Download Parham C, et al. Design and implementation of a compact low

New Developments
in Medicine
Design and Implementation of a Compact Low-Dose
Diffraction Enhanced Medical Imaging System1
Christopher Parham, MD, PhD, Zhong Zhong, PhD, Dean M. Connor, PhD, L. Dean Chapman, PhD, Etta D. Pisano, MD
Rationale and Objectives. Diffraction-enhanced imaging (DEI) is a new x-ray imaging modality that differs from conventional
radiography in its use of three physical mechanisms to generate contrast. DEI is able to generate contrast from x-ray absorption,
refraction, and ultra-small-angle scatter rejection (extinction) to produce high-contrast images with a much lower radiation dose
compared to conventional radiography.
Materials and Methods. A prototype DEI system was constructed using a 1-kW tungsten x-ray tube and a single silicon
monochromator and analyzer crystal. The monochromator crystal was aligned to reflect the combined Ka1 (59.32 keV) and Ka2
(57.98 keV) characteristic emission lines of tungsten using a tube voltage of 160 kV. System performance and demonstration of
contrast were evaluated using a nylon monofilament refraction phantom, full-thickness breast specimens, a human thumb, and
a live mouse.
Results. Images acquired using this system successfully demonstrated all three DEI contrast mechanisms. Flux measurements
acquired using this 1-kW prototype system demonstrated that this design can be scaled to use a more powerful 60-kW x-ray tube
to generate similar images with an image time of approximately 30 seconds. This single–crystal pair design can be further
modified to allow for an array of crystals to reduce clinical image times to <3 seconds.
Conclusions. This paper describes the design, construction, and performance of a new DEI system using a commercially
available tungsten anode x-ray tube and includes the first high-quality low-dose diffraction-enhanced images of full-thickness
human tissue specimens.
Key Words. Diffraction enhanced imaging; analyzer based imaging; low dose; refraction; soft tissue; mammography.
ª AUR, 2009
Since Roentgen’s discovery in 1895, x-ray imaging has relied
on the varied attenuation of photons to generate images.
Diffraction-enhanced imaging (DEI) is a new x-ray imaging
modality that uses two additional contrast mechanisms,
Acad Radiol 2009; 16:911–917
1
From the University of North Carolina at Chapel Hill School of Medicine,
Departments of Radiology and Biomedical Engineering, the UNC-Lineberger
Comprehensive Cancer Center, and the UNC Biomedical Research Imaging
Center, 4030 Bondurant Hall, Campus Box 7000, Chapel Hill, NC 27599-7000
(C.P., E.D.P.); the National Synchrotron Light Source, Brookhaven National
Laboratory, Upton, NY (Z.Z., D.M.C.); and the Department of Anatomy and Cell
Biology, University of Saskatchewan College of Medicine, Saskatoon, SK,
Canada (L.D.C.). This research was supported by contract DE-AC0298CH10886 from the US Department of Energy (Washington, DC) and by
laboratory-directed research and development grant 05-057 from Brookhaven
National Laboratory (Upton, NY). Received October 9, 2008; accepted
February 3, 2009. Address correspondence to: C.P. e-mail: caparham@
gmail.com
ª AUR, 2009
doi:10.1016/j.acra.2009.02.007
refraction and ultra-small-angle scatter rejection (extinction)
(1). Previous imaging of breast cancer specimens (1–5),
articular cartilage (6,7), and small animals (8–10) with synchrotron-based DEI systems has shown improvement in fine
detail in images compared to conventional radiography, with
applications ranging from industrial inspection to medical
imaging (3–6,9,11–13).
In conventional radiography, contrast between different
tissues is provided by their differential x-ray attenuation. For
soft tissues, with adjacent materials having similar electron
densities, achieving adequate contrast with conventional
x-ray imaging requires imaging at low x-ray energies (approximately 20 keV) to take maximal advantage of small
differences in the effective atomic number through the photoelectric effect. Unfortunately, the attenuation of x-rays
increases as energy is lowered. To obtain adequate transmission through the body and produce an image with an
acceptable signal-to-noise ratio, more incident radiation must
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be used when imaging at low energies, resulting in a higher
absorbed dose to patients.
