phases preceding and following directional inversion of flow (ie, where the slope is maximal) (Fig 6e, 6j, 6n). Subsequently, the flow becomes predominantly laminan, as indicated by a more parabolic profile (Fig 6f-6h, 6k, 6o). During meversal of the flow direction, flow near the vessel wall exceeds the flow centrally (Fig 6i, 6m) (25,26), that is, opposite to the behavior found duning laminar flow. The multiple reversal of the flow direction is a consequence of the yessel’s elastic properties (vascular compliance) causing damped oscillation (22), a property characteristic of a penipherab arterial circuit (19). It is well known impedance
that,
during is lowered
stress, vascular (through an
Radiology
a
439
a.
b.
C.
d.
e.
f.
8
h.
i.
j-
k.
1.
q.
r.
m.
n.
Figure bution field
of view
selected
Fourier flow-encoding lumen during complete
was
10 cm,
corresponding
except during phases of strong acceleration flow near the vessel wall exceeds the central particularly head in
well
seen
in
p.
0.
6. High-resolution across the vascular
f
(straight
arrow).
images cardiac
of popliteal artery cycle. The images
subject showing temporal changes in velocity distriwere obtained with the protocol outlined earlier, except that the pixel size of 0.39 mm. During most of the cardiac cycle flow is laminar (f, h, k, o), where it is more pluglike (e, j, n). Further, during reversal of flow direction,
to a spatial
and deceleration, velocity (i, m). Note The
popliteal
vein
in normal
also the superposition
is displayed
of a small
as a signal
of constant
adapting process to satisfy the increased demand for oxygen and thus increased blood flow to the muscles), which causes virtual disappearance of the reversal in flow direction
flow
(F) on the
rate
of the popliteal
highly
laminar
artery, flow
(arrow-
assumption CONCLUSIONS
of radial
symmetry:
F
27
=
i:
rv1(r)dr,
(26,27).
The practical implications knowledge of the intmaluminal
of the veboc-
where
ity
permit-
beats
distributions
the
are
calculation
in milliliters
per
in their
of true second,
flow
rates
which
is
difficult to achieve with Doppler US because of the limitations imposed on sample gate size (28). Flow rates are more significant than flow velocities since the former obviously function of the vessel’s radius vessel’s diameter, which varies
among individuals) and and pathologic changes. Since we locity profile
crement
branch and
C).
the
ting
collateral amplitude
and,
further,
since
the
yein-
yes-
can fairly readily be from the axial spin-echo
Of course, the rates derived in ically depend on precision of the ments.
For
surements further
that
the
following
nc.
440
a
Radiology
example,
the
gives
for each requirement
is difficult
vessels.
which
number
of
accuracy of flow this manner will cnitthe accuracy and geometric measure-
nor
bution
the
cardiac is that
neither
intraluminal
be fully
diam-
phase. A the pixel
the luminal a condition
allow which
quantitation cannot
fied phy.
of true flow rates, accurately be quanti-
by means The short
of Doppler examination
for
small
well
vessel
study of the hemodynamics tracranial vasculature.
velocity
radially
distri-
sonogratime
(2-3
minutes) and the excellent neproducibibity (no dependence on transducer placement) are further benefits. Also, unlike in US, a single data set can provide velocity data for more than one vessel. Another potential application in which US fails to perform
the
to satisfy
Finally, may
vessel
the cardiac cycle compliance, separate mea-
size be much less than diameter of the vessel,
bocahizem image, we can, in principle, determine the flow mate (in milliliters per minute) by numerically solving equation,
the
n the number of temporal phases, r0 the vessel radius, and v1(r) the velocity at radial position r for the ith temporal increment.
eter changes during as a result of vascular which would require
physiologic
directly observe the for every temporal
sel’s diameter determined
are a (ie, the
N represents per minute,
The present study indicates that measuring blood flow with MR imaging is likely to create a role for this modality in vascular examinations, as it is capable of providing both functional and morphologic information. Further, MR imaging is shown to have the potential for providing detailed hemodynamic information, which includes the intrabuminal yebocity distribution, and thus should
due
On
symmetnot
the
real
to lack
other time
of a bone
hand, and
thus
window of the
MR
imaging
the
quality
May
is in-
is of
1991
the data is sensitive to variations in RR interval. Other potential problems relate to the possible noncoincidence of the flow-encoding gradient and vessel orientation, which requimes application of an angle correction or, alternatively, oblique orientation of the flow-encoding gradient, both of which can be achieved by prior acquisition of an angiographic localizer image. Finally, the data analysis currently is labor-intensive, requiring measurement of the bolus excursion
by
means
of a cross-hair
6.
