Many improvements have been made to amorphous silicon (a-Si) flat panel detectors (FPDs) to meet the market needs for different x-ray imaging applications. With the current generation of a-Si FPDs the performance is limited by the a-Si thin film transistors (TFTs). The low electron mobility of a-Si necessitates large TFT’s with large parasitic dataline capacitance, which increases electronic noise and reduces the pixel fill factor (FF). In other words, large TFT’s negatively impact Signal-to-Noise Ratio (SNR). CMOS FPDs were introduced to provide improved low dose imaging performance and faster readout times, but the increase in cost can be prohibitive. IGZO TFTs have an electron mobility that is <10x higher than a-Si, which facilitates a reduction in the size of the TFT while also reducing the pixel discharge time, resulting in an increase to both the detector readout rate and the SNR. Reducing the TFT size is particularly important in achieving adequate low dose performance in dynamic detectors with pixels approaching 100μm. A 31cm x 31cm (100μm) FPD using IGZO TFTs was evaluated at 25 frames/second (fps) in 1x1, 2x2, 3x3, and 4x4 binning. In the 1x1 standard noise configuration, the noise equivalent dose (NED) was 24nGy with a max linear dose (MLD) of 10uGy. The NED was reduced to 6.6nGy in the 2x2, 3.4nGy in the 3x3, and 2.4nGy in the 4x4 mode. The linearity of the IGZO imager was comparable to a-Si imager. The 1x1 MTF was 57.5% at 1 lp/mm and 28.5% at 2lp/mm. The quantum limited DQE in the 1x1 binning mode was 79% at 0 lp/mm and 47% at 1 lp/mm. The 1x1 DQE measured at NED was 71% at 0 lp/mm, 29% at 1 lp/mm. This paper will explore how to optimally employ IGZO and present data from a first IGZO imager, showing that IGZO is an excellent technology for the future of FPDs.
Complementary metal-oxide-semiconductors (CMOS) flat panel detectors (FPD) have steadily gained acceptance into medical imaging applications1-15. Selecting the proper detector technology for the imaging task requires optimization to balance the cost and the image quality. To facilitate this, fundamental detector performance of CMOS and a-Si panels were evaluated using the following quantitative imaging metrics: X-ray sensitivity, Noise Equivalent Dose (NED,) Noise Power Spectrum (NPS), Modulation Transfer Function (MTF), and Detective Quantum Efficiency (DQE). Imaging task measurements involved high-contrast and low-contrast resolution assessment. Varex FPDs evaluated for this study included: CMOS 3131 (150 μm pixel), a-Si 3030X (194 μm pixel), a-Si XRpad2 3025 (100 μm) and CMOS 2020 (100 μm pixel). Performance comparisons were organized by pixel size: large pixels, 150 μm CMOS and 194 μm a-Si, and small pixels, 100 μm in a-Si and CMOS technology. The results showed high dose DQE of the a-Si 3030X was about 10% higher than the CMOS 3131 between 0 - 1.8 cycles/mm, while beyond 1.8 cycles/mm, the CMOS performed better. The 3030X low dose DQE was higher than the 3131 between 0-1.3 cycles/mm, while the CMOS performance was higher beyond 1.3 cycles/mm. The high dose DQE of 100 μm a-Si was higher than the 100 μm CMOS for all frequencies. However, the low dose DQE of 100 μm CMOS was higher beyond 0.6 cycles/mm, while the 100 μm a-Si pixel had higher DQE only between 0 – 0.6 cycles/mm. Large pixel image quality (IQ) assessment favored a-Si pixel with 7% higher Contrast-to-Noise-Ratio (CNR) results for both high and low contrast-detail at 500 nGy. Small pixel CNR favored CMOS with ~38% better high contrast-detail and 12% greater low contrast-detail at ~500 nGy. Through these measurements that combine imaging metrics and image quality, we demonstrated a practical method for selecting the appropriate detector technology based on the requirements of the imaging applications.
Mammography systems demand high quality imaging at reduced acquisition times. The Varex 3024MX imager was designed specifically with the demands of mammographic imaging in mind: high spatial resolution, excellent low contrast resolution as well as excellent low dose performance, and acquisition speeds capable of tomography. This paper will describe the details of the next generation a-Si mammography sensor array and contrast the predicate product, PS3024M. The Varex 3024MX imager delivers a 3584x2816 matrix with a pixel pitch of 83um resulting in an active area of 297.5mm x 233.7mm., optimized for mammography applications. A 250um thick deposited columnar CsI(Tl) layer is used as the scintillator. The development of a new pixel architecture and charge amplifier ASIC allows for faster readout of the sensor array at 16 bit pixel depth. The faster readout of the Varex 3024MX enables readout speeds up to 10fps. In addition to the faster frame rates, the combination of the new pixel architecture and ASIC, result in a very low electronic noise floor and improved ghosting behavior. The results, as outlined below, will show that the 3024MX design targeted improvements to detective quantum efficiency (DQE), maximum linear dose (MLD), quantum-limited dose (QLD), ghosting, and image readout time.
