Solid-state detectors collect the charge generated by x rays in a solid semiconducting material (Fig. 4-3). The key clinical feature of these detectors is the rapid availability of the image after exposure. The matrix and its associated readout and amplifying electronics of intraoral detectors are enclosed within a plastic housing to protect them from the oral environment. These elements of the detector consume part of the real estate of the sensor so that the active area of the sensor is smaller than its total surface area. Sensor bulk, although reduced by continued miniaturization of the electronic components, is a potential drawback of intraoral solid-state detectors. In addition, most detectors incorporate an electronic cable to transfer data to the computer. The presence of a cable can make positioning of the sensor more challenging and requires some adaptation. It also results in increased vulnerability of the device. Manufacturers have addressed these issues in various ways. Some manufacturers have changed the location of the cable attachment to the corner of the sensor. Others offer sensors with magnetic connectors, induction connectors, or reinforced cables to reduce accidental damage to the device. Wireless radiofrequency transmission also has been introduced to eliminate the cable altogether. Wireless radiofrequency transmission frees the detector from a direct tether to the computer, but it necessitates some additional electronic components, thus increasing the overall bulk of the sensor.

Many manufacturers produce detectors with varying active sensor areas roughly corresponding to the different sizes of intraoral film. Detectors without flaws are relatively expensive to produce, and the expense of the detector increases with increasing matrix size (total number of pixels). Pixel size ranges from less than 20 to 70 micrometers (µm). Three types of solid-state sensors are in common use.

Charge-Coupled Device

The charge-coupled device (CCD), introduced to dentistry in 1987, was the first digital image receptor to be adapted for intraoral imaging. The CCD uses a thin wafer of silicon as the basis for image recording. The silicon crystals are formed in a pixel matrix (Fig. 4-4). When exposed to radiation, the covalent bonds between silicon atoms are broken, producing electron-hole pairs (Fig. 4-5). The number of electron-hole pairs that are formed is proportional to the amount of exposure that an area receives. The electrons are attracted toward the most positive potential in the device, where they create “charge packets.” Each packet corresponds to one pixel. The charge pattern formed from the individual pixels in the matrix represents the latent image (Fig. 4-6). The image is read by transferring each row of pixel charges from one pixel to the next in a “bucket brigade” fashion. As a charge reaches the end of its row, it is transferred to a readout amplifier and transmitted as a voltage to the analog-to-digital converter located within or connected to the computer. Voltages from each pixel are sampled and assigned a numeric value representing a gray level (ADC). Because CCD detectors are more sensitive to light than to x rays, most manufacturers use a layer of scintillating material coated directly on the CCD surface or coupled to the surface by fiberoptics. This scintillating material increases the x-ray absorption efficiency of the detector. Gadolinium oxybromide compounds similar to those used in rare earth radiographic screens or cesium iodide are examples of scintillators that have been used for this purpose.

CCDs have also been made in linear arrays of a few pixels wide and many pixels long for panoramic and cephalometric imaging. In the case of panoramic units, the CCD is fixed in position opposite to the x-ray source with the long axis of the array oriented parallel to the fan-shaped x-ray beam. Some manufacturers provide CCD sensors that may be retrofitted to older panoramic units. In contrast to film imaging, the mechanics for cephalometric imaging are different. Construction of a single CCD of a size that could simultaneously capture the area of a full skull would be prohibitively expensive. Combining a linear CCD array and a slit-shaped x-ray beam with a scanning motion permits scanning of the skull over several seconds. One disadvantage of this approach is the increased possibility of patient movement artifacts during the several seconds required to complete a scan.

Complementary Metal Oxide Semiconductors

Complementary metal oxide semiconductor (CMOS) technology is the basis for typical consumer-grade digital cameras. These detectors are also silicon-based semiconductors but are fundamentally different from CCDs in the way that pixel charges are read. Each pixel is isolated from its neighboring pixels and is directly connected to a transistor. Similar to the CCD, electron-hole pairs are generated within the pixel in proportion to the amount of x-ray energy that is absorbed. This charge is transferred to the transistor as a small voltage. The voltage in each transistor can be addressed separately, read by a frame grabber, and stored and displayed as a digital gray value. CMOS technology is widely used in the construction of computer central processing unit chips and digital camera detectors, and the technology is less expensive than that used in the manufacturing of CCDs. Several manufacturers are using this technology for intraoral imaging applications (Fig. 4-7).

