Download Magnetic Resonance Angiography Techniques

Magnetic Resonance Angiography Techniques
G. Laub1, J. Gaa2, M. Drobnitzky1
1
2
Siemens AG, Medical Engineering Group, Erlangen, Germany
Institute for Clinical Radiology, University Hospital of Mannheim, Germany
Introduction
MR Imaging (MRI) depends on relaxation times T1
and T2 and spin density r. Furthermore, MRI is very
sensitive to motion resulting in a variety of flow effects.
On one hand, flow effects are responsible for a number
of artifacts which can drastically impair the diagnostic
value of the images; on the other hand, flow effects
can be used advantageously to develop non-invasive
techniques for imaging of the vascular anatomy. These
techniques are referred to as Magnetic Resonance
Angiography (MRA). The techniques most commonly
used for MRA can be classified into two major categories: Time-of-flight and phase contrast MR angiography. Both techniques rely on separate physical effects,
and will result in images with different information
about the vasculature. More recently, the use of contrast
agents in combination with ultra-fast T1-weighted
imaging sequences has shown significant improvements
in the delineation of the vessel lumen. It is important to
note that a proper use of MRA techniques and correct
interpretation of MR angiographic images requires
a knowledge of the underlying physical mechanisms of
flow sensitivity in MRI [1, 2].
Inflow Enhancement
The basic idea of Time-of-flight MR Angiography
(ToF MRA) is demonstrated in Figure 1a. The blood
flow is assumed to be perpendicular to the imaging
plane (or volume in the case of a 3D study). For repetition times TR shorter than the longitudinal relaxation
time T1 of the stationary spins within the slice, the signal
will be reduced due to partial saturation effects. Blood
flow in the vessel will move spins from outside the slice
which have not been subjected to the spatially selective
rf pulses into the imaging slice. These unsaturated or
fully relaxed spins have full equilibrium magnetization
and, therefore, upon entering the slice will produce a
much stronger signal than stationary spins assuming
that a gradient echo sequence is applied. This effect
is referred to as “entry slice phenomenon”, or “inflow
enhancement”, or “flow related enhancement”.
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electromedica 66 (1998) no. 2
a
Ds = v · TR
TR << T1
saturated spins
flow
v
unsaturated spins
vessel
S
slice
1a
1b
Figure 1a
Basic idea of ToF MR
Angiography.
For pulse repetition times TR
less than T1 the stationary spins
within the slice are partially
saturated and, therefore, will
produce only little signal.
Spins moving into the slice
are fully relaxed, resulting in
a high signal intensity.
Figure 1b
Cross-section in the neck.
High vessel contrast is achieved
due to saturation of stationary
spins and inflow of fully relaxed
spins. Signal from the veins is
suppressed by an additional saturation pulse which is placed on
the cranial side of the imaging
slice.
A typical application of this effect is demonstrated
in Figure 1b which shows a thin section acquired in
the neck. Due to the relatively large flip angle of 40°
in combination with a short pulse repetition rate of
30 msec, the signal from stationary tissue is almost
saturated. In general, blood flow signal from the veins
can be sufficiently removed by the application of additional saturation pulses which are applied on the cranial
side of the slice. A repetition of the same sequence at
different slice positions results in a series of images
which show the vascular tree as a sum of individual slices.
I
Imax
The amount of inflow enhancement depends on
several factors, including tissue specific parameters like
T1, sequence specific parameters (flip angle a, and TR),
and geometrical parameters like slice thickness and
orientation, or blood flow velocity.
Postprocessing of MR Angiography Data
In principle, with this technique any vessel segment
can be imaged by cutting through the vessel perpendicularly with regard to the flow direction. With repetitive
increments of the slice position a three-dimensional data
set of the complete vascular structure can be measured
[3, 4]. For the observer, this form of representation
requires experience in order to obtain the correct threedimensional spatial impression. Obviously, postprocessing methods should be used to extract two-dimensional
projections of the vasculature from the three-dimensional volume data [5]. With these methods threedimensional perception can be obtained in two ways – by
showing a sequence of projective images with different
projection angles or by coding of the depth information
onto the surface of the displayed objects.
