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356 15 Imaging Blood Circulation Using Nuclear Magnetic Resonance
Differing methods are available for hyperpolarization, producing polarization of
up to 40% in some C-labeled substances [14]. Dilution in the vascular system causes
the concentration of a hyperpolarized imaging agent to decrease. This effect can be
cancelled out through hyperpolarization because the signal can be enhanced up to
10
6
times. This allows the imaging of changes in molecular structure, which would
not have been previously possible (e.g., metabolic processes).
At present, the T
1
of the available
13
C and
15
N substances are short when
compared with nuclei, such as
18
fluorodeoxyglucose (FDG), used in clinical PET
procedures. This means that after polarization, advanced chemical synthesis cannot
be performed. Clinically acceptable relaxation times may be attainable with an
injectable water solution obtained through the application of a general polarization
process to a wide range of small organic molecules. Developments in low-field MRI
might enable longer relaxation rates in vivo [15].
Reported T
1
and T
2
values of the
13
C-containing molecules, used in in vivo
imaging experiments, are typically 10 –100 times longer than those of
1
H resulting
in in vivo values of T
1
20–45 s and T
2
2–5 s [16, 17]. Because the signal from thermal
equilibrium polarized
13
C nuclei is far below the limit of detection, following an
administration of hyperpolarized
13
C, the MR signal that is detected can only come
from the injected substance. This differs from standard proton imaging where
the contrast substance injected affects the relaxation times of protons in proximity
to the paramagnetic molecules. Hyperpolarized
13
C imaging has no background
signal apart from patient and coil/receiver system noise, making subtraction of
images unnecessary. The T
1
value of a given substance dictates the expected signal
amplitude in the area of imaging at a defined time after injection. Imaging pulse
methods take advantage of the long T
2
values for multiecho imaging strategies;
however, the imaging agent and the clinical application are the determining factors
in selecting the applied pulse sequence.
As hyperpolarized
13
C MRI gives direct information regarding the molecule, to
which hyperpolarized atoms are attached, it makes tissue and cell viability research
possible. NMR’s capacity for distinguishing signals from the tracer in different
molecules [16, 17] has more potential for direct molecular imaging than SPECT or
PET, which are limited to imaging only the distribution of the active nuclei. The
visualization of differing molecules containing
13
C with chemical shift imaging
makes it feasible to generate frequency and spatial information simultaneously.
15.2
MR-Angiography
In the early days of MRI, the signal from flowing blood was considered a nuisance
as it generated frequent artifacts, which detracted from image quality. It was some
time before it was realized that the signal could be exploited to generate diagnostic
quality images of blood vessels. Although magnetic resonance-angiography (MRA)
is frequently regarded as a specific subdiscipline, it is important to remember that
blood vessels can be identified on virtually all MR images.
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15.2 MR-Angiography 357
15.2.1
MR-Angiography Techniques
In the following sections, we describe the important techniques used in MRA.
15.2.1.1 Time-of-Flight MR Angiography (TOF MRA)
In time-of-flight magnetic resonance-angiography (TOF MRA), flow-compensated
GE imaging sequences are optimized to favor the vascular signal over that of the
surrounding tissue. This can be achieved by using RF pulses with a high flip
angle while using a short repetition time (TR). The stationary tissue signal will get
saturated due to short TR because the longitudinal magnetization does not have
time to relax back to equilibrium magnetization. The signal in a train of RF pulses
reduces to a low steady state of magnetization.
5)
At the same time, this method favors the inflow effect because the blood flowing
into the imaging plane has not been affected and its longitudinal magnetization is
maximal. The signal from the nonsaturated blood flow is higher than that of the
saturated tissues.
The signal in TOF MRA depends on hemodynamic, geometrical, and tissue
conditions as well as sequence parameters. The hemodynamics of the inflow alter
the amount of nonsaturated magnetization entering the imaging region. Complex
or turbulent (stenosis) flow, which is linked to spin dephasing and slow flow, limits
the TOF MRA sensitivity.
The length and orientation of the vessel is an important factor in the display
of blood vessels. The vascular signal will be optimized by placing the slice
perpendicular to the axis of the vessel, while vessels oriented parallel to the slice
plane may not be discernable due to saturation effects. Either arteries or veins
can be selectively visualized by placing a presaturation band on one side of the
imaging slice, thus suppressing unwanted arterial or venous flow. The suppression
of static tissue signal can be improved on with preparation pulses. These employ
magnetization transfer (MT) saturation, which affects mainly tissue as it contains
more large molecules or selective excitation of the water or fat saturation. Poor
suppression of the signal from stationary tissues can happen in tissue with short T
1
relaxation, for example, those containing fat, atheroma, hematoma, and thrombi.
