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Review |
1 Both authors: Department of Radiology, University of Michigan Medical Center, 1500 E Medical Center Dr., TC-2910A, Ann Arbor, MI 48109-0326.
Received May 5, 2003;
accepted after revision October 21, 2003.
Address correspondence to B. Desjardins.
Introduction
 Top Introduction Historical BackgroundMDCT: General PrinciplesCardiac GatingTechnical ConsiderationsClinical IndicationsCoronary Artery Calcium ImagingCoronary Artery Lumen ImagingCardiac FunctionOther IndicationsConclusionReferences |
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The definitive diagnosis of coronary artery disease is made by catheter coronary angiography, an invasive and expensive procedure with associated morbidity and mortality [2]. Therefore, catheter coronary angiography is often used as second-line diagnostic test after noninvasive diagnostic procedures, such as a treadmill exercise tolerance test, that provide indirect evidence of coronary artery stenosis [3]. In the past few years, progress has been made in the development of noninvasive diagnostic alternatives to directly image the coronary arteries. The two most promising noninvasive approaches are CT and MR coronary angiography.
This review describes the state of the art in cardiac imaging with CT, with emphasis on coronary artery imaging. After a more detailed description of the two principal CT technologies used in cardiac imaging—electron beam CT and MDCT—the technical aspects of ECG-gated cardiac MDCT are described. Elements relevant to the day-to-day clinical practice of cardiac imaging with CT are emphasized, with focus on how and why things are done, the clinical uses of this modern technology, its current limitations, and a comparison with other imaging techniques such as catheter angiography.
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TopIntroductionHistorical BackgroundMDCT: General PrinciplesCardiac GatingTechnical ConsiderationsClinical IndicationsCoronary Artery Calcium ImagingCoronary Artery Lumen ImagingCardiac FunctionOther IndicationsConclusionReferences |
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CT technology has made remarkable progress during the past three decades, with successive generations of scanners providing faster imaging at progressively higher spatial resolution. Following is a short overview on the history of the developments in CT technology as it relates to cardiac imaging.
Axial CT
In 1972, the first generation of CT scanners was introduced
[6]. This invention earned
Godfrey Hounsfield and Allan Cormack the Nobel Prize for Medicine in 1979
[7]. Such scanners were not
very efficient by modern standards. They used a pencil-beam X-ray source. One
or two detectors acquired data while the source was translated linearly across
the field of view. The whole system was then rotated by steps of 1° for
each source and detector translation, and this was repeated to span 180°.
Approximately 5 min was needed to acquire one or two tomographic images (the
earliest CT scanners were already dual-slice scanners). This approach was
rapidly improved by using a narrow fan-beam X-ray source and more detectors
per row (second generation)
[8], and then by using a wide
fan-beam X-ray source with rotation of both the source and the detectors
(third generation) [8]. These
latter axial CT scanners could acquire a tomographic image in less than 5 sec.
Cardiac imaging by CT was first reported in 1981
[9] on a third-generation CT
scanner with a 2-sec gantry rotation time, providing an effective temporal
resolution of 500–1,000 msec, still insufficient to achieve motion-free
images of the heart.
Fast Axial CT
In the late 1970s, a stationary detector ring was introduced to remove ring
artifacts in the axial CT images (fourth generation)
[8]. In 1982, both the source
and the detector were made stationary on electron beam CT scanners
[8], which were specifically
created for cardiac imaging and could acquire a tomographic image in 100 msec.
For the first time, electron beam CT provided temporal resolution sufficient
for motion-free imaging of the heart during diastole
[10].
Electron beam CT scanners differ from axial CT scanners. In electron beam CT, electrons are accelerated in a vacuum funnel and are precisely focused toward and swept across a 210° tungsten ring anode placed under the patient. After hitting the target ring, a cone beam of X-ray photons is emitted. The X-ray photons go through the patient and are captured by two 240° detector rows above the patient. Slice collimation is 3 mm, so 40 slices are needed to cover the entire heart (12 cm), for a total imaging time of 30 sec. Imaging can usually be performed in one breath-hold. ECG-based triggering is used for motion-free imaging during diastole. However, significant motion artifacts still remain in approximately 20% of patients [11]. Although in-plane spatial resolution is submillimeter (0.8 x 0.8 mm), the resolution along the longitudinal z-axis remains limited (3 mm) in order to image the entire heart in a single breath-hold.
