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Copyright © 2017 The Korean Society of Radiology
INTRODUCTION
CT is a cross-sectional, high-resolution, three-dimensional
diagnostic imaging modality that generally uses single-
energy polychromatic X-rays. Its recently increased clinical
utility is primarily attributed to significantly increased scan
Dual-Energy CT: New Horizon in Medical Imaging
Hyun Woo Goo, MD, PhD1, Jin Mo Goo, MD, PhD2
1Department of Radiology and Research Institute of Radiology, Asan Medical Center, University of Ulsan College of Medicine, Seoul 05505, Korea;
2Department of Radiology, Seoul National University College of Medicine, Seoul 03080, Korea
Dual-energy CT has remained underutilized over the past decade probably due to a cumbersome workflow issue and current
technical limitations. Clinical radiologists should be made aware of the potential clinical benefits of dual-energy CT over
single-energy CT. To accomplish this aim, the basic principle, current acquisition methods with advantages and disadvantages,
and various material-specific imaging methods as clinical applications of dual-energy CT should be addressed in detail. Current
dual-energy CT acquisition methods include dual tubes with or without beam filtration, rapid voltage switching, dual-layer
detector, split filter technique, and sequential scanning. Dual-energy material-specific imaging methods include virtual
monoenergetic or monochromatic imaging, effective atomic number map, virtual non-contrast or unenhanced imaging, virtual
non-calcium imaging, iodine map, inhaled xenon map, uric acid imaging, automatic bone removal, and lung vessels analysis.
In this review, we focus on dual-energy CT imaging including related issues of radiation exposure to patients, scanning and
post-processing options, and potential clinical benefits mainly to improve the understanding of clinical radiologists and thus,
expand the clinical use of dual-energy CT; in addition, we briefly describe the current technical limitations of dual-energy CT
and the current developments of photon-counting detector.
Keywords: Dual-energy CT; CT imaging techniques; Spectral CT; Virtual monoenergetic imaging; Effective atomic number;
Material decomposition; Photon-counting detector
Received January 28, 2017; accepted after revision February 23,
2017.
Corresponding author: Hyun Woo Goo, MD, PhD, Department
of Radiology and Research Institute of Radiology, Asan Medical
Center, University of Ulsan College of Medicine, 88 Olympic-ro 43-
gil, Songpa-gu, Seoul 05505, Korea.
• Tel: (822) 3010-4388 • Fax: (822) 476-0090
• E-mail: hwgoo@amc.seoul.kr
This is an Open Access article distributed under the terms of
the Creative Commons Attribution Non-Commercial License
(http://creativecommons.org/licenses/by-nc/4.0) which permits
unrestricted non-commercial use, distribution, and reproduction in
any medium, provided the original work is properly cited.
Korean J Radiol 2017;18(4):555-569
speed as a synergic effect of increased gantry rotation speed
and increased longitudinal detector coverage, as well as the
development of various radiation-lowering techniques for
favorable patient risk-to-benefit ratio (1). In contrast, CT has
an inherent limitation in soft tissue differentiation because
the pixel value or CT number entirely depends on the linear
attenuation coefficient (μ) which has considerable overlap
between different body materials. The linear attenuation
coefficient is a result of two physical interactions between
X-ray photons, i.e., the sum of photoelectric absorption that
is predominant under low energy and Compton scattering
that is predominant under high energy. Compton scattering
strongly depends on the electron density of the material.
The photoelectric effect is proportional to the cube of
the atomic number (Z) and inversely proportional to the
cube of the incident photon energy (E). Only a few heavy
atoms, such as calcium, iodine, barium, and xenon, having
https://doi.org/10.3348/kjr.2017.18.4.555
pISSN 1229-6929 · eISSN 2005-8330
Review Article | Experiment, Engineering, and Physics
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strong photoelectric effect can be easily differentiated
from other body tissue having similarly weak photoelectric
effect. In this regard, dual-energy CT, introduced as a first-
generation dual-source CT system in 2006, can improve
material differentiation by using two different X-ray energy
spectra (2). The concept of dual-energy CT was initially
described in 1973 (3) and re-emerged in the field of clinical
radiology with the recent technical developments in CT. The
technical varieties and clinical applications of dual-energy
CT are continuously expanding (4). Moreover, multi-energy
CT using the so-called photon-counting detector technology
is shedding new light on CT imaging (5). This review is
targeted to clinical radiologists with an interest in dual-
energy CT imaging, hence, the viewpoint may be slightly
different from that for CT physicists or manufacturers
involved in technical developments of dual-energy CT.
