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729

IMAGING PHYSICS
Photon-counting Detector CT:
System Design and Clinical Applica-
tions of an Emerging Technology
Shuai Leng, PhD
Michael Bruesewitz, RT Photon-counting detector (PCD) CT is an emerging technology
Shengzhen Tao, PhD that has shown tremendous progress in the last decade. Various
Kishore Rajendran, PhD types of PCD CT systems have been developed to investigate the
Ahmed F. Halaweish, PhD benefits of this technology, which include reduced electronic noise,
Norbert G. Campeau, MD increased contrast-to-noise ratio with iodinated contrast material
Joel G. Fletcher, MD and radiation dose efficiency, reduced beam-hardening and metal
Cynthia H. McCollough, PhD artifacts, extremely high spatial resolution (33 line pairs per centi-
meter), simultaneous multienergy data acquisition, and the ability
Abbreviations: ASIC = application-specific in- to image with and differentiate among multiple CT contrast agents.
tegrated circuit, EID = energy-integrating detec-
tor, PCD = photon-counting detector, UHR = PCD technology is described and compared with conventional CT
ultra-high-resolution mode detector technology. With the use of a whole-body research PCD
RadioGraphics 2019; 39:729–743 CT system as an example, PCD technology and its use for in vivo
high-spatial-resolution multienergy CT imaging is discussed. The
https://doi.org/10.1148/rg.2019180115
potential clinical applications, diagnostic benefits, and challenges
Content Codes:
associated with this technology are then discussed, and examples
From the Department of Radiology, Mayo Clinic, with phantom, animal, and patient studies are provided.
200 First St SW, Rochester, MN 55905 (S.L.,
M.B., S.T., K.R., N.G.C., J.G.F., C.H.M.); and ©
RSNA, 2019 • radiographics.rsna.org
Siemens Healthcare, Malvern, Pa (A.F.H.). Pre-
sented as an education exhibit at the 2017 RSNA
Annual Meeting. Received March 26, 2018; revi-
sion requested May 18 and received June 15; ac-
cepted June 26. For this journal-based SA-CME
activity, the authors A.F.H. and C.H.M. have Introduction
provided disclosures; all other authors, the edi-
tor, and the reviewers have disclosed no relevant Since its introduction in 1971, CT has been used widely in the di-
relationships. Address correspondence to S.L. agnostic and therapeutic medical arenas because of its fast scanning
(e-mail: leng.shuai@mayo.edu).
speed, high spatial resolution, and broad availability. In the United
Supported by the National Institutes of Health States alone, more than 80 million CT examinations are performed
(EB016966, RR018898).
every year, which makes CT one of the most important and wide-
The content is solely the responsibility of the
authors and does not necessarily represent the spread imaging modalities used for patient care (1). The wide range
official views of the National Institutes of Health. of available CT applications is attributed mainly to extensive and
The device described is a research scanner and is
not commercially available.
continual technological innovation during its history of more than 40
©
years, including improvements in CT detector technology.
RSNA, 2019
The x-ray detector is a major component of a CT scanner that is
critical to image formation and has a substantial effect on image qual-
SA-CME LEARNING OBJECTIVES ity and radiation dose. All current commercial CT scanners use solid-
After completing this journal-based SA-CME state detectors and share similar third-generation rotate-rotate designs,
activity, participants will be able to: with minor implementation and design differences according to the
■■Identify the fundamental principles
scanner model and vendor (2). Dual-energy CT systems have been
of PCDs and recognize the differences
between them and conventional energy- available commercially and have been used in routine clinical practice
integrating detectors. for approximately a decade. Different approaches have been used to
■■Describe the current status, benefits, perform dual-energy CT, such as dual-source, fast-kilovolt-switching,
and challenges of PCD CT. dual-layer detector, dual-filter, and two-consecutive-scan techniques
■■Discuss
potential clinical applications (3,4). With the acquisition of two measurements corresponding to two
of PCD CT. different spectra, dual-energy CT enables material differentiation and
See rsna.org/learning-center-rg. quantification and has generated a range of new applications beyond
the scope of conventional single-energy CT (4–6). In recent years,
photon-counting detectors (PCDs), which are capable of spectral
imaging, have been introduced as part of an experimental CT system
(7,8). PCDs possess many inherent advantages over conventional CT
detectors because of the fundamental differences in the physical mech-
anism responsible for photon detection and signal generation.
730  May-June 2019 radiographics.rsna.org

