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(BQ) Part 2 book “Physics for diagnostic radiology” has contents: Diagnostic imaging with radioactive materials, positron emission tomographic imaging, radiobiology and generic radiation risks, diagnostic ultrasound, magnetic resonance imaging, multiple choice questions,… and other contents.
10 Diagnostic Imaging with Radioactive Materials F I McKiddie SUMMARY This chapter covers the following aspects of imaging with radioactive materials: • Requirements of imaging systems and techniques for obtaining accurate data • Principles of operation of the gamma camera • Additional features of modern gamma camera systems • Parameters influencing image quality • Gamma camera performance • Data display and storage • Methods of data acquisition • Quality control of the gamma camera and other aspects of nuclear medicine CONTENTS 10.1 Introduction 338 10.2 Principles of Imaging 339 10.2.1 The Gamma Camera 340 10.2.1.1 The Detector System 341 10.2.1.2 The Collimator 342 10.2.1.3 Pulse Processing .345 10.2.1.4 Correction Circuits 346 10.2.1.5 Image Display 347 10.2.2 Additional Features on the Modern Gamma Camera 347 10.2.2.1 Dual Headed Camera 347 10.2.2.2 Whole Body Scanning 347 10.2.2.3 Tomographic Camera 349 10.2.2.4 The Cardiac Camera 349 10.3 Factors Affecting the Quality of Radionuclide Images 349 10.3.1 Information in the Image and Signal to Noise Ratio 350 10.3.2 Choice of Radionuclide 351 10.3.3 Choice of Radiopharmaceutical 353 10.3.4 Performance of the Imaging Device 354 10.3.4.1 Collimator Design 354 10.3.4.2 Intrinsic Resolution 354 337 338 Physics for Diagnostic Radiology 10.3.4.3 System Resolution 355 10.3.4.4 Spatial Linearity and Non-uniformity 355 10.3.4.5 Effect of Scattered Radiation 356 10.3.4.6 High Count Rates 358 10.3.5 Data Display 358 10.3.5.1 Persistence Monitor 358 10.3.5.2 Display and Hard Copy 358 10.3.5.3 Grey Scale versus Colour Images 359 10.4 Dynamic Investigations 359 10.4.1 Data Analysis 359 10.4.1.1 Cine Mode 360 10.4.1.2 Time-Activity Curves 360 10.4.1.3 Deconvolution 361 10.4.1.4 Functional Imaging 361 10.4.2 Camera Performance at High Count Rates 363 10.5 Single Photon Emission Computed Tomography (SPECT) 364 10.6 Quality Standards, Quality Assurance and Quality Control 366 10.6.1 Radionuclide Calibrators and Accuracy of Injected Doses 367 10.6.2 Gamma Camera and Computer 369 10.7 Conclusions 370 References 371 Exercises 372 10.1 Introduction Nuclear medicine is popularly understood to be the use of radioactive materials to produce diagnostic images of biochemical processes within the body Although the wider term includes all applications of radioactivity in diagnosis and treatment, excluding sealed source radiotherapy, the general perception is taken to mean diagnostic imaging in vivo However, this does not mean that the in vitro, or non-imaging techniques are insignificant These involve the measurement of samples taken from the patient and are around 7% of the workload in a typical UK department (Hart and Wall 2005) The samples can be blood, breath, urine or faeces and are labelled with both gamma and beta emitting radionuclides The requirement for accurate mathematical models of the processes under investigation in many in vitro tests ensures that the results are absolute measures of physiological processes such as glomerular filtration rate For further details of the range of in vitro tests see Elliott and Hilditch (2005) The primary requirement in in vivo diagnostic imaging is the ability to obtain information concerning the spatial distribution of activity within the patient This chapter deals with the physical principles involved in obtaining diagnostic quality images after a small quantity of radioactive material has been administered to the patient in a suitable form The basic requirements of a good imaging system are as follows: A device that is able to use the radiation emitted from the body to produce high resolution images, supported by electronics, computing facilities and displays that 339 Diagnostic Imaging with Radioactive Materials will permit the resulting image to be presented to the clinician in the manner most suitable for interpretation A radionuclide that can be administered to the patient at sufficiently high activity to give an acceptable number of counts in the image without delivering an unacceptably high dose of radiation to the patient A radiopharmaceutical, that is a radionuclide firmly attached