Since the surfaces of most vessels are relatively small,
the first method – multiple projections with different
angles – has proven more useful in practice. The starting
point for this method must be a three-dimensional data
set in which the structures to be extracted are associated
with a characteristic range of signal intensity levels. In
this case a projective image can be calculated by penetrating the data volume with a set of parallel projection
rays and selecting along each of these rays only the data
point that represents the maximum intensity as demonstrated in Figure 2. The inflow enhancement and proper
pulse-sequence parameters (flip angle, pulse repetition
time, and flow-compensation) assure that the maximum
intensity is always associated with a blood vessel, as long
as the projection ray intersects at least one [6, 7]. All of
the other projection rays will just pick up a background
pixel intensity of the three-dimensional data set. As a
result, Figure 3 demonstrates a complete projection
image at different viewing angle calculated from one
single 3D data set.
By varying the projection angle multiple projective
images can be obtained retrospectively which allow the
observer to obtain the correct spatial perception of the
Figure 2
Principle of Maximum Intensity
Projection (MIP). Along each ray
only the pixel with the highest
signal intensity I max is projected
onto the imaging plane.
A
A
B
C
B
D
C
E
D
E
Figure 3
Application of MIP post-processing to a 3D data set in the neck.
Different views can be calculated
retrospectively to create a 3D
perception of the carotid arteries.
three-dimensional data set. By displaying a number of
projections with projection increments of only a few
degrees in a rapid fashion, the perception of a continuously rotated object will be generated which allows a
correct three-dimensional visualization of such complex
structures as a vessel tree.
In some instances, particularly for the best assessment
of a vessel narrowing, it is recommended to also evaluate
the source images in addition to the MIP results. In complex anatomical situations as in the case of the depiction
electromedica 66 (1998) no. 2
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of the renal arteries it is clearly favorable to restrict the
MIP algorithm to those parts of the whole data volume
which covers the vasculature of interest thereby avoiding
projection-related misinterpretations [9].
ToF MR Angiography Techniques
ToF MR Angiography can be classified into three
major categories as demonstrated in Figure 4. On the left
side (Figure 4a), a sequential two-dimensional technique
is shown which provides multiple thin sections of the
vessels. Alternatively, a three-dimensional technique
can be applied. As demonstrated in Figure 4b, the whole
volume is excited simultaneously, and will then be
subdivided into thin partitions, or slices by using an
additional phase encoding scheme in the slice select
direction. Unlike in 2D imaging, where the slice resolution is defined by the excitation profile of the radiofrequency pulse, the slice resolution is defined by spatially
encoding magnetic field gradients and can be less than
1 mm. Both techniques – 2D and 3D – are currently used
in clinical applications. There are specific advantages,
and disadvantages related to each of the techniques.
While 2D techniques offer a high vessel/background
contrast which is pretty much constant over the coverage
of the vasculare structure, the vessel/background contrast
in 3D techniques is typically lower, and is progressively
getting smaller when spins penetrate through the
imaging volume. The slab thickness, or vessel coverage
in 3D techniques is therefore limited to a distance at
which blood signal approaches a steady state signal.
Typically, 3D techniques are applicable in combination
with fast flow situations, while 2D techniques may be
applied for the visualization of slower flow as well.
Another aspect in selecting two- or three-dimensional
techniques is related to the spatial resolution. In 2D techniques the spatial resolution is defined by the in-plane
resolution (FOV divided by the matrix size, or number
of lines, respectively), and the slice thickness of the
sequence. Typically, in-plane resolution may be isotropic, like 0.8 mm x 0.8 mm, but slice thickness is more
than that, like 2-3 mm, resulting in an anisotropic volume
data set. Isotropic resolution, i. e. the voxel size has equal
length in all directions, can be obtained with 3D techniques, which also offer a better signal-to-noise ratio due
to the averaging effect of the phase encoding in the slab
direction. Small vessels, as they typically occur in the
intracranial vasculature are generally better visualized
with a 3D technique, while larger vessel with a predominantly uni-directional flow, e. g. the common carotid
artery, may be visualized very well with 2D techniques.