Herein, we describe the current technical options for dual-
Table 1. Current Dual-Energy CT Acquisition Methods with Technical Specifications
CT Acquisition
Methods X-Ray Tube Detectors Gantry Rotation
Time (s)
Temporal
Offset (ms)
Z Coverage
(cm)
Field of View
(cm)
Contrast per
Dose Efficiency†
Dual tubes with
or without
beam
filtration*
Two X-ray tubes
with or without
tin filter, and with
independently
selected tube
voltage pairs
Two sets
of energy
integrating
detector
0.33, 0.28,
0.25*
83, 75,
66*
1.9, 3.8,
4.8*
26, 33,
35.5* 100%
Rapid voltage
switching with
single tube
One X-ray tube
with rapidly
changing tube
voltage between
80 and 140 kVp
One set
of energy
integrating
detector
0.5 0.5 4.0 50 35%
Dual-layer
detector with
single tube
One X-ray tube
with 120 kVp
One set of
dual-layer
energy-
resolving
detector
0.27 Negligible 4.0 50 22–45%
Single tube with
split filter
One X-ray tube
with split gold/
tin filter, and with
120 kVp
One set
of energy
integrating
detector
0.28 280‡3.8 50 Not available
Single tube with
sequential dual
scans
One X-ray tube;
first at low kV,
second at high kV
One set of
energy
integrating
detector
0.27–0.28 > One scan
time 4.0–16.0 50 70%
*Two X-ray tubes and two detector arrays almost orthogonally oriented each other in dual-source CT system; three values in gantry
rotation time, temporal offset, z coverage, and field of view represent those for first, second, and third generations, respectively,
†Relative contrast-to-noise ratio per radiation dose normalized to dual-source dual-energy technique (21), ‡Temporal offset probably
caused by pitch factor limited to 0.5.
1
x x x
y
z z
y y
4
2
5
X-ray tube
High kV or
energy
Low kV or
energy
1st scan 2nd scan
3
Fig. 1. Illustration of five different methods of dual-energy
CT data acquisition. 1 = dual tubes with or without beam filtration,
2 = rapid voltage switching with single tube, 3 = dual-layer detector
with single tube, 4 = single tube with split filter, 5 = single tube with
sequential dual scans
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energy CT with their advantages and disadvantages, the
diverse spectrum of clinical applications of dual-energy CT,
and the current technical limitations and future directions
of dual-energy CT.
Current Technical Options for Dual-Energy CT
We define the technical principles underlying the
currently available dual-energy techniques to facilitate
ease of understanding and avoid conceptual confusion.
Five technical options are illustrated in Figure 1; and their
technical specifications are summarized in Table 1.
Dual Tubes with or without Beam Filtration
This method requires a dual-source CT system in which
each X-ray tube produces different X-ray energy spectra.
The most striking advantage of this method is that the
tube voltage, tube current, and filter are adjustable to
maximize dual-energy spectral contrast and radiation dose
efficiency based on the patients’ body size and diagnostic
purpose. Different combinations of tube voltages with or
without a tin filter are available for the first (80 kVp/140
kVp), second (additionally available pairs: 80 kVp/140 Sn
kVp, 100 kVp/140 Sn kVp), and third (additionally available
pairs: 70 kVp/150 Sn kVp, 80 kVp/150 Sn kVp, 90 kVp/150
Sn kVp, 100 kVp/150 Sn kVp) generations of the dual-
source CT system, with gradual increases in the magnitude
of dual-energy spectral separation from the first to the third
generation. Among these, the combination of 70 kVp and
150 kVp with a tin filter available in the third generation
dual-source CT system currently provides the highest dual-
energy spectral contrast and seems to be particularly useful
in evaluating small body parts, such as the whole body
of children and the extremities of adults and children.
However, cross-scatter radiation inevitably degrades the
dual-energy CT image quality due to the unique orthogonal
geometry between the two tube-detector pairs; moreover,
the adverse effect may not be completely eliminated
despite the use of a small portion of detector elements
to measure and correct the cross-scatter radiation. The
angular offset (approximately 90° for the first generation,
and 95° for the second and third generation) between the
two tubes results in a small temporal difference that may
be recognized as motion artifacts in and around rapidly
moving structures, such as the heart. Projection-domain
dual-energy processing is difficult to perform due to the
temporal difference between the two projection data sets;
therefore, an image-based algorithm is required for dual-
energy image reconstruction in the method. Due to the
smaller detector, the field of view (FOV) of dual-energy CT
is limited to 26, 33, or 35 cm depending on the generation
of dual-source CT system. Nonetheless, the target organ
or structure is usually within the dual-energy FOV, and the
anatomy outside the dual-energy FOV can be evaluated
because the larger detector data is available for single-
energy image reconstruction.