Today, electron beam CT is predominantly used for the detection of calcium load in the coronary arteries, a prognostic factor in patients with coronary artery disease [12, 13]. It is also used to a lesser extent to perform CT coronary angiography [14] and to evaluate cardiac function [15]. The spectrum of applications for electron beam CT is limited, and its physical setup is cumbersome. It appears to have lost its competitive edge over the latest all-purpose MDCT scanners, although current technologic improvement in electron beam CT technique, such as the integration of 4-slice technology, thinner (1.5-mm) collimation, 30- to 50-msec temporal resolution, and improved user interfaces could reposition electron beam CT in the future.
Single-Slice Helical CT
Slip-ring technology was first introduced in CT scanners in 1989
[8] and enabled the gantry to
rotate continuously without fixed wires, offering continuous helical scanning
of entire volumes in less than 30 sec, with a 1-sec gantry rotation time and
subsecond image acquisition
[16]. Scanning and image
positions were decoupled, allowing images to be reconstructed at any arbitrary
position along the longitudinal axis. Z-axis resolution was also
improved, enabling improved multiplanar and 3D image reconstruction. General
image quality for a given slice thickness was not significantly different from
that of axial CT, and the radiation dose was also similar, given appropriate
pitch values [17].
Single-slice helical CT could be used for cardiac imaging with ECG-synchronized protocols [18]. Temporal resolution was improved over nonhelical scanners but, at 250–500 msec, remained insufficient for motion-free cardiac imaging. Almost all patients had motion artifacts [18]. In-plane spatial resolution was 0.6 x 0.6 mm. However, to cover the entire volume of the heart in a single breath hold, 3-mm collimation was required, severely limiting the z-axis resolution and the ability to evaluate the small coronary arteries.
Fan-Beam MDCT
CT scanners have been introduced with increasing numbers of detector rows,
referred to as MDCT or multislice CT. In 1993, CT units were introduced with
two rows of detectors and 1-sec rotation time
[19]. In 1994, units with a
0.75-sec gantry rotation time were introduced
[20], which improved volume
coverage and z-axis resolution. In 1998, units with four rows of
detectors were introduced, with 0.5-sec rotation time, in-plane resolution of
0.6 x 0.6 mm, and 1.0- to 1.25-mm collimation
[21]. These units allowed an
eightfold increase in performance compared with single-slice helical CT (four
times the rows, twice the rotation speed). This gain in performance could be
used for faster coverage of the heart by increasing the table feed per
rotation, for covering a greater volume, or for improved temporal resolution
to 125–250 msec. Thinner collimation (1-mm detector) could be used to
improve z-axis resolution. General image quality was better than with
single-slice CT and remained excellent at higher scanning speeds, thereby
reducing motion artifacts.
The first MDCT scanners provided true volume image data and have enabled motion-free cardiac imaging in diastole for heart rates up to 70 beats per minute in a single 20- to 40-sec breath-hold. These scanners provided the first viable alternative to electron beam CT. These systems have superior spatial resolution to electron beam CT (0.6 x 0.6 x 1.0 mm vs 0.8 x 0.8 x 3 mm) in a single breath-hold, but inferior temporal resolution (125–250 msec vs 100 msec). ECG gating is used to improve temporal resolution, but motion artifacts are still present in approximately 50% of patients [22]. Radiation dose issues became more important, with doses up to 10 mSv for cardiac MDCT versus 1.1 mSv for electron beam CT [23], and systems were subsequently designed with variable tube output to decrease total radiation exposure during the CT examination. In such systems, tube current is modulated in two ways: by projection angle, with tube current being much lower in frontal projections than lateral projections because total attenuation (thickness of the chest) varies according to projection angle, and by ECG, with tube current being lower during systole because most of the diagnostic information comes from diastolic images, which present fewer motion artifacts.
Cone-Beam MDCT
In 2002, the latest generation of MDCT scanners was introduced, with 16
rows of detectors and even thinner collimation (0.625–0.75 mm),
providing for the first time isotropic resolution of 0.5 x 0.5 x
0.6 mm and gantry rotation times of 400–500 msec. Isotropic resolution
means similar resolution along all axes. Arrays of detectors with variable
geometry (isotropic arrays, adaptive arrays) are used. For example, this can
be 24 rows of detectors along the z-axis, with 16 central rows of
0.625-mm collimation flanked on each side by four rows of 1.25-mm collimation,
and 888 columns of detectors along the plane of the gantry (Fig.
