X-ray Detector Array Replaces Film in Radiographic Imaging
For nearly 100 years, conventional projection radiography has used film to capture x-ray images, chemically processing the exposed film to create visible images for diagnosis. The functional utility and perceived high image quality of x-ray film in combination with various intensifying screens has made it the standard for medical imaging. Although it is used for image capture, display, storage, and transfer, the technique has its disadvantages. Processing time is required to produce images, which cannot then be electronically transferred without the additional steps of scanning the films. Moreover, the technique suffers from image degradation caused by optical scatter inherent to the process. Such considerations have created interest in digital imaging.
In the 1970s, digital imaging modalities such as computed tomography, ultrasound, and nuclear medicine gained widespread acceptance. The trend toward digital imaging was pushed further in the 1980s by magnetic resonance imaging (MRI) and digital subtraction angiography techniques. Efforts to integrate conventional radiography into the digital environment through film digitizers, storage phosphor-based computed radiography, and digital conversion of image intensifier video outputs have been less than satisfactory, however, involving both intermediate conversion and additional work by technologists. The result is that an estimated 70 percent of all diagnostic examinations are still performed using conventional x-ray techniques.
In recent years it has become technically and economically feasible to use electronic technologies to display, store, and transfer x-ray images. Digital image capture has remained elusive, however. Now an x-ray detector array has been developed in which amorphous selenium, a semiconductor material performs direct x-ray detection, opening the door for real time digital x-ray detection and imaging.
Indirect versus direct imaging
Conventional screen film systems image x-rays indirectly onto film. The radiation exiting from the patient impinges on material in an intensifying screen, causing it to fluoresce and emit visible light that exposes the emulsion to create an image. The significant flaw in any x-ray capture system involving a light conversion step like this is that the emitted light scatters through the media, with some photons traveling directly to the detector plane and some undergoing diversion en route to the detector plane (see Figure 1). Such scatter can decrease image resolution and image contrast.
FIGURE 1. In indirect detection methods such as film (above, from left), computed radiography, or cesium-iodide-based solid state detection, x-rays generate photons of visible light that are captured by film or detectors. This approach suffers from light scattering that can degrade image quality. In the direct detection approach (above right), x-rays generate electron hole pairs in amorphous selenium directly, and the charges can be read out by a thin film transistor array (below).
Photo courtesy of Sterling Diagnostic Imaging
Solid-state alternatives to film-based systems typically involve a scintillation material such as cesium iodide that once again converts the x-ray energy to visible light. An array of thin film photodiodes captures this energy and generates an electrical signal that is read out by an array of thin film transistors. As a result of light scatter, such approaches to date have been characterized by poor contrast and low spatial resolution.
The x-ray detector array technology developed by Sterling Diagnostic Imaging Inc. (Greenville, SC) performs direct x-ray detection without any light-emitting conversion steps. It produces high contrast, high-resolution digital images while also eliminating the processing time, image storage, and image transfer drawbacks associated with screen film-based imaging.
System design and performance
A 500-mm-thick layer of amorphous selenium is deposited atop a 2560 x 3072 lithographically fabricated array of thin film transistors (see Figure 2). The selenium layer is topped by a dielectric layer capped by a top electrode made of chromium. This sandwich structure effectively forms a 2560 x 3072 detector element array in which each 139 µm x 139 µm pixel detector element is electrically equivalent to a circuit of three capacitors in series. The gate control lines are arranged in rows and the image charge output lines are arranged in columns.
FIGURE 2. A layer of amorphous selenium applied over an array of thin film transistors creates a 2560 x 3072 element detector array for direct, digital x-ray imaging (left). Each detector element is equivalent to three capacitors in series (right). When a bias voltage is applied, x-rays impinging on the selenium photoconductive layer generate electron hole pairs that are separated by an applied electrical field and collected by the detector elements.
In operation, x-rays pass through the top electrode and the dielectric layers to impinge upon the selenium. Under a bias voltage of up to 5 KV, electron hole pairs proportional to the intensity of x-rays absorbed are generated in the selenium. The applied electric field separates the charges, which are collected by each detector element. The current form of the detector array consists of two 14 x 8.5 inch panels butt-coupled together to form a 14 x 17 inch detector array (see Figure 3). Data is captured at 14 bit depth and digitized to 12 bits per pixel detector element; the resulting image is 16 megabytes.
FIGURE 3. Detector array for direct x-ray radiography consists of two 14 x 8.5 inch panels butt-coupled together to form a 14 x 17 inch panel that can be installed in a medical x-ray table or wall unit.
Photo courtesy of Sterling Diagnostic Imaging
In tests, system modulation transfer function over the frequency range covering normal diagnostic imaging operations was greater than 0.7 (see Figure 4). The detector array performed appreciably better than conventional screen film systems operating even at slow speeds.
FIGURE 4. Modulation transfer function of the detector array is greater than 0.7 over the frequency range applicable to diagnostic imaging.
The quantum efficiency of the detector array was measured at approximately 35 percent. This compares favorably to screen/film technology at less than 25 percent. Comparison to other solid-state detectors has yet to be made; published data for prototype solid-state detectors have indicated higher quantum efficiencies but with differing measurement criteria. Radiologists may have difficulty understanding what quantum efficiency really means. A more useful measure of detector efficiency may be to equate patient radiation exposure factors of one technology to another. With the detector technology, the exposure factors are equivalent to those that are used if the patient were imaged with a conventional 400-speed screen/film system.
The detector array is designed for integration with new and existing general radiographic equipment. It can be installed in either a conventional x-ray table or a wall unit. To minimize image degradation caused by secondary x-ray scatter from the patient, a reciprocating grid or an air gap exposure technique must be used. The technology is available as either an OEM component or in complete chest and general-purpose x-ray systems.
About the author: James Culley and J.E.D. Armstrong are with Sterling Diagnostic Imaging Inc., P.O. Box 6101, Newark, DE 19714-6101. Phone: 302-631-3446; Fax: 302-631-3483; e-mail: armstrj@sterlingdi.com.