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162
The actual arteriovenous connection may be too small to be seen
in postcatheterization AVF.
Evaluation of dialysis fistulae use specific criteria for the Doppler
spectra obtained from arterial inflow and venous outflow.
systolic velocity is recorded throughout the native system and the
graft. The arterial limb should demonstrate high velocities and con-
CH
tinuous forward flow with a low-resistance waveform. The venous
12
limb is expected to have slightly lower velocities. The normal PSV
at the anastomosis is 300 cm/sec. The normal outflow vein has a
PSV greater than 180 cm/sec and appears distended. Peak systolic
velocity less than 150 cm/sec indicates a fistula in jeopardy of failure. The fistula may result in arterial steal from the distal circulation. If this is suspected, direction of distal flow should be evaluated
before and after compression of the AVF.
Venous Duplex Ultrasound
45,46
Peak
FIGURE 12-30 Transverse gray-scale imaging of internal jugular vein
(IJV) and common carotid artery (CCA) without compression (left) and with
gentle compression (right) obliterating vein lumen.
Duplex examination of the extremity veins enables accurate noninvasive evaluation for deep vein thrombosis (DVT).
47
Normal
veins have thin walls and an echo-free lumen. The vein lumen
can be obliterated (compressed) with a small amount of extrinsic
pressure (
Fig. 12-29). The walls do not co-apt, however, when the
lumen contains thrombus, even when enough pressure is applied
to distort the shape of an adjacent artery. Vein compressibility is
best tested in an image plane transverse to the vein axis. Veins are
characterized by anatomical location as deep or superficial, and as
proximal or distal. The major veins of the thigh and arm are larger
in diameter than the corresponding arteries. Extremity veins have
valves that permit only cephalad flow, and these increase in number from proximal to distal. Valve sinuses are widened areas of the
lumen that accommodate the valve cusps.
Doppler evaluation of flow in normal veins has four important
characteristics: (1) respirophasic variation, (2) augmentation with
distal compression, (3) unidirectional flow toward the heart, (4) and
abrogation of flow in the lower extremities by the Valsalva maneuver. Complete analysis of venous spectral waveforms requires
comparison of the waveforms from both right and left limbs.
Presence of a flattened, unvarying waveform (loss of respirophasic
variation in flow) on one side compared with the other suggests
the presence of more proximal obstruction of venous return
proximal to the site of the Doppler interrogation.
Neck and Upper-Extremity Venous
Duplex Ultrasound
Neck and upper-extremity duplex evaluation includes assessments
of the internal jugular, subclavian, axillary, brachial, cephalic,
and basilic veins.
cava cannot be evaluated with duplex ultrasound because of
their location within the bony thorax. Examination begins with
evaluation of the internal jugular (Fig. 12-30) and subclavian veins.
The subclavian vein can be imaged from a supraclavicular or
48
The innominate veins and the superior vena
subclavicular approach. The arm is extended in a comfortable position for the evaluation of the axillary vein, paired brachial veins,
basilic vein (medial), and cephalic vein (lateral). Examination
includes color and spectral Doppler evaluation of flow in all these
veins. Loss of respirophasic variation in the waveform in the subclavian or axillary veins suggests the presence of more proximal
venous obstruction (due to thrombosis or extrinsic compression)
(
Fig. 12-31). The subclavian vein cannot be compressed where
it lies directly below the clavicle, and venous thrombosis is suspected when flow is absent or echogenic material is seen within
the lumen.
Loss of compressibility is the pathognomonic feature of venous
thrombosis. As the thrombus progresses from acute to chronic,
there is increased echogenicity of the thrombus and decreased
diameter of the vein. Over time, collateral veins may develop and
recanalization may occur in the thrombosed vessel. In the upper
arm, both superficial and deep venous systems have a significant
role in venous drainage. The majority of upper-extremity DVTs are
secondary to indwelling venous catheters, pacemaker leads, or
hypercoagulability. Primary upper-extremity DVT is a rare disorder
that is idiopathic, attributed to effort thrombosis (Paget-Schrötter's
syndrome), or related to thoracic outlet obstruction.
49
An unusual
etiology of noncompressible veins is intravascular tumor. This is
suspected when the echogenic material within the lumen appears
to extend through the vessel wall and may contain arterial flow
signals.
Lower-Extremity Venous Duplex Ultrasound
The venous ultrasound examination to evaluate the presence or
absence of leg DVT begins at the inguinal ligament with identification of the common femoral vein and extends to the calf.
The proximal deep veins evaluated include the common femoral, femoral (previously known as superficial femoral), and popliteal veins. The deep calf veins include posterior tibial, peroneal,
gastrocnemius (sural), and soleal veins. Special attention is given
50
FIGURE 12-29 Left, Transverse gray-scale imaging of superficial femoral
artery (SFA) (A) and vein (V). Right, With gentle compression, artery is unchanged
and vein is obliterated.
RT AX V
FIGURE 12-31 Spectral Doppler evaluation of axillary vein (AX)
demonstrating loss of phasic variation in flow with respiration. This finding
suggests more proximal venous obstruction. Increase in flow on right results
from compression of forearm (augmenting venous return).