to a pharmaceutical, that shows high specificity for the organ or region of interest in the body It is important to recognise that, when detecting in vivo radioactivity, sensitivity and spatial resolution are mutually exclusive (see Figure 10.1) The arrangement on the left (Figure 10.1a) has high sensitivity because a large amount of radioactivity is in the field of view of the detector, but poor resolution The arrangement on the right (Figure 10.1b) has better resolution but correspondingly lower sensitivity Since gamma rays are emitted in all directions, the collimator ensures that the image is only made up of those events travelling perpendicular to the detector This preserves the relationship between the position within the patient from which the gamma ray was emitted, and its position of interaction in the detector In diagnostic imaging spatial resolution is important and sensitivity must be sacrificed A modern gamma camera (see Section 10.2.1) records no more than in 104 of the gamma rays emitted from that part of the patient within the field of view of the camera Furthermore, any additional loss of counts in the complete system will result in an image of inferior quality unless the imaging time is extended to compensate Therefore this chapter also considers the factors that limit image quality and the precautions that must be taken to optimise the images obtained using strictly controlled amounts of administered activity and realistic imaging times 10.2 Principles of Imaging Medium energy gamma rays in the range 100–200 keV are most suitable for in vivo imaging Lower energy gamma rays are stopped in the body resulting in an undesirable patient dose, whilst higher energy gamma rays are difficult to stop in the detector This will be discussed further in Section 10.3 where factors affecting the quality of radionuclide images are considered (a) (b) Detectors Collimators probably of lead Extended sources of radioactivity FIGURE 10.1 Collimator design showing conflicting requirements of sensitivity and resolution Arrangement (a) where the detector has a wide acceptance angle will have high sensitivity but poor resolution, whereas arrangement (b) will have much better resolution but greatly reduced sensitivity 340 Physics for Diagnostic Radiology In all commercial equipment currently available, the radiation detector is a scintillation crystal of sodium iodide doped with about 0.1% by weight of thallium-NaI (Tl) The fundamental interaction process in a scintillation detector is fluorescence which was discussed in Section 5.3 The sodium iodide has a high density (3.7 × 103 kg m–3) and since iodine has a high atomic number (Z = 53) the material has a high stopping efficiency for gamma rays Furthermore, provided the gamma ray energy is not too high, most of the interactions are by the photoelectric effect (see Section 3.4.2) and result in a light pulse proportional to the gamma ray energy This is important for discriminating against scatter (see Section 10.3.4) The thallium increases the light output from the scintillant, because the traps generated by thallium in the NaI lattice are about eV above the band of valence electrons so the emitted photon is in the visible range and about 10% of the gamma ray energy is converted into light This yields about 4000 light photons at a wavelength of 410 nm from a 140 keV gamma ray Note that whereas the number of photons emitted is a function of the energy imparted by the interaction, the energy or wavelength of the photons depends only on the positions of the energy levels in the scintillation crystal Finally, the light flashes have a short decay time, of the order of 0.2 µs Thus the crystal has only a short dead time and can be used for quite high counting rates One disadvantage of the NaI (Tl) detector is that it is hygroscopic and thus must be placed in a hermetically sealed container Also the large crystals in gamma cameras are easily damaged by thermal or physical shocks Alternative scintillation detectors are caesium iodide doped with thallium, and bismuth germanate Like NaI (Tl), the latter has a high detection efficiency, and is the commonest detector in positron emission tomography (PET) systems It has the higher stopping power required for the high energy gamma rays and it has a short decay time, allowing it to cope with the high count rates encountered in the absence of a collimator Bismuth germanate detectors also exhibit a good dynamic range and long-term