Rapid Voltage Switching with Single Tube
In this method, tube voltage is rapidly changed between
80 and 140 kVp, and the two projection data sets are
collected separately for subsequent use in a projection-
based dual-energy reconstruction algorithm. The rise and
fall times required for voltage modulation limit the quality
of two voltage-specific projection data; and reduced
gantry rotation speed (0.5 second or longer) is required
to allow dual-energy CT scanning. Slow gantry rotation
introduces considerable motion artifacts that nullify the
small temporal offset (0.5 ms) between the two X-ray
energy spectra. Difference in photon output between high
and low voltages is another critical problem of this method,
leading to high radiation exposure to compensate for
degraded image quality. Recently, this problem has been
addressed by increasing the low-voltage exposure time
ratio from 50 to 65%, but the dwell time ratio (65:35)
cannot be further increased without increasing the angular
mismatch between the two energy projections (6, 7).
In addition, the reduced number of projections for each
energy spectra may compromise the overall image quality.
Other disadvantages include a limited dual-energy spectral
contrast and non-availability of tube current modulation for
radiation dose reduction. A potential benefit of projection-
based algorithm, such as reduction of beam-hardening
artifact and accurate CT densitometry, is not confirmed in
this method (8, 9). Only 140 kVp images with high image
noise are available for diagnostic imaging immediately after
dual-energy scanning requiring additional reconstruction of
virtual monoenergetic imaging, e.g., 70 keV imaging; this
results in improved image quality for diagnostic imaging
despite a minor practical limitation in workflow (Fig. 2).
Dual-Layer Detector with Single Tube
In this method, the unique dual-layer energy-resolving
detector is used for dual-energy data acquisition.
Polychromatic X-ray photons are generated by one tube;
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thus, dual-energy scan is performed at a single fixed-tube
voltage, generally 120 kVp, unlike other methods using
two different tube voltages. The inner thin layer consisting
of yttrium-based scintillator absorbs low-energy photons
selectively, while the outer thick layer consisting of Gd2O2S2
absorbs high-energy photons. Temporal difference between
the dual-energy data is almost negligible. Projection-based
algorithm used in the method has potential advantage
over image-based algorithm, particularly in beam-
hardening correction at the expense of a higher noise level
for material decomposition images (7). Projection-based
method generally involves difficult calibration process,
scatter problem, and intense computation, as compared
with the image-based approach (7). As an important merit
in workflow, dual-energy evaluation can be performed
retrospectively after CT scanning in all clinical cases, but at
the expense of a relatively long dual-energy reconstruction
time. Dual-energy scanning can be performed with full
rotation speed (0.27 second) and full FOV (50 cm). However,
the dual-energy spectral contrast is lower than that of
dual tubes with beam filtration because the sensitivity
profiles of the scintillator materials between the two layers
are considerably overlapped. Disadvantages related to the
complex detector design include a lower sensitivity to
optical photons and cross-talk between the two detector
layers (7). Further clinical investigation is required to fully
define dual-energy performance.
Single Tube with Split Filter
In this method, a split filter is applied to a single X-ray
tube at 120 kVp to obtain two separated but overlapped
X-ray energy spectra (the so-called twin beam), limiting
dual-energy spectral contrast to lower levels than that
achieved by a combination of 80 and 140 kVp. The split
filter consists of 0.05-mm thick gold filter to decrease
X-ray photon energy and 0.6-mm thick tin filter to increase
X-ray photon energy. As compared with the sequential dual-
data acquisition method, this method enables dual-energy
evaluation of enhanced or moving structures by minimizing
the temporal difference between the two X-ray energy
spectra equivalent to single gantry rotation time. Pitch
factor is limited to 0.5 to maintain gapless imaging volume.
Radiation dose is almost neutral to single-energy CT;
however, greater X-ray output is necessary because the pre-
filtration absorbs approximately two-thirds of the radiation.
Single Tube with Sequential Dual Scans
In this method, dual-energy CT data with spiral or
sequential scanning are acquired simply twice sequentially
with two different tube voltages, usually 80 and 140 kVp.
