Dr. Seibert
is a Professor and Assistant Chair of Informatics, Department of
Radiology, University of California Davis Medical Center,
Sacramento, CA.
Digital radiography detector systems were first implemented for
medical applications in the mid 1980s, but the promise of
digitalimaging was not realized until the early 1990s, in
conjunction with the establishment of first generation picture
archiving and communications systems (PACS). Atthe time, there was
only one technology available to replace the analog screen-film
detector-a cassette-based, passive acquisition photostimulable
storage phosphor (PSP) and plate reader system, known as "computed
radiography" (CR). This system closely emulated the screen-film
paradigm. Alternate technologies for digital image acquisition
appeared in the mid 1990s with the use of large field of view (FOV)
X-ray phosphors and optical lens assemblies to focus the X-ray
induced light output onto a small-area charge coupled device (CCD)
photodetector array, as well as rectangular CCD arrays used with
slot-scan geometries. Active-matrix flat-panel imager (AMFPI)
systems appeared in the late 1990s and early 2000s, employing
either an X-ray-to-light converter with photodiodearray, or a
semiconductor material to directly convert incident X-rays into
signals. Both CCD and flat-panel based detectors use an
"active"readout of the image following acquisition to present the
image immediately without further interaction by the technologist.
Other technologies such as complementary metal-oxide semiconductor
(CMOS) detectors were introduced in the same time frame as AMFPI's,
but have not as of yet been successful, mainly due to problems with
excessive electronic noise.
Beyond the digital detector characteristics are considerations
for software for pre- and postprocessing of the digital image data,
the user and modality interfaces, display monitors and
calibrations. Many unique acquisition capabilities, such as
dual-energy image tissue decomposition and limited-angle digital
tomosynthesis, are important when considering future applications
of specific importance.
The contents and discussion within this article are based on a
"generic" description of PSP (CR) and DR detectors. Suffice it to
say, there are a wide range and capability of detector systems
within each "class" of digital detector technology, and many of the
sweeping statements made may or may not be accurate with respect to
a specific digital radiography device.
Digital radiography technology
Photostimulable storage phosphor (PSP) detectors
More commonly recognized as CR, the PSP detector "system" is
comprised of 2 main components. The detector is usually a
cassette-based storage phosphor that absorbs X-ray energy
transmitted through the patient and temporarily stores the X-ray
latent image as a 2-dimensional array of electrons trapped in
semistable energy wells. The imaging plate "reader" is comprised of
a scanning laser beam to stimulate the trapped electrons and
produce "photostimulated luminescence" of a different wavelength
that is optically separable from the stimulation wavelength. The
reader also includes a light guide and photomultiplier assembly
that extracts and processes the stored X-ray content to a sampling
resolution on the order of 100 microns (0.1 µm), digital
electronics to create the corresponding digital image, and an
erasure stage to eliminate any residual signal and prepare for the
next exposure. From the reader, all images proceed to a quality
control workstation for image evaluation, annotation and transfer
to PACS (Figure 1). Most often, the storage phosphor is layered on
a flexible or solid substrate in a cassette enclosure, which allows
for the ability to directly replace a screen-film cassette in a
conventional radiography room. Thus there is the flexibility and
portability of a cassette with digital radiography acquisition
capability using existing X-ray equipment; this is CR's greatest
asset. CR cassettes of various size and number, together with a
high-speed imaging plate reader can service multiple X-ray rooms,
resulting in a relatively low initial acquisition cost. However,
the technologist must handle the cassettes and process the imaging
plates in a manner similar to film, which can reduce patient
throughput in a busy clinical room and increase labor costs, as the
time to handle the exposed imaging plate to the reader and output
of the X-ray image can often take several minutes (about 45 to 60
seconds to "read" the plate with the moving laser beam). Other
expenses to consider are the need for high-frequency (e.g., 170
lines per inch) antiscatter grids for stationary (non-Bucky)
applications such as portable bedside imaging and fixed grid
cassette holders, and for readjusting phototimer sensitivity to
account for the lower detection efficiency of the conventional PSP
imaging plate compared to 400 speed screen-film detectors.
Long-term costs include cassette and imaging platelongevity,
maintenance and cleaning of the imaging plates and reader assembly,
replacement of damaged detectors, and continuous oversight with a
quality control program.