Figure 4
Techniques for
ToF MR Angiography:
sequential 2D (a),
volume imaging (b),
multi-slab (c)
slice #
1
2
3
slab 4
slab 3
single slab
(3DFT)
slab 2
slab 1
N
flow
flow
a
b
a
small
flow
c
a
water
large
MT pulse
Magnetization reduction
as a result of
magnetization transfer
flow
macromolecules
frequency
imaging
volume
5a
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dipolar interaction
between water
and macromolecules
5b
Figure 5
Techniques to improve vessel
contrast in ToF MR Angiography.
5a: Ramped rf pulses are used to
partially compensate for the progressive saturation of spins in 3D
ToF MRA.
5b: Magnetization transfer
effects are used to improve the
vessel-background contrast.
One of the major limitation in three-dimensional
ToF MRA is the loss of vessel contrast as spins are
penetrating into the imaging volume. This effect is due
to progressive saturation when spins are experiencing
the rf excitation pulses in the imaging volume.
A technique which is used to address the spin saturation in 3D ToF MRA techniques is referred to as multislab MRA, or MOTSA (see Figure 4c) [8]. Several slabs
are used to create sufficient coverage over the vasculature. Each of the slabs is thin enough to avoid significant
spin saturation within this slab. All of the slabs are
acquired in a sequential mode, i.e. one slab is acquired
after the other. The advantage of this technique is that
it combines the isotropic resolution capability of the
3D technique with a relatively small amount of spin
saturation similar to the 2D techniques. The disdvantage
of the multi-slab technique is related to the imperfections
of the slice profile. Therefore, all of the slabs need
to overlap by about 20-30% to avoid venetian blind
effects in areas in between the slabs.
A more effective use of the magnetization is possible
when using variable flip angles across the slab as shown
in Figure 5a. At the entrance plane when spins enter with
equilibrium magnetization a relatively small flip angle
is used which still provides sufficient signal, with only
little effects on the longitudinal magnetization. The flip
angle is increased deeper in the imaging volume to
compensate for the reduction of the longitudinal
magnetization of the spins on their way through the
volume, and so forth. In principle, it is possible to shape
the flip angle distribution over the entire imaging volume
according to the specific flow velocity and vessel
coverage.
Another improvement is possible by the application of
magnetization transfer pulses (MTC) as demonstrated in
Figure 5b. The idea is to use off-resonance rf pulses
which do not directly affect the mobile protons which are
used to create the signal in MRI. Protons with restricted
mobility, however, do get saturated, and because of cross
correlation or chemical exchange processes the magnetization will be transfered to some biological tissue such
as grey, or white matter resulting in a partial saturation.
Blood will not be affected by the MTC pulses, and as a
result, there will be more contrast between blood and
background.
Flow-Induced Phase Shift
The second class of MR Angiography techniques is
based on the changes in the phase of the transverse
magnetization. These phase shifts occur when the spins
move along the magnetic field gradients which are used
for position encoding in 2D and 3D Fourier imaging.
For a gradient pulse pair as shown in Figure 6 the flowinduced phase shift FV is given as
FV = g • t • A • v
t
FV = g • A • t • v My
flow
FV
+A
-A
time
stationary
Mx
Figure 6
Motion along a bipolar gradient
pulse pair results in a flow-induced
phase shift of the transverse
magnetization. Stationary spins
will not be affected by the bipolar
gradient pulse pair. The amount
of phase shifts depends on the
area of each gradient pulse, and
distance between the pulses.
which shows a linear relationship between the phase shift
of a moving spin and its velocity. The flow-induced
phase shift can be used for flow imaging and flow
quantification techniques.
Phase Contrast MR Angiography
The basic idea of phase-contrast MRA is to acquire
two data sets with a different amount of flow sensitivity
as shown in Figure 7 [10]. The first data set S1 is
acquired with a flow compensated sequence, i.e. without flow sensitivity. The second data set S2 is acquired
with a flow sensitive sequence. The amount of flow
sensitivity is controlled by the strength of the gradient
pulse pair which is incorporated into the sequence.