system in which in vivo imaging of human subjects

TEACHING POINTS has been performed (7,18,19,21). This provided
■■ PCDs use a direct conversion technology for x-ray detection researchers with a working system to explore the
that does not require a scintillator layer as in EIDs. The semi-
conductor detector material directly converts x-ray photons
benefit of PCD CT in clinical settings. In these
into electron hole pairs. With a bias voltage applied through- different systems, studies with phantoms, animals,
out the semiconductor, electrons travel to and are collected and human subjects have shown various benefits
by the anode to generate electronic signals. of PCD CT (Table 1), such as increased con-
■■ Because electronic noise usually is detected as a low-ampli- trast-to-noise ratio and radiation dose efficiency,
tude signal, it is interpreted by a PCD as a photon with energy reduced electronic noise and beam-hardening and
located at the lower end of a typical x-ray spectrum. Thus, by
metal artifacts, and extremely high spatial reso-
setting the low energy threshold to be slightly higher than the
energy level associated with the electronic noise signal am- lution (>33 line pairs per centimeter) (22), and
plitude (eg, 25 keV), electronic noise can be excluded readily simultaneous multienergy data acquisition, which
from the measured count data. Since a signal with an energy enables the ability to differentiate contrast agents
level lower than this threshold is very unlikely to be caused by and tissue types with a single scan.
a primary photon transmitted through the imaging object of
Given the development of PCD CT technology
interest, it typically does not contain meaningful information
vital to any clinical task. However, electronic noise can have and the promising results from recent studies, this
some effect on the detected energy spectrum, because its sig- article reviews this important emerging technol-
nal amplitude is added to that of a detected photon, which ogy and its potential effect on medical imaging
consequently artificially increases the energy of the detected and patient care. We discuss PCD technology; the
photon.
key differences between PCD and conventional
■■ PCD CT systems count each individual photon equally, re- energy-integrating detector (EID) technologies;
gardless of the measured photon energy. Hence, the low-
energy photons contribute more to the image contrast for
and the clinical applications, benefits, and chal-
PCD CT than for EID CT, which improves the image contrast lenges of PCD CT.
and contrast-to-noise ratio of iodine contrast material. The in-
creased iodine contrast can be observed for different patient Conventional and PCD Technologies
sizes, ranging from newborns to large adults, and is more pro-
nounced for higher tube potentials, such as 120 or 140 kV.
Conventional EID Technology
■■ Since PCDs use direct conversion technology, detector pixels
Conventional CT detectors are based on indirect
can be designed without a mechanical separation (septum),
which inherently improves the geometric dose efficiency. The conversion technology, which uses a layer of scin-
use of smaller PCD pixels eliminates the need for high-spatial- tillators to convert x-ray photons into visible light
resolution comb or grid filters, thereby enabling dose-efficient that is consequently detected by a photodiode and
high-spatial-resolution imaging. converted into electronic signals (Fig 1a). The out-
■■ One specific aspect of PCD CT is its ability to allow simulta- put signal of the detector is proportional to the to-
neous acquisition of high-spatial-resolution and multienergy tal energy deposited by all detected x-ray photons;
images.
therefore, this type of detector usually is referred
to as an EID. As the x-ray sources used at CT
generate polyenergetic beams, and an EID weights
Although PCDs have been used widely in the measured signal according to the energy of the
SPECT and PET, their application at x-ray detected photon, higher-energy photons generate
CT has been hampered, because the photon- stronger signals compared with lower-energy pho-
count rate used at CT is higher than the PCDs tons. In addition, because the detector integrates
used in nuclear medicine imaging can handle the energy from all detected photons, the detector
(9–14). However, the potential clinical benefits signal does not carry any information regarding
from the inherent advantages of this technology the energy of individual photons.
have continued to motivate extensive research
efforts in the past decade to bring PCDs into PCD Technology
CT practice, especially with the development of PCDs use a direct conversion technology for x-ray
higher atomic number detector materials and detection that does not require a scintillator layer
high-spatial-resolution fast application-specific as in EIDs. The semiconductor detector material
integrated circuits (ASICs). directly converts x-ray photons into electron hole
Benchtop PCD CT systems have been built pairs. With a bias voltage applied throughout the
in multiple research laboratories, and promis- semiconductor, electrons travel to and are col-
ing phantom and small-animal data have been lected by the anode to generate electronic signals
reported in the literature (15–17). In addition, two (Fig 1b).The most common semiconductor
PCD CT systems have been built on a commercial materials used in PCDs are cadmium telluride or
CT gantry to explore the potential of translat- cadmium zinc telluride, although other materials
ing this technology into clinical practice (18–20). such as silicon and gallium arsenide also have been
Of these, one is a whole-body research PCD CT used (9–12,14,23–25). In contrast to the conven-
RG  •  Volume 39  Number 3 Leng et al  731

Table 1: Potential Benefits and Clinical Applications of PCD CT

Features and Benefits of

PCD CT Potential Clinical Applications
High spatial resolution Temporal bone imaging
Musculoskeletal imaging
Lung imaging
Cardiovascular imaging (stent imaging)
Simultaneous multienergy Material decomposition
acquisition Plaque removal
Bone removal
Virtual noncontrast imaging
Virtual noncalcium imaging
Virtual monoenergetic synthesis
Energy binning
  Low electronic noise Imaging large patients
Low-radiation-dose examinations
  K-edge imaging Dual-contrast imaging*
  Metal artifact reduction Metal implant imaging (musculoskeletal,
neurologic)
Uniform photon weighting Contrast material–enhanced CT (ab-
and improved contrast-to- dominal, neurologic, cardiovascular)
noise ratio with iodinated
contrast material
*Use is limited to research in animals.