1
An alternate description of CR as "cassette radiography" in lieu
of "computed" is a sign of the technological changes that are
occurring-PSP technologies are now being implemented in enclosed
housings, with high-speed parallel laser beam stimulation and
photodiode array acquisition that fully reads the exposed storage
phosphor in as little as 5 seconds, comparable to many "direct" DR
detectors. Available a reportable flat-panel radiography detectors
in a cassette form, some with wireless technology, which can
provide active readout at the point of service without subsequent
user interactions.
Technological developments of storage phosphors include
compounds with less intrinsic lag during stimulation for faster
readout times,"dual-side" phosphor deposition on a transparent
substrate to improve X-ray detection and stimulated luminescence
efficiency for a higher signal-to-noise ratio (SNR) with the same
exposure, and "structured PSP" materials such as cesium bromide
(CsBr) that simultaneously improve spatial resolution and detection
efficiency.
2
Cassette-based CR detectors are available for digital mammography
with special adjustments to the reader device and laser beam (e.g.,
50 micron pixel sampling), and are a viable consideration for
converting from screen-film particularly for operations that have a
low patient volume and/or have relatively new mammography X-ray
equipment. Initial capital outlay is certainly lower than a
corresponding integrated digital radiography (or digital
mammography) flat-panel detector system.
Charge-coupled device detectors
The design of a charge-coupled device (CCD)-based DR system is
straightforward. The detector is comprised of a large FOV (e.g., 43
cm by 43 cm) scintillator that converts absorbed X-ray energy into
light. It also includes an optical lens assembly to focus the light
onto the photosensitive CCD array, and a CCD camera to integrate,
scan and output the corresponding light image. While there were
initially several configurations in early systems, today's
CCD-based detector is typically comprised of a single-compound
optical lens and a high-resolution CCD camera comprised of 9
million pixels (3000× 3000 pixels) to 16 million pixels (4000 ×
4000 pixels). When referred back to the image plane, this results
inimage pixel sizes of ~0.10 to ~0.14 µm (Figure 2). The
photosensitive area of the CCD chip is actually quite small, on the
order of 2.5 cm × 2.5cm to 4.0 cm × 4.0 cm, which is required to
maintain extremely high charge-coupling efficiency and low-noise
operation during the readout of the image. Thus, there is a large
optical demagnification that is necessary to focus the full FOV
light image onto the CCD sensor. One physical difficulty is the
inefficiency of light collection caused by the dispersed light
emission from the phosphor, resulting in only a small fraction that
can be focused onto the CCD, thus potentially reducing the
statistical integrity of information carried by the X-ray photons
and increasing overall noise in the image. This is determined by
the demagnification factor, conversion efficiency, luminance and
directionality of the light emission. A non-structured phosphor
such as gadolinium oxysulfide has a high light dispersion and
corresponding low fraction of light that can be focused on the CCD,
while a structured phosphor such as cesium iodide (CsI) produces a
more forward-directed light output, so that the lens-light
collection efficiency, and thus the SNR in the output image, is
better for a given incident X-ray exposure. Newer, advanced CCD
systems with a CsI phosphor have proven to be reasonably efficient,
particularly when using higher kilovolt peak (kVp) techniques that
produce more light photons per absorbed X-ray photon. One minor
disadvantage in some positioning situations is the relatively large
and bulky enclosure of a CCD-based DR system, necessitated by
placing the CCD out of the direct X-ray beam and using mirror
optics to reflect the light to the photosensor array.
Linear CCD arrays optically coupled to a scintillator by
fiberoptic channel plates (often with a demagnification taper of
2:1 to 3:1) are used in slot-scan geometries (Figure 3). A
significant advantage is pre- and postpatient collimation that
limits X-ray scatter and allows grid-free operation with equivalent
image quality (in terms of SNR) of a large area FOV at 2 to 4 times
less patient dose. Disadvantages include the extended exposure time
required for image acquisition with potential motion artifacts and
reduced X-ray tube efficiency. Nevertheless, imaging systems based
on slot-scan acquisition have provided excellent clinical results
for dedicated chest and full-body trauma imaging.
Complementary metal-oxide semiconductor (CMOS)
detectors
CMOS light-sensitive arrays are based upon a crystalline silicon
matrix and are essentially random access memory "chips" with
built-in photodiodes, storage capacitors, and active readout
electronics, operating at low voltage (3 to 5 volts) for image
acquisition and readout. The ability to randomly address any
detector element on the chip enables opportunities for automatic
exposure control (AEC) capabilities that are not easily performed
with a CCD photo detector. However, electronic noise has been a
problem that has slowed the introduction of this technology.