1A,
1B). The X-ray tubes are more
powerful, with increased tube current, to provide faster scanning with
improved effective slice thickness.
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These systems provide motion-free cardiac imaging in diastole for heart rates up to 80 beats per minute, in a 20-sec breath-hold. With segmented reconstruction algorithms (discussed later in this article), a temporal resolution of up to 65 msec can be achieved, and motion-free imaging can be performed for heart rates up to 100 beats per minute. These MDCT scanners can therefore offer both higher spatial resolution and higher temporal resolution than electron beam CT and use a small form factor, and they are multipurpose clinical units. Radiation exposure is relatively high, ranging from 5 to 10 mSv for coronary CT angiography [23]. This exposure can be reduced using tube current ECG-synchronized techniques [24]. Cone-beam artifacts are present on MDCT with more than four rows of detectors, requiring complex cone-beam algorithms for image reconstruction [25–27].
Volume CT
To further improve the current generation of MDCT scanners requires
increasing both temporal and spatial resolution. Increased temporal resolution
can be achieved by faster gantry rotation. This is an engineering challenge
because the gantry experiences centrifugal forces of 10 G
(gravitational constant) on 0.5-sec rotations
[21]. A realistic limit for
rotation times appears to be approximately 50–200 msec (Kalendar WA,
presented at the 2002 meeting of the European Congress of Radiology). Temporal
resolution can also be improved using better and more complex segmentation
algorithms; however, rotation time must be variable to desynchronize rotation
and cardiac cycle (discussed later in this article), limiting the
applicability of segmented algorithms
[28].
Higher isotropic spatial resolution requires more numerous and smaller detectors, with thinner, submillimeter slice collimation. Large-area detectors (40 x 30 cm) or volume CT may replace detector arrays and are currently under development [29]. A pig heart imaged on a prototype device (General Electric Medical Systems) with a flat-panel detector made of a grid of 1024 x 1024 cells each having a resolution of 200 µm in each axis [30] shows increased anatomic detail (Fig. 2). Early results show that even the fifth-degree coronary artery branches can be visualized with such a system [31]. However, higher isotropic spatial resolution comes at a cost. As detector dimension decreases by N, contrast resolution decreases and noise increases by N2, which is compensated for by an increase in radiation dose of N4 [21]. Thus, noise and dose factors may limit future increases in spatial resolution.
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Effective slice thickness using a single-slice detector scanner is determined by collimation, whereas in multidetector scanners it is determined by both the detector configuration and the reconstruction method. The table travel per complete rotation of the gantry divided by the X-ray beam width is called the "beam pitch." The beam pitch conforms to the current industry standard of pitch and is used in this article. In multidetector scanners, another definition of pitch is sometimes used: the table travel per complete rotation of the gantry divided by the detector width, called the "detector pitch." With a (beam) pitch of 0, there is no table motion and scanning is axial. With a pitch of 1, the table displacement for each rotation is equal to the z-axis dimension of the array of active rows of detectors. A pitch greater than 1 implies gaps in the helix of projections, whereas a pitch between 0 and 1 implies overlap between the projections (Fig. 4A, 4B, 4C).
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Images are reconstructed from linear interpolation of projection data from rays that are the closest to the image plane (Fig. 3B), using algorithms to correct for cone-beam geometry. A minimum of 180° of projection data are mathematically required to reconstruct a complete image. Thus, slice thickness increases with pitch, which determines how much the table travels during a 180° rotation. The thinnest slice is equal to the height of a single row of detectors (the detectors are actually wider than the slice to compensate for geometric magnification). Thicker slices can be generated mathematically by combining the thinner slices. The determinants of in-plane resolution are focal spot size, detector width and geometry, reconstruction algorithm, and image matrix size. The determinants of longitudinal z-axis resolution are focal spot size, detector height and collimation, and reconstruction algorithm. The determinants of temporal resolution are gantry rotation speed and reconstruction algorithm.
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The least cardiac motion occurs during diastole, when the ventricles are passively filling. Prospective ECG triggering uses the ECG signal to control scanning, so that X rays are generated and projection data are acquired only during cardiac diastole, more than half the rotation of the gantry. The total number of slices produced per heartbeat during this half rotation of the gantry is proportional to the number of rows of active detectors. Because an axial scanning method (rather than helical) is typically used and the table has to move by the total collimation width after each acquisition, one heartbeat typically has to be skipped between each acquisition. About 12 cm of scanning is required to cover most heart sizes, which requires approximately 48 heartbeats for single-slice CT (5-mm collimation), 24 heartbeats for 4-slice MDCT (2.5-mm collimation each row), and 12 heartbeats for 16-slice MDCT (1.25-mm collimation each row). Thus, multidetector technology can obtain the entire scan during one breath-hold. Electron beam CT requires about 24 heartbeats and can scan the entire heart in one breath-hold. The start of the diastolic phase of the cardiac cycle is estimated from the prior three to seven consecutive heartbeats and occurs approximately 450 msec before the R wave on the ECG.