163
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Gray scale
A
Spectral Doppler
LT SFV
Color Doppler
*
CH
12
VASCULAR LABORATORY TESTING
SFV
LT SFV
B
SFA
SFA
SFV
SFV
CD
FIGURE 12-32 Ultrasound evaluation of femoral vein thrombosis. A, Gray-scale imaging with echogenic material (asterisk) seen within lumen of superficial
femoral vein (SFV). B, Doppler in SFV shows a flow void within lumen. C, Spectral Doppler demonstrates diminished flow with respirophasic variation. D, Superficial
femoral artery (SFA) and SFV without (left) and with (right) compression; vein is only partially obliterated due to thrombus.
to the saphenofemoral and saphenopopliteal junctions because
thrombus in the superficial veins of these regions deserve more
aggressive treatment than thrombus limited to other parts of the
superficial venous system. Examination includes color and spectral Doppler evaluation of flow in all these veins
Loss of respirophasic variation in the waveform of the common
femoral vein suggests presence of obstruction proximal to the
site of Doppler interrogation that is preventing venous return.
Augmentation of flow with calf compression is not prevented
by proximal venous obstruction. Proximal obstruction may be
caused by extrinsic compression or venous thrombosis.
Bright-mode transverse images are used to determine compressibility along the entire course of the veins examined. Normally the
vein walls fully coapt with gentle pressure. Lack of compressibility,
which occurs because of a thrombosis in the vein, is the most reliable finding for determining venous thrombosis. With acute thrombosis, there is low echogenicity of the intraluminal thrombus, and
the vein is dilated. As the thrombus ages, it becomes more echogenic and less central within the lumen. Vein diameter decreases
as the thrombus retracts. Recanalization occurs, and flow can be
detected by pulsed or color Doppler. The thrombus often appears
eccentric and adjacent to the vein wall. When image quality is
poor because of significant soft-tissue edema, it is not possible to
exclude the presence of small nonocclusive thrombi. Sensitivity for
detection of common femoral vein thrombosis is 91%, and for both
the femoral and popliteal veins is 94%.
Duplex ultrasound is accurate for diagnosing deep calf vein
thrombosis in symptomatic patients, so long as the calf veins
51,52
(Fig. 12-32) .
FIGURE 12-33 Transverse gray-scale imaging of peroneal (deep) veins
of calf. Vein appears dilated in noncompression image on left. Vein lumen is not
obliterated by gentle compression, indicating presence of calf vein thrombosis.
A, artery; V, vein.
can be seen clearly (
Fig. 12-33). When compared with angiog-
raphy, sensitivity of compression ultrasound for deep calf vein
thrombosis is 94% and specificity is 100%. Small calf veins cannot be visualized well in all patients. However, specificity and
positive predictive value are high even when individuals with
poor calf vein images are included in the evaluation. Thus, the
diagnosis of DVT is made when calf veins are seen and cannot
be compressed.
Lower-extremity venous ultrasound testing is often ordered when
a patient is undergoing evaluation for pulmonary embolism (PE).
164
( ) ( )
This test will provide information about presence or absence of
venous thrombosis, but will not determine whether or not someone has had a PE.
Duplex Ultrasound Evaluation of Venous
CH
Insufficiency
12
Use of duplex ultrasound has been extended to detect reflux or
obstruction and determine the anatomical extent of venous disease in patients with chronic venous insufficiency.
ment has been facilitated by color Doppler imaging to provide
instant determination of the direction of blood flow. A 4- to 7-MHz
linear array transducer is used. The saphenofemoral junction is
examined first with the patient standing, and then in the supine
reverse-Trendelenburg position. Compressibility is determined in
transverse views of the veins. A longitudinal view of the saphenofemoral junction is then obtained.
One of two maneuvers can be used to elicit reflux. The first
is the Valsalva maneuver. Intraabdominal pressure increases
as the patient bears down, and venous outflow from the legs
decreases. Venous return from the legs increases with release of
the maneuver.54 The second is thigh cuff inflation and deflation.
Venous return is stopped with inflation of a cuff, typically to a
level approximating arterial diastolic pressure. There is a transient
increase in venous return that accompanies cuff deflation. Color
flow is evaluated before and after one of these two maneuvers to
elicit reflux. Baseline antegrade flow is displayed by blue color
Doppler. Red color after the maneuver indicates retrograde flow.
Reflux is present if red color persists for more than 0.5 seconds
after either maneuver. Spectral Doppler can also be used to evaluate reflux. The Doppler cursor is placed midstream with an angle
of 60 degrees with respect to the wall. Reverse flow over 0.5 seconds in duration is consistent with reflux (
remainder of the examination is performed with the patient standing with the weight on the leg not being examined.
The extent of reflux can be determined by repeating this
assessment throughout the deep and superficial veins of the
leg. For evaluation of the small saphenous and popliteal veins,
the patient sits on the edge of the examination table with his/
her foot resting on a stool. The probe is placed over the popliteal
fossa. The gastrocnemius veins can be seen between the popliteal (which is deep) and small saphenous (which is superficial)
vein. Compressibility and reflux following a Valsalva maneuver are
determined in these veins. The posterior tibial and peroneal veins
are assessed for reflux using the posteromedial and anterolateral
views.
53
This develop-
Fig. 12-34). Ideally, the
Perforating veins are vessels connecting superficial and deep
veins. Incompetent perforating veins are identified by sliding the
transducer up and down dilated superficial varicose veins.