7.
8.
9.
10.
cursor on each of the images. Howevem, this may be a temporary deficiency, since opportunities exist for automating the data analysis procedune. U
11.
12.
References 1.
2.
3.
Axel L. Blood flow effects in magnetic resonance imaging. AJR 1984; 143:11571166. Axel L, Shimakawa A, MacFall JR. A time-of-flight method of measuring flow velocity by magnetic resonance imaging. Magn Reson Imaging 1986; 4:199-205. Wehrli
FW,
Shimakawa
A, MacFall
JR.
sist
5.
Tomogr
1985;
K, Matsudav
Visualization tive analysis MR imaging. 199.
T, Sakurai
15.
16.
T, et al.
of moving fluid: quantitaof blood flow velocity using Radiology 1986; 159:195-
17.
18.
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179
a
Number
2
Wehrli
FW.
Time-of-flight
effects
in
MR
imaging of flow. Magn Reson Med 1990; 14:187-193. Edelman RR, Mattle HP, Kleefield HP, Silver MS. Quantification of blood flow with dynamic MR imaging and presaturation bolus tracking. Radiology 1990; 171:551-556. Van Dijk P. Direct cardiac NMR imaging of heart wall and blood flow velocity. Comput Assist Tomogr 1984; 8:429-436. Nayler G, Firmin DN, Longmore DBJ. Blood flow imaging by cine magnetic resonance. J Comput Assist Tomogr 1986; 10:715-722. Walker MF, Souza SP, Dumoulin CL. Quantitative flow measurement in phase contrast MR angiography. J Comput Assist Tomogr 1988; 12:304-313. Hennig J, Mueri M, Brunner P. Friedburg H. Fast and exact flow measurements with the fast Fourier flow technique. Magn Reson Imaging 1988; 6:369-372. Merboldt KD, H#{227}nicke W, Frahm J. Rapid Fourier
9:537-545.
Wehrli FW, Shimakawa A, Gullberg GT, MacFall JR. Time-of-flight MR flow imaging: selective saturation recovery with gradient refocusing. Radiology 1986; 160:781-785. Shimizu
14.
et
al. MR imaging of venous and arterial flow by a selective saturation-recovery spin-echo (SSRSE) method. J Comput As4.
13.
Hennig J, Mueri M, Brunner P. Friedburg H. MR imaging of flow using the steadystate selective saturation method. J Comput Assist Tomogr 1987; 11:872-877. Matsuda T, Shimizu K, Sakurai T, et al. Measurement of aortic blood flow with MR imaging: comparative study with Doppler US. Radiology 1987; 162:857-861.
flow
imaging
using
FLASH
se-
quences. Magn Reson Med Biol 1988; 1:137-145. Maier SE, Meier D, Boesiger P. Moser UT, Vieli A. Human abdominal aorta: cornparative measurements of blood flow with MR imaging and Doppler US. Radiology 1990; 171:487-492. Moran PR. A flow velocity zeugmatographic interlace for NMR imaging in hu. mans. Magn Reson Imaging 1982; 1:197203. Feinberg
A, Mark
AS.
Human
brain
mo-
tion and cerebrospinal fluid circulation demonstrated with MR velocity imaging. Radiology 1987; 163:793-799. O’Donnell M. NMR blood flow imaging using multi-echo phase contrast sequences. Med Phys 1985; 12:59-64.
19.
Standness DEJ. niques in vascular EF, ed. Noninvasive in vascular
20.
21.
22.
24.
25.
26.
27.
St Louis:
Mosby,
1985;
waveforms:
problems
and
solu-
tions. In: Bernstein EF, ed. Noninvasive diagnostic techniques in vascular disease. St Louis: Mosby, 1985; 40-57. Pourcelot L. Applications cliniques de l’examen Doppler transcutan#{233}. In: Peroneau L, ed. Velocimetrie ultrasonore Doppler. Paris: INSERM, 1974. Glover GH, Pelc NJ. A rapid-gated cine MRI technique. Magn Reson Annu 1988; 299-234. Milnor RM. Pulsatile pressure and flow. In: Milnor WR, ed. Hemodynamics. Baltimore: Williams & Wilkins, 1982; 97-134. Edelman RR, Rubin JB, Buxton RB. Physics and instrumentation, In: Edelman RR, Hesselink JR. eds. Clinical magnetic resonance imaging. Philadelphia: Saunders, 1990; 109-182. Caro CG. The mechanics of the circulation.
