Download Relationship between Noise, Dose, and Pitch in Cardiac Multi

RadioGraphics
EDUCATION EXHIBIT
1785
Relationship between
Noise, Dose, and Pitch
in Cardiac Multi–
Detector Row CT1
TEACHING
POINTS
See last page
Andrew N. Primak, PhD ● Cynthia H. McCollough, PhD ● Michael R.
Bruesewitz, RT(R) ● Jie Zhang, PhD ● Joel G. Fletcher, MD
In spiral computed tomography (CT), dose is always inversely proportional to pitch. However, the relationship between noise and pitch (and
hence noise and dose) depends on the scanner type (single vs multi–
detector row) and reconstruction mode (cardiac vs noncardiac). In
single detector row spiral CT, noise is independent of pitch. Conversely, in noncardiac multi– detector row CT, noise depends on pitch
because the spiral interpolation algorithm makes use of redundant data
from different detector rows to decrease noise for pitch values less than
1 (and increase noise for pitch values ⬎ 1). However, in cardiac spiral
CT, redundant data cannot be used because such data averaging
would degrade the temporal resolution. Therefore, the behavior of
noise versus pitch returns to the single detector row paradigm, with
noise being independent of pitch. Consequently, since faster rotation
times require lower pitch values in cardiac multi– detector row CT,
dose is increased without a commensurate decrease in noise. Thus, the
use of faster rotation times will improve temporal resolution, not alter
noise, and increase dose. For a particular application, the higher dose
resulting from faster rotation speeds should be justified by the clinical
benefits of the improved temporal resolution.
©
RSNA, 2006
Abbreviations: CAC ⫽ quantitation of coronary artery calcium, ECG ⫽ electrocardiographically
RadioGraphics 2006; 26:1785–1794 ● Published online 10.1148/rg.266065063 ● Content Codes:
1From
the CT Clinical Innovation Center, Department of Radiology, Mayo Clinic College of Medicine, 200 First Street SW, Rochester, MN 55905.
Recipient of a Certificate of Merit award for an education exhibit at the 2005 RSNA Annual Meeting. Received April 12, 2006; revision requested
May 30 and received July 6; accepted July 10. A.N.P., C.H.M., J.Z., and J.G.F. receive research grants from Siemens Medical Solutions, Malvern, Pa;
C.H.M. also receives a research grant from GE Healthcare, Waukesha, Wis; J.G.F. has an educational license from GE Healthcare and participates in
a CME course sponsored by E-Z-Em, Lake Success, NY; M.R.B. has no financial relationships to disclose. Address correspondence to A.N.P.
(e-mail: [email protected]).
©
RSNA, 2006
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Glossary of Technical Terms
Term
Definition
N
Noise
Variable representing the total number of photons used to form an image
Standard deviation of pixel values measured within a uniform region of
interest in the image (expressed in Hounsfield units)*
Ratio of the table increment to the total nominal beam width†
Set of x-ray transmission data through an object collected at one angular
position of the x-ray tube around the object
Process by which spiral CT projection data, which are collected at continuously varying z-axis positions, are transformed into projection data
at one specific z-axis location for the purpose of reconstructing an axial
image
Distance that the table is advanced per rotation of the x-ray tube
Total width of the imaged tomographic section (the sum of all individual
detector widths), measured at isocenter of the scanner‡
Pitch
Projection
Spiral interpolation algorithm
Table increment
Total nominal beam width
*Also known as quantum mottle, this is a quantitative measure of the amount of statistical variation in the image.
†For example, for a 15-mm per rotation table increment and 4 ⫻ 5-mm (20-mm) beam width, the pitch ⫽ 0.75.
‡For example, for a four– detector row system acquisition with a 5-mm detector row width, the total nominal
beam width ⫽ 4 ⫻ 5 mm ⫽ 20 mm.
Introduction
The improved diagnostic capabilities of multi–
detector row computed tomographic (CT) technology have resulted in an increasing number of
CT examinations, which represent a significant
portion of the radiation dose received from all
medical procedures (1–3). The potential radiation risk from this increased use of CT makes it
important that CT doses be kept as low as reasonably achievable. To fulfill this goal, it is important
to understand the relationship between dose and
image noise, as noise is a major factor in determining acceptable image quality and often dictates the dose for a particular CT protocol.