information about each individually detected pho-

ton. The output signal from a PCD is processed
by multiple electronic comparators and counters,
where the number of comparators and counters
depends on the electronic design of the PCD and
its ASICs. Each detected signal is compared with a
voltage that has been calibrated to reflect a speci-
fied photon energy level, referred to as an energy
threshold. When the energy level of a detected
photon exceeds an energy threshold associated
with a counter, the photon count is increased by
one. In this manner, the number of photons that
have energy equal to or greater than a specified
energy level is measured. This process is enabled
by the very fast ASIC, a key element in PCDs.
The detector absorption efficiency depends
on the detector material used and its thickness.
High-atomic-number sensor materials such as
cadmium telluride and cadmium zinc telluride
have higher absorption efficiency per unit thick-
ness and are the most common detector materials
Figure 1.  Schematic drawings illustrate the principles of
EID (a) and PCD (b) technologies.
in PCDs (25). In general, the absorption effi-
ciency is comparable between EIDs and PCDs
(8); however, it may vary, depending on the
tional EIDs, which integrate the energy levels of specific design of each detector.
all detected photons, PCDs count the number of A detailed comparison between conventional
individual photons that exceed a specified energy EIDs and PCDs is given in Table 2.
level. For a given x-ray photon, the pulse height of
the signal created by the charge deposition at the PCD Datasets
anode is proportional to the energy of the photon. The number of energy thresholds of a PCD
Thus, the signal from PCDs carries with it energy depends on the ASIC design, usually between
732  May-June 2019 radiographics.rsna.org

Table 2: Comparisons between Conventional EIDs and PCDs

Properties and
Characteristics EIDs PCDs
Detector material Cadmium tungstate, gadolinium oxide, Cadmium telluride, cadmium zinc telluride, sili-
gadolinium oxysulfide con, gallium arsenide, chromium compensated
gallium arsenide
Detection mecha- Energy integrating with a two-step Energy resolving with the use of a single-step pro-
nism process involving an x-ray scintilla- cess involving a semiconductor and the direct
tor and a photodiode, which converts conversion of x-rays to electric signal
x-rays to visible light and visible light
to electric signal
Spectral abilities Inherently no x-ray energy-resolving ca- Photon events are counted and binned in digital
pabilities owing to charge integration counters with user-defined energy thresholds.
Energy-resolving X-rays converted to visible photons and X-ray photon interactions form charge clouds
mechanism visible photons converted to electron- (electron hole pairs) in the semiconducting
ic signal; photon energy information layer, creating a signal whose magnitude is pro-
is lost during this process portional to the photon energy.
Spatial resolution Smaller detector pixels become dose inef- Smaller detector pixel sizes are possible because
ficient owing to the finite-width septum septa between detector pixels are not needed.
required between detector pixels.
Electronic noise Noticeable on conventional CT images Can be excluded from the measured signal by
properties at low doses or at scanning of obese selecting an energy threshold higher than the
patients electronic noise floor
Photon weighting High-energy x-rays receive more All photon energy levels receive uniform weight-
weighting than do low-energy x-rays, ing, allowing improved contrast between soft
consequently deteriorating contrast tissue and iodinated contrast material.
between soft tissue and iodinated
contrast material.
Reduction of Metal artifact reduction algorithms can The high-energy bin images are much less af-
beam-hardening be used to mitigate these artifacts. fected by these artifacts. Algorithms can be
and metal arti- used to reduce any residual artifacts.
facts
Multienergy imag- Requires dual-source, dual-tube poten- Single-source, single-tube-potential, single
ing tials, dual acquisitions, dual detector- acquisition, single–detector layer, single-filter
layers, or dual beam filters to acquire simultaneous multienergy acquisition is inher-
the needed dual-energy data ently possible.
High-resolution Radiation dose inefficient owing to Radiation dose efficient high-spatial-resolution
imaging comb or grid filters or decreased imaging is possible owing to inherently smaller
detector fill factor due to requirement detector pixels.
of additional septa
Energy-selective Limited options owing to lack of energy Energy binning allows K-edge imaging custom-
imaging discrimination ized to gadolinium, gold, bismuth, ytterbium,
and other high-Z contrast agents.

two and eight thresholds. Users must set up the datasets, with photon energy ranges of 25–140
energy thresholds (in kiloelectron volts) before keV and 65–140 keV, respectively. Two energy bin
data acquisition. The output data at each energy datasets are then generated by the subtraction of
threshold represent the number of photons with the two threshold data, which gives 25–65 keV
energy equal to or higher than the threshold, and 65–140 keV. Note that the highest energy
up to the maximum photon energy determined bin data (ie, 65–140 keV) cover the same en-
according to the tube potential (in kilovolts). ergy range as the highest energy threshold data
Subtraction of the counts associated with two (65–140 keV), making them identical datasets.
different energy thresholds creates energy bin
data, which represent the number of photons PCD CT
between two different energy thresholds (Fig 2). Micro CT and preclinical CT scanners that
For example, a PCD CT scan performed at are capable of phantom and small-animal
140 kV with energy thresholds of 25 and 65 keV studies have been developed and evaluated
facilitates the creation of two energy-threshold (15–17,26–33). With smaller detector sizes
RG  •  Volume 39  Number 3 Leng et al  733