28.
disease.
11-18. Nelson TR, Pretorius DH. The Doppler signal: where does it come from and what does it mean? AJR 1988; 151:439-447. Taylor KJW, Holland S. Doppler US. I. Basic principles, instrumentation, and pitfalls. Radiology 1990; 174:297-307. Johnston KW, Kassam MS. Processing Doppler signals and analysis of peripheral arterial
23.
Doppler ultrasonic techdisease. In: Bernstein diagnostic techniques
New
1978. Fronek ments ity.
In:
agnostic Louis:
York:
Oxford
University
A. Quantitative in arterial disease Bernstein
EF,
techniques Mosby,
1985;
ed.
velocity of lower
Press,
measureextrem-
Noninvasive
in vascular 554-562.
di-
disease.
St
Dousset,
MD
a
Felix
W. Wehrli,
Popliteal Artery MR Imaging-US
terms:
Arteries,
Arteries,
popliteal,
namics,
924.12984,
MR
studies,
ies,
924.12984
parative
extremities,
924.12984 924.1299
924.1299 a
a
a
Blood
Magnetic
924.91 Blood, flow Blood vessels,
a
vessels,
resonance
US (MR),
a
dystudcom-
studies
Radiology
1991;
a
Adehine
Louie,
MD
John
a
Listerud,
179:437-441
T
HE
feasibility
of measuring
hemo-
without
dynamic parameters by means magnetic resonance (MR) imaging has been demonstrated for a variety
of time-of-flight 14) methods.
(1-9) However,
none
of which
planar nique,
Fourier which,
stantial
was
is based
accuracy (5,7,15),
on the
flow-encoding in our view,
advantages
niques. The
objective
to assess
oven of the
the
of
and phase (10only a few
studies have assessed the any one of these methods
against
bi-
other
an accepted
standard,
em realized
in the
form
sional embodiments spin-echo detection
repro-
duplex was and
lat-
of one-dirnen-
by means by Feinberg
of and
Mark (17) and a gradient-echo implementation by Hennig et al (13). A second objective of the study was to the
technique’s
potential
detailed information velocity distribution,
not obtainable with Doppler the purpose of this validation, normal popliteal artery was
US. For the chosen.
The rationale for this choice was as follows: (a) The popliteal artery is readily accessible to duplex scanning. (b) It provides a large mange of both forward and reverse velocities covering the entire spectrum of physiobogic flow velocities. (c) It provides a straight vessel segment where flow is
expected tumbed. required parallel
to be relatively (d) No angular because the to the z axis.
encoding
undiscorrection is vessel is nearby
first
Market
St.
Ste
420,
Philadelphia,
PA
19104.
RSNA,
1991
METHODS
Subjects Popliteal artery flow dynamics were studied in 11 male and female volunteers (age range, 21-53 years; mean, 37 years)
(13),
where
the
in a plane
parallel
of the
velocity-encoding
moment
vent
velocity
olution
aliasing.
of 2 cm/sec
locity-encoding
spins
With per
steps
a velocity
pixel for
and
each
res-
64 ye-
velocity
di-
rection (128 total), a range of velocities from 0 to 128 cm/sec is covered. Temporal resolution was achieved by electrocardiographically triggering the pulse sequence. The intersequence delay (temporal resolution) chosen was 43 msec. Further, the pulse sequence, which is illustrated
pensated
in
on the
Figure
2, is velocity
com-
section-selection/readout
axis.
The protocol used consisted lowing steps: A series of axial bocalizer images were obtained
by means
Re-
ceived October 30, 1990; revision requested December 28; revision received January 17, 1991; accepted January 22. Address reprint requests to F.W.W. C
AND
vascular
gradient is governed by its amplitude and the spacing between the two lobes of opposite polarity. Spacing and step size thus determine the velocity resolution, which needs to be appropriately selected to pre-
of a transmit/receive
to determine
velocity-encoding
MATERIALS
peripheral
to the flow direction and subsequently velocity-sensitized by means of stepping a bipolar gradient in the direction of flow (Fig 1). Velocity is encoded by stepping a bipolam balanced gradient (18), which takes the place of the spatial phase-encoding gradient and imparts a first moment on the spins, as originally suggested by Moran (16), followed by the collection of a gradient echo. In contrast to Moran’s approach, however, velocity is encoded in one spatial direction only to achieve clinically practical imaging times. The
coil Hospi3440
known
Imaging
are excited
study
and
MR
velocity
tech-
present
accuracy
any
abnormalities, recruited in the Department of Radiology in the investigators’ institution.