55
Color
Doppler is then used while distal compression of the superficial
vein is performed. The presence of different colors during compression and release indicates that the direction of venous flow
changes with compression and relief. This finding is diagnostic of
reflux in the perforator veins.
Plethysmographic Evaluation
of Venous Reflux
Duplex ultrasound identifies reflux in individual veins, and
plethysmographic methods evaluate the volume of venous
reflux in the limb.56 Air or strain gauge plethysmography is a simple screening test that has the potential to provide a complete
analysis of venous hemodynamics (also see Chapter 55) . The air
chamber is filled with air to 6 mmHg and connected to a pressure
transducer and recorder. Changes in the volume of the leg as a
result of emptying or filling veins produce changes in the pressure
of the air chamber. Recordings are made with the patient supine,
and the leg elevated at a 45-degree angle. The patient then stands
with the leg flexed slightly and bearing weight on the nonstudy
leg. Venous filling time and venous volume are determined. The
time until the volume plateaus after the raised limb is dropped is
the venous filling time. Venous volumes of 80 to 150 mL are normal.
The venous filling index (VFI) correlates best with the clinical
severity of reflux.
VFI .90 venous volume .9 venous filling time÷=
Values less than 2 mL/sec indicate the absence of significant reflux,
whereas values over 10 mL/sec indicate high risks of edema, skin
changes, and ulceration.
57
Post-exercise plethysmography can be used to evaluate the ejecting capacity of the calf muscle pump. Venous volume is measured
at rest, and again post-exercise. Rest volume minus post-exercise
volume equals ejected volume. Ejection fraction is the ejected volume/rest volume × 100. Calf ejection fractions below 40% indicate
patients most likely to benefit from deep vein reconstruction.
Vascular Laboratory Accreditation
Laboratory accreditation is obtained through organizations such
as the Intersocietal Commission for the Accreditation of Vascular
Laboratories (www.icavl.org) and the American College of
Radiology (www.acr.org). The accreditation process reviews the
educational credentials of the interpreting physicians and sonographers, as well as laboratory procedures. It provides excellent standards for setting up examination protocols and quality assurance
programs.
RT LSV Back of prox calf
FIGURE 12-34 Spectral Doppler of venous insufficiency. Lesser saphenous
vein is imaged at saphenopopliteal junction. Spectral Doppler demonstrates
prolonged retrograde flow at this site.
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165
CH
12
VASCULAR LABORATORY TESTING
CHAPTER
13 Magnetic Resonance Imaging
Cihan Duran, Piotr S. Sobieszczyk, Frank J. Rybicki
Magnetic resonance angiography (MRA) is widely accepted for
the majority of vascular imaging applications and is the modality of
choice for many of them, particularly peripheral and renal arterio-
1–3
graphy.
high-quality spatial resolution images with relatively short acquisition times. Detailed comparison with other imaging technologies is
beyond the scope of this chapter, but general advantages and disadvantages appear in
applications such as aortography, where multiple modalities (e.g.,
computed tomography [CT] and MRA) are routinely diagnostic
and capable of providing images for planning most interventions.
Gadolinium-based contrast-enhanced (CE) MRA yields
Table 13-1. These become important for many
Basic Principles
Magnetic resonance imaging (MRI) relies upon the inherent magnetic properties of human tissue and the ability to use these properties to produce tissue contrast. Magnetic resonance imaging
detects the magnetic moment created by single protons in omnipresent hydrogen atoms. Because any moving electric charge produces a magnetic field, spinning protons produce small magnetic
fields and can be thought of as little magnets or “spins.” When a
patient is placed in the bore of a large magnet (i.e., MRI scanner),
hydrogen protons align with the externally applied static magnetic
field (B
most protons will distribute randomly, either with or against the
scanner's B
causing net tissue magnetization. The time required for this alignment is denoted by the longitudinal relaxation time, T1. T1 variations between tissues is used to provide contrast.
frequency of the wobble is proportional to the strength of B
radiofrequency (RF) pulse is applied at the resonance frequency
of the wobble, protons can absorb energy and jump to a higher
energy state. This RF pulse deflects the protons, creating a new net
magnetization vector distinct from the major axis of the applied
magnetic field. The net magnetization vector tips from the longitudinal to the transverse plane (transverse magnetization). The protons are “flipped” by the RF pulse, and the net magnetization vector
is defined by a “flip angle.” The stronger the RF pulse applied, the
greater the angle of deflection for the magnetization. Common flip
angles for spin echo are 90° and 180°. For gradient echo (GRE)
MRI, flip angles typically range between 10° and 70°. After the RF
pulse tips the spinning protons out of alignment with the main
magnetic field, new protons begin to align with the main magnetic
field at a rate determined by the T1 relaxation time.
states. The absorbed RF energy is retransmitted at the resonance
frequency and can be detected with RF antennas or “coils” placed
around the patient. These signals are compiled, and after mathematical processes become the MR images. Proton excitation
with an externally applied RF field is repeated at short intervals
to obtain signals. This MR parameter is referred to as repetition
time (TR). For conventional MRI, TR is typically 0.5 to 2 seconds,
whereas for MRA, TR ranges from 30 to less than 5 milliseconds.