could be optimized further for PCDs to improve

system performance.
The PCD technique provides great flexibility
in adjusting imaging resolution and the number
of energy thresholds by combining detector pix-
els and using different pairings of energy thresh-
olds. For example, on the current whole-body
research PCD CT system, there are four data
acquisition modes (Fig 3b): macro mode, chess
mode, UHR mode, and sharp mode, depending
on the energy-threshold and pixel configuration
(8,36).
In the macro mode, the 4 3 4 neighboring sub-
pixels (native detector elements) are grouped into
one macro pixel, resulting in an effective pixel size
of 0.5 mm 3 0.5 mm at the isocenter. The same
two energy thresholds are applied to all subpixels,
resulting in a total of two energy bin datasets.
Figure 2.  PCD datasets for the x-ray tube potential or energy
In the chess mode, the energy thresholds as-
maximum (EM) and two energy thresholds (low-energy [EL] and
high-energy [EH] thresholds): the two threshold datasets (low signed to each subpixel in a macro pixel follow a
threshold [TL] and high threshold [TH]) correspond to photons chessboard pattern, effectively resulting in four
with energy levels higher than the respective thresholds (EL, EH) energy thresholds. The chess mode is intended
but lower than the tube potential (EM), and two bin data (Bin 1
as a convenient way to provide four threshold
and Bin 2) corresponding to photons with energy levels between
(arrow between energy levels) the two energy thresholds (Bin 1) datasets with ASICs that offer only two energy
or between the high-energy threshold and the tube potential thresholds for each pixel. However, it is radia-
(Bin 2). Note that high-threshold and Bin 2 data are identical. tion dose inefficient, because only one-half of
(Adapted and reprinted, with permission, from reference 8.)
the delivered photons are used to create any
given threshold dataset.
In UHR mode, the detector elements are
and a limited field of view, these systems have grouped into a 2 3 2 pattern (instead of the 4 3
enabled investigators to understand the basic 4 pattern as in macro and chess modes), resulting
principles governing the technology and the in an effective pixel size of 0.25 mm 3 0.25 mm
underlying physics. A preclinical PCD scanner at the isocenter. The UHR mode was imple-
using cadmium telluride was built on the basis mented to take advantage of the smaller subpixel
of a clinical CT system gantry (Brilliance iCT; size and to meet the clinical need for high spatial
Philips Healthcare, Haifa, Israel), with a 16.8- resolution in lung, vascular, musculoskeletal,
cm in-plane field of view and 2.5-mm longitu- and temporal bone applications. In the current
dinal coverage. Results of phantom and animal implementation, only two energy thresholds are
studies (20,34,35) performed with this system enabled for the UHR mode, which is similar to
have been reported, including applications such the macro mode.
as dual–contrast agent and K-edge imaging. In the sharp mode, the low-energy-threshold
In 2010, a whole-body research PCD CT sys- data are acquired with the same pixel dimension
tem was introduced (Somatom Count; Siemens as that in the UHR mode, while the high-energy-
Healthineers, Forchheim, Germany) that was threshold data are acquired with the pixel size of
the first PCD CT system capable of performing the macro mode. With the sharp mode, the low-
studies in human subjects (8,19). This system energy-threshold dataset permits high-spatial-
was built with a second-generation dual-source resolution anatomic imaging. Recall that the low-
CT platform in which one source was coupled threshold data are generated with the use of all
to an EID and the other to a PCD (Fig 3a). The available photons, which, therefore, have the low-
PCD subsystem allows for high photon flux val- est noise and can tolerate a small pixel size. On
ues, with tube current up to 550 mAs, which is the other hand, high-threshold data are generated
sufficient for most clinical examinations (8). The from relatively fewer photons, and therefore, have
ASICs associated with each subpixel allow for stronger noise. Since the high-threshold data are
two energy thresholds (low and high thresholds). only required for dual-energy processing, which
In the current implementation, the x-ray tube is susceptible to noise amplification, higher spa-
coupled with the PCD subsystem is the same as tial resolution is less desirable and the larger pixel
that of the EID subsystem. However, it is pos- size for high-threshold data can help to improve
sible that the x-ray source and beam filtration the signal-to-noise ratio.
734  May-June 2019 radiographics.rsna.org

Figure 3.  A whole-body research PCD

CT scanner capable of human imaging at
a clinical dose rate was built on the basis
of a dual-source CT system with one of
the EID arrays replaced with a cadmium
telluride–based PCD array (a). The PCD
CT system has four different acquisition
modes (macro, chess, ultra-high-resolu-
tion [UHR], and sharp modes), each cor-
responding to a specific detector configu-
ration (b). (Fig 3a adapted and reprinted,
with permission, from reference 8.)