Prospective triggering techniques have important limitations. They are sensitive to heart rate changes and arrhythmias, but they have limited spatial z-axis resolution in order to cover the entire heart in a single breath-hold. They are effective only for heart rates of less than 90 beats per minute and perform poorly with arrhythmias, such as in atrial fibrillation. To overcome these limitations, retrospective ECG gating techniques are commonly used, at the expense of a higher radiation dose. Retrospective gating techniques allow faster continuous cardiac volume coverage, improved z-axis resolution, and imaging of the entire cardiac cycle for functional analysis.
In retrospective techniques, partially overlapping MDCT projections are continuously acquired, and the ECG signal is simultaneously recorded. Algorithms are then used to sort the data from different phases of the cardiac cycle by progressively shifting the temporal window of acquired helical projection data relative to the R wave (Fig. 5). Every position of the heart must be covered by a detector row at every point during the cardiac cycle. This means that the scanner table must not advance more than the total width of the active detectors for each heartbeat. Helical pitch can be varied proportionally to the heart rate to achieve continuous volume coverage. Typical pitch for an average heart rate of 70 beats per minute is 0.3:1, with a total scanning time of about 20 sec for a 16-slice MDCT scanner using 0.625-mm collimation.
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Two main algorithmic approaches are used to perform retrospective cardiac gating: partial scanning and segmented adapted scanning. To reconstruct an image, a minimum helical projection data segment of 180° must be available for every fan angle, corresponding to a rotation of 180° plus the breadth of the fan beam, so approximately two thirds of a full gantry rotation is required (Fig. 6A). This technique is called partial scanning. Temporal resolution is therefore two thirds of rotation time. Parallel beam geometries and rebinning techniques can be used to decrease the minimum data segment to 180°, with a temporal resolution of half the rotation time, or 250 msec for 16-slice MDCT scanners [32]. The partial scanning technique is typically used in patients with heart rates between 40 and 75 beats per minute.
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To improve temporal resolution, segmented adaptive reconstruction can be used, which involves combining shorter segments of projection data from two of more subsequent cardiac cycles [33] (Fig. 6B). Temporal resolution is equal to that of the longest projection data segment. Maximum values for temporal resolution are 125 msec for segmentation over two cardiac cycles and 65 msec over four cardiac cycles. Either volume coverage or longitudinal resolution may need to be reduced to maintain a low pitch and still scan in a single breath-hold.
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TopIntroductionHistorical BackgroundMDCT: General PrinciplesCardiac GatingTechnical ConsiderationsClinical IndicationsCoronary Artery Calcium ImagingCoronary Artery Lumen ImagingCardiac FunctionOther IndicationsConclusionReferences |
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Patient Position
The position of the patient in the CT scanner is important, because a
patient placed off-center in the scanner produces images with suboptimal
temporal resolution. The temporal resolution of any given pixel in an image
depends on the temporal window of the projection data used to reconstruct that
specific pixel. In helical CT using a partial scanning algorithm, a minimum
helical projection data segment of 180° must be available for every fan
angle, so each image is reconstructed from a segment of projection data longer
than 180°. However, not every detector contributes data at each projection
angle to reconstruct an image. For example, if a patient is scanned clockwise
from the 11-o'clock to the 7-o'clock position to produce an image (Fig.
7A,
7B,
7C,
7D), the left side of the
chest in the image has a much tighter temporal window than the right side. The
temporal resolution of the points of the image in this example increases
almost linearly between the edges of the field of view going from the left
side of the chest to the right side
[34]. By positioning the heart
near the center of rotation of the gantry, the temporal resolution of its
elements remains constant and average, rather than going from bad to better as
the X-ray tube rotates on the gantry.