Attenuation through the photoelectric effect or Compton
scattering is not the only interaction that x-ray photons can
have with tissue. X-rays passing through an interface between adjacent materials of different densities can be refracted. In biologic materials, this refraction amount is very
small (on the order of tenths of microradians) and thus does
not contribute to the contrast in conventional x-ray imaging.
With the use of x-ray optics, DEI can resolve these very small
angular deviations of the x-ray beam and thus provide a new
contrast mechanism.
On the basis of Bragg’s law of x-ray diffraction, DEI uses
a silicon monochromator crystal to create a collimated,
monochromatic x-ray beam. The beam then passes through
the sample, where the photons can be absorbed, scattered,
refracted, or have no interaction with the sample. The transmitted beam is then incident upon a silicon analyzer crystal.
The analyzer crystal has an angular reflectivity profile,
known as its rocking curve, which is a function of x-ray incident angle (1). At the Bragg angle, or peak of the rocking
curve, the reflectivity is close to unity. Reflectivity decreases
rapidly with angle, approaching zero within about 1 microradian when imaging at 59 keV, which is used to convert
angular change into intensity differences in the resulting
image.
Contrast generated from the refraction of photons is
known as refraction contrast (1). Materials with minimal
absorption contrast (eg, soft tissues such as human breast
specimens) can have a large amount of refraction contrast
(3,5). Refracted photons can interact without being absorbed
and thus do not add to the absorbed radiation dose. Refraction
contrast is a function of both the amount of refraction generated in an object and the slope of the rocking curve. The
amount of refraction varies as the inverse square of the energy, whereas the slope of the rocking curve is proportional to
the energy. Thus, refraction contrast varies as the inverse of
the energy. Interactions that cause photons to deviate outside
the narrow angular acceptance window of the analyzer
crystal create ultra-small-angle scatter contrast, also known
as extinction, a reduction analogous to the removal of photons through coherent scattering or Compton scattering in
conventional radiography (10). Extinction contrast can be
present in combination with absorption and refraction in
a single diffraction-enhanced image, depending on its material properties. Pairs of images on each side of the analyzer
rocking curve can be acquired and processed to generate
a refraction image and apparent absorption image, which is
a combination of both absorption and extinction contrast (1).
DEI is a type of phase-contrast x-ray imaging. Several
other phase-contrast-based technologies are being developed.
Studies have shown that because propagation-based phasecontrast imaging generates contrast from the second
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derivative of the phase and DEI from the first derivative, DEI
generates greater image contrast than propagation-based
phase contrast for a given energy as well as a weaker energy
dependency than propagation-based imaging (8,14). Thus,
DEI is able to provide images that have simultaneously higher
contrast with lower dose compared to propagation-based
phase-contrast imaging. Two other phase-contrast techniques
are analogous to DEI in that they measure the first derivative
of the phase. The first was developed by Ingal et al (15); it uses
asymmetrically cut monochromator and analyzer crystals
along with an x-ray tube source to generate its images.
Although the asymmetrically cut crystals allow the use of
a large-area x-ray beam, this configuration greatly limits the
ability to use high-energy sources. Thus, these investigators’
studies have only shown the ability to image small samples of
thickness 10 to 30 mm at an x-ray energy of 22.2 keV.
Most recently, Pfeiffer et al (16,17) developed a gratingbased phase-contrast imaging technique. This result was
significant because the system both uses a polychromatic
x-ray beam (there is no need to use a monochromator) and
gets its contrast from the first derivative of the phase. To date,
these investigators’ studies have shown the ability to image
with low-energy x-ray tube sources, but they have shown that
there are technical hurdles that will have to be overcome to
image using higher energy tube sources (18). Advancements
have also been made by Myers et al (19) in the area of phasecontrast tomography to generate phase contrast in low-attenuation test phantoms.