stability The light signal produced by a scintillation crystal is too small to be used until it has been amplified and this is almost invariably achieved by using a photomultiplier tube (PMT) The main features of the PMT coupled to a scintillation crystal were discussed in Section 4.8 To isolate the output pulses from the PMT corresponding to the photopeak energy of the radionuclide being imaged, the technique of pulse height analysis is used (see Section 4.9) For a radionuclide emitting monoenergetic gamma rays, pulse height analysis should, in principle, discriminate completely between scattered and unscattered rays When a 140 keV gamma ray from technetium-99m interacts with an NaI (Tl) crystal, it does so primarily by the photoelectric effect This produces a number of visible photons and, hence, a final signal that is proportional to the gamma ray energy Any photon that has been scattered in the patient by the Compton effect will be of lower energy and will produce a smaller pulse that can be identified and rejected If an incident pulse is accepted by the pulse height analyser a signal is passed to the computer system and a ‘count’ is registered Note that there is further discussion on this point in Section 10.3.4.5 10.2.1 The Gamma Camera Modern gamma camera systems consist of one or two collimated detectors mounted on a gantry connected to a desktop-computer (PC) based acquisition and processing terminal The gantry is also intricately linked to the patient couch and the combination is designed to allow the detectors to manoeuvre freely around the patient This allows the detectors to obtain static images of any part of the body, or to track over the entire length of the patient’s body to obtain what are known as whole body images The commonest gantry design is the ring gantry which was developed from the slip-ring technology introduced 341 Diagnostic Imaging with Radioactive Materials in computed tomography (CT) This also allows the detectors to be rotated around the patient in up to a 540o arc to obtain tomographic image data The collimation of the detectors allows the spatial relationship between the point of emission of a gamma ray in the patient and the point at which it strikes the crystal to be established (see Figure 10.2) Note that unlike a grid in conventional radiology, the collimator in radionuclide imaging has no role in discriminating against scatter within the patient The function is purely to ensure that all photons incident on the crystal are travelling perpendicular to the crystal (or nearly so) when they interact The detectors on modern gamma cameras are generally rectangular with a crystal of approximately 400 mm × 500 mm Up to 100 PMTs will be arranged in a close packed hexagonal array behind the crystal to improve spatial resolution As shown in Figure 10.3, the number of photons reaching each PMT, and hence the strength of the signal, will be determined by the solid angle subtended by the event at that PMT Hence, by analysing all the PMT signals, it is possible to determine the position of the gamma ray interaction in the crystal Essential features of the gamma camera may be considered under five headings 10.2.1.1 The Detector System Components of the detector system are shown in Figure 10.4 In the gamma camera, crystal thickness must be a compromise A very thin crystal reduces sensitivity whereas a very thick crystal degrades resolution (see Figure 10.5) A camera crystal is typically 6–12 mm thick, with most manufacturers now choosing a mm thickness as optimal Light scintillation NaI (TI) crystal Parallel hole collimator Small radioactive source Emitted gamma rays FIGURE 10.2 Use of a collimator to encode spatial information In the absence of the collimator radiation from the source may strike any point in the crystal A ΩA Photoelectric event in crystal B C D PMTs ΩC NaI(TI) crystal Incident gamma rays FIGURE 10.3 Use of an array of PMTs to obtain spatial information about an event in an NaI (TI) crystal Light photons spread out in all directions from an interaction and the signal from each PMT is proportional to the solid angle subtended by the PMT at the event The signal from PMT A is proportional to ΩA and much greater for the event shown than the signal from PMT C which is proportional to Ω C 342 Physics for Diagnostic Radiology Display Accept Amplified PMT signals X Y Pulse height analyser Z pulse Correction circuits Pulse processing/ electronics Lead shielding PMTs Light