To develop a dual-energy X-ray CT (DE-CT) system, we have performed investigation of high-speed dual-energy photon counting using two comparators and a low-dark-counting LSO-MPPC (multipixel photon counter) detector. To measure X-ray spectra, electric charges produced in the MPPC are converted into voltages and amplified by a highspeed current-voltage amplifier, and the event pulses are sent to a multichannel analyzer. The MPPC was driven under pre-Geiger mode at an MPPC bias voltage of 70.7 V. The event pulses are sent to two high-speed comparators for selecting two threshold energies to perform DE-CT. The ED-CT is accomplished by repeated linear scans and rotations of the object, and two sets of projection curves of the object are obtained simultaneously by the linear scan. In the DECT, two different-energy tomograms are obtained simultaneously, and photon-count energy subtraction imaging was carried out.
X-ray photon counting was performed using a readymade silicon-PIN photodiode (Si-PIN-PD) at tube voltages ranging
from 42 to 60 kV, and X-ray photons are directly detected using the 100 MHz Si-PIN-PD without a scintillator.
Photocurrent from the diode is amplified using charge-sensitive and shaping amplifiers. Using a multichannel analyzer,
X-ray spectra at a tube voltage of 60 kV could easily be measured. The photon-counting computed tomography (PCCT)
is accomplished by repeated linear scans and rotations of an object, and projection curves of the object are obtained
by the linear scan. In the PC-CT, we confirmed the energy-dispersive effect with changes in lower-level voltage of the
event pulse using a comparator.
X-ray photon counting was performed using a silicon X-ray diode (Si-XD) at a tube current of 2.0 mA and tube voltages
ranging from 50 to 70 kV. The Si-XD is a high-sensitivity Si photodiode selected for detecting X-ray photons, and Xray
photons are directly detected using the Si-XD without a scintillator. Photocurrent from the diode is amplified using
charge-sensitive and shaping amplifiers. To investigate the X-ray-electric conversion, we performed the event-pulseheight
(EPH) analysis using a multichannel analyzer. Photon-counting computed tomography (PC-CT) is accomplished
by repeated linear scans and rotations of an object, and projection curves of the object are obtained by the linear scan.
The exposure time for obtaining a tomogram was 10 min at a scan step of 0.5 mm and a rotation step of 1.0°. In PC-CT
at a tube voltage of 70 kV, the image contrast of iodine media fell with increasing lower-level voltage of the event pulse
using a comparator.
X-ray photons are detected using an Lu2(SiO4)O [LSO] single-crystal scintillator with a decay time of 40 ns and a multipixel photon counter (MPPC). The photocurrent from the MPPC is amplified by a high-speed current-voltage
amplifier with an 80 MHz-gain-band operational amplifier, and the 200-ns-width event pulses are sent to a multichannel
analyzer to measure X-ray spectra. The MPPC is driven in the pre-Geiger mode at a bias voltage of 70.7 V and a
temperature of 23°C. Photon-counting computed tomography (PC-CT) is accomplished by repeated linear scans and
rotations of an object, and projection curves of the object are obtained by linear scanning. The exposure time for
obtaining a tomogram was 10 min with scan steps of 0.5 mm and rotation steps of 1.0°. At a tube voltage of 100 kV, the
maximum count rate was 350 kcps/pixel. We carried out PC-CT using gadolinium media and confirmed the energydispersive
effect with changes in the lower level voltage of event pulses using a comparator.
A low-dose-rate X-ray computed tomography (CT) system is useful for reducing absorbed dose for patients. The CT
system with a tube current of 1.91 mA was developed using a silicon-PIN X-ray diode (Si-PIN-XD). The Si-PIN-XD is
a selected high-sensitive Si-PIN photodiode (PD) for detecting X-ray photons. X-ray photons are detected directly using
the Si-PIN-XD without a scintillator, and the photocurrent from the diode is amplified using current-voltage and
voltage-voltage amplifiers. The output voltage is converted into logical pulses using a voltage-frequency converter with maximum frequency of 500 kHz, and the frequency is proportional to the voltage. The pulses from the converter are sent to differentiator with a time constant of 1 μs to generate short positive pulses for counting, and the pulses are counted using a counter card. Tomography is accomplished by repeated linear scans and rotations of an object, and projection curves of the object are obtained by the linear scan. The exposure time for obtaining a tomogram was 5 min at a scan step of 0.5 mm and a rotation step of 3.0°. The tube current and voltage were 1.91 mA and 100 kV, respectively, and gadolinium K-edge CT was carried out using filtered X-ray spectra with a peak energy of 52 keV.