All MR data were obtained on a 1 .5-T clinical imager (Signa; GE Medical Systems, Milwaukee) with use of biplanar
techhas sub-
ultrasound (US). The method first suggested by Moran (16)
for providing on intmaluminal
the Department of Radiology, University of Pennsylvania,
PhD
of
ducibility of the Fourier flow-encoding technique by calibrating it
highlight
I From tal of the
MD,
Hemodynamics: Correlation’
Temporally resolved velocity measurements in the pophiteal arteries of 11 healthy subjects were obtained by means of magnetic resonance (MR) imaging with use of the Fourier flow-encoding technique. Excellent agreement with corresponding Doppler ultrasonography (US) data (r .97, slope 0.99, intercept -1.5 cm/sec) was demonstrated over the entire velocity range from 50 to -20 cm/sec. The method was rapid and its implementation straightforward. Further, MR imaging was shown to provide the intraluminal velocity distribution relevant for the determination of true flow rates, not obtainable with Doppler US. Index
PhD
Radiology
the
level
image
of the folspin-echo bilaterally
extremity at which
would
be
the ob-
tamed. The site of measurement selected was an arbitrary but easily reproducible landmark-the intercondylar space-to ensure that the site of analysis was the same in all subjects. The parameters of the velocity-encoding pulse sequence were as follows: repetition time, 43 msec (mini-
437
mum
intersequence
delay);
echo
time,
30
msec; flip angle, 30#{176}; field of view, 16 cm; section thickness, 5 mm; number of signals averaged, one; velocity resolution, 20 mm/sec per pixel; and number of velocity encodings, 128. Under these conditions, 15-25
temporal
commodated, heart rate. velocity
increments
could
placement. Peak forward and revelocities of the three wave compoderived from the two modalities
were
then
identity
correlated, was
encoding
was
the
Doppler
of
2 mm-
duplex
subjected to The Doppon an ATL 9
instrument
(Advanced
Technology Laboratories, Bothell, Wash) with use of a 5-MHz transducer. The cursor was placed at the location of interest (ie, at the center of the vessel) with the US beam at an angle of 45#{176}-60#{176} (19-21).
The velocity
measured
regres-
Figure 3 shows a typical set of MR data, consisting of a transaxial bocalizem image (Fig 3a) with which the velocity analysis was performed, as
(Fig 3b-3d). phase-dependent velocity
From
waveform
meimages
an analysis flow images, was
with multiple flow (A, B, and
(Fig
4c). This
to result of the
from pressure
behavthe
opgradi-
ent built up along the wave propagation front following ventricular contraction, and vascular compliance (20,21). The correlation of the maximum central velocity data from the two modalities (peak forward and reverse velocities for the peaks labeled A, B, and C) is illustrated in Figure 5. A correlation coefficient of .97, a slope of 0.99, and an intercept of -1.5 cm! sec (suggesting that MR overesti-
RESULTS
of temporally flow-encoding
(19,22)
ion is known posing effects
sion.
well as a series solved Fourier
US
Mark
by linear
oscillation and reverse
C peaks)
ac-
equivalent
The same volunteers were Doppler (duplex) US analysis. ler waveforms were obtained
damped forward
and the line of
determined
depending on the subject’s The total imaging time for the
128 heartbeats or approximately utes per extremity.
Ultra
be
cursor verse nents
of the the
determined
(Fig 4). Note the close correspondence of the MR-derived waveform with the Doppler velocity waveform, which exhibits the characteristic RF
was the maxi-
-411)
A V
mum (central) velocity, representing the fastest-moving red blood cells, which is also the quantity measured with MR imaging.
TE
-5
____
a-
Vessels
Data
Analysis
The Fourier tamed
maximum central velocity in the flow-encoding images was obby measuring the maximum excur-
sion
of the
in
each
resolved was
velocity data, ated no more tion
C F E
bolus
temporally correction
of the
direction). Doppler
SIIca
Imaging
made
of the
images. to the
No angular tllpate
MR-derived
as the vessel typically devithan 10% from the omienta-
velocity-encoding
gradient
(z
US velocities were read off the velocity waveform by means of
Figure 1. Principle ing technique. Spins
of flow-encoding imagare excited in a plane
perpendicular
velocity
signal
is read
to
the
out in the same
vector,
and
3.