guide NaI(TI) scintillation crystal Lead multiparallel hole collimator 2 Localised source of activity Tissue equivalent absorbing and scattering material FIGURE 10.4 Basic components of a gamma camera detector system The fates of photons emitted from the source may be classified as follows: (1) useful photon, (2) oblique photon removed by collimator, (3) scattered photon removed by pulse height analyser, (4) absorbed photon contributing to patient dose but giving no information, (5) wasted photons emitted in the wrong direction Thin crystal P C P P Thick crystal P P2 C P P Flux of gamma ray photons from patient FIGURE 10.5 Interactions of gamma rays with thin and thick NaI (TI) crystals P = photoelectric absorption C = Compton scattering With a thin crystal, many photons may pass through undetected, thereby reducing sensitivity With a thick crystal the image is degraded for two reasons First, the distribution of light photons to the PMTs for an event at the front of the crystal such as P1 will be different from the distribution for an event at the rear of the crystal such as P2 Second, scatter in the crystal degrades image quality since the electronics will position ‘the event’ somewhere between the two points of interaction in the crystal As shown in Table 10.1 a 12.5 mm crystal stops most of the 140 keV photons from technetium-99m (Tc-99m), the most widely used radionuclide in nuclear medicine (see Section 10.3.2) However, it can also be seen that these crystals are less well suited to higher energies The detector system is protected by lead shielding to stop stray radiation 10.2.1.2 The Collimator The most common type of collimator, which has parallel holes, is shown in Figure 10.6a It consists of a thick lead plate in which a series of small holes has been microcast or 343 Diagnostic Imaging with Radioactive Materials TABLE 10.1 Stopping Capability of a 12.5 mm Thick NaI (Tl) Crystal for Photons of Different Energy Photon Energy keV Interactions % 80 140 200 350 500 100 89 60 23 15 Object plane 10 10 (c) 5 10 10 10 (b) 5 (a) 10 Image plane FIGURE 10.6 The effect of different collimator designs on image appearance (a) The parallel hole collimator produces the most faithful reproduction of the object (b) The diverging collimator produces a minified image but is useful when the required field of view is bigger than the detector area (c) The pinhole collimator produces an enlarged inverted image and is useful for very small fields of view constructed from stacks of corrugated foil The axes of the holes are perpendicular to the face of the collimator and parallel to each other Performance of the collimator will be determined primarily by its resolution and sensitivity As shown in Figure 10.7 long narrow holes will produce high resolution but low sensitivity so these two variables work against each other A typical low energy general purpose collimator will have a resolution of mm and a sensitivity of around 150 cps per megabecquerel However, a typical low energy high resolution collimator will have a resolution of mm and a sensitivity of around 100 cps per megabecquerel This emphasises the non-linear relation between resolution and sensitivity in parallel hole collimators The general purpose and high resolution collimator pairs are the most widely used in routine diagnostic imaging The ‘low energy’ in their name refers to the fact that the thickness of the septa and the size of the holes are optimised for gamma rays in the 120–140 keV range As the object is moved away from the face of a parallel hole collimator, resolution deteriorates markedly so all imaging should be done with the relevant part of the patient as close as possible to the collimator face Sensitivity is relatively independent of distance from the collimator face, only decreasing if additional attenuating material is interposed Figure 10.7 also illustrates another problem Higher energy gamma rays may be able to penetrate the septa and this will cause serious image degradation Thicker septa are now required and for adequate sensitivity this also means larger holes and correspondingly poorer resolution 344 Physics for Diagnostic Radiology NaI (TI) crystal l Collimator s 2r Gamma ray Patient FIGURE 10.7 Diagram showing that oblique gamma rays will pass through many lead strips, or septa, before reaching the detector Typical dimensions for a low energy collimator are l = 25 mm, 2r = mm, s = 0.2 mm The number of holes will be approximately 15,000 