In spiral CT, both dose and noise depend on
pitch, but not in the same way. While dose is always inversely proportional to pitch, the behavior
of noise as a function of pitch depends on scanner
type (single vs multi– detector row) and reconstruction mode (cardiac vs noncardiac). Further,
in cardiac CT, the relationship between pitch and
patient heart rate must also be taken into account
(4,5).
In this article, we review the underlying principles and demonstrate the relationships between
image noise, dose, and pitch for cardiac and noncardiac multi– detector row CT. We also discuss
the clinical implications of these relationships.
The Table provides a glossary of several technical
terms used in these discussions.
Radiation Dose
Radiation dose is related to the amount of energy
that x-ray photons deliver during a CT scan. It
depends on the total number of photons and their
individual energies. The energy distribution of
these photons depends on the x-ray tube potential
and the spectral (bowtie) filter. The total number
of photons is proportional to the tube current (in
milliamperes) and the “x-ray on” time (in seconds) during a single gantry rotation, and hence it
is proportional to the tube current–time product
(in milliampere-seconds). Thus, for a single axial
scan, or the most common scenario of contiguous
scans (where table increment ⫽ total nominal
beam width), dose is directly proportional to the
tube current–time product.
In the case of multiple axial scans or spiral
scans with the table increment ⫽ total nominal
beam width, an accumulation of dose due to x-ray
beam overlap, or a decrease in dose due to gaps in
the radiation beams, from successive gantry rotations will occur. The quantitative measure of such
overlap or gap is given by the ratio of table increment per rotation to total nominal beam width,
which is known as pitch. Therefore, the general
relationship exists that dose is proportional to
tube current–time product/pitch (which on systems manufactured by Siemens Medical Solutions [Forchheim, Germany] is called “effective
mAs” and on systems manufactured by Philips
Medical Systems [Andover, Mass] is called “mAs
per slice”) (6 – 8).
For examinations where the patient table is not
incremented between successive rotations of the
x-ray tube (eg, dynamic CT, CT perfusion, or
interventional CT examinations), dose will accumulate in the irradiated section of tissue over the
multiple exposures at that location. Thus, for
these examinations, dose is proportional to the
tube current–time product multiplied by the
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number of scans in the examination (ie, the number of tube rotations where the x-ray tube is energized).
Noise
In general, noise in CT depends on the number
of x-ray photons reaching the detector (quantum
noise), the electronic noise of the detector system,
and the reconstruction kernel (sharper kernels
give noisier images). Unless images suffer from
severe photon starvation (eg, in morbidly obese
patients), quantum noise plays the dominant role.
Since x-ray photon statistics obey the Poisson
distribution, quantum noise is proportional to
公N and the corresponding image noise is approximately proportional to 1/公N, where N is the
number of photons that have contributed to the
reconstructed image. Since the number of photons reaching the detector depends on the object
attenuation, which in turn depends on photon
energies, N is strongly dependent on tube potential. In addition, N is proportional to section
width, tube current, and the amount of time necessary to acquire all the projection data needed
for the reconstruction. In sequential mode, this
time equals the “x-ray on” time per rotation, so
image noise is approximately proportional to
1/公mAs.
In spiral mode, however, the interpolation algorithm, which transforms the projection data
acquired at various z-axis locations into projection data at one specific z-axis location, must be
taken into account. Because the spiral interpolation algorithm is inherently different for multi–
detector row CT compared to single detector row
CT (9), the relationship between noise and pitch
in spiral CT depends on the scanner type (single
vs multi– detector row CT). In addition, because
cardiac spiral reconstructions are optimized to
decrease motion artifact (4,5) (ie, provide the best
possible temporal resolution), the relationship
between noise and pitch also depends on the
multi– detector row CT reconstruction mode
(cardiac vs noncardiac).
Single Detector Row Spiral CT
Teaching
Point
In single detector row spiral CT, the gantry always has to rotate through a certain angle (dependent on the reconstruction algorithm) in order to
acquire the projection data needed for an axial
image (10). Thus, N depends only on tube current–time product and not on the table speed or
pitch. An intuitive way of understanding this effect is to realize that regardless of the spacing of
the consecutive x-ray tube rotations (ie, pitch),
the same amount of projection data is required to
reconstruct an image (either 720° or 360° of data,
depending on the spiral interpolation algorithm
implemented on the system; 360° is the most
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common). Because the same amount of image
data (photons) is used to reconstruct the image,
the image noise remains constant even as pitch is
changed.