The combinations of a 60 fps kV x-ray flat panel imager, a 19 focal spot kV x-ray tube enabled by a steered electron beam, plus SART or SIRT sliding reconstruction via GPUs, allow real time 6 fps 3D-rendered digital tomosynthesis tracking of the respiratory motion of lung cancer lesions. The tube consists of a “U” shaped vacuum chamber with 19 tungsten anodes, spread uniformly over 3 sides of a 30 cm x 30 cm square, each attached to a cylindrical copper heat sink cooled by flowing water. The beam from an electron gun was steered and focused onto each of the 19 anodes in a predetermined sequence by a series of dipole, quadrupole and solenoid magnets. The imager consists of 0.194 mm pixels laid out in 1576 rows by 2048 columns, binned 4x4 to achieve 60 fps projection image operation with 16 bits dynamic range. These are intended for application with free breathing patients during ordinary linac C-arm radiotherapy with modest modifications to typical system hardware or to standard clinical treatment delivery protocols. The sliding digital tomosynthesis reconstruction is completed after every 10 projection images acquired at 60 fps, but using the last 19 such projection images for each such reconstruction at less than 8 mAs exposure per 3D rendered frame. Comparisons, to “ground truth” optical imaging and to diagnostic 4D CT (10 phase) images, are being used to determine the accuracy and limitations of the various versions of this new “19 projection image x-ray tomosynthesis fluorooscopy” motion tracking technique.
The focus of this work was to improve the DQE performance of a
full-field digital mammography (FFDM) system by
means of selecting an optimal X-ray tube anode-filter combination in conjunction with an optimal scintillator
configuration. The flat panel detector in this work is a Varian PaxScan 3024M. The detector technology is comprised of
a 2816 row × 3584 column amorphous silicon (a-Si) photodiode array with a pixel pitch of 83μm. The scintillator is
cesium iodide and is deposited directly onto the photodiode array and available with configurable optical and x-ray
properties. Two X-ray beam spectra were generated with the anode/filter combinations, Molybdenum/Molybdenum
(Mo/Mo) and Tungsten/Aluminum (W/Al), to evaluate the imaging performance of two types of scintillators, high
resolution (HR) type and high light output (HL) type. The results for the HR scintillator with W/Al anode-filter (HRW/
Al) yielded a DQE(0) of 67%, while HR-Mo/Mo was lower with a DQE(0) of 50%. In addition, the DQE(0) of the
HR-W/Al configuration was comparable to the DQE(0) of the HL-Mo/Mo configuration. The significance of this result
is the HR type scintillator yields about twice the light output with the W/Al spectrum, at about half the dose, as
compared to the Mo/Mo spectrum. The light output or sensitivity was measured in analog-to-digital convertor units
(ADU) per dose. The sensitivities (ADU/uGy) were 8.6, 16.8 and 25.4 for HR-Mo/Mo, HR-W/Al, HL-Mo/Mo,
respectively. The Nyquist frequency for the 83 μm pixel is 6 lp/mm. The MTF at 5 lp/mm for HR-Mo/Mo and HR-W/Al
were equivalent at 37%, while the HL-Mo/Mo MTF was 24%. According to the DQE metric, the more favorable anodefilter
combination was W/Al with the HR scintillator. Future testing will evaluate the HL-W/Al configuration, as well as
other x-ray filters materials and other scintillator optimizations. While higher DQE values were achieved, the more
general conclusion is that the imaging performance can be tuned as required by the application by modifying optical and
x-ray properties of the scintillator to match the spectral output of the chosen anode-filter combination.
Lag and sensitivity modulation are well known temporal artifacts of a-Si photodiode based flat panel detectors. Both
effects are caused by charge carriers being trapped in the semiconductor. Trapping and releasing of these carriers is a
statistical process with time constants much longer than the frame time of flat panel detectors. One way to reduce these
temporal artifacts is to keep the traps filled by applying a pulse of light over the entire detector area every frame before
the x-ray exposure. This paper describes an alternative method, forward biasing the a-Si photodiodes and supplying free
carriers to fill the traps. The array photodiodes are forward biased and then reversed biased again every frame between
the panel readout and x-ray exposure. The method requires no change to the mechanical construction of the detector,
only minor modifications of the detector electronics and no image post processing. An existing flat panel detector was
modified and evaluated for lag and sensitivity modulation. The required changes of the panel configuration, readout
scheme and readout timing are presented in this paper. The results of applying the new technique are presented and
compared to the standard mode of operation. The improvements are better than an order of magnitude for both
sensitivity modulation and lag; lowering their values to levels comparable to the scintillator afterglow. To differentiate
the contribution of the a-Si array, from that of the scintillator, a large area light source was used. Possible
implementations and applications of the method are discussed.