Crystal c 2r t s d P FIGURE 10.8 Diagram showing the physical proportions and geometry of a parallel hole collimator with a point source positioned at P Insight Resolution and Sensitivity of a Collimator The spatial resolution of a parallel hole collimator depends on the geometry of the holes, corrected for any septal penetration If the resolution RP of the image of a point source at P (see Figure 10.8) is measured by its full width at half maximum height (FWHM) then RP = r (t e + d + c ) te where r is the hole radius and te the effective collimator thickness after septal penetration has been accounted for ⎛ 2⎞ te = t − ⎜ ⎟ ⎝ µ⎠ Diagnostic Imaging with Radioactive Materials 345 where μ is the linear attenuation coefficient for gamma rays in the collimator material The sensitivity (or geometric efficiency) of the collimator is given by ⎡ Kr ⎤ Sens = ⎢ ⎥ ⎣ te ( 2r + s) ⎦ where K is a factor dependent on the shape and pattern of the holes These equations demonstrate that with increasing distance d from the collimator face the resolution deteriorates, that is, RP increases, but the sensitivity is unaffected (assuming no attenuation) Other collimator designs are used for special purposes A converging collimator will magnify the image of a small organ (Figure 10.6b) A variation sometimes used to image the brain is a cone-beam collimator This gives improved sensitivity and resolution However, these collimators introduce distortion because the magnification factor depends on the distance from the object plane to the collimator and is therefore different for activity in different planes in the object There are also variations in resolution and sensitivity across the field of view as the hole geometry varies from being almost parallel at the centre to highly angled near the edge To image small objects a pinhole collimator which functions in a manner analogous to the pinhole camera may be useful (Figure 10.6c) The pinhole is a few millimetres in diameter and effectively limits the gamma rays to those passing through a point The ratio of the size of image to the size of object will depend on the ratio of the distance of the image plane from the hole to the distance of the object plane from the hole The latter distance must be small if reasonable magnification is to be achieved The thyroid gland is the organ most frequently imaged in this way Note that the pinhole collimator suffers from the same distortions as converging collimators 10.2.1.3 Pulse Processing Pulse arithmetic circuits convert the outputs from the PMTs into three signals, two of which give the spatial co-ordinates of the scintillation, usually denoted by X and Y, and the third the energy of the event Z (see Figure 10.4) Each PMT has two weighting factors applied to its output signal, one producing its contribution to the X co-ordinate, the other to the Y co-ordinate Several different mathematical expressions have been suggested for the shape of the weighting factors Those which give the greatest weight to PMTs nearest to the event are to be preferred since they will be the largest signals and hence least susceptible to statistical fluctuations due to noise (for fuller discussion see Sharp et al 1985) The final X and Y signals are obtained by summing the contributions from all tubes Insight Positional Signal Calculation A simple method of demonstrating the positional calculation is shown in Figure 10.9 This assumes that the field of view of each PMT is triangular, dropping to zero at the centre of each adjacent tube (see Figure 10.9a) If the signals from all the tubes are simply summed, this produces the output shown in Figure 10.9b This is the energy signal Z To obtain useful positional information, the output must vary linearly with x Therefore, the weighting factors ωj are used In the case shown in Figure 10.9c the weighting factors are ω1 = 2, ω2 = 1, ω3 = 0, ω4 = –1, ω5 = –2 346 Physics for Diagnostic Radiology PMT ω1 ω2 ω3 ω4 ω5 (a) Σj PMTj (b) ΣωjPMTj (c) j FIGURE 10.9 The use of weighting factors in a positional signal calculation This example shows a calculation for the x-axis A similar calculation would be carried out for the y-axis (a) A linear array of PMTs; (b) Simply summing the signals produces an output which is independent of x, except at the edges of the array; (c) The sum of the weighted signals produces an output which varies linearly with x As the weighting factors are energy dependent, allowance must be made for this by using a ratio