Pausas y silencios

With regard to digital detector types, digital mammography systems are classified into two subsections: computed radiography (CR) systems; and digital radiography (DR) systems. The latter is subdivided into indirect and direct digital radiography (Lança & Silva, 2009a). However, other researchers who classified digital systems into direct digital radiography (DR) and indirect digital radiography systems which also include computed radiography (CR) (Mothiram, Brennan, Lewis, Moran, & Robinson, 2014).

3.3.1.1 Computed Radiography (CR)

CR was first introduced in the early 1980s (Lança & Silva, 2009a). This was the earliest digital technology employed for mammography (James, 2004). Here the photostimulable phosphor detectors, known as storage phosphor screens (SPS), were used instead of traditional films inside special cassettes. They were still standard size of film-screen cassettes (Lança & Silva, 2009a). These image detectors could be thus used with conventional X-ray systems (Testagrossa et al., 2012). Exposing SPS to X-rays resulted in an excitation process through the movement of valence band electrons to the conduction band, thereby forming the latent image as an electronic signal. The SPS was then scanned with a laser in the reader; this converted the latent image to blue light which was proportional to the amount of incident X- ray on the SPS. Finally, a photomultiplier system in the reader converted the blue light to an

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electronic signal which was then available as a digital dataset. To ensure that the SPS was free from any residual charge, it was then scanned with high intensity white light. Typically 25% of the latent image was lost from the SPS within 10 minutes to 8 hours (James, 2004; Lança & Silva, 2009a). The main disadvantage of this type of digital detector is the lack of spatial resolution due to light scattering. Thicker phosphor layer detectors are more sensitive to radiation but more light scattering is produced (Testagrossa et al., 2012).

3.3.1.2 Indirect Digital Mammography

The initial digital mammography (DM) systems utilised indirect conversion detectors. In this type of detector the image capturing process was achieved in two steps, using charge couple devices (CCD). The first step includes the X-ray energy to light photon conversion, which is then converted to an electronic signal in the second step (Smith, 2005). The CCD technology involves the use of phosphor, and millions of optic fibers on coupling plates. The main function of the optic fibers is to transfer light from the phosphor to the CCD. This is then digitised (Pisano & Yaffe, 2005). The early CCD was limited to a small field size of 5 cm X 5 cm. The success of these limited field size systems in digital spot mammography, for stereoscopic needle biopsy, led to the development of this technology in larger field sizes of 1cm X 22 cm, using an array of four phosphor-CCD assemblies. These detectors were synchronised with a slit collimated X-ray beam to scan the breast perpendicularly to a patient‘s body producing 22 cm X 30 cm images (James, 2004). The required image acquisition time was about 6 seconds longer than that required by large area detectors. Due to slit collimation, the scattered radiation was reduced significantly thereby eliminating the need for a grid (Pisano & Yaffe, 2005).

Amorphous silicon flat-panel detectors, commonly known as large area detectors, are another form of indirect digital radiography detectors. They were introduced to clinical use in the late 1990s (Lança & Silva, 2009a). They were made of a thallium activated caesium iodide (CsI:Tl) phosphor layer acting as an X-ray absorber, and a light sensitive two-dimensional (rectangular) array of photo-diodes. The incident X-ray beam on the detector is absorbed by the CsI:Tl layer releasing light photons. This light is then converted into an electronic signal by the photo-diodes. The electronic signal is then finally captured by thin film transistors (TFTs). Since the CsI:Tl crystals are designed in needle like channels and both the Cs and I

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have high atomic numbers (for Cs Z=55 and for I Z=53), this technique exhibits high efficiency for X-ray absorption (80-90%). CsI:Tl light is in the green area of the light spectrum where the photo-diodes have relatively high absorption efficiency (approximately 80%) (Cowen, Kengyelics, & Davies, 2008). The principle advantage of this detector type is that it can be used for radiographic procedures which require rapid sequence imaging. However, its main limitations are the high cost, that format changing is difficult, and that the detector‘s element size cannot be easily reduced (Pisano & Yaffe, 2005).

3.3.1.3 Direct Digital Mammography

Amorphous selenium-based detectors avoid multiple conversions of X-ray photon energy, from light photons to an electronic signal, via direct X-ray conversion (James, 2004). This conversion process eliminates the light scattering problem associated with indirect conversion detectors. Direct conversion digital detectors are made from a layer of amorphous selenium (a-Se) mounted on the top of the image plate which consists of a regular matrix of storage capacitors and thin film transistors (Cowen et al., 2008). Before exposing the a-Se to X-ray, its surface is charged with a uniform positive charge. The uniform surface charge pattern is partially discharged when the X-rays are absorbed by the a-Se. The amount of discharge is proportional to the energy of the absorbed X-ray photons. This charge distribution forms the latent image as an electronic signal (James, 2004). This electronic signal is read by the thin film transistor array (NHSBSP, 2009). This type of detector is exemplary for use in digital mammography because its absorption efficiency is high, it produces high resolution images, and its dose efficiency is good (Qian, 2013).