After completion of the measurement, both signals are
subtracted to create a complex difference ÆS. The length
of the difference ÆS depends on the phase shift FV
within each pixel. An image which shows the signal
intensity ÆS represents the velocity of spins at each point
within the field of view. Both 2D and 3D acquisition
techniques can be applied with phase contrast MRA
[11-15]. Alternatively, the phase shift FV can be evaluated from the measurement of S1 and S2. As shown
above, the phase shift FV is proportional to the spin’s
velocity, and therefore the measurement of FV allows
the quantitative assessment of flow velocities.
The difference signal ÆS has a maximum value for
opposite directions of S1 and S 2, i. e. for FV = 180°. This
velocity is typically referred to as venc, and depends on
the sequence design, i.e. pulse amplitude and distance
between the gradient pulse pair which is used for flow
encoding. For velocities larger than venc the difference
signal is decreased constantly until it gets zero for
FV = 360°. Therefore, in phase contrast (PC) MRA it is
important to correctly set the venc of the sequence to the
maximum flow velocity which is expected during the
measurement.
electromedica 66 (1998) no. 2
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My
My
My
FV
Mx
Mx
flow compensated
flow sensitive
A bipolar gradient pulse pair produces a phase shift
which depends on the velocity component along this
gradient. For a complete measurement of the flow velocity more measurements are necessary with orthogonal
flow sensitization, e. g.
image 1: flow compensated
image 2: flow sensitization along x (e.g. read gradient)
image 3: flow sensitization along y (e. g. phase encode
gradient)
image 4: flow sensitization along z (e. g. slice select
gradient)
postprocessing is done in the following way:
v read = image 1 – image 2
v phase = image 1 – image 3
v slice = image 1 – image 4
V = = (v2read + v2phase + v2slice)
Gslice
Gphase
Gread
Figure 8
Typical sequence diagram for
phase contrast MR Angiography.
A bipolar flow encoding pulse
pair is applied in subsequent
measurements in slice, phase,
and read direction, respectively.
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S2
ÆS
FV
S1
Mx
Figure 7
Basic idea of phase contrast
Angiography. The complex
subtraction ÆS of two data sets
which are acquired with a
different amount of flow sensitivity (e. g. flow compensated
S1 and flow sensitized S2)
produces an image with signal
intensities depending on local
flow velocities.
subtraction
Contrast-Enhanced MR Angiography
The signal response of a Flash-type sequence as a
function of the flip angle a is shown in the following
diagram (Fig. 8) for tissues of various longitudinal
relaxation times T1 assuming a sequence with a pulse
repetition time TR of 10 msec.
In contrast-enhanced MR Angiography (CE-MRA),
the use of a paramagnetic extracellular contrast agent
(such as Gd-DTPA) increases the blood signal [16, 17].
This is due to shortening of the T1 relaxation time of
blood after contrast injection. The amount of signal
enhancement with T1 shortening is as shown in Figure 9.
Depending on the actual concentration of Gd-DTPA
the arterial blood T1 can be as short as 50 ... 100 msec,
substantially shorter than the T1 of fat. Therefore, blood
produces the largest signal, and thus the vessel lumen
will be picked up with the maximum intensity program
(MIP) to create an MR angiogram [18-20].
Following the intravenous (IV) injection of the contrast agent, it is delivered to the vasculature of interest.
Local blood signal is substantially enhanced. This
result is optimal when data collection (in particular the
center of k-space) occurs right when the contrast agent
arrives at the vessels being imaged as shown in Figure
10. Therefore, timing of the scan with respect to the
intravenous injection of contrast agent is very important
in CE-MRA. In a clinical setting the contrast arrival
time is best determined by administering a small testbolus and continuously monitoring for vascular signal
enhancement in the region of interest using a very fast
gradient-echo technique with fluoroscopic image update. Furthermore, there will be venous enhancement
visible when the acquisition window is long enough to
catch the venous phase as well. Therefore, the acquisition window is limited to about 10 to 15 seconds in the
carotid artery to avoid excessive signal enhancement in
the jugular vein [22]. Similar constraints exist for the
first-pass bolus width for imaging of the abdominal
vasculature [21, 23].