Sophisticated CT hardware is not required, which may be
regarded as a merit. However, the method is greatly limited
by the greatest temporal difference between the two X-ray
energy spectra precluding many dual-energy evaluations
involved in contrast enhancement and moving body
parts. As a result, its clinical application is restricted to
unenhanced studies, such as kidney stone differentiation,
gout, and metal artifact reduction in metal implants. This
method uses an image-based dual-energy reconstruction
algorithm; and radiation-lowering technique such as tube
current modulation can be used.
Dual-Energy Applications
Dual-energy CT applications can largely be divided into
exploration of material-nonspecific and material-specific
energy-dependent information. Both evaluations can be
qualitative or quantitative. The former includes virtual
A B C
Fig. 2. Contrast-enhanced axial abdominal CT images using rapid voltage switching with single tube.
A. Image generated immediately after dual-energy scanning by using 140 kVp projections only shows high image noise. B. Virtual monoenergetic
image at 70 keV showing improved image quality needs to be additionally reconstructed for diagnostic imaging. C. Iodine map demonstrates
improved iodine contrast-to-noise ratio. Of note, patient skin, cloth, and CT table appear artifactually bright on iodine map.
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monoenergetic imaging, effective atomic map, and electron
density map. The latter includes material decomposition,
material labeling, and material highlighting.
Virtual Monoenergetic or Monochromatic Imaging
Linear or nonlinear (combining high iodine contrast and
low noise to provide optimal image contrast) blending of
dual-energy CT data is a simple approach to generate CT
images for routine diagnosis (Fig. 3A). Virtual monoenergetic
images are synthesized by decomposing two basis materials,
reconstructing the bone and water density map at the
projection domain, and combining the mass density maps
linearly at each energy; or generated simply by combining
the low and high kVp CT images linearly at the image domain
(9). Previously, uncompensated higher noise at lower keV
image hampered the optimal use of virtual monoenergetic
imaging for general contrast-enhanced CT exams (Fig. 3B)
(9). In contrast, comparable or higher iodine contrast-to-
noise ratio can be achieved in the recently introduced virtual
monoenergetic imaging techniques with energy domain
noise reduction, as compared with single-energy scan at
optimal kVp and linearly-blended techniques (Fig. 3C) (10,
11). Energy domain noise reduction technique reduces image
noise by exploiting information redundancy between low-
and high-energy images with the same anatomic details (10).
The benefits of the energy domain noise reduction technique
A B C
Fig. 3. Contrast-enhanced chest volume-rendered CT images with cropped posterior chest wall to unveil cardiovascular structures.
A, B. Compared with volume-rendered image reconstructed from linearly mixed dual-energy images with ratio of 0.8 (A), volume-rendered 40 keV
virtual monoenergetic image (B) shows further increase in cardiovascular opacification, but simultaneously increased noise compromises iodine
contrast-to-noise ratio and image quality. C. On volume-rendered noise-optimized 40 keV virtual monoenergetic image, image noise reduction
decoupled with increased iodine contrast leads to improved iodine contrast-to-noise ratio and image quality.
A B
Fig. 4. Coronal chest noise-optimized virtual monoenergetic dual-energy CT imaging.
Beam-hardening and/or photon starvation artifacts in thoracic inlet and shoulder pronounced in 40 keV image (A) are reduced in 60 keV
images (B). Because iodine contrast is progressively reduced at higher keV images, overall optimal image quality can be achieved around 60 keV
depending on patients’ size as well as body region.
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include preserved spatial resolution and synergistic effect
with spatial domain noise reduction technique, i.e.,
iterative reconstruction. The increased iodine contrast-to-
noise ratio of noise-optimized virtual monoenergetic low
keV images is useful to reduce the amount of intravenous
iodine contrast agent, salvage enhanced CT studies with
suboptimal enhancement, or increase small-vessel visibility
(Fig. 3) (12). However, beam-hardening and/or photon-
starvation artifacts at thick body regions, such as shoulder
and pelvis, are still pronounced in noise-optimized low keV
images (Fig. 4). It is difficult to compensate for photon-
starvation artifacts with high keV images. In unenhanced
brain CT, the image quality can be maximized at 65–75 keV
images, as compared with single-energy 120 kV images (13).
Materials can be qualitatively and graphically differentiated
by spectral attenuation curve as a function of energy (Fig. 5)
(9, 14). Metal artifacts can be reduced at high keV images
(95–150 keV) at the expense of loss of iodine enhancement
(Fig. 6) (9); whereas, iterative metal artifact reduction
algorithm offers greater reduction of metal artifacts with
relatively preserved iodine contrast and CT numbers than
high keV images and the combination of the two provides an
incremental benefit compared to both single methods (15).