Construction of a large-area detector is also a hurdle for detector
arrays directly coupled to the X-ray converter, due to the maximum
detector element size currently achievable (on the order of 50 µm).
CMOS detector technology applications for radiography are currently
unavailable, but recent development of a CMOS-based digital
mammography detector compatible with conventional mammography
systems has beenintroduced for small FOV (10 cm × 10 cm)
application.
3
Active-matrix, flat-panel imagers
AMFPI technologies are based on thin-film-transistor (TFT)
arrays, made from amorphous silicon, upon which lithographic
etching and material evaporation on the micron scale produces the
electronic components and connections necessary for detector
operation. The flat-panel substrate is divided into individual
detector element (del) compartments, arranged in a row and column
matrix, typically with a spacing dimension of 70 microns to 250
microns, depending on the detector specifications. Components
within each del include a thin-film-transistor (essentially an
electronic switch), a charge collection electrode and a storage
capacitor. Electronic interconnections including gate lines (rows)
and drain lines (columns) are connected to each of the TFTs to
control the on/off status of each of the dels and to provide the
conduction paths to the charge amplifiers, respectively. All of the
TFTs are closed during an exposure to collect X-ray-induced charge
proportional to the incident X-ray fluence in each del. After the
exposure, active readout of the array occurs one row at a time by
activating the respective gate line, which turns on the TFTs and
allows the stored charge to flow along the columns from each del
capacitor via drain lines to the corresponding charge amplifier.
Banks of amplifiers simultaneously amplify the charge, convert to a
proportional voltage, digitize signals in parallel from each row of
the detector matrix and produce a corresponding row of integer
values in the digital image matrix (Figure 4). This process is
repeated for each row in the matrix. Detector readout speed is
governed by the intrinsic lag characteristics of the X-ray
converter material,switching speed of the TFT array electronics and
the number of independent charge amplifier arrays that can function
in parallel.
Limits to spatial resolution for TFT arrays include fill factor,
electronic interconnections and manufacturing yield. Fill-factor is
a term describing the active charge collection area to the total
area of the del,
4
and ideally is 100% for most efficient collection of X-ray
information. However, because electronic components and connection
lines occupy space on the substrate, with conventional TFT designs
the fill factor is less than ideal, this is more severe for smaller
del areas of ~0.1 µm (e.g., <50%), and can ultimately limit the
minimum size of the del (highest spatial resolution achievable). To
increase the spatial resolution by a factor of 2 requires a 4 times
increase in the number of electronic interconnections and a much
greater complexity in the design of the detector. The probability
of malfunctioning dels, gates and drains increases with higher
spatial resolution requirements, which can drastically lower the
yield. While there is no "perfect" TFT, small defects are corrected
by mapping bad dels and using interpolation methods of nearby
elements to "fill-in" the expected response. This is successful as
long as the interpolation does not exceed more than a couple
elements in the local area that require repair. In addition,
flat-field preprocessing methods (beyond the scope of the topic
presented here) are common and required for all digital radiography
detectors (DR and CR) to compensatefor variations in detector gain
and response across the detector.
5
The frequency of flat-field calibration suggested by manufacturers
(weekly, monthly or annually) is dependent on the detector and the
propensity for going out of calibration.
AMFPIs are classified into "indirect" and "direct" X-ray
converters, which are based on the physics of X-ray charge
conversion (Figure 5). Indirect AMFPIs use an X-ray phosphor
material layered on a TFT-photodiode array substrate, where light
photons are the intermediary between X-ray absorption and charge
generation on the TFT array detector element. Structured CsI is
currently the most widely used phosphor, mainly for excellent X-ray
detection efficiency and good spatial resolution capabilities.
Incident X-rays are absorbed, proportional light intensity is
produced, proportional charge amplitude is generated by the
photodiode upon exposure to the light, and the resultant charge
(X-ray signal) is stored at the local capacitor. Several additional
lithography steps are needed for including the photodiode
components during the manufacturing process of the TFT array
compared to TFT detectors without photodiodes.
6
Manufacturers also create arrays in two different ways: monolithic,
whereby the whole detector is created from one substrate material,
and smaller subdetector arrays that are joined together as
quadrants to create the full FOV detector. Both manufacturing
methods produce excellent detectors, but there is a limit on the
size of the monolithic production that is slightly smaller than the
conventional 43 cm x 43 cm large FOV. On the other hand,the process
of joining four independent detector subarrays into one large array
can have meshing defects, and while not visible on inspection of
the image, certainly have a noticeable impact on the noise
characteristics of the detector as determined by sophisticated
noise-power spectrum analyses.