Clinical Benefits and tum noise can be traced back to the random
Applications of PCD CT nature of x-ray photon interactions and is de-
In this section, we discuss in detail the benefits of termined by the number of photons detected.
the PCD CT system and potential clinical appli- The electronic noise, on the other hand, mainly
cations, with sample cases associated with each of originates from the analog electronic circuits in
the applications. A general summary of the clini- the x-ray detection system and is independent
cal applications of PCD CT is found in Table 1. of the number of detected photons. The relative
effect of these two noise sources is determined
Reduced Electronic Noise by the incident photon flux. In the realm of high
Noise on CT images originates from two sources: photon flux, the quantum noise dominates the
quantum noise and electronic noise. The quan- total noise magnitude, and the effect of electronic
RG  •  Volume 39  Number 3 Leng et al  735

Figure 4.  The shoulder section of a thorax phantom reconstructed from data acquired with EID CT (a) and with PCD CT (b)
using the same x-ray tube potential and radiation dose. Compared with the image acquired with EID CT, the PCD CT image
has noticeably fewer horizontal streaking artifacts and an overall more uniform appearance, which indicates that electronic
noise has a more noticeable effect on the EID image than on the PCD image. (Reprinted, with permission, from reference 39.)

noise is negligible. However, when the number from a PCD CT system, which have been shown
of detected photons is low, the magnitude of the to be more immune to the effect of electronic
electronic noise can be similar to or higher than noise (Fig 4) (39). This rejection of electronic
that of the quantum noise. With modern EID CT noise in a PCD CT system can be used to
systems, the electronic noise is usually negligible reduce overall image noise and improve image
for protocols in which clinical radiation dose lev- diagnostic quality for low-dose examinations.
els are used in average-sized patients. However, For low-dose lung cancer screening, PCD CT
for low-dose scans or in morbidly obese patients, was shown to provide better CT number stabil-
electronic noise can degrade image quality, ity and scan reproducibility, as well as reduced
particularly along highly attenuating path lengths image noise when compared with EID CT (21).
(eg, through the shoulders) (Fig 4) (37,38). Alternatively, a lower radiation dose can be used
Because electronic noise usually is detected as with a PCD system to reach the same noise
a low-amplitude signal, it is interpreted by a PCD level as that achieved with an EID CT system
as a photon with energy located at the lower end and a higher dose level, thereby improving dose
of a typical x-ray spectrum. Thus, by setting the efficiency.
low energy threshold to be slightly higher than
the energy level associated with the electronic Improved Iodine Contrast-to-Noise Ratio
noise signal amplitude (eg, 25 keV), electronic and Radiation Dose Efficiency
noise can be excluded readily from the measured In the diagnostic x-ray energy range, the x-ray
count data (10), as demonstrated in Figure 5. photons primarily interact with the scanned
Since a signal with an energy level lower than object through two physical mechanisms: the
this threshold is very unlikely to be caused by a photoelectric effect (μPE) and Compton scatter-
primary photon transmitted through the imag- ing. The attenuation due to photoelectric effect is
ing object of interest, it typically does not contain proportional to the effective atomic number (Z)
meaningful information vital to any clinical task. of a material and is inversely proportional to the
However, electronic noise can have some effect energy (E) of incident x-ray photons:
on the detected energy spectrum, because its
signal amplitude is added to that of a detected µ PE ∝ Z 3 / E 3,
photon, which consequently artificially increases while the Compton scattering effect is relatively
the energy of the detected photon. independent of material composition, and its
The intrinsic advantage of a PCD system energy dependence is relatively flat at a diagnos-
for excluding electronic noise can be beneficial tic energy range. Photoelectric effect contributes
particularly for examinations with low detector more to the attenuation of low-energy photons,
signal intensity, such as those in morbidly obese while Compton scattering is the dominant effect
patients and those performed with a low radia- that accounts for high-energy photons. Conse-
tion dose (39). In these scenarios, the electronic quently, high-Z materials such as iodine allow
noise can reduce image uniformity and cause better contrast on CT images because of stron-
noticeable streak artifacts on images acquired ger photon attenuation in the low-energy range
with an EID CT system compared with images through the photoelectric effect.
736  May-June 2019 radiographics.rsna.org