Effective Atomic Number (Zeff) Map and Electron Density
(ρe) Map
Effective atomic number (Zeff) and electron density (ρe)
can be calculated from dual-energy CT data with small errors
of 1.7 and 4.1%, respectively (16). Effective atomic number
map is a quantitative approach in material differentiation
A B
Fig. 5. Contrast-enhanced axial chest virtual monoenergetic dual-energy CT imaging.
A. Three round regions of interest are placed in left atrium, back muscle, and subcutaneous fat in anterior chest wall, respectively, on axial chest
CT image. B. Graph illustrating changes in CT value in three regions of interest as function of energy. Iodine in blood (white line) shows higher CT
values at lower keV, while fat (orange line) reveals lower CT values at lower keV. In contrast, muscle (yellow line) demonstrates almost constant
CT values in range of 40–190 keV.
Fig. 6. Contrast-enhanced axial chest dual-energy CT imaging with posterior spinal fixation for scoliosis.
A. Linearly mixed image with ratio of 0.4 shows beam-hardening artifacts caused by pedicle screws. B. Beam-hardening artifacts become less
prominent on 130 keV image at expense of reduced iodine enhancement in vessels.
A B
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by analyzing attenuation changes as a function of energy.
CT number of water is zero irrespective of X-ray energy and
likewise, Zeff of water is approximately 7.4–7.5. Of interest,
lung perfusion defects caused by pulmonary embolism can
be distinguished from the normally perfused lung more
clearly on effective atomic number map than on iodine map
(Fig. 7).
Material Decomposition
In the three-material decomposition used for a dual-
source system in the image domain, iodine map is
generated from the iodine concentration quantified in each
voxel based on two basis materials, fat and soft tissue;
and virtual non-contrast (VNC) or unenhanced image is
produced by subtracting the iodine map from the dual-
energy enhanced CT image. In contrast, the two-material,
water and iodine, decomposition is used for a single-source
system in the projection domain.
Virtual Non-Contrast or Unenhanced Imaging and Iodine
Map
In all body regions, VNC imaging may replace a pre-
contrast scan and substantially reduce radiation exposure,
which is particularly useful in children. For example,
Chen et al. (17) reported that split-bolus dual-energy CT
urography allowed 56.4% reduction of radiation dose by
eliminating the need for a pre-contrast scan. However, the
size of calcification tends to be smaller on VNC imaging,
as compared with that on true non-contrast imaging (14,
17-20), and tiny calcifications or calcified stones may be
overlooked on VNC imaging. On the contrary, incompletely
removed iodine areas result in false positive findings
(17). Usually, CT numbers of soft tissues are slightly
overestimated on VNC imaging (14, 17, 18). The noise
levels of the VNC images as well as iodine maps are strongly
correlated with the inversion of the dual-energy ratio,
emphasizing the importance of spectral separation (7,
21). Furthermore, greater spectral separation reduces the
erroneous discrimination between iodine and calcium on
VNC images.
In the abdominal region, VNC images allow better
visualization of isodense cholesterol gallstones with
accentuated contrast chiefly due to increased attenuation
value of fat at higher tube voltage (20). Liver iron
overload may be quantified by using an iron-specific three-
material decomposition algorithm with similar diagnostic
performance to MRI (22) and, therefore, it may be used
as an alternative method when MRI is not available or
contraindicated.
Similarly, synovial hemosiderin deposits can be
identified on dual-energy CT in patients with pigmented
villonodular synovitis (23). As in intravenously enhanced
CT, VNC imaging may be applied to CT arthrography in the
musculoskeletal region (23). In the musculoskeletal region,
virtual non-calcium imaging may be used to evaluate the
bone marrow that is beyond the scope of CT evaluation.
Fig. 7. Contrast-enhanced sagittal chest dual-energy CT imaging acquired with dual-layer detector technique.
A. 70 keV image reveals subsegmental embolus (arrow) in anterior basal segment of left lower lobe. B, C. Wedge-shaped perfusion defect (arrows)
is seen on iodine map (B) and more conspicuously on effective atomic number map (C).
A B C
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Virtual non-calcium imaging subtracts calcium from
cancellous bone and allows detection of acute traumatic
bone marrow lesions including occult fractures and bone
bruises, which cannot be clearly visualized on single-energy
CT (23-25). Diagnostic performance of virtual non-calcium
imaging tends to be better for small appendicular bones
than axial skeletons. Adjustment of material decomposition
ratio (r) is necessary for different tube voltage settings
(1.45 for 80 and 140 kVp; 1.3 for 100 and 140 kVp) to
eliminate bone mineral completely. Nevertheless, virtual
non-calcium imaging is limited in evaluating bone marrow
alterations close to the cortical bone or in sclerotic bone.