Each of these studies has been limited in that they have
been successful only at imaging thin objects with low-energy
x-rays, whereas this DEI prototype has been used to generate
high-quality images in thick specimens with higher energy
x-rays up to 59 keV. The demonstrated ability to image
thick objects combined with the ability to use commercially
available high-power x-ray tubes and an optical design that
can be expanded to use an array of crystals to dramatically
decrease imaging time makes this design plausible for clinical phase-contrast x-ray imaging.
In DEI, a collimated, monochromatic x-ray beam must be
used. By making an x-ray beam collimated and monochromatic, the x-ray flux is reduced by 5 orders of magnitude (10),
thereby severely limiting the number and types of x-ray
sources that can generate sufficient flux to create a clinically
useful DEI device. To date, DEI of clinically relevant specimens has been available only through the use of high-flux
x-ray sources present at synchrotron research facilities.
A tungsten anode x-ray tube generates a high flux of
characteristic Ka1 x-rays at 59.318 keV and Ka2 x-rays at
57.982 keV. These x-rays only weakly attenuate through thick
soft-tissue specimens, such as the human breast. For example,
35% of 59.318 keV x-rays transmit through 5 cm of water,
compared to 0.65% of 18 keV x-rays. This dramatic increase
in transmission at high x-ray energy results in a reduction of
Academic Radiology, Vol 16, No 8, August 2009 COMPACT DIFFRACTION ENHANCED MEDICAL IMAGING SYSTEM
the incident x-ray intensity required for imaging. Factoring in
the high brightness of the Ka1 and Ka2 emission lines of
tungsten, the weak absorption of the tungsten Ka1 and Ka2
x-rays in soft tissue, and the persistence of refraction contrast
even at high energies makes a tungsten anode x-ray tube
a strong candidate for a future clinical x-ray system.
In this paper, we report on the design and performance
testing of a compact DEI device that uses a conventional
tungsten anode x-ray tube to image clinically relevant specimens at an ultra low dose. An argument is presented on how
to transition from this experimental prototype system to
a DEI system with clinically relevant imaging times.
MATERIALS AND METHODS
Figure 1 presents the system design used for this prototype
compact-source DEI device. The x-ray source used for this
DEI system was a Comet MXR-160HP/20 x-ray tube (Comet
AG, Flamatt, Switzerland) with a stationary tungsten anode
with a focal spot size of 0.4 mm. A Titan 160 x-ray generator
(GE Inspection Technologies, Ahrensburg, Germany) with
a maximum voltage of 160 kV and 1 kW total power was
used to power the x-ray tube. A 2.0-mm-thick tantalum collimator with a 25-mm-wide by 1-mm-high aperture was
placed over the exit window of the x-ray tube to create a fan
beam. A 250-mm copper filter was placed in the beam to
remove low-energy x-rays from the beam (specifically, the
19.77-keV x-rays that would be passed by the monochromator’s 111 reflection). A monochromator was built using
a single perfect float zone silicon crystal (Shaw Monochromators, Riverton, KS) of the 333 reflection type measuring
70 35 10 mm. With a Darwin width of 0.8 microradians
and a Bragg angle of approximately 0.1 radians, the intrinsic
energy resolution, or DE/E, of this silicon monochromator is
0.47 eV at 59 keV. The monochromator was placed at a distance of 100 mm from the x-ray tube at an angle of 5.7 with
respect to the incident x-ray beam to select the Ka1 (59.318
keV) characteristic emission line of tungsten. Because of the
source’s divergence and the monochromator’s crystal size,
the monochromator also reflected the Ka2 (57.982 keV)
emission line. The Ka1 portion of the postmonochromator
fan beam was parallel to the ground.