(a) Transaxial
spin-echo
localizer
tion
and
readout
compensated.
the
direction.
echo
time,
gradients
RF
GFE
are
first-moment
radio frequency, flow-encoded gradient. =
TE
=
C.
b.
a. Figure
Figure 2. Flow-encoding pulse sequence. Instead of spatial phase encoding, velocity is encoded by means of a bipolar gradient (G) coincident with the section-selection orientation. This gradient is stepped to encode velocity. Note further that the section-selec-
15-25
image,
showing the flow bolus (white arrow) from popliteal set of temporally resolved velocity-encoding images tion of the velocity during cardiac cycle. Annotations R wave (in percent).
together
with
flow-encoding
artery (black arrow). (b-d) in normal subject showing pertain to the phase delay
image, Typical evolufrom the
d.
438
a
Radiology
May
1991
0
10
20
30
40
50
60
70
80
9(
% PHASE b.
C.
Figure
4. (a) Velocity waveform, derived from the temporally resolved MR images of Figure 3b-3d, plotted as a percentage of the delay from the R wave. Note the tripolar waveform characteristic of the damped oscillation resulting from the vascular resistance to the pressure wave and the vessel’s compliance. Peaks labeled A, B, and C are the first forward, first reverse, and second reverse maximum velocities, mespectively. LPA = left popliteal artery. (b) Duplex localizer image with cursor indicating site of velocity measurement. (C) Doppler US waveform.
ple the entire RR interval. Finally, the temporal resolution achievable with MR imaging may be inferior that achievable with US, although
MR is capable
of correctly
delineat-
ing the complex bidirectional form. On the positive side,
MR imaging
as well as the Pouncebot index (also denoted “resistive index”), defined as (VA VB)/ VA, where VA and VB represent the velocities pertaining to peaks A and B, respectively. The latter is a measure of vascular impedance (20,23). The remarkable degree of mutual agreement observed is indicative of the high accuracy of MR imaging in hemodynamic analysis. :60.40.20
0 Maximum
20 Velocity
40
60
80
1001201s0
us
Figure 5. Correlation MR imaging velocity
DISCUSSION
of Doppler US and data in 11 subjects. Ve-
locities compared are the maximum central velocities. Note the three clusters of data points pertaining to the A, B, and C peaks, respectively (see Fig 4a). The correlation coefficient was .97 with a slope of 0.99 and an intercept of -1.5 cm/sec. indicating the excellent accuracy of the MR measurements. The high-velocity A peak (arrow) found in one of the test subjects with both modalities is abnormal.
Hemodynamic niphenal vascular flow encoding
mates
flow
found.
velocity
by
less
than
5%)
The data is summarized in the Table, which lists means and standard deviations for the three peaks as denived by means of the two modalities,
Volume
179
a
Number
2
distinguishes
ages
itself
in
several respects from Doppler US. It has, of course, the disadvantage of not being real time and therefore of being sensitive to variations in heart rate. A further limitation (which can be circumvented by using retrospective gating) (24) is the failure to sam-
information
of Figure
6, which
were
ac-
quired to illustrate the significant changes in velocity profile occurring during different phases of the cardiac cycle. Whereas the profile is Poiseuibhan during most of the cardiac cycle, this is not the case during the occurnence of strong acceleration or decel-
flow of the pewith MR means of
assessing the hemodynamic panametens of the peripheral arterial system. The method is found to be accurate oven a wide range of velocities and its implementation straightforward. It thus lends itself as an add-on to a morphologic imaging examination, with only insignificantly prolonged total examination time.
MR imaging
were
analysis system is a viable
wavehowever,
not obtainable with Doppler US in that it allows the complete distribution of velocities across the vessel lumen. This becomes evident in the higherresolution Fourier flow-encoding im-
eration, when vail, causing
(cm/sec)
provides
to
behavior
the flow
inertial forces to be pluglike.
is present,
preThis
for exam-
ple, during the phases preceding and following directional inversion of flow (ie, where the slope is maximal) (Fig 6e, 6j, 6n). Subsequently, the flow becomes predominantly laminan, as indicated by a more parabolic profile (Fig 6f-6h, 6k, 6o). During meversal of the flow direction, flow near the vessel wall exceeds the flow centrally (Fig 6i, 6m) (25,26), that is, opposite to the behavior found duning laminar flow. The multiple reversal of the flow direction is a consequence of the yessel’s elastic properties (vascular compliance) causing damped oscillation (22), a property characteristic of a penipherab arterial circuit (19). It is well known impedance
that,
during is lowered
stress, vascular (through an
Radiology
a
439
a.
b.