In addition, false-positive results may occur due to the
presence of normal red marrow, and other pathologies,
Fig. 9. Axial brain dual-energy CT imaging in patient with recurrent primitive neuroectodermal tumor.
A-C. Linearly mixed image, iodine map, virtual non-contrast image. Larger anterior hyperdense lesion, pure intracerebral hemorrhage, shows
no enhancement on iodine map (B) and hyperdensity suggesting recent hemorrhage on virtual non-contrast image (C). In contrast, smaller
heterogenous lesion (arrows) reveals enhancing areas suggesting viable tumor on iodine map (B).
A B C
Fig. 8. Coronal abdominopelvic dual-energy CT imaging in patient with Hodgkin lymphoma.
A, B. Linearly mixed image, iodine map. Right renal artery (long arrows), left renal vein (short arrows), inferior vena cava (asterisks) are
displaced or encased by extensive, necrotic retroperitoneal lymphadenopathy. Lymphadenopathy shows subtle peripheral enhancement on iodine
map (B). Multiple hypodense small nodules are noted in spleen.
B
***
A
***
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such as osteonecrosis or degenerative changes, may mimic
post-traumatic bone marrow lesions. A combined review of
gray-scale images and color-coded images can facilitate
the identification of small attenuation changes within the
bone marrow (23-25). Of note, automatic color-coded bone
marrow segmentation is incomplete, especially in the head
region.
In the breast region, a color-coded map of unenhanced
dual-energy CT may be used to depict silicone breast
implant leaks from normal surrounding soft tissue with
similar CT numbers due to the large differences in atomic
numbers between the two (5).
Using contrast-enhanced dual energy CT data in all body
regions, iodine map specifically shows iodine distribution
in tissues with improved iodine contrast-to-noise ratio, but
the bone and calcium are also included in the map (Figs.
8, 9). In the thoracic region, dual-energy lung parenchymal
iodine or pulmonary blood volume (PBV) map, as a
surrogate of lung perfusion, is mainly used to improve the
diagnosis of pulmonary thromboembolism. By this method,
characteristically wedge-shaped iodine-deficient lung lesions
are detected, which are not apparent on conventional
pulmonary CT angiography (Fig. 7) (26-28). In pulmonary
thromboembolism, dual-phase dual-energy PBV map can be
used to differentiate between acute and chronic phases by
identifying delayed systemic collateral flow at the expense
of higher radiation dose (29). Dual-energy PBV map can
demonstrate that endothelial dysfunction represented with
hypoxic peripheral arteriolar vasoconstriction is reversible
after administration of oral sildenafil, supported by
reduced PBV coefficients of variation due to lung perfusion
heterogeneity, in smoking-associated emphysema (30).
In all body regions, due to improved lesion-to-
background contrast, enhancing lesions or vessels are more
conspicuous on the iodine map (Fig. 8). Iodine map is
helpful not only to distinguish a particularly hyperdense,
cystic lesion or hematoma from enhancing lesion, but
also to clearly delineate the extent of bowel ischemia
(14, 18, 25); in addition, malignant tumors may be more
accurately differentiated from benign tumors based on the
degree of iodine enhancement (19). Treatment response
may be assessed quantitatively by measuring the iodine
concentration in enhancing tumors in oncologic patients
(14, 18, 28, 31).
In the head region, dual-energy iodine map is useful
for differentiating between tumor bleeding and pure
intracerebral hemorrhage (Fig. 9) or between contrast
extravasation and intracerebral hemorrhage after intra-
arterial revascularization in patients with acute ischemic
stroke (32, 33).
In the cardiovascular region, static dual-energy stress
myocardial perfusion CT is more useful than coronary
CT angiography for the detection of hemodynamically
significant coronary artery stenosis by providing myocardial
iodine distribution during the early arterial phase (25, 34).
In cardioembolic stroke, dual-energy cardiac CT with dual-
phase (arterial and 3 minutes delayed) intravenous injection
of contrast agent can be used to accurately differentiate
between left atrial appendage thrombi and circulatory
stasis (35). Iodine map increases confidence in detecting
endoleaks after aortic stent graft placement (36).
Fig. 10. Axial chest xenon-inhaled dual-energy CT imaging in patient with post-infectious bronchiolitis obliterans.