When the spins are tipped to the transverse plane, they all precess
in phase. The speed of wobbling depends on the strength of the
magnetic field each proton experiences. Some protons spin faster
while others spin slower, and they quickly get out of phase relative
to one another. Throughout the dephasing process, the MR signal
decays. This loss of phase is termed T2 relaxation time or transverse
relaxation. T2, like T1, is unique among tissues and is used for image
contrast. In addition to the intrinsic T2 of tissue, inhomogeneity of
B
time that reflects the sum of these random defects with tissue T2 is
) to create a net magnetization vector. On a quantum level,
0
. However, a slight excess of spins aligns with the field,
0
Spinning protons wobble or “precess” about the axis of B
Energy is given off as the spins move from high to low energy
results in rapid loss of transverse magnetization. The relaxation
0
. The
0
. If a
0
called T2*. To obtain an MRI signal, these spins must be brought
back in phase and produce a signal or echo. The time at which it
happens is referred to as echo time (TE). In spin echo imaging technique, the echo is obtained by using a refocusing 180° RF pulse,
after which the spins begin to dephase. Another 180° RF pulse can
be applied to generate a second echo and so on. Signal loss at
longer echo times reflects tissue T2. In GRE imaging, the echo is
obtained by gradient reversal rather than RF pulse. Because this
includes effects from tissue homogeneity, TE-dependent signal loss
reflects T2*. Recently, GRE sequences (balanced GRE steady-state
free precession [SSFP]) have been developed that are insensitive
to magnet field inhomogeneities and reflective of actual tissue T2.
Longitudinal and transverse magnetizations occur simultaneously but are two different processes that reflect properties of various tissues in the body. Since T1 measures signal recovery, tissues
with short T1 are bright, whereas tissues with long T1 are dark.
Fat has a very short T1. In contrast, T2 is a measure of signal loss.
Therefore, tissues with short T2 are dark, and those with long T2
are bright. Simple fluids, such as cerebrospinal fluid and urine,
have long T2. To differentiate between the tissues based on these
relaxation times, MR images can be designed to be T1-weighted,
T2-weighted, or proton-density weighted. Exogenous contrast
such as gadolinium-based agents are routinely used to alter tissue conspicuity. Spatial encoding of signals obtained from tissues
is required for imaging. Additional external time-varying magnetic fields are applied to spatially encode the MR signal. Spatially
dependent gradients are used to locate the MR signal in space. In
two-dimensional (2D) MRI, these are slice-selection, frequencyencoding, and phase-encoding gradients. In three-dimensional
(3D) MRI, the slice-selection gradient is replaced by a second
phase-encoding gradient.
Magnetic resonance echoes are digitized and stored in “k-space”
composed of either two axes (for 2D imaging) or three axes (for 3D
imaging). K-space represents frequency data and is related to image
space by Fourier transformation. An important feature of k-space
is that tissue contrast is determined by the center of k-space (central phase encoding lines), whereas the periphery of the k-space
encodes the image detail. The order in which k-space lines are
collected can be varied, strongly influencing tissue contrast. For
example, in CE-MRA, the central contrast-defining portion of k-space
may be acquired early in the scan (centric acquisition) during peak
intraarterial contrast concentration to maximize arterial contrast. In
addition to simple line-by-line k-space acquisition schemes, more
complex schemes have been described. In spiral imaging, data
acquisition begins at the center of k-space and spirals to the periphery. Slice-selective gradients applied along the z-axis will form axial
images. Those along the y-axis will yield coronal images, and the
x-axis gradients will provide sagittal images. An oblique slice can be
selected by a combination of two or more gradients.
Magnetic Resonance Angiography
Techniques
Magnetic resonance imaging relies on selective imaging of moving
blood where signals from blood vessels are maximized, whereas
signals from the stationary tissues are suppressed. Algorithms then
enable reformatted images similar to those found in conventional
x-ray angiography (Table 13-2).
Magnetic resonance angiography methods can depict blood
as either black or white. For those “black-blood” methods that use
standard spin echo (SE) sequences (Fig. 13-1A), the excitation RF
pulse is applied at 90° and followed by a refocusing pulse at 180°. If
the imaging slice cuts across a vessel, then depending on the flow
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TABLE 13-1 Advantages and Disadvantages of Vascular Imaging Methods
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METHOD ADVANTAGES DISADVANTAGES
MR No ionizing radiation Expensive hardware
CE High signal from gadolinium-based agents
Non-contrast Eliminates toxicity concern Longer acquisition times compared to CE protocol
CT Rapid, high signal, high-quality image acquisition
DSA Intervention can be performed at time of diagnosis
Sonography
CE, contrast-enhanced; CNR, contrast-to-noise ratio; CT, computed tomography; DSA, digital subtraction angiography; MRA, magnetic resonance angiography; MRI, magnetic resonance imaging;
SNR, signal-to-noise ratio.