This relationship was counterintuitive to many
users, since it has been well (and correctly) publicized that increasing pitch decreases patient dose
(11). Since intuition teaches that noise increases
when dose decreases, the incorrect assumption
that noise depended on pitch resulted. Studies
have shown that indeed, noise is independent of
pitch, while dose decreases with increasing pitch
in single detector row CT (12).
This raises the obvious question as to how one
can achieve a dose decrease at the same image
noise level, as typically in imaging, one does not
gain an advantage (decreased dose) without paying some price. The price of an increased pitch is
a widening of the section sensitivity profile (ie, the
width of the reconstructed image) (13). This widening occurs because, at higher pitch values, the
360° of required projection data are obtained
from z-axis locations that are spread farther apart
(relative to use of a small pitch value). Thus, data
are averaged over a wider z-axis distance, widening the effective thickness of the reconstructed
image (13).
Multi–Detector Row Spiral CT
Multi– detector row spiral CT reconstruction algorithms differ profoundly from single detector
row spiral CT reconstruction algorithms because
of the flexibility provided from having data collected simultaneously at multiple z-axis locations
(multiple detector rows). The interpolation process in multi– detector row CT can select and utilize data from the detector row nearest to the requested z-axis image location, regardless of which
row those data come from. Using a modified 360°
interpolation approach, there are redundant data
points (ie, data passing through the same z-axis
position) from different detectors as they rotate
about the patient. The frequency of these redundancies is determined by the pitch values, with
greater redundancy at lower pitch values and less
redundancy at higher pitch values.
To make use of all the photons (dose) applied
to the patient, multi– detector row CT interpolation algorithms make use of the redundant data
by averaging redundant projection values, which
effectively decreases image noise. As the pitch is
increased, there are more gaps in the data with
fewer redundant points to average, so the image
noise increases. Thus, the paradigm from single
detector row CT is altered in multi– detector row
CT, and noise becomes dependent on pitch. The
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Figure 1. Graphs of the interpolation algorithm used to generate planar data from the measured spiral data.
The dotted lines show the center of every detector row, whereas the solid lines indicate the detector boundaries. (See the text for additional details.)
relationship of dose to pitch is unchanged. However, this paradigm is valid only for routine spiral multi– detector row CT reconstruction algorithms, as the special algorithms for cardiac
multi– detector row CT alter the noise versus
pitch relationship for yet a third time.
Noncardiac Mode.—As discussed earlier, in
noncardiac spiral multi– detector row CT, noise
depends on pitch through an interpolation algorithm used to generate a set of planar projection
data (Fig 1). All spiral data within a predefined
spiral interpolation window (horizontal blue
dashes) are used to generate planar data for the
image plane (vertical red line).
When pitch is less than 1, the measured spiral
data partially overlap (shaded areas) in the z direction (perpendicular to the gantry plane) (Fig
1b), so some portions of the planar projection
data for the same axial image can be generated
more than once (from the spiral data acquired by
different detector rows). These redundant data
result in more x-ray photons contributing to the
reconstruction of the axial images compared to
when pitch ⫽ 1, reducing the image noise and
making it pitch dependent. For example, a scan
with pitch ⫽ 0.5 needs twice the number of gantry rotations to cover the same distance compared
to when pitch ⫽ 1. This results in fully overlapped spiral data, which can generate two completely redundant sets of planar projection data
for every axial image (Fig 1c). Combining these
two redundant data sets into one allows reconstruction of axial images where noise is reduced
by a factor of 公2.
When pitch is more than 1, the measured spiral data have gaps in the z direction (Fig 1d).
These missing data result in less x-ray photons
contributing to the reconstruction of each axial
image compared to when pitch ⫽ 1, increasing
the image noise and making it pitch dependent.
Generally speaking, to offset the increase in
noise as pitch is increased (or offset the increase
in dose as pitch is decreased), CT system manufacturers provide some mechanism to adjust the
tube current so that tube current–time product is
increased approximately proportional to the increase in pitch (or decreased approximately proportional to the decrease in pitch). Thus, as long
as effective mAs (defined as tube current–time
product/pitch) is held constant, both dose and
noise remain constant.