A. Linearly mixed image shows bronchial wall thickenings and mosaic lung hyperlucency in right middle and lower lobes. Collapse of anterior
basal segment of right lower lobe is also noted. B. Xenon map demonstrates severely reduced xenon enhancement in right lower lobe and mildly,
heterogeneously decreased xenon enhancement in right middle lobe.
A B
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Lung Ventilation Map
Recently, xenon gas (atomic number 54) has been used
for inhalation contrast agent for dual energy lung CT (37-
47). Volumetric whole lung or dynamic focal lung scan
protocol can be used during the xenon wash-in and wash-
out periods (46). Xenon-inhaled dual-energy CT has been
applied to various pulmonary diseases including chronic
obstructive pulmonary disease (39, 47, 48), asthma (40-42),
bronchiolitis obliterans (Fig. 10) (43), and bronchial atresia
(37). Due to the anesthetic effect caused by increased
blood xenon concentration, single inspiration technique
of high-concentration xenon gas (45) or alternative use
of krypton gas having no anesthetic effect (48, 49) is
suggested. Lung density enhancement by krypton gas
inhalation is lower than that by xenon gas inhalation (48).
As in radionuclide ventilation-perfusion scans, dual-energy
CT may be used to depict ventilation-perfusion mismatch
specifically caused by pulmonary embolism (46, 49).
Material Differentiation or Labeling
In material differentiation or labeling, two materials
with different dual-energy slopes caused by different
photoelectric effects can be differentiated by using a pre-
defined separation line.
Urinary Stone Differentiation
In nephrolithiasis, dual-energy CT can be used to reliably
distinguish uric acid-containing stones from calcium-
containing stones because the former consists of materials
with significantly smaller atomic numbers than the latter
(25). Dual-energy CT also can differentiate different types
of non-uric acid calculi (50, 51).
Gout Imaging
Dual-energy CT can differentiate monosodium urate
crystals from calcium-containing compounds within joints
and periarticular soft tissues such as tendons with a
sensitivity of 87% (95% confidence interval, 0.79–0.93)
and a specificity of 84% (0.75–0.90) in a meta-analysis
(Fig. 11) (25). In gout, dual-energy CT is particularly useful
for cases with unusual locations, negative or inconclusive
arthrocentesis, and other concomitant arthropathies, and
evaluating total gout burden and treatment-response
(23). For complete evaluation, the use of a standardized
4-limb dual-energy CT is suggested. Early crystals below
threshold density for inclusion may result in false negative
interpretations, while various artifacts can result in false
positives (52).
Dual-Energy Bone Removal
Spectral information obtained with dual-energy CT can be
Fig. 11. Dual-energy CT imaging of right foot in patient with gout.
Color-coded map (A) and volume-rendered image (B) show periarticular green foci (arrows) suggesting monosodium urate deposits and
associated soft tissue swelling. False-positive artifacts are noted in typical location around nail bed and skin of great toe on volume-rendered
image (B).
A B
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used to separate iodine from bone in CT angiography (Fig.
12). For whole-body CT angiography, dual-energy automatic
bone removal offers significantly less errors in bone
segmentation than threshold-based single-energy technique
at equivalent radiation exposure and postprocessing time
(53). The small errors in dual-energy bone removal can be
minimized with greater spectral separation or higher dual-
energy ratio of dual-energy CT data. Nevertheless, the
quality of dual-energy bone removal is often inferior to that
of subtraction technique using pre- and post-contrast scans.
Lung Vessels Analysis
The lung vessels analysis tool can be used to distinguish
enhanced lung vessels from unenhanced lung vessels
based on different slopes between the two on the dual-
energy plot, irrespective of the vessel diameter (Fig. 13A)
(46, 54). As a result, dual-energy lung vessels analysis
improves the detection of small peripheral pulmonary
embolism (54). However, the ability to correctly identify
peripheral pulmonary embolism may be compromised by
insufficient pulmonary arterial enhancement (Fig. 13B).
Fig. 12. Head CT angiographic volume-rendered imaging.
A. Three-dimensional dual-energy angiographic image after automatic dual-energy bone removal shows residual bone at skull base due to
incomplete dual-energy iodine-bone separation. B. Three-dimensional dual-energy angiographic image after detailed manual bone removal
improves quality of head angiography but is time-consuming.
A B
Fig. 13. Dual-energy chest CT imaging demonstrating lung vessels analysis.