Each sample, at a distance of 650 mm from the x-ray tube,
was moved vertically through the x-ray beam using a translation stage (Newport Corporation, Irvine, CA). A silicon
analyzer crystal measuring 150 60 10 mm was placed
behind the sample and tuned to an angle of 5.7 with respect to
the imaging beam. The analyzer is the same type as used for
synchrotron DEI studies at the National Synchrotron Light
Source (Upton, NY) reported elsewhere (10).
All images were acquired using a Fuji ST-VI general
purpose image plate (Fuji Medical Systems, Stamford, CT)
Digital detector
or image plate
Lead collimators
Sample
stage
Sodium iodide Silicon analyzer
crystal
detectors
Tantalum
collimator
Copper
Silicon filter
monochromator
crystal
X-ray tube
with
tungsten
anode
Figure 1. X-ray tube–based diffraction-enhanced imaging
prototype design schematic.
that was placed at an angle twice that of the analyzer Bragg
angle (11.4 ). The image plate was digitized using a Fuji
BAS-2500 image plate reader (Fuji Medical Systems) at
a resolution of 50 mm. This detector plate was selected because of its fixed noise for long exposure times and its
detection efficiency at 59 keV. The image plate was scanned
using another translation stage (Newport Corporation) in the
opposite direction of the sample stage to form a radiograph of
the sample using the fan beam. The detector and sample
scanning methods have been previously described (10).
Several measurements were made to quantify the system
flux and beam properties. With the analyzer crystal set to the
Bragg peak, the beam was imaged using an image plate. A line
scan across the image of the beam was used to determine the
beam height. To measure the system’s reflectivity profile, or
rocking curve, the photon counts were measured as a function
of the analyzer crystal’s detuning angle with respect to the
Bragg peak for values from 20 to +20 microradians in 0.1microradian steps, with a tube voltage of 160 kVp and current
of 6.2 mA and 1-second integration time per angular step. The
photon counts were measured using a sodium iodide detector
in photon-counting mode with an opening of 1 cm high by 6.5
mm wide. The detector was placed at the normal image plate
position (with the image plate removed), so the flux would be
recorded at the image plate position. To determine the beam
flux, a rocking curve scan of the analyzer crystal was performed repeatedly using a tube voltage of 160 kVp and current of 6.2 mA. The photon counts for a fully detuned analyzer
crystal (background) were subtracted from the photon counts
at the peak of the rocking curve (the Bragg angle) to calculate
the photon counts per second. The beam height and detector
width (6.5 mm) were then used to calculate the in-beam flux
(in photons per second per square centimeter). To ensure that
the beam energy was indeed 59.318 keV, a stepped-wedge
acrylic phantom was created, with each step being 11.1 mm.
Rocking curve scans were done through zero to seven steps of
the phantom to measure the transmitted flux.
Nylon monofilaments, because of their low x-ray attenuation and their similar refracting properties to soft tissues,
have been used as a standard phantom for characterizing DEI
refraction contrast (20). A refraction phantom containing
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Figure 2. Nylon refraction phantom and measured postanalyzer rocking curves. The diameters of the nylon monofilaments from largest to
smallest are 560, 360, 200, and 100 mm. Phantom images were acquired at 0.6 microradians (a) and +0.6 microradians (c) on the rocking
curve using a tube voltage of 160 kV and 6.2 mA with a surface dose of 0.04 mGy. A measured postanalyzer rocking curve (b) using the Bragg
[111] and Bragg [333] reflections is presented to demonstrate system performance.
nylon monofilaments of various diameters was placed behind
a 4.5-cm water bath and imaged at analyzer rocking curve
positions of +0.6 and 0.6 microradians.