circuit for the final positional calculation The positional signal for X is then expressed as X= ∑ j ω jPMTj( x,y ) ∑ jPMTj( x,y ) where the denominator is the energy signal Z The energy signal Z is produced by summing all the unweighted PMT signals This signal is then subjected to pulse height analysis as described earlier in this section and the XY signal is only allowed to pass to the processing system if the Z signal falls within the preselected energy window 10.2.1.4 Correction Circuits Image quality has been improved considerably in recent years by using microprocessor technology to minimise some of the defects that are inherent in a gamma camera Exact methods vary from one manufacturer to another, the examples given below illustrate possible approaches Spatial distortion may be corrected by imaging a set of accurately parallel straight lines aligned with either the X- or Y-axis The deviation of the measured position of each point on a line from its true position can be measured and stored as a correction matrix which may then be applied to any subsequent clinical image Similarly any variation in the energy signal with the position of the scintillation in the crystal can be determined by imaging a flood source and recording the counts in two narrow energy windows situated symmetrically on either side of the photopeak If the measured photopeak coincides exactly with the true photopeak, the counts in each energy 665 Multiple Choice Questions Question a) b) c) d) e) 4.5 4.6 4.7 4.8 4.9 4.10 5.1 5.2 5.3 5.4 5.5 5.6 5.7 5.8 5.9 6.1 6.2 6.3 6.4 6.5 6.6 6.7 6.8 6.9 6.10 7.1 7.2 7.3 7.4 7.5 7.6 7.7 7.8 7.9 8.1 8.2 T T T T T T F F T F F T T T F F T F T F T T T F T F T F T F T T T F T T T F T T T F F T T T T F T F T T F F T F F F F T T T T T F F T F T T T F T F F F F T T F F T F F F F T T F T F F F T T T T F T T T F F F T T T T T F T F T T T T F F T F T T F T T T T T T T T F F T F F T T F F T F T F T T F T T F T F F F F T T T F T F T F T T F F F F T F F F F T T F T T F 8.3 8.4 8.5 8.6 8.7 8.8 8.9 8.10 9.1 9.2 F T F T T T T T F T T F F F F F F F T F T T T F F T F T F T F F T F F T T T F T F T T F T F T F F T (continued) 666 Physics for Diagnostic Radiology Question a) b) c) d) e) 9.3 9.4 9.5 9.6 9.7 9.8 9.9 10.1 10.2 10.3 10.4 10.5 10.6 10.7 10.8 10.9 10.10 11.1 11.2 11.3 11.4 11.5 11.6 12.1 12.2 12.3 12.4 12.5 12.6 12.7 12.8 13.1 13.2 13.3 13.4 13.5 13.6 13.7 T F F T T F F F F F T T F F T F F F F F T F T T T T T F T T T F T T F F T T T T T T F T T T T F F T F F T T T T F T T T F F F T F T T F F T F F T F F T F F T F T T T F F T T T T F T F T T T T T F T T T T T F F T T F T T T F T F T T F T T F T T F F F F F T T T F T T F F F T T T F F T F F T F F F F T F F T T F F F T F T F F F T T F T T T F F T F F F F F F T F F T F F F F F T F F 13.8 13.9 13.10 13.11 13.12 13.13 14.1 14.2 14.3 14.4 F T F T T T F T T F F F T T F F T T F T T T F T T F F T F F T T T F F F T F F T T T T T T F T T T T 667 Multiple Choice Questions Question 14.5 14.6 14.7 14.8 14.9 14.10 14.11 14.12 14.13 14.14 15.1 15.2 15.3 15.4 15.5 15.6 15.7 15.8 15.9 15.10 15.11 15.12 16.1 16.2 16.3 16.4 16.5 16.6 16.7 16.8 16.9 17.1 17.2 17.3 17.4 17.5 17.6 a) b) c) d) e) F T T T F T F T F F T F T F F T F T F F T F F T F T T T F F F F F F F F T T T T F T T F T T F F F F T T T T T T F T T T T F T T F T T T T T T F T F T F T F T F T T F T F F F T T T T T F F F F T T T T F T T F T T F F T F T F F F F T F T T T F F T T T F F T T F T T T T F F T F T T T F F F F T T T T F F F T T T F F F T T F T F F F T F T T T F F T T T T T T T F T T F T T 18.3 Notes Some of these notes illustrate one of the weaknesses of MCQs in this subject area It is often very difficult to set non-trivial questions that are not potentially ambiguous with deeper knowledge Thus MCQs are a good teaching aide for testing an understanding 668 Physics for Diagnostic Radiology of the subject and sometimes promote further discussion They are less satisfactory as a method of examination 1.3d) As Auger electrons (see Persson L, The Auger effect in radiation dosimetry, Health Phys 67, 471–6, 1994 1.7e) Ionising radiations are oxidising agents, for example, ferrous ions to ferric ions 2.2b) The high density prevents too much electron penetration into the anode 2.3d) kV and mAs will affect the effective spot size because they will influence the performance of the cathode focussing cup in forming a small target on the anode surface 2.6d) There is substantial self-absorption of X-rays in the anode and