The sequence parameters available to optimize contrast between static and flowing
magnetization are TR, flip angle, echo time (TE), and slice thickness. Two imaging
modalities are used for TOF MRA: 2D and 3D imaging. In 2D acquisition,
time-of-flight imaging allows the use of a set of thin slices. Thin slices have better
sensitivity to slow flows (which do not remain in the slice for long and will therefore
5) In many applications, fast MRI methods
are used to achieve high temporal and spa-
tial resolution. In order to achieve this, the
repetition time has to be decreased. The
repetition time has an impact on the signal
intensity. To optimize the signal for this
purpose, one can use the so-called Ernst
angle [18]. This repetitive excitation with
RF pulses leads to a steady state of longi-
tudinal magnetization. The Ernst angle is
derived by optimizing TR, T
1
, and the flip
angle relative to each other. One finds for
the Ernst angle α = arccos(exp(−TR/T
1
).
The Ernst angle is generally important in
achieving good S/N ratio.
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358 15 Imaging Blood Circulation Using Nuclear Magnetic Resonance
not be saturated). It also gives the possibility of using high flip angles (giving better
stationary tissue saturation and an increased vascular signal). However, this has
the disadvantage of poor spatial resolution along the axis of the slice stack as they
are stacked up to reconstruct a pseudo-volume.
On the contrary, 3D TOF MRA gives good spatial resolution in all three spatial
directions, with a better signal-to-noise (S/N) ratio for fast flow. However, each
repetition excites the total volume, producing a progressive saturation of the flows,
even more so when they are slow. The slowest flows may even disappear entirely
due to saturation. Flow saturation can be reduced as it passes through the explored
volume by dividing 3D acquisition into slabs and by using ramped RF excitation, a
technique that progressively the flip angle changes from one side of the imaging
volume to the other [19].
15.2.1.2 Phase Contrast Angiography
PCA describes a method whereby phase shifts are experienced by protons moving
through a bipolar gradient. For a bipolar gradient of an imparted length and
strength, the moving spins dephase in proportion to their velocity v
enc
as shown in
Eqs. (15.4a,b). The potential values of the phases range from −π to +π. However,
beyond this range of possible values, aliasing occurs, which causes the velocity to
be poorly encoded. Hence, the characteristics of the encoding gradient only encode
flows within specific velocity parameters from −v
enc
to +v
enc
as defined by the user.
PCA consists of two imaging sequences, both containing bipolar velocity-encoding
gradients. They differ by the polarity of the gradients. This results in two images
with opposite phase encoding, which allows differentiation between flowing and
static magnetization. The acquisition in 2D PCA imaging is fast, making it useful
for testing differing encoding speeds. The slice plane is perpendicular to the flow
direction for the quantitative study of flows. In 3D PCA imaging, each partition
is encoded in three directions, which results in better quality images than in 2D
PCA due to reconstruction of the images in maximum intensity projection (MIP).
Three-dimensional PCA is a more sensitive imaging technique for slow flows than
Figure 15.4 Carotids – MIP of the
whole data set showing normal appear-
ance of the head and neck arteries on
contrast-enhanced MRA.
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15.2 MR-Angiography 359
TOF where there is a tendency toward saturation. Parallel imaging can reduce the
scan time of these sequences enormously [20].
15.2.1.3 Contrast-Enhanced MR-Angiography (CEMRA)
Contrast-enhanced methods have better signal-to-noise ratios and are faster than
noncontrast agent injection techniques. There are many benefits in Gadolinium
chelate-induced T
1
shortening, such as marked increase in the vascular signal
that is independent of inflow effects and virtual complete of saturation of the
blood signal that allows larger volumes to be examined in any plane. Furthermore,
turbulent flow imaging is improved. A 4D dynamic MRA is also possible, enabling
an injected bolus to be observed while flowing through the system. However,
accurate timing is necessary when injecting the bolus if the acquisition is to occur
simultaneously with the intravascular passage of the contrast agent. Inaccuracy in
the timing can lead to either ringing artifact if central k-space data are acquired too
early while the T
1
value is rapidly changing or contamination by venous signal if
acquired too late (Figure 15.4).
