Principles of X-Ray Imaging

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1 Principles of X-Ray Imaging 1 Already a few weeks after the discovery of X-rays in 1895 by Wilhelm Conrad R ontgen the first medical images with photographic plates and fluorescent screens were made. This was the origin of projection radiography and fluoroscopy. The greatest steps forward in X-ray diagnostic radiology since Roentgen s observations were the development of the image intensifier systems and then above all the announcement of computed tomography (CT) in a clinical environment by Hounsfield at the 1972 British Institute of Radiology annual conference. A further important step was the introduction of digital image receptors in projection radiography during the last years. Compared to conventional film-screen systems these receptors allow the separate optimisation of photon detection and image processing, resulting in significant advantages for image quality and dose. Although today projection radiography is still the most frequent examination with X-rays the use of computed tomography increases rapidly, and because it involves larger radiation doses than the conventional imaging procedures (cf. Table 10.1) contributes significantly to the annual collective dose (see Fig. 1.1). Therefore CT also obtains growing attention in radiation protection (Brenner and Hall 2007). In X-ray diagnostic radiology the image is generated by the interaction of X-ray photons, which have transmitted the patient, with a photon detector. These photons can either be primary photons, which have passed through the tissue without interacting, or secondary photons, which result from an interaction along their path through the patient. The secondary photons will in general be deflected from their original direction and result in scattered radiation. The basic principles of projection radiography/fluoroscopy and CT are shortly explained in Sects. 1.1 and 1.2 respectively. Although totally different in image character, both imaging systems have in common certain features, which can be recognised in Fig. 1.2: 1. X-rays are produced in an X-ray tube. 2. The energy distribution of the photons is modified by inherent and additional filtration. 3. The X-rays are attenuated differently by the various body tissues. 4. Scattered radiation, which impairs image contrast, is reduced. 5. The transmitted photons are detected. 6. The image is processed and in the case of CT reconstructed. This makes it possible to discuss the aspects of image quality and radiation exposure for both systems together in the main parts of the book (cf. Chap. 2). In radiography/fluoroscopy with digital image receptors and in computed tomography the digital image consists of a (typically square) matrix of picture elements (pixels) which represent the corresponding volume elements (voxels) and after the exposure carry the local intensity information (gray scale value). Quality of digital images depends primarily on the image matrix size, i.e. the pixel size (cf. Chap. 9). As the matrix size is increased resolution improves but the number of photons in each pixel must be increased in order to maintain a certain minimum noise level. 1.1 Projection Radiography and Fluoroscopy In projection radiography and fluoroscopy the image is a two-dimensional projection of the attenuating properties of all the tissues along the paths of the X-rays. The components of a typical radiographic/fluoroscopic system are shown in Fig The photons emitted by the X-ray tube are collimated by a beam-limiting device. Then they enter the patient, where they may be scattered, absorbed or transmitted without interaction. The primary photons recorded by the image receptor form the image. The secondary photons create a certain amount of background radiation which degrades contrast. If necessary, the majority of the scattered photons can be removed by placing an anti-scatter device between the patient and the image receptor. This device can simply be H. Aichinger et al., Radiation Exposure and Image Quality in X-Ray Diagnostic Radiology, DOI / _1, # Springer-Verlag Berlin Heidelberg

2 4 1 Principles of X-Ray Imaging Angiography and intervention 2% Mammography 4% 3% GI and urogen and bile tract Remainder CT 0.7% 7% Dental 37% Dental 0.2% Remainder 1% 3% Thorax 9% Skeleton GI and urogen and bile tract 8% 1% Mammography Skeleton 33% CT 60% 18% Angiography and intervention Thorax 13% Fig. 1.1 Contribution of various examination types to total frequency (left) and to collective effective dose (right) in 2006 for Germany adapted from BMU (2009) X-ray tube x Collimation Production of X-rays Filtration Object transmission Scatter reduction Patient Photon detection Image reconstruction and processing Fig. 1.2 Basic principles of radiography/fluoroscopy and CT imaging Patient support Anti-scatter device AEC system Image receptor Fig. 1.3 Typical arrangement of a radiography/fluoroscopy system