For a conventional EID CT system, the detec-

tor signal is proportional to the total energy of all
detected x-ray photons. Therefore, lower-energy
photons contribute relatively less to the detector
signal than do higher-energy photons, which have
less information content from high-Z materials
such as iodine. Consequently, the underweight-
ing of signal produced by lower-energy photons
reduces the contrast-to-noise ratio of iodine
signal in an EID CT system. However, PCD CT
systems count each individual photon equally,
regardless of the measured photon energy.
Hence, the low-energy photons contribute more
to the image contrast for PCD CT than for EID
CT, which improves the image contrast and Figure 5.  Graph shows signals of individual x-ray photons
contrast-to-noise ratio of iodine contrast material detected with the use of a PCD, with additive electronic noise.
(9–13,40–48). The increased iodine contrast can The PCD is able to discriminate the energy of each incident
x-ray photon. Since the electronic noise usually is detected as
be observed for different patient sizes, ranging a low-amplitude signal, it can be excluded readily from the
from newborns to large adults, and is more pro- measured counting data by setting a proper counting energy
nounced for higher tube potentials, such as 120 threshold to be slightly higher than the energy level associated
or 140 kV (19). Higher tube potentials are espe- with the electronic noise amplitude.
cially relevant in imaging of moderately sized to
large patients. The improved iodine contrast on
a PCD CT system results in an improved iodine Since each individual photon is sorted accord-
contrast-to-noise ratio if radiation dose is main- ing to its energy level at PCD CT, an energy bin
tained. Alternatively, if the same iodine contrast- image can be reconstructed with the use of only
to-noise ratio is desired, the PCD CT system higher-energy photons (51). Compared with the
acquisition can be altered to reduce either the conventional EID CT image or the low-energy-
radiation dose or the volume of iodine contrast threshold image in a PCD acquisition, the high-
agent compared with those with conventional energy-bin image is more immune to beam-hard-
EID CT. The contrast-to-noise ratio for a given ening effects in areas around dense bones, such as
material can be increased further by designing the area around the posterior fossa (Fig 6). The
an optimal weighting scheme for the differ- high-energy-bin image also demonstrates reduced
ent narrow energy bin datasets available from a calcium blooming, which typically is observed
single PCD scan (49,50). Such schemes typi- around the interface between the calvarium and
cally weight each energy bin dataset according to the brain (Fig 6) (8,19). However, the use of high-
its image contrast level and noise variance. The energy signal to reduce beam-hardening effects is
weighted energy bin images are then combined at the cost of increased image noise and reduced
to yield a final image with an improved contrast- radiation dose efficiency, because the lower energy
to-noise ratio. The amount of improvement in photons are not used in the generation of the
contrast-to-noise ratio depends on the object high-energy-bin image. To improve dose effi-
size and material, the x-ray tube potential, the ciency, the incident x-ray beam can be hardened
PCD energy threshold settings, and the energy further with additional filtration such as that with
response of the EID detector. a tin filter, which increases the relative percentage
of photons in the high-energy bin. The combina-
Reduction of Beam-hardening tion of using the high-energy-bin image and tin
and Metal Artifacts beam filtration can further reduce metal artifacts
As photons pass through the scanned object, and image noise, thus providing improved delin-
low-energy photons are preferentially attenuated eation for tissue regions that would otherwise be
compared with high-energy photons. Since poly- affected by prominent metal artifacts (Fig 7) (52).
energetic beams are used in CT, this causes the Unlike true monochromatic x-ray imaging (such
effective photon energy to be shifted toward the as that based on a synchrotron source) (53), the
higher end of the spectrum (a phenomenon known use of the PCD high-energy bin and additional
as beam hardening). This introduces artifacts, filtration can reduce beam-hardening and metal
which are typically dark areas adjacent to highly artifacts but cannot completely eliminate them,
attenuating objects such as cortical bone and metal because the quasi-monochromatic narrow energy
implants, affecting the image appearance and CT bin images are reconstructed from a polychro-
number accuracy for nearby soft tissues. matic x-ray spectrum.
RG  •  Volume 39  Number 3 Leng et al  737

Figure 6.  Artifacts on EID and low- and high-energy PCD im-
ages. Axial EID (a), low-energy-threshold PCD (b), and high-
energy-threshold PCD (c) CT images in a cadaver head (left im-
ages) and magnified images (right images) of the areas outlined
in red squares show that the high-energy-threshold PCD CT
images (c) have fewer blooming artifacts (arrows) and beam-
hardening artifacts (arrowheads) than do the EID CT images (a)
and low-energy-threshold PCD CT images (b). (Reprinted, with
permission, from reference 19.)

detector, the x-rays blocked by the filter have al-

ready passed through the patient and contributed
to the radiation dose but have not been detected
by the detector; therefore, the use of this filter
reduces radiation dose efficiency.
Although high-spatial-resolution CT is used
currently in clinical practice for specific tasks
such as temporal bone and extremity imaging,
its use throughout the body has been restricted
because of the dose inefficiency of this approach.
Since PCDs use direct conversion technology,
detector pixels can be designed without a me-
chanical separation (septum), which inherently
High-Spatial-Resolution CT improves the geometric dose efficiency. The
Conventional EIDs require a finite-width reflec- use of smaller PCD pixels eliminates the need
tive layer (septum) between detector pixels to for high-spatial-resolution comb or grid filters,
avoid signal (visible light) leakage to neighboring thereby enabling dose-efficient high-spatial-
detector pixels. Designing smaller detector pixels resolution imaging. Typical detector pixel sizes of
leads to an increase in the relative area of the 55–1000 mm (pixel pitch) have been reported in
septum compared with the scintillator material the literature (8,55,56).
and a corresponding decrease in the fill factor, The sharp and UHR modes on the whole-
which consequently reduces the geometric dose body PCD CT system have an effective detector
efficiency. Another approach to high-spatial-reso- pixel size of 0.25 mm 3 0.25 mm. The small
lution imaging is to place a dedicated attenuating detector pixel size allows spatial resolution to
filter (in the form of a comb or grid) in front of be limited to 150 μm (22). Numerous clinical
the detector array to reduce the detector pixel applications may benefit from this improved
aperture, consequently reducing the effective spatial resolution, including lung, temporal bone,
detector pixel size (54). Because the high-spatial- musculoskeletal, and vascular imaging. Figure
resolution filter is placed directly in front of the 8 shows sample images of lung and wrist CT
738  May-June 2019 radiographics.rsna.org

Figure 7.  Reduction of metal

artifacts at PCD CT. Axial EID (a)
and PCD (b) CT images of a fused
lumbar spine in a 55-year-old
man show that the reduction of
metal artifacts arising from the
posterior metallic hardware en-
ables visualization of the spinal
canal, which was obscured by the
metal artifacts on a.