7
Direct AMFPIs use a semiconductor X-ray converter that directly
produces electron-hole pairs in proportion to the incident X-ray
intensity. Amorphous selenium (a-Se) is currently the semiconductor
of choice, and is layered between two electrodes connected to the
bias voltage and a dielectric layer. The ion pairs are collected
under a high voltage (10 to 50 volts per micron of thickness) in
order to prevent recombination,and to prevent lateral spreading of
the information carriers within the a-Se material. A simple
TFT-array structure is used to collect the resultant charge, and
excellent spatial resolution is achieved, chiefly limited by the
size of the del. Fill-factor penalties are less of an issue with
directAMFPI devices, as electrode designs can funnel charges along
electric field lines. However, charge trapping and residual charge
lag can occur within the semiconductor layer, which can be quite
problematic for image ghosting from previous exposures. Detector
"relaxation" after an exposure is often performed to release
trapped charges within the substrate. Additionally, if the a-Se
substrate beginsto crystallize, the semiconducting properties of
the converter are reduced and the detector must be replaced.
Improved manufacturing methods have reduced trapping and lag to low
levels, and the propensity for crystallization is minimized by
placing a small amount of arsenic doping to the semiconductor layer
during manufacturing.
Which is better: indirect or direct AMFPI detectors? There are
no clear answers to this often-asked question; both technologies
have advantages and disadvantages. With current technological
implementations, in general, it appears that indirect AMFPIs have
fared better with conventional radiography applications and with
high-speed image acquisition and readout (e.g., dual-energy
radiography and digital tomosynthesis). Direct AMFPIs have
performed better in high-resolution environments such as digital
mammography. Of course, as technology changes, so do the
answers.
When deciding on an AMFPI detector technology for clinical
implementation, important considerations are the warranty and
guarantees for detector longevity and uptime, as replacement of the
detector electronics can be a very expensive proposition.
Comparing passive PSP (CR) and active DR
technologies
Advantages for cassette-based PSP (CR)
CR represents the digital technology that can directly replace
screen-film imaging at a low initial investment cost and still use
existing X-ray equipment with minimal changes or calibration
requirements. Multiple size cassettes allow the proper detector
size to be used for various procedures, and provide the ultimate in
positioning flexibility for any examination. New "point of service"
imaging plate readers onboard with portable X-ray systems can
provide rapid turnaround beside radiography examinations and
verification of positioning before leaving the patient area.
Smaller and highly capable in-room dedicated single-plate readers
can achieve good throughput by allowing the technologist to acquire
the next image in a series while processing the previously exposed
imaging plate. Capital outlay required to introduce digital
radiography into the clinical environment is a clear advantage of
CR systems over most DR technologies, due in large part to the
ability to use existing X-ray infrastructure. While proponents of
DR technology point towards the lower detection efficiency of
standard PSPimaging plates and required higher dose to achieve a
given SNR, introduction of dual-side readout phosphors and CsBr
structured storage phosphors have resulted in a significant
increase in the detection efficiency on par with some DR-based
detectors.
Bottom line: Flexibility, convenience, ease of implementation
and initial (low) costs are the hallmarks of cassette-based CR
detectors.
Advantages for active, integrated DR detectors
Integrated DR detector systems have a great advantage with
technologist workflow and patient throughput, particularly for
ambulatory outpatient imaging systems. This is chiefly due to the
rapid display of the acquired image seconds after the exposure. A
specific DR system for dual-energy chest imaging at UC Davis
provides a patient turnaround for a 2-view chest in under 2 minutes
from patient positioning to image acquisition to
review/verification of the images and sending to the PACS. A
corresponding room using CR cassettes easily takes 5 to 7 minutes
or more, much of the time is spent handling cassettes and waiting
for the image readout and processing. AMFPIs typically have a
significantly higher detective quantum efficiency compared to a
conventional PSP/CR cassette (e.g., ~60% compared to ~30%), meaning
that reduced X-ray techniques with lower patient doses can be used
for similar image quality and SNR.
Recent entry of wired, and now wireless, portable DR detectors
are treading on the domain of PSP cassettes in many instances,
making a strong case for the possible elimination of PSP detectors
in the future, but there is still a way to go before that will
happen. Certainly the retrofit market for implementation of add-on
DR detectors is poised for rapid growth as users want the benefits
of DR workflow without the huge capital investment of a new,
integrated DR system. Because of high acquisition speed, advanced
acquisition technologies, such as dual-energy radiography and
digital tomosynthesis, are possible in situations that can benefit
from these current, and future, cutting-edge capabilities.