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Figure 2. Noise and dose versus tube current–time product and effective mAs for noncardiac (a, b) and cardiac (c, d) spiral CT modes at four different pitch (p) values and two different gantry rotation times (0.33 seconds
and 0.37 seconds). The left y axis corresponds to noise curves; the right y axis corresponds to dose curves. Note that
in cardiac mode, noise is dependent on tube current–time product and not on pitch (c). CTDIvol ⫽ Volume CT Dose
Index.
This is demonstrated in
Figure 2a and 2b, where noise and dose are
shown as a function of tube current–time product and as a function of effective mAs.
For Siemens Medical Solutions scanners, the
manufacturer adopted the term “effective mAs”
to help users realize that as long as the ratio of
tube current–time product to pitch was held constant, they could reliably predict a constant noise
level for a given object size. Philips Medical Systems similarly used the term “mAs per slice” to
accomplish the same task. On GE Healthcare
Technologies (Waukesha, Wis) systems, when the
user alters the pitch values, the user interface automatically alters the prescribed tube current to a
value that will maintain a constant noise. The
new (suggested) tube current is highlighted with
orange to alert the user that this change has occurred. If the user does not desire to hold the
image noise constant at the new pitch value, the
user can manually override the suggested value.
Toshiba Medical Systems (Tokyo, Japan) systems
have a similar mechanism of suggesting to the
operator the appropriate tube current value when
the pitch value is changed.
In Figure 2b, we observe that the noise is constant for a given effective mAs, completely independent of pitch. This complete “pitch independence” is a result of the chosen spiral interpolation algorithm and the manufacturer’s conscious
decision to allow the user to alter pitch in a continuous fashion to prescribe the desired scan acquisition time without having to worry about the
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Figure 3. (a) Anthropomorphic cardiac phantom with small calcified cylinders used for calcium quantitation
(arrows). (b) Axial image shows the water-equivalent cylindrical insert used for noise versus tube current–time
product measurements (arrow).
effect of the pitch on image quality. On Siemens
Medical Solutions systems, both noise and section width are essentially independent of pitch for
a constant effective mAs (9,14). Other manufacturers have implemented spiral interpolation and
reconstruction algorithms that have preferred
pitch values with regard to both image width and
image noise. For example, GE Healthcare Technologies systems allow the user four discrete pitch
values from which to choose, with the two lowest
values providing the narrowest actual image
thickness (for a selected nominal value) at the
expense of higher noise. Thus, data such as presented in Figure 2b show comparable trends but
do not coalesce into one curve. This makes the
precise tube current value needed at a given pitch
(in order to maintain constant noise) a bit more
difficult to predict; thus, the user should rely on
the suggested tube current values provided by the
scanner interface (15).
Teaching
Point
Cardiac Mode.—In cardiac spiral multi– detector row CT, the best possible temporal resolution
is required to minimize artifacts resulting from
cardiac motion. This goal is achieved by minimizing the number of projections used to reconstruct
the image to those projections gathered in the
shortest possible time window. The use of redundant data would not be acceptable because it
would degrade the temporal resolution by averaging data over one or more rotations of the x-ray
tube. The minimum amount of data required to
reconstruct a CT image is 180° plus the angle (in
degrees) of the x-ray beam in the plane of the image (known as the fan angle). Hence, cardiac algorithms use partial reconstruction techniques
(180° ⫹ fan angle) to reconstruct an image.
These data are collected either during a single
cardiac cycle (single-segment reconstruction) or
during two or more consecutive heartbeats (multisegment reconstruction). In both cases, the
number of photons N contributing to the cardiac
reconstruction depends only on the tube current
and the time it takes for the gantry to rotate
through 180° plus the fan angle, and not pitch.
Since this time is proportional to the rotation
time, N (and hence noise) is dependent only on
tube current–time product and is not affected by
pitch. This is analogous to the case of single detector row CT, which uses exactly 360° of data to
reconstruct an image and hence has noise that is
independent of pitch. Figure 2c and 2d demonstrates the independence of noise on pitch for cardiac reconstructions, even though dose remains
dependent on pitch. Hence, a constant effective
mAs in cardiac multi– detector row CT does not
guarantee equivalent noise, but rather only
equivalent dose.