Less nephrotoxic than comparable iodine-based imaging (CT)
Less technical expertise required compared to MR
Highest spatial resolution
No ionizing radiation
Less expensive
Flow information readily obtained
Portable
Acquisition requires technical expertise
Nephrogenic systemic fibrosis can occur rarely in patients with
severe renal insufficiency
Increased risk of artifacts that are largely mitigated with CE-MRA
Ionizing radiation
Nephrotoxicity of iodine
Invasive
Nephrotoxicity of iodine
Projectional data may be inferior to volumetric acquisitions (CT, MR)
that can be viewed in any plane
Better for superficial imaging
Limited by artifacts from bone, air, and sonographic interfaces
Lower CNR, SNR compared to CT and MRI
Operator-dependent
TABLE 13-2 Types of Magnetic Resonance Angiography Sequences
NAME OF
SEQUENCE
TOF Bright vessels produced by inflow of
PC Bright vessels produced by applica-
Dynamic CE Bright vessels produced by rapid infusion
Postcontrast Bright vessels produced by equilibrium
Black blood
2D, two-dimensional; 3D, three-dimensional; CE, contrast-enhanced; PC, phase-contrast; RF, radiofrequency; TOF, time-of-flight.
blood with full magnetization into a
slice or volume where magnetization
has been reduced by RF saturation
tion of flow encoding radiofrequency
pulses that produce a phase image
where intensity is proportional to
velocity
of gadolinium contrast, with timing of
scan to arterial and/or venous transit;
mask subtraction may be used
enhancement of gadolinium contrast
with fat saturation
Dark vessels produced by use of an
inversion prepulse to null the signal of
blood, based on its T1 recovery time
DESCRIPTION
2D OR 3D GATING USEFUL
Both Occasionally No Intracranial, carotid, pedal
Both Occasionally Yes None in current practice
3D No No Carotid, chest, abdomen,
Both No No Veins, aorta
2D
Necessary
FLOW QUANTIFICATION
POSSIBLE
No
APPLICATIONS
extremity runoff
Plaque imaging
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velocity and time interval between the pulses, the blood volume
originally excited by the first pulse may not “see” the second pulse.
This results in a black-appearing signal void in the vessel lumen.
The use of thin sections or long echo times can further emphasize
this flow void. This technique allows detailed examination of arterial wall morphology. Fast spin echo (FSE) sequences, in which a
long train of echoes is obtained by use of repeated 180° pulses, produce images more rapidly. The double inversion recovery (DIR)
FSE offers a new approach to enhance black-blood sequences. The
technique uses two consecutive inversion pulses: the first nulls or
blackens the blood everywhere in the coil, and the second restores
magnetization in the slice being imaged. Between these pulses
and image production, blood within the slice is replaced by nulled
blood from outside. This produces more reliable black blood than
conventional approaches, making this sequence ideal for examining wall thickness, dissection flaps, and the presence of mural
thrombus or inflammation.
4
This provides a clear advantage over
traditional x-ray angiography.
“Bright blood” MRA techniques use GRE sequences and are generally divided into those measuring signal amplitude (time-of-flight
[TOF]) and those based on phase effects (phase-contrast [PC]). In
each GRE sequence, a single RF pulse is applied in short time intervals, eliminating signal loss due to flow void. The stationary protons
occupying a given tissue slice do not have sufficient time to relax
to their equilibrium state.
TOF-MRA techniques depend on the inflow of unsaturated
protons in blood from outside the field of view (FOV) into the
stationary tissue within a section already saturated by its exposure to repeated RF pulses. These “saturated” protons are unable
to contribute signal to the image. The signals in the stationary tissues of GRE images used in MRA are therefore typically low. The
“unsaturated” protons in blood flowing into the imaging plane
have not experienced the RF pulses and yield maximum signal.
The unsaturated blood appears bright compared to background
tissue (
Fig. 13-1B). The time required for blood to flow through an
image slice and its effect on the resulting signal is known as TOF.
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A
B
FIGURE 131 “Black-blood” and “bright-blood” imaging techniques. A, Cross-sectional T1-weighted image of ascending and descending aorta; lumen appears
black. B, Time-of-flight (TOF) image of carotid-vertebral system. C, Maximum-intensity projection (MIP) reconstruction of arterial phase of contrast-enhanced magnetic
resonance angiography (CE-MRA) images in same patient shows normal carotid-vertebral arteries at higher spatial resolution than corresponding TOF acquisition.
Saturation of signals can occur in vessels with slow-moving blood
as a result of repeated RF excitation in the acquisition plane. This
can create artifacts in vessels with stenotic lesions or reduced
blood flow.
Time-of-flight techniques can be obtained in 2D or 3D. The 2D
TOF utilizes multiple sequentially acquired, overlapping thin slices
to form an image. The patient is instructed to hold his/her breath
to minimize motion artifact. However, spatial misregistration may
occur if patients cannot hold their breath at the same level each
time. Thus, only one or two slices are typically acquired per breathhold. 2D TOF has good sensitivity for identifying vessels with slow
flow because blood must move only 3 to 5 mm to refresh a slice.
3D TOF consists of GRE acquisition of a volume into which blood
is flowing. The advantage of this technique is higher signal-to-noise
ratio (SNR) and improved resolution. The thick volumes of tissue
imaged require rapid flow to fully refresh signals within the arteries. The technique is flow dependent and superior in vessels with
rapid steady flow without respiratory motion. Additional saturation
pulses can be applied to eliminate signal from veins. Segmented
GRE sequence with cardiac triggering can be used to eliminate
arterial pulsation artifacts.
A successful TOF image requires the section to be thin enough
to allow for sufficient inflow between RF pulse repetitions, but
thick enough to ensure adequate SNR and anatomical coverage.