A high-sensitive X-ray computed tomography (CT) system is useful for decreasing absorbed dose for patients, and a
dark-count-less photon-counting CT system was developed. X-ray photons are detected using a YAP(Ce) [cerium-doped
yttrium aluminum perovskite] single crystal scintillator and an MPPC (multipixel photon counter). Photocurrents are
amplified by a high-speed current-voltage amplifier, and smooth event pulses from an integrator are sent to a high-speed comparator. Then, logical pulses are produced from the comparator and are counted by a counter card. Tomography is accomplished by repeated linear scans and rotations of an object, and projection curves of the object are obtained by the linear scan. The image contrast of gadolinium medium slightly fell with increase in lower-level voltage (Vl) of the comparator. The dark count rate was 0 cps, and the count rate for the CT was approximately 250 kcps.
We developed an embossed radiography system utilizing single- and dual-energy subtractions for decreasing the
absorption contrast of unnecessary regions, and contrast resolution of a target region was increased using image-shifting
subtraction and a linear-contrast system in a flat panel detector (FPD). To carry out embossed radiography, we
developed a computer program for two-dimensional subtraction, and a conventional x-ray generator with a 0.5-mm-focus tube was used. Energy subtraction was performed at tube voltages of 42.5 and 70.0 kV, a tube current of 1.0 mA, and an x-ray exposure time of 5.0 s. Embossed radiography was achieved with cohesion imaging by use of the
FPD with pixel sizes of 48 ×48 μm, and the shifting dimension of an object in the horizontal and vertical directions
ranged from 48 to 144 μm. We obtained high-contrast embossed images of fine bones and coronary arteries approximately 100 μm in diameter.
An energy-discriminating x-ray camera is useful for performing monochromatic radiography using polychromatic x rays. This x-ray camera was developed to carry out K-edge radiography using iodine-based contrast media. In this camera, objects are exposed by a cone beam from a cerium x-ray generator, and penetrating x-ray photons are detected by a cadmium telluride detector with an amplifier unit. The optimal x-ray photon energy and the energy width are selected out using a multichannel analyzer, and the photon number is counted by a counter card. Radiography was performed by the detector scanning using an x-y stage driven by a two-stage controller, and radiograms obtained by energy discriminating are shown on a personal computer monitor. In radiography, the tube voltage and current were 60 kV and 36 µA, respectively, and the x-ray intensity was 4.7 µGy/s. Cerium K-series characteristic x rays are absorbed effectively by iodine-based contrast media, and iodine K-edge radiography was performed using x rays with energies just beyond iodine K-edge energy 33.2 keV.
Embossed radiography is an important technique for imaging target region by decreasing absorption contrast of objects. The ultra-high-speed embossed radiography system consists of a computed radiography system, an intense flash x-ray generator, and a computer program for shifting the image pixel. In the flash x-ray generator, a high-voltage condenser of 200 nF was charged to 50 kV, and the electric charges in the condenser were discharged to the flash x-ray tube after triggering the cathode electrode. The molybdenum-target evaporation lead to the formation of weakly ionized linear plasma, and intense molybdenum K-series x-rays were produced. High-speed radiography was performed using molybdenum K-rays, and the embossed radiography was carried out utilizing single-energy subtraction after the image shifting. The minimum spatial resolution was equal to the sampling pitch of the CR system of 87.5 μm, and concavoconvex radiography such as phase-differential imaging was performed with an x-ray duration of approximately 0.5 Μs.
Characteristic x-ray generator consists of a constant high-voltage power supply, a filament power supply, a
turbomolecular pump, and an x-ray tube. The x-ray tube is a demountable diode which is connected to the
turbomolecular pump and consists of the following major devices: a pipe-shaped molybdenum hole target, a tungsten
hairpin cathode (filament), a focusing (Wehnelt) electrode, a polyethylene terephthalate x-ray window 0.25 mm in
thickness, and a stainless-steel tube body. In the x-ray tube, the positive high voltage is applied to the anode (target)
electrode, and the cathode is connected to the tube body (ground potential). In this experiment, the tube voltage applied
was from 25 to 35 kV, and the tube current was regulated to within 10 μA by the filament temperature. The exposure
time is controlled in order to obtain optimum x-ray intensity. The electron beams from the cathode are converged to the
target by the focusing electrode, and sharp K-series characteristic x-rays are produced through the focusing electrode at
a tube voltage of 35 kV. Using this generator, we performed monochromatic radiography, monochromatic x-ray
computed tomography, and x-ray fluorescence analysis.