Noise versus Tube Current–
Time Product Measurements
Noise versus tube current–time product data were
obtained on a 64-channel CT scanner (Sensation
64; Siemens Medical Solutions) by using an anthropomorphic cardiac CT phantom (16) (QRM,
Möhrendorf, Germany) (Fig 3a). The phantom
was scanned by using both electrocardiographically (ECG) gated (with ECG signal provided by
an ECG simulator on the CT system) and nongated spiral modes with four different pitch values
(0.18, 0.20, 0.24, and 0.27 for gated and 0.5, 0.6,
0.7, and 0.8 for nongated) and two rotation times
(0.33 and 0.37 seconds). Scanning was done at
120 kVp and variable tube current–time product
values. We used the 32 ⫻ 0.6-mm collimation
with a z-flying focal spot technique (17), which
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Figure 4. Visual demonstration of the relationships between noise, dose, and pitch. Cardiac phantom images
are shown for low and high pitch values. (a) Noncardiac images obtained at the same effective mAs have the
same noise. (b) Cardiac images obtained at the same effective mAs do not have the same noise (but do have the
same dose). (c) Conversely, cardiac images obtained at the same tube current–time product (not effective mAs)
do have the same noise (but do not have the same dose). CTDIvol ⫽ Volume CT Dose Index, SD ⫽ standard
deviation.
resulted in 64 overlapping projections per rotation. Contiguous 3-mm-thick axial sections were
reconstructed with a 300-mm field of view and
B35 kernel. Noise was measured as the standard
deviation of the pixel values within a waterequivalent cylinder embedded in the central portion of the cardiac phantom (Fig 3b). Dose was
assessed by using the Volume CT Dose Index (in
milligrays), per International Electrotechnical
Commission publication 60601-2-44 (18).
The results of these measurements are shown
in Figure 2 to demonstrate the quantitative relationships described earlier. A visual demonstration of these relationships is given in Figure 4,
where for low and high pitch values, phantom
images obtained at the same effective mAs are
shown. In Figure 4a, the noncardiac images have
the same noise texture independent of pitch because the effective mAs was held constant. In
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Figure 4b, the cardiac images obtained at the
same effective mAs do not have the same image
noise (though they do have the same dose). In
Figure 4c, the cardiac images obtained at the
same tube current–time product (not effective
mAs) do have the same image noise (but not the
same dose).
Pitch in Cardiac
Multi–Detector Row CT
Temporal resolution is of fundamental importance to cardiac multi– detector row CT and was
a driving force behind making gantries rotate
faster and faster. However, faster gantry rotation
requires a slower pitch in cardiac mode to avoid
discontinuities in the anatomic coverage of the
heart between images reconstructed from consecutive cardiac cycles (Fig 5) (4,5,19). For example, on a CT scanner with a rotation time of
0.5 seconds (Volume Zoom, Siemens Medical
Solutions), the necessary pitch was 0.375. On a
CT scanner with a rotation time of 0.33 seconds
(Sensation 64, Siemens Medical Solutions), the
necessary pitch is 0.2.
We can derive a good approximation of the
relationship between pitch, heart rate, and the
rotation time by taking into account that for
single-segment reconstruction (all data acquired
in one cardiac cycle), the table should not move
more than the total nominal beam width W during the time of one heart cycle (R-R interval time
TRR). This allows all phases of the heart to be
“seen” by some part of the detector at any z-axis
location. This means that the table speed V
should be less than or equal to W/TRR. Keeping in
mind the definition of pitch (table increment/total
nominal beam width), pitch ⫽ VTrot/W, where
Trot is the rotation time. With some minor algebraic manipulation, we obtain the following general requirement:
pitch ⱕ
Trot
.
TRR
(1)
To derive the exact relationship, one has to consider the details of a particular cardiac reconstruction algorithm (19,20). For example, according
to Ohnesorge et al (19), the maximum pitch for
single-segment reconstruction is given by the following formula:
pitch ⱕ
M ⫺ 1 Trot
,
M TRR
(2)
where M is the number of detector rows in the
cardiac mode. For modern state-of-the-art multi–
detector row CT scanners with M ⱖ 32, this pitch
Figure 5. Diagram of an ECG-gated spiral CT scan
with pitch that is too high for the heart rate. Continuous anatomic coverage of the heart is not possible because of the volume gaps between the images reconstructed from the consecutive cardiac cycles. R ⫽ R
wave, Recon ⫽ reconstruction.
restriction is very close to our intuitive approximation given in Equation (1).