One major technical prerequisites for CE-MRA is the
availability of a high-performance gradient system to
reach the very short echo and repetition times needed
to cover a resonable 3D-volume in the time given by
the arterio-venous transit time. At the same time it is
mandatory to use CP-phased array coils for complete
volume coverage, fitted to the anatomy, to optimize
signal-to-noise.
The clinical significance of CE-MRA is expected to
grow due to a number of reasons: T1-shortening as the
source of signal increase is virtually immune against
signal loss due to flow-related dephasing especially for
in-plane flow as encountered in ToF- and PC-MR
angiography. This is clearly advantageous in situations
of tortuous vessels and stenotic regions. Due to the fact
that Gd-DTPA has no nephrotoxic potential its use in
kidney studies is widely accepted. Furthermore it is
reported in the literature that the depiction of large aneu-
risms is markedly simplified with CE-MRA compared
to DSA because filling and subsequent washout of
contrast material is not immediate but can be traced
over a longer period of time. The main advantage
of CE-MRA compared to ToF- and PC-MRA is its intrinsic advantage in acquisition speed, if sufficient
strong gradients are available. This results from the fact
that the slice orientation can now be chosen independent
of the primary vessel direction and must not longer be
perpendicular to the vessel. Instead, the image plane can
now be chosen to be along any major vessel axis.
We consider CE-MRA as an emerging new vascular
application of MR. Its specific advantages with respect
to conventional DSA is currently been studied at major
research centers. Outcome analysis and its potential
to reduce costs are currently considered by a growing
number of research groups.
Figure 9
Relative signal intensity versus
flip angle for a spoiled gradientecho sequence (FLASH) for
tissues of various T1, and a TR of
10 msec. The flip angle a is
selected to produce the largest
signal for the short T1 tissues
signal intensity
TR = 10 ms
T1 = 50 ms
T1 = 100 ms
T1 = 260 ms (fat)
T1 = 1200 ms (blood)
arterial phase
venous phase
contrast
injection
8 sec
12 sec
16 sec
20 sec
24 sec
28 sec
Figure 10
Timing of contrast-enhanced
MR Angiography. In this study a
series of ultrafast 3D measurements with only 4 sec scan time
was acquired right after bolus
injection of a single-dose
Gd-DTPA. Note rapid filling
of the aorta and subsequent
filling of the renal vein, as an
example.
32 sec
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Figure 11
Dual-phase contrast-enhanced
MRA study. The first measurement (left) is done during the
arterial phase. Twenty seconds
later a second measurement
is performed to get an image
during the venous phase. Each
measurement is done during
a breath-hold of 24 seconds.
Figure 12
Carotid MR Angiography.
Left: 2D-ToF technique.
The stripe artifact is due to
swallowing right during the
acquisition of this particular
slice.
Right: Contrast-enhanced 3D
MRA technique. A 3D volume is
acquired with a scan time of
only 10 seconds. Due to the short
acquisition time the venous
enhancement is relatively small.
A stenosis is demonstrated in
the left internal carotid artery.
Figure 13
Thoracic MR Angiography. Two
maximum intensity projections
are shown from the same 3D
data set. A 1.5 mm isotropic
resolution is acquired in a single
breath-hold of 23 seconds
after injection of a double-dose
Gd-DTPA.
Figure 14
Application of contrastenhanced MR Angiography in
the abdominal vasculature.
Left: Demonstration of
aneurysms in the abdominal
aorta.
Right: Demonstration of multivessel disease in the aorta and
femoral artery.
electromedica 66 (1998) no. 2
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Author’s address:
Dr. Gerhard Laub
Siemens AG, Medical Engineering Group
Henkestrasse 127
D-91052 Erlangen, Germany
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