15.2.1.4 Choice of the Underlying Technique
Blood vessels are usually recognized as a function of the presence and velocity
of flowing blood contained therein, which present a characteristic signal on MR
images. In the case of SE images, blood usually appears dark, while on GE images,
vessels usually appear bright. Alteration of the contour of the vessel, the presence of
abnormal intraluminal content, and alterations in blood flow velocities compared
to adjacent or paired vessels disrupts the normal signal from flowing blood and
usually gives information about the state of the blood vessel. GE sequences that
demonstrate bright blood vessels (bright blood MRA) form the basis of most MRA
approaches looking at the vessel lumen. However, there are exceptions to this
generalization, for example, for delineation of pathology in the blood vessel wall
where SE images play a larger role. Similarly, dark blood images are routinely used
to accurately delineate the cardiac wall in patients in whom abnormality of the
cardiac muscle is suspected.
15.2.1.5 Fast Imaging
Since most MRA techniques employ fast acquisitions, systems equipped with
fast gradients are necessary. Progressive improvements in gradient performance,
parallel imaging, partial k-space (e.g., keyhole imaging), time-resolved imaging,
(e.g., TRICKS), and so on have resulted in scanning techniques with sufficient
spatial and temporal resolution to address most clinical issues.
15.2.2
Clinical Applications
15.2.2.1 Different Diseases Usually Affect Arteries and Veins
Although many different disorders can affect blood vessels, diseases usually affect
arteries by causing stenosis (narrowing) within them or by causing abnormal
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360 15 Imaging Blood Circulation Using Nuclear Magnetic Resonance
dilatation of the vessel (aneurysm formation). Although the pathology causing
these end effects usually starts within the blood vessel wall, the wall is usually too
thin to image reliably and a ‘‘road map’’ of the lumen of the vessel is acquired.
This is usually performed by acquiring relatively thin cross-sectional images of the
blood vessel by using a technique that portrays the blood vessel as the brightest
structure on the image, and then usually a computer-based algorithm such as MIP
to reproject the relatively large number of slices into a (pseudo) three-dimensional
image that portrays the arterial tree as if it were extracted from the other anatomy.
As a high degree of vessel to background contrast improves the efficiency of
this extraction algorithm, efforts are aimed at boosting intravascular signal while
minimizing background signal. Although imaging of the blood vessel wall, and
particularly the factors that lead to narrowing of the lumen are highly sought after
prizes, delineation of the components of atherosclerosis (the most common cause
of narrowing) remains elusive. Improvements in wall delineation will depend on
development of improved surface coils, higher resolution, and better methods for
flow-related artifact and motion correction.
15.2.2.2 Requirement for MR Imaging of Blood Vessels
As MRI employs neither ionizing radiation nor invasive arterial catheterization, it
potentially offers an ideal or near ideal modality for blood vessel imaging. However,
in the early days of MRA when techniques were limited to TOF and PCA methods, it
rapidly became apparent that long examination times and frequent artifacts would
limit widespread use of MRA in clinical practice despite the fact that the study was
virtually devoid of side effects (occasional contraindications of claustrophobia and
the presence of implanted devices excepted). The introduction of contrast-enhanced
techniques into clinical practice overcame many of the limitations of TOF and PCA.
The benefit of contrast-enhanced magnetic resonance-angiography (CEMRA) lay
in the fact that intravascular signal was no longer dependent on flow character-
istics (which varied between patients and between different vessels in the same
patient), and blood flow velocities that frequently declined with a parallel decline
in intravascular signal now became largely irrelevant. Intravascular signal relied
primarily on the degree of T
1
shortening which was determined by the injection
rate of gadolinium. CEMRA became the standard bearer for MRA because of its
inherent high vessel contrast, unrivalled CNR, freedom from artifacts, and speed
of performance.
For CEMRA, images are acquired during first arterial passage of an intravenously
injected bolus by synchronizing peak arterial enhancement with central k-space
data. As the arterial phase (before circulation of the contrast agent into the veins)
is relatively short, this mandates that image data is rapidly collected at the correct
time. This is facilitated by either measuring the circulation time before the MRA
with a test bolus or using ‘‘bolus detection,’’ a term to denote the process whereby
the circulation of contrast into the region of interest (ROI) can be visualized in
real time by acquiring a low-resolution thick-slice image that is reconstructed and
displayed in real time on the monitor, which allows the operator to commence the
3D scan immediately on arrival of the contrast agent within the artery of interest.
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15.2 MR-Angiography 361
Even by exploiting fast 3D GRE acquisitions with acceleration techniques (parallel
imaging), the acquisition time may extend well beyond the imaging window
(often as short as 10 s) of arterial enhancement. However, the unique nature of
k-space data facilitates the development of ‘‘pure’’ arteriograms despite this fact by
synchronizing the contrast-defining central k-space data with the arterial phase of
contrast enhancement.