The clinical implications of Equation (2) are
quite important. At faster gantry rotation times, a
lower pitch value is required. However, to achieve
the same noise, one has to use the same tube current–time product value (recall that noise is independent of pitch in cardiac multi– detector row
CT). The use of the same tube current–time
product value (relative to slower rotation time
scans), but smaller pitch values, results in higher
radiation doses (dose is always inversely proportional to pitch). Thus, for single-source (ie, one
x-ray tube) multi– detector row CT systems, better temporal resolution in cardiac spiral CT
comes with the price of higher dose. (These conclusions do not necessarily apply to the recently
introduced dual-source CT system [21], which
allows pitch values to be increased as heart rate
increased, thus offsetting the increased dose that
accompanies improved temporal resolution in
single-source cardiac multi– detector row CT.)
Clinical Implications
In noncardiac multi– detector row CT, the implications of the relationship between noise, pitch,
and dose are that as long as the ratio of tube current–time product to pitch is held relatively constant, a relatively constant image noise will result.
For the Siemens Medical Solutions systems, as
shown earlier, the relationship is exact. For other
manufacturers, the relationship holds in general,
although small variations in noise (10%–20%)
may occur even if the ratio of tube current–time
product to pitch is held constant. This relation-
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Figure 6. Percentage of area stenosis versus rotation
time for a moving stenotic vessel phantom scanned on
16 – and 64 – detector row CT systems with coronary
CT angiography protocols and multiple gantry rotation
times (0.33, 0.42, and 0.5 seconds). The measurements were performed for 0.6-, 0.75-, and 1-mm section widths. (See the text for additional details.)
ship can be used to the advantage of the operator
when long scan ranges are required yet the total
scan time is desired to be kept short (ie, a short
breath hold for a thoracic-abdominal-pelvic CT
angiography examination). The operator can increase the pitch (table speed) in order to keep the
examination time short. Image noise will not be
compromised as long as the system can provide a
sufficient tube current–time product value to keep
the ratio of tube current–time product/pitch constant (relative to a lower pitch examination over a
short anatomic range).
In cardiac spiral multi– detector row CT, faster
rotation times deliver improved temporal resolution, but require higher radiation dose to achieve
equivalent image noise (compared to slower rotation times). Thus, the use of faster rotation times
for a particular application should be justified by
clinical benefits. We have examined in a phantom
model two clinical cardiac CT applications where
these issues may be relevant: coronary artery angiography (CT angiography) and quantitation of
coronary artery calcium (CAC).
In our CT angiography phantom study (22),
an anthropomorphic cardiac CT phantom containing an iodine-filled vessel phantom with a stenosis was scanned by using retrospective ECGgated CT angiography protocols with 16- and
64-channel multi– detector row CT systems (Sensation 16 and 64; Siemens Medical Solutions).
Data were acquired at rest (0.33 seconds rotation
time only) and while the vessel phantom was
moving at physiologic coronary artery velocities
for multiple gantry rotation times (0.33, 0.42, and
Figure 7. Calcium mass score measured in a rotating
insert of an anthropomorphic cardiac phantom scanned
on 16 – and 64 – detector row CT systems with CAC
protocols and multiple gantry rotation times (0.33,
0.37, 0.42, and 0.5 seconds). HA ⫽ hydroxyapatite
concentration in milligrams per cubic centimeter; the
diameter of the calcification is given in millimeters.
(For example, 400 HA 3 mm corresponds to 400 mg/
cm3 HA density and a 3-mm-diameter cylinder.)
0.5 seconds). The stenoses were measured by
using a semiautomated tool (VesselView; Siemens
Medical Solutions) for section widths of 0.6 mm
(64-channel system only), 0.75 mm, and 1.0 mm.
The accuracy of percent area stenosis measurements in a moving phantom improved significantly (P ⬍ .01) as the rotation time decreased
from 0.5 to 0.33 seconds (Fig 6). Hence, for CT
angiography, the benefit of faster rotation time
appears to outweigh the disadvantage of the increased dose. These findings in a simple phantom
model are consistent with reported clinical experiences (23–26).