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Heart Rate
The heart rate of the patient is an important determinant of image
reconstruction quality and for selecting scanning parameters such as pitch and
reconstruction algorithms (pitch is covered in the next section). If the heart
rate is too high (> 100 beats per minute), a temporal resolution of less
than 150 msec becomes necessary for motion-free cardiac imaging during
diastole. Segmented adaptive reconstruction can be used to improve temporal
resolution, but at the expense of volume coverage or longitudinal resolution,
to keep scanning time in one breath-hold. If neither can be sacrificed,
ß-blocker medication can be administered orally or IV 1 hr before
scanning to reduce the heart rate
[35].
If segmented adaptive reconstruction is used, improved temporal resolution can be achieved only if the patient's heart rate is not synchronized with the rotation cycle of the gantry. If they are synchronized, the same cardiac phase corresponds to the same angular segment of projection data at every rotation, and it is therefore not possible to build the 180° span of projection data required to reconstruct individual images.
A patient's heart rate may vary during scanning. The heart rate decreases after breath-holding [36]. If the heart rate varies during scanning after the rotation time and pitch are selected by the operator, the temporal resolution will vary from image to image.
Pitch
Helical pitch is a parameter that is selected before the examination is
acquired. Helical pitch is varied proportionally to the heart rate to achieve
continuous volume coverage (Fig.
8). If the heart rate increases, the pitch can increase. If the
heart rate decreases, the pitch must decrease. If the pitch is too high given
the heart rate, gaps in the image data set are present. A pitch that is too
low implies increased radiation exposure and increased duration of the
breath-hold. Before scanning, the heart rate of the patient should be assessed
under breath-holding conditions, and the lowest expected heart rate should be
used to select the pitch. The latest scanners can automatically adjust the
pitch when variations in the heart rate are detected.
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Collimation
Scans should be acquired at the best possible spatial resolution for the
specific task at hand. Coronary angiography requires the best resolution. The
left main artery has a diameter of 4 mm, and the distal left anterior
descending and circumflex arteries have diameters of 1 mm
[37]. Submillimeter spatial
resolution, both in-plane and longitudinally, is therefore required for
coronary CT angiography. Other examinations, such as calcium scoring, do not
require longitudinal spatial resolution this fine and are performed with
slightly larger collimation. It is always possible to reconstruct thicker
slices from the projection data of thinner slices, but thinner slices require
higher radiation exposure.
Dosimetry
Radiation exposure for cardiac CT is relatively high because of continuous
overlapping scanning and the use of retrospective cardiac gating. Coronary CT
angiography has a 5- to 10-mSv exposure, which is more than the typical 2 mSv
of traditional diagnostic catheter coronary angiography
[23]. Calcium scoring has an
exposure of 1–2 mSv
[23]. Modern scanners
incorporate fine control of tube current with respect to cardiac cycle, and
this will become more important as spatial resolution is improved in future
units. Radiation dose should be high enough to maintain an appropriate
contrast-to-noise ratio for diagnostic quality images, but no higher.
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Coronary Artery Calcium Imaging
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TopIntroductionHistorical BackgroundMDCT: General PrinciplesCardiac GatingTechnical ConsiderationsClinical IndicationsCoronary Artery Calcium ImagingCoronary Artery Lumen ImagingCardiac FunctionOther IndicationsConclusionReferences |
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Acquisition
A typical protocol used for 16-slice MDCT is as follows: axial 16-row mode,
1.25-mm collimation, 20-mm table feed between each acquisition, reconstruction
of eight images at 2.5-mm thickness per table location, 0.4- to 0.5-sec
rotation, scanning from the carina to cardiac apex, X-ray tube at 120 kV and
300–320 mAs, field of view of 25 cm, and prospective gating with a delay
of 70%. The protocol may be executed twice for reproducibility assessment.
Ideally, a continuous volume containing the heart should be scanned as rapidly
as possible in one breath-hold to minimize motion, with thin slices and the
lowest radiation exposure providing a sufficiently high contrast-to-noise
ratio. Cardiac gating should be used to avoid gaps. In this case, prospective
cardiac gating is used. A CT phantom can be placed underneath the patient to
calibrate the attenuation values in the final images
[13]. Initial scout
posteroanterior and lateral views are obtained to locate the position of the
heart. After a period of hyperventilation, the patient is asked to
breath-hold. The approximate image acquisition time by axial cine protocol is
15–16 sec, the exposure time is 3.4 sec, and the total radiation dose is
1–2 mSv [23].