Since it was so widely believed that gadolinium-based contrast agents were
almost devoid of serious side effects, the penalty of intravenous cannulation and
the minimal risk of adverse reaction to the injected contrast agent were dismissed as
largely irrelevant. The fact that gadolinium-based contrast agents were substantially
safer than iodinated contrast agents (injected for X-ray catheter angiography and
computed tomographic angiography – CTA – the main competitor to MRA) was
noted. For these reasons, CEMRA proliferated widely in clinical practice at the
expense of noncontrast techniques that were used less often. However, increasing
awareness of therare but potentiallyfatal condition ofnephrogenic systemic fibrosis
(NSF), which is strongly associated with high-dose gadolinium usage especially in
patients with kidney failure, has prompted renewed interest in noncontrast MRA
techniques.
15.2.2.3 Difference in Gadolinium Contrast Agents
When intravenous contrast agents are injected for MRA, it is important to remem-
ber that there are some differences between the agents, the majority of which are
gadolinium-based. Different contrast agents for MRA differ in their relaxivity, for-
mulation strength, and degree of protein binding. The ‘‘standard’’ contrast agents
are formulated at 0.5 mmol l
−1
and are well proven as effective agents in clinical
practice. A single agent (Gadovist, Berlex) is formulated at 1.0 mmol l
−1
. Also,
different contrast agents differ in their degree of protein binding, which affects the
relaxivity. For example, one of the ‘‘standard’’ 0.5 mmol l
−1
agents exhibits weak
protein binding (most of the other agents have essentially no protein binding),
which improves its’ relaxivity and therefore gives greater T
1
shortening on a volume
comparison with other agents (Multihance, Bracco). More recently, a contrast agent
with strong protein binding has been introduced into clinical practice. Because of
the strong protein binding in plasma, it does not rapidly diffuse into the tissue
and thus persists for a long period within the circulation. The long intravascular
persistence of agents such as MS325 (the only one launched as a clinical product)
leads to its classification as a blood-pool contrast agent (Ablavar, Lantheus).
15.2.2.4 Noncontrast MRA in Clinical Practice
TOF and PCA methods have been rapidly displaced by CEMRA for many applica-
tions. In some instances, for example, for evaluation of the intracranial arteries,
TOF techniques remain the approach of choice (Figure 15.6). Also, although results
have been variable in clinical practice, coronary MRA remains the preserve of non-
contrast approaches to date. For the majority of vascular territories where CEMRA
had become the technique of choice, renewed interest in noncontrast techniques
was driven by concerns about safety of the contrast agent (gadolinium) injection.
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362 15 Imaging Blood Circulation Using Nuclear Magnetic Resonance
(a) (b)
Figure 15.5 (a,b) TOF MRA in two orthogonal planes
demonstrating normal appearance of the intracranial arteries
at 3.0 T in (a) lateral and (b) craniocaudal projection.
Although a noncontrast technique to replace CEMRA as the prime technique for
noninvasive vascular assessment has not been found, steady progress in newer
noncontrast techniques has resulted in some competing techniques that may ad-
equately allow decision making in clinical practice. For example, ECG-gated 3D
partial-Fourier fast spin echo (FSE) methods and 3D balanced steady-state free
precession (SSFP) methods have been reported for imaging of the larger vessels
(thoracic aorta and pulmonary arteries). Data from the peripheral circulation sug-
gest that delays in vascular enhancement on noncontrast time-resolved images
may reflect stenosis severity. In addition, although time-resolved imaging is tradi-
tionally thought as the preserve of fast contrast-enhanced techniques, ECG-gated
partial-Fourier FSE method, and adiabatic fast passage (AFP) inversion at 3 T
(longer T
1
relaxation times at 3 T benefit AFP by prolonging the duration of tagging
effect) with SSFP offer a promising noninvasive MR angiographic method for
time-resolved MRA.
15.2.2.5 From Anatomy to Function: MRA as Part of a Multisequence Approach
Usually, MRA depicts the blood vessel lumen adequately and in many instances this
is sufficient for clinical management. However, in many instances, it is important
not only to detect the degree of vessel wall narrowing (usually by CEMRA) but
also to detect the effect that the stenosis is having on the organ being supplied
by the abnormal vessel. Although there is some correlation between the degree of
narrowing of the vessel and the impairment in flow to the organ beyond, because
of many physiological compensatory mechanisms, the effect is often not well
correlated with the degree of narrowing. As a result, many methods have been
proposed to measure flow into organs. For example, in the case of narrowing of the
arteries supplying the kidneys, flow can be measured by 2D and 3D PCA methods
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15.2 MR-Angiography 363
T
1
– no fatsat T
1
– fatsat
PD – no fatsat PD – fatsat
Figure 15.6 Dark blood imaging of the heart using
spin-echo T
1
and proton-density-weighted imaging without
and with fat saturation.
and also by measuring excretion of a small amount of the contrast agent injected
before the MRA (Figure 15.6).
