Large area flat panel solid-state detectors are being studied for digital radiography and fluoroscopy. Such systems use active matrix arrays to readout latent charge images created either by direct conversion of x-ray energy to charge in a photoconductor or indirectly using a phosphor and individual photodiodes on the active matrix array. Our work has utilized the direct conversion method because of its simplicity and the higher resolution possible due to the electrostatic collection of secondary quanta. Aliasing of noise occurs in current designs of direct detectors based on amorphous selenium (a-Se) because of its high intrinsic resolution. This aliasing leads to a decrease in detective quantum efficiency (DQE) as frequency increases. It has been predicted, using a previously developed model of the complete imaging system, that appropriately controlled spatial filtration can reduce this aliased noise and hence increase DQE at the Nyquist frequency,f%y. Our purpose is to experimentally verify this concept by implementing presampling filtration in a practical flat panel system. An a-Se based flat panel imager is modified by incorporating an insulating layer between the active matrix and the a-Se layer to introduce a predetermined amount of presampling burring. The modified imager is evaluated using standard linear analysis tools, modulation transfer function (MTF), noise power spectra (NPS) and DQEW, and the results are compared to theoretical predictions.

Amorphous selenium direct-conversion x-ray detectors have been used successfully for full field digital mammography (FFDM) and digital radiography (DR). Such detectors characteristically exhibit high spatial resolution and conversion efficiency that is a function of the applied electric field [1]. At an electric field of 10 volts per micron, about 50 electron volts of photon energy are required to generate an electron-hole pair in a typical amorphous selenium x-ray conversion layer. At FFDM and DR imaging x-ray energies each absorbed photon can generate only about 250 to 1000 electronhole pairs. Each absorbed x-ray photon is only contributing 4x10 -17 to 1.6x10 -16 coulombs of imaging charge. On the noise side, detectors operating at room temperature have a basic thermal (kTC) noise of 300 to 600 electrons per pixel from the image charge storage capacitor. Electronic noise from the front-end charge amplifier is also amplified by one plus the ratio of the TFT data line capacitance and the feedback capacitance of the charge amplifier. Medical imaging applications must therefore employ low noise thin film transistor (TFT) arrays, low data line capacitance and low noise charge integration amplifiers to achieve high signal-to-noise ratio (SNR) and detective quantum efficiency (DQE).To achieve quantum-noise limited imaging results with the lowest practical x-ray exposure dose, it is desirable to include an additional low-noise gain stage in the x-ray conversion layer. This is particularly important for the application of dynamic imaging or for tomosynthesis where x-ray dose per frame is very limited. A new structure for an amorphous selenium detector that employs an internal biased gain grid to cause avalanche-gain within the x-ray conversion layer is being proposed. A signal charge amplification of at least 10X can be achieved without introducing excessive noise. Quantum-limited image detection should then be attainable for even very low exposures.