The PSP material used for radiographic imaging is “europium-doped” barium fluorohalide. Barium in combination with iodine, chlorine, or bromine forms a crystal lattice. The addition of europium (Eu⁺²) creates imperfections in this lattice. When exposed to a sufficiently energetic source of radiation, valence electrons in europium can absorb energy and move into the conduction band. These electrons migrate to nearby halogen vacancies (F-centers) in the fluorohalide lattice and may become trapped there in a metastable state. While in this state, the number of trapped electrons is proportional to x-ray exposure and represents a latent image. When stimulated by red light of around 600 nm, the barium fluorohalide releases trapped electrons to the conduction band. When an electron returns to the Eu⁺³ ion, energy is released in the green spectrum between 300 and 500 nm (Fig. 4-8). Fiberoptics conduct light from the PSP plate to a photomultiplier tube. The photomultiplier tube converts light into electrical energy. A red filter at the photomultiplier tube selectively removes the stimulating laser light, and the remaining green light is detected and converted to a varying voltage. The variations in voltage output from the photomultiplier tube correspond to variations in stimulated light intensity from the latent image. The voltage signal is quantified by an analog-to-digital converter and stored and displayed as a digital image. In practice, the barium fluorohalide material is combined with a polymer and spread in a thin layer on a base material to create a PSP. For intraoral radiography, a polyester base similar to radiographic film is used.

When they are manufactured in standard intraoral sizes, these plates provide handling characteristics similar to intraoral film. PSP plates are also made in sizes commonly used for panoramic and cephalometric imaging. Some PSP processors accommodate a full range of intraoral and extraoral plate sizes. Other processors are limited to intraoral or extraoral formats.

Before exposure, PSP plates must be erased to eliminate residual images from prior exposures. This erasure is accomplished by flooding the plate with a bright light; this is done by placing plates on a dental viewbox with the phosphor side of the plates facing the light for 1 or 2 minutes. More intense light sources can be used for shorter periods of time. Inadequate plate erasure results in double images and usually renders the image undiagnostic. Some PSP systems integrate automatic plate-erasing lights. Erased plates are placed in light-tight containers before exposure. In the case of intraoral plates, sealable polyvinyl envelopes that are impervious to oral fluids and light are used for packaging. For large-format plates, conventional cassettes without intensifying screens are used. After exposure, plates should be processed as soon as possible because trapped electrons spontaneously release over time. The rate of loss of electrons is greatest shortly after exposure. The rate varies depending on the composition of the storage phosphor and the environmental temperature. Some phosphors lose 23% of their trapped electrons after 30 minutes and 30% after 1 hour. Because loss of trapped electrons is uniform across the plate surface, early loss of charge does not typically result in clinically meaningful image deterioration. However, underexposed images may have noticeable image degradation. Adequately exposed images may be stored for 12 to 24 hours and retain acceptable image quality. A more important source of latent image fading is exposure to ambient light during plate preparation for processing. A semidark environment is recommended for plate handling. The more intense the background light and the longer the exposure of the plate to this light, the greater the loss of trapped electrons and the more degraded the resultant image. Red safelights found in most darkrooms are not safe for exposed PSP plates, which are most sensitive to the red light spectrum.

Contrast resolution is the ability to distinguish different densities in the radiographic image; this is a function of the interaction of the following:

Current digital detectors capture data at 8, 10, 12, or 16 bits. The bit depth is a power of 2 (Fig. 4-9). This means that the detector can theoretically capture 256 (2⁸) to 65,536 (2¹⁶) different attenuation levels. In practice, the actual number of meaningful attenuation levels that can be captured is limited by inaccuracies in image acquisition—that is, noise. Regardless of the number of attenuation differences that a detector can capture, conventional computer monitors are capable of displaying a gray scale of only 8 bits. Because operating systems such as Windows reserve many gray levels for the display of system information, the actual number of gray levels that can be displayed on a monitor is 242. A more important limiting factor is the human visual system, which is capable of distinguishing only about 60 gray levels at any time under ideal viewing conditions. Considering the typical viewing environment in the dental operatory, the actual number of gray levels that can be distinguished decreases to less than 30. Human visual limitations are also present for film viewing; however, the luminance (brightness) of a typical radiograph viewbox is much greater than that of a typical computer display. Therefore the ambient lighting of the room in which the image is viewed theoretically has a lower impact on film than on digital displays.

Spatial resolution is the capacity for distinguishing fine detail in an image. Resolution is often measured and reported in units of line pairs per millimeter. Test objects consisting of sets of very fine radiopaque lines separated from each other by spaces equal to the width of a line are constructed with a variety of line widths (Fig. 4-10). A line and its associated space are called a line pair (lp). At least two columns of pixels are required to resolve a line pair, one for the dark line and one for the light space. Typical observers are able to distinguish about 6 lp/mm without benefit of magnification. Intraoral film is capable of providing more than 20 lp/mm of resolution. Unless a film image is magnified, the observer is unable to ap/>