A 12 cm, a full-thickness human breast mastectomy
specimen (obtained through an institutional review board–
approved protocol at the University of North Carolina at
Chapel Hill) was compressed to 7.0 cm and mounted to approximate a clinical mediolateral oblique view and imaged at
+0.5 and 0.5 microradian points on the rocking curve with
the new device, using a surface dose of 0.040 mGy and mean
glandular dose of 0.035 mGy for each image. The mean
glandular dose was calculated using the monoenergetic normalized glandular dose reported previously (21). The specimen was placed in a water bath during imaging to prevent
drying during the procedure and to reduce the presence of
reftraction artifacts at the interface between air and the
specimen. Imaging time using the compact-source DEI
system was 23 hours. This specimen, still under the same
compression and placed in the same water bath, was imaged
using a GE Senographe 2000D digital mammography system
(GE, Paris, France) at the University of North Carolina at
Chapel Hill at 31 kVp and 98 mAs (using the rhodium/rhodium setting) which yielded a surface dose of 12 mGy and
mean glandular dose of 1.9 mGy for this 50:50 fat/glandular
tissue specimen.
A cadaveric human thumb (acquired through an institutional review board–approved protocol at Rush University
School of Medicine) was also imaged using the new device
at a surface dose of 0.07 mGy and at the +0.5 and 0.5
microradian points on the rocking curve. The specimen was
placed in a 4.5-cm-thick water bath container during imaging
to prevent drying during the procedure and imaged in the
lateral projection. Image acquisition time was 11 hours.
After approval by the Brookhaven National Laboratory’s
Institutional Animal Care and Use Committee, a nude mouse
(model NCRNu-M; Taconic Farms, Germantown, NY) was
anesthetized using an intraperitoneal injection of 75 mg/kg
sodium pentobarbital and placed between two acrylic plates
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to minimize its motion during scanning. Imaging was made
in the anteroposterior direction, with a maximum thickness of
the mouse of about 3 cm. Acquisition time was 40 minutes
total, with a surface dose of 0.004 mGy.
RESULTS
A refraction phantom using nylon monofilament with
decreasing dimensions of 560, 360, 200, and 100 mm at analyzer positions of +0.6 and 0.6 microradians is presented
in Figure 2. Because x-ray attenuation by nylon at 60 keV is
negligible, contrast seen in the images of the fibers is due to
refraction contrast. The black-over-white contrast feature of
each fiber in the +0.6-microradian image is characteristic of
DEI refraction contrast. This contrast is reversed to whiteover-black when imaged on the opposite side of the rocking
curve, further demonstrating the presence of refraction
contrast.
The Ka1 and Ka2 emission lines were imaged using
a stationary Fuji ST-VI image plate at a distance of 960 mm
from the x-ray source to allow for adequate divergence, with
a measured height of 600 mm each. The system’s rocking
curve, shown in Figure 2b, demonstrates experimental
agreement with the theoretically predicted rocking curve for
the 333 reflection at 59 keV. At the detector position (where
the image plate is placed for imaging), the x-ray flux in the
postmonochromator fan beam was measured to be 4.4 105
photons/s/cm2.
Figure 3 is a plot of the measured flux through the system
as a function of stepped-wedge phantom thickness. It shows
a log-linear relationship between the transmitted flux and the
step thickness. This log-linear relationship means that little to
no beam hardening occurred as more steps of acrylic were
added; thus, the incident beam can be inferred to be monochromatic or nearly monochromatic. From the slope of the
graph, the absorption coefficient was found to be 0.222 0.015 cm 1 which agrees with the accepted value for the
Academic Radiology, Vol 16, No 8, August 2009 COMPACT DIFFRACTION ENHANCED MEDICAL IMAGING SYSTEM
ln(I/I0)
1.11 cm
-1.8
-1.2
-0.6
0
0
1
3
4
333, 160 kVp
333, 100 kVp
Linear (333, 160
kVp)
Linear (333, 100
kVp)
5
Thickness (cm)
2
6
7
8
Figure 3. Measured photon flux at the detector position as a function of the thickness
of an acrylic wedge phantom for tube settings of both 160 kVp (blue diamonds) and 100
kVp (red squares), demonstrating a log-linear relationship between the transmitted flux
and the step thickness.