Postprocessing
The total coronary calcium load is determined semiautomatically. A
threshold of 130 H and higher is used, and all clusters of calcifications in
the coronary arteries are identified by a postprocessing algorithm and
quantified. Clusters smaller than two pixels are considered to be noise and
are not included in the final assessment. For semiquantitative assessment of
coronary artery calcium, the Agatston calcium score is the most widely used
[41]. The total score is the
sum of all scores of calcium clusters on all images of the scan (Fig.
9A,
9B,
9C,
9D). The score of each single
lesion on each image is the product of its area multiplied by a cofactor
reflecting calcium density in each lesion. More precisely, the cofactor
indicates maximal attenuation in the lesion (1, 130–199 H; 2,
200–299 H; 3, 300–399 H; and 4, 
400 H). The Agatston score is
highly dependent on the imaging parameters and shows variable reproducibility,
especially for small amounts of calcium, with variations in the scores by 20%
or more from scan to scan
[42]. Volumetric
quantification algorithms are also available for calcium scoring
[43]. They have a higher
reproducibility (7–25% variability)
[43] and could replace the
Agatston score in the future.
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Comparison with Electron Beam CT
Although calcium scoring has been performed since the 1990s using electron
beam CT, it is slowly being replaced by scoring with MDCT because of the
greater availability of the latter type of units. MDCT has a poorer temporal
resolution than does electron beam CT (250 vs 100 msec), so images by MDCT
have more motion artifacts. However, MDCT has higher spatial resolution than
electron beam CT, lower image noise, and similar radiation exposure (1.0 vs
0.7 mSv) [23]. Noise can be
decreased in MDCT by varying tube current and voltage, which is useful for
obese patients. Increase in overlap of projection data (decreasing pitch)
results in higher reproducibility than electron beam CT
[44].
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Today most coronary CT angiography is still experimental. CT angiography is clinically indicated for evaluation of the patency of coronary vessels after intervention, for evaluation for a suspected anomalous coronary artery origin, and in patients with atypical chest pain. CT angiography may be useful for follow-up of patients undergoing statin therapy [46], to assess cardiac valves, and in other cardiac diseases such as tumors and pericardial disease.
Electron beam CT has been reported to have high sensitivity, specificity, and negative predictive value for assessment of the proximal and mid segments of the coronary arteries [47, 48]. The limited longitudinal spatial resolution of electron beam CT (1.5–3.0 mm) in a single breath-hold limits its assessment of distal coronary arteries. Images are also degraded by motion, and imaging of several coronary segments (e.g., the mid right coronary artery and the circumflex artery) is often of poor quality because of the atrial contraction at end-diastole. Studies have found that approximately 25% of coronary artery segments must be excluded because of poor image quality [47, 49]. The same is true for 4-slice MDCT, for which visibility of coronary segments is optimal for heart rates only below 65 beats per minute [50].
Because of its shorter acquisition time, superior spatial resolution (0.625-mm longitudinal resolution), and temporal resolution approaching that of electron beam CT, 16-slice MDCT is currently superior to electron beam CT and 4-slice MDCT for the evaluation of coronary artery disease [51], especially in coronary segments with diameters smaller than 2 mm.
At least four elements affect the quality of the images produced by CT angiography. First, the presence of heavy calcification creates streak artifacts and markedly limits the assessment of vessel lumen. The possible development of calcium subtraction techniques may minimize or eliminate this artifact. Second, excessive motion can markedly degrade image quality. The duration of the period of low motion decreases at high heart rates, making imaging more difficult. A heart rate greater than 75 beats per minute and arrhythmia are therefore limiting. The motion of the left main, left anterior descending, and circumflex coronary arteries mainly follows the motion of the left ventricle. The motion of the right coronary artery is related to the contraction of the right atrium and ventricle and is not negligible at end-diastole. Good assessment of the mid portion of the right coronary artery is especially problematic in many cases [52], a problem also seen with coronary MR angiography. Third, coronary CT angiography becomes less reliable for small vessels with a lumen diameter of 0.5–1.5 mm. Fourth, a breath-hold of approximately 30 sec may be required for adequate coronary imaging, depending on the heart size and heart rate. Although such a breath-hold is not a problem for normal test subjects (or with CT phantoms [45]), it can prove difficult for patients with severe cardiac disease.
