Although it is possible to extend the approach described above to create a full-area detector, Equation below

indicates that the large demagnification factor that would be required would result in unacceptably low coupling efficiency and a secondary quantum sink. This problem can be circumvented by designing the detector as a mosaic of smaller modules. One manufacturer, Trex, uses a mosaic of 3 x 4 detectors, each one employing a demagnification on the order of 2x. This system produces a full-size image of the breast with approximately 40 micron pixels (Figure 1).

The x-ray absorbing phosphor in this system is composed of thallium-activated cesium iodide (CsI:TI). The advantage of utilizing CsI as the x-ray absorber is that it can be grown in columnar crystals which act as fiber optics. When coupled to the photodiode pixels, there is little lateral spread of light and, therefore, high spatial resolution can be maintained. In addition, unlike conventional phosphors in which diffusion of light and loss of resolution become worse when the thickness is increased, CsI phosphors can be made thick enough to ensure a high value of n while maintaining high spatial resolution.

Fig. 1 Full field detector formed as a mosaic of small-field devices. Each of the 12 modules consists of a demagnifying fiber-optic taper which couples light from a Csl:Ti phosphor layer to a full frame CCD.

Active matrix LCDs(AMLCDs)have been made using amorphous (hydrogenated amorphous silicon. The active matrix is a large-area integrated circuit consisting of a large number of thin-film field-effect transistors (TFTs) connected to individual photodetector elements in a matrix.

The potential advantages of such self-scanned, compact readout systems include their compactness, freedom from veiling glare, geometric uniformity, and immunity to stray magnetic fields. To produce an x-ray detector, Csl:Ti is evaporated directly onto the active matrix.

The principle of operation of an amorphous silicon detector is shown schematically in Figure 2. The dels are configured as photodiodes (Figure 2 (a)) which convert the optical signal from the phosphor to charge and store that charge on the capacitance of the element. Being low-noise devices, the photodiodes provide a very large dynamic range, on the order of 40,000:1. A typical thin-film transistor readout array is shown in Figure 2 (b). The signal is read out by activation of scanning control lines for each row of the device, connected to the gates of TFTs located on each detector pixel. An entire row of the detector array is activated simultaneously and the signal is read on lines for each column in the array which connect all the TFT sources in that column to a low-noise charge amplifier. The amplified signals from the columns are then multiplexed and digitized. This al-lows fast detector readout and requires a number of electronic channels equal to the number of columns of the array.

Alternatively, instead of TFT readout various diode switching schemes can be used. The advantage of the diode approach is that, because the photodiode has to be made anyway, the switching diode can be made at the same time without increase in the number of material-processing steps. The disadvantages of diode readout is a strong nonlinearity and large charge injection.

The area allocated to each pixel of the array must contain the photodiode, switching device, and control and signal lines so that the fill factor is less than 100%. This potential loss of x-ray utilization efficiency becomes proportionately greater as the pixel size is decreased and provides a challenge for the application of this technology to very high resolution applications. Currently dels of 10 μm have been produced and new techniques should allow sizes down to 50 or 60μm.

Fig. 2 Amorphous silicon full-field detector. a)Light from a Csl:Tl layer is direct coupled to a photodiode on each del. b) The readout array on an amorphous silicon plate uses thin film transistor switches to multiplex the charge stored on the capacitance of each del to readout lines.

General Electric has produced a system using CsI on a-Si with a del of 100 μm. The detector assembly fits onto a modified GE conventional mamography unit.

Probably the most widespread detectors for digital radiography to date have been photostimulable phosphors, also known as storage phosphors. These phosphors are commonly in the barium fluorohalide family, typically BaFBr:Eu²⁺, where the atomic energy levels of the europium activator determines the characteristics of light emission. X-ray absorption mechanisms are identical to those of conventional phosphors. They differ in that the useful optical signal is not derived from the light that is emitted in prompt response to the incident radiation, but rather from subsequent emission when electrons and holes are released from traps in the material. The initial x-ray interaction with the phosphor crystal causes electrons to be excited. Some of these produce light in the phosphor in the normal manner, however, the phosphor is intentionally designed to contain traps which store the charges. By stimulating the crystal by irradiation with red light, electrons are released from the traps and raised to the conduction band of the crystal, subsequently triggering the emission of shorter wavelength (blue) light. This process is called photostimulated luminescence. The physics of photostimulable phosphor imaging has been reviewed in more detail elsewhere.

Fig. 3 Digital mammography system based on a photostimulable phosphor with laser raster readout

In the digital radiography application, the imaging plate is positioned in a light-tight cassette or enclosure, exposed and then read by raster scanning the plate with a laser to release the luminescence (Figure 3). The emitted light is collected and detected with a photomultiplier tube whose output signal is digitized to form the image.

The energy levels in the crystal are critical to the effective operation of the detector. The energy difference between the traps and the conduction band E_T must be small enough so that stimulation with laser light is possible, yet sufficiently large to prevent random thermal release of the electron from the trap. The energetics should also provide for wavelength of the emitted light that can be efficiently detected by a photomultiplier and for adequate wavelength separation between the stimulating and emitted light quanta to avoid contaminating the measured signal. The electrons liberated during irradiation either produce light promptly or are stored in traps. Because the “prompt” light is not of interest in this application, the efficiency of the storage function can be improved by increasing the probability of electron trapping. On the other hand, when these electrons are released by the stimulating light during readout, the probability of their being retrapped instead of producing light would then be higher, thus reducing the efficiency the readout. The optimum balance occurs where the probabilities of an excited electron being retrapped or stimulating fluorescence are equal. This causes the conversion efficiency to be reduced by a factor of 4 compared to the same phosphor without traps, that is, a factor of 2 from the prompt light given off during x-ray exposure and another factor of 2 from unwanted retrapping of the electrons during readout.

In addition, the decay characteristics of the emission must be sufficiently fast that the image can be read in a conveniently short time while capturing an acceptable fraction of the emitted energy. In practice, depending on the laser intensity, the readout of a stimulable phosphor plate yields only a fraction of the stored signal. This is a disadvantage with respect to sensitivity and readout noise, however, it can be helpful by allowing the plate to be “pre-read,” in other words, read out with only a small part of the stored signal to allow automatic optimization of the sensitivity of the electronic circuitry for the main readout.

The photostimulable phosphor is a convenient detector for digital radiography in that, when placed in a cassette, it can be used with conventional x-ray machines. Large-area plates are conveniently produced, and because of this format, images can be acquired quickly. The plates are reusable, have linear response over a wide range of x-ray intensities, and are erased simply by exposure to a uniform stimulating light source to release any residual traps.

One limitation of this type of detector is that because the traps are located throughout the depth of the phosphor material, the laser beam providing the stimulating light must penetrate into the phosphor. Scattering of the light within the phosphor causes release of traps over a greater area of the image than the size of the incident laser beam. This results in loss of spatial resolution, which is aggravated if the plate is made thicker to increase η. An ideal solution to this problem would be a phosphor which was nonscattering for the stimulating light and both non-scattering and nonabsorbing for the emitted light.Limitations can arise in the readout stage, where efficient collection of the emitted light requires great attention to design. This can result in a secondary quantum sink, especially at high spatial frequencies, causing a reduction of DQE(f). Fuji is currently performing clinical evaluation of a photostimulable phosphor plate system for use in digital mammography. The system is designed to provide data at a sampling interval of 100μm.