A photon-counting K-edge x-ray Computed Tomography (CT) system is useful for discriminating photon energy and for
decreasing absorbed dose for patients. The CT system is of the first generation type and consists of an x-ray generator, a
turn table, a translation stage, a two-stage controller, a multipixel photon counter (MPPC) module, a 0.5-mm-thick
zinc oxide (ZnO) scintillator, a counter board (CB), and a personal computer (PC). Tomography is accomplished by
repeating the translation and rotation of an object. Penetrating x-ray photons from the object are detected by the
scintillator in conjunction with the MPPC module, and the event signals are counted by the CB. Without using energy
discriminating, photon counting CT was carried out by controlling x-ray spectra.
Energy-discriminating x-ray camera is useful for performing monochromatic radiography using polychromatic x-rays.
The x-ray camera was developed to carry out K-edge radiography using iodine-based contrast media. In this camera,
objects are exposed by a cerium x-ray generator, and penetrating x-rays are detected by a cadmium telluride (CdTe)
detector with an amplifier unit. The optimal x-ray photon energy and energy width are selected out using a multichannel
analyzer (MCA), and the photon number is counted by a counter board (CB). Radiography was performed by the
detector scanning using an x-y stage driven by a two-stage controller, and x-ray images obtained by energy
discriminating are shown in a personal-computer (PC) monitor. Cerium K-series characteristic x-rays are absorbed
effectively by iodine based contrast media, and iodine K-edge radiography was performed using x-rays with photon
energies just beyond K-edge energy 33.2 keV.
An x-ray fluorescence (XRF) computed tomography (CT) system utilizing a cadmium telluride (CdTe) detector is
described. The CT system is of the first generation type and consists of a cerium x-ray generator, a turn table, a
translation stage, a two-stage controller, a CdTe spectrometer, a multichannel analyzer (MCA), a counter board (CB),
and a personal computer (PC). When an object is exposed by the x-ray generator, iodine K-series fluorescences are
produced and are detected from vertical direction to x-ray axis using the spectrometer. Fluorescent photons are selected
out using the MCA and are counted by the PC via CB, and XRF CT is performed by repeating translation and rotation
of an object.
X-Ray Fluorescence (XRF) analysis is useful for measuring density distributions of contrast media in vivo. An XRF
camera was developed to carry out mapping for iodine-based contrast media used in medical angiography. In this
camera, objects are exposed by an x-ray beam formed using a 3.0-mm-diameter lead hole. Next, cerium K-series
characteristic x-rays are absorbed effectively by iodine media in objects, and iodine fluorescences are produced from
the objects. Iodine Kα fluorescences are selected out using a 58-μm-thick stannum filter and are detected by a cadmium
telluride (CdTe) detector. Kα rays are discriminated out by a multichannel analyzer (MCA), and photon number is
counted by a counter board (CB). The objects are moved and scanned using an x-y stage driven by a two-stage
controller, and x-ray images obtained by iodine mapping are shown in a personal computer (PC) monitor. In particular,
iodine fluorescences were produced from remanent iodine elements in a cancer region of a rabbit ear.
Digital subtraction is useful for carrying out embossed radiography by shifting an x-ray source, and energy subtraction
is an important technique for imaging target region by deleting unnecessary region in vivo. X-ray generator had a
100-μm-focus tube, energy subtraction was performed at tube voltages of 40 and 60 kV, and a 3.0-mm-thick aluminum
filter was used to absorb low-photon-energy bremsstrahlung x-rays. Embossed radiography was achieved with cohesion
imaging using a flat panel detector (FPD) with pixel sizes of 48×48 μm, and the shifting distance of the x-ray source in
horizontal direction and the distance between the x-ray source and the FPD face were 5.0 mm and 1.0 m, respectively.
At a tube voltage of 60 kV and a tube current of 0.50 mA, x-ray intensities without filtering and with filtering were 307
and 28.4 μGy/s, respectively, at 1.0 m from the source. In embossed radiography of non-living animals, the spatial
resolution measured using a lead test chart was approximately 70 μm, and we observed embossed images of fine bones,
soft tissues, and coronary arteries of approximately 100 μm.
An energy-discriminating K-edge x-ray Computed Tomography (CT) system is useful for increasing contrast resolution
of a target region and for diagnosing cancers utilizing a drug delivery system. The CT system is of the first generation
type and consists of an x-ray generator, a turn table, a translation stage, a two-stage controller, a cadmium telluride
(CdTe) detector, a charge amplifier, a shaping amplifier, a multi-channel analyzer (MCA), a counter board (CB), and a
personal computer (PC). The K-edge CT is accomplished by repeating translation and rotation of an object. Penetrating
x-ray spectra from the object are measured by a spectrometer utilizing the CdTe detector, amplifiers, and MCA. Both
the photon energy and the energy width are selected by the MCA for discriminating photon energy. Enhanced iodine
K-edge x-ray CT was performed by selecting photons with energies just beyond iodine K-edge energy of 33.2 keV.
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