Figure 4. Full-thickness breast specimen image acquired using a GE Senographe
2000D (GE, Paris, France) conventional digital mammographic system (a) and a corresponding image acquired using the compact diffraction-enhanced imaging system (b).
The 12.0-cm-thick specimen containing an invasive cancer was compressed to 7.0 cm
and mounted to provide a mediolateral view of the breast. Each image was acquired
using a tube voltage of 160 kV and 6.2 mA with a surface dose of 0.04 mGy and an
absorbed dose 0.035 mGy. The spiculated masses and skin evident in the conventional
digital radiograph are visible in the diffraction image.
absorption coefficient of acrylic, which is 0.2295 cm 1 at
59.318 keV.
Figure 4 demonstrates a full-thickness breast specimen
acquired using this prototype DEI system compared to
a conventional digital radiograph of the same specimen.
Qualitative visualization of breast cancer features, such as
spiculations with high refraction contrast, are similar to
those reported in previous studies using synchrotron x-ray
sources (3–5). A diffraction-enhanced image of the human
thumb is presented in Figure 5, with simultaneous visualization of bone, muscle, tendons, and other soft-tissue
structures not possible with conventional radiography. Five
small soft-tissue calcifications are identified in the figure by
an arrow. Qualitative visualization of all three DEI contrast
mechanisms in both bone and soft tissue is similar to previously published diffraction-enhanced images acquired
using a synchrotron (6,7,13). Figure 6 presents the first live
animal image acquired using an x-ray tube–based phasecontrast imaging system. Clinically relevant features such
as air in the bowel, extinction contrast in the aerated lung,
trachea, and other soft-tissue details are visualized. In addition, this live-animal image demonstrates that in vivo DEI
contrast can be generated even with image times of 40
minutes.
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Figure 5. A human thumb image acquired using the compactsource diffraction system. Note the soft-tissue detail, including the
tendon with its attachment to the bone, the fingernail, and the interface of the skin edge and subcutaneous tissue. An area of five
small soft-tissue calcifications is identified in the image by an
arrow.
DISCUSSION
Using the data and design information acquired from this
prototype, a plausible extrapolation to a full-power DEI system can be made. The measured in-beam flux at the image
plate was 4.4 105 photons/s/cm2 using a tube voltage of 160
kVp, current of 6.2 mA, and a beam height of 0.6 mm. There
are commercially available tubes with power output of up to
80 kW. Because flux scales linearly with current, an 80-kW
tube running with a tube voltage of 160 kVp and maximized
current would generate 80 times the flux of this 1-kW prototype system. In addition, the source-to-detector distance was
approximately 2 times what a likely clinical system would be.
This reduction in the source to detector distance will increase
the flux by a factor of 1/r, decreasing the imaging time by
a factor of 2. The 1/r relationship of flux to distance in DEI
differs from the conventional relationship of 1/r2 because of
the vertical collimation of the x-ray beam where Ka1 and Ka2
maintain their sizes at all distances. With these two system
changes in tube power and source-to-detector distance, a flux
of 7 107 photons/s/cm2 would be possible.
What would this mean for imaging time in a DEI mammographic system? For a compressed breast of average
thickness (4 cm), the vertical region to be imaged is approximately 15 cm. If a pixel size of 125 mm with 300 photons/pixel is used in combination with a 60-kW-rated x-ray
tube, images acquired at the 60% reflectivity point of the
rocking curve would require approximately 30 seconds. It is
also important to remember that this is a line-scan system
with a geometry similar to slot-scan digital mammographic
systems (22,23). For the above configuration, the scan speed
would be approximately 0.5 cm/s. The beam height, including both the Ka1 and Ka2 lines, would be about 1.5 mm.
Thus, a given feature edge will be in the beam for only 1% of
the imaging time or, for this example, about 0.5 seconds.