this will be affected by anode angle 2.8e) Patient dose may be unacceptably high 2.10c) The extra heat is lost mainly by conduction 10e) Black bodies are good emitters of radiation 2.12e) Actual kV, and hence keV, is above the K shell energy for longer 3.1b) Because of absorption edges 3.1d) Any such radiation will be of such low energy that it is absorbed within the body, not scattered 3.2b) The second half value thickness will be greater because of beam hardening 3.3d) This statement would normally be true for the whole body because of the effects of attenuation, but is not true for the Compton process itself 3.4c) Healthy lung tissue is less dense than water 3.5b) There will be a small amount of elastic scattering 3.5d) Characteristic radiation will be produced if a K shell (or higher shell) vacancy is created If this occurs in the body (with low atomic number elements) the radiation may be of too low energy to escape 3.7b) Scatter may be reduced but is never eliminated 3.8a) Some X-rays will be Compton scattered Note also that many photons will not interact at all, but these are excluded by the stem of the question 3.9b) See Section 9.2.3 3.9e) There are several reasons but the most important is that the filter, in practice, would be too thin 4.2b) Some energy will be deposited as excitation 4.3e) The voltage does not need to be very constant on the ionisation plateau 4.7d) Although the electric field will be much higher than for an ionisation chamber, the operating voltage may be similar 5.4e) Although the intensifying screen stops a higher fraction of photons than film, the number of incident photons will be greatly reduced 5.9d) Both outputs are digital images so this is not an advantage Multiple Choice Questions 6.3e) 6.5c) 669 If the penumbra are excluded the larger focal spot will give a smaller image, if the penumbra are included it will give a larger image This assumes the field size is set at the cassette, if it were set on the patient surface there might be less scatter because of less beam divergence 6a and d) Both reduce the scatter reaching the film 6.10c) A major contribution to the variations in grey scale will be quantum mottle 7.1d) Quantum noise = N ½ so it increases with dose Signal to noise ratio (usually the key consideration in imaging) also equals N ½ and improves with increasing dose 7.9e) The dominant factor will be the way in which the quantum efficiency of the receptor varies with keV This will depend on the K shell absorption edges of the constituent material 8.7c) The two are unrelated Quantum noise is caused by too few photons, afterglow is a property of the receptor 9.7b) Increasing the concentration of iodine contrast is more effective 9.7e) Each frame needs the same photon density so photon flux must be increased 9.9e) 50 µGy should not normally be exceeded, 30 µGy should suffice 10.2d) days is less than two half-lives for Mo-99 10.2e) Note the question says ‘equilibrium’, by the time the generator has returned to equilibrium, it will not matter whether the generator has been eluted or not 10.10d) This thickness of crystal will cause loss of resolution with minimal increase in sensitivity for Tc-99m 11.2b) Compton scattering, not elastic scattering 11.3a) The contributions are added in quadrature 11.3d) It is better for the PET imager 11.5d) PET imaging time is much longer 12.5e) X-rays of all energies are low LET radiations, so the quadratic term in the model for double strand breaks (see Section 12.7.1) is important for both 12.8b) At Hiroshima there was a substantial neutron component 13.1e) Filtration cannot increase any component of the spectrum, so the exit dose cannot increase with the specified conditions 13.2b) Variation is in input dose, exit dose is governed by the receptor sensitivity 13.2e) For this to be true the half value thickness would have to be 10 cm—it is much less 13.6c) See general instructions at the beginning of the questions Compression will actually reduce the amount of tissue in the beam 13.7 Many values of wT are only given to one significant figure 13.8d) See general instructions 13.9c) A reasonable mean from quite a wide range of quoted values 670 Physics for Diagnostic Radiology 13.11e) Interventional procedures may cause skin reactions but they are an adjunct to therapy, not diagnostic 13.13e) Although hereditary effects cannot be positively excluded, the risk will be very low and no higher than for adults 14.1e) The mean atomic number