Figure 8.  CT imaging in UHR

mode. In vivo axial CT image of the
lung in a 69-year-old woman (a)
and coronal CT image of the
wrist in a 59-year-old man (b)
show that high spatial resolution
enables accurate delineation of
lung fissures (arrow in a), airway
walls (arrowhead in a), and tra-
becular bone.

examinations performed in UHR mode. For the

lung image, the benefit of high spatial resolution
is shown with accurate delineation of the lung
fissures and airway walls. In the wrist image, the
high spatial resolution enables excellent visual-
ization of the trabecular bone structures. Figure
9 shows sample EID and PCD CT images of a
coronary stent. Individual struts are clearly de-
lineated on the PCD CT images because of their
high spatial resolution. Example EID images of
the temporal bone acquired with high-spatial-
resolution comb filters and PCD images acquired
with UHR mode are shown in Figure 10. Note
the higher conspicuity of the stapes superstruc-
ture on the PCD CT image compared with that
on the EID CT image.
Compared with EID CT images acquired
with UHR mode and comb or grid filters, PCD Figure 9.  Coronal EID (a) and PCD CT images acquired in UHR
mode (b) show examples of a coronary stent. The PCD image
CT images acquired with high spatial resolution
clearly shows the individual struts, which are evident on the cor-
show lower image noise at the same radiation responding three-dimensional rendering (c) of the PCD CT image.
dose because of improved dose efficiency. Phan-
tom and cadaveric studies have shown a 29%
reduction in noise with PCD CT high-spatial- PCD CT with high spatial resolution and EID
resolution acquisition compared with EID CT CT with high spatial resolution.
high-spatial-resolution acquisition at matched
radiation dose levels (Fig 10). This translates to Simultaneous Multienergy Acquisition
a 50% reduction in radiation dose with PCD One of the main driving forces of the develop-
CT to achieve the same noise levels between ment of PCD CT is that it allows the acquisition
RG  •  Volume 39  Number 3 Leng et al  739

Figure 10.  Axial EID image (a) and PCD CT image acquired in UHR mode (b) of a cadaveric
temporal bone specimen. The conspicuity of the stapes superstructure (arrows) is higher on
the PCD CT image, and noise reduction of 29% was achieved with PCD CT at matched dose
levels. (Reprinted, with permission, from reference 66.)

Figure 11.  Sagittal single-energy CT angiogram (a) and iodine map (b) after material decomposition
acquired with a whole-body PCD CT system in a living swine.

of simultaneous multienergy CT images. PCD the best image quality or lowest radiation dose.
CT inherently allows dual-energy or multie- Also, the K-edge imaging ability of PCD CT
nergy (>2) acquisitions at a single x-ray tube has provided the opportunity for development of
potential owing to its energy-discriminating abil- new types of contrast agents and imaging tech-
ity. This unique feature enables single-source, niques such as nanoparticle-based blood pool
single-tube-potential, single-acquisition, single imaging (57,58) or targeted imaging (29,59,60)
detector-layer, and single-filter multienergy CT with multiple-contrast-agent multiphase im-
imaging with perfect temporal and spatial regis- aging (61–63). However, these are primarily
tration in the acquired multienergy data, thereby research methods that are yet to be translated to
eliminating many sources of artifacts. The mainstream clinical practice.
user-defined energy threshold selection provides Multienergy PCD CT provides opportunities
the freedom to select the correct energy thresh- for advanced data processing, such as creation
olds tailored to the specific diagnostic task. For of virtual monoenergetic images, virtual non-
example, by selecting energy thresholds lower contrast images, and images with automated
and higher than the K edge of a specific contrast bone removal. Figure 11 shows sample CT
agent, PCD CT may enable K-edge imaging. angiograms in a living swine with the PCD CT
This task-driven energy-threshold selection system, both a conventional single-energy image
helps resolve different tissue types or contrast (Fig. 11a) and an iodine map image (Fig. 11b)
media with optimal imaging settings to achieve after material decomposition. PCD CT has been
740  May-June 2019 radiographics.rsna.org

Figure 12.  Sample PCD CT images, which

were acquired with simultaneous high-spatial-
resolution and multienergy capabilities, of the
pelvis in a 72-year-old man. The high-spatial-
resolution low-energy-threshold image (a) is
used as a standard single-energy image for rou-
tine diagnosis, while dual-energy images can be
processed for bone removal (b) or ruling out
the presence of gout (c).