Section thickness of 3 to 4 mm is used for large vessels, and 1 to
2 mm for smaller vessels. Spatial presaturation pulses are applied
above or below the imaged slice or volume to eliminate unwanted
signal from arteries or veins, depending on which part of the vascular tree is being imaged. Optimal TR for TOF is 20 to 50 milliseconds. Short TR keeps background tissues saturated, but it must be
C
long enough to allow for satisfactory inflow of unsaturated blood
between successive repetitions. The best flip angle is usually 30° to
60°. With phasic flow in the extremities, systolic flow signal may be
increased (because of greater transverse magnetization created),
and distal flow may be decreased, creating view-to-view intensity
changes and phase artifacts from pulsatile variations. This pulsation artifact is greatest at higher flip angles. Cardiac gating can be
used to minimize these artifacts at the expense of increased imaging time.
Whereas TOF uses differences in signal amplitude to differentiate between stationary and flowing spins, the PC technique
observes the phase shifts of signals. Moving spins experience different phase shifts in the presence of the applied magnetic fields
used in MRA. Strength and orientation of the applied magnetic
field are varied to encode different phase shifts for flowing protons relative to stationary protons. The faster the spins are moving,
the greater their phase shift, and protons of flowing blood may
be discriminated from stationary protons. The phase shifts result
in a contrast between moving and stationary tissues and form
the basis for PC imaging. Pairs of images are acquired that have
different sensitivities to flow and are then subtracted to cancel
background signal, leaving only the signal from flowing blood.
Phase shift is proportional to velocity, allowing flow quantification
with this modality. Phase-contrast acquisitions may be acquired
in two or three dimensions; although used rarely in angiography
today, Phase-contrast offers a reliable way to quantify amount and
direction of flow. It requires long imaging times: two data sets in
each direction are acquired by using flow-encoding gradients of
opposite polarity, and up to three measurements in the orthogonal
planes are needed to image flow in all directions.
Visualization of the arterial system with PC and TOF is adequate5
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but has limitations. Acquisition times can be long and prevent
imaging within the time span of a single breath-hold. This increases
the chance of movement artifacts. Some of the limitations are
caused by flow-related artifacts such as in-plane saturation and
phase dispersion. Flow-based imaging also has limits in areas
of slow flow, such as aneurysms. Overgrading of stenotic lesions
is most commonly a manifestation of signal loss in the areas of
complex flow. Undergrading is a matter of inadequate spatial resolution. Complex turbulent flow patterns in areas of stenoses can
create signal loss and mimic a critical lesion. This is due to “intravoxel dephasing.” An accelerated flow across a stenosis consists of
a wide distribution of velocities and thus a large distribution of proton phases. In the smallest volume element, a “voxel” of the image,
this distribution of phases can result in cancellation rather than
coherent addition of signals, accounting for the presence of signal
voids at the site of stenosis. A short TE minimizes flow-phase dispersion artifacts. Phase dispersion is further decreased when voxel
size is minimized using thin sections. Small voxels and short TE are
most easily obtained with 3D TOF methods. The biggest drawback
of the thick volumes used with 3D techniques is that slow or recirculating flow can become saturated. The MOTSA (multiple overlapping thin-slab acquisitions) technique of sequential 3D TOF gives
better flow enhancement than single-slab 3D TOF techniques and
less dephasing than 2D techniques. However, the need for substantial overlap of adjacent slabs increases acquisition time.
Contrast-Enhanced Magnetic Resonance
Angiography
The introduction of CE-MRA has revolutionized MRA.
nique overcomes many of the limitations of traditional bright
blood modalities: respiratory motion artifacts, poor SNR, and
flow and saturation-related artifacts (Fig. 13-1C). Gadolinium
increases the signal intensity of blood on contrast-enhanced 3D
T1-weighted (spoiled) GRE images. Blood contrast is not flow
dependent. It is determined by the concentration of contrast
agent within the arterial system while imaging data are being collected. Reliable images can be acquired irrespective of whether
flow is laminar, turbulent, or stagnant. This technique acquires
large-volume data sets in coronal or sagittal orientation within
a single breath-hold during the first pass of the contrast material.
The contrast agent, gadolinium, is a heavy metal but becomes
inert when bound to a chelator. Intravenous (IV) administration of
gadolinium-diethylenetriamine pentaacetic acid (DTPA) results
in a marked reduction of the T1 or longitudinal relaxation time
of blood, therefore reducing the effects of spin saturation. Signal
reduction is also problematic in 3D TOF sequences. Moreover,
the very short TE reduces spin dephasing and allows accurate
evaluation of vascular stenoses.
Multiple refinements have resulted in a technique that is much
faster than TOF-MRA. The development of high-performance
gradient systems with ultra-short repetition TR and TE has shortened acquisition time in CE-MRA to allow imaging within a single
breath-hold and minimize motion artifacts. Administration of agents
shortening T1 allows selective visualization of contrast- containing
structures and better visualization of circuitous collaterals. Digital
subtraction, spoiling, and fat saturation techniques suppress background signal and enhance signal from the contrast agent in the
vessels. The subtracted data sets can be postprocessed to provide
3D projectional images. CE-MRA still provides a luminogram, and
conventional or FSE images are needed for a complete study so that
true lumen diameter and presence of thrombus can be established.