Acquisition
A typical protocol for 16-slice MDCT is as follows: helical 16-row mode,
0.625-mm collimation (1.25-mm collimation if scanning time is > 30 sec),
400- to 600-msec rotation, pitch of 0.275:1–0.3:1, scanning from the
carina to the cardiac apex, X-ray tube at 120 kV and 340–400 mAs, field
of view of 25 cm, and retrospective gating. Initial scout frontal and lateral
images are obtained and the position of the heart is identified. After a
period of hyperventilation, the patient is asked to breath-hold. Unenhanced CT
through the coronary arteries is performed for the purpose of calcium scoring
and to confirm location of the heart. This is followed by a timing bolus to
determine the scanning delay for coronary CT angiography. A single axial scan
at 5-mm collimation approximately 1 cm below the carina is obtained every 2
sec to determine peak aortic root enhancement (Fig.
10A,
10B,
10C,
10D,
10E). A region of interest is
placed in the root of the aorta, near the origin of the left main coronary
artery. A test bolus of a total of 15–20 mL of iohexol 300
(Omnipaque-300, Nycomed Imaging) is administered at a rate of 4 mL/sec.
Additional parameters are field of view, 25 cm; and X-ray tube, 120 kV and 40
mAs. The image acquisition and exposure time is 12–15 sec, and the total
radiation dose is 0.33 mSv (Goodsitt MM, unpublished data). The scanning delay
is typically equal to the time from the start of the injection to peak
enhancement of the ascending aorta, plus 3–5 sec for filling of the
distal coronary arteries.
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The next step is the actual coronary CT angiography. A total dose of 100–150 mL of iohexol 300 is administered through an antecubital vein using a power injector at a rate of 4 mL/sec, using the scanning delay determined from the timing bolus. The rapid scanning time of the 16-slice MDCT scanner reduces the duration of the single breath-hold scan and therefore reduces the amount of IV contrast material necessary for the examination. This step is also performed after a period of hyperventilation. The image acquisition and exposure time is 15–30 sec.
Postprocessing
Postprocessing steps include retrospective reconstruction of the images at
a resolution of 0.6 x 0.6 x 0.625 mm, at phases 70%, 75%, and 80%
of the R-R interval. A partial scanning algorithm is used if the heart rate is
less than 60 beats per minute; otherwise, segmented adaptive scanning is used.
The sequences of images at all three reconstructed phases are transferred to a
processing workstation for further analysis with specialized software. The
three sequences are initially compared, and the sequence showing the least
amount of motion is selected for further processing. In practice, the sequence
of images reconstructed at 75% is the most often used for further processing,
so the other two sequences are examined only if the former is contaminated by
too much motion.
Initial processing of the acquired data produces an image of the global coronary tree, using 3D surface shading and 3D volume rendering. Examination of the coronary arteries is then performed, with the assistance of coronary artery analysis software, using 2D multiplanar reformations, 2D maximum intensity projections, and curved coronary artery reformations (Fig. 11A, 11B, 11C, 11D). The American College of Cardiology–American Heart Association guidelines for segmental anatomy [53] and lesion morphology [54] are used for lesion characterization. Each stenosis is graded by percentage of diameter, with reference to the coronary artery diameter proximal and distal to the stenosis. Bypass grafts are included in the scan and are evaluated from their proximal to distal anastomoses. Multiple findings can be identified in the coronary arteries, such as stents, aneurysms, or an aberrant origin (Figs. 12, 13, 14, 15, 16).
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Comparison with Catheter Angiography
Most studies comparing MDCT with catheter angiography for the evaluation of
coronary artery stenosis have been performed using 4-slice MDCT
[22,
55–60].
Image quality and spatial resolution for proximal and mid segments of the
coronaries have been good; however, as many as 30% of coronary artery segments
have been excluded because of motion
[22,
55,
60]. For stenosis exceeding
50% in diameter, Achenbach et al.
[22,
55] found a sensitivity of 85%
and a specificity of 76%; and Nieman et al.
[60] found sensitivity,
specificity, and positive and negative predictive values to be 82%, 93%, 66%,
and 97% respectively. Vogl et al.
[61] found a sensitivity of
73.4% using multiplanar reconstruction.
In recent series of patients, researchers compared 16-slice MDCT with coronary angiography [35, 62] (Fig. 17A, 17B). Image quality and spatial resolution of even distal segments and side branches were excellent [51], and temporal resolution and longitudinal spatial resolution improved compared with 4-slice MDCT [63]. For stenosis exceeding 50% in diameter, Nieman et al. [62] found sensitivity, specificity, and positive and negative predictive values to be 95%, 86%, 80%, and 97%, respectively. In that research, accuracy for detection of lesions was 100% for the left main artery, 91% for the left anterior descending artery, 81% for the circumflex artery, and 86% for the right coronary artery. Ropers et al. [35] obtained similar results for stenosis greater than 50%: sensitivity of 92%, specificity of 93%, accuracy of 93%, and positive and negative predictive values of 79% and 97%, respectively.