A. Axial image with lung vessels analysis shows normal enhancing pulmonary vessels in light blue in both lungs and limited dual-energy field of
view (arrows) typically seen in dual-energy technique using dual X-ray tubes. B. In patient with dextrocaria, pulmonary atresia, ventricular septal
defect, right aortic arch, and Eisenmenger syndrome, unobstructed pulmonary vessels in both lungs are red, secondary to very slow pulmonary
circulation.
A B
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Material Highlighting
Compared with single energy CT, dual-energy CT can
improve the visualization of tendons and ligaments in the
extremities by using the differences in atomic numbers
between these structures and surrounding tissues, such
as bone and fat (23, 25). As a result, abnormalities in
large tendons and ligaments, such as avulsion, atrophy,
compression, and thickening, can be better visualized.
Current Technical Limitations and Future
Directions
In most dual-energy approaches available for current CT
systems, motion artifacts and beam-hardening artifacts
may deteriorate the image quality and need to be reduced
(Fig. 14) (28, 46, 52). As in virtual monoenergetic imaging,
further noise optimization is necessary for material
decomposition images to improve their image quality in
all dual-energy approaches. Development of more versatile
dual-energy application algorithms is required to expand the
clinical utility of dual-energy CT imaging. Imperfect bone
or calcium segmentation or removal needs to be improved
(Fig. 12A). Workflow issues including difficulties in CT
scheduling, increased reconstruction time, increased number
of images, and increased interpretation time, substantially
increase the radiologists’ workload and hinder the use of
dual-energy CT imaging in daily routine (28).
In a photon-counting detector CT using cadmium-based
semiconductors (CdTe or CdZnTe), single X-ray energy
spectrum simultaneously acquired at a fixed tube voltage
can be split into more than two photon energy bins, and
is expected to be a promising future solution to overcome
limitations of current CT imaging techniques including dual-
energy scanning. In an optimized system, it has potential
to not only eliminate electronic noise and misregistration
between different energy bins, but also provide higher
contrast-to-noise ratio, higher spatial resolution, higher
radiation dose efficiency, and better spectral information
(7, 55, 56). Also, a photon-counting detector is inherently
suitable for projection-based material decomposition
requiring perfectly registered X-ray spectra as in a dual-
layer detector (7).
A photon-counting detector CT has several challenges
including pulse pile-up, charge sharing, K-escape, Compton
scattering, and charge trapping and causes non-ideal
detector responses and data overlap in energy bins and
eventually degrades the energy resolution of the CT system
(7, 56). For example, the pulse pile-up effect that occurs in
high tube currents when two or more photons are detected
as one higher-energy photon due to their proximity in
time, is regarded as one of challenges of photon-counting
detector for clinical whole-body CT scanning requiring
sufficiently high photon flux to provide high image quality.
The charge sharing effect occurs due to incorrect detection
of a photon by neighboring detector pixels at lower energy
levels and occurs predominantly at very low tube current
Fig. 14. Dual-energy pulmonary blood volume map demonstrating cardiac motion and beam-hardening artifacts.
A. On axial pulmonary blood volume map, cardiac motion artifacts appear as red color areas around heart as well as blue areas in right middle
lobe. B. On sagittal pulmonary blood volume map, beam-hardening artifacts appear as pattern of oblique stripes parallel to ribs.
A B
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settings in contrast to the pulse pile-up effect.
Initial phantom and human studies on photon-counting
detector in abdominal CT (56) and whole-body CT (57) have
been performed using a prototype built on the platform of
a second-generation dual-source CT system. This facilitates
direct comparison of imaging performance between
an energy-integrating detector and a photon-counting
detector. In these studies, the photon-counting detector
showed comparable image quality to the conventional
energy-integrating detector (56, 57). For the photon-
counting detector, the pulse pile-up effect of high photon
flux was negligible and loss of spatial resolution caused
by charge sharing and K-escape did not occur (57). In
addition, the photon-counting detector could provide multi-
energy and material-specific information (56, 57). Imaging
performance is anticipated to be continuously improved
by tailoring and optimizing calibration, artifact correction
algorithms, and dual- or multi-energy application algorithms
for photon-counting detector in the near future.
CONCLUSION
Dual-energy CT enhances the diagnostic performance and
confidence of CT by increasing iodine contrast-to-noise
ratio, decreasing metal or beam-hardening artifacts, and
providing material-specific information. In addition, patient
safety is increased by the reduction of required contrast
agent and by omitting true unenhanced CT. Radiologists
should explore the various clinical benefits of dual-energy
CT, an emerging technology in medical imaging.
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