is similar to that of air and soft tissue 14.7e) Although this is recommended good practice, it may not be feasible—for example, mobile radiography 14.9e) See Section 9.6.4 14.10d) it is preferable to retain dexterity and work quickly 14.12e) Only if a patient received a dose ‘much greater than intended’ 15.6e) Full 2D is not necessary Either a lens or a few rows of detectors covering the slice-width direction (to allow focussing) will suffice 15.9b) It is the difference in frequency between the outgoing and returning signals that is in the audible range 15.9e) The word ‘power’ in the name of this mode refers to the measurement of echo power and is not a reference to the acoustic power used, which is less than in most other Doppler modes 15.12a) Improving the quality of the received image reduces scan time 16.7a) The amount of C-11 is too small to give an adequate signal, C-12 gives no signal (b) FIGURE 10.19 Functional image of a normal MUGA study (a) The phase image represents the phase (or timing) of the contraction of the heart chambers (b) The amplitude image represents the amplitude of the contraction In this case it can be seen that the largest contraction occurs apically and along the lateral wall Both of these parameters are derived from the Fourier fitted curve (Figure 10.18) (a) CT transaxial CT Scout view CT coronal PET coronal Fused coronal PET transaxial Fused transaxial CT sagittal PET sagittal Fused sagittal PET MIP View FIGURE 11.1 Typical PET/CT review screen showing CT, PET and fused image data sets The bottom right hand image shows a rotating maximum intensity reprojection image FIGURE 15.38 A 3D ultrasound rendering of a foetal face (b) FIGURE 15.44 Doppler ultrasound images showing vasculature in the kidney There are two main ways of mapping Doppler shifts detected within the ‘colour box’ on the B-mode image: (a) A colour flow map (CFM) shows mean Doppler shifts at each point; (b) A power Doppler map shows the strength of Doppler-shifted echoes at each point Notice the reference colour bars to the left of each image (a) Physics SerieS editorS: John G WebSter, Slavik tabakov, kWan-hoonG nG PHYSICS FOR DIAGNOSTIC RADIOLOGY THIRD EDITION “This is the third edition of a well-established and popular textbook on physics of diagnostic radiology It is a textbook written in a clear and concise style, supported by excellent illustrations The textbook describes recent state-of-the-art advances in medical imaging in a way radiologists, radiographers and medical physicists will find easy to understand It is internationally recognised as one of the key textbooks in its field.” —Dr Keith Faulkner, North East Strategic Health Authority, UK With every chapter revised and updated, Physics for Diagnostic Radiology, Third Edition continues to emphasise the importance of physics education as a critical component of radiology training This bestselling text helps readers understand how various imaging techniques work, from planar analogue and digital radiology to computed tomography (CT), nuclear medicine and positron emission tomography (PET) to ultrasound imaging and magnetic resonance imaging (MRI) New to the Third Edition • Material on digital receptors • Emphasis on the differences between analogue and digital images • Coverage of multi-slice CT and three-dimensional resolution, dual energy applications and cone beam CT • Special radiographic techniques, including subtraction techniques and interventional radiology • New chapter on PET, with discussion of multi-modality imaging (PET/CT) • Additional material on radiation doses and risks to patients • New chapter covering the picture archiving and communication system (PACS), teleradiology, networks, archiving and related factors • A summary of the main teaching points at the beginning of each chapter 83155 ISBN: 978-1-4200-8315-6 90000 781420 083156 ... is imaged, or in non-uniformity for a uniform extended source Correction circuits for non-linearity were discussed in Section 10 .2. 1 Non-uniformity in the image of a uniform flood source is a... interest with time t is given by C(t) = a + b sin (2 ft + ϕ) 3 62 Total counts over left ventricle (arbitrary units) Physics for Diagnostic Radiology 1.0 0.5 20 0 400 Time (msec) 600 FIGURE 10.18 Time-activity... or septa, before reaching the detector Typical dimensions for a low energy collimator are l = 25 mm, 2r = mm, s = 0 .2 mm The number of holes will be approximately 15,000 Crystal c 2r t s d P FIGURE