Motion blurring of feature edges would be comparable to that
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Figure 6. In vivo mouse image acquired using the compact
source. This image of a live anesthetized mouse reveals features
including air in the bowel and other soft-tissue details acquired
with a much lower dose than conventional absorption x-ray imaging. Note that the lungs show increased density compared to
their usual appearance on a conventional radiograph, which is due
to extinction contrast.
of a 0.5-second conventional mammogram. Although this is
a significant improvement from the current prototype, an
image time of 30 seconds is still not acceptable for clinical
imaging. Image time can be further reduced by the addition of
multiple crystal pairs in an array to capture more of the diverging fan beam of the x-ray tube. The narrow x-ray beam
generated using a synchrotron has a nominal horizontal and
vertical divergence, and thus, only a single monochromator
crystal and analyzer crystal can be used. This is not the case
with an x-ray tube that has a significant horizontal and vertical divergence, allowing for multiple monochromator and
analyzer crystal pairs to be used. Wide divergence in the
horizontal plane increases the field of view to a width sufficient for general-purpose imaging, while divergence in the
vertical plane allows multiple crystals to be used, capturing
a larger portion of the emitted photons and maximizing
photon efficiency. For each analyzer and monochromator
crystal pair, there is an estimated twofold reduction in image
time than can reduce the estimated imaging time to <3
seconds. A practical consideration in increasing the number
of crystals is the size of the array, especially crystal depth at
the monochromator, where the beam size is narrow. The
depth of the crystals used in this design is 10 mm, but the
crystal depth needed for x-ray diffraction is <1 mm, allowing
the use of thin silicon wafers. A multiple-crystal design also
increases the object thickness that can be imaged for a given
image time, which is essential for general-purpose imaging
such as chest, abdominal, and musculoskeletal imaging.
Another primary contribution of this work compared to
conventional radiography and previous phase-contrast imaging systems using low-energy x-rays is radiation dose reduction. The persistence of refraction and extinction contrast
at high x-ray beam energies allows soft-tissue imaging
without complete reliance on absorption, which means that
fewer incident and absorbed photons are needed to generate
useful image contrast. The importance of reducing the dose
of radiation to the population, especially for younger people
and children, has recently been emphasized (24,25). The dose
to the human population due to medical imaging tests has
Academic Radiology, Vol 16, No 8, August 2009 COMPACT DIFFRACTION ENHANCED MEDICAL IMAGING SYSTEM
gradually been increasing over the past 30 years, largely
because of advances in the imaging technologies themselves,
particularly computed tomography (25). We believe that
DEI, when optimized for each individual application, is
likely to provide equivalent or greater diagnostic information
at substantially lower doses compared to conventional radiography (3–5,7,11).
In addition, each component of the device must be optimized for the specific intended application. X-ray sources
with different spectra, different crystals, and specialized detectors will be needed for each unique clinical application.
Diffraction-enhanced computed tomography has been used
to reconstruct DEI contrast mechanisms in three dimensions
(2,3,26,27), and the design reported here could be adapted for
diffraction-enhanced computed tomography. Because of the
importance of reducing x-ray exposure to the youngest and
most radiation-vulnerable patients, we are planning next to
adapt this device for the imaging of infants and children and
to develop a diffraction-enhanced computed tomographic
system. Shortly thereafter, we hope to use it for breast cancer
imaging.
CONCLUSION
Given the evident advantages of improved object feature
visibility at a reduced dose to patients, we believe that this
compact DEI x-ray system is just the first of many tools that
will be produced using our methodology that will use both
x-ray refraction and extinction in addition to absorption to
produce high-quality medical and industrial images.
ACKNOWLEDGMENTS
We wish to thank William Thomlinson, Martin Yaffe,
Ann Sherman, Carol Muehleman, Jun Li, Charles Livasy,
Laura Faulconer, Elodia Cole, and the late Dale Sayers for
their contributions to our work. The rights to the intellectual
property described in this paper have been licensed to NextRay, Inc. Drs Parham, Zhong, Chapman, Connor, and
Pisano all hold stock in NextRay.
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