C.
d.
e.
f.
8
h.
i.
j-
k.
1.
q.
r.
m.
n.
Figure bution field
of view
selected
Fourier flow-encoding lumen during complete
was
10 cm,
corresponding
except during phases of strong acceleration flow near the vessel wall exceeds the central particularly head in
well
seen
in
p.
0.
6. High-resolution across the vascular
f
(straight
arrow).
images cardiac
of popliteal artery cycle. The images
subject showing temporal changes in velocity distriwere obtained with the protocol outlined earlier, except that the pixel size of 0.39 mm. During most of the cardiac cycle flow is laminar (f, h, k, o), where it is more pluglike (e, j, n). Further, during reversal of flow direction,
to a spatial
and deceleration, velocity (i, m). Note The
popliteal
vein
in normal
also the superposition
is displayed
of a small
as a signal
of constant
adapting process to satisfy the increased demand for oxygen and thus increased blood flow to the muscles), which causes virtual disappearance of the reversal in flow direction
flow
(F) on the
rate
of the popliteal
highly
laminar
artery, flow
(arrow-
assumption CONCLUSIONS
of radial
symmetry:
F
27
=
i:
rv1(r)dr,
(26,27).
The practical implications knowledge of the intmaluminal
of the veboc-
where
ity
permit-
beats
distributions
the
are
calculation
in milliliters
per
in their
of true second,
flow
rates
which
is
difficult to achieve with Doppler US because of the limitations imposed on sample gate size (28). Flow rates are more significant than flow velocities since the former obviously function of the vessel’s radius vessel’s diameter, which varies
among individuals) and and pathologic changes. Since we locity profile
crement
branch and
C).
the
ting
collateral amplitude
and,
further,
since
the
yein-
yes-
can fairly readily be from the axial spin-echo
Of course, the rates derived in ically depend on precision of the ments.
For
surements further
that
the
following
nc.
440
a
Radiology
example,
the
gives
for each requirement
is difficult
vessels.
which
number
of
accuracy of flow this manner will cnitthe accuracy and geometric measure-
nor
bution
the
cardiac is that
neither
intraluminal
be fully
diam-
phase. A the pixel
the luminal a condition
allow which
quantitation cannot
fied phy.
of true flow rates, accurately be quanti-
by means The short
of Doppler examination
for
small
well
vessel
study of the hemodynamics tracranial vasculature.
velocity
radially
distri-
sonogratime
(2-3
minutes) and the excellent neproducibibity (no dependence on transducer placement) are further benefits. Also, unlike in US, a single data set can provide velocity data for more than one vessel. Another potential application in which US fails to perform
the
to satisfy
Finally, may
vessel
the cardiac cycle compliance, separate mea-
size be much less than diameter of the vessel,
bocahizem image, we can, in principle, determine the flow mate (in milliliters per minute) by numerically solving equation,
the
n the number of temporal phases, r0 the vessel radius, and v1(r) the velocity at radial position r for the ith temporal increment.
eter changes during as a result of vascular which would require
physiologic
directly observe the for every temporal
sel’s diameter determined
are a (ie, the
N represents per minute,
The present study indicates that measuring blood flow with MR imaging is likely to create a role for this modality in vascular examinations, as it is capable of providing both functional and morphologic information. Further, MR imaging is shown to have the potential for providing detailed hemodynamic information, which includes the intrabuminal yebocity distribution, and thus should
due
On
symmetnot
the
real
to lack
other time
of a bone
hand, and
thus
window of the
MR
imaging
the
quality
May
is in-
is of
1991
the data is sensitive to variations in RR interval. Other potential problems relate to the possible noncoincidence of the flow-encoding gradient and vessel orientation, which requimes application of an angle correction or, alternatively, oblique orientation of the flow-encoding gradient, both of which can be achieved by prior acquisition of an angiographic localizer image. Finally, the data analysis currently is labor-intensive, requiring measurement of the bolus excursion
by
means
of a cross-hair
6.
7.
8.
9.
10.
cursor on each of the images. Howevem, this may be a temporary deficiency, since opportunities exist for automating the data analysis procedune. U
11.
12.
References 1.
2.
3.
Axel L. Blood flow effects in magnetic resonance imaging. AJR 1984; 143:11571166. Axel L, Shimakawa A, MacFall JR. A time-of-flight method of measuring flow velocity by magnetic resonance imaging. Magn Reson Imaging 1986; 4:199-205. Wehrli
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