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8 2. X-ray imaging (a) Fig. 2.5: (a): CT image of a human heart including the coronary artery (taken with Philips Brilliance CT 40-channel) [14]. (b): Photograph of an open CT gantry. The picture shows the X-ray tube (T), the detector banana (D) and the fan-shaped X-ray beam (X). The direction of rotation is also indicated (R) [15]. detector rows (currently up to 256 slices [13]) placed next to each other. Furthermore, the patient is moved slowly through the center of the gantry resulting in a helical motion of the detector-tube assembly. Like this it is possible to acquire large volume scans of the patient at a time-scale of seconds, keeping motion artifacts, e.g. respiratory motion, to a minimum. However, this increased resolution comes at the cost of a significantly increased dose. A typical chest CT scan exposes the patient to approximately 8 mSv compared to the 0.02 mSv of conventional two-dimensional chest projection image [16]. In comparison, the accumulated average yearly dose including natural as well as anthropogenic sources is about 2.14 mSv. These dose considerations have a major impact on the quality of medical X-ray examinations because they directly affect the achievable image contrast. It can be shown that the X-ray dose necessary to obtain a given contrast and signal to noise ratio for a quadratic object placed in front of a homogeneous background is inversely proportional to the squared area of that object [16]. Therefore the design goal of any X-ray imaging system is always to find a balance between the lowest possible patient dose and the minimum image contrast necessary to identify anomalies in the human body. The specifications of different medical X-ray imaging modalities in Tab. 2.2 reflect these dose and image contrast considerations in terms of the average photon flux and the typical energies of the X-rays. The table illustrates that mammography systems for example need only to count 5 · 10 7 photons/(mm 2 s) at relatively low energies below 40 keV. In contrast to the static image acquisition of the mammography system, a CT imager has to acquire its single scans in rapid succession due to the quickly revolving gantry. This results in photon fluxes which are typically 20 times higher. Furthermore, in order to penetrate the thicker and denser objects a CT scanner is operating at higher X-ray energies compared to a mammography imager. Regardless of the rate considerations, the demand for high resolution projection images in mammography requires typical pixel sizes smaller than 100 µm, whereas the three dimensional reconstruction algorithm allows modern CT scanners to operate with 1 mm pixels. D R (b) X T
2.3. X-ray detector concepts 9 Count rate (photons/(mm2 ·s)) Equiv. current per pixel (nA) Pixel size (µm) Detector area (m2 ) Highest energy (kVp) Mammography General X-ray Computed radiography Tomography 5 · 10 7 10 6 to 5 · 10 8 typ. 10 9 0.2 0.03 - 20 2000 typ. 85 typ. 150 typ. 1000 0.072 0.12 0.14 (128 slices) 28 - 40 70 - 120 80 - 140 Tab. 2.2: Typical requirements of different X-ray imaging applications [4]. Tube energies given in kVp refer to the acceleration voltage of the X-ray tube. This value sets the upper boundary for the photon energies in the generated X-ray spectrum. Typical tube settings: Mammography: 28 kVp, 144 mA, 65 cm; Radiography: 70 - 120 kVp, 2 - 300 mA, 1.5 m; Tomography: 140 kVp, 400 mA, 1.04 m, including filters. 2.3 X-ray detector concepts Medical X-ray detectors have diversified into a number of different concepts. The selection of a particular type of detector for a special application like mammography, dental imaging or computed tomography is influenced by essential features like photon detection efficiency, image noise and real time capability. Nevertheless other aspects like the price of the imager and the ease of use have to be considered as well. Photographic film The earliest detector concept, which is still widely used, is the X-ray sensitive photographic film. These relatively cheap and easy to handle detectors are based on the darkening of a X-ray sensitive emulsion, which is applied to a base carrier. X-ray images are obtained through chemical processing of the films, which also prevents real time imaging with these detectors. Although the quantum detection efficiency of photographic film is only on the order of a few percent, very high spatial resolutions (µm) can be achieved by using fine grained emulsions. For reduced patient exposure it is possible to increase the X-ray sensitivity of the film by placing a thin scintillator foil on top of it. This however reduces the spatial resolution of the detectors. Apart from the lacking real time capability, the major drawback of photographic films is their non-linear response characteristic, i.e. the detectors are prone to suffer from under- and over-exposure. Typical dynamic ranges are on the order 40 dB [17]. Image intensifiers Image intensifiers are real-time imaging devices in which incident X-rays are converted first into visible light with the help of a scintillating material. A commonly used substance is CsI, which can be grown in columnar structures that effectively guide the generated optical photons onto a photocathode. This second conversion layer is placed directly on top of the scintillator and converts the incoming photons into electrons. The following amplification stage accelerates the electrons onto a small anode target, which converts the
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120 8. X-ray images C o n tr a s t
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122 8. X-ray images
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124 9. Conclusion and outlook • S
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126 9. Conclusion and outlook princ
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Bibliography [1] E. Kraft, Counting
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Bibliography 137 [32] A. E. Bolotni
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Acknowledgements Looking back at my
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