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TopIntroductionHistorical BackgroundMDCT: General PrinciplesCardiac GatingTechnical ConsiderationsClinical IndicationsCoronary Artery Calcium ImagingCoronary Artery Lumen ImagingCardiac FunctionOther IndicationsConclusionReferences |
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Wall motion is qualitatively assessed by the radiologist from cine loops in the long- and short-axis planes. Automated left ventricular endocardial and epicardial contour detection is performed using specialized software on the short-axis images of the left ventricle for all 20 phases of the cardiac cycle. Following conventions used for echocardiography, the aortic outflow tract is not included in the volume, and the papillary muscles are included inside the endocardial borders [64]. The contours are then verified by a CT technologist or a radiologist, and fine manual adjustments can be made, if necessary, to ensure correct placement.
From the endocardial and epicardial contours at each phase of the cardiac cycle (Fig. 19A, 19B), the software computes two sets of measurements. First, the volumes bounded by the endocardium and epicardium are computed using Simpson's rule [64–66], and the variation of volumes throughout the cardiac cycle is determined [67]. From these basic measurements, the end-systolic and end-diastolic volumes, ejection fraction, stroke volume, cardiac output, and cyclic variations in the myocardial wall thickness are derived. The results are displayed graphically. Second, the radial excursion of each point of the endocardium throughout the cardiac cycle is computed and displayed graphically for three levels (base, mid left ventricle, apex) to determine regional left ventricular wall motion.
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TopIntroductionHistorical BackgroundMDCT: General PrinciplesCardiac GatingTechnical ConsiderationsClinical IndicationsCoronary Artery Calcium ImagingCoronary Artery Lumen ImagingCardiac FunctionOther IndicationsConclusionReferences |
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Several new experimental but promising roles are possible for cardiac CT with the latest scanners. Plaque characterization seems promising. Not all coronary artery plaques are calcified: some are soft and fatty, and others are more fibrous. Lipid-rich noncalcified plaques are most prone to spontaneous rupture, leading to acute cardiac events and death, and are known as "vulnerable plaque" [69]. Until now, only intravascular sonography could be used to differentiate lipid-rich and fibrous plaques. The noninvasive accurate identification of vulnerable plaque may lead to preemptive intervention before an acute coronary event, either by invasive means such as angioplasty and stenting, or by pharmacologic means such as statin therapy [46]. Recent research finds that MDCT enables not only identification but also characterization of coronary plaques. Schroeder et al. [70], using intravascular sonography as the reference test, showed that the average density of coronary plaques correlated well with their echogenicity on sonography. Average density of the plaques was 14 ± 26 H for soft plaque (low echogenicity), 91 ± 21 H for intermediate plaque (bright echoes), and 419 ± 194 H for calcified plaque (bright echoes with acoustic shadowing). Nikolaou et al. [71] found that lipid-rich plaques have a density of 50 ± 12 H, whereas fibrous plaques have a density of 89 ± 31 H, using histopathology as the reference test. Morphologic criteria can also be used to determine the presence of thrombus [72]. Collimation is of major importance in plaque characterization, and quantitative density measurements are affected by volume averaging from the contrast-enhanced vessel lumen with thicker collimation.
Another area of investigation is the identification and characterization of myocardial perfusion defects after an infarct [73]. Differences in myocardial wall enhancement after the administration of IV contrast material occur as a result of vascular obstruction and local edema, and have been detected by CT for years. The improved temporal and spatial resolution of MDCT will likely allow more precise characterization and follow-up. Similar assessment of myocardial perfusion is currently done by MRI with the patient at rest and under stress, and enables detection of infarction and ischemia [74].
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TopIntroductionHistorical BackgroundMDCT: General PrinciplesCardiac GatingTechnical ConsiderationsClinical IndicationsCoronary Artery Calcium ImagingCoronary Artery Lumen ImagingCardiac FunctionOther IndicationsConclusionReferences |
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TopIntroductionHistorical BackgroundMDCT: General PrinciplesCardiac GatingTechnical ConsiderationsClinical IndicationsCoronary Artery Calcium ImagingCoronary Artery Lumen ImagingCardiac FunctionOther IndicationsConclusionReferences |
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