Optimal images are generated when gadolinium concentration
is highest in the vessel of interest. To make blood bright compared
to background tissues, the gadolinium bolus must be administered in a way that ensures the majority of the contrast to be present in the arterial tree. This requires exact timing of the arrival of
the gadolinium bolus. Acquisition prior to contrast arrival creates
6,7
This tech-
a “ringing” artifact, whereas late acquisition creates venous and
tissue enhancement, contaminating the arterial signal. This is especially problematic in MRA of the extremities, where the images
are obtained in multiple segments. Contrast transit time can be
affected by low cardiac output, valvular regurgitation, large abdominal aneurysms, and flow-limiting stenoses. Proper timing can be
achieved by empirical estimation of transit time or a test bolus in
the anatomical field of interest. Alternatively, with automated triggering, a pulse sequence can be designed to sense the arrival of
contrast and automatically trigger image acquisition. Magnetic
resonance fluoroscopy allows the user to visualize arrival of the
contrast bolus directly on the image and manually trigger the start
of the scan. Areas that require higher spatial resolution, such as
the lower extremities, also need larger doses of contrast for longer
acquisition times.
Imaging during the arterial phase of gadolinium infusion takes
advantage of higher arterial SNR and eliminates overlapping
venous enhancement. This is a brief moment in time, but several methods allow slower MR image acquisition to capture that
moment. The phase reordering (mapping k-space) technique
acquires central k-space data (i.e., the low spatial frequency data)
when contrast concentration is high in arteries but lower in veins.
This allows a relatively long MR acquisition to achieve the image
contrast associated with the shorter arterial phase of the contrast
bolus. It is critical to time the contrast bolus to achieve maximum
arterial gadolinium concentration during acquisition of central
k-space data.
CE-MRA is limited by venous and soft-tissue enhancement.
Contrast media not only passes into venous structures, dependent on the arteriovenous transit time of the tissue, but also rapidly leaks out of the vascular compartment, creating significant
tissue enhancement. New “blood pool” agents, which are currently
undergoing clinical trials, are retained within blood vessels and
selectively enhance the blood pool on T1-weighted MR images.
These use either gadolinium compounds that bind to albumin, or
are large enough to stay within the vascular space or ultra-small
iron particle. Another agent, gadobenate, has a higher T1 relaxation
time because of its capacity for weak and transient interaction
with serum albumin. This may enhance vascular signal intensity
and thus increase diagnostic efficacy at doses comparable to those
used for current gadolinium agents. It is approved for imaging use
in Europe but is under clinical investigation in the United States. It
provides a higher and longer-lasting vascular signal enhancement
in the abdominal aorta compared with gadolinium, which does
not interact with proteins.
Gadolinium-based contrast has a very favorable safety profile. However, gadolinium is nephrotoxic. For patients with underlying chronic renal insufficiency, gadolinium chelates can cause
acute renal failure. Nephrogenic systemic fibrosis (NSF) is linked
to gadolinium-based contrast agents9 and largely involves the
skin, though it may also affect the muscle, joints, or internal organs
such as the lungs, liver, and heart in patients with renal failure.
Nephrogenic systemic fibrosis occurs in patients with severe renal
disease who are exposed to high doses of gadolinium agents, or
in patients who receive multiple standard doses of contrast agents
in a short period of time. The reported prevalence of NSF among
patients with glomerular filtration rate (GFR) less than 30 mL/min
is 3% to 5%.10 Therefore, the MR protocol should aim to minimize
contrast volume, especially in patients with moderate to severe
renal failure.
Metal objects, such as surgical clips, lead to susceptibility
artifacts in MRA. The increasing use of stents has important
implications for MRA. Cavagna et al. evaluated CE-MRA of seven
stent types in the aortic, iliac, and popliteal positions.12 Few of
the commonly used stents permitted visualization of the lumen.
Susceptibility artifact results in significant signal loss that can
preclude proper visualization of the stent lumen even with
gadolinium-enhanced MRA. Some nitinol, tantalum, or polytetrafluoroethylene (PTFE)-based devices, on the other hand, cause
less artifact on CE-MRA.
11
8
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170
Postprocessing Techniques
Magnetic resonance data can be viewed as source images or be
displayed in projections with any orientation. Image postprocessing allows reformation in any desired plane to improve conspicuity of overlapping vessels (
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common carotid artery (CCA) and left subclavian artery, for exam-
13
ple, can overlap in coronal projections, whereas the origins of the
right subclavian artery and right CCA can overlap in some oblique
views. The renal ostia are usually best seen in either coronal or
slightly oblique view. The celiac axis and superior mesenteric
artery (SMA) are best depicted on sagittal projections. One advantage of MR versus digital subtraction angiography (DSA) is that the
latter may require multiple injections to assess the origins of these
vessels. The details of image interpretation are beyond the scope
of this review, but source image data, multiplanar reconstruction
(MPR), maximum-intensity projection (MIP), and volume rendering (VR) are used (
Fig. 13-2). Source images are the initial recon-
structions and should be used for problem solving and to confirm
Table 13-3). The origins of the left
TABLE 13-3 Types of Postprocessing Techniques
TECHNIQUE DESCRIPTION
MPR Production of cross-sectional images in planes
MIP projection Production of full- or partial-volume images
Volume rendering
MIP, maximum intensity projection; MPR, multiplanar reconstruction; MRI, magnetic
resonance imaging.