Bottom line: Image acquisition speed, advanced acquisition
technologies and excellent image quality are the hallmarks of DR
systems.
Disadvantages for cassette-based PSP (CR) detectors
The delay between exposure and readout with labor-intensive
handling is a major disadvantage of cassette-based PSP detectors,
particularly in situations that require the technologist to be away
from the patient. Readout and processing time of imaging plates can
be long, and for single plate readers, overexposures can take a
long time to "erase" the residual signals before another imaging
plate can be inserted. PSP detectors are always "on," meaning that
they are prone to background radiation and scattered radiation if
improperly stored next to, or in, a radiographic room. It is very
important to implement an erasure of imaging plates that have not
been used frequently, particularly after a long weekend as a
precaution to eliminate any potential contrast degrading background
radiation signals on patient images. Multiple PSP detectors, while
a nice "redundancy" feature, can also be a chore in terms of
inventory, cleaning, quality control testing and general
maintenance.
Disadvantages for active, integrated DR detectors
Positioning flexibility (e.g., cross-table lateral acquisitions,
exotic views) is often difficult or impossible to achieve with many
DR systems (although newer integrated designs and portable DR
detectors are overcoming some of the early system design
limitations). Initial capital costs are extremely high (on the
order of 3 to 5 times the cost of a high-quality CR reader and
imaging plates), making the case for return on investment crucial
if the workload can't justify the throughput advantages. Redundancy
concerns and the cost of detector replacement are serious issues,
and any down time in a busy, high-throughput room can be
detrimental to operating costs and timely delivery of patient
care.
Best of both: Combining cassette-based CR and integrated
DR
In most major hospitals, there will always be the need for
flexibility and for efficient throughput for radiography, and with
the current status of both technologies, they nicely coexist.
Integration of CR and DR in a single room with similar image
processing and concatenation of studies under a single accession
number is an attribute that allows for maximum productivity under
one acquisition console, which has been successfully implemented in
an outpatient clinic in the UC Davis Health System. In low-volume
situations that do not require high throughput for radiography
examinations, the implementation of cassette-based CR has been the
choice for replacement of screen-film technology. However, with
advancements and lower costs likely, the potential of using a
robust, low-cost portable DR detector in lieu of CR is very
possible in the coming years.
Replacement strategies for historically cassette-based CR
systems at the UC Davis Health System are focused on implementation
of AMFPIindirect detectors in targeted service areas, side by side
with radiographic rooms that continue to use cassette-based CR.
This has proven to be an effective way of providing the best of
both technologies by taking advantage of the benefits of each.
While the investment in CR technology remains strong, there is no
doubt that advanced imaging capabilities such as dual-energy
radiography and digital tomosynthesis volume imaging have resulted
in a focus on and greater emphasis of flat-panel DR technologies
for targeted procedures that can reduce the need for computed
tomography, MRI and ultrasound examinations.
The American College of Radiology has recently published
practice guidelines for digital radiography
8,9
including a review of digital radiography technologies, clinical
considerations for implementation and quality control
recommendations. This information is an excellent resource for all
individuals responsible for digital radiography in the clinical
environment.
Conclusion
Digital radiography detectors (CR and DR detector systems) are
now in the majority. Cassette-based photostimulable storage
phosphor detectors comprise the largest segment of digital
detectors; however, cassetteless, integrated detectors are
increasing in areas that demand high efficiency and high
throughput. The historical comparison of "CR vs. DR" while still a
consideration for making an informed decision for choice of
adigital radiography system, is of less importance, as the advance
of technology has blurred the differences. Image acquisition speed
of AMFPIs is enabling advanced acquisition and processing
capabilities such as dual-energy radiography and digital
tomosynthesis, which will likely become common methods that are
used to overcome the superimposition limitations of conventional
projection radiography and result in more accurat ediagnoses. Just
expect to pay more for advanced technology capabilities, and be
sure to justify the expense in terms of realistic workloads and
throughput requirements; otherwise do not expect a decent return on
investment. Knowledge of DR system characteristics, advantages,
disadvantages and operational details provide the insight and
confidence to make a reasonable, informed decision on equipment
selection and optimization of digital radiography implementation
for a specific purpose.
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