Digital flat panel a-Si x-ray detectors can exhibit image lag of several percent. The image lag can limit the temporal
resolution of the detector, and introduce artifacts into CT reconstructions. It is believed that the majority of image lag is
due to defect states, or traps, in the a-Si layer. Software methods to characterize and correct for the image lag exist, but
they may make assumptions such as the system behaves in a linear time-invariant manner. The proposed method of
reducing lag is a hardware solution that makes few additional hardware changes. For pulsed irradiation, the proposed
method inserts a new stage in between the readout of the detector and the data collection stages. During this stage the
photodiode is operated in a forward bias mode, which fills the defect states with charge. Parameters of importance are
current per diode and current duration, which were investigated under light illumination by the following design
parameters: 1.) forward bias voltage across the photodiode and TFT switch, 2.) number of rows simultaneously forward
biased, and 3.) duration of the forward bias current. From measurements, it appears that good design criteria for the
particular imager used are 8 or fewer active rows, 2.9V (or greater) forward bias voltage, and a row frequency of 100
kHz or less. Overall, the forward bias method has been found to reduce first frame lag by as much as 95%. The panel
was also tested under x-ray irradiation. Image lag improved (94% reduction), but the temporal response of the
scintillator became evident in the turn-on step response.
A unique 64-row flat panel (FP) detector has been developed for sub-second multidetector-row CT (MDCT). The intent
was to explore the image quality achievable with relatively inexpensive amorphous silicon (a-Si) compared to existing
diagnostic scanners with discrete crystalline diode detectors. The FP MDCT system is a bench-top design that consists
of three FP modules. Each module uses a 30 cm x 3.3 cm a-Si array with 576 x 64 photodiodes. The photodiodes are
0.52 mm x 0.52 mm, which allows for about twice the spatial resolution of most commercial MDCT scanners. The
modules are arranged in an overlapping geometry, which is sufficient to provide a full-fan 48 cm diameter scan. Scans
were obtained with various detachable scintillators, e.g. ceramic Gd2O2S, particle-in-binder Gd2O2S:Tb and columnar
CsI:Tl. Scan quality was evaluated with a Catphan-500 performance phantom and anthropomorphic phantoms. The FP
MDCT scans demonstrate nearly equivalent performance scans to a commercial 16-slice MDCT scanner at comparable
10 - 20 mGy/100mAs doses. Thus far, a high contrast resolution of 15 lp/cm and a low contrast resolution of 5 mm @
0.3 % have been achieved on 1 second scans. Sub-second scans have been achieved with partial rotations. Since the
future direction of MDCT appears to be in acquiring single organ coverage per scan, future efforts are planned for
increasing the number of detector rows beyond the current 64- rows.
The dynamic range of many flat panel imaging systems are fundamentally limited by the dynamic range of the charge amplifier and readout signal processing. We developed two new flat panel readout methods that achieve extended dynamic range by changing the read out charge amplifier feedback capacitance dynamically and on a real-time basis. In one method, the feedback capacitor is selected automatically by a level sensing circuit, pixel-by-pixel, based on its exposure level. Alternatively, capacitor selection is driven externally, such that each pixel is read out two (or more) times, each time with increased feedback capacitance. Both methods allow the acquisition of X-ray image data with a dynamic range approaching the fundamental limits of flat panel pixels. Data with an equivalent bit depth of better than 16 bits are made available for further image processing. Successful implementation of these methods requires careful matching of selectable capacitor values and switching thresholds, with the imager noise and sensitivity characteristics, to insure X-ray quantum limited operation over the whole extended dynamic range. Successful implementation also depends on the use of new calibration methods and image reconstruction algorithms, to insure artifact free rebuilding of linear image data by the downstream image processing systems.
The multiple gain ranging flat panel readout method extends the utility of flat panel imagers and paves the way to new flat panel applications, such as cone beam CT. We believe that this method will provide a valuable extension to the clinical application of flat panel imagers.
This paper describes a new flat panel imager designed for use in cardiovascular and mobile C-arm imaging systems. The a-Si sensor array has a 1024 x 1024 matrix with a pixel pitch of 194 μm, resulting in an active area of 198.7 mm x 198.7 mm. The imager allows frame rates of up to 30 fps in full resolution fluoroscopy mode
and up to 60 fps in a 2 x 2 binned low dose fluoroscopy mode. Typically, a 600 μm thick deposited columnar CsI(Tl) layer is used as the scintillator.
Improvements in the pixel architecture, charge amplifier ASICs, and system level electronics resulted in a very low electronic noise floor, such that both the fluoroscopy and low dose fluoroscopy modes of the panel are x-ray quantum limited below 1 μR/frame.
Low power consumption electronics combined with a mechanical design optimized for heat transfer and dissipation makes air-cooling sufficient for most environments. The small size of 24.1 x 24.1 x 6 cm and the weight of only 4.1 kg meet the requirements of C-Arm systems. Special consideration was given to the border around the active area, which has been reduced to 2 cm. Reported performance parameters include linearity, lag, contrast ratio, MTF, and DQE. For the full resolution mode, the MTF is greater than 0.53 and 0.21 at 1
and 2 lp/mm, respectively. DQE measured at 22 nGy/frame was greater than 0.68, 0.50, and 0.23 at 0, 1, and 2 lp/mm, respectively.
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