In other instances, for example, in patients with suspected stroke, MRA is part of
a wider imaging workup. For example, the detection of the findings of infarction on
SE imaging (T
2
and FLAIR) and diffusion-weighted imaging triggers a search for
an abnormality within the arterial tree, the occlusion of which could be responsible
for the infarct. Under such circumstances, anatomical imaging (usually SE and
diffusion-weighted imaging) is typically combined with vascular imaging (e.g., TOF
MRA of the intracranial arteries) and possibly with perfusion imaging to detect
the degree of insult to the brain parenchyma. Although this has traditionally been
performed by susceptibility imaging following injection of the contrast agent, the
requirement to scan dynamically in the arterial phase precludes first-pass MRA of
the carotid and vertebral arteries. However, by adding techniques such as arterial
spin labeling (ASL) to detect cerebral blood flow (CBF), comprehensive assessment
of the arterial tree and the effect of arterial narrowing become possible during a
single examination.
15.2.2.6 Effect of Field Strength: 1.5 T vs. 3 T
As the difference between parallel and antiparallel protons is 10 per million at 3 T
compared to 5 per million at 1.5 T imaging at 3 T anticipates a doubling of SNR
over 1.5 T. However, as a host of additional interrelating factors, such as T
1
, T
2
,
and T
2
∗
, affect SNR in addition to field strength SNR increases of less than double
(∼1.70–1.8) can be anticipated at 3 T. CEMRA at 3 T imaging benefits not only from
the inherent improved SNR but also from the ability to use higher acceleration
factors and from the fact that T
1
values are longer at 3 T, so that for any dose
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364 15 Imaging Blood Circulation Using Nuclear Magnetic Resonance
Figure 15.7 Whole-body MRA using a multistation
stepping-table approach following a single injection of con-
trast agent.
of contrast agent the observed tissue contrast (difference between unenhanced
and contrast agent-enhanced images) is better at 3 T compared to 1.5 T. From a
practical viewpoint, the anticipated SNR boost at 3 T can be expected to improve
image quality, improve temporal, or spatial resolution using the same volume (and
injection rate) of contrast agent, or alternatively, a smaller volume of contrast agent
can be injected for a similar effect as achieved at 1.5 T.
15.2.3
Conclusion
Many methods have been validated in clinical practice to evaluate the vascular tree
noninvasively. Contrast-enhanced (Figure 15.7) techniques offer unrivalled spatial
and temporal resolution and speed of performance. In some clinical instances and
where it is inadvisable to use gadolinium-based contrast agents, there are many
noncontrast techniques that can be exploited to good effect.
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15.3 Perfusion Imaging 365
1.
2.
3.
4.
Labeled image
Label - Control = Perfusion
Arterial spin labeling
Control image
∝ CBF
Figure 15.8 Generalized schematical
drawing of an ASL areas. (1) Labeling
plane (labeling of inflowing blood by in-
version), (2) imaging plane (labeling), (3)
repeat experiment without labeling, (4)
imaging plane same position as in (2)
but without labeling, and (5) subtracting
results in an image with signal intensity
depending on CBF. (Picture is taken from
http://www.umich.edu/fmri/asl.html). (Source:
Images Courtesy of University of Michigon
(Educational Resources).)
15.3
Perfusion Imaging
Blood vessels and capillaries, which are beyond the resolution of normal MRI
sequences, can be visualized as the so-called perfusion images, where the signal
intensity in a voxel is related to the amount of flow and volume of those small
vessels. In this section, we discuss the most common approaches.
15.3.1
Time-of-Flight Perfusion Imaging
The TOF methods primarily optimized to measure perfusion are called arterial spin
labeling or arterial spin tagging (AST) techniques. These techniques can be divided
into two groups, which are referred to as continuous and pulsed ASL.
15.3.1.1 Continuous Arterial Spin Labeling (CASL)
In 1992, this technique based on an idea for MRA [21] was used for the first time
for imaging brain perfusion [22, 23]. In continuous arterial spin labeling (CASL),
an inversion slice or labeling plane is located proximal to the imaging plane. The
location is chosen in such a way that the slice cuts through big arteries. For a few
seconds, magnetization flowing through this plane is continuously inverted by the
principle of a flow-driven version of the AFP
6)
[21]. The inverted magnetization
leaves the labeling plane flowing into the imaging plane, where it can be detected
with MRI. A number of pulse sequences can be used as the imaging sequence,
6) The principle of AFPisexplainedinRef. [6].
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