shown to provide accurate iodine quantifica- tions, high-spatial-resolution images with 0.25-
tion (root mean square error of 0.5 mg of iodine mm image thickness demonstrated previously
per milliliter) for concentrations of 2–20 mg of unobservable morphologic features, such as a thin
iodine per milliliter at different phantom sizes highly attenuating outer shell and internal tex-
(64). In addition, accurate CT numbers across ture information (Fig 13a), while the dual-energy
different object sizes can be achieved on virtual analysis allowed determination of stone composi-
monoenergetic images, with a percentage of er- tion (Fig 13b, 13c).
ror of 8.9% (64).
One specific aspect of PCD CT is its ability Challenges
to allow simultaneous acquisition of high-spatial- The PCD technique is not without limitations.
resolution and multienergy images. Current There are still several technical challenges that
dual-energy EID CT systems can be used to scan must be addressed to realize the full potential of
in either dual-energy mode or a UHR mode, but PCD CT. These include photon flux–indepen-
not both simultaneously. This creates a dilemma at dent effects such as charge sharing, charge trap-
clinical examinations in which both high-spatial- ping, and k fluorescence escape, as well as photon
resolution and multienergy imaging are needed, flux–dependent effects such as pulse pile-up (13).
such as musculoskeletal examinations. Figure 12 These nonideal effects can degrade the spatial
shows sample images of the pelvis acquired with a resolution and energy-resolving ability of PCD,
PCD CT system. The low-energy-threshold image leading to compromised system performance.
shows clear delineation of fine bony structures and We refer interested readers to technical literature
is used for standard diagnostic tasks (Fig 12a). (13,30,65) for detailed discussions on these ef-
With the use of low- and high-energy data that are fects. In addition, the current PCD CT research
simultaneously acquired, dual-energy postprocess- systems have a limited field of view; however, this
ing can be performed to remove bony anatomy is not a fundamental limitation of the PCD tech-
from images or to evaluate for the presence of gout nology and can be addressed readily by increasing
(Fig 12b, 12c). Similarly, in renal stone applica- the detector units mounted onto the scanner.
RG  •  Volume 39  Number 3 Leng et al  741

Figure 13.  Renal stones in a 64-year-old man. (a) High-resolution single-energy PCD CT image
shows the morphologic features of the stones. (b) Dual-energy image shows the stone composi-
tion. (c) Cinematic three-dimensional volume-rendered image enhances the visualization of stone
morphology and composition.

Figure 14.  Axial reconstructed images in a swine. (a, b) Low-threshold CT image (corresponding to 20–140 keV) (a) and energy bin
CT image (20–25 keV) (b) reconstructed with standard filtered back projection. Note the higher noise in the energy bin image (b)
compared with that in the low-threshold image (a). (c) Energy bin CT image reconstructed with the spectral prior image constrained
compressed sensing technique from the same dataset as that for b shows reduced image noise and preserved spatial and spectral
information. (Reprinted, with permission, from reference 70.)

For a PCD CT acquisition, the transmitted x- tial and spectral resolution for multienergy PCD
ray spectrum is divided into a number of different datasets (Fig 14) (70).
energy bins, and the number of photons within
any energy bin is less than the total number of Conclusion
photons reaching the detector. Reduced photon PCD CT is an emerging technology that, com-
counts can substantially increase the noise on the pared with conventional EID detector technol-
energy bin images (4). In addition, the reduced ogy, has multiple advantages owing to its unique
detector size for UHR mode PCD CT also leads interaction physics, especially its dose-efficient,
to increased image noise, because the number of high-spatial-resolution, and energy-discrimination
photons hitting each detector cell is decreased abilities. In addition to multiple preclinical sys-
(66). To address these challenges, various noise tems, a whole-body research PCD CT system has
reduction and iterative image reconstruction been used to demonstrate a number of clinical
algorithms have been developed (67–72). Iterative benefits in human subjects. Large-scale produc-
methods involving statistical and physical models tion of high-quality PCDs at an affordable price
for PCD-based x-ray measurements have allowed remains the primary limit to commercialization of
reconstruction of low-noise PCD CT images, this technology. However, research into and devel-
reduced spectral distortion, and direct material opment of PCD technology are sure to continue
reconstruction abilities (73–77). In addition to the so that the theoretical and demonstrated benefits
conventional CT denoising strategies, techniques of PCDs can be translated into clinical practice.
that exploit data redundancy that exist within a
PCD dataset have been developed and can allow Disclosures of Conflicts of Interest.—A.F.H. Activities related to
reduction in image noise while preserving the spa- the present article: disclosed no relevant relationships. Activities not
742  May-June 2019 radiographics.rsna.org

related to the present article: employed by Siemens Healthineers. 21. Symons R, Cork TE, Sahbaee P, et al. Low-dose lung cancer
Other activities: disclosed no relevant relationships. C.H.M. Ac- screening with photon-counting CT: a feasibility study. Phys
tivities related to the present article: disclosed no relevant relation- Med Biol 2017;62(1):202–213.
ships. Activities not related to the present article: royalties from Bayer 22. Leng S, Rajendran K, Gong H, et al. 150-μm spatial resolu-
Healthcare, grants/grants pending from Siemens Healthineers, tion using photon-counting detector computed tomography
intellectual property rights for patents owned by Mayo Clinic. technology: technical performance and first patient images.
Other activities: disclosed no relevant relationships. Invest Radiol. 2018 Nov;53(11):655–662.
23. Xu C, Chen H, Persson M, et al. Energy resolution of a
segmented silicon strip detector for photon-counting spectral
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TM
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