In our CAC phantom study (27), the central
portion of an anthropomorphic cardiac phantom
containing nine cylinders of varying amounts of
hydroxyapatite (Fig 3a) was attached to a rotational motion device and imaged by using retrospective ECG-gated CAC protocols with Sensation 16 and 64 CT systems (Siemens Medical
Solutions). The data were acquired during rotational motion corresponding to the mean velocity
of coronary arteries by using multiple gantry rotation times (0.33, 0.37, 0.42, and 0.5 seconds).
Calcium scores were obtained by using an automated tool (CaScoring; Siemens Medical Solutions). The accuracy of CAC measurements was
not statistically different (P ⬎ .05) for calcifications greater than or equal to 3 mm and densities greater than or equal to 400 mg/cm3 (Fig 7)
among the four rotation times tested. Thus, for
screening purposes, slower rotation times may be
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more appropriate, as they appear equally accurate
yet deliver lower radiation dose to the patient.
These findings are also consistent with reported
clinical experiences of reasonably robust CAC
with multi– detector row CT systems having both
slower and faster rotation times.
Summary
The relationship between image noise and radiation dose is not the same for cardiac and noncardiac spiral scanning with multi– detector row CT
systems. In noncardiac spiral multi– detector row
CT, noise depends on pitch, which results in
comparable noise when the ratio of tube current–
time product to pitch is held constant. In cardiac
spiral multi– detector row CT, noise depends only
on tube current–time product and not on pitch,
while radiation dose remains inversely proportional to pitch. Since faster rotation speeds require lower pitch values (relative to slower rotation speed), they result in higher radiation dose
to achieve equivalent image noise, and their use
should be justified by clinical benefits, such as in
the case of coronary CT angiography.
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RadioGraphics
RG
Volume 26 • Volume 6 • November-December 2006
Primak et al
Relationship between Noise, Dose, and Pitch in Cardiac Multi–
Detector Row CT
Andrew N. Primak, PhD, et al
RadioGraphics 2006; 26:1785–1794 ● Published online 10.1148/rg.266065063 ● Content Codes:
Pages 1787
In single detector row spiral CT, the gantry always has to rotate through a certain angle (dependent
on the reconstruction algorithm) in order to acquire the projection data needed for an axial image
(10). Thus, N depends only on tube current–time product and not on the table speed or pitch.
Page 1788
Generally speaking, to offset the increase in noise as pitch is increased (or offset the increase in dose
as pitch is decreased), CT system manufacturers provide some mechanism to adjust the tube current
so that tube current–time product is increased approximately proportional to the increase in pitch (or
decreased approximately proportional to the decrease in pitch). Thus, as long as effective mAs
(defined as tube current–time product/pitch) is held constant, both dose and noise remain constant.
Page 1790
In cardiac spiral multi–detector row CT, the best possible temporal resolution is required to minimize
artifacts resulting from cardiac motion. This goal is achieved by minimizing the number of projections
used to reconstruct the image to those projections gathered in the shortest possible time window. The
use of redundant data would not be acceptable because it would degrade the temporal resolution by
averaging data over one or more rotations of the x-ray tube. The minimum amount of data required
to reconstruct a CT image is 180• plus the angle (in degrees) of the x-ray beam in the plane of the
image (known as the fan angle). Hence, cardiac algorithms use partial reconstruction techniques
(180• + fan angle) to reconstruct an image.
Page 1792
However, faster gantry rotation requires a slower pitch in cardiac mode to avoid discontinuities in the
anatomic coverage of the heart between images reconstructed from consecutive cardiac cycles (Fig 5)
(4,5,19).
Page 1792
At faster gantry rotation times, a lower pitch value is required. However, to achieve the same noise,
one has to use the same tube current–time product value (recall that noise is independent of pitch in
cardiac multi–detector row CT). The use of the same tube current–time product value (relative to
slower rotation time scans), but smaller pitch values, results in higher radiation doses (dose is always
inversely proportional to pitch). Thus, for single-source (ie, one x-ray tube) multi–detector row CT
systems, better temporal resolution in cardiac spiral CT comes with the price of higher dose.