3 1.2 Computed Tomography 5 an air gap or a so-called anti-scatter grid formed from a series of parallel metal strips. An automatic exposure control system (AEC) provides for the correct exposure of the image receptor. Today digital image receptors predominate in radiography and fluoroscopy, but film-screen systems and image intensifiers are also still in use. 1.2 Computed Tomography Whereas it is not possible in projection radiography to gain any depth information from a single image, computed tomography separates the superimposed anatomical details and produces sectional or axial slice images with excellent soft tissue contrast. Compared to projection radiography and fluoroscopy computed tomography is a rather new imaging technique. Therefore it seems to be reasonable to present its fundamental principles in some more detail. The principle of computed tomography is illustrated in Fig A well-collimated X-ray pencil beam is attenuated by the tissues along its path and the transmitted radiation is detected. In order to generate one projection the tube-detector assembly scans the object in a linear translatory motion. This procedure is repeated at many viewing angles (typically at least 180 projections are received with a rotational increment of 1 ). From these projections a two-dimensional discrete distribution of the linear attenuation coefficients m tissue is reconstructed as image signal by computation. In practice CT numbers or Hounsfield units are used instead of m tissue where the Hounsfield unit HU is defined by: HU ¼ 1000 ðm tissue m water Þ m water (1.1) where m water is the linear attenuation coefficient of water. The experimental set-up of Hounsfield corresponded largely to the arrangement sketched in Fig This set-up was termed the first generation of CT (Kalender 2006). To speed up scanning and to utilise the available X-ray power more efficiently the first commercial scanners (the second generation ) used some more detectors and a small fan beam. The typical scan time for an image matrix was 5 min (Kalender 2006). Continuously rotating CT systems ( third generation ) according to Fig. 1.5 with a fan beam covering the total patient cross-section and a corresponding detector array, consisting of gas proportional detectors or scintillation detectors (cf. Sect. 8.2), were introduced in the 1980s. Continuous rotation was made possible by a slip-ring technology for electrical power supply and data acquisition. Scan time was reduced down to 2 s for a single slice with a matrix. A major step forward in CT technology was the introduction of spiral or helical CT by Kalender and Vock in 1989 (Kalender et al. 1989; Vock et al. 1989): Slice-by-slice imaging was replaced by volume scanning. The principle X-ray tube Detector Fig. 1.4 Principle of data acquisition in CT imaging (Adapted from Bunke 2003) Fig. 1.5 Continuously rotating CT system with a fan beam and corresponding detector array (Adapted from Bunke 2003)

4 6 1 Principles of X-Ray Imaging Fig. 1.6 Principle of spiral CT imaging (From Bunke 2003) of this method is illustrated in Fig. 1.6: While the fan beam is continuously rotating the patient is moved with constant velocity along his body axis (the z-axis) through the gantry; this results in a spiral track of the focal spot around the patient and accordingly in a spiral data set. Direct image reconstruction from these data would give rise to image artefacts (similar to motion artefacts). This can mostly be avoided by data interpolation. The interpolation method developed at first was the 360 linear interpolation (LI) algorithm, which used data from a full rotation of the tube-detector assembly. Since for a complete interpolated data set at a definite slice position two successive 360 rotations on either side of the selected plane were necessary, considerable widening of the slice profile resulted, thus reducing image quality. Therefore the 360 LI was soon replaced by a 180 LI where interpolation from opposing 180 points reduces the spiral range used for reconstruction. This is possible since X-ray beam attenuation at a distinct rotation angle j is equivalent to the X-ray beam attenuation traversing the body from the opposite side, at j. As the distance of the data points is now smaller, effective slice width will be less. In 1992 CT scanners were introduced, which used two parallel banks of detectors. This was followed by multirow detector CT scanning in 1998 using solid detectors and simultaneously imaging four slices in each rotation of the X-ray tube (Kalender 2006). A great advantage of multislice CT (MSCT) scanners over single section spiral CT is the opportunity for longer anatomic coverage during the same scanning time. [ mm ] Fig. 1.7 Adaptive array detector with detector combinations for different slice thicknesses (from Bunke 2003), e.g. the uppermost combination allows slice widths in the longitudinal direction from 1 to 5 mm at the isocentre The MSCT detector arrays could be divided into two groups: Those with detector elements of unequal width along the z-axis (adaptive array detector) and those with elements of equal width (linear or matrix detector). Figure 1.7 shows as an example an adaptive array detector with the possibility of the setting of different slice thicknesses. At present 64-slice scanning represents the state of the art, allowing the imaging of all body regions with submillimetre isotropic spatial resolution and scan times of 5 15 s (Kalender 2006).

5 References 7 Scan time can be further reduced with recent developments of CT such as dual source CT or cone beam CT with C-arm systems (Kalender 2006). This is especially interesting for cardiac imaging, angiography and interventions. Dual source CT scanners are equipped with an ultrafast dual detector system and two X-ray tube assemblies. Cone beam CT (see Sect. 8.2) uses a flat-panel-detector with up to 1,920 rows and 2,480 columns (Oppelt 2005). This enables enhanced use of X-ray quanta, but also leads to a higher fraction of scattered radiation (see Sect ). Kalender WA (2006) X-ray computed tomography. Phys Med Biol 51: R29 R43 Kalender WA, Seissler W, Vock P (1989) Single-breath-hold spiral volumetric CT by continuous patient translation and scanner rotation. Radiology 173:414 BMU (Bundesministerium f ur Umwelt, Naturschutz und Reaktorsicherheit) (2009) Umweltradioaktiuit at und Strahlenbelastung im Jahr 2008: Unterrichtung durch die Bundesregierung Oppelt A (ed) (2005) Imaging systems for medical diagnostics. Publicis, Erlangen Vock P, Jung H, Kalender WA (1989) Single breathhold spiral volumetric CT of the lung. Radiology 173:400 References Brenner DJ, Hall EJ (2007) Computed tomography an increasing source of radiation exposure. N Engl J Med 357: Bunke J (2003) Computertomographie. In: Schmidt T (ed) Strahlenphysik Strahlenbiologie Strahlenschutz. Springer, Berlin, pp 84 98

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