different from acquisition plane
along any desired axis from a stack of image
slices
Manipulation of MRI slices to produce full
volumetric images; structures segmented for
viewing by application of intensity thresholds
and removal of unwanted structures
findings. Interpretation often begins with a vascular survey using
MPR and MIP data sets. Multiplanar reconstructions are very useful in volumetric acquisitions because the desired imaging plane
B
A
C
FIGURE 132 Postprocessing techniques; contrast-enhanced magnetic resonance angiography (CE-MRA) of abdominal aorta and branches. A, Coronal
thin-section image of abdominal aorta. Summation of these images, projected with maximum intensity, is used for (B) coronal maximum-intensity projection (MIP)
that includes normal renal and mesenteric arteries. C, Axial multiplanar reconstruction (MPR) image at level of left renal ostium does not include entire extent of both
renal arteries, giving false impression of a proximal right renal artery occlusion (arrow). D, Three-dimensional volume-rendered (3D-VR) image shows entire course of
abdominal aorta and its branches.
D
can be prescribed to enhance vascular separations. Subtracted
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MIPs are routinely created from CE-MRA. The noncontrast (mask)
images are subtracted from the enhanced images, and resulting
high SNR data sets undergo projection of maximum intensity. By
performing the projections of all angles around the z
-axis of the
patient, the data sets can be viewed in cine. These projections are
referred to as rotating MIPs.
Clinical Applications
Extracranial Carotid and Vertebral Arteries
Atherosclerosis, dissection, and inflammatory diseases affect the
extracranial carotid and vertebral arteries. Time-of-flight methods (2D and 3D) have been largely replaced by CE-MRA because
CE-MRA allows imaging of the entire course of these vessels
(Fig. 13-3). A problem with TOF imaging is turbulent flow at or near
the carotid bifurcation where most of the lesions occur; this may
lead to overestimation of lesion severity. Gadolinium-enhanced
MRA has greater SNR than noncontrast imaging, is less sensitive
to intravoxel dephasing from turbulence, and does not have signal
loss from saturation effects. Phase-contrast sequences can supplement anatomical data for flow direction or quantification.
Transcranial flow has a rapid arteriovenous transit time, and
thus venous contamination can limit image quality; therefore, timeresolved imaging may be required. Assessing the test characteristics
of carotid MRA is challenging because technologies evolve, and
the patients and methods are both heterogeneous. In a 41-study
meta-analysis, carotid CE-MRA had high sensitivity (94%) and high
specificity (93%) for diagnosis of severe (70%-99 %) carotid artery
stenosis.13 The sensitivity and specificity of CE-MRA for ostial stenosis also is very high.
The intimal flap of CCA dissection can be visualized with either
CT or MRA. For a patient with poor or uncertain hemodynamic stability, CT is the preferred modality because it is more rapid and has
better patient monitoring capabilities. However, MR is preferred for
stable patients, since it does not impart ionizing radiation to the
thyroid gland for either initial or follow-up studies.
For vertebral artery dissection, MRA can detect the level of stenosis or obstruction and distinguish residual flow from intramural
hematoma.
14
FIGURE 134 Maximum-intensity projection (MIP) image shows normal
thoracic aorta and origin and course of supraaortic vessels.
Thoracic Aorta and Its Branches
Both CT and MRI provide comprehensive imaging of the aorta.
Single breath-hold 3D CE-MRA imaging of the thoracic aorta
(Fig. 13-4) is typically performed with electrocardiographic (ECG)
gating to eliminate pulsation artifact.
AORTIC DISSECTION
A properly performed modern MRA is 100% sensitive for dissection
and intramural hematoma.
and extent of the intimal tear, and the relationship of the tear to the
branch vessels are readily tracked in multiple planes. Cine images
of the proximal aorta can identify aortic regurgitation complicating
type A dissection (
Fig. 13-5). Delayed phase images allow identifica-
tion of intramural hematoma, ulceration, and complications including rupture. On T1-weighted SE sequences, intramural hematoma is
seen as a concentric thickening of the wall, with increased intramural signal intensity.
16
Inflammatory changes are seen as arterial wall
thickening and enhancement. When used with clinical parameters
and other testing (e.g., blood pressure changes), MRA is useful to triage patients into either medical or surgical management options.
15
The true and false lumens, location
17
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THORACIC AORTIC ANEURYSM
Aneurysm imaging should consider slow blood flow through the
lesion. Magnetic resonance angiography can demonstrate the
location and size of an aneurysm, presence of a mural thrombus, and the relationship of the aneurysm to the branch vessels.
Because MIP images from 3D CE-MRA are designed to highlight
the lumen of the aorta, the source (including noncontrast) images
should be evaluated to determine the extent of mural thrombus
and the actual aneurysm size for accurate measurements. Timeresolved imaging shows delayed enhancement from slower flow.18
Patients with a clinically suspected dissection and contraindication to gadolinium (e.g., severe allergy, acute renal failure) can
undergo DIR imaging or often noncontrast techniques to evaluate
the thoracic aorta.
19
ARCH VESSEL DISEASE
Occlusive disease of the great vessels is usually due to atheroscle-
FIGURE 133 Carotid artery disease. Left internal carotid artery (ICA)
dissection resulted in a thrombotic occlusion of proximal vessel. Internal carotid
artery reconstitutes more distally.
rosis. Vasculitis, fibromuscular dysplasia (FMD), and radiation arteriopathy also can cause branch vessel stenoses. CE-MRA is an
established